research papers on drug delivery

An official website of the United States government

The .gov means it’s official. Federal government websites often end in .gov or .mil. Before sharing sensitive information, make sure you’re on a federal government site.

The site is secure. The https:// ensures that you are connecting to the official website and that any information you provide is encrypted and transmitted securely.

• Publications
• Account settings

Preview improvements coming to the PMC website in October 2024. Learn More or Try it out now .

• Advanced Search
• Journal List
• v.9(6); 2023 Jun
• PMC10320272

Advances in drug delivery systems, challenges and future directions

Tobechukwu christian ezike.

a Department of Biochemistry, Faculty of Biological Sciences, University of Nigeria, Nsukka, 410001, Enugu State, Nigeria

b Department of Genetics and Biotechnology, Faculty of Biological Sciences, University of Nigeria, Nsukka, 410001, Enugu State, Nigeria

Ugochukwu Solomon Okpala

Ufedo lovet onoja, chinenye princess nwike, emmanuel chimeh ezeako, osinachi juliet okpara, charles chinkwere okoroafor, shadrach chinecherem eze.

c Department of Clinical Pharmacy and Pharmacy Management, Faculty of Pharmaceutical Sciences, University of Nigeria, Nsukka, 410001, Enugu State, Nigeria

Onyinyechi Loveth Kalu

Evaristus chinonso odoh.

d Department of Pharmacy, Federal Medical Center Bida, Niger State, Nigeria

Ugochukwu Gideon Nwadike

John onyebuchi ogbodo.

e Department of Science Laboratory Technology, Faculty of Physical Sciences, University of Nigeria, Nsukka, 410001, Enugu State, Nigeria

Bravo Udochukwu Umeh

Emmanuel chekwube ossai, bennett chima nwanguma, associated data.

No data was used for the research described in the article.

Advances in molecular pharmacology and an improved understanding of the mechanism of most diseases have created the need to specifically target the cells involved in the initiation and progression of diseases. This is especially true for most life-threatening diseases requiring therapeutic agents which have numerous side effects, thus requiring accurate tissue targeting to minimize systemic exposure. Recent drug delivery systems (DDS) are formulated using advanced technology to accelerate systemic drug delivery to the specific target site, maximizing therapeutic efficacy and minimizing off-target accumulation in the body. As a result, they play an important role in disease management and treatment. Recent DDS offer greater advantages when compared to conventional drug delivery systems due to their enhanced performance, automation, precision, and efficacy. They are made of nanomaterials or miniaturized devices with multifunctional components that are biocompatible, biodegradable, and have high viscoelasticity with an extended circulating half-life. This review, therefore, provides a comprehensive insight into the history and technological advancement of drug delivery systems. It updates the most recent drug delivery systems, their therapeutic applications, challenges associated with their use, and future directions for improved performance and use.

1. Introduction

Drug delivery systems are technological systems that formulate and store drug molecules into suitable forms like tablets or solutions for administration. They hasten the reach of drugs to the specific targeted site in the body, thereby maximizing therapeutic efficacy and minimizing off-target accumulation in the body [ 1 , 2 ]. Drugs have various routes through which they can be introduced into the body, they include but are not limited to the oral route of administration [ 3 , 4 ], buccal and sublingual routes of administration [ 5 ], nasal and ophthalmic [ 6 , 7 ], transdermal and subcutaneous [ 8 , 9 ], anal and transvaginal [ 10 , 11 ] and intravesical [ 12 , 13 ]. The components of the drug account for its physiochemical properties and are responsible for the changes it influences in the body system when taken.

Over the past few decades, DDS have been applied effectively in the treatment of diseases and improvement of health due to increased systemic circulation and control of the pharmacological effect of the drug. The advancement of pharmacology and pharmacokinetics showed the importance of drug release in determining therapeutic effectiveness, giving rise to the concept of controlled release [ 14 ]. The controlled-release formulation of a drug was first approved in the 1950s, and it has since attracted considerable attention due to its significant advantages over conventional drugs. It releases drugs at a predetermined rate and for a specific period of time. In addition, controlled drug delivery systems are not affected by physiological conditions and can thus last for days to years. It also provides spatial control over drug release, with constant or variable release rates [ 15 ]. Furthermore, it improves drug solubility, target site accumulation, efficacy, pharmacological activity, pharmacokinetic properties, patient acceptance, and compliance, and reduces drug toxicity [ 2 ].

Recently, several drug delivery systems (NDDS) have been developed using advanced systems for more convenient, controlled, and targeted delivery. Each drug delivery system has its own peculiarities that determine its release rate and mechanism. This is largely due to the differences in the physical, chemical, and morphological characteristics which will ultimately affect their affinities for various drug substances [ 16 ]. Studies on these have identified diffusion, chemical reaction, solvent reaction, and stimuli control as major release mechanisms [ 17 , 18 ]. For instance, since most cancer cells can proliferate the porous blood vessels and lymphatic system, the drug can easily permeate through this opening to reach the target tissues. This is referred to as Enhanced Permeability and Retention (EPR) [ 19 ]. EPR is a passive diffusion mechanism well researched and applied in the delivery of many chemotherapeutic agents. Although EPR is a controversial concept, this effect has been observed by many researchers in various types of human tumors and is the basis for the use of nanomedicine in cancer treatment. Though it has a drawback of lack of selectivity and increased toxicity [ 20 ]. Active targeting overcomes the lack of specificity and selectivity found in passive targeting. It involves attaching to the carriers, certain ligands, and molecules that can actively bind to the surface of target tissues. This prevents uptake by non-target cells thereby reducing side effects and toxicity [ 21 , 22 ]. Selectivity of ligands to target cells, immunogenicity, and chances of lysosomal degradation after macrophage endocytosis still pose solid challenges to the full development of actively targeting Drugs [ 23 ]. These delivery systems can also reach the target cells through the control of one or more physical or chemical properties in the process of responsive stimuli targeting [ 24 , 25 ]. These physical properties include pH, temperature, ultrasound, magnetic and electric field.

2. The early period of drug delivery systems

In the ancient period, people depended on medicinal plants. Although they were beneficial, they lacked consistency, homogeneity, and specificity in drug delivery [ 26 ]. Before the use of controlled drug delivery, all pharmaceuticals were produced and stored in pill or capsule formulations. It is dissolved when it comes in contact with gastrointestinal fluids, permeates the gut wall, and is then absorbed into the bloodstream through blood capillaries. There was no capacity to control the drug release kinetics. With the aim to hide the bitter taste of drugs, Rhazes and Avicenna, introduced coated technology. This coating method altered the rate of release of the drug itself. It was adopted in the 10th century however in the form of gold, silver, and pearl-coated tablets.

In the 20th century, advanced coating technology with keratin, shellac, sugar, enteric coating, and pearl coating, was also introduced, but keratin and shellac were ineffective due to storage instability and high pH for adequate dissolution in the small intestine. Malm et al. [ 27 ] introduced an enteric-coating material with polymeric cellulose acetate phthalate that is dissolved at a very weak alkaline pH, like that of the small intestine, which made it highly suitable to be applied for enteric controlled release.

The first generation was extremely productive, it focused on the development of numerous oral and transdermal controlled-release formulations for clinical use and the establishment of controlled drug-release mechanisms. In 1951, Lipowski first introduced a patent oral sustained-release formulation, when he coated pills with enteric polymers (like beads) such that the drug and coat were layered alternatively, resulting in slow release of the drug, regularly, and periodically [ 28 ]. This was further developed by Smith, Klein Beecham, and French (SKF) in 1952, they developed Spansule technology, an oral predetermined-release formulation that sustains and controls the kinetic release of a drug gradually [ 29 ]. This formula is composed of hundreds of micro pellets drug loaded beads with variable layers of natural water-soluble wax with dissimilar thicknesses on individual pellets. On ingestion, the outer capsule rapidly disintegrates, the waxy coating around the beads gradually dissolves as they transit down the GI tract, and liberates the drug-loaded beads. This improved patient compliance and convenience by reducing the dosing schedule, resulting in great popularity [ 30 ]. This technology was further developed by replacing the wax with more reproducible synthetic polymers [ 31 ].

In 1955, the first nanoparticle therapeutic was reported by Jatzkewitz when he prepared the first polymer-drug conjugate. In the 1960s, the first nanotechnology known as liposome (lipid vesicles) was discovered [ 32 , 33 ]. Polymer-drug conjugates and liposomes mark the birth of nanocarriers. In this decade, the ALZA Corporation did not create drugs they specialized in targeting and controlling the release of drugs at the right place and time [ 34 ]. In 1972, Scheffel and his colleague prepared the first protein-based microspheres. In 1976, “micelle” and “emulsion” polymerization techniques were used to prepare drug-loaded nanoparticles and microcapsules by Peter Paul Speiser's research group [ 35 ]. In 1977, Couvreur et al. [ 36 ] reported the lysosomotropic effects of the nanoparticles, and they produced the first rapidly biodegradable acrylic nanoparticles.

The drug delivery formulations developed during the second generation (2G) were impressive, but they did not produce the expected clinical results [ 37 ]. The researchers were interested in developing drug delivery systems with constant drug release rates, self-regulating, long-term depot formulations, and nanotechnology-based formulations, particularly nanoparticle formulations. In this era, long-term depot-sustained drug-release formulations of peptide/protein drugs were developed [ 38 ]. In addition, smart polymers and hydrogels were developed to stabilise drug delivery systems that are affected by physiological changes such as pH, temperature, electric field, and glucose. Furthermore, efforts were made to develop targeted nanotechnology DDS for tumors and gene delivery using biodegradable polymers in nanoparticle structures such as polymeric micelles, chitosan, lipids, and dendrimers. The idea was to modify the nanoparticles so that they could be administered directly into the body for increased drug accumulation at the site of action. Although this nanotechnology-based DDS demonstrated high efficacy in controlling tumor growth in animal models, the FDA only approved a few drugs [ 31 , 39 ].

The third generation of drug delivery systems is the modern era of controlled release technology. For it to be beneficial and successful, it has to overcome the hurdles of both physicochemical and biological barriers, associated with the earlier drug delivery systems. The physiochemical challenges are caused by poor water solubility, the high molecular weight of therapeutic proteins and peptides, and the difficulty in achieving targeted and controlled drug release, whereas the biological barrier challenges are associated with systemic drug distribution issues [ 31 , 40 ]. Many new drug delivery systems must be developed during this time period to meet the challenges associated with earlier forms of drug delivery in order to improve performance and sustainability. However, designing a suitable carrier system is often very difficult due to the challenges associated with targeting a drug to a specific site and continuous release over a specified period of time.

3. Recent drug delivery systems and applications

Significant progress has been made in recent years toward the successful development of drug delivery systems based on organic, inorganic, and hybrid nanoparticles as drug carriers for active targeting, particularly in chemotherapy. Recent drug delivery systems (DDS) are formulated with improved properties such as smaller particle size, increased permeability, increased solubility, efficacy, specific site targeting, stability, toxicity, and sustained delivery. They can significantly improve therapeutic agent performance over conventional dosage forms [ 19 , 41 ].

In the development of an optimal drug delivery system, recent drug delivery systems are recognized to be the newest developments and innovative understanding of the pharmacokinetic and pharmacodynamic behavior of pharmaceuticals. Because these DDS are transporters, they can keep medication concentrations in the therapeutic range for a long time while also delivering material to the site of action. The adoption of the delivery mechanism is directly tied to the commercial and therapeutic success of the innovation. This would entail involving patients early in the development process, recognizing any problems, and ensuring they receive the most out of the device. Improving delivery systems that reduce toxicity while increasing efficacy. The different types of drug delivery systems are depicted in Fig. 1 .

Several types of recent drug delivery systems for different therapeutic purposes.

3.1. Red blood cell membrane-camouflaged nanoparticles drug delivery system

Researchers have recognized the potential benefits of nanotechnology in vastly improving medicine delivery methods throughout time. Red blood cell membrane-camouflaged nanoparticles are a new class of drug delivery systems. The nature and biological significance of red blood cells (RBCs) allow for their use as an efficient system as a nanoparticle camouflaging material [ 42 ]. Because red blood cells (RBCs) are the most abundant circulating cells in the body, their biocompatibility (non-immunogenic), biodegradability, and extended circulating half-life, making them an ideal vehicle for drug delivery [ 43 ]. Engineered RBCs have been investigated and found to be an excellent carriers for a variety of bioactive chemicals, including enzymes, medications, proteins, and macromolecules [ 44 ]. Because of their abundance, red blood cell membranes serve as a “camouflage,” allowing nanoparticles to combine the benefits of native red blood cell membranes with those of the nanomaterial. Several strategies have been developed to load therapeutic agents onto RBCs without comprising the structure and the physiological function of RBCs. The coated nanoparticles will mimic RBCs and interact with the environment to establish long systemic circulation when injected. Sonication is the most common method for creating RBC camouflaged nanoparticles. Other methods of RBC fusion with nanoparticles include in-situ polymerization, microfluidic electroporation, and extrusion. However, each has advantages and disadvantages in terms of synthesis, scale-up challenges, reproducibility, and the nature of the final product [ 42 ]. Prior to the fusion, the RBC membrane-derived vesicle is obtained through hypotonic treatment (dialysis, hemolysis or dilutions) of fresh whole blood from an organism. The hypotonic treatment will help to remove unwanted cells and plasma ( Fig. 2 ).

Synthetic strategies for RBCM-NPs preparation for antitumor activity. Whole fresh blood is centrifuged and washed multiple times to remove the plasma and other unwanted cells. The resulting pure red blood cells are subjected to hypotonic hemolysis and are used to coat selected nanoparticles which are intravenously injected into the blood to maintain long systemic circulation. The RBCM-NPs permeate the tumor tissue via the EPR effect and finally enter into the tumor cells by endocytosis for therapeutic effect. (For interpretation of the references to colour in this figure legend, the reader is referred to the Web version of this article.)

The use of RBCM-NPs drug delivery systems is extremely promising and offers numerous benefits due to their low immunogenicity and ability to maintain long systemic circulation (a lifespan of 120 days). Furthermore, because of the large number of cell membranes, RBC vesicles are inherently biocompatible and biodegradable, and can easily achieve high load capacity, resulting in higher accumulation at the target site. Remarkably, erythrocyte membrane-coated nano-formulations have been extensively applied in anticancer research to substantial accomplishment [ 45 , 46 ], cardiovascular diseases [ 47 ], and encephalopathy [ 48 ].

3.2. Hyaluronic acid-based drug nanocarriers drug delivery systems

The usage of hyaluronic acid is one of the drug delivery techniques. Hyaluronic acid is a novel polymer that can be used to make medication delivery systems [ 49 ]. It has a linear macromolecular mucopolysaccharide made up of proportionately connected glucuronic acid and N -acetylglucosamine saccharide units [ 50 ]. It exhibits biocompatibility, biodegradability, and high viscoelasticity, and it can be coupled with a specific cell surface receptor [ 51 ]. Because Hyaluronic acid is a natural component of eye tissue and plays an important function in wound healing, it makes sense to use it as a carrier for ocular drug delivery as long as the integrated pharmaceuticals are released consistently. They aid in the thickening, sustained release, and transdermal absorption of drugs, as well as improving drug targeting. Drug distribution to cancer cells was significantly improved using active targeted HA-based drug nanocarriers. In addition, lipid nanoparticles with an appropriate HA coating have been developed as biocompatible drug carriers with a great potential for targeted drug delivery to the target tissue while minimizing side effects and harming other tissues. Benefits of utilizing HA-based nanocarriers for cancers with elevated expression of the CD44 receptor include improved drug delivery, increased therapeutic efficacy, higher cytotoxicity, and considerable reduction of tumor development, as well as a high potential for targeted chemotherapy [ 52 ].

Another application is when an HA-based nanocarrier is combined with doxorubicin (DOX) and cisplatin (CDDP) and manufactured as a CD44-targeting anti-cancer drug delivery system, as well as its tumor inhibition activities in vitro and in vivo against CD44 + breast cancer cells. In 4T1 (CD44 + ) breast cancer cells, these dual drug-loaded HA micelles (HA-DOX-CDDP) showed significantly improved drug release under acidic circumstances, as well as higher cellular uptake and stronger cellular growth suppression than free medicines. HA-DOX-CDDP micelles are a potential drug delivery system with acid-sensitive drug release, CD44-targeted delivery, and high biocompatibility and biodegradability. These characteristics resulted in excellent tumor accumulation and less side effects, indicating that HA-DOX-CDDP micelles could be useful in breast cancer chemotherapy [ 53 , 54 ].

Hyaluronic acid and its derivatives are incorporated into various drug delivery systems (DDS) such as nanoparticle DDS, cationic polymer DDS, and gel DDS for active targeting of cancer cell CD44 receptors ( Fig. 3 ). According to studies, the use of HA and drug conjugates after administration aggregates at the tumor site, where it maintains sustained drug release. The surface of HA-based nanocarriers is generally negative, which helps in blocking the systemic clearance of nanocarriers by the reticuloendothelial system (RES). The Hyaluronic acid-based drug nanocarriers NDDs selectively enter into the cancer cells through the EPR effect and active targeting of CD44 receptors [ 55 , 56 ].

The application of hyaluronic acid-based nanocarriers in cancer treatment. (a) and (b) Direct conjugation of cytotoxic drug with HA or hydrophobic moiety results in self-assembly of nanoparticles (NPs) that can be administered intravenously for cancer cell targeting; (b) HA hydrogel formation using a cationic polymer; (c) Surface coating of NP with HA; (e) Hyaluronic acid-based drug nanocarriers permeate cancerous tissues via EPR effect and binds to the CD44 receptor site to elicit anticancer activity.

3.3. Hexagonal Boron Nitride nanosheet drug delivery system

As technology continues to advance and science evolves, more and more materials are being researched and studied to help improve drug delivery. Among these materials is Boron nitride (BN) which is a crystalline material with a balanced stoichiometry of nitrogen (N) and boron (B) atoms. This material occurs in various forms such as cubic BN (c-BN), hexagonal BN (h-BN), wurtzite BN (w-BN), and rhombohedral BN (r-BN). Hexagonal Boron Nitride is a two-dimensional (2D) layered-dense structure with sp2 hybridized B–N bonds. It can also be called white graphene sometimes regarded as an analogue of graphite [ 57 ]. The B–N atoms substitute the carbon atoms and are held together by a strong covalent bond forming interlocking rings. The layers of the compound are held together by van der Waals forces with a bond length of 1.466 Å and an interlayer space of 3.331 Å. This compound is partially ionic and as a result of this unique characteristic has its B–N bonds to be polar. H-BN is an insulator that has gained wide applications in various fields of life such as cosmetics, dental, cement, ceramics, and most especially in medicine as a drug carrier similar to graphene or graphene oxide [ 58 ].

Hexagonal Boron Nitride has been proven to be useful in drug research and delivery systems ( Fig. 4 ). The research by Jedrzejczak-Silicka and her colleagues reported a reduction in the proliferation of MCF-7 cell line cultures when compared with the normal L929 cell lines upon exposure to H-BN loaded with gold particles. H-BN was exfoliated via chemical treatment using modified Hummers' method, sonication treatment, and was finally functionalized with gold particles for the studies and analyzed using Neutral Red (NR) uptake assay [ 59 ]. In another study, H-BN nanosheets were conferred photothermal properties as a result of the in-situ deposition of Pd on its surface. This made the compound have a high loading capacity for doxorubicin which is an anticancer drug as well as functions effectively as a drug delivery carrier. The study reported a remarkable inhibition of tumor growth as the drug was administered in mice for two weeks. This was possible as a result of the decrease in pH which led to the release of doxorubicin from the nanohybrids and a concomitant increase in glutathione concentration as well as near infrared radiation (NIR) [ 60 ]. Another successful study showed that H-BN conjugated with DNA oligonucleotide and copper (II) phthalocyanine (CuPc) was effective as a therapeutic agent in photodynamic therapy (PDT) and in situ monitoring and miR-21 imaging [ 57 ]. Boron compound is being recognized today as an effective chemotherapeutic agent.

Schematic explanation of H-BN nanosheet drug delivery system. H-BN was exfoliated via chemical treatment and functionalized with Au particles resulting in high loading capacity of DOX, a chemotherapy drug which is released in tumor cells by decrease in pH.

3.4. Polymer-lipid hybrid nanoparticles drug delivery system

Nanocarriers are gaining wide usage as drug delivery systems because of their increased stability in storage, improved targeting ability on disease cells, sustained drug release, and higher encapsulation ability [ 61 ]. Amongst the widely accepted nanoparticles being used today for drug delivery, liposomes and polymeric nanoparticles are the most widely accepted. Liposomes which is a lipid-based nanoparticles though showed excellent biocompatibility still suffered drug leakage and instability upon storage while the polymeric nanoparticle which is a polymer-based nanoparticle was able to curb this limitation by exhibiting high encapsulation/drug loading ability as well as stability. However, it had its own shortfalls in that it showed lower biocompatibility [ 62 , 63 ]. In order to overcome these shortcomings and obtain an effective nanomaterial, researchers sought and developed a hybrid system that will combine the unique properties of the two classes of nanoparticles, which is known as Polymer-lipid hybrid nanoparticles (plhnps). This hybrid system was able to satisfy the requirements of biocompatibility, high storage stability, sustained drug release, minimal drug leakage, small particle size, and high encapsulation [ 64 ]. As a result of its efficacy, this system is being used today for different therapeutic purposes as well as diagnostic applications. Plhnps is made up of three distinct components which include: A polymeric core that encapsulates both hydrophilic and hydrophobic drugs effectively. This is possible as a result of the hydrophilic and hydrophobic nature of the core and results in a high sustained release, a lipid shell that provides biocompatibility and high stability and a lipid-polyethylene glycol (PEG) that is found in the outer part and covered by a lipid layer to provide increased steric stability, prevent immune recognition and increase time for circulation ( Fig. 5 ). Plhnps has wide applications such as in the delivery of various chemotherapeutic agents, in gene transfer (sirna, DNA) and in photothermal, photodynamic therapy and ultrasound. Studies have shown that they can be used in the delivery of vaccines and immune activation as well as in imaging and alternative magnetic field (AMF). Hence its wide application in the fast-growing medical environment [ 65 ].

The formation of a Polymeric-Lipid Hybrid Nanoparticle. The Hybrid contains three distinct components which include: A polymeric core that encapsulates both hydrophilic and hydrophobic drugs effectively. A lipid shell that provides biocompatibility and high stability and a lipid-polyethylene glycol (PEG) that is found in the outer part and covered by a lipid layer to provide increased steric stability, prevent immune recognition and increase time for circulation.

3.5. Self-microemulsifying drug-delivery system

Recently, lipid-based drug preparations have received a lot of interest, with a specific focus on self-microemulsifying drug-delivery systems (SMEDDS) [ 66 ]. Inadequate bioavailability is one of the most difficult aspects of developing oral dosage forms of drugs [ 67 ]. Accordingly, minimal hydrophilicity is an essential factor for bioavailability in this context, because drugs cannot be absorbed via the gastrointestinal tract (GIT) except it exists in solution forms [ 68 ]. Aqueous solubility is a problem for many chemical compounds having notable and favorable pharmacological effects [ 69 ]. Furthermore, almost 30% of widely marketed medicinal entities and nearly 50% of innovative drug compounds accessible for product manufacture are hydrophobic in nature, meaning they have low water solubility [ 70 ]. The utilization of a lipid-based carrier system to boost the bioavailability of less water-soluble medications has grown in popularity in recent years [ 71 ]. The main rationale behind this formulation is to sustain the hydrophobic components in solution all through the digestive system [ 70 ].

Lipid-based carriers come in a variety of forms, including suspensions, dry emulsions, microemulsions, and self-emulsifying drug-delivery systems (SEDDS) [ 72 ]. SEDDS' ability to incorporate hydrophobic drugs was previously reported. SEDDS has also been revised as self-microemulsifying drug-delivery systems (SMEDDS) and self-nanoemulsifying drug-delivery systems (SNEDDS). Emulsions, on the other hand, are created by dispersing a liquid phase containing macroscopic particles in a different liquid phase composed of surfactant [ 73 ]. They are also thermodynamically unstable solution that is semitransparent (occasionally hazy) and has properties that resemble viscous liquids [ 74 ]. Emulsions are of three types, viz; water-in-oil, oil-in-water, and multiple emulsions. Additionally, conventional micro- or nanoemulsions differ from SMEDDS, in that following oral ingestion, they self-emulsify [ 75 ].

Surfactants (S) and Co-surfactants (CoSs) are the two types of emulsifying agents used in microemulsions ( Fig. 6 ). However, surfactant is predominantly soluble in water, while CoS is mostly soluble in the oil phase. CoSs are essential for lowering the tension that exists between the two liquid phases to the optimal level required for the formation of a microemulsion [ 76 , 77 ]. The production of nanoemulsions with droplet sizes smaller than 100 nm, on the other hand, requires either mechanical or chemical energy [ 78 ]. Although nanoemulsions are classified as kinetically stable due to their extremely low destabilization rate, their long-duration stability (in months) is notable [ 79 ]. Consequently, nanoemulsion globules are shown to be stable across a variety of situations, including varied dilutions and temperatures, while microemulsions are mostly impacted by factors including dilutions and temperature [ 74 , 80 ]. The key distinctions between SMEDDS, SNEDDS, and SEDDS are shown in Table 1 .

Mechanism of Self-emulsification in aqueous environment. SEDDS comprise a mixture drugs, surfactants, oil, stabilizers, and cosolvents. Like conventional emulsification, SEDDS (ionotropic mixture) form (o/w) nano or microemulsion within the gastrointestinal tract with very small energy input.

Four basic types of lipid-based drug-delivery systems (LBDDS) with their merits and demerits.

HLB=Hydrophilic/lipophilic balance; LFCS = lipid formulation classification system; SEDDS = self-emulsifying drug-delivery systems; SMEDDS = self-microemulsifying drug-delivery systems; SNEDDS = self-nanoemulsifying drug-delivery system [ 70 ] (Rajpoot, Kuldeep 2020).

3.6. In-situ gel drug delivery system

The foremost goal of any drug delivery system is to change the drug's pharmacokinetic characteristics and tissue distribution in a meaningful way [ 81 ]. Over the last 60 years, there has been a lot of focus on developing controlled and well reliable drug delivery systems [ 82 ]. In-situ gel medication administration has emerged as one of the most innovative drug delivery systems. By virtue of its unique property of transitioning from Sol to Gel, the in-situ gel drug delivery system aids in the prolonged and regulated release of medications, as well as increased patient compliance and comfort [ 83 ]. In most case, formulations in solution form become transformed into gel form under certain physiological conditions before it enters the body [ 84 ]. A variety of stimuli, such as pH change, temperature modulation, and solvent exchange, combine to transform a solution into a gel form [ 85 ]. Oral, nasal, injectable, vaginal, rectal ocular, intraperitoneal, and parenteral routes have all been used in various research. Many polymeric methods capable of delivering pharmaceuticals have been created [ 86 ]. When these polymers come into touch with physiological stimuli, they go through a sol-gel transition. In situ gel drug delivery systems are made from a variety of natural and synthetic polymers [ 83 ]. Four processes are known to produce the formation of in-situ gel biomaterials, viz; (1) temperature and pH variations, (2) variations of the physical properties of the biomaterials such as solvent exchange and swelling, (3) Biochemical modification such as enzymatic and chemical reactions, and (4) photo-polymerization [ 87 ].

3.6.1. Applications of in-situ gel delivery system

3.6.1.1. oral in-situ gel delivery systems.

The application of pH-sensitive hydrogels to particular parts of the gastrointestinal tract for site-specific medicine delivery is the focus of this approach. With different amounts of Polyacrylic acid (PAA) derivatives and cross-linked Polyethylene glycol (PEG), silicone microsphere hydrogels that liberate prednisolone into the stomach media or exhibit gastroprotective properties have been produced. A possible colon-specific medicine delivery approach for decreasing edema at high pH has been developed using dextran hydrogels cross-linked with polysaccharides such as guar gum, inulin, and amidade pectin. Researchers produced gellan gum and sodium alginate formulations that utilized calcium ions as complexing agents and were gelatinized by releasing these ions into the stomach's acidic medium. The oral in situ gel distribution procedures employ natural polymers such as xyloglucan, pectin, and gellan gum. The pectin-based formulation was created to ensure that Metformin loaded pectin (PCM) was distributed continuously. Because pectin is water soluble, an organic solvent is not required [ 88 , 89 ].

3.6.1.2. Opthalmic in-situ gelling systems

Gellanic gum, alginic acid, and xyloglucan are examples of natural polymers employed in the ocular delivery system. Several combinations of anti-inflammatory, antibacterial, and autonomic medicines are utilized to reduce intraocular glaucoma stress in the local ophthalmic administration approach. Tear fluid has been devised to overcome the ocular in-situ gel bioavailability problem due to its rapid turnover and dynamics. Traditional delivery methods also result in limited availability and therapeutic response, making it easier to remove the medication from the eye. In ocular preparations, viscosity enhancers such as Carboxymethyl Cellulose, Polyvinyl alcohol, Carbomers, and Hydroxypropylmethyl cellulose are deployed to improve the viscosity of the drug formulations, resulting in enhanced bioavailability and precorneal residence duration. Chelating agents' penetration enhancers are utilized to promote the infiltration of corneal substances such as surfactants and preservers [ 90 , 91 ].

3.6.1.3. Nasal in-situ gelling systems

In nasal in-situ production, polymers such as gum gelan and gum xanthan are employed. The efficacy of Momethasone furoate as an in situ gel in the management of allergic rhinitis has been studied. The impact of in-situ gel on antigen-induced nasal symptoms was shown in vivo using sensitized rats as a model of allergic rhinitis. In situ gel was proven to resist an increase in nasal complications when compared to the commercialized nosonexex preparation [ 92 , 93 ].

3.6.1.4. Rectal in-situ gelling systems

This method can be used to prescribe a variety of pharmaceuticals that are packed as liquid, semi-solid (liniments, emulsions, and froths), or suppositories in solid administration formulations. During penetration, traditional suppositories can cause discomfort. Furthermore, because suppositories cannot be effectively kept at a single rectum site, they can migrate upward into the gut, allowing the drug's first-pass impact to be experienced. Indomethacin-loaded xyloglucan-based device administration to rabbits revealed a substantial medication immersion and a prolonged drug residence period when compared to commercial suppository administration [ 94 ].

3.6.1.5. Vaginal in-situ gelling systems

The vaginal canal is a possible route for medication delivery. A delivery system based on a thermoplastic graft copolymer undergoing in situ gelation has been formulated for the continuous liberation of active substances such as estrogens, peptides, progestins, and proteins. In a recent study, clotrimazole medication antifungal efficiency was boosted and sustained using a mix of poloxamers and polycarbophils mucoadhesive thermosensitive gels in contrast to conformist polyethylene glycol-based formulations [ 95 ].

3.6.1.6. Injectable in-situ gelling system

Formulating dose forms like injectable or implanted delivery systems is one of the only obvious ways to give drugs for prolonged release. Thermo-reversible gels, typically made of poloxamers, are the most commonly utilized. These dosage forms might be useful in the manufacture of controlled drug delivery for systemic absorption. Pluronic F127 gels with insulin or insulin-PLGA nanoparticles have been tested. Poloxamer gels have also been used to subcutaneously and intramuscularly deliver human growth hormone as well as to generate a single-dose lidocaine injectable with a longer duration of action. Pluronics recently developed a new type of depot protein injectables controlled release formulation made up of blends of poly (D, l -lactide)/1-methyl-2-pyrrolidone solutions. At a temperature of roughly 45 °C, the hydrogel develops into a gel after being injected subcutaneously, and fast body temperature cooling was achieved by employing biodegradable polymers of poly (ethylene oxide) and poly (propylene oxide) ( l -lactic acid). The injectable drug delivery method is utilized to cross-link pluronic acid-modified hydrazide with aldehyde-modified cellulose derivatives such as hydroxypropylmethyl cellulose, carboxymethyl cellulose, and methylcellulose, among others. In order to avoid postoperative peritoneal adhesion, this in-situ forming gel has been used to reduce pelvic discomfort, bowel blockage, and infertility [ 96 , 97 ].

3.7. Micro electro mechanical systems (MEMS) for drug delivery

MEMS technology has vast applications in fields such as actuators, drug delivery, motion sensing, accelerometers, and inkjet printing [ 98 ]. The devices produced through this technology incorporate microfabrication techniques to produce micro/nano-sized electromechanical and mechanical devices or implants [ 99 ]. Interestingly, the use of these techniques enhances the efficacy of these devices by allowing considerable control over their topography, microarchitecture, and size of the resulting devices [ 100 ].

Among the numerous materials and processes available for designing these MEMS-based devices, the most commonly used, incorporate creative blends of varying micromachining techniques such as; deposition (an addictive process), etching (a subtractive process), lithography (a patterning process), ink jetting, ion implantation, oxidation, and micromolding [ [101] , [102] , [103] , [104] ]. As drug delivery systems, MEMS technology fabricates miniaturized systems comprising of various materials such as silicon, glass, metals, and nitrides, as well as polymers, micropumps, sensors, microvalves, reservoirs, actuators, and high-performance processors [ 100 , 103 , 105 , 106 ]. These distinct components function synergistically, to provide the broadly reported multi-functionality and precision of MEMS devices, relative to other conventional drug delivery systems. Each of these features works strategically, for instance; actuators mostly play a vital role in the drug release process, as it pressurizes the drug reservoir to facilitate drug release [ 103 ]. Reservoirs provide port(s) to house the drug(s) and can be singular or multiple [ 107 ]. Single reservoir architecture comprises a relatively large port that can contain a single drug. It is thus able to contain a relatively larger amount of drug and is also suited for long-term usage as it can be refilled. On the other hand, multi-reservoirs comprise different ports (within the same substrate) that separately store drugs, and as such different drugs can be incorporated into these. However, they are less suitable for long-term usage as they would require repetitive replacement surgeries, due to a lack of refilling methods. Moreover, microvalves are employed to control fluid flow rate, sealing, as well as switching on/off of the delivery device [ 108 ]. Silicon is popularly used as a substrate or structural material during fabrication, due to its favorable mechanical and electrical properties [ 109 ]. Sensors, on their part, utilize electrical radiation, mechanical, thermal, magnetic, or biochemical mechanisms to monitor the flow measurements of fluid or gas being delivered [ [110] , [111] , [112] ]. As such, the choice of each feature during the design process is crucial to the overall functionality of the MEMS-based delivery device.

MEMS-based devices play essential roles in targeted and precise drug delivery by facilitating controlled and pulsatile release of enclosed pharmaceuticals [ 113 ]. For this purpose, these devices could be designed either as electric-powered or non-electric powered. The electric-powered devices allow selective drug release from the reservoir(s) by electric potential, while the non-powered devices utilize diffusion, and osmotic environmental stimulus mechanisms to enable drug release [ 105 ]. The most popular type of MEMS technology applicable in drug delivery is microchips, followed by microfluidic devices, majorly micropumps. These microchips are reservoir-based implantable devices, which are capable of delivering pharmaceutics in solid, gel, or liquid forms through trans/intradermal delivery [ 109 ]. On the other hand, micropumps which are either classified as mechanical or non-mechanical, depending on the presence of moving parts, are distinctly limited to the delivery of drug suspensions or solutions [ 106 ]. The major components of MEMs are shown in Fig. 7 .

Schematic representation of MEMS components. Generally, MEMS is made up of mechanical microstructures, microactuators, microsensors, and microelectronics all integrated onto a single silicon chip. (For interpretation of the references to colour in this figure legend, the reader is referred to the Web version of this article.)

Drug delivery devices fabricated by MEMS technologies offer several advantages over conventional delivery methods, such as; enhanced performance, automation, precision, and efficacy due to the integration of their miniaturized sizes with multi-functional components. It also contributes to less painful and invasive attributes of the devices [ 105 ]. In addition, MEMS-based devices have the ability to maintain drug stability during encapsulation, adjustable and continuous delivery, and also facilitate the automated release of multiple drugs from reservoirs [ 114 ]. Furthermore, it enhances bioavailability and localized release of medication [ 115 ], sustainability over a long period of time for medications requiring complex dosing, as well as personalized dosing profile [ 109 ], and has the ability to function following sustained zero-order kinetics. However, MEMS-drug delivery devices may pose technical challenges; because the incorporation of wireless electronics to remotely control and track the device's operation, as well as patient's response, has also been projected to be capable of increasing device security risks, medical packaging, and regulatory complexities [ 105 ]. Moreover, surgeries would be required for the implantation and removal of these devices; highly stable products would be needed for long-term uses; and the technologies required for their fabrication are relatively expensive [ 114 ].

3.8. Combined drug delivery approach

Resistance to drugs has been a recurrent issue in medical therapy. Today, combination therapy proves more viable as a result of wider target specificity and complementarity in enhancing treatment efficiency and increasing clinical results. Combined drug delivery approach has been widely adopted in cancer research and therapy as a means to overcome multidrug resistance. It has been reported that combination drug delivery approach reduces therapeutic dosage as well as adverse reactions while efficiency and decrease in drug resistance are maintained [ 116 ].

A study conducted by Zamora-Mera et al. reported positive results in the use of combination therapy for magnetic hyperthermia therapy. They crosslinked chitosan nanoparticles (CSNPs) ionically with tripolyphosphate salts (TPP). The magnetic CSNPs were obtained by encapsulation with three different ferrofluid concentrations and a constant 5-Fluorouracil (5-FU) concentration. They used normal cells, fibroblasts (FHB) and cancer cells, human glioblastoma A-172 cells [ 117 ]. The CSNPs showed dose-dependent cytotoxicity and were successfully up-taken in both cell lines. The study reported that the MH-treatment in the A-172 cells produced a 67–75% cell viability whereas no cell viability was noticed in FHB. The study equally reported a 4-h regrowth of the population upon MH treatment with CSNPs loaded only with ferrofluid but a decreased amount of released 5-FU upon combination with the MH treatment and 5-FU demonstrating a positive result using a combination approach [ 117 ].

In a study conducted by Silva et al., they reported an increase in the application of combination approach in drug research and in therapeutic studies. They investigated the use of a combined method to remove endotoxins from protein nanocages for drug delivery approaches. They combined an affinity purification with Endotrap-HD resin and treatment with Triton X-144. The study yielded good results that showed combination treatment as a good potential in chemotherapy [ 118 ].

A review article conducted by Pang et al. on a particular combination drug delivery approach that focused on exploiting cells in combination with nanoparticles reported that nanoparticles loaded in cells were more effective than the nanoparticle drug delivery system. The cell-based therapy showed improved drug efficacy, extended half-lives, sustained drug release, and limited immunogenicity and cytotoxicity [ 119 ]. Combining nanoparticles with exploit cells did not affect its migration or chemotaxic ability. As a result, combination drug delivery approach is viewed as a promising approach in drug research and medical therapy [ 119 ].

3.9. Targeted drug delivery system

This approach is an advanced technique employed recently due to its efficiency and reduced side effects. It is a system that delivers drugs in a targeted sequence which in turn leads to an increase in the drug concentration as it is being delivered to its target site [ 120 ]. The dosage of the drug is reduced to minimize side effects but its efficacy and strength remain untouched. This approach employs other drug carriers such as soluble polymers, biodegradable microsphere polymers, neutrophils, liposomes, micelles, and artificial cells amongst others [ 120 ]. This technique is gaining wide acceptance as it proves useful, especially in the fight against cancer.

A study conducted by Murugun showed effectiveness in the use of this drug delivery system. Topotecan (TPT) and quercetin (QT) were delivered using polyacrylic acid chitosan surface-modified mesoporous silica nanoparticle (MSN) to target negative breast cancer cells (TNBC) (MDA-MB-231) and multidrug-resistant breast cancer cells (MCF-7) [ 121 ]. The surface of the nanoparticles was grafted with RGD-peptides which is an amino acid made up of Arg-Gly-Asp sequences. This was done to effectively target αvβ3 integrin. The RGD peptide led to an effective release of encapsulated drugs as well as cellular uptakes by the cancer cells. Both cell lines showed cell death, molecular and structural changes of cellular nucleus, endoplasmic reticulum, and mitochondria. A synergistic antiproliferative effect was also observed [ 121 ].

Another study conducted by Wu et al. showed an enhanced release of methrotrexate (MTX) from Fe 3 O 4 MgAl-LDH (layered double hydroxide) nanoparticles of ∼230 nm [ 122 ]. They reported 84.94% release in the tumor with a pH of 3.5 within 48 h. Their study showed higher antitumor activities across the cell lines that were investigated. Lin et al. used this approach to target HeLa cells. Mitomycin C (MMC) and 10-hydroxycamptothecin (HCPT) were co-delivered using a folate-functionalized soybean phosphatidylcholine micellar Nano formulation to determine their therapeutic effect on the HeLa Cells [ 123 ]. The study reported enhanced cellular uptake both in vitro and in vivo and an enhanced decrease in tumor burden compared to free drugs. These findings and much more suggest that a targeted drug delivery system is an area researchers’ ought to also pay more attention to. A comprehensive summary of recent drug delivery systems including their therapeutic uses, advantages, and disadvantages were presented in Table 2 .

The summary of recent drug delivery systems, their uses and advantages, and disadvantages.

4. Challenges associated with current drug delivery systems

Many delivery systems have been used successfully recently, there has been a great development in the quest to deliver drugs from various plant sources to their target sites for treatment in the body, however there are numerous limitations and challenges to what these systems can achieve in treatment, some of which are discussed below.

A major challenge facing the advancement of drug delivery systems is the limited amount of literature and variation in the available literature. Literature provides important information for the advancement of any research and in this case, nanomedicine approaches to treatment. The variation in published studies in relation to recorded characterization of reported experimental details is seen to also be a major difficulty in the progression of nanotechnology application in medicine [ 134 ]. The limited and varied information that should be a guide for industries could impede future breakthroughs of nanomedicines and delay the transformation from research and experimentation to clinical application [ 135 ]. Many researchers agree that nanoparticles can be either good or bad, the benefits of nanoparticles are more widely known and recognized but there is a scarcity of information when it comes to how safe these particles are, their level of interactions with proteins that are not specific to them and their movement and interaction with other organs that are not their target organs [ 136 ].

Some of these delivery systems make use of large particles as carriers which are not particularly favorable for treatment because they can constitute challenges such as poor absorption and solubility, in vivo instability , poor bioavailability, target-specific delivery complications, and several adverse side effects upon administration [ 23 ], the use of much smaller particles for delivery to the human biological system is a way out that solves the issues that come with using much larger particles.

Target-specific delivery complication is a challenge that faces all delivery systems. Even though target-specific delivery has been found to reduce toxicity and shows more effective treatment, its efficacy cannot be assured until it is able to reach the targeted site in sufficient amount, this is seen when siRNA is given systemically, rarely do they get to their target cell/organ because they are easily degraded by body enzymes, and when administered in large quantity their negative charge becomes an obstacle to absorption by the cells [ 137 ] resulting in little or no absorption by the body. Micelles and liposomes which are lipid nanoparticles are being studied for target drug delivery, but the downside to this is the lowering of their efficiency by their reaction with the body, these reactions include phagocytic absorption and hepatic filtration, which could lead to failure in target delivery, also the nanoparticles could show signs of toxicity [ 138 ]. The challenge to targeted delivery is that a patient that is unconscious cannot take a dose, there is low solubility and permeability at the target site, they could interact with food and may be degraded by gastro intestinal flora [ 139 ].

The toxicity of particles used in delivery is another great challenge that faces drug delivery systems in general, some of the nanomaterials used can be harmful to human health and also the environment [ 140 ]. In vivo and in vitro experimentation has shown the harmful effect of silver, gold, silica, and titanium used as nanoparticles for coupling and delivering drugs [ 136 ]. Carbon nanotubes (CNTs) have become widely used in gene therapy, bio-imaging, and drug delivery [ 141 ]. because they have been found to have the ability to cross cell membranes even when they are used as carriers for biomolecules [ 142 ] but the properties of carbon nanotubes have raised concerns among researchers especially in its use in drug delivery because experiments show that they can cause harm to embryos, genes, liver, heart, neurons and immune system [ 136 ]. Although carbon nanotubes have shown favorable results in their use, it is important to carry out crucial toxicity tests to ensure their safety before a widespread application in treatment [ 141 ]. In application, their effects have become an obstacle in their use for cancer treatment [ 143 ].

Scientists have successfully produced drugs that can serve as carriers and drugs as well, one of the major challenges facing drug delivery systems is biocompatibility (their ability to function with the body in specific situations) and acceptability (being received by the body without triggering the immune system), this is a problem because the way the body reacts to biological materials are very different from the way they react to synthetic materials [ 140 ]. In addition, because of the complex structure of the human system, there could be natural barriers to the functions of these delivery systems for example the blood brain-barrier (BBB) has a selectively permeable feature that makes it difficult in achieving therapeutic drug concentration in the brain tissues, the BBB prevents the entry of carrier particles into the brain and entire central nervous system causing ineffectiveness of therapeutic agents in the treatment of cerebral diseases due to inability to deliver and sustain intended drugs within the brain efficiently [ 144 ]. Also, one of the most abundant carriers in the body are the monoclonal antibodies (mAb), because they form immunoliposomes by binding to liposome surfaces however the functions of these immunoliposomes are limited because they can trigger an immune response and low levels of absorption, distribution, metabolism, and elimination by the body, this presents a challenge in the use of liposomes to achieve efficacy as a site-specific drug carrier [ 145 ].

The kidney and liver also have a natural ability to detox the body, these organs can treat the nanoparticles as potential waste products. Their function can constitute obstruction in drug delivery and lead to the accumulation of nanoparticles in these organs. In the liver, nanomaterials accumulate primarily in the Kupffer cells, macrophages in the liver, sinusoidal endothelial cells, hepatic stellate cells, and a little in the hepatocytes. Meanwhile the size, charge and shape of the kidney decide the fate of the nanomaterials once they get into the renal system [ 146 ].

5. Future directions and conclusion

Drug delivery and nanomedicine have become a very fascinating area of research in modern science, in the past years, it has gotten a lot of attention in both research, experimentation and in a number of clinical trials [ 147 ]. Despite the setback that have limited the clinical application of these delivery systems, the recent drug delivery system holds great potential, it would require a collaboration across academic theory, laboratory experimentation, the knowledge of medicine, pharmaceuticals, and great research to help achieve the efficiency we need to take, findings from bench to bedside [ 148 ]. Vargason et al. [ 2 ] believe that the use of cell therapies can go a long way to solve the bio-acceptability issues that drug delivery systems face, they also think it will create a single dose that is effective that avoids high accumulation of drugs in the system. As a matter of fact, cell therapies promise a seeming sustained source of complex biologics, break down innate biological barriers and create responses that appear natural within the system. Adepu [ 149 ] have suggested the use of inorganic mesoporous nanoparticles, micro fluids, and molecular imprinting polymers as some of the ways to combat some of the challenges that face drug delivery. According to Khalid et al. [ 137 ] a way to improve drug delivery efficacy is by the use of priming agents that can influence the biological environment where they are administered, especially those that can change the form and function of tissues in a way that makes the administered drug favorable without harming the patient. Also, cell-based drug systems should be considered in the field of biomaterials, this means the use of cells coupled with nano biomaterials since cells are indigenous to the human system, this is a novel method and still theoretical but appears to be most creative, encourages drug delivery method with hopes to achieve maximum drug delivery pattern. A lot of research and clinical trials are still needed to foster the efficiency of these modern drug delivery systems and the challenges that face their usage.

Author contribution statement

All authors listed have significantly contributed to the development and the writing of this article.

Data availability statement

Declaration of competing interest.

The authors declare that they have no known competing financial interests or personal relationships that could have appeared to influence the work reported in this paper.

Drug Delivery and Translational Research

An Official Journal of the Controlled Release Society

• Explores the effect of rationally developed, effective delivery systems in different disease conditions.
• Considers preclinical and clinical data related to drug delivery systems.
• Devices for drug delivery and drug/device combination products.
• Provides a unique blend of original full-length papers, communications, reviews, and more.
• Includes reports of future meetings, research highlights, and announcements pertaining to activities of the Controlled Release Society (CRS).

This is a transformative journal , you may have access to funding.

• Maria Jose Alonso

Societies and partnerships

Latest issue

Volume 14, Issue 5

Latest articles

Enhancing the hypotensive effect of latanoprost by combining synthetic phosphatidylcholine liposomes with hyaluronic acid and osmoprotective agents.

• Marco Brugnera
• Marta Vicario-de-la-Torre
• Rocío Herrero-Vanrell

Anti-lung cancer synergy of low-dose doxorubicin and PD-L1 blocker co-delivered via mild photothermia-responsive black phosphorus

• Hua-Zhen Xu
• Fei-Xiang Chen

Progress of antibody–drug conjugates (ADCs) targeting c-Met in cancer therapy; insights from clinical and preclinical studies

• Ali Hussein Mer
• Yousef Mirzaei
• Meghdad Abdollahpour‐Alitappeh

Targeting strategies with lipid vectors for nucleic acid supplementation therapy in Fabry disease: a systematic review

• Julen Rodríguez-Castejón
• Marina Beraza-Millor
• Ana del Pozo-Rodríguez

Nanomedicines via the pulmonary route: a promising strategy to reach the target?

• Mélina Guérin
• Elise Lepeltier

Journal updates

Featured article: april 2024.

Read the featured article from the April 2024 issue!  

Featured Article: March 2024

Read the featured article from the March 2024 issue!

Featured Article: February 2024

Read the featured article from the February 2024 issue!

Featured Article: January 2024

Read the featured article from the January 2024 issue!  

Journal information

• Chemical Abstracts Service (CAS)
• Google Scholar
• Japanese Science and Technology Agency (JST)
• Norwegian Register for Scientific Journals and Series
• OCLC WorldCat Discovery Service
• Science Citation Index Expanded (SCIE)
• TD Net Discovery Service
• UGC-CARE List (India)

Rights and permissions

Springer policies

© Controlled Release Society

• Find a journal
• Publish with us
• Track your research

• Open access
• Published: 19 September 2018

Nano based drug delivery systems: recent developments and future prospects

• Jayanta Kumar Patra   ORCID: orcid.org/0000-0003-4118-4355 1 ,
• Gitishree Das 1 ,
• Leonardo Fernandes Fraceto 2 , 3 ,
• Estefania Vangelie Ramos Campos 2 , 3 ,
• Maria del Pilar Rodriguez-Torres   ORCID: orcid.org/0000-0001-9107-247X 4 ,
• Laura Susana Acosta-Torres   ORCID: orcid.org/0000-0002-5959-9113 4 ,
• Luis Armando Diaz-Torres   ORCID: orcid.org/0000-0002-1281-9916 5 ,
• Renato Grillo 6 ,
• Mallappa Kumara Swamy 7 ,
• Shivesh Sharma 8 ,
• Solomon Habtemariam 9 &
• Han-Seung Shin 10  

Journal of Nanobiotechnology volume  16 , Article number:  71 ( 2018 ) Cite this article

448k Accesses

3607 Citations

62 Altmetric

Metrics details

Nanomedicine and nano delivery systems are a relatively new but rapidly developing science where materials in the nanoscale range are employed to serve as means of diagnostic tools or to deliver therapeutic agents to specific targeted sites in a controlled manner. Nanotechnology offers multiple benefits in treating chronic human diseases by site-specific, and target-oriented delivery of precise medicines. Recently, there are a number of outstanding applications of the nanomedicine (chemotherapeutic agents, biological agents, immunotherapeutic agents etc.) in the treatment of various diseases. The current review, presents an updated summary of recent advances in the field of nanomedicines and nano based drug delivery systems through comprehensive scrutiny of the discovery and application of nanomaterials in improving both the efficacy of novel and old drugs (e.g., natural products) and selective diagnosis through disease marker molecules. The opportunities and challenges of nanomedicines in drug delivery from synthetic/natural sources to their clinical applications are also discussed. In addition, we have included information regarding the trends and perspectives in nanomedicine area.

Since ancient times, humans have widely used plant-based natural products as medicines against various diseases. Modern medicines are mainly derived from herbs on the basis of traditional knowledge and practices. Nearly, 25% of the major pharmaceutical compounds and their derivatives available today are obtained from natural resources [ 1 , 2 ]. Natural compounds with different molecular backgrounds present a basis for the discovery of novel drugs. A recent trend in the natural product-based drug discovery has been the interest in designing synthetically amenable lead molecules, which mimic their counterpart’s chemistry [ 3 ]. Natural products exhibit remarkable characteristics such as extraordinary chemical diversity, chemical and biological properties with macromolecular specificity and less toxicity. These make them favorable leads in the discovery of novel drugs [ 4 ]. Further, computational studies have helped envisage molecular interactions of drugs and develop next-generation drug inventions such as target-based drug discovery and drug delivery.

Despite several advantages, pharmaceutical companies are hesitant to invest more in natural product-based drug discovery and drug delivery systems [ 5 ] and instead explore the available chemical compounds libraries to discover novel drugs. However, natural compounds are now being screened for treating several major diseases, including cancer, diabetes, cardiovascular, inflammatory, and microbial diseases. This is mainly because natural drugs possess unique advantages, such as lower toxicity and side effects, low-price, and good therapeutic potential. However, concerns associated with the biocompatibility, and toxicity of natural compounds presents a greater challenge of using them as medicine. Consequently, many natural compounds are not clearing the clinical trial phases because of these problems [ 6 , 7 , 8 ]. The use of large sized materials in drug delivery poses major challenges, including in vivo instability, poor bioavailability, and poor solubility, poor absorption in the body, issues with target-specific delivery, and tonic effectiveness, and probable adverse effects of drugs. Therefore, using new drug delivery systems for targeting drugs to specific body parts could be an option that might solve these critical issues [ 9 , 10 ]. Hence, nanotechnology plays a significant role in advanced medicine/drug formulations, targeting arena and their controlled drug release and delivery with immense success.

Nanotechnology is shown to bridge the barrier of biological and physical sciences by applying nanostructures and nanophases at various fields of science [ 11 ]; specially in nanomedicine and nano based drug delivery systems, where such particles are of major interest [ 12 , 13 ]. Nanomaterials can be well-defined as a material with sizes ranged between 1 and 100 nm, which influences the frontiers of nanomedicine starting from biosensors, microfluidics, drug delivery, and microarray tests to tissue engineering [ 14 , 15 , 16 ]. Nanotechnology employs curative agents at the nanoscale level to develop nanomedicines. The field of biomedicine comprising nanobiotechnology, drug delivery, biosensors, and tissue engineering has been powered by nanoparticles [ 17 ]. As nanoparticles comprise materials designed at the atomic or molecular level, they are usually small sized nanospheres [ 18 ]. Hence, they can move more freely in the human body as compared to bigger materials. Nanoscale sized particles exhibit unique structural, chemical, mechanical, magnetic, electrical, and biological properties. Nanomedicines have become well appreciated in recent times due to the fact that nanostructures could be utilized as delivery agents by encapsulating drugs or attaching therapeutic drugs and deliver them to target tissues more precisely with a controlled release [ 10 , 19 ]. Nanomedicine, is an emerging field implementing the use of knowledge and techniques of nanoscience in medical biology and disease prevention and remediation. It implicates the utilization of nanodimensional materials including nanorobots, nanosensors for diagnosis, delivery, and sensory purposes, and actuate materials in live cells (Fig.  1 ). For example, a nanoparticle-based method has been developed which combined both the treatment and imaging modalities of cancer diagnosis [ 20 ]. The very first generation of nanoparticle-based therapy included lipid systems like liposomes and micelles, which are now FDA-approved [ 21 ]. These liposomes and micelles can contain inorganic nanoparticles like gold or magnetic nanoparticles [ 22 ]. These properties let to an increase in the use of inorganic nanoparticles with an emphasis on drug delivery, imaging and therapeutics functions. In addition, nanostructures reportedly aid in preventing drugs from being tarnished in the gastrointestinal region and help the delivery of sparingly water-soluble drugs to their target location. Nanodrugs show higher oral bioavailability because they exhibit typical uptake mechanisms of absorptive endocytosis.

Application and goals of nanomedicine in different sphere of biomedical research

Nanostructures stay in the blood circulatory system for a prolonged period and enable the release of amalgamated drugs as per the specified dose. Thus, they cause fewer plasma fluctuations with reduced adverse effects [ 23 ]. Being nanosized, these structures penetrate in the tissue system, facilitate easy uptake of the drug by cells, permit an efficient drug delivery, and ensure action at the targeted location. The uptake of nanostructures by cells is much higher than that of large particles with size ranging between 1 and 10 µm [ 17 , 24 ]. Hence, they directly interact to treat the diseased cells with improved efficiency and reduced or negligible side effects.

At all stages of clinical practices, nanoparticles have been found to be useful in acquiring information owing to their use in numerous novel assays to treat and diagnose diseases. The main benefits of these nanoparticles are associated with their surface properties; as various proteins can be affixed to the surface. For instance, gold nanoparticles are used as biomarkers and tumor labels for various biomolecule detection procedural assays.

Regarding the use of nanomaterials in drug delivery, the selection of the nanoparticle is based on the physicochemical features of drugs. The combined use of nanoscience along with bioactive natural compounds is very attractive, and growing very rapidly in recent times. It presents several advantages when it comes to the delivery of natural products for treating cancer and many other diseases. Natural compounds have been comprehensively studied in curing diseases owing to their various characteristic activities, such as inducing tumor-suppressing autophagy and acting as antimicrobial agents. Autophagy has been observed in curcumin and caffeine [ 25 ], whereas antimicrobial effects have been shown by cinnamaldehyde, carvacrol, curcumin and eugenol [ 26 , 27 ]. The enrichment of their properties, such as bioavailability, targeting and controlled release were made by incorporating nanoparticles. For instance, thymoquinone, a bioactive compound in Nigella sativa , is studied after its encapsulation in lipid nanocarrier. After encapsulation, it showed sixfold increase in bioavailability in comparison to free thymoquinone and thus protects the gastrointestinal stuffs [ 28 ]. It also increased the pharmacokinetic characteristics of the natural product resulting in better therapeutic effects.

Metallic, organic, inorganic and polymeric nanostructures, including dendrimers, micelles, and liposomes are frequently considered in designing the target-specific drug delivery systems. In particular, those drugs having poor solubility with less absorption ability are tagged with these nanoparticles [ 17 , 29 ]. However, the efficacy of these nanostructures as drug delivery vehicles varies depending on the size, shape, and other inherent biophysical/chemical characteristics. For instance, polymeric nanomaterials with diameters ranging from 10 to 1000 nm, exhibit characteristics ideal for an efficient delivery vehicle [ 7 ]. Because of their high biocompatibility and biodegradability properties, various synthetic polymers such as polyvinyl alcohol, poly- l -lactic acid, polyethylene glycol, and poly(lactic- co -glycolic acid), and natural polymers, such as alginate and chitosan, are extensively used in the nanofabrication of nanoparticles [ 8 , 30 , 31 , 32 ]. Polymeric nanoparticles can be categorized into nanospheres and nanocapsules both of which are excellent drug delivery systems. Likewise, compact lipid nanostructures and phospholipids including liposomes and micelles are very useful in targeted drug delivery.

The use of ideal nano-drug delivery system is decided primarily based on the biophysical and biochemical properties of the targeted drugs being selected for the treatment [ 8 ]. However, problems such as toxicity exhibited by nanoparticles cannot be ignored when considering the use of nanomedicine. More recently, nanoparticles have mostly been used in combination with natural products to lower the toxicity issues. The green chemistry route of designing nanoparticles loaded with drugs is widely encouraged as it minimises the hazardous constituents in the biosynthetic process. Thus, using green nanoparticles for drug delivery can lessen the side-effects of the medications [ 19 ]. Moreover, adjustments in nanostructures size, shape, hydrophobicity, and surface changes can further enhance the bioactivity of these nanomaterials.

Thus, nanotechnology offers multiple benefits in treating chronic human diseases by site-specific, and target-oriented delivery of medicines. However, inadequate knowledge about nanostructures toxicity is a major worry and undoubtedly warrants further research to improve the efficacy with higher safety to enable safer practical implementation of these medicines. Therefore, cautiously designing these nanoparticles could be helpful in tackling the problems associated with their use. Considering the above facts, this review aims to report different nano based drug delivery systems, significant applications of natural compound-based nanomedicines, and bioavailability, targeting sites, and controlled release of nano-drugs, as well as other challenges associated with nanomaterials in medicines.

Nano based drug delivery systems

Recently, there has been enormous developments in the field of delivery systems to provide therapeutic agents or natural based active compounds to its target location for treatment of various aliments [ 33 , 34 ]. There are a number of drug delivery systems successfully employed in the recent times, however there are still certain challenges that need to be addresses and an advanced technology need to be developed for successful delivery of drugs to its target sites. Hence the nano based drug delivery systems are currently been studied that will facilitate the advanced system of drug delivery.

Fundamentals of nanotechnology based techniques in designing of drug

Nanomedicine is the branch of medicine that utilizes the science of nanotechnology in the preclusion and cure of various diseases using the nanoscale materials, such as biocompatible nanoparticles [ 35 ] and nanorobots [ 36 ], for various applications including, diagnosis [ 37 ], delivery [ 38 ], sensory [ 39 ], or actuation purposes in a living organism [ 40 ]. Drugs with very low solubility possess various biopharmaceutical delivery issues including limited bio accessibility after intake through mouth, less diffusion capacity into the outer membrane, require more quantity for intravenous intake and unwanted after-effects preceding traditional formulated vaccination process. However all these limitations could be overcome by the application of nanotechnology approaches in the drug delivery mechanism.

Drug designing at the nanoscale has been studied extensively and is by far, the most advanced technology in the area of nanoparticle applications because of its potential advantages such as the possibility to modify properties like solubility, drug release profiles, diffusivity, bioavailability and immunogenicity. This, can consequently lead to the improvement and development of convenient administration routes, lower toxicity, fewer side effects, improved biodistribution and extended drug life cycle [ 17 ]. The engineered drug delivery systems are either targeted to a particular location or are intended for the controlled release of therapeutic agents at a particular site. Their formation involves self-assembly where in well-defined structures or patterns spontaneously are formed from building blocks [ 41 ]. Additionally they need to overcome barriers like opsonization/sequestration by the mononuclear phagocyte system [ 42 ].

There are two ways through which nanostructures deliver drugs: passive and self-delivery. In the former, drugs are incorporated in the inner cavity of the structure mainly via the hydrophobic effect. When the nanostructure materials are targeted to a particular sites, the intended amount of the drug is released because of the low content of the drugs which is encapsulated in a hydrophobic environment [ 41 ]. Conversely, in the latter, the drugs intended for release are directly conjugated to the carrier nanostructure material for easy delivery. In this approach, the timing of release is crucial as the drug will not reach the target site and it dissociates from the carrier very quickly, and conversely, its bioactivity and efficacy will be decreased if it is released from its nanocarrier system at the right time [ 41 ]. Targeting of drugs is another significant aspect that uses nanomaterials or nanoformulations as the drug delivery systems and, is classified into active and passive. In active targeting, moieties, such as antibodies and peptides are coupled with drug delivery system to anchor them to the receptor structures expressed at the target site. In passive targeting, the prepared drug carrier complex circulates through the bloodstream and is driven to the target site by affinity or binding influenced by properties like pH, temperature, molecular site and shape. The main targets in the body are the receptors on cell membranes, lipid components of the cell membrane and antigens or proteins on the cell surfaces [ 43 ]. Currently, most nanotechnology-mediated drug delivery system are targeted towards the cancer disease and its cure.

Biopolymeric nanoparticles in diagnosis, detection and imaging

The integration of therapy and diagnosis is defined as theranostic and is being extensively utilized for cancer treatment [ 44 , 45 ]. Theranostic nanoparticles can help diagnose the disease, report the location, identify the stage of the disease, and provide information about the treatment response. In addition, such nanoparticles can carry a therapeutic agent for the tumor, which can provide the necessary concentrations of the therapeutic agent via molecular and/or external stimuli [ 44 , 45 ]. Chitosan is a biopolymer which possesses distinctive properties with biocompatibility and presence of functional groups [ 45 , 46 , 47 ]. It is used in the encapsulation or coating of various types of nanoparticles, thus producing different particles with multiple functions for their potential uses in the detection and diagnosis of different types of diseases [ 45 , 47 ].

Lee et al. [ 48 ] encapsulated oleic acid-coated FeO nanoparticles in oleic acid-conjugated chitosan (oleyl-chitosan) to examine the accretion of these nanoparticles in tumor cells through the penetrability and holding (EPR) consequence under the in vivo state for analytical uses by the near-infrared and magnetic resonance imaging (MRI) mechanisms. By the in vivo evaluations, both techniques showed noticeable signal strength and improvement in the tumor tissues through a higher EPR consequence after the injection of cyanine-5-attached oleyl-chitosan nanoparticles intravenously (Cyanine 5).

Yang et al. [ 49 ] prepared highly effective nanoparticles for revealing colorectal cancer (CC) cells via a light-mediated mechanism; these cells are visible owing to the physical conjugation of alginate with folic acid-modified chitosan leading to the formation of nanoparticles with enhanced 5-aminolevulinic (5-ALA) release in the cell lysosome. The results displayed that the engineered nanoparticles were voluntarily endocytosed by the CC cells by the folate receptor-based endocytosis process. Subsequently, the charged 5-ALA was dispersed into the lysosome which was triggered by less desirability strength between the 5-ALA and chitosan through deprotonated alginate that gave rise to the gathering of protoporphyrin IX (PpIX) for photodynamic detection within the cells. As per this research, chitosan-based nanoparticles in combination with alginate and folic acid are tremendous vectors for the definite delivery of 5-ALA to the CC cells to enable endoscopic fluorescent detection. Cathepsin B (CB) is strongly associated with the metastatic process and is available in surplus in the pericellular areas where this process occurs; thus, CB is important for the detection of metastasis. Ryu et al. [ 50 ] designed a CB-sensitive nanoprobe (CB-CNP) comprising a self-satisfied CB-CNP with a fluorogenic peptide attached to the tumor-targeting glycol chitosan nanoparticles (CNPs) on its surface. The designed nanoprobe is a sphere with a diameter of 280 nm, with spherical structure and its fluorescence capacity was completely extinguished under the biological condition. The evaluation of the usability of CB-sensitive nanoprobe in three rat metastatic models demonstrated the potential of these nonoprobes in discriminating metastatic cells from healthy ones through non-invasive imaging. Hyaluronic acid (HA) is another biopolymeric material. This is a biocompatible, negatively charged glycosaminoglycan, and is one of the main constituents of the extracellular matrix [ 51 , 52 ]. HA can bind to the CD44 receptor, which is mostly over articulated in various cancerous cells, through the receptor-linker interaction. Thus, HA-modified nanoparticles are intriguing for their use in the detection and cure of cancer [ 53 , 54 , 55 ]. Wang et al. [ 56 ], coated the surface of iron oxide nanoparticles (IONP) with dopamine-modified HA. These nanoparticles have a hydrophilic exterior and a hydrophobic interior where the chemotherapeutic homocamptothecin is encapsulated [ 56 ]. The biopotential of this process was investigated in both laboratory and in the live cells. Increased uptake of nanoparticles by tumor cells was observed by MRI when an external magnetic field was employed [ 56 ]. After the intravenous administration of the nano-vehicle in 3 mg/kg (relative to the free drug) rats, a large tumor ablation was observed and after treatment, the tumors almost disappeared [ 56 ].

Choi et al. [ 53 ] also synthesized nanoparticles of hyaluronic acid with different diameters by changing the degree of hydrophobic replacement of HA. The nanoparticles were systemically administered in the mice with tumor, and then, its effect was studied. This same research group developed a versatile thermostatic system using poly (ethylene glycol) conjugated hyaluronic acid (P-HA-NPs) nanoparticles for the early detection of colon cancer and targeted therapy. To assess the effectiveness of the nanoparticles, they were first attached to the near-infrared fluorescent dye (Cy 5.5) by chemical conjugation, and then, the irinotecan anticancer drug (IRT) was encapsulated within these systems. The therapeutic potential of P-HA-NP was then investigated in different systems of the mice colon cancer. Through the intravenous injection of the fluorescent dye attached nanoparticles (Cy 5.5-P-HA-NPs), minute and initial-stage tumors as well as liver-embedded colon tumors were efficiently pictured using an NIRF imaging method. Due to their extraordinary capability to target tumors, drug-containing nanoparticles (IRT-P-HA-NP) showed markedly decreased tumor development with decreased systemic harmfulness. In addition, healing effects could be examined concurrently with Cy 5.5-P-HA-NPs [ 57 ].

Another option that can be used is alginate, which is a natural polymer derived from the brown seaweed and has been expansively scrutinized for its potential uses in the biomedical field because of its several favorable characteristics, such as low cost of manufacture, harmonious nature, less harmfulness, and easy gelling in response to the addition of divalent cations [ 58 , 59 ]. Baghbani et al. [ 60 ] prepared perfluorohexane (PFH) nanodroplets stabilized with alginate to drive doxorubicin and then evaluated their sensitivity to ultrasound and imaging as well as their therapeutic properties. Further found that the ultrasound-facilitated treatment with PFH nanodroplets loaded with doxorubicin exhibited promising positive responses in the breast cancer rat models. The efficacy was characterized by the deterioration of the tumor [ 60 ]. In another study, Podgorna et al. [ 61 ] prepared gadolinium (GdNG) containing nanogels for hydrophilic drug loading and to enable screening by MRI. The gadolinium alginate nanogels had an average diameter of 110 nm with stability duration of 60 days. Because of their paramagnetic behavior, the gadolinium mixtures are normally used as positive contrast agents (T1) in the MRI images. Gadolinium nanogels significantly reduce the relaxation time (T1) compared to controls. Therefore, alginate nanogels act as contrast-enhancing agents and can be assumed as an appropriate material for pharmacological application.

Also, the polymeric material dextran is a neutral polymer and is assumed as the first notable example of microbial exopolysaccharides used in medical applications. A remarkable advantage of using dextran is that it is well-tolerated, non-toxic, and biodegradable in humans, with no reactions in the body [ 62 ]. Photodynamic therapy is a site-specific cancer cure with less damage to non-cancerous cells. Ding et al. [ 63 ] prepared a nanoparticulate multifunctional composite system by encapsulating Fe 3 O 4 nanoparticles in dextran nanoparticles conjugated to redox-responsive chlorine 6 (C6) for near infrared (NIR) and magnetic resonance (MR) imaging. The nanoparticles exhibited an “off/on” behavior of the redox cellular response of the fluorescence signal, thus resulting in accurate imaging of the tumor. In addition, excellent in vitro and in vivo magnetic targeting ability was observed, contributing to the efficacy of enhanced photodynamic therapy. Hong et al. [ 64 ] prepared theranostic nanoparticles or glioma cells of C6 mice. These particles comprised of gadolinium oxide nanoparticles coated with folic acid-conjugated dextran (FA) or paclitaxel (PTX). The bioprotective effects of dextran coating and the chemotherapeutic effect of PTX on the C6 glioma cells were evaluated by the MTT assay. The synthesized nanoparticles have been shown to enter C6 tumor cells by receptor-mediated endocytosis and provide enhanced contrast (MR) concentration-dependent activity due to the paramagnetic property of the gadolinium nanoparticle. Multifunctional nanoparticles were more effective in reducing cell viability than uncoated gadolinium nanoparticles. Therefore, FA and PTX conjugated nanoparticles can be used as theranostic agents with paramagnetic and chemotherapeutic properties.

Drug designing and drug delivery process and mechanism

With the progression of nanomedicine and, due to the advancement of drug discovery/design and drug delivery systems, numerous therapeutic procedures have been proposed and traditional clinical diagnostic methods have been studied, to increase the drug specificity and diagnostic accuracy. For instance, new routes of drug administration are being explored, and there is focus on ensuring their targeted action in specific regions, thus reducing their toxicity and increasing their bioavailability in the organism [ 65 ].

In this context, drug designing has been a promising feature that characterizes the discovery of novel lead drugs based on the knowledge of a biological target. The advancements in computer sciences, and the progression of experimental procedures for the categorization and purification of proteins, peptides, and biological targets are essential for the growth and development of this sector [ 66 , 67 ]. In addition, several studies and reviews have been found in this area; they focus on the rational design of different molecules and show the importance of studying different mechanisms of drug release [ 68 ]. Moreover, natural products can provide feasible and interesting solutions to address the drug design challenges, and can serve as an inspiration for drug discovery with desired physicochemical properties [ 3 , 69 , 70 ].

Also, the drug delivery systems have been gaining importance in the last few years. Such systems can be easily developed and are capable of promoting the modified release of the active ingredients in the body. For example, Chen et al. [ 70 ] described an interesting review using nanocarriers for imaging and sensory applications and discussed the, therapy effect of these systems. In addition, Pelaz et al. [ 71 ] provided an up-to-date overview of several applications of nanocarriers to nanomedicine and discussed new opportunities and challenges for this sector.

Interestingly, each of these drug delivery systems has its own chemical, physical and morphological characteristics, and may have affinity for different drugs polarities through chemical interactions (e.g., covalent bonds and hydrogen bonds) or physical interactions (e.g., electrostatic and van der Waals interactions). As an example, Mattos et al. [ 72 ] demonstrated that, the release profile of neem bark extract-grafted biogenic silica nanoparticles (chemical interactions) was lower than neem bark extract-loaded biogenic silica nanoparticles. Hence, all these factors influence the interaction of nanocarriers with biological systems [ 73 ], as well as the release kinetics of the active ingredient in the organism [ 68 ]. In addition, Sethi et al. [ 74 ] designed a crosslinkable lipid shell (CLS) containing docetaxel and wortmannin as the prototypical drugs used for controlling the drug discharge kinetics; then, they studied, its discharge profile, which was found to be affected in both in vivo and in vitro conditions. Apart from this, other parameters, such as the composition of the nanocarriers (e.g., organic, inorganic, and hybrid materials) and the form in which drugs are associated with them (such as core–shell system or matrix system) are also fundamental for understanding their drug delivery profile [ 75 , 76 ]. Taken together, several studies regarding release mechanisms of drugs in nanocarriers have been conducted. Diffusion, solvent, chemical reaction, and stimuli-controlled release are a few mechanisms that can represent the release of drugs in nanocarriers as shown in Fig.  2 [ 77 , 78 ]. Kamaly et al. [ 79 ] provided a widespread review of controlled-release systems with a focus on studies related to controlling drug release from polymeric nanocarriers.

Mechanisms for controlled release of drugs using different types of nanocarriers

Although there are several nanocarriers with different drug release profiles, strategies are currently being formulated to improve the specificity of the nanostructures to target regions of the organism [ 80 ], and to reduce the immunogenicity through their coating or chemical functionalization with several substances, such as polymers [ 81 ], natural polysaccharides [ 82 , 83 ], antibodies [ 84 ], cell-membrane [ 85 ], and tunable surfactants [ 86 ], peptides [ 87 ], etc. In some cases where drugs do not display binding and affinity with a specific target or do not cross certain barriers (e.g. blood–brain barrier or the blood–cerebrospinal fluid barrier) [ 88 ], these ligand-modified nanocarriers have been used to pass through the cell membrane and allow a programmed drug delivery in a particular environment. For example, hyaluronic acid (a polysaccharide found in the extracellular matrix) has been used as a ligand-appended in several nanocarriers, showing promising results to boost antitumor action against the melanoma stem-like cells [ 89 ], breast cancer cells [ 90 ], pulmonary adenocarcinoma cells [ 91 ], as well as to facilitate intravitreal drug delivery for retinal gene therapy [ 83 ] and to reduce the immunogenicity of the formed protein corona [ 82 ]. However, the construction of the ligand-appended drug delivery systems is labor-intensive, and several targeting designs must be performed previously, taking into account the physiological variables of blood flow, disease status, and tissue architecture [ 92 ]. Moreover, few studies have been performed to evaluate the interaction of the ligand-appended in nanocarriers with cell membranes, and also their uptake mechanism is still unclear. Furthermore, has been known that the uptake of the nanoparticles by the cells occurs via phagocytic or non-phagocytic pathways (e.x. clathrin-mediated endocytosis, caveolae-mediated endocytosis, and others) [ 93 , 94 ], meanwhile due some particular physicochemical characteristics of each delivery systems have been difficult to standardize the mechanism of action/interaction of these systems in the cells. For example, Salatin and Khosroushahi [ 95 ], in a review highlighted the main endocytosis mechanisms responsible for the cellular uptake of polysaccharide nanoparticles containing active compounds.

On the other hand, stimuli-responsive nanocarriers have shown the ability to control the release profile of drugs (as a triggered release) using external factors such as ultrasound [ 96 ], heat [ 97 , 98 , 99 ], magnetism [ 100 , 101 ], light [ 102 ], pH [ 103 ], and ionic strength [ 104 ], which can improve the targeting and allow greater dosage control (Fig.  2 ). For example, superparamagnetic iron oxide nanoparticles are associated with polymeric nanocarriers [ 105 ] or lipids [ 106 ] to initially stimulate a controlled release system by the application of external magnetic field. In addition, Ulbrich et al. [ 107 ] revised recent achievements of drug delivery systems, in particular, on the basis of polymeric and magnetic nanoparticles, and also addressed the effect of covalently or noncovalently attached drugs for cancer cure [ 107 ]. Moreover, Au/Fe 3 O 4 @polymer nanoparticles have also been synthesized for the use in NIR-triggered chemo-photothermal therapy [ 108 ]. Therefore, hybrid nanocarriers are currently among the most promising tools for nanomedicine as they present a mixture of properties of different systems in a single system, thus ensuring materials with enhanced performance for both therapeutic and diagnostic applications (i.e., theranostic systems). Despite this, little is known about the real mechanisms of action and toxicity of drug delivery systems, which open opportunity for new studies. In addition, studies focusing on the synthesis of nanocarriers based on environmentally safe chemical reactions by implementing plant extracts and microorganisms have increased [ 10 ].

Nanoparticles used in drug delivery system

Biopolymeric nanoparticles.

There are numerous biopolymeric materials that are utilized in the drug delivery systems. These materials and their properties are discussed below.

Chitosan exhibits muco-adhesive properties and can be used to act in the tight epithelial junctions. Thus, chitosan-based nanomaterials are widely used for continued drug release systems for various types of epithelia, including buccal [ 109 ], intestinal [ 110 ], nasal [ 111 ], eye [ 112 ] and pulmonary [ 113 ]. Silva et al. [ 114 ] prepared and evaluated the efficacy of a 0.75% w/w isotonic solution of hydroxypropyl methylcellulose (HPMC) containing chitosan/sodium tripolyphosphate/hyaluronic acid nanoparticles to deliver the antibiotic ceftazidime to the eye. The rheological synergism parameter was calculated by calculating the viscosity of the nanoparticles in contact with mucin in different mass proportions. A minimum viscosity was observed when chitosan nanoparticles were placed in contact with mucin. However, the nanoparticles presented mucoadhesion which resulted in good interaction with the ocular mucosa and prolonged release of the antibiotic, and therefore, the nanoparticles can enhance the life span of the drug in the eyes. The nanoparticles did not show cytotoxicity for two cell lines tested (ARPE-19 and HEK 239T). The nanoparticles were also able to preserve the antibacterial activity, thus making them a promising formulations for the administration of ocular drugs with improved mucoadhesive properties.

Pistone et al. [ 115 ] prepared nanoparticles of chitosan, alginate and pectin as potential candidates for the administration of drugs into the oral cavity. The biocompatibility of the formulations was estimated based on the solubility of the nanoparticles in a salivary environment and its cytotoxicity potential was estimated in an oral cell line. Alginate nanoparticles were the most unwavering in the artificial saliva for at least 2 h, whereas pectin and especially chitosan nanoparticles were unstable. However, the chitosan nanoparticles were the most cyto-competitive, whereas alginate and pectin nanoparticles showed cytotoxicity under all tested conditions (concentration and time). The presence of Zn 2+ (cross-linking agent) may be the cause of the observed cytotoxicity. Each formulation presented advantage and limitations for release into the oral cavity, thus necessitating their further refinement.

In addition, Liu et al. [ 116 ] prepared nanoparticles of carboxymethyl chitosan for the release of intra-nasal carbamazepine (CBZ) to bypass the blood–brain barrier membrane, thus increasing the amount of the medication in the brain and refining the treatment efficacy, thereby reducing the systemic drug exposure. The nanoparticles had a mean diameter of 218.76 ± 2.41 nm, encapsulation efficiency of 80% and drug loading of 35%. Concentrations of CBZ remained higher (P < 0.05) in the brain than the plasma over 240 min.

In another example, Jain and Jain [ 117 ] investigated the discharge profile of 5-fluorouracil (5-FU) from hyaluronic acid-coated chitosan nanoparticles into the gut, via oral administration. Release assays in conditions mimicking the transit from the stomach to the colon indicated the release profile of 5-FU which was protected against discharge in the stomach and small intestine. Also, the high local concentration of drugs would be able to increase the exposure time and thus, enhance the capacity for antitumor efficacy and decrease the systemic toxicity in the treatment of colon cancer.

Another biopolymeric material that has been used as a drug delivery is alginate. This biopolymer presents final carboxyl groups, being classified as anionic mucoadhesive polymer and presents greater mucoadhesive strength when compared with cationic and neutral polymers [ 59 , 118 ]. Patil and Devarajan [ 119 ] developed insulin-containing alginate nanoparticles with nicotinamide as a permeation agent in order to lower the serum glucose levels and raise serum insulin levels in diabetic rats. Nanoparticles administered sublingually (5 IU/kg) in the presence of nicotinamide showed high availability pharmacology (> 100%) and bioavailability (> 80%). The fact that NPs are promising carriers of insulin via the sublingual route have been proved in case of the streptozotocin-induced diabetic mouse model by achieving a pharmacological high potential of 20.2% and bio-availability of 24.1% compared to the subcutaneous injection at 1 IU/kg [ 119 ].

Also, Haque et al. [ 120 ] prepared alginate nanoparticles to release venlafaxine (VLF) via intranasal for treatment of depression. The higher blood/brain ratios of the VLF concentration to the alginate nanoparticles administered intra-nasally when compared to the intranasal VLF and VLF solution intravenously indicated the superiority of the nano-formulation in directly transporting the VLF to the brain. In this way, these nanoparticles are promising for the treatment of depression. In another example, Román et al. [ 121 ] prepared alginate microcapsules containing epidermal growth factor bound on its exterior part to target the non-small cell lung cancer cells. Cisplatin (carcinogen drug) was also loaded in the nanoparticles. The addition of EGF significantly increased specificity of carrier systems and presented kinetics of cell death (H460-lung cancer strain) faster than the free drug.

In addition, Garrait et al. [ 122 ] prepared nanoparticles of chitosan containing Amaranth red (AR) and subsequently microencapsulated these nanoparticles in alginate microparticles and studied the release kinetics of this new system in simulated gastric and intestinal fluids. The microparticles had a mean diameter of 285 μm with a homogeneous distribution; it was observed that there was a release of less than 5% of the AR contained in the systems in the gastric pH conditions, whereas the discharge was fast and comprehensive in the intestinal pH conditions. Thus, the carrier showed promise to protect molecules for intestinal release after oral administration.

Costa et al. [ 123 ] prepared chitosan-coated alginate nanoparticles to enhance the permeation of daptomycin into the ocular epithelium aiming for an antibacterial effect. In vitro permeability was assessed using ocular epithelial cell culture models. The antimicrobial activity of nanoencapsulated daptomycin showed potential over the pathogens engaged in bacterial endophthalmitis. Also, the ocular permeability studies demonstrated that with 4 h of treatment from 9 to 12% in total of daptomycin encapsulated in chitosan/alginate nanoparticles, these were able to cross the HCE and ARPE-19 cells. These results indicated that with this system an increasing in the drug retention in the ocular epithelium has occurred.

Xanthan gum

Xanthan gum (XG) is a high molecular weight heteropolysaccharide produced by Xanthomonas campestris . It is a polyanionic polysaccharide and has good bioadhesive properties. Because it is considered non-toxic and non-irritating, xanthan gum is widely used as a pharmaceutical excipient [ 124 ].

Laffleur and Michalek [ 125 ] have prepared a carrier composed of xanthan gum thiolated with l -cysteine to release tannin in the buccal mucosa to treat sialorrhea. Thiolation of xanthan gum resulted in increased adhesion on the buccal mucosa when compared to native xanthan gum. In addition, xanthan gum thiolate has a higher uptake of saliva whereas tannic acid ad-string and dry the oral mucosa. In this way, this system would be an efficient way of reducing the salivary flow of patients with sialorrhea. Angiogenesis is an important feature in regeneration of soft tissues.

Huang et al. [ 126 ] prepared injectable hydrogels composed of aldehyde-modified xanthan and carboxymethyl-modified chitosan containing potent angiogenic factor (antivascular endothelial growth factor, VEGF) to improve abdominal wall reconstruction. The hydrogel presented release properties mainly in tissues like digestive tract and open wounds. The hydrogel containing VEGF was able to accelerate the angiogenesis process and rebuild the abdominal wall. Menzel et al. [ 127 ] studied a new excipient aiming the use as nasal release system. Xanthan gum was used as a major polymer in which the-((2-amino-2-carboxyethyl) disulfanyl) nicotinic acid (Cys-MNA) was coupled. Characteristics, such as amount of the associated binder, mucoadhesive properties and stability against degradation, were analyzed in the resulting conjugate. Each gram of polymer was ligated with 252.52 ± 20.54 μmol of the binder. The muco-adhesion of the grafted polymer was 1.7 fold greater than that of thiolated xanthan and 2.5 fold greater than, that of native xanthan. In addition, the frequency of ciliary beating of nasal epithelial cells was poorly affected and was reversible only upon the removal of the polymer from the mucosa.

Cellulose and its derivatives are extensively utilized in the drug delivery systems basically for modification of the solubility and gelation of the drugs that resulted in the control of the release profile of the same [ 128 ]. Elseoud et al. [ 129 ] investigated the utilization of cellulose nanocrystals and chitosan nanoparticles for the oral releasing of repaglinide (an anti-hyperglycemic—RPG). The chitosan nanoparticles showed a mean size distribution of 197 nm while the hybrid nanoparticles of chitosan and cellulose nanocrystals containing RPG. Chitosan hybrid nanoparticles and oxidized cellulose nanocrystals containing RPG had a mean diameter of 251–310 nm. The presence of the hydrogen bonds between the cellulose nanocrystals and the drug, resulted in sustained release of the same, and subsequently the nanoparticles made with oxidized cellulose nanocrystals presented lower release when compared to the nanoparticles produced with native cellulose nanocrystals.

Agarwal et al. [ 130 ] have developed a drug targeting mechanism which is based on the conjugation of calcium alginate beads with carboxymethylcellulose (CMC) loaded 5-fluoroacyl (5-FU) and is targeted to the colon. The beads with lower CMC proportions presented greater swelling and muco-adhesiveness in the simulated colonic environment. With existence of colonic enzymes there was a 90% release of 5-FU encapsulated in the beads. Hansen et al. [ 131 ] investigated four cellulose derivatives, including, meteylcellulose, hydroxypropyl methylcellulose, sodium carboxymethylcellulose and cationic hydroxyethyl cellulose for application in drug release into the nasal mucosa. The association of these cellulose derivatives with an additional excipient, was also evaluated. The drug model employed in this process was acyclovir. The viability of the polymers as excipients for nasal release applications was also scrutinized for its ciliary beat frequency (CBF) and its infusion through the tissue system of the nostril cavity. An increase in thermally induced viscosity was observed when the cellulose derivatives were mixed with polymer graft copolymer. Further an increased permeation of acyclovir into the nasal mucosa was detected when it was combined with cationic hydroxyethylcellulose. None of the cellulose derivatives caused negative effects on tissues and cells of the nasal mucosa, as assessed by CBF.

They were discovered by Alec Bangham in 1960. Liposomes are used in the pharmaceutical and cosmetics industry for the transportation of diverse molecules and are among the most studied carrier system for drug delivery. Liposomes are an engrained formulation strategy to improve the drug delivery. They are vesicles of spherical form composed of phospholipids and steroids usually in the 50–450 nm size range [ 132 ]. These are considered as a better drug delivery vehicles since their membrane structure is analogous to the cell membranes and because they facilitate incorporation of drugs in them [ 132 ]. It has also been proved that they make therapeutic compounds stable, improve their biodistribution, can be used with hydrophilic and hydrophobic drugs and are also biocompatible and biodegradable. Liposomes are divided into four types: (1) conventional type liposomes: these consists of a lipid bilayer which can make either anionic, cationic, or neutral cholesterol and phospholipids, which surrounds an aqueous core material. In this case, both the lipid bilayer and the aqueous space can be filled with hydrophobic or hydrophilic materials, respectively. (2) PEGylated types: polyethylene glycol (PEG) is incorporated to the surface of liposome to achieve steric equilibrium, (3) ligand-targeted type: ligands like antibodies, carbohydrates and peptides, are linked to the surface of the liposome or to the end of previously attached PEG chains and (4) theranostic liposome type: it is an amalgamation kind of the previous three types of liposomes and generally consists of a nanoparticle along with a targeting, imaging and a therapeutic element [ 133 ].

The typical synthesis procedure for liposomes are as follows, thin layer hydration, mechanical agitation, solvent evaporation, solvent injection and the surfactant solubilization [ 134 ]. One aspect to point out on liposomes is that the drugs that are trapped within them are not bioavailable until they are released. Therefore, their accumulation in particular sites is very important to increase drug bioavailability within the therapeutic window at the right rates and times. Drug loading in liposomes is attained by active (drug encapsulated after liposome formation) and passive (drug encapsulated during liposome formation) approaches [ 135 ]. Hydrophilic drugs such as ampicillin and, 5-fluoro-deoxyuridine are typically confined in the aqueous core of the liposome and thus, their encapsulation does not depend on any modification in the drug/lipid ratio. However, the hydrophobic ones such as Amphotericin B, Indomethacin were found in the acyl hydrocarbon chain of the liposome and thus their engulfing are subjected to the characteristics of the acyl chain [ 136 ]. Among the passive loading approaches the mechanical and the solvent dispersion method as well as the detergent removal method can be mentioned [ 135 ].

There are obstacles with the use of liposomes for drug delivery purposes in the form of the RES (reticuloendothelial system), opsonization and immunogenicity although there are factors like enhanced permeability and EPR (retention effect) that can be utilized in order to boost the drug delivery efficiency of the liposomes [ 133 , 135 ]. Once liposomes get into the body, they run into opsonins and high density lipoproteins (HDLs) and low density lipoproteins (LDLs) while circulating in the bloodstream by themselves. Opsonins (immunoglobulins and fibronectin, for example) assist RES on recognizing and eliminating liposomes. HDLs and LDLs have interactions with liposomes and decrease their stability. Liposomes tends to gather more in the sites like the liver and the spleen, this is an advantage because then a high concentration of liposomes can help treat pathogenic diseases, although in the case of cancers this can lead to a delay in the removal of lipophilic anticancer drugs. This is the reason why as mentioned at the beginning, different types of liposomes have been developed, in this case PEGylated ones. Dimov et al. [ 137 ] reported an incessant procedure of flow system for the synthesis, functionalization and cleansing of liposomes. This research consists of vesicles under 300 nm in a lab-on-chip that are useful and potential candidates for cost-intensive drugs or protein encapsulation development [ 137 ]. This is very important because costs of production also determine whether or not a specific drug can be commercialized. Liposome-based systems have now been permitted by the FDA [ 133 , 135 , 138 , 139 , 140 ].

Polymeric micelles

Polymeric micelles are nanostructures made of amphiphilic block copolymers that gather by itself to form a core shell structure in the aqueous solution. The hydrophobic core can be loaded with hydrophobic drugs (e.g. camptothecin, docetaxel, paclitaxel), at the same time the hydrophilic shell makes the whole system soluble in water and stabilizes the core. Polymeric micelles are under 100 nm in size and normally have a narrow distribution to avoid fast renal excretion, thus permitting their accumulation in tumor tissues through the EPR effect. In addition, their polymeric shell restrains nonspecific interactions with biological components. These nanostructures have a strong prospective for hydrophobic drug delivery since their interior core structure permits the assimilation of these kind of drugs resulting in enhancement of stability and bioavailability [ 141 , 142 ].

Polymeric micelles are synthesized by two approaches: (1) convenient solvent-based direct dissolution of polymer followed by dialysis process or (2) precipitation of one block by adding a solvent [ 142 , 143 ]. The factors like, hydrophobic chain size in the amphiphilic molecule, amphiphiles concentration, solvent system and temperature, affects the micelle formation [ 144 ]. The micelle assembly creation starts when minimum concentration known as the critical micelle concentration (CMC) is reached by the amphiphilic molecules [ 143 ]. At lower concentrations, the amphiphilic molecules are indeed small and occur independently [ 143 ]. Drugs are loaded within polymeric micelles by three common methodologies such as direct dissolution process, solvent evaporation process, and the dialysis process. As of the direct dissolution process, the copolymer and the drugs combine with each other by themselves in the water medium and forms a drug loaded with the micelles. While in the solvent evaporation process, the copolymer and the intended drug is dissolved using a volatile organic solvent and finally, in case of the dialysis process, both the drug in solution and the copolymer in the organic solvent are combined in the dialysis bag and then dialyzed with the formation of the micelle [ 145 ].

The targeting of the drugs using different polymeric micelles as established by various mechanism of action including the boosted penetrability and the holding effect stimuli; complexing of a definite aiming ligand molecule to the surface of the micelle; or by combination of the monoclonal antibodies to the micelle corona [ 146 ]. Polymeric micelles are reported to be applicable for both drug delivery against cancer [ 143 ] and also for ocular drug delivery [ 147 ] as shown in Fig.  3 in which a polymeric micelle is used for reaching the posterior ocular tissues [ 147 ]. In the work by Li et al. [ 148 ], dasatinib was encapsulated within nanoparticles prepared from micellation of PEG-b-PC, to treat proliferative vitreoretinopathy (PVR), their size was 55 nm with a narrow distribution and they turned out to be noncytotoxic to ARPE-19 cells. This micellar formulation ominously repressed the cell proliferation, attachment and relocation in comparison to the free drugs [ 148 ]. The polymeric micelles is habitually get into the rear eye tissues through the transcleral pathway after relevant applications (Fig.  3 ; [ 147 ]).

(the figure is reproduced from Mandal et al. [ 147 ] with required copyright permission)

Polymeric micelles used for reaching the posterior ocular tissues via the transcleral pathway after topical application

Dendrimers are highly bifurcated, monodisperse, well-defined and three-dimensional structures. They are globular-shaped and their surface is functionalized easily in a controlled way, which makes these structures excellent candidates as drug delivery agents [ 149 , 150 , 151 ]. Dendrimers can be synthesized by means of two approaches: The first one is the different route in which the dendrimer starts formation from its core and then it is extended outwards and the second is the convergent one, starts from the outside of the dendrimer [ 152 ]. Dendrimers are grouped into several kinds according to their functionalization moieties: PAMAM, PPI, liquid crystalline, core–shell, chiral, peptide, glycodendrimers and PAMAMOS, being PAMAM, the most studied for oral drug delivery because it is water soluble and it can pass through the epithelial tissue boosting their transfer via the paracellular pathway [ 153 ]. Dendrimers are limited in their clinical applications because of the presence of amine groups. These groups are positively charged or cationic which makes them toxic, hence dendrimers are usually modified in order to reduce this toxicity issue or to eliminate it. Drug loading in dendrimers is performed via the following mechanisms: Simple encapsulation, electrostatic interaction and covalent conjugation [ 154 ].

Drug is basically delivered by the dendrimers following two different paths, a) by the in vivo degradation of drug dendrimer’s covalent bonding on the basis of availability of suitable enzymes or favorable environment that could cleave the bonds and b) by discharge of the drug due to changes in the physical environment like pH, temperature etc., [ 154 ]. Dendrimers have been developed for transdermal, oral, ocular, pulmonary and in targeted drug delivery [ 155 ].

Jain et al. [ 156 ] have described the folate attached poly- l -lysine dendrimers (doxorubicin hydrochloride) as a capable cancer prevention drug carrier model for pH dependent drug discharge, target specificity, antiangiogenic and anticancer prospective, it was shown that doxorubicin-folate conjugated poly- l -lysine dendrimers increased the concentration of doxorubicin in the tumor by 121.5-fold after 24 h compared with free doxorubicin. Similarly, (Kaur et al. [ 157 ] developed folate-conjugated polypropylene imine dendrimers (FA-PPI) as a methotrexate (MTX) nanocarrier, for pH-sensitive drug release, selective targeting to cancer cells, and anticancer treatment. The in vitro studies on them showed sustained release, increased cell uptake and low cytotoxicity on MCF-7 cell lines [ 157 ]. Further, it has to be pointed out that the developed formulations, methotrexate (MTX)-loaded and folic acid-conjugated 5.0G PPI (MTX-FA-PPI), were selectively taken up by the tumor cells in comparison with the free drug, methotrexate (MTX).

Inorganic nanoparticles

Inorganic nanoparticles include silver, gold, iron oxide and silica nanoparticles are included. Studies focused on them are not as many as there are on other nanoparticle types discussed in this section although they show some potential applications. However, only few of the nanoparticles have been accepted for its clinical use, whereas the majority of them are still in the clinical trial stage. Metal nanoparticles, silver and gold, have particular properties like SPR (surface plasmon resonance), that liposomes, dendrimers, micelles do not possess. They showed several advantages such as good biocompatibility and versatility when it comes to surface functionalization.

Studies on their drug delivery-related activity have not been able to clear out whether the particulate or ionized form is actually related to their toxicity, and even though two mechanisms have been proposed, namely paracellular transport and transcytosis, there is not enough information about their in vivo transport and uptake mechanism [ 158 ]. Drugs can be conjugated to gold nanoparticles (AuNPs) surfaces via ionic or covalent bonding and physical absorption and they can deliver them and control their release through biological stimuli or light activation [ 159 ]. Silver nanoparticles exhibited antimicrobial activity, but as for drug delivery, very few studies have been carried out, for example, Prusty and Swain [ 160 ] synthesized an inter-linked and spongy polyacrylamide/dextran nano-hydrogels hybrid system with covalently attached silver nanoparticles for the release of ornidazole which turned out to have an in vitro release of 98.5% [ 160 ]. Similarly in another study, the iron oxide nanoparticles were synthesized using laser pyrolysis method and were covered with Violamycine B1, and antracyclinic antibiotics and tested against the MCF-7 cells for its cytotoxicity and the anti-proliferation properties along with its comparison with the commercially available iron oxide nanoparticles [ 161 ].

Nanocrystals

Nanocrystals are pure solid drug particles within 1000 nm range. These are 100% drug without any carriers molecule attached to it and are usually stabilized by using a polymeric steric stabilizers or surfactants. A nanocrystals suspension in a marginal liquid medium is normally alleviated by addition of a surfactant agent known as nano-suspension. In this case, the dispersing medium are mostly water or any aqueous or non-aqueous media including liquid polyethylene glycol and oils [ 162 , 163 ]. Nanocrystals possesses specific characters that permit them to overcome difficulties like increase saturation solubility, increased dissolution velocity and increased glueyness to surface/cell membranes. The process by which nanocrystals are synthesized are divided into top-down and bottom-up approaches. The top-down approach includes, sono-crystallization, precipitation, high gravity controlled precipitation technology, multi-inlet vortex mixing techniques and limited impinging liquid jet precipitation technique [ 162 ]. However, use of an organic solvent and its removal at the end makes this process quite expensive. The bottom-up approach involves, grinding procedures along with homogenization at higher pressure [ 162 ]. Among all of the methods, milling, high pressure homogenization, and precipitation are the most used methods for the production of nanocrystals. The mechanisms by which nanocrystals support the absorption of a drug to the system includes, enhancement of solubility, suspension rate and capacity to hold intestinal wall firmly [ 162 ]. Ni et al. [ 164 ] embedded cinaciguat nanocrystals in chitosan microparticles for pulmonary drug delivery of the hydrophobic drug. The nanoparticles were contrived for continuous release of the drug taking advantage of the swelling and muco-adhesive potential of the polymer. They found that inhalation efficacy might be conceded under the disease conditions, so more studies are needed to prove that this system has more potential [ 164 ].

Metallic nanoparticles

In recent years, the interest of using metallic nanoparticles has been growing in different medical applications, such as bioimaging, biosensors, target/sustained drug delivery, hyperthermia and photoablation therapy [ 35 , 165 ]. In addition, the modification and functionalization of these nanoparticles with specific functional groups allow them to bind to antibodies, drugs and other ligands, become these making these systems more promising in biomedical applications [ 166 ]. Although the most extensively studied, metallic nanoparticles are gold, silver, iron and copper, a crescent interest has been exploited regarding other kinds of metallic nanoparticles, such as, zinc oxide, titanium oxide, platinum, selenium, gadolinium, palladium, cerium dioxide among others [ 35 , 165 , 166 ].

Quantum dots

Quantum dots (QDs) are known as semiconductor nanocrystals with diameter range from 2 to 10 nm and their optical properties, such as absorbance and photoluminescence are size-dependent [ 167 ]. The QDs has gained great attention in the field of nanomedicine, since, unlike conventional organic dyes, the QDs presents emission in the near-infrared region (< 650 nm), a very desirable characteristic in the field of biomedical images, due to the low absorption by the tissues and reduction in the light scattering [ 167 , 168 ]. In addition, QDs with different sizes and/or compositions can be excited by the same light source resulting in separate emission colors over a wide spectral range [ 169 , 170 ]. In this sense, QDs are very appealing for multiplex imaging. In the medicine field QDs has been extensively studied as targeted drug delivery, sensors and bioimaging. A large number of studies regarding the applications of QDs as contrast agents for in vivo imaging is currently available in literature [ 168 , 171 , 172 , 173 ]. Han et al. [ 172 ] developed a novel fluorophore for intravital cytometric imaging based on QDs-antibodies conjugates coated with norbornene-displaying polyimidazole ligands. This fluorophore was used to label bone marrow cells in vivo. The authors found that the fluorophore was able to diffuse in the entire bone marrow and label rare populations of cells, such as hematopoietic stem and progenitor cells [ 172 ]. Shi et al. [ 171 ] developed a multifunctional biocompatible graphene oxide quantum dot covered with luminescent magnetic nanoplatform for recognize/diagnostic of a specific liver cancer tumor cells (glypican-3-expressing Hep G2). According to the authors the attachment of an anti-GPC3-antibody to the nanoplataform results in selective separation of Hep G2 hepatocellular carcinoma cells from infected blood samples [ 171 ]. QDs could also bring benefits in the sustained and/or controlled release of therapeutic molecules. Regarding the controlled release, this behavior can be achieved via external stimulation by light, heat, radio frequency or magnetic fields [ 170 , 174 , 175 ]. Olerile et al. [ 176 ] have developed a theranostic system based on co-loaded of QDs and anti-cancer drug in nanostructured lipid carriers as a parenteral multifunctional system. The nanoparticles were spherical with higher encapsulation efficiency of paclitaxel (80.7 ± 2.11%) and tumor growth inhibition rate of 77.85%. The authors also found that the system was able to specifically target and detect H22 tumor cells [ 176 ]. Cai et al. [ 177 ] have synthesized pH responsive quantum dots based on ZnO quantum dots decorated with PEG and hyaluronic acid for become stable in physiological conditions and for targeting specific cells with HA-receptor CD44, respectively. This nanocarrier was also evaluated for doxorubicin (DOX) sustained release. The nanocarrier was stable in physiological pH and DOX was loaded in the carrier by forming complex with Zn 2+ ions or conjugated to PEG. The DOX was released only in acidic intracellular conditions of tumor cells due to the disruption of ZnO QDs. The authors found that the anticancer activity was enhanced by the combination of DOX and ZnO QDs [ 177 ].

Protein and polysaccharides nanoparticles

Polysaccharides and proteins are collectively called as natural biopolymers and are extracted from biological sources such as plants, animals, microorganisms and marine sources [ 178 , 179 ]. Protein-based nanoparticles are generally decomposable, metabolizable, and are easy to functionalize for its attachment to specific drugs and other targeting ligands. They are normally produced by using two different systems, (a) from water-soluble proteins like bovine and human serum albumin and (b) from insoluble ones like zein and gliadin [ 180 ]. The usual methods to synthesize them are coacervation/desolvation, emulsion/solvent extraction, complex coacervation and electrospraying. The protein based nanoparticles are chemically altered in order to combine targeting ligands that identify exact cells and tissues to promote and augment their targeting mechanism [ 180 ]. Similarly, the polysaccharides are composed of sugar units (monosaccharides) linked through O-glycosidic bonds. The composition of these monomers as well as their biological source are able to confer to these polysaccharides, a series of specific physical–chemical properties [ 126 , 179 , 181 ]. One of the main drawback of the use of polysaccharides in the nanomedicine field is its degradation (oxidation) characteristics at high temperatures (above their melting point) which are often required in industrial processes. Besides, most of the polysaccharides are soluble in water, which limits their application in some fields of nanomedicine, such as tissue engineering [ 182 , 183 ]. However, techniques such as crosslinking of the polymer chains have been employed in order to guarantee stability of the polysaccharide chains, guaranteeing them stability in aqueous environments [ 182 , 183 ]. In Fig.  4 , examples of some polysaccharides used in nanomedicine obtained from different sources are summarized. The success of these biopolymers in nanomedicine and drug delivery is due to their versatility and specified properties such as since they can originate from soft gels, flexible fibers and hard shapes, so they can be porous or non-porous; they have great similarity with components of the extracellular matrix, which may be able to avoid immunological reactions [ 179 , 184 ].

Different sources of natural biopolymers to be used in nanomedicine applications. Natural biopolymers could be obtained from higher plants, animals, microorganisms and algae

There is not much literature related to these kind of nanoparticles, however, since they are generated from biocompatible compounds they are excellent candidates for their further development as drug delivery systems. Yu et al. [ 185 ] synthesized Bovine serum albumin and tested its attachment and/or infiltration property through the opening of the cochlea and middle ear of guinea pigs. The nanoparticles considered as the drug transporters were tested for their loading capacity and release behaviors that could provide better bio-suitability, drug loading capacity, and well-ordered discharge mechanism [ 185 ].

Natural product-based nanotechnology and drug delivery

As per the World Health Organization (WHO) report, in developing countries, the basic health needs of approximately 80% of the population are met and/or complemented by traditional medicine [ 186 ]. Currently, the scientific community is focusing on the studies related to the bioactive compounds, its chemical composition and pharmacological potential of various plant species, to produce innovative active ingredients that present relatively minor side effects than existing molecules [ 5 , 187 ]. Plants are documented as a huge sources of natural compounds of medicinal importance since long time and still it holds ample of resources for the discovery of new and highly effective drugs. However, the discovery of active compounds through natural sources is associated with several issues because they originate from living beings whose metabolite composition changes in the presence of stress. In this sense, the pharmaceutical industries have chosen to combine their efforts in the development of synthetic compounds [ 187 , 188 , 189 ]. Nevertheless, the number of synthetic molecules that are actually marketed are going on decreasing day by day and thus research on the natural product based active compounds are again coming to the limelight in spite of its hurdles [ 189 , 190 ]. Most of the natural compounds of economic importance with medicinal potential that are already being marketed have been discovered in higher plants [ 187 , 191 ]. Several drugs that also possess natural therapeutic agents in their composition are already available commercially; their applications and names are as follows: malaria treatment (Artemotil ® derived from Artemisia annua L., a traditional Chinese medicine plant), Alzheimer’s disease treatment (Reminyl ® , an acetylcholinesterase inhibitor isolated from the Galanthus woronowii Losinsk), cancer treatment (Paclitaxel ® and its analogues derived from the Taxus brevifolia plant; vinblastine and vincristine extracted from Catharanthus roseus ; camptothecin and its analogs derived from Camptotheca acuminata Decne), liver disease treatment (silymarin from Silybum marianum ) [ 187 ].

The composition and activity of many natural compounds have already been studied and established. The alkaloids, flavonoids, tannins, terpenes, saponins, steroids, phenolic compounds, among others, are the bioactive molecules found in plants. However in most of the cases, these compounds have low absorption capacity due to the absence of the ability to cross the lipid membranes because of its high molecular sizes, and thus resulting in reduced bioavailability and efficacy [ 192 ]. These molecules also exhibit high systemic clearance, necessitating repeated applications and/or high doses, making the drug less effective for therapeutic use [ 189 ]. The scientific development of nanotechnology can revolutionize the development of formulations based on natural products, bringing tools capable of solving the problems mentioned above that limits the application of these compounds in large scale in the nanomedicine [ 7 , 189 ]. Utilization of nanotechnology techniques in the medical field has been extensively studied in the last few years [ 193 , 194 ]. Hence these can overcome these barriers and allow different compounds and mixtures to be used in the preparation of the same formulation. In addition, they can change the properties and behavior of a compound within the biological system [ 7 , 189 ]. Besides, bringing benefits to the compound relative to the solubility and stability of the compounds, release systems direct the compound to the specific site, increase bioavailability and extend compound action, and combine molecules with varying degrees of hydrophilicity/lipophilicity [ 7 ]. Also, there is evidence that the association of release systems with natural compounds may help to delay the development of drug resistance and therefore plays an important role in order to find new possibilities for the treatment of several diseases that have low response to treatment conventional approaches to modern medicine [ 7 , 189 ].

The natural product based materials are of two categories, (1) which are targeted to specific location and released in the specific sites to treat a number of diseases [ 43 , 195 ] and (2) which are mostly utilized in the synthesis process [ 196 ]. Most of the research is intended for treatment against the cancer disease, since it is the foremost reason of death worldwide nowadays [ 197 , 198 ]. In case of the cancer disease, different organs of the body are affected, and therefore the need for the development of an alternative medicine to target the cancerous cells is the utmost priority among the modern researchers, however, a number of applications of nanomedicine to other ailments is also being worked on [ 199 , 200 ]. These delivery systems are categorized in terms of their surface charge, particle size, size dispersion, shape, stability, encapsulation potential and biological action which are further utilized as per their requirements [ 33 ]. Some examples of biological compounds obtained from higher plants and their uses in the nanomedicine field are described in Fig.  5 . Pharmaceutical industries have continuously sought the development and application of new technologies for the advancement and design of modern drugs, as well as the enhancement of existing ones [ 71 , 201 ]. In this sense, the accelerated development of nanotechnology has driven the design of new formulations through different approaches, such as, driving the drug to the site of action (nanopharmaceutics); image and diagnosis (nanodiagnostic), medical implants (nanobiomaterials) and the combination diagnosis and treatment of diseases (nanotheranostics) [ 71 , 202 , 203 ].

Examples of natural compounds extracted from higher plants used in nanomedicine aiming different approaches. Some of these extracts are already being marketed, others are in clinical trials and others are being extensively studied by the scientific community

Currently, many of the nanomedicines under development, are modified release systems for active ingredients (AI) that are already employed in the treatment of patients [ 203 , 204 ]. For this type of approach, it is evaluated whether the sustained release of these AIs modifies the pharmacokinetic profile and biodistribution. In this context, it can be ascertained that the nano-formulation offers advantages over the existing formulation if the AI is directed towards the target tissue shows increased uptake/absorption by the cells and lower toxicity profile for the organism [ 205 , 206 ]. This section is focused on berberine, curcumin, ellagic acid, resveratrol, curcumin and quercetin [ 8 ]. Some other compounds mentioned are doxorubicin, paclitaxel and vancomycin that also come from natural products.

Nanoparticles have been synthesized using natural products. For example, metallic, metal oxide and sulfides nanoparticles have been reported to be synthesized using various microorganisms including bacteria, fungi, algae, yeast and so on [ 207 ] or plant extracts [ 208 ]. For the first approach, the microorganism that aids the synthesis procedure is prepared in the adequate growth medium and then mixed with a metal precursor in solution and left for incubation to form the nanoparticles either intracellularly or extracellularly [ 209 , 210 , 211 ]. As for the second approach, the plant extract is prepared and mixed afterwards with the metal precursor in solution and incubated further at room temperature or boiling temperature for a definite time or exposed to light as an external stimulus to initiate the synthesis of nanoparticles [ 212 ].

Presently, these natural product based materials are considered as the key ingredients in the preparation and processing of new nano-formulations because they have interesting characteristics, such as being biodegradable, biocompatible, availability, being renewable and presenting low toxicity [ 178 , 179 , 213 ]. In addition to the aforementioned properties, biomaterials are, for the most part, capable of undergoing chemical modifications, guaranteeing them unique and desirable properties for is potential uses in the field of nanomedicine [ 45 , 214 ]. Gold, silver, cadmium sulfide and titanium dioxide of different morphological characteristics have been synthesized using a number of bacteria namely Escherichia coli , Pseudomonas aeruginosa , Bacillus subtilis and Klebsiella pneumoniae [ 211 ]. These nanoparticles, especially the silver nanoparticles have been abundantly studied in vitro for their antibacterial, antifungal, and cytotoxicity potential due to their higher potential among all metal nanoparticles [ 215 , 216 ]. In the event of microorganism mediated nanoparticle synthesis, maximum research is focused on the way that microorganisms reduce metal precursors and generate the nanoparticles. For instance, Rahimi et al. [ 217 ] synthesized silver nanoparticles using Candida albicans and studied their antibacterial activity against two pathogenic bacteria namely Staphylococcus aureus and E. coli. Similarly, Ali et al. [ 218 ] synthesized silver nanoparticles with the Artemisia absinthium aqueous extract and their antimicrobial activity was assessed versus Phytophthora parasitica and Phytophthora capsici [ 218 ]. Further, Malapermal et al. [ 219 ] used Ocimum basilicum and Ocimum sanctum extracts to synthesize nanoparticles and studied its antimicrobial potential against E. coli , Salmonella spp., S. aureus , and P. aeruginosa along with the antidiabetic potential. Likewise, Sankar et al. [ 220 ] also tested the effect of silver nanoparticles for both antibacterial and anticancer potential against human lung cancer cell line. Besides the use of microorganism, our group has synthesized silver, gold and iron oxide nanoparticles using various food waste materials such as extracts of Zea mays leaves [ 221 , 222 ], onion peel extract [ 223 ], silky hairs of Zea mays [ 224 ], outer peel of fruit of Cucumis melo and Prunus persica [ 225 ], outer peel of Prunus persica [ 226 ] and the rind extract of watermelon [ 227 ], etc. and have tested their potential antibacterial effects against various foodborne pathogenic bacteria, anticandidal activity against a number of pathogenic Candida spp., for their potential antioxidant activity and proteasome inhibitory effects.

For drug delivery purposes, the most commonly studied nanocarriers are crystal nanoparticles, liposomes, micelles, polymeric nanoparticles, solid lipid nanoparticles, superparamagnetic iron oxide nanoparticles and dendrimers [ 228 , 229 , 230 ]. All of these nanocarriers are formulated for natural product based drug delivery. For applications in cancer treatment, Gupta et al. [ 231 ] synthesized chitosan based nanoparticles loaded with Paclitaxel (Taxol) derived from Taxus brevifolia , and utilized them for treatment of different kinds of cancer. The authors concluded that the nanoparticle loaded drug exhibited better activity with sustained release, high cell uptake and reduced hemolytic toxicity compared with pure Paclitaxel [ 231 ]. Berberine is an alkaloid from the barberry plant. Chang et al. [ 232 ] created a heparin/berberine conjugate to increase the suppressive Helicobacter pylori growth and at the same time to reduce cytotoxic effects in infected cells [ 232 ] which is depicted in Fig.  6 .

(the figure is reproduced from Chang et al. [ 232 ] with required copyright permission)

a Structure of berberine/heparin based nanoparticles and berberine/heparin/chitosan nanoparticles. b TEM images of the berberine/heparin nanoparticles and berberine/heparin/chitosan nanoparticles

Aldawsari and Hosny [ 233 ] synthesized ellagic acid-SLNs to encapsulate Vancomycin (a glycopeptide antibiotic produced in the cultures of Amycolatopsis orientalis ). Further, its in vivo tests were performed on rabbits and the results indicated that the ellagic acid prevented the formation of free oxygen radicals and their clearance radicals, thus preventing damages and promoting repair [ 233 ]. Quercetin is a polyphenol that belongs to the flavonoid group, it can be found in citrus fruits and vegetables and it has antioxidant properties. In a study by Dian et al. [ 234 ], polymeric micelles was used to deliver quercetin and the results showed that such micelles could provide continuous release for up to 10 days in vitro, with continuous plasma level and boosted complete accessibility of the drug under in vivo condition [ 234 ].

Daunorubicin is a natural product derived from a number of different wild type strains of Streptomyces , doxorubicin (DOX) is a hydrolated version of it used in chemotherapy [ 213 ]. Spillmann et al. [ 235 ] developed a multifunctional liquid crystal nanoparticle system for intracellular fluorescent imaging and for the delivery of doxorubicin in which the nanoparticles were functionalized with transferrin. Cellular uptake and sustained released were attained within endocytic vesicles in HEK 293T/17 cells. Perylene was used as a chromophore to track the particles and to encapsulate agents aimed for intracellular delivery [ 235 ]. Purama et al. [ 236 ] extracted dextran from two sucrose based lactic acid bacteria namely Streptococcus mutans and Leuconostoc mesenteroides . Agarwal et al. [ 237 ] formulated a dextran-based dendrimer formulation and evaluated its drug discharge capacity and haemolytic activity under in vitro condition. They concluded that the dendritic structure selectively enters the highly permeable portion of the affected cells without disturbing the healthy tissues thereby making more convenient for its application in the biomedical field [ 237 ]. Folate- functionalized superparamagnetic iron oxide nanoparticles developed previously for liver cancer cure are also been used for the delivery of Doxil (a form of doxorubicin which was the first FDA-approved nano-drug in 1995) [ 238 ]. The in vivo studies in rabbits and rats showed a two- and fourfold decrease compared with Doxil alone while folate aided and enhanced specific targeting [ 239 ]. Liposomes are the nanostructures that have been studied the most, and they have been used in several formulations for the delivery of natural products like resveratrol [ 240 ]. Curcumin, a polyphenolic compound obtained from turmeric, have been reported to be utilized in the cure of cancers including the breast, bone, cervices, liver, lung, and prostate [ 241 ]. Liposomal curcumin formulations have been developed for the treatment of cancer [ 242 , 243 ]. Cheng et al. [ 244 ] encapsulated curcumin in liposomes by different methods and compared the outcomes resulting that the one dependent on pH yielded stable products with good encapsulation efficiency and bio-accessibility with potential applications in cancer treatment [ 244 ].

Over all, it can be said that the sustained release systems of naturally occurring therapeutic compounds present themselves as a key tools for improving the biological activity of these compounds as well as minimizing their limitations by providing new alternatives for the cure of chronic and terminal diseases [ 8 , 245 ]. According to BBC Research, the global market for plant-derived pharmaceuticals will increase from $29.4 billion in 2017 to about $39.6 billion in 2022 with a compound annual growth rate (CAGR) of 6.15% in this period (BCC-RESEARCH). Some of nanostructure-based materials covered in this section have already been approved by the FDA. Bobo et al. [ 255 ] has provided the information on nanotechnology-based products already approved by the FDA (Table  1 ).

Regulation and reality: products now on the market

In the current medical nanotechnology scenario, there are 51 products based on this technology [ 204 , 246 , 247 , 248 ] which are currently being applied in clinical practice (Table  2 ). Notably, such nanomedicines are primarily developed for drugs, which have low aqueous solubility and high toxicity, and these nanoformulations are often capable of reducing the toxicity while increasing the pharmacokinetic properties of the drug in question.

According to a recent review by Caster et al. [ 249 ], although few nanomedicines have been regulated by the FDA there are many initiatives that are currently in progress in terms of clinical trials suggesting many nanotechnology-based new drugs will soon be able to reach the market. Among these nanomaterials that are in phase of study, 18 are directed to chemotherapeutics; 15 are intended for antimicrobial agents; 28 are for different medical applications and psychological diseases, autoimmune conditions and many others and 30 are aimed at nucleic acid based therapies [ 249 ]. The list of nanomedicine approved by FDA classified by type of carrier/material used in preparation of the formulation is shown in Table  2 .

Nanotechnology has dynamically developed in recent years, and all countries, whether developed or not, are increasing their investments in research and development in this field. However, researchers who work with practical applications of the nano-drugs deal with high levels of uncertainties, such as a framing a clear definition of these products; characterization of these nanomaterials in relation to safety and toxicity; and the lack of effective regulation. Although the list of approved nanomedicine is quite extensive, the insufficiency of specific regulatory guidelines for the development and characterization of these nanomaterials end up hampering its clinical potential [ 250 ]. The structure/function relationships of various nanomaterials, as well as their characteristics, composition and surface coating, interacts with the biological systems. In addition, it is important to evaluate the possibility of aggregate and agglomerate formation when these nanomedicines are introduced into biological systems, since they do not reflect the properties of the individual particle; this may generate different results and/or unexpected toxic effects depending on the nano-formulation [ 250 ].

The lack of standard protocols for nanomedicines characterization at physico-chemical and physiological/biological levels has often limited the efforts of many researchers to determine the toxic potential of nano-drugs in the early stages of testing, and that resulted in the failures in late-phase clinical trials. To simplify and/or shorten the approval process for nano based medicines/drugs, drug delivery system etc., a closer cooperation among regulatory agencies is warranted [ 204 , 251 ].

As a strategy for the lack of regulation of nanomedicines and nano drug delivery system; the safety assessment and the toxicity and compatibility of these are performed based on the regulations used by the FDA for conventional drugs. After gaining the status of a new research drug (Investigational New Drug, IND) by the FDA, nanomedicines, nano-drug delivery systems begin the clinical trials phase to investigate their safety and efficacy in humans. These clinical trials are divided into three phases: phase 1 (mainly assesses safety); phase 2 (mainly evaluates efficacy) and phase 3 (safety, efficacy and dosage are evaluated). After approval in these three phases the IND can be filed by the FDA to request endorsement of the new nanomedicine or nano drug delivery systems. However, this approach to nanomedicine regulation has been extensively questioned [ 204 , 246 , 252 ].

Due to the rapid development of nanotechnology as well as its potential use of nanomedicine, a reformed and more integrated regulatory approach is urgently required. In this regard, country governments must come together to develop new protocols that must be specific and sufficiently rigorous to address any safety concerns, thus ensuring the release of safe and beneficial nanomedicine for patients [ 204 , 252 , 253 ].

Future of nanomedicine and drug delivery system

The science of nanomedicine is currently among the most fascinating areas of research. A lot of research in this field in the last two decades has already led to the filling of 1500 patents and completion of several dozens of clinical trials [ 254 ]. As outlined in the various sections above, cancer appears to be the best example of diseases where both its diagnosis and therapy have benefited from nonmedical technologies. By using various types of nanoparticles for the delivery of the accurate amount of drug to the affected cells such as the cancer/tumour cells, without disturbing the physiology of the normal cells, the application of nanomedicine and nano-drug delivery system is certainly the trend that will remain to be the future arena of research and development for decades to come.

The examples of nanoparticles showed in this communications are not uniform in their size, with some truly measuring in nanometers while others are measured in sub-micrometers (over 100 nm). More research on materials with more consistent uniformity and drug loading and release capacity would be the further area of research. Considerable amount of progress in the use of metals-based nanoparticles for diagnostic purposes has also been addressed in this review. The application of these metals including gold and silver both in diagnosis and therapy is an area of research that could potentially lead to wider application of nanomedicines in the future. One major enthusiasm in this direction includes the gold-nanoparticles that appear to be well absorbed in soft tumour tissues and making the tumour susceptible to radiation (e.g., in the near infrared region) based heat therapy for selective elimination.

Despite the overwhelming understanding of the future prospect of nanomedicine and nano-drug delivery system, its real impact in healthcare system, even in cancer therapy/diagnosis, remains to be very limited. This attributes to the field being a new area of science with only two decades of real research on the subject and many key fundamental attributes still being unknown. The fundamental markers of diseased tissues including key biological markers that allow absolute targeting without altering the normal cellular process is one main future area of research. Ultimately, the application of nanomedicine will advance with our increasing knowledge of diseases at molecular level or that mirrors a nanomaterial-subcellular size comparable marker identification to open up avenues for new diagnosis/therapy. Hence, understanding the molecular signatures of disease in the future will lead to advances in nanomedicine applications. Beyond what we have outlined in this review using the known nanoprobes and nanotheragnostics products, further research would be key for the wider application of nanomedicine.

The concept of controlled release of specific drugs at the beleaguered sites, technology for the assessment of these events, drug’s effect in tissues/cellular level, as well as theoretical mathematical models of predication have not yet been perfected. Numerous studies in nanomedicine areas are centered in biomaterials and formulation studies that appear to be the initial stages of the biomedicine applications. Valuable data in potential application as drug therapeutic and diagnosis studies will come from animal studies and multidisciplinary researches that requires significant amount of time and research resources. With the growing global trend to look for more precise medicines and diagnosis, the future for a more intelligent and multi-centered approach of nanomedicine and nano-drug delivery technology looks bright.

There has been lots of enthusiasm with the simplistic view of development of nanorobots (and nanodevices) that function in tissue diagnosis and repair mechanism with full external control mechanism. This has not yet been a reality and remains a futuristic research that perhaps could be attained by mankind in the very near future. As with their benefits, however, the potential risk of nanomedicines both to humans and the environment at large require long term study too. Hence, proper impact analysis of the possible acute or chronic toxicity effects of new nanomaterials on humans and environment must be analyzed. As nanomedicines gain popularity, their affordability would be another area of research that needs more research input. Finally, the regulation of nanomedicines, as elaborated in the previous section will continue to evolve alongside the advances in nanomedicine applications.

The present review discusses the recent advances in nanomedicines, including technological progresses in the delivery of old and new drugs as well as novel diagnostic methodologies. A range of nano-dimensional materials, including nanorobots and nanosensors that are applicable to diagnose, precisely deliver to targets, sense or activate materials in live system have been outlined. Initially, the use of nanotechnology was largely based on enhancing the solubility, absorption, bioavailability, and controlled-release of drugs. Even though the discovery of nanodrugs deal with high levels of uncertainties, and the discovery of pharmacologically active compounds from natural sources is not a favored option today, as compared to some 50 years ago; hence enhancing the efficacy of known natural bioactive compounds through nanotechnology has become a common feature. Good examples are the therapeutic application of nanotechnology for berberine, curcumin, ellagic acid, resveratrol, curcumin and quercetin. The efficacy of these natural products has greatly improved through the use of nanocarriers formulated with gold, silver, cadmium sulphide, and titanium dioxide polymeric nanoparticles together with solid lipid nanoparticles, crystal nanoparticles, liposomes, micelles, superparamagnetic iron oxide nanoparticles and dendrimers.

There has been a continued demand for novel natural biomaterials for their quality of being biodegradable, biocompatible, readily availability, renewable and low toxicity. Beyond identifying such polysaccharides and proteins natural biopolymers, research on making them more stable under industrial processing environment and biological matrix through techniques such as crosslinking is among the most advanced research area nowadays. Polymeric nanoparticles (nanocapsules and nanospheres) synthesized through solvent evaporation, emulsion polymerization and surfactant-free emulsion polymerization have also been widely introduced. One of the great interest in the development of nanomedicine in recent years relates to the integration of therapy and diagnosis (theranostic) as exemplified by cancer as a disease model. Good examples have been encapsulated such as, oleic acid-coated iron oxide nanoparticles for diagnostic applications through near-infrared; photodynamic detection of colorectal cancer using alginate and folic acid based chitosan nanoparticles; utilization of cathepsin B as metastatic processes fluorogenic peptide probes conjugated to glycol chitosan nanoparticles; iron oxide coated hyaluronic acid as a biopolymeric material in cancer therapy; and dextran among others.

Since the 1990s, the list of FDA-approved nanotechnology-based products and clinical trials has staggeringly increased and include synthetic polymer particles; liposome formulations; micellar nanoparticles; protein nanoparticles; nanocrystals and many others often in combination with drugs or biologics. Even though regulatory mechanisms for nanomedicines along with safety/toxicity assessments will be the subject of further development in the future, nanomedicine has already revolutionized the way we discover and administer drugs in biological systems. Thanks to advances in nanomedicine, our ability to diagnose diseases and even combining diagnosis with therapy has also became a reality.

Abbreviations

Amaranth red

ciliary beat frequency

carbamazepine

colorectal cancer

carboxymethylcellulose

((2-amino-2-carboxyethyl) disulfanyl) nicotinic acid (Cys-MNA)

penetrability and holding

folic acid-conjugated dextran

Food and Drug Administration

ferrous oxide

hyaluronic acid

high density lipoproteins

hydroxypropylmethylcellulose

low density lipoproteins

magnetic resonance

near infrared

nanoparticle

perfluorohexane

repaglidine

antivascular endothelial growth factor

venlafaxine

xanthan gum

Swamy MK, Sinniah UR. Patchouli (Pogostemon cablin Benth.): botany, agrotechnology and biotechnological aspects. Ind Crops Prod. 2016;87:161–76.

Article   CAS   Google Scholar  

Mohanty SK, Swamy MK, Sinniah UR, Anuradha M. Leptadenia reticulata (Retz.) Wight & Arn. (Jivanti): botanical, agronomical, phytochemical, pharmacological, and biotechnological aspects. Molecules. 1019;2017:22.

Google Scholar  

Rodrigues T, Reker D, Schneider P, Schneider G. Counting on natural products for drug design. Nat Chem. 2016;8:531.

Article   CAS   PubMed   Google Scholar  

Siddiqui AA, Iram F, Siddiqui S, Sahu K. Role of natural products in drug discovery process. Int J Drug Dev Res. 2014;6(2):172–204.

CAS   Google Scholar  

Beutler JA. Natural products as a foundation for drug discovery. Curr Prot Pharmacol. 2009;46(1):9–11.

Thilakarathna SH, Rupasinghe H. Flavonoid bioavailability and attempts for bioavailability enhancement. Nutrients. 2013;5:3367–87.

Article   PubMed   PubMed Central   CAS   Google Scholar  

Bonifácio BV, da Silva PB, dos Santos Ramos MA, Negri KMS, Bauab TM, Chorilli M. Nanotechnology-based drug delivery systems and herbal medicines: a review. Int J Nanomed. 2014;9:1.

Watkins R, Wu L, Zhang C, Davis RM, Xu B. Natural product-based nanomedicine: recent advances and issues. Int J Nanomed. 2015;10:6055.

Martinho N, Damgé C, Reis CP. Recent advances in drug delivery systems. J Biomater Nanobiotechnol. 2011;2:510.

Jahangirian H, Lemraski EG, Webster TJ, Rafiee-Moghaddam R, Abdollahi Y. A review of drug delivery systems based on nanotechnology and green chemistry: green nanomedicine. Int J Nanomed. 2017;12:2957.

Liu Z, Tabakman S, Welsher K, Dai H. Carbon nanotubes in biology and medicine: in vitro and in vivo detection, imaging and drug delivery. Nano Res. 2009;2:85–120.

Article   CAS   PubMed   PubMed Central   Google Scholar  

Orive G, Gascon AR, Hernández RM, Domı́nguez-Gil A, Pedraz JL. Techniques: new approaches to the delivery of biopharmaceuticals. Trends Pharmacol Sci. 2004;25:382–7.

Razzacki SZ, Thwar PK, Yang M, Ugaz VM, Burns MA. Integrated microsystems for controlled drug delivery. Adv Drug Deliv Rev. 2004;56:185–98.

Article   PubMed   CAS   Google Scholar  

Arayne MS, Sultana N, Qureshi F. nanoparticles in delivery of cardiovascular drugs. Pak J Pharm Sci. 2007;20:340–8.

CAS   PubMed   Google Scholar  

Patra JK, Baek K-H. Green nanobiotechnology: factors affecting synthesis and characterization techniques. J Nanomater. 2014;2014:219.

Joseph RR, Venkatraman SS. Drug delivery to the eye: what benefits do nanocarriers offer? Nanomedicine. 2017;12:683–702.

Mirza AZ, Siddiqui FA. Nanomedicine and drug delivery: a mini review. Int Nano Lett. 2014;4:94.

Rudramurthy GR, Swamy MK, Sinniah UR, Ghasemzadeh A. Nanoparticles: alternatives against drug-resistant pathogenic microbes. Molecules. 2016;21:836.

Article   CAS   PubMed Central   Google Scholar  

Lam P-L, Wong W-Y, Bian Z, Chui C-H, Gambari R. Recent advances in green nanoparticulate systems for drug delivery: efficient delivery and safety concern. Nanomedicine. 2017;12:357–85.

Haba Y, Kojima C, Harada A, Ura T, Horinaka H, Kono K. Preparation of poly (ethylene glycol)-modified poly (amido amine) dendrimers encapsulating gold nanoparticles and their heat-generating ability. Langmuir. 2007;23:5243–6.

Shi X, Sun K, Baker JR Jr. Spontaneous formation of functionalized dendrimer-stabilized gold nanoparticles. J Phys Chem C. 2008;112:8251–8.

Park S-H, Oh S-G, Mun J-Y, Han S-S. Loading of gold nanoparticles inside the DPPC bilayers of liposome and their effects on membrane fluidities. Coll Surf B. 2006;48:112–8.

de Villiers MM, Aramwit P, Kwon GS. Nanotechnology in drug delivery. New York: Springer; 2008.

Kabanov AV, Lemieux P, Vinogradov S, Alakhov V. Pluronic ® block copolymers: novel functional molecules for gene therapy. Adv Drug Deliv Rev. 2002;54:223–33.

Wang N, Feng Y. Elaborating the role of natural products-induced autophagy in cancer treatment: achievements and artifacts in the state of the art. BioMed Res Int. 2015;2015:934207.

PubMed   PubMed Central   Google Scholar  

Ouattara B, Simard RE, Holley RA. Piette GJ-P, Bégin A: Antibacterial activity of selected fatty acids and essential oils against six meat spoilage organisms. Int J Food Microbiol. 1997;37:155–62.

Sharma G, Raturi K, Dang S, Gupta S, Gabrani R. Combinatorial antimicrobial effect of curcumin with selected phytochemicals on Staphylococcus epidermidis . J Asian Nat Prod Res. 2014;16:535–41.

Abdelwahab SI, Sheikh BY, Taha MME, How CW, Abdullah R, Yagoub U, El-Sunousi R, Eid EE. Thymoquinone-loaded nanostructured lipid carriers: preparation, gastroprotection, in vitro toxicity, and pharmacokinetic properties after extravascular administration. Int J Nanomed. 2013;8:2163.

Krauel K, Pitaksuteepong T, Davies NM, Rades T. Entrapment of bioactive molecules in poly (alkylcyanoacrylate) nanoparticles. Am J Drug Deliv. 2004;2:251–9.

Tan Q, Liu W, Guo C, Zhai G. Preparation and evaluation of quercetin-loaded lecithin-chitosan nanoparticles for topical delivery. Int J Nanomed. 2011;6:1621.

Sanna V, Roggio AM, Siliani S, Piccinini M, Marceddu S, Mariani A, Sechi M. Development of novel cationic chitosan-and anionic alginate–coated poly ( d, l- lactide-co-glycolide) nanoparticles for controlled release and light protection of resveratrol. Int J Nanomed. 2012;7:5501.

Casettari L, Illum L. Chitosan in nasal delivery systems for therapeutic drugs. J Control Release. 2014;190:189–200.

Obeid MA, Al Qaraghuli MM, Alsaadi M, Alzahrani AR, Niwasabutra K, Ferro VA. Delivering natural products and biotherapeutics to improve drug efficacy. Ther Deliv. 2017;8:947–56.

Miele E, Spinelli GP, Miele E, Di Fabrizio E, Ferretti E, Tomao S, Gulino A. Nanoparticle-based delivery of small interfering RNA: challenges for cancer therapy. Int J Nanomed. 2012;7:3637.

McNamara K, Tofail SA. Nanosystems: the use of nanoalloys, metallic, bimetallic, and magnetic nanoparticles in biomedical applications. Phys Chem Chem Phys. 2015;17:27981–95.

Saadeh Y, Vyas D. Nanorobotic applications in medicine: current proposals and designs. Am J Robot Surg. 2014;1:4–11.

Article   PubMed   PubMed Central   Google Scholar  

Oliveira ON Jr, Iost RM, Siqueira JR Jr, Crespilho FN, Caseli L. Nanomaterials for diagnosis: challenges and applications in smart devices based on molecular recognition. ACS Appl Mater Interfaces. 2014;6:14745–66.

De Jong WH, Borm PJ. Drug delivery and nanoparticles: applications and hazards. Int J Nanomed. 2008;3:133.

Article   Google Scholar  

Holzinger M, Le Goff A, Cosnier S. Nanomaterials for biosensing applications: a review. Front Chem. 2014;2:63.

Golovin YI, Gribanovsky SL, Golovin DY, Klyachko NL, Majouga AG, Master AM, Sokolsky M, Kabanov AV. Towards nanomedicines of the future: remote magneto-mechanical actuation of nanomedicines by alternating magnetic fields. J Control Release. 2015;219:43–60.

Lu H, Wang J, Wang T, Zhong J, Bao Y, Hao H. Recent progress on nanostructures for drug delivery applications. J Nanomater. 2016;2016:20.

Blanco E, Shen H, Ferrari M. Principles of nanoparticle design for overcoming biological barriers to drug delivery. Nat Biotechnol. 2015;33:941.

Kumari A, Kumar V, Yadav S. Nanotechnology: a tool to enhance therapeutic values of natural plant products. Trends Med Res. 2012;7:34–42.

Chen F, Ehlerding EB, Cai W. Theranostic nanoparticles. J Nucl Med. 2014;55:1919–22.

Swierczewska M, Han H, Kim K, Park J, Lee S. Polysaccharide-based nanoparticles for theranostic nanomedicine. Adv Drug Deliv Rev. 2016;99:70–84.

Chen K, Chen X. Design and development of molecular imaging probes. Curr Top Med Chem. 2010;10:1227–36.

Yhee JY, Son S, Kim SH, Park K, Choi K, Kwon IC. Self-assembled glycol chitosan nanoparticles for disease-specific theranostics. J Control Release. 2014;193:202–13.

Lee C-M, Jang D, Kim J, Cheong S-J, Kim E-M, Jeong M-H, Kim S-H, Kim DW, Lim ST, Sohn M-H, et al. Oleyl-Chitosan nanoparticles based on a dual probe for Optical/MR imaging in vivo. Bioconjug Chem. 2011;22:186–92.

Yang S-J, Lin F-H, Tsai H-M, Lin C-F, Chin H-C, Wong J-M, Shieh M-J. Alginate-folic acid-modified chitosan nanoparticles for photodynamic detection of intestinal neoplasms. Biomaterials. 2011;32:2174–82.

Ryu JH, Na JH, Ko HK, You DG, Park S, Jun E, Yeom HJ, Seo DH, Park JH, Jeong SY. Non-invasive optical imaging of cathepsin B with activatable fluorogenic nanoprobes in various metastatic models. Biomaterials. 2014;35:2302–11.

Lapčík L, Lapcik L, De Smedt S, Demeester J, Chabrecek P. Hyaluronan: preparation, structure, properties, and applications. Chem Rev. 1998;98:2663–84.

Article   PubMed   Google Scholar  

Kim H, Kim Y, Kim I-H, Kim K, Choi Y. ROS-responsive activatable photosensitizing agent for imaging and photodynamic therapy of activated macrophages. Theranostics. 2014;4:1.

Choi KY, Chung H, Min KH, Yoon HY, Kim K, Park JH, Kwon IC, Jeong SY. Self-assembled hyaluronic acid nanoparticles for active tumor targeting. Biomaterials. 2010;31:106–14.

Kamat M, El-Boubbou K, Zhu DC, Lansdell T, Lu X, Li W, Huang X. Hyaluronic acid immobilized magnetic nanoparticles for active targeting and imaging of macrophages. Bioconjug Chem. 2010;21:2128–35.

Arpicco S, Lerda C, Dalla Pozza E, Costanzo C, Tsapis N, Stella B, Donadelli M, Dando I, Fattal E, Cattel L. Hyaluronic acid-coated liposomes for active targeting of gemcitabine. Eur J Pharm Biopharm. 2013;85:373–80.

Wang G, Gao S, Tian R, Miller-Kleinhenz J, Qin Z, Liu T, Li L, Zhang F, Ma Q, Zhu L. Theranostic hyaluronic acid-iron micellar nanoparticles for magnetic-field-enhanced in vivo cancer chemotherapy. ChemMedChem. 2018;13:78–86.

Choi KY, Jeon EJ, Yoon HY, Lee BS, Na JH, Min KH, Kim SY, Myung S-J, Lee S, Chen X. Theranostic nanoparticles based on PEGylated hyaluronic acid for the diagnosis, therapy and monitoring of colon cancer. Biomaterials. 2012;33:6186–93.

Gombotz WR, Wee S. Protein release from alginate matrices. Adv Drug Deliv Rev. 1998;31:267–85.

Lee KY, Mooney DJ. Alginate: properties and biomedical applications. Prog Polym Sci. 2012;37:106–26.

Baghbani F, Moztarzadeh F, Mohandesi JA, Yazdian F, Mokhtari-Dizaji M. Novel alginate-stabilized doxorubicin-loaded nanodroplets for ultrasounic theranosis of breast cancer. Int J Biol Macromol. 2016;93:512–9.

Podgórna K, Szczepanowicz K, Piotrowski M, Gajdošová M, Štěpánek F, Warszyński P. Gadolinium alginate nanogels for theranostic applications. Coll Surf B. 2017;153:183–9.

Moscovici M. Present and future medical applications of microbial exopolysaccharides. Front Microbiol. 1012;2015:6.

Ding Z, Liu P, Hu D, Sheng Z, Yi H, Gao G, Wu Y, Zhang P, Ling S, Cai L. Redox-responsive dextran based theranostic nanoparticles for near-infrared/magnetic resonance imaging and magnetically targeted photodynamic therapy. Biomater Sci. 2017;5:762–71.

Hong S-P, Kang SH, Kim DK, Kang BS. Paramagnetic nanoparticle-based targeting theranostic agent for c6 rat glioma cell. J Nanomater. 2016; 2016:7617894. https://doi.org/10.1155/2016/7617894 .

Mignani S, El Kazzouli S, Bousmina M, Majoral JP. Expand classical drug administration ways by emerging routes using dendrimer drug delivery systems: a concise overview. Adv Drug Deliv Rev. 2013;65:1316–30.

Lounnas V, Ritschel T, Kelder J, McGuire R, Bywater RP, Foloppe N. Current progress in structure-based rational drug design marks a new mindset in drug discovery. Comput Struc Biotechnol J. 2013;5:e201302011.

Mavromoustakos T, Durdagi S, Koukoulitsa C, Simcic M, Papadopoulos M, Hodoscek M, Golic Grdadolnik S. Strategies in the rational drug design. Curr Med Chem. 2011;18:2517–30.

Wong PT, Choi SK. Mechanisms of drug release in nanotherapeutic delivery systems. Chem Rev. 2015;115:3388–432.

Prachayasittikul V, Worachartcheewan A, Shoombuatong W, Songtawee N, Simeon S, Prachayasittikul V, Nantasenamat C. Computer-aided drug design of bioactive natural products. Curr Top Med Chem. 2015;15:1780–800.

Chen G, Roy I, Yang C, Prasad PN. Nanochemistry and nanomedicine for nanoparticle-based diagnostics and therapy. Chem Rev. 2016;116:2826–85.

Pelaz B, Alexiou C, Alvarez-Puebla RA, Alves F, Andrews AM, Ashraf S, Balogh LP, Ballerini L, Bestetti A, Brendel C, Bosi S. Diverse applications of nanomedicine. Acs Nano. 2017;11:2313–81.

Mattos BD, Rojas OJ, Magalhaes WLE. Biogenic silica nanoparticles loaded with neem bark extract as green, slow-release biocide. J Clean Prod. 2017;142:4206–13.

Kinnear C, Moore TL, Rodriguez-Lorenzo L, Rothen-Rutishauser B, Petri-Fink A. Form follows function: nanoparticle shape and its implications for nanomedicine. Chem Rev. 2017;117:11476–521.

Sethi M, Sukumar R, Karve S, Werner ME, Wang EC, Moore DT, Kowalczyk SR, Zhang L, Wang AZ. Effect of drug release kinetics on nanoparticle therapeutic efficacy and toxicity. Nanoscale. 2014;6:2321–7.

Mattos BD, Tardy BL, Magalhaes WLE, Rojas OJ. Controlled release for crop and wood protection: recent progress toward sustainable and safe nanostructured biocidal systems. J Control Release. 2017;262:139–50.

Siepmann F, Herrmann S, Winter G, Siepmann J. A novel mathematical model quantifying drug release from lipid implants. J Control Release. 2008;128:233–40.

Ding CZ, Li ZB. A review of drug release mechanisms from nanocarrier systems. Mater Sci Eng. 2017;76:1440–53.

Lee JH, Yeo Y. Controlled drug release from pharmaceutical nanocarriers. Chem Eng Sci. 2015;125:75–84.

Kamaly N, Yameen B, Wu J, Farokhzad OC. Degradable controlled-release polymers and polymeric nanoparticles: mechanisms of controlling drug release. Chem Rev. 2016;116:2602–63.

Torchilin VP. Multifunctional nanocarriers. Adv Drug Deliv Rev. 2012;64:302–15.

Pelaz B, del Pino P, Maffre P, Hartmann R, Gallego M, Rivera-Fernandez S, de la Fuente JM, Nienhaus GU, Parak WJ. Surface functionalization of nanoparticles with polyethylene glycol: effects on protein adsorption and cellular uptake. Acs Nano. 2015;9:6996–7008.

Almalik A, Benabdelkamel H, Masood A, Alanazi IO, Alradwan I, Majrashi MA, Alfadda AA, Alghamdi WM, Alrabiah H, Tirelli N, Alhasan AH. Hyaluronic acid coated chitosan nanoparticles reduced the immunogenicity of the formed protein corona. Sci Rep. 2017;7:10542.

Martens TF, Remaut K, Deschout H, Engbersen JFJ, Hennink WE, van Steenbergen MJ, Demeester J, De Smedt SC, Braeckmans K. Coating nanocarriers with hyaluronic acid facilitates intravitreal drug delivery for retinal gene therapy. J Control Release. 2015;202:83–92.

Kolhar P, Anselmo AC, Gupta V, Pant K, Prabhakarpandian B, Ruoslahti E, Mitragotri S. Using shape effects to target antibody-coated nanoparticles to lung and brain endothelium. Proc Natl Acad Sci USA. 2013;110:10753–8.

Gao WW, Zhang LF. Coating nanoparticles with cell membranes for targeted drug delivery. J Drug Target. 2015;23:619–26.

Muller J, Bauer KN, Prozeller D, Simon J, Mailander V, Wurm FR, Winzen S, Landfester K. Coating nanoparticles with tunable surfactants facilitates control over the protein corona. Biomaterials. 2017;115:1–8.

Gao H, Yang Z, Zhang S, Cao S, Shen S, Pang Z, Jiang X. Ligand modified nanoparticles increases cell uptake, alters endocytosis and elevates glioma distribution and internalization. Sci Rep. 2013;3:2534.

Jain A, Jain SK. Ligand-appended BBB-targeted nanocarriers (LABTNs). Crit Rev Ther Drug Carrier Syst. 2015;32:149–80.

Shen HX, Shi SJ, Zhang ZR, Gong T, Sun X. Coating solid lipid nanoparticles with hyaluronic acid enhances antitumor activity against melanoma stem-like cells. Theranostics. 2015;5:755–71.

Gao X, Zhang J, Xu Q, Huang Z, Wang YY, Shen Q. Hyaluronic acid-coated cationic nanostructured lipid carriers for oral vincristine sulfate delivery. Drug Dev Ind Pharm. 2017;43:661–7.

Wang T, Hou JH, Su C, Zhao L, Shi YJ. Hyaluronic acid-coated chitosan nanoparticles induce ROS-mediated tumor cell apoptosis and enhance antitumor efficiency by targeted drug delivery via CD44. J Nanobiotechnol. 2017;15:7.

Muro S. Challenges in design and characterization of ligand-targeted drug delivery systems. J Control Release. 2012;164:125–37.

Kou L, Sun J, Zhai Y, He Z. The endocytosis and intracellular fate of nanomedicines: implication for rational design. Asian J Pharm Sci. 2013;8:1–10.

Li Z, Zhang Y, Zhu D, Li S, Yu X, Zhao Y, Ouyang X, Xie Z, Li L. Transporting carriers for intracellular targeting delivery via non-endocytic uptake pathways. Drug delivery. 2017;24:45–55.

Salatin S, Yari Khosroushahi A. Overviews on the cellular uptake mechanism of polysaccharide colloidal nanoparticles. J Cell Mol Med. 2017;21:1668–86.

Anirudhan TS, Nair AS. Temperature and ultrasound sensitive gatekeepers for the controlled release of chemotherapeutic drugs from mesoporous silica nanoparticles. J Mater Chem B. 2018;6:428–39.

Al-Ahmady Z, Kostarelos K. Chemical components for the design of temperature-responsive vesicles as cancer therapeutics. Chem Rev. 2016;116:3883–918.

Bai Y, Xie FY, Tian W. Controlled self-assembly of thermo-responsive amphiphilic h-shaped polymer for adjustable drug release. Chin J Polym Sci. 2018;36:406–16.

Zhang Z, Zhang D, Wei L, Wang X, Xu YL, Li HW, Ma M, Chen B, Xiao LH. Temperature responsive fluorescent polymer nanoparticles (TRFNPs) for cellular imaging and controlled releasing of drug to living cells. Coll Surf B. 2017;159:905–12.

Guo Y, Zhang Y, Ma J, Li Q, Li Y, Zhou X, Zhao D, Song H, Chen Q, Zhu X. Light/magnetic hyperthermia triggered drug released from multi-functional thermo-sensitive magnetoliposomes for precise cancer synergetic theranostics. J Control Release. 2017;272:145–58.

Hervault A, Thanh NT. Magnetic nanoparticle-based therapeutic agents for thermo-chemotherapy treatment of cancer. Nanoscale. 2014;6:11553–73.

Mathiyazhakan M, Wiraja C, Xu CJ: A Concise Review of Gold Nanoparticles-Based Photo-Responsive Liposomes for Controlled Drug Delivery. Nano - Micro Letters 2018, 10.

Xu L, Qiu LZ, Sheng Y, Sun YX, Deng LH, Li XQ, Bradley M, Zhang R. Biodegradable pH-responsive hydrogels for controlled dual-drug release. J Mater Chem B. 2018;6:510–7.

Ma GL, Lin WF, Yuan ZF, Wu J, Qian HF, Xua LB, Chen SF. Development of ionic strength/pH/enzyme triple-responsive zwitterionic hydrogel of the mixed l -glutamic acid and l -lysine polypeptide for site-specific drug delivery. J Mater Chem B. 2017;5:935–43.

Grillo R, Gallo J, Stroppa DG, Carbo-Argibay E, Lima R, Fraceto LF, Banobre-Lopez M. Sub-micrometer magnetic nanocomposites: insights into the effect of magnetic nanoparticles interactions on the optimization of SAR and MRI performance. Acs Appl Mater Interfaces. 2016;8:25777–87.

Alonso J, Khurshid H, Devkota J, Nemati Z, Khadka NK, Srikanth H, Pan JJ, Phan MH. Superparamagnetic nanoparticles encapsulated in lipid vesicles for advanced magnetic hyperthermia and biodetection. J Appl Phys. 2016;119:083904.

Ulbrich K, Hola K, Subr V, Bakandritsos A, Tucek J, Zboril R. Targeted drug delivery with polymers and magnetic nanoparticles: covalent and noncovalent approaches, release control, and clinical studies. Chem Rev. 2016;116:5338–431.

Chen CW, Syu WJ, Huang TC, Lee YC, Hsiao JK, Huang KY, Yu HP, Liao MY, Lai PS. Encapsulation of Au/Fe 3 O 4 nanoparticles into a polymer nanoarchitecture with combined near infrared-triggered chemo-photothermal therapy based on intracellular secondary protein understanding. J Mater Chem B. 2017;5:5774–82.

Portero A, Remunan-Lopez C, Criado M, Alonso M. Reacetylated chitosan microspheres for controlled delivery of anti-microbial agents to the gastric mucosa. J Microencapsul. 2002;19:797–809.

Artursson P, Lindmark T, Davis SS, Illum L. Effect of chitosan on the permeability of monolayers of intestinal epithelial cells (Caco-2). Pharm Res. 1994;11:1358–61.

Fernández-Urrusuno R, Calvo P, Remuñán-López C, Vila-Jato JL, Alonso MJ. Enhancement of nasal absorption of insulin using chitosan nanoparticles. Pharm Res. 1999;16:1576–81.

De Campos AM, Sánchez A, Alonso MJ. Chitosan nanoparticles: a new vehicle for the improvement of the delivery of drugs to the ocular surface. Application to cyclosporin A. Int J Pharm. 2001;224:159–68.

Al-Qadi S, Grenha A, Carrión-Recio D, Seijo B, Remuñán-López C. Microencapsulated chitosan nanoparticles for pulmonary protein delivery: in vivo evaluation of insulin-loaded formulations. J Control Release. 2012;157:383–90.

Silva MM, Calado R, Marto J, Bettencourt A, Almeida AJ, Gonçalves L. Chitosan Nanoparticles as a mucoadhesive drug delivery system for ocular administration. Mar Drugs. 2017;15:370.

Article   PubMed Central   Google Scholar  

Pistone S, Goycoolea FM, Young A, Smistad G, Hiorth M. Formulation of polysaccharide-based nanoparticles for local administration into the oral cavity. Eur J Pharm Sci. 2017;96:381–9.

Liu S, Yang S, Ho PC. Intranasal administration of carbamazepine-loaded carboxymethyl chitosan nanoparticles for drug delivery to the brain. Asian J Pharm Sci. 2018;13:72–81.

Jain A, Jain SK. Optimization of chitosan nanoparticles for colon tumors using experimental design methodology. Artif Cells Nanomed Biotechnol. 2016;44:1917–26.

Sosnik A. Alginate particles as platform for drug delivery by the oral route: state-of-the-art. ISRN Pharm. 2014;2014:926157.

Patil NH, Devarajan PV. Insulin-loaded alginic acid nanoparticles for sublingual delivery. Drug Deliv. 2016;23:429–36.

Haque S, Md S, Sahni JK, Ali J, Baboota S. Development and evaluation of brain targeted intranasal alginate nanoparticles for treatment of depression. J Psychiatr Res. 2014;48:1–12.

Román JV, Galán MA, del Valle EMM. Preparation and preliminary evaluation of alginate crosslinked microcapsules as potential drug delivery system (DDS) for human lung cancer therapy. Biomed Phys Eng Expr. 2016;2:035015.

Garrait G, Beyssac E, Subirade M. Development of a novel drug delivery system: chitosan nanoparticles entrapped in alginate microparticles. J Microencapsul. 2014;31:363–72.

Costa J, Silva N, Sarmento B, Pintado M. Potential chitosan-coated alginate nanoparticles for ocular delivery of daptomycin. Eur J Clin Microbiol Infect Dis. 2015;34:1255–62.

Goswami S, Naik S. Natural gums and its pharmaceutical application. J Sci Innovative Res. 2014;3:112–21.

Laffleur F, Michalek M. Modified xanthan gum for buccal delivery—a promising approach in treating sialorrhea. Int J Biol Macromol. 2017;102:1250–6.

Huang J, Deng Y, Ren J, Chen G, Wang G, Wang F, Wu X. Novel in situ forming hydrogel based on xanthan and chitosan re-gelifying in liquids for local drug delivery. Carbohydr Polym. 2018;186:54–63.

Menzel C, Jelkmann M, Laffleur F, Bernkop-Schnürch A. Nasal drug delivery: design of a novel mucoadhesive and in situ gelling polymer. Int J Pharm. 2017;517:196–202.

Sun B, Zhang M, Shen J, He Z, Fatehi P, Ni Y. Applications of cellulose-based materials in sustained drug delivery systems. Curr Med Chem. 2017. https://doi.org/10.2174/0929867324666170705143308 .

Elseoud WSA, Hassan ML, Sabaa MW, Basha M, Hassan EA, Fadel SM. Chitosan nanoparticles/cellulose nanocrystals nanocomposites as a carrier system for the controlled release of repaglinide. Int J Biol Macromol. 2018;111:604–13.

Agarwal T, Narayana SGH, Pal K, Pramanik K, Giri S, Banerjee I. Calcium alginate-carboxymethyl cellulose beads for colon-targeted drug delivery. Int J Biol Macromol. 2015;75:409–17.

Hansen K, Kim G, Desai KG, Patel H, Olsen KF, Curtis-Fisk J, Tocce E, Jordan S, Schwendeman SP. Feasibility investigation of cellulose polymers for mucoadhesive nasal drug delivery applications. Mol Pharm. 2015;12:2732–41.

Bozzuto G, Molinari A. Liposomes as nanomedical devices. Int J Nanomed. 2015;10:975.

Sercombe L, Veerati T, Moheimani F, Wu SY, Sood AK, Hua S. Advances and challenges of liposome assisted drug delivery. Front Pharm. 2015;6:286.

Kotla NG, Chandrasekar B, Rooney P, Sivaraman G, Larrañaga A, Krishna KV, Pandit A, Rochev Y. Biomimetic lipid-based nanosystems for enhanced dermal delivery of drugs and bioactive agents. ACS Biomater Sci Eng. 2017;3:1262–72.

Akbarzadeh A, Rezaei-Sadabady R, Davaran S, Joo SW, Zarghami N, Hanifehpour Y, Samiei M, Kouhi M, Nejati-Koshki K. Liposome: classification, preparation, and applications. Nanoscale Res Lett. 2013;8:102.

Mohan A, Narayanan S, Sethuraman S, Krishnan UM. Novel resveratrol and 5-fluorouracil coencapsulated in PEGylated nanoliposomes improve chemotherapeutic efficacy of combination against head and neck squamous cell carcinoma. BioMed res int. 2014;2014:424239.

Dimov N, Kastner E, Hussain M, Perrie Y, Szita N. Formation and purification of tailored liposomes for drug delivery using a module-based micro continuous-flow system. Sci Rep. 2017;7:12045.

Zylberberg C, Matosevic S. Pharmaceutical liposomal drug delivery: a review of new delivery systems and a look at the regulatory landscape. Drug Deliv. 2016;23:3319–29.

Sapsford KE, Algar WR, Berti L, Gemmill KB, Casey BJ, Oh E, Stewart MH, Medintz IL. Functionalizing nanoparticles with biological molecules: developing chemistries that facilitate nanotechnology. Chem Rev. 2013;113:1904–2074.

Zhang L, Gu F, Chan J, Wang A, Langer R, Farokhzad O. Nanoparticles in medicine: therapeutic applications and developments. Clin Pharmacol Ther. 2008;83:761–9.

Miyata K, Christie RJ, Kataoka K. Polymeric micelles for nano-scale drug delivery. React Funct Polym. 2011;71:227–34.

Xu W, Ling P, Zhang T. Polymeric micelles, a promising drug delivery system to enhance bioavailability of poorly water-soluble drugs. J Drug Deliv. 2013;2013:340315.

Kulthe SS, Choudhari YM, Inamdar NN, Mourya V. Polymeric micelles: authoritative aspects for drug delivery. Design Monomers Polym. 2012;15:465–521.

Devarajan PV, Jain S. Targeted drug delivery: concepts and design. Berlin: Springer; 2016.

Mourya V, Inamdar N, Nawale R, Kulthe S. Polymeric micelles: general considerations and their applications. Ind J Pharm Educ Res. 2011;45:128–38.

Wakaskar RR. Polymeric micelles for drug delivery. Int J Drug Dev Res. 2017;9:1–2.

Mandal A, Bisht R, Rupenthal ID, Mitra AK. Polymeric micelles for ocular drug delivery: from structural frameworks to recent preclinical studies. J Control Release. 2017;248:96–116.

Li Q, Lai KL, Chan PS, Leung SC, Li HY, Fang Y, To KK, Choi CHJ, Gao QY, Lee TW. Micellar delivery of dasatinib for the inhibition of pathologic cellular processes of the retinal pigment epithelium. Coll Surf B. 2016;140:278–86.

Kesharwani P, Xie L, Banerjee S, Mao G, Padhye S, Sarkar FH, Iyer AK. Hyaluronic acid-conjugated polyamidoamine dendrimers for targeted delivery of 3, 4-difluorobenzylidene curcumin to CD44 overexpressing pancreatic cancer cells. Coll Surf B. 2015;136:413–23.

Zhu J, Shi X. Dendrimer-based nanodevices for targeted drug delivery applications. J Mater Chem B. 2013;1:4199–211.

Madaan K, Kumar S, Poonia N, Lather V, Pandita D. Dendrimers in drug delivery and targeting: drug-dendrimer interactions and toxicity issues. J Pharm Bioallied Sci. 2014;6:139.

Cheng Y, Xu Z, Ma M, Xu T. Dendrimers as drug carriers: applications in different routes of drug administration. J Pharm Sci. 2008;97:123–43.

Noriega-Luna B, Godínez LA, Rodríguez FJ, Rodríguez A, Larrea G, Sosa-Ferreyra C, Mercado-Curiel R, Manríquez J, Bustos E. Applications of dendrimers in drug delivery agents, diagnosis, therapy, and detection. J Nanomater. 2014;2014:39.

Tripathy S, Das M. Dendrimers and their applications as novel drug delivery carriers. J Appl Pharm Sci. 2013;3:142–9.

Kesharwani P, Jain K, Jain NK. Dendrimer as nanocarrier for drug delivery. Progr Polym Sci. 2014;39:268–307.

Jain K, Gupta U, Jain NK. Dendronized nanoconjugates of lysine and folate for treatment of cancer. Eur J Pharm Biopharm. 2014;87:500–9.

Kaur A, Jain K, Mehra NK, Jain N. Development and characterization of surface engineered PPI dendrimers for targeted drug delivery. Artif Cells Nanomed Biotechnol. 2017;45:414–25.

Choi S-J, Lee JK, Jeong J, Choy J-H. Toxicity evaluation of inorganic nanoparticles: considerations and challenges. Mol Cell Toxicol. 2013;9:205–10.

Kong F-Y, Zhang J-W, Li R-F, Wang Z-X, Wang W-J, Wang W. Unique roles of gold nanoparticles in drug delivery, targeting and imaging applications. Molecules. 2017;22:1445.

Prusty K, Swain SK. Nano silver decorated polyacrylamide/dextran nanohydrogels hybrid composites for drug delivery applications. Mater Sci Eng. 2018;85:130–41.

Marcu A, Pop S, Dumitrache F, Mocanu M, Niculite C, Gherghiceanu M, Lungu C, Fleaca C, Ianchis R, Barbut A. Magnetic iron oxide nanoparticles as drug delivery system in breast cancer. Appl Surf Sci. 2013;281:60–5.

Junyaprasert VB, Morakul B. Nanocrystals for enhancement of oral bioavailability of poorly water-soluble drugs. Asian J Pharm Sci. 2015;10:13–23.

Du J, Li X, Zhao H, Zhou Y, Wang L, Tian S, Wang Y. Nanosuspensions of poorly water-soluble drugs prepared by bottom-up technologies. Int J Pharm. 2015;495:738–49.

Ni R, Zhao J, Liu Q, Liang Z, Muenster U, Mao S. Nanocrystals embedded in chitosan-based respirable swellable microparticles as dry powder for sustained pulmonary drug delivery. Eur J Pharm Sci. 2017;99:137–46.

McNamara K, Tofail SA. Nanoparticles in biomedical applications. Adv Phys. 2017;2:54–88.

Kudr J, Haddad Y, Richtera L, Heger Z, Cernak M, Adam V, Zitka O. Magnetic nanoparticles: from design and synthesis to real world applications. Nanomaterials. 2017;7:243.

Article   PubMed Central   CAS   Google Scholar  

Prasad PN. Nanophotonics. New York: Wiley; 2004.

Book   Google Scholar  

Volkov Y. Quantum dots in nanomedicine: recent trends, advances and unresolved issues. Biochem Biophys Res Commun. 2015;468:419–27.

Liu J, Lau SK, Varma VA, Moffitt RA, Caldwell M, Liu T, Young AN, Petros JA, Osunkoya AO, Krogstad T. Molecular mapping of tumor heterogeneity on clinical tissue specimens with multiplexed quantum dots. ACS Nano. 2010;4:2755–65.

Xu G, Zeng S, Zhang B, Swihart MT, Yong K-T, Prasad PN. New generation cadmium-free quantum dots for biophotonics and nanomedicine. Chem Rev. 2016;116:12234–327.

Shi Y, Pramanik A, Tchounwou C, Pedraza F, Crouch RA, Chavva SR, Vangara A, Sinha SS, Jones S, Sardar D. Multifunctional biocompatible graphene oxide quantum dots decorated magnetic nanoplatform for efficient capture and two-photon imaging of rare tumor cells. ACS Appl Mater Interfaces. 2015;7:10935–43.

Han H-S, Niemeyer E, Huang Y, Kamoun WS, Martin JD, Bhaumik J, Chen Y, Roberge S, Cui J, Martin MR. Quantum dot/antibody conjugates for in vivo cytometric imaging in mice. Proc Natl Acad Sci. 2015;112:1350–5.

So M-K, Xu C, Loening AM, Gambhir SS, Rao J. Self-illuminating quantum dot conjugates for in vivo imaging. Nat Biotechnol. 2006;24:339.

Zheng F-F, Zhang P-H, Xi Y, Chen J-J, Li L-L, Zhu J-J. Aptamer/graphene quantum dots nanocomposite capped fluorescent mesoporous silica nanoparticles for intracellular drug delivery and real-time monitoring of drug release. Anal Chem. 2015;87:11739–45.

Huang C-L, Huang C-C, Mai F-D, Yen C-L, Tzing S-H, Hsieh H-T, Ling Y-C, Chang J-Y. Application of paramagnetic graphene quantum dots as a platform for simultaneous dual-modality bioimaging and tumor-targeted drug delivery. J Mater Chem B. 2015;3:651–64.

Olerile LD, Liu Y, Zhang B, Wang T, Mu S, Zhang J, Selotlegeng L, Zhang N. Near-infrared mediated quantum dots and paclitaxel co-loaded nanostructured lipid carriers for cancer theragnostic. Coll Surf B. 2017;150:121–30.

Cai X, Luo Y, Zhang W, Du D, Lin Y. pH-Sensitive ZnO quantum dots–doxorubicin nanoparticles for lung cancer targeted drug delivery. ACS Appl Mater Interfaces. 2016;8:22442–50.

Balaji AB, Pakalapati H, Khalid M, Walvekar R, Siddiqui H. Natural and synthetic biocompatible and biodegradable polymers. In: Shimpi NG (ed) Biodegradable and biocompatible polymer composites: processing, properties and applications. Woodhead Publishing series in composites science and engineering. Duxford: Woodhead Publishing; 2017. p. 3–32.

Bassas-Galia M, Follonier S, Pusnik M, Zinn M. Natural polymers: a source of inspiration. In: Bioresorbable polymers for biomedical applications. New York: Elsevier; 2017. p. 31–64.

Chapter   Google Scholar  

Lohcharoenkal W, Wang L, Chen YC, Rojanasakul Y. Protein nanoparticles as drug delivery carriers for cancer therapy. BioMed Res Int. 2014;2014:180549.

Liu Z, Jiao Y, Wang Y, Zhou C, Zhang Z. Polysaccharides-based nanoparticles as drug delivery systems. Adv Drug Deliv Rev. 2008;60:1650–62.

Poole-Warren L, Patton A. Introduction to biomedical polymers and biocompatibility. In: Biosynthetic polymers for medical applications. New York: Elsevier; 2016. p. 3–31.

Pertici G. Introduction to bioresorbable polymers for biomedical applications. In: Biosynthetic polymers for medical applications. New York: Elsevier; 2016. p. 3–29.

Cardoso MJ, Costa RR, Mano JF. Marine origin polysaccharides in drug delivery systems. Mar Drugs. 2016;14:34.

Yu Z, Yu M, Zhang Z, Hong G, Xiong Q. Bovine serum albumin nanoparticles as controlled release carrier for local drug delivery to the inner ear. Nanoscale Res Lett. 2014;9:343.

Robinson M, Zhang X. The world medicines situation. Traditional medicines: global situation, issues and challenges. Geneva: World Health Organization; 2011. p. 1–12.

Atanasov AG, Waltenberger B, Pferschy-Wenzig E-M, Linder T, Wawrosch C, Uhrin P, Temml V, Wang L, Schwaiger S, Heiss EH. Discovery and resupply of pharmacologically active plant-derived natural products: a review. Biotechnol Adv. 2015;33:1582–614.

David B, Wolfender J-L, Dias DA. The pharmaceutical industry and natural products: historical status and new trends. Phytochem Rev. 2015;14:299–315.

Namdari M, Eatemadi A, Soleimaninejad M, Hammed AT. A brief review on the application of nanoparticle enclosed herbal medicine for the treatment of infective endocarditis. Biomed Pharm. 2017;87:321–31.

Heinrich M. Ethnopharmacology in the 21st century-grand challenges. Front Pharm. 2010;1:8.

Kinghorn AD, Pan L, Fletcher JN, Chai H. The relevance of higher plants in lead compound discovery programs. J Nat Prod. 2011;74:1539–55.

Yuan H, Ma Q, Ye L, Piao G. The traditional medicine and modern medicine from natural products. Molecules. 2016;21:559.

Patra JK, Das G, Baek K-H. Towards a greener environment: synthesis and applications of green nanoparticles. Pak J Agric Sci. 2016;53:59–79.

Duncan R, Gaspar R. Nanomedicine (s) under the microscope. Mol Pharm. 2011;8:2101–41.

Ramana KV, Singhal SS, Reddy AB. Therapeutic potential of natural pharmacological agents in the treatment of human diseases. BioMed Res Int. 2014;2014:573452.

Guo W. Green technology for nanoparticles in biomedical applications. In: Rai M, Posten C, editors. Green biosynthesis of nanoparticles: mechanisms and applications. Wallington: CABI; 2013.

Wicki A, Witzigmann D, Balasubramanian V, Huwyler J. Nanomedicine in cancer therapy: challenges, opportunities, and clinical applications. J Control Release. 2015;200:138–57.

Brigger I, Dubernet C, Couvreur P. Nanoparticles in cancer therapy and diagnosis. Adv Drug Deliv Rev. 2002;54:631–51.

Yohan D, Chithrani BD. Applications of nanoparticles in nanomedicine. J Biomed Nanotechnol. 2014;10:2371–92.

Ambesh P, Campia U, Obiagwu C, Bansal R, Shetty V, Hollander G, Shani J. Nanomedicine in coronary artery disease. Indian Heart J. 2017;69:244–51.

Grazu V, Moros M, Sánchez-Espinel C. Nanocarriers as nanomedicines: design concepts and recent advances. In: Frontiers of nanoscience. Vol. 4, New York: Elsevier; 2012. p. 337–440.

Rizzo LY, Theek B, Storm G, Kiessling F, Lammers T. Recent progress in nanomedicine: therapeutic, diagnostic and theranostic applications. Curr Opin Biotechnol. 2013;24:1159–66.

Devasena T. Diagnostic and therapeutic nanomaterials. In: Therapeutic and diagnostic nanomaterials. New York: Springer; 2017. p. 1–13.

Ventola CL. Progress in nanomedicine: approved and investigational nanodrugs. Pharm Ther. 2017;42:742.

Havel H, Finch G, Strode P, Wolfgang M, Zale S, Bobe I, Youssoufian H, Peterson M, Liu M. Nanomedicines: from bench to bedside and beyond. AAPS J. 2016;18:1373–8.

Kumar A, Chen F, Mozhi A, Zhang X, Zhao Y, Xue X, Hao Y, Zhang X, Wang PC, Liang X-J. Innovative pharmaceutical development based on unique properties of nanoscale delivery formulation. Nanoscale. 2013;5:8307–25.

Boroumand Moghaddam A, Namvar F, Moniri M, Md Tahir P, Azizi S, Mohamad R. Nanoparticles biosynthesized by fungi and yeast: a review of their preparation, properties, and medical applications. Molecules. 2015;20:16540–65.

Metz KM, Sanders SE, Pender JP, Dix MR, Hinds DT, Quinn SJ, Ward AD, Duffy P, Cullen RJ, Colavita PE. Green synthesis of metal nanoparticles via natural extracts: the biogenic nanoparticle corona and its effects on reactivity. ACS Sustain Chem Eng. 2015;3:1610–7.

Paul D, Sinha SN. Extracellular synthesis of silver nanoparticles using Pseudomonas aeruginosa KUPSB12 and its antibacterial activity. JJBS. 2014;7:245–50.

Kushwaha A, Singh VK, Bhartariya J, Singh P, Yasmeen K. Isolation and identification of E. coli bacteria for the synthesis of silver nanoparticles: characterization of the particles and study of antibacterial activity. Eur J Exp Biol. 2015;5:65–70.

Iravani S. Bacteria in nanoparticle synthesis: current status and future prospects. Int Sch Res Notices. 2014;2014:359316.

Mittal AK, Chisti Y, Banerjee UC. Synthesis of metallic nanoparticles using plant extracts. Biotechnol Adv. 2013;31:346–56.

Khan HA, Sakharkar MK, Nayak A, Kishore U, Khan A. 14-nanoparticles for biomedical applications: an overview. In: Narayan R, editor. Nanobiomaterials. Cambridge: Woodhead Publishing; 2018. p. 357–84.

Aravamudhan A, Ramos DM, Nada AA, Kumbar SG. Natural polymers: polysaccharides and their derivatives for biomedical applications. In: Natural and synthetic biomedical polymers. New York: Elsevier; 2014. p. 67–89.

Franci G, Falanga A, Galdiero S, Palomba L, Rai M, Morelli G, Galdiero M. Silver nanoparticles as potential antibacterial agents. Molecules. 2015;20:8856–74.

Pajardi G, Rapisarda V, Somalvico F, Scotti A, Russo GL, Ciancio F, Sgrò A, Nebuloni M, Allevi R, Torre ML. Skin substitutes based on allogenic fibroblasts or keratinocytes for chronic wounds not responding to conventional therapy: a retrospective observational study. Int Wound J. 2016;13:44–52.

Rahimi G, Alizadeh F, Khodavandi A. Mycosynthesis of silver nanoparticles from Candida albicans and its antibacterial activity against Escherichia coli and Staphylococcus aureus . Trop J Pharm Res. 2016;15:371–5.

Ali M, Kim B, Belfield KD, Norman D, Brennan M, Ali GS. Inhibition of Phytophthora parasitica and P. capsici by silver nanoparticles synthesized using aqueous extract of Artemisia absinthium . Phytopathology. 2015;105:1183–90.

Malapermal V, Botha I, Krishna SBN, Mbatha JN. Enhancing antidiabetic and antimicrobial performance of Ocimum basilicum , and Ocimum sanctum (L.) using silver nanoparticles. Saudi J Biol Sci. 2017;24:1294–305.

Sankar R, Karthik A, Prabu A, Karthik S, Shivashangari KS, Ravikumar V. Origanum vulgare mediated biosynthesis of silver nanoparticles for its antibacterial and anticancer activity. Coll Surf B. 2013;108:80–4.

Patra JK, Ali MS, Oh I-G, Baek K-H. Proteasome inhibitory, antioxidant, and synergistic antibacterial and anticandidal activity of green biosynthesized magnetic Fe3O4 nanoparticles using the aqueous extract of corn ( Zea mays L.) ear leaves. Artif Cells Nanomed Biotechnol. 2017;45:349–56.

Patra JK, Baek K-H. Antibacterial activity and synergistic antibacterial potential of biosynthesized silver nanoparticles against foodborne pathogenic bacteria along with its anticandidal and antioxidant effects. Front Microbiol. 2017;8:167.

Patra JK, Kwon Y, Baek K-H. Green biosynthesis of gold nanoparticles by onion peel extract: synthesis, characterization and biological activities. Adv Powder Technol. 2016;27:2204–13.

Patra JK, Baek K-H. Biosynthesis of silver nanoparticles using aqueous extract of silky hairs of corn and investigation of its antibacterial and anticandidal synergistic activity and antioxidant potential. IET Nanobiotechnol. 2016;10:326–33.

Patra JK, Baek K-H. Comparative study of proteasome inhibitory, synergistic antibacterial, synergistic anticandidal, and antioxidant activities of gold nanoparticles biosynthesized using fruit waste materials. Int J Nanomed. 2016;11:4691.

Patra JK, Baek K-H. Green synthesis of silver chloride nanoparticles using Prunus persica L. outer peel extract and investigation of antibacterial, anticandidal, antioxidant potential. Green Chem Lett Rev. 2016;9:132–42.

Patra JK, Das G, Baek K-H. Phyto-mediated biosynthesis of silver nanoparticles using the rind extract of watermelon ( Citrullus lanatus ) under photo-catalyzed condition and investigation of its antibacterial, anticandidal and antioxidant efficacy. J Photochem Photobiol B. 2016;161:200–10.

Wilczewska AZ, Niemirowicz K, Markiewicz KH, Car H. Nanoparticles as drug delivery systems. Pharmacol Rep. 2012;64:1020–37.

Zhu Z, Li Y, Yang X, Pan W, Pan H. The reversion of anti-cancer drug antagonism of tamoxifen and docetaxel by the hyaluronic acid-decorated polymeric nanoparticles. Pharmacol Res. 2017;126:84–96.

Dias DA, Urban S, Roessner U. A historical overview of natural products in drug discovery. Metabolites. 2012;2:303–36.

Gupta U, Sharma S, Khan I, Gothwal A, Sharma AK, Singh Y, Chourasia MK, Kumar V. Enhanced apoptotic and anticancer potential of paclitaxel loaded biodegradable nanoparticles based on chitosan. Int J Biol Macromol. 2017;98:810–9.

Chang C-H, Huang W-Y, Lai C-H, Hsu Y-M, Yao Y-H, Chen T-Y, Wu J-Y, Peng S-F, Lin Y-H. Development of novel nanoparticles shelled with heparin for berberine delivery to treat Helicobacter pylori. Acta Biomaterialia. 2011;7:593–603.

Aldawsari HM, Hosny KM. Solid lipid nanoparticles of Vancomycin loaded with Ellagic acid as a tool for overcoming nephrotoxic side effects: preparation, characterization, and nephrotoxicity evaluation. J Drug Deliv Sci Technol. 2018;45:76–80.

Dian L, Yu E, Chen X, Wen X, Zhang Z, Qin L, Wang Q, Li G, Wu C. Enhancing oral bioavailability of quercetin using novel soluplus polymeric micelles. Nanoscale Res Lett. 2014;9:684.

Spillmann CM, Naciri J, Algar WR, Medintz IL, Delehanty JB. Multifunctional liquid crystal nanoparticles for intracellular fluorescent imaging and drug delivery. ACS Nano. 2014;8:6986–97.

Purama RK, Goswami P, Khan AT, Goyal A. Structural analysis and properties of dextran produced by Leuconostoc mesenteroides NRRL B-640. Carbohydr Polym. 2009;76:30–5.

Agarwal A, Gupta U, Asthana A, Jain NK. Dextran conjugated dendritic nanoconstructs as potential vectors for anti-cancer agent. Biomaterials. 2009;30:3588–96.

Barenholz YC. Doxil ® —the first FDA-approved nano-drug: lessons learned. J Control Release. 2012;160:117–34.

Maeng JH, Lee D-H, Jung KH, Bae Y-H, Park I-S, Jeong S, Jeon Y-S, Shim C-K, Kim W, Kim J. Multifunctional doxorubicin loaded superparamagnetic iron oxide nanoparticles for chemotherapy and magnetic resonance imaging in liver cancer. Biomaterials. 2010;31:4995–5006.

Bonechi C, Martini S, Ciani L, Lamponi S, Rebmann H, Rossi C, Ristori S. Using liposomes as carriers for polyphenolic compounds: the case of trans-resveratrol. PLoS ONE. 2012;7:e41438.

Noorafshan A, Ashkani-Esfahani S. A review of therapeutic effects of curcumin. Curr Pharm Des. 2013;19:2032–46.

Wei X, Senanayake TH, Bohling A, Vinogradov SV. Targeted nanogel conjugate for improved stability and cellular permeability of curcumin: synthesis, pharmacokinetics, and tumor growth inhibition. Mol Pharm. 2014;11:3112–22.

Feng T, Wei Y, Lee RJ, Zhao L. Liposomal curcumin and its application in cancer. Int J Nanomed. 2017;12:6027.

Cheng C, Peng S, Li Z, Zou L, Liu W, Liu C. Improved bioavailability of curcumin in liposomes prepared using a pH-driven, organic solvent-free, easily scalable process. RSC Adv. 2017;7:25978–86.

Bilia AR, Guccione C, Isacchi B, Righeschi C, Firenzuoli F, Bergonzi MC. Essential oils loaded in nanosystems: a developing strategy for a successful therapeutic approach. Evid Based Complement Alternat Med. 2014;2014:651593.

Sainz V, Conniot J, Matos AI, Peres C, Zupanǒiǒ E, Moura L, Silva LC, Florindo HF, Gaspar RS. Regulatory aspects on nanomedicines. Biochem Biophys Res Commun. 2015;468:504–10.

Hassan S, Prakash G, Ozturk AB, Saghazadeh S, Sohail MF, Seo J, Dokmeci MR, Zhang YS, Khademhosseini A. Evolution and clinical translation of drug delivery nanomaterials. Nano Today. 2017;15:91–106.

Agrahari V, Agrahari V. Facilitating the translation of nanomedicines to a clinical product: challenges and opportunities. Drug Discov Today. 2018;23(5):974–91.

Caster JM, Patel AN, Zhang T, Wang A. Investigational nanomedicines in 2016: a review of nanotherapeutics currently undergoing clinical trials. Wiley Interdiscip Rev. 2016;2017:9.

Wacker MG, Proykova A, Santos GML. Dealing with nanosafety around the globe—regulation vs. innovation. Int J Pharm. 2016;509:95–106.

Lin P-C, Lin S, Wang PC, Sridhar R. Techniques for physicochemical characterization of nanomaterials. Biotechnol Adv. 2014;32:711–26.

Grossman JH, Crist RM, Clogston JD. Early development challenges for drug products containing nanomaterials. AAPS J. 2017;19:92–102.

Tinkle S, McNeil SE, Mühlebach S, Bawa R, Borchard G, Barenholz YC, Tamarkin L, Desai N. Nanomedicines: addressing the scientific and regulatory gap. Ann NY Acad Sci. 2014;1313:35–56.

Pandit A, Zeugolis DI. Twenty-five years of nano-bio-materials: have we revolutionized healthcare? Fut Med. 2016;11(9):985–7.

Bobo D, Robinson KJ, Islam J, Thurecht KJ, Corrie SR. Nanoparticle-based medicines: a review of FDA-approved materials and clinical trials to date. Pharm Res. 2016;33:2373–87.

Tran S, DeGiovanni P-J, Piel B, Rai P. Cancer nanomedicine: a review of recent success in drug delivery. Clin Transl Med. 2017;6:44.

Anselmo AC, Mitragotri S. Nanoparticles in the clinic. Bioeng Transl Med. 2016;1:10–29.

Grumezescu AM. Nanoscale fabrication, optimization, scale-up and biological aspects of pharmaceutical nanotechnology. New York: William Andrew; 2017.

Caster JM, Patel AN, Zhang T, Wang A. Investigational nanomedicines in 2016: a review of nanotherapeutics currently undergoing clinical trials. Wiley Interdiscip Rev. 2017;9:e1416.

Drug approvals and databases. https://www.fda.gov/Drugs/InformationOnDrugs/default.htm . Accessed 16 Aug 2018.

D’Mello SR, Cruz CN, Chen M-L, Kapoor M, Lee SL, Tyner KM. The evolving landscape of drug products containing nanomaterials in the United States. Nat Nanotechnol. 2017;12:523.

Download references

Authors’ contributions

JKP, GD, LFF, EVRC, MDPRT, LSAT, LADT, RG, MKS, SS and SH wrote different sections of the manuscript. JKP, LFF, MDPRT, RG, SS, SH and HSS edited the manuscript. All authors read and approved the final manuscript.

Acknowledgements

Jayanta Kumar Patra and Gitishree Das are grateful to Dongguk University, Republic of Korea for support. Leonardo Fernandes Fraceto and Estefânia V.R. Campos are grateful for the financial support provided by the São Paulo State Research Foundation (FAPESP) and National Council for Scientific and Technological Development (CNPQ). Maria del Pilar Rodriguez-Torres wishes to thank particularly DGAPA UNAM for the postdoctoral scholarship granted. Maria del Pilar Rodriguez-Torres, Laura Susana Acosta-Torres and Luis Armando Diaz-Torres thank Red de Farmoquimicos-CONACYT and DGAPA-UNAM PAPIIT-IN225516 project for support. Renato Grillo would like to thanks the São Paulo State Science Foundation (FAPESP, Grants #2015/26189-8). Han-Seung Shin thank Korea Environmental Industry & Technology Institute (A117-00197-0703-0) and Korea Institute of Planning and Evaluation for Technology in Food, Agriculture, Forestry and Fisheries (IPET) through Agricultural-BioTechnology Development Program funded by Ministry of Agriculture, Food and Rural Affairs (MAFRA)(710 003-07-7- SB120, 116075-3) for funding.

Competing interests

The authors declare that they have no competing interests.

Availability of data and materials

Not applicable.

Consent for publication

We have included two figures (Figs.  3 and 6 ) and one table (Table  1 ) from previously published literature with required copyright permission from the copyright holder. We have mentioned this in the manuscript with proper citation.

Ethics approval and consent to participate

This work was supported by the Korea Environmental Industry & Technology Institute (A117-00197-0703-0) and Korea Institute of Planning and Evaluation for Technology in Food, Agriculture, Forestry and Fisheries (IPET) through Agricultural-BioTechnology Development Program funded by Ministry of Agriculture, Food and Rural Affairs (MAFRA)(710 003-07-7- SB120, 116075-3).

Publisher’s Note

Springer Nature remains neutral with regard to jurisdictional claims in published maps and institutional affiliations.

Author information

Authors and affiliations.

Research Institute of Biotechnology & Medical Converged Science, Dongguk University-Seoul, Goyang-si, 10326, Republic of Korea

Jayanta Kumar Patra & Gitishree Das

Sao Paulo State University (UNESP), Institute of Science and Technology, Sorocaba, São Paulo, Zip Code 18087-180, Brazil

Leonardo Fernandes Fraceto & Estefania Vangelie Ramos Campos

Department of Biochemistry and Tissue Biology, Institute of Biology, State University of Campinas, Campinas, São Paulo, Zip code 13083-862, Brazil

Laboratorio de Investigación Interdisciplinaria, Área de Nanoestructuras y Biomateriales, Escuela Nacional de Estudios Superiores, Unidad Leon, Universidad Nacional Autonóma de México (UNAM), Boulevard UNAM No 2011. Predio El Saucillo y El Potrero, 37684, León, Guanajuato, Mexico

Maria del Pilar Rodriguez-Torres & Laura Susana Acosta-Torres

Centro de Investigaciones en Óptica, A.P. 1-948, C.P. 37000, León, Guanajuato, Mexico

Luis Armando Diaz-Torres

Department of Physics and Chemistry, School of Engineering, São Paulo State University (UNESP), Ilha Solteira, SP, 15385-000, Brazil

Renato Grillo

Department of Crop Science, Faculty of Agriculture, Universiti Putra Malaysia, 43400, Serdang, Selangor, Malaysia

Mallappa Kumara Swamy

Department of Biotechnology, Motilal Nehru National Institute of Technology Allahabad, Allahabad, Uttar Pradesh, 211004, India

Shivesh Sharma

Pharmacognosy Research Laboratories & Herbal Analysis Services UK, University of Greenwich, Medway Campus-Science, Grenville Building (G102/G107), Central Avenue, Chatham-Maritime, Kent, ME4 4TB, UK

Solomon Habtemariam

Department of Food Science and Biotechnology, Dongguk University, Ilsandong-gu, Goyang, Gyeonggi-do, 10326, Republic of Korea

Han-Seung Shin

You can also search for this author in PubMed   Google Scholar

Corresponding author

Correspondence to Han-Seung Shin .

Rights and permissions

Open Access This article is distributed under the terms of the Creative Commons Attribution 4.0 International License ( http://creativecommons.org/licenses/by/4.0/ ), which permits unrestricted use, distribution, and reproduction in any medium, provided you give appropriate credit to the original author(s) and the source, provide a link to the Creative Commons license, and indicate if changes were made. The Creative Commons Public Domain Dedication waiver ( http://creativecommons.org/publicdomain/zero/1.0/ ) applies to the data made available in this article, unless otherwise stated.

Reprints and permissions

About this article

Cite this article.

Patra, J.K., Das, G., Fraceto, L.F. et al. Nano based drug delivery systems: recent developments and future prospects. J Nanobiotechnol 16 , 71 (2018). https://doi.org/10.1186/s12951-018-0392-8

Download citation

Received : 19 July 2018

Accepted : 25 August 2018

Published : 19 September 2018

DOI : https://doi.org/10.1186/s12951-018-0392-8

Share this article

Anyone you share the following link with will be able to read this content:

Sorry, a shareable link is not currently available for this article.

Provided by the Springer Nature SharedIt content-sharing initiative

• Nanomedicine
• Nanomaterials
• Drug delivery
• Drug targeting
• Natural products

Journal of Nanobiotechnology

ISSN: 1477-3155

• Submission enquiries: [email protected]

Thank you for visiting nature.com. You are using a browser version with limited support for CSS. To obtain the best experience, we recommend you use a more up to date browser (or turn off compatibility mode in Internet Explorer). In the meantime, to ensure continued support, we are displaying the site without styles and JavaScript.

• View all journals
• Explore content
• About the journal
• Publish with us
• Sign up for alerts
• Review Article
• Published: 04 December 2020

Engineering precision nanoparticles for drug delivery

• Michael J. Mitchell   ORCID: orcid.org/0000-0002-3628-2244 1 , 2 , 3 , 4 , 5 ,
• Margaret M. Billingsley 1 ,
• Rebecca M. Haley   ORCID: orcid.org/0000-0001-7322-7829 1 ,
• Marissa E. Wechsler 6 ,
• Nicholas A. Peppas 6 , 7 , 8 , 9 , 10 &
• Robert Langer   ORCID: orcid.org/0000-0003-4255-0492 11  

Nature Reviews Drug Discovery volume  20 ,  pages 101–124 ( 2021 ) Cite this article

285k Accesses

3124 Citations

202 Altmetric

Metrics details

• Biomedical engineering
• Biotechnology
• Drug delivery
• Nanoparticles

In recent years, the development of nanoparticles has expanded into a broad range of clinical applications. Nanoparticles have been developed to overcome the limitations of free therapeutics and navigate biological barriers — systemic, microenvironmental and cellular — that are heterogeneous across patient populations and diseases. Overcoming this patient heterogeneity has also been accomplished through precision therapeutics, in which personalized interventions have enhanced therapeutic efficacy. However, nanoparticle development continues to focus on optimizing delivery platforms with a one-size-fits-all solution. As lipid-based, polymeric and inorganic nanoparticles are engineered in increasingly specified ways, they can begin to be optimized for drug delivery in a more personalized manner, entering the era of precision medicine. In this Review, we discuss advanced nanoparticle designs utilized in both non-personalized and precision applications that could be applied to improve precision therapies. We focus on advances in nanoparticle design that overcome heterogeneous barriers to delivery, arguing that intelligent nanoparticle design can improve efficacy in general delivery applications while enabling tailored designs for precision applications, thereby ultimately improving patient outcome overall.

Similar content being viewed by others

A framework for designing delivery systems

Wilson Poon, Benjamin R. Kingston, … Warren C. W. Chan

The ancillary effects of nanoparticles and their implications for nanomedicine

Evan P. Stater, Ali Y. Sonay, … Jan Grimm

The dose threshold for nanoparticle tumour delivery

Ben Ouyang, Wilson Poon, … Warren C. W. Chan

Introduction

Engineered nanomaterials hold significant promise to improve disease diagnosis and treatment specificity. Nanotechnology could help overcome the limitations of conventional delivery — from large-scale issues such as biodistribution to smaller-scale barriers such as intracellular trafficking — through cell-specific targeting, molecular transport to specific organelles and other approaches. To facilitate the realization and clinical translation of these promising nano-enabled technologies, the US National Science and Technology Council (NSTC) launched the National Nanotechnology Initiative (NNI) in 2000 and outlined well-defined initiatives and grand challenges for the field 1 . These initiatives have supported the recent efforts to investigate and improve nanotechnology, of which nanoparticles (NPs) constitute a significant portion of reported research and advancement.

NPs have the potential to improve the stability and solubility of encapsulated cargos, promote transport across membranes and prolong circulation times to increase safety and efficacy 2 , 3 . For these reasons, NP research has been widespread, generating promising results in vitro and in small animal models 4 . However, despite this extensive research motivated by the NNI, the number of nanomedicines available to patients is drastically below projections for the field, partially because of a translational gap between animal and human studies 4 , 5 . This gap comes from a lack of understanding of the differences in physiology and pathology between animal model species and humans, specifically how these differences influence the behaviour and functionality of nanomedicines in the body 6 . The differences across species are not the only factor that limits clinical translation. Heterogeneity amongst patients can also limit the success of nanomedicines, and there is currently only limited research on the interactions between nanomedicines and in stratified patient populations. Thus, of the nanomedicines that are approved, few are recommended as first-line treatment options, and many show improvements in only a small subset of patients 7 . This is due, in part, to the underexplored heterogeneity both in the biological underpinnings of diseases and amongst patients, which alters NP efficacy because the growth, structure and physiology of diseased tissue alter NP distribution and functionality.

Many early NP iterations were unable to overcome these biological barriers to delivery, but more recent NP designs have utilized advancements in controlled synthesis strategies to incorporate complex architectures, bio-responsive moieties and targeting agents to enhance delivery 8 , 9 , 10 , 11 , 12 . These NPs can therefore be utilized as more complex systems — including in nanocarrier-mediated combination therapies — to alter multiple pathways, maximize the therapeutic efficacy against specific macromolecules, target particular phases of the cell cycle or overcome mechanisms of drug resistance.

This new focus on generating NPs to overcome biological barriers specific to patient subsets or disease states can be attributed, in part, to the increasing prevalence of precision, or personalized, medicine and the creation of the Precision Medicine Initiative (PMI) in 2015 (ref. 13 ). The goal of precision medicine is to utilize patient information — such as genetic profile, environmental exposures or comorbidities — to develop an individualized treatment plan. The use of precision minimizes the impact of patient heterogeneity and allows for more accurate patient stratification, improved drug specificity and optimized dosing or combinatorial strategies. However, precision therapies are subject to the same biological barriers to delivery as other medicines, which limits their clinical potential. As such, new NP designs, informed by patient data and engineered to overcome particular barriers in a stratified patient population, could greatly improve the delivery of and response to precision medicine therapies.

This Review focuses on advances in nanomedicine that could facilitate clinical translation of precision medicines and improve patient-specific therapeutic responses, with an emphasis on leveraging biomaterials and biomedical engineering innovations to overcome biological barriers and patient heterogeneity. The Review presents the progress made towards goals set forth by the NNI and the PMI to improve disease treatment for the individual. Although NPs have been used successfully in precision diagnostic applications, this Review focuses on the delivery of precision medicine therapeutics, as we believe that these medicines will greatly influence precision NPs in the future. Further, we discuss the biological barriers that have limited the widespread success of NP applications and critically review rational NP designs that have aimed to overcome these obstacles. The distribution and delivery trends from decades of NP research are also covered, as the impact of NP characteristics on therapeutic responses are explored. These emerging topics — as well as advances in engineering NPs for specific applications — are of particular importance as new opportunities arise for the clinical translation of NP-based precision therapies in cancer medicine, immunotherapy and in vivo gene editing (Fig.  1 ).

Overview highlighting some of the biological barriers that nanoparticles (NPs) can overcome (inner ring) and precision medicine applications that may benefit from NPs (outer ring). As explored in this Review, intelligent NP designs that improve delivery have the potential to enhance the performance of precision medicines and, thus, accelerate their clinical translation. CAR, chimeric antigen receptor; EGFR, epidermal growth factor receptor; EPR, enhanced permeation and retention; gRNA, guide RNA; RNP, ribonucleoprotein.

Lipid-based NPs

Lipid-based NPs include various subset structures but are most typically spherical platforms comprising at least one lipid bilayer surrounding at least one internal aqueous compartment (Fig.  2 ). As a delivery system, lipid-based NPs offer many advantages including formulation simplicity, self-assembly, biocompatibility, high bioavailability, ability to carry large payloads and a range of physicochemical properties that can be controlled to modulate their biological characteristics 14 , 15 . For these reasons, lipid-based NPs are the most common class of FDA-approved nanomedicines 7 , 16 (Table  1 ).

Each class of nanoparticle (NP) features multiple subclasses, with some of the most common highlighted here. Each class has numerous broad advantages and disadvantages regarding cargo, delivery and patient response.

For liposomes — one of the subsets of lipid-based NPs that has the most members — the NPs are typically composed of phospholipids, which can form unilamellar and multilamellar vesicular structures. This allows the liposome to carry and deliver hydrophilic, hydrophobic and lipophilic drugs, and they can even entrap hydrophilic and lipophilic compounds in the same system, thereby expanding their use 17 . Their in vitro and in vivo stability are altered by NP size, surface charge, lipid composition, number of lamellae and surface modifications (with ligands or polymers), which can be altered during synthesis 15 , 18 . Because they can be rapidly taken up by the reticuloendothelial system , liposomes often include surface modifications to extend their circulation and enhance delivery, which has enabled their clinical use 14 , 19 .

Another notable subset of lipid-based NPs is commonly referred to as lipid nanoparticles (LNPs) — liposome-like structures widely used for the delivery of nucleic acids. They differ from traditional liposomes primarily because they form micellar structures within the particle core, a morphology that can be altered based on formulation and synthesis parameters 20 . LNPs are typically composed of four major components: cationic or ionizable lipids that complex with negatively charged genetic material and aid endosomal escape, phospholipids for particle structure, cholesterol for stability and membrane fusion, and PEGylated lipids to improve stability and circulation 21 , 22 . The efficacy of their nucleic acid delivery along with their simple synthesis, small size and serum stability have made LNPs particularly important in personalized genetic therapy applications 12 , 23 . Ionizable LNPs are an ideal platform for the delivery of these nucleic acid therapies as they have a near-neutral charge at physiological pH but become charged in acidic endosomal compartments, promoting endosomal escape for intracellular delivery 24 , 25 . However, despite these advantages, LNP systems can still be limited by low drug loading and biodistribution that results in high uptake to the liver and spleen 16 .

Polymeric NPs

Polymeric NPs can be synthesized from natural or synthetic materials, as well as monomers or preformed polymers — allowing for a wide variety of possible structures and characteristics (Fig.  2 ). They can be formulated to enable precise control of multiple NP features and are generally good delivery vehicles because they are biocompatible and have simple formulation parameters. Polymeric NPs are synthesized using various techniques such as emulsification (solvent displacement or diffusion) 26 , nanoprecipitation 27 , 28 , ionic gelation 29 and microfluidics 30 , which all result in different final products. Polymeric NPs also have variable drug delivery capabilities. Therapeutics can be encapsulated within the NP core, entrapped in the polymer matrix, chemically conjugated to the polymer or bound to the NP surface. This enables delivery of various payloads including hydrophobic and hydrophilic compounds, as well as cargos with different molecular weights such as small molecules, biological macromolecules, proteins and vaccines 30 , 31 , 32 , 33 , 34 , 35 , making polymeric NPs ideal for co-delivery applications 36 . By modulating properties such as composition, stability, responsivity and surface charge, the loading efficacies and release kinetics of these therapeutics can be precisely controlled 37 , 38 .

The most common forms of polymeric NPs are nanocapsules (cavities surrounded by a polymeric membrane or shell) and nanospheres (solid matrix systems). Within these two large categories, NPs are divided further into shapes such as polymersomes, micelles and dendrimers. Polymersomes are artificial vesicles, with membranes made using amphiphilic block copolymers. They are comparable to liposomes, and are often locally responsive, but are reported to have improved stability and cargo-retention efficiency 39 , making them effective vehicles for the delivery of therapeutics to the cytosol 40 , 41 . Some polymers which are commonly copolymerized for these applications include poly(ethylene glycol) (PEG) and poly(dimethylsiloxane) (PDMS). Polymeric micelles, which are also typically responsive block copolymers, self-assemble to form nanospheres with a hydrophilic core and a hydrophobic coating: this serves to protect aqueous drug cargo and improve circulation times. Polymeric micelles can load various therapeutic types — from small molecules to proteins 35 — and have been used for the delivery of cancer therapeutics in clinical trials 42 .

Dendrimers are hyperbranched polymers with complex three-dimensional architectures for which the mass, size, shape and surface chemistry can be highly controlled. Active functional groups present on the exterior of dendrimers enable conjugation of biomolecules or contrast agents to the surface while drugs can be loaded in the interior. Dendrimers can hold many types of cargo, but are most commonly investigated for the delivery of nucleic acids and small molecules 43 , 44 . For these applications, charged polymers such as poly(ethylenimine) (PEI) and poly(amidoamine) (PAMAM) are commonly used. Several dendrimer-based products are currently in clinical trials as theranostic agents, transfection agents, topical gels and contrast agents 44 , 45 , 46 . Charged polymers can be used to form non-dendrimer NPs as well. Polyelectrolytes are one such example: these polymers have a repeating electrolyte group, giving them charge that varies with pH. Polyelectrolytes have been incorporated in numerous NP formulations to improve properties such as bioavailability 47 and mucosal transport 48 . They are also inherently responsive, and can be useful for intracellular delivery.

Overall, polymeric NPs are ideal candidates for drug delivery because they are biodegradable, water soluble, biocompatible, biomimetic and stable during storage. Their surfaces can be easily modified for additional targeting 49 — allowing them to deliver drugs, proteins and genetic material to targeted tissues, which makes them useful in cancer medicine, gene therapy and diagnostics. However, disadvantages of polymeric NPs include an increased risk of particle aggregation and toxicity. Only a small number of polymeric nanomedicines are currently FDA approved and used in the clinic (Table  1 ), but polymeric nanocarriers are currently undergoing testing in numerous clinical trials 7 .

Inorganic NPs

Inorganic materials such as gold, iron and silica have been used to synthesize nanostructured materials for various drug delivery and imaging applications (Fig.  2 ). These inorganic NPs are precisely formulated and can be engineered to have a wide variety of sizes, structures and geometries. Gold NPs (AuNPs), which are the most well studied, are used in various forms such as nanospheres, nanorods, nanostars, nanoshells and nanocages 50 . Additionally, inorganic NPs have unique physical, electrical, magnetic and optical properties, due to the properties of the base material itself. For example, AuNPs possess free electrons at their surface that continually oscillate at a frequency dependent on their size and shape, giving them photothermal properties 51 . AuNPs are also easily functionalized, granting them additional properties and delivery capabilities 50 .

Iron oxide is another commonly researched material for inorganic NP synthesis, and iron oxide NPs make up the majority of FDA-approved inorganic nanomedicines 52 (Table  1 ). Magnetic iron oxide NPs — composed of magnetite (Fe 3 O 4 ) or maghemite (Fe 2 O 3 ) — possess superparamagnetic properties at certain sizes and have shown success as contrast agents, drug delivery vehicles and thermal-based therapeutics 53 . Other common inorganic NPs include calcium phosphate and mesoporous silica NPs, which have both been used successfully for gene and drug delivery 54 , 55 . Quantum dots — typically made of semiconducting materials such as silicon — are unique NPs used primarily in in vitro imaging applications, but they show promise for in vivo diagnostics 56 , 57 .

Due to their magnetic, radioactive or plasmonic properties, inorganic NPs are uniquely qualified for applications such as diagnostics, imaging and photothermal therapies. Most have good biocompatibility and stability, and fill niche applications that require properties unattainable by organic materials. However, they are limited in their clinical application by low solubility and toxicity concerns, especially in formulations using heavy metals 53 , 58 .

NPs in precision medicine

Precision medicine pushes for the development of patient-specific treatments in a clinical setting, to overcome the many limitations of traditional one-size-fits-all approaches and improve therapeutic outcomes 59 . In oncology, patient stratification through biomarkers and companion diagnostics has become the norm for drug development, as most cancer nanomedicines fail to produce positive results in unstratified studies 60 . Even though patient stratification has been essential in the clinical development of several precision medicines for cancer, NP-based clinical trials are currently conducted in unstratified patient populations 61 . However, this will likely change in the near future, as the importance of stratification becomes more apparent and NPs begin to be developed with specific patient populations in mind. The progression of NPs through clinical trials may similarly be hastened by the incorporation of stratified patient populations, as these populations will likely respond more uniformly to treatment. Furthermore, NPs are particularly well placed to broaden the potential patient populations that qualify for precision medicine therapies by neutralizing factors, such as comorbidities or heterogeneous biological barriers, that may have made patients previously unqualified. As NPs overcome many of the current limitations to delivery — potentially improving the potency and therapeutic efficacy of precision medicines — they may allow more patients to qualify for clinical trials and benefit from individualized therapies.

Since the launch of the PMI in 2015, several applications have incorporated nanomaterials in precision medicine 59 . For example, a blood test for the early detection of pancreatic cancer analyses the personalized biomolecular corona that adsorbs onto graphene oxide nanoflakes 62 . The unique property of graphene oxide, which binds low amounts of albumin, allows strong adsorption of proteins that are present in the plasma at low levels 62 . Other studies use magnetic NPs 63 or AuNPs 64 , 65 , which are simple to use, in biomarker detection assays, thereby saving time and money if compared with existing methods that require substantial sample processing. In addition to diagnostic screening, some therapeutic applications of NPs aim to remodel the tumour microenvironment to promote particle accumulation and penetration, and thus increase drug efficacy, and/or to sensitize tumours to a particular therapy 66 , 67 , 68 . For example, tumour-associated endothelial cells can be manipulated by a NP-delivered microRNA, which alters the tumour vasculature and thereby sensitizes the tumour to traditional cancer therapies 67 . Similarly, bio-inspired lipoproteins have been used to remodel tumours, and can improve NP accessibility to cancer cells 27-fold 66 . The usage of photothermal NPs can improve the infiltration and activity of chimeric antigen receptor (CAR) T cells against solid tumours 69 . NPs can also be used to modulate immune activation or suppression to sensitize cancer cells to therapeutics, helping to homogenize these currently heterogeneous environments in an attempt to increase the number of patients who respond to or qualify for precision treatments 40 , 70 .

In summary, combining NPs and precision medicine has the potential to advance both fields. Because NPs are currently screened in unstratified patient populations, the introduction of NPs developed for specific patient populations could allow for the accelerated clinical translation of numerous nanomaterials. Conversely, the success of precision medicine relies on strictly stratified patient populations, and the use of NPs to improve delivery across heterogeneous biological barriers could increase the efficacy of precision medicines, allowing for more patients to be included in the stratified population, as well as increasing the likelihood of successful translation to the clinic. Advances in genome sequencing and biomarker detection allow for the appropriate selection of cargo for the treatment of patient-specific diseases. Although it is not the focus of this Review, there are several diagnostic applications that may be improved by NP technologies. Development of nanobiomaterials for precision medicine is a highly customizable process. This careful design approach enables adjustments of the therapeutics’ pharmacokinetics to match requirements for solubility, administration or biodistribution and has seen success in research settings (Table  2 ).

Biological barriers

Even under normal physiological conditions, effective biodistribution and drug delivery are difficult to achieve as NPs face both physical and biological barriers — including shear forces, protein adsorption and rapid clearance — that limit the fraction of administered NPs that reach the target therapeutic site 71 . These barriers are often altered in disease states and can be even more difficult to overcome with a generalized, one-size-fits-all approach 3 , 72 , 73 , 74 . Furthermore, these changes in biological barriers vary not just across diseases but also on a patient-to-patient basis, and they can occur at the systemic, microenvironmental and cellular levels, making them hard to isolate and characterize broadly. Understanding the biological barriers faced both generally and on a patient-specific level allows for the design of optimally engineered NP platforms. In this section, we discuss strategies used by NPs to overcome biological barriers on the systemic, local and cellular scale (Fig.  3 ).

Factors such as size, shape, charge and surface coating determine what happens to nanoparticles (NPs) in the circulation, including clearance, and how the NPs interact with local barriers such as the tumour microenvironment or mucus layers. A few general trends are highlighted here: spherical and larger NPs marginate more easily during circulation, whereas rod-shaped NPs extravasate more readily (top left); and uncoated or positively charged NPs are cleared more quickly by macrophages (top right). In terms of local distribution, in general, rod-shaped, neutral and targeted NPs penetrate tumours more readily (bottom left) whereas positively charged, smaller and coated NPs more easily traverse mucosal barriers (bottom right).

Systemic delivery and biodistribution

The biological barriers that NPs encounter depend on the route of administration as well as the patient’s disease type and progression 3 . Although local delivery methods may allow NPs to circumvent some of the obstacles faced by systemic delivery, they often involve more invasive procedures and complex techniques that present other limitations. Furthermore, local delivery may only be useful in diseases where the pathology is restricted to known, accessible sites — such as certain solid cancers or traumatic injuries — so systemic administration is more common in NP applications 75 . Thus, this section explores the most prominent barriers to delivery faced by systemic administration.

Circulation, stability and clearance

While in circulation, factors such as excretion, blood flow, coronas and phagocytic cells can reduce NP stability and delivery (Fig.  3 ). The specific effects of each of these environmental factors is dependent upon the physiochemical properties of the NP platform, which has led to general design principles aimed to manipulate these characteristics to achieve favourable outcomes. In size, for example, NPs with a diameter less than 10 nm have generally been shown to be rapidly eliminated by the kidneys, whereas NPs larger than 200 nm risk activating the complement system , if not otherwise engineered 76 . Furthermore, to avoid rapid excretion based on surface properties, many NP formulations incorporate PEG as a stealth coating. PEGylation improves the circulation time by altering the NP size and solubility while shielding the NP surface from enzymes and antibodies that may induce degradation, secretion and clearance, but this physical barrier does not completely prevent recognition by macrophages or other cells of the immune system. Additionally, exposure to PEG results in the production of anti-PEG antibodies that, when present in high concentrations, can induce the rapid clearance of PEGylated NPs 77 , 78 . Clinical studies have also shown that these anti-PEG antibodies can be present in humans who have been exposed to PEG through means other than PEGylated medicines, indicating that even the first dose of PEGylated NPs would not necessarily circulate for long in all patients 79 , 80 .

Another option for stealth is platelet membrane cloaking, which reduces cell uptake and complement activation 81 . Although this cloaking avoids the macrophage-based immune issues associated with PEGylation, the NPs may still be recognized by other cell populations 82 . However, these platelet-based cell interactions can also help with targeting: platelet membrane-cloaked NPs feature the ligands present on native platelets — including mediators of adhesion to von Willebrand factor and collagen — allowing the wrapped NPs to target injury sites and accumulate around activated platelets 83 .

Surface modifications and cloaking techniques allow NPs to avoid the recognition and clearance systems that may lead to rapid NP degradation and instability, and there are also numerous NP design strategies that specifically focus on improving stability. NP stability is greatly affected by how its composition material interacts with the environment, and lipid-based and polymer-based NPs are the most susceptible to instability and aggregation both in circulation and in storage. Thus, to improve the robustness of these softer NPs, excipients such as helper lipids, cholesterol and PEGylated lipids 18 , 21 can be formulated with lipid-based NPs to increase their stability, whereas polymer NPs may utilize cross-linking techniques 84 , 85 . For storage and transport, many NPs are lyophilized to improve stability, although this does not affect the NP stability once administered 21 . However, as NP designs aim to increase stability, the balance between stabilization and effective intracellular delivery — which typically requires carrier degradation — must be considered.

In the bloodstream, NPs experience varying flow rates that induce shear stress and may damage the platforms or their cargo and prevent extravasation 86 , 87 . These fluid forces can strip NPs of their surface coatings and can prevent NPs from localizing to vessel walls to extravasate — either transcellularly or paracellularly — to reach target tissues 3 , 86 , 87 , 88 , 89 , 90 . Larger (microscale) particles have a higher probability of localizing to the vessel walls, and non-spherical particles show better margination 89 . Specifically, ellipsoids, discoid shapes and nanorods with higher aspect ratios localize to blood vessels better than spheres do 89 , 91 , 92 . This is caused by flow-induced rolling in shapes with high ratios, which results in edge margination at a speed proportional to the NP aspect ratio 3 . Even after vessel localization, architecture-dependent drag force from blood flow may rip NPs from cell membranes if they lack sufficient binding affinity for endothelial cells 88 . Thus, the haemodynamics experienced by systemically administered NPs — which are often altered in vascular pathologies such as stenoses and hypertension 93 , 94 — greatly influence NP distribution and delivery.

In addition to their interactions with vessel walls, circulating NPs come into contact with biomolecules and cells suspended in blood. The non-specific adherence of serum proteins and lipids forms a corona on the surface of NPs 61 , 95 , 96 . The composition of the corona depends on the biomolecules present in blood as well as the physicochemical characteristics of the NP surface, as this dictates the adsorption or desorption of proteins from biological fluids 61 , 96 . At times, the engineered surface properties meant to enhance NP targeting — such as conjugated ligands or modified surface charge — may encourage corona formation through charge-specific interactions 97 . Once formed, this corona will dictate the distribution of the NP, and can compromise stability of both the NP and its cargo 61 , 96 , 98 . Recent investigations have sought to determine how the specific corona biomolecules alter NP distribution and tissue-specific targeting 99 , 100 , 101 . For example, coronas containing apolipoprotein E (ApoE) act as targeting moieties for low-density lipoprotein receptors, which leads to NP delivery to hepatocytes and, in some instances, across the blood–brain barrier (BBB) 99 , 102 , 103 , 104 . If the corona contains opsonin or ligands for pattern recognition receptors, it can cause rapid clearance of the NPs via contact with cells of the innate immune system 98 .

Clearance of NPs from the circulation can be influenced by their physicochemical properties, but often results from interactions with the mononuclear phagocytic system (MPS) or reticuloendothelial system 3 , 97 (Fig.  3 ). These systems feature phagocytes (predominantly macrophages), monocytes and dendritic cells, which take up NPs and accumulate in the spleen and liver 71 , 97 , 98 . This clearance tends to happen more rapidly in stiffer NPs 105 , 106 . In terms of surface charge, cationic NPs are generally most rapidly cleared, followed by anionic NPs, whereas neutral and slightly negative NPs have the longest half-lives in circulation 2 , 3 . To minimize clearance, some NP designs implement surface modifications — such as PEG, ‘self’ peptides (including CD47) or cell membrane coatings — that aim to reduce these interactions with phagocytic cells of the MPS 3 , 76 .

In addition to clearance, interactions with the MPS can cause toxicity, as these cells trigger immune responses involving the secretion of tumour necrosis factors, interleukins and interferons that cause inflammation or tissue damage 98 , 107 . The type and magnitude of immune response to NPs is greatly affected by NP size, shape and surface properties. For example, in a mouse ovalbumin model, spherical NPs cause a T helper 1 cell-biased (cell-mediated) response, micrometre-length rods cause a T helper 2 cell-biased (humoral) response and spherical NPs induce a stronger immune response overall 108 . Furthermore, uptake by phagocytic cells has been related to the NP curvature and aspect ratio: triangular and rod-shaped NPs show more uptake than star-shaped or spherical NPs, and rod-shaped NPs induce more inflammation in macrophages 109 , 110 , 111 . Certain surface properties induce inflammation; some PEGylated NPs have caused severe allergic reactions or anaphylaxis in a small subset of patients in clinical trials 112 , 113 . Although the steric effects of PEG on the surface of NPs typically prevent interactions with MPS cells, anti-PEG antibodies developed from previous PEG exposure undermine this stealth property and promote MPS interactions 77 , 78 . At high concentrations, anti-PEG antibodies most commonly cause rapid clearance, but they are also thought to contribute to these uncommon but severe allergic reactions. In all, these immune responses to NP architecture and surface modifications can induce inflammation and adverse reactions, which emphasizes the importance of tailoring NP design to minimize these risks 61 , 82 .

Barriers to biodistribution

Extravasation is the first step for a NP in circulation to reach the target tissue 89 , 90 . Extravasation can be altered by NP characteristics, including size: for example, small NPs generally cross capillary walls more easily than large NPs 3 , 71 , 98 (Fig.  3 ). Thus, NPs tend to distribute across organs in a size-dependent manner, with the highest accumulation often in the liver and spleen 3 , 76 . However, size-dependent distribution can be altered by pathological environments such as the tumour vasculature, in which larger than normal intercellular gaps allow for larger NPs to exit the vessels 71 . Overall, extravasation leads to non-specific distribution, which presents a translational challenge for applications that require specific localization 3 .

Optimizing the administration route can improve biodistribution. The means of administering any drug may alter its fate and efficacy in vivo, and numerous studies have explored how these routes impact the fate of NPs specifically 114 , 115 . For example, polymeric (poly(lactic-co-glycolic) acid (PLGA)) NPs that are intravenously injected accumulate primarily in the liver and spleen, whereas if these NPs are subcutaneously or intranodally injected, they are more likely to accumulate in local lymph nodes 116 . These alternate administration routes enable NPs to reach the lymphatic system prior to systemic circulation, which could be beneficial in certain immunotherapeutic applications 116 , 117 . Another method for bypassing extravasation that has been increasingly explored for NP delivery is pulmonary administration, specifically NP inhalation. This route avoids exposure to the systemic circulation prior to lung delivery, thus avoiding hepatic first-pass metabolism and increasing the delivery of dendrimer-based NPs to the lung and lymph node as compared with intravenous delivery 114 . However, despite their improved delivery to lung tissue, inhaled NPs face the unique obstacles of mucus and pulmonary surfactant, which act as physical barriers to lung delivery (discussed further in later sections) and can vary greatly across patients and pathologies 118 , 119 . Furthermore, a recent comparison of three widely used routes of pulmonary administration in mouse models — intratracheal instillation, intratracheal spraying and intranasal instillation — revealed different rates of polymeric (PLGA) NP deposition in the lungs and heterogeneous distributions overall, suggesting the need for validated and consistent delivery methods when assessing pulmonary administration routes for NPs 120 . Clinically, approved NP formulations (such as ONPATTRO, VYXEOS and NBTXR3) are either intratumourally or intravenously administered, with limited optimization of the administration route 7 . Although preclinical work is being done to explore alternate routes, these studies are still ongoing. In all, selecting the optimal administration route for NPs may allow for more desirable distribution, but many current administration routes still, ultimately, result in widespread distribution of NPs and fail to provide the level of targeting and specificity desired.

To further prevent non-specific distribution, many NP platforms have added targeting moieties to their surface to direct their delivery. Most targeting moieties — including antibodies 121 , glucose 122 , transferrin 34 , folate 123 , transporters 2 and integrin ligands 124 — use interactions with molecules on the target cell’s surface, such as ligand–receptor, enzyme–substrate or antibody–antigen-mediated interactions 73 . Thus, targeted NPs must be engineered with a targeting moiety density that allows for these cell surface interactions, making it important to understand the ratio of receptors to ligands and the number of interactions needed to overcome the initial energy barrier to NP uptake 2 , 76 . Active targeting may also improve NP distribution within a target tissue: binding peptides for collagen type III increased NP accumulation in joints and enabled preferential NP association with osteoarthritic cartilage over healthy tissue 26 . Additionally, use of the tumour-targeting peptide CREKA allowed for enhanced permeation and uniform distribution of NPs in a mouse model of breast cancer with solid tumours 125 . However, despite the benefits of active targeting, the process of target selection is limiting. Because disease markers can vary among diseased cells within a patient as well as among patients in a population, target selection is a personalized process 73 , 76 . Furthermore, although antibodies can be engineered with high specificity, their conjugation to NPs may increase MPS interactions and result in rapid NP clearance 76 . Although the selection of less specific targets, such as broadly expressed transporters, may reduce immunogenicity compared with antibodies, they face additional obstacles associated with off-target delivery 2 , which occurs if the marker or target is expressed on both diseased and healthy cells. Off-target delivery is further complicated if the diseased cells are widely distributed throughout normal tissues, which precludes local delivery 61 , 73 , 97 , 126 . Overall, although targeting NPs to disease markers aids in specific delivery, active targeting is not currently an ideal solution.

Physical barriers to NP distribution include tight junctions among the endothelial and epithelial cells of the BBB (in intravenous delivery) and the gastrointestinal tract (in oral delivery), respectively. For NPs to reach the central nervous system (CNS), they must utilize receptor-mediated endocytosis to be taken up by endothelial cells of the BBB and exocytosed to the other side 97 , 98 , 127 . Receptor-mediated transcytosis is an effective way to deliver therapeutics to the brain or to infiltrate tumour tissue 128 , 129 . However, this method of crossing the BBB is complicated by the heterogeneity of plasma membrane transporters on endothelial cells 2 . However, some transporters — such as glucose transporters — are consistently highly expressed on the BBB, and some common targets — such as vascular cell adhesion molecule 1 — can increase NP transport across the BBB 2 , 91 . These two molecules could be harnessed to deliver NPs. Other targeting routes have been explored, including the transferrin receptor, which has theoretical advantages over other transporter types but has yet to see clinical success 130 . With transferrin receptor systems, only approximately 5% of the systemically administered NP dose reaches the CNS and even less reaches target cells 97 , 127 . However, a recent investigation characterizing AuNPs that had crossed the BBB revealed that the composition of the NP corona was altered but stable after crossing: investigations to better understand these altered coronas could help develop future strategies for CNS targeting 96 . Overall, the BBB remains a major challenge for systemically administered NPs attempting to reach tissues of the CNS. Thus, intranasal administration has been increasingly explored as an option for NP delivery to the brain as it bypasses the BBB and avoids many of the limitations of systemic delivery 115 , 131 . However, factors such as a limited dosing volume and variables attributed to patient congestion and mucus have presented notable obstacles to the intranasal route 132 , 133 .

Although oral delivery is the most widely used and readily accepted form of drug administration, the gastrointestinal tract presents numerous barriers for NPs 72 . For NPs that rely on passive diffusion, crossing the endothelium is restricted by concentration gradients and P-glycoproteins that excrete drugs from the vasculature into the intestinal lumen. However, some NP properties may encourage transport across the gastrointestinal tract. In a recent screen of inorganic NPs for the oral delivery of protein drugs, smaller, negatively charged silica NPs enhanced intestinal permeation by opening tight junctions, thus avoiding the need for cellular uptake for transport across the epithelial barrier 134 . However, for platforms that rely on endocytosis and subsequent exocytosis to cross the gastrointestinal tract, size remains an important factor. For example, the large surface area of polymeric NPs (as compared with soluble drug) has been beneficial as it increases the number of interactions with the gastrointestinal tract following oral delivery 135 . Overall, the average optimal reported size for NP transcytosis in gastrointestinal applications seems to be around 100 nm 28 , 48 , 135 , 136 , 137 . This size range allows for both enterocytes and M cells — which preferentially take up NPs 20–100 nm and 100–500 nm in diameter, respectively — to transport NPs across the gastrointestinal tract 47 . Rod-shaped NPs generally outperform spherical particles, which aligns with trends showing that nanorods are internalized into epithelial cells more efficiently than spheres are 135 , 138 , 139 . However, even when NPs are internalized by intestinal epithelial cells, only a small percentage undergo exocytosis 140 . Thus, even when utilizing these NP design elements to optimize transport, passive diffusion across the gastrointestinal tract is limited, so active targeting methods have been explored.

The transferrin pathway can be exploited for trans-epithelial movement in the intestine, using a transferrin-coated NP 136 . This target may be especially useful in the treatment of colon cancer and irritable bowel disease, which both cause overexpression of the transferrin receptor in the intestinal mucosa 136 . However, in addition to the limitations of active targeting described above, targeting strategies in the gastrointestinal tract are frustrated by the formation of coronas in gastrointestinal fluids, which vary with diet, and goblet cells that produce mucus to coat the endothelial surface. Both of these issues limit the interactions between NPs and the intestinal walls 72 , 141 . These barriers are made heterogeneous by pathologies, such as inflammatory diseases, that may increase epithelial permeability and alter mucus production, pH and the gastrointestinal microbiome 72 , 142 . Thus, the challenges presented by the gastrointestinal tract, and heterogenized by patient pathologies, present substantial barriers to achieving therapeutically desired biodistribution via oral delivery.

Microenvironmental barriers

Once at the target site, NPs must navigate the local microenvironment. Here, obstacles may include changes in chemical conditions or physical barriers to penetration. Thus, to successfully engineer NPs that reach the desired tissues or cells, a fundamental understanding of the microenvironments they will encounter is critical.

Microenvironment variability

Microenvironments often feature conditions that are substantially different from those in the circulation, which can greatly alter the physical properties and stability of NPs. For example, the gastrointestinal tract includes areas of extreme pH variation and acidity 72 . These conditions, in addition to the presence of enzymes that induce degradation, make the gastrointestinal tract an unstable environment for many NPs 72 , 74 . Furthermore, the gastrointestinal microenvironments can be diversely altered by disease states, resulting in heterogeneous reactions to biomaterials 74 . For example, a comparison of microenvironments in colon cancer and colitis, which feature different amine surface group densities on colon tissue, determined that the pathologies resulted in disease-dependent compatibility with dendrimer/dextran biomaterials 74 .

Numerous diseased microenvironments feature variations in pH, such as the low pH observed in many tumours or the fluctuating pH observed across stages of wound healing 90 , 143 . Some pH-sensitive NP platforms (detailed below) have been developed that allow the release of the drug only in specific pH conditions. Wound sites are often hyperthermic, so temperature-responsive systems can react to this local environment and provide targeted delivery 144 . In the case of stenosis and atherosclerosis, narrowed arteries result in elevated shear stresses that can be exploited to increase therapeutic release from NPs that break down under these conditions 145 .

Local NP distribution

Barriers to local distribution have been explored in depth in the tumour microenvironment, as NP penetration and stability are challenging in solid tumours 107 , 146 . Many characteristics of the tumour microenvironment — including the vasculature, interstitial fluid pressure and extracellular matrix (ECM) density — contribute to the limited permeation and penetration of NPs 3 , 147 , 148 , 149 , 150 . Thus, the exact cause of successful NP accumulation in tumours has been highly debated, with only a few established trends correlating NP design to tumour delivery. Some of these NP properties that can promote accumulation in tumours (Fig.  3 ) include hydrodynamic diameters above 100 nm, rod-shaped architectures, near-neutral charges or inorganic material compositions — all of which may be optimal for tumour accumulation 71 .

The tumour microenvironment also plays a key role in determining NP fate. As the vasculature within tumours is heterogeneous and abnormal, NPs can accumulate in tumours as the leaky vessels enable NP extravasation, a phenomenon often referred to as the enhanced permeation and retention (EPR) effect 61 . Reports vary on the role of the EPR effect in NP accumulation in tumours. Up to 10–15% of injected NPs accumulate at the tumour site, as compared with 0.1% of free drug, and some studies attribute this to the EPR effect 61 . In contrast, recent work utilizing a combination of computational analysis and imaging techniques in a mouse tumour model has determined that only a fraction of NP accumulation in tumours can be attributed to passive transport, including the EPR effect. Instead, the work suggests that other mechanisms such as immune cell interactions, protein coronas and molecular mechanisms may contribute substantially to the enhanced tumour accumulation of NPs 129 . These conclusions seem to be supported by meta-analysis: one study reviewed 232 data sets and determined that, on average, only 0.7% of injected NP doses reach tumours — a finding that greatly de-emphasized the impact of the EPR effect 71 . However, it is important to note the limitations of these generalized findings, as a recent investigation has highlighted the potentially misleading results from quantifying NP distribution using non-standard calculations, which may have led to biased results 151 . Thus, while continuing to explore the broad implications of the EPR effect for NP accumulation, future investigations must critically evaluate the metrics used to quantify delivery and distribution.

The EPR effect relies on the heterogeneous formation of vasculature throughout tumours that can be altered by individual patient factors such as age, genetics, lifestyle and even previous antitumour treatments 61 , 98 , 146 . Thus, to select the appropriate delivery platform for a specific patient, their individual tumour and its vasculature should be assessed for EPR effects that alter NP accumulation and permeation 61 , 106 . This is a promising diagnostic application for tagged NPs, which have been used in preliminary studies to quantify the level of the EPR effect at the tumour site seen in individual patients in an attempt to identify patient populations that are well suited for NP-based therapies 152 .

The heterogeneity of tumour microenvironments generates many obstacles to successful NP delivery, including reduced permeation. Within the tumour environment, cells may overproduce or generate altered ECM components that result in a dense ECM that physically hinders NP delivery 147 , 148 , 149 . This is especially true for cationic NPs as they adhere to the negatively charged tumour ECM, reducing permeation 71 , 153 . In addition, abnormalities in tumour lymphatic vasculature can result in decreased interstitial fluid drainage, which increases intertumoural interstitial pressure and prevents effective NP perfusion 3 , 107 , 148 , 150 . These barriers can prevent most tumour cells from interacting with NPs; one study found that antibody-targeted NPs interacted with only 2% of tumour cells — a number far below the level required for therapeutic efficacy 148 .

Limited perfusion is a therapeutic obstacle for NPs delivered to the brain as well. After crossing the BBB via systemic delivery or local administration, NPs in the brain microenvironment often fail to permeate the tissue because of the limited extracellular space and non-specific adherence to the ECM 154 , 155 . Thus, advanced delivery methods such as convection-enhanced delivery, and NP surface modifications, such as dense PEG coatings, have been explored. These methods may aid in more widespread and evenly distributed delivery across the brain, as well as improved permeation in glioblastomas 154 , 155 , 156 .

NPs face additional barriers to local distribution, including biofilms and mucus 72 , 149 . Within mucus layers, the distances between adjacent polymer links determine the mesh pore size, which can vary from 10 to 1000 nm, so smaller objects diffuse through whereas larger objects are trapped 72 , 149 . In addition to filtering by size, mucus may trap objects via non-specific interactions that lead to their rapid clearance from epithelial surfaces 72 . Although mucus throughout the body shares a similar function, its behaviour varies depending on its physiological location because of differences in its composition, hydration and viscoelasticity 72 , 119 , 157 . For example, mucus in the gastrointestinal tract acts as an adherent, thick layer whereas mucus in the lungs tends to be thinner and more mobile, making it a heterogeneous barrier 72 , 119 , 157 (Fig.  3 ).

Although mucus behaviour can be generalized within each of these physiological environments, there are areas of disparity within the mucus of an organ system, and these barriers are dynamic. In the gastrointestinal tract, the thickness of the mucus barrier can range from 40 to 450 μm in the stomach and from 110 to 160 μm in the colon, and factors such as fibre intake affect both mucus thickness and the turnover rate 72 , 119 . Additionally, as the mucosal barrier transitions between the near-neutral endothelial cell surfaces and the acidic intestinal lumen, a steep pH gradient is present across its micrometre-scale thickness, creating a very unstable environment for NP platforms 72 , 74 . Changes to these properties of the mucus in the gastrointestinal tract are also observed in pathologies that change glycosylation patterns, pH and the mucus layer thickness 72 , 142 .

Similarly, pathologies of the lungs change mucus behaviour in that tissue. Mucus in the lungs — a barrier that greatly impacts inhaled NPs — is characterized by high concentrations of MUC5AC and MUC5B polymers 118 , 157 . However, in cystic fibrosis, increased MUC5B expression and excessive cross-linking of polymers in the mucus results in decreased pore size and low rates of mucus clearance because this mucus has higher viscosity, which encourages biofilm formation by entrapping pathogens and limiting the mobility of neutrophils 157 , 158 . MUC5B concentrations are also elevated in cases of primary ciliary dyskinesia and cigarette smoke-induced chronic bronchitis; MUC5AC is elevated in asthma 158 . In all, the properties of mucus have been found to vary greatly based on patient factors such as diet, lifestyle and disease, making it a complex environment for inhaled NP delivery.

Cellular and intracellular barriers

When NPs make contact with their target cells, there are still numerous barriers to the uptake and intracellular trafficking that determine their functional delivery 3 . This section explores the barriers NPs must overcome to achieve cellular uptake and proper internal trafficking and discuss how cellular heterogeneity affects these NP interactions.

NP uptake and internalization

The corona, in combination with the NP characteristics it alters, such as hydrophilicity and charge, alters cellular uptake in numerous cell types including macrophages and cancer cells 61 , 159 , 160 . This corona-covered NP interacts with the surface of the cell, which consists of a negatively charged, selectively permeable phospholipid bilayer with biomolecules incorporated throughout in a fluid mosaic structure 75 , 160 . Cell membranes vary widely and membrane components such as lipid rafts and transmembrane proteins are heterogeneously distributed; over 400 cell surface transporter types have been identified in human cells 2 , 75 , 160 . Furthermore, the exact stiffness of the cell membrane and its compositional fluidity are determined, in part, by the cytoskeleton, which can respond to external cues, making these characteristics dynamic 161 . Thus, NPs interacting with the same cell may experience different interactions depending on their location on the cell’s membrane or their time of contact. Anionic NPs may struggle to make contact with the cell surface due to repulsive forces, whereas cationic NPs, if too positively charged, may damage the cell membrane and even cause cytotoxicity 3 , 76 , 159 , 162 . Thus, the first contact between a NP and a cell — which varies with NP and cell properties — may determine the NP fate and, therefore, its therapeutic potential.

For the next step in delivery — cell uptake — few definitive trends have been established concerning the optimal NP shape and size; some models and studies indicate that, in non-phagocytic cells, spherical NPs have improved uptake over rod-shaped particles 163 , 164 , but other studies show the opposite effect 165 , 166 . Similarly, many in vitro studies have shown that non-phagocytic cells only take up NPs that are 10–60 nm in size, and that smaller NPs internalize better, whereas other investigations indicate that smaller NPs are more likely to cause cytotoxicity 2 , 76 , 167 . The process of NP uptake can be broken down into passive and active methods 75 . Because the cell membrane is selectively permeable, passive diffusion is predominantly limited to small, uncharged molecules that travel down concentration gradients 162 . Thus, NPs most commonly rely on active transport to cross the cell membrane 3 , 75 . Specifically, NPs tend to utilize endocytic pathways, in which the plasma membrane is folded into vesicles to engulf NPs on the cell surface, and then release them intracellularly 3 , 75 , 90 , 160 . The type of endocytosis a NP undergoes can affect its fate within the cell and is determined by numerous factors including cell type, NP size and receptor interactions 3 , 90 , 160 . For example, in non-specific cell membrane interactions, smaller or larger NPs will be taken up by either phagocytosis or pinocytosis, respectively 90 .

However, more specific interactions — often with negatively charged NPs — may result in caveolin-mediated or clathrin-mediated endocytosis 162 . Caveolin-mediated endocytosis can occur in molecules smaller than approximately 60 nm and utilizes lipid rafts to create specialized vesicles after engulfment 90 . This form of endocytosis is more common for nanorods; nanosphere uptake is usually clathrin-mediated 138 . Clathrin-mediated endocytosis — the most common route for NP uptake in non-specialized mammalian cells — relies on receptor-mediated, hydrophobic or electrostatic interactions between NPs and the cell membrane in areas of clathrin expression 90 , 160 . The induction of these endocytic pathways is influenced by NP properties such as stiffness and size. Although results vary, stiffer NPs are generally more easily taken up, and both experimental and theoretical analyses indicate that endocytosis of rigid particles requires less energy 162 , 168 . Additionally, NPs that are too small (<30 nm) may not be capable of driving membrane wrapping enough to activate endocytic processes 76 . Multiple studies report good cellular uptake and intracellular delivery when particles ~50 nm in diameter are used 27 , 76 , 169 , 170 , 171 , 172 . Thus, the process used for NP uptake is determined by numerous factors including characteristics of the cell membrane as well as properties of the NP, which also influence the subsequent endocytic process (Fig.  4 ).

a | Upon interaction with the cell surface, nanoparticles (NPs) — depending on their surface, size, shape and charge — are taken up by various types of endocytosis or pinocytosis via non-specific interactions, such as membrane wrapping, or specific interactions, such as with cell surface receptors. b | Once they have entered the cell, NPs remain trapped within vesicular compartments, or endosomes, that feature various characteristics such as internal or external receptors. To achieve functional delivery, most NPs must escape from these compartments before they acidify. Thus, responsive NPs — such as ionizable NPs that become charged in low-pH environments — aid in endosomal escape and allow for intracellular delivery whereas unresponsive NPs often remain trapped and are destroyed by lysosome acidity and proteolytic enzymes.

During endocytic processes, the vesicles, or endosomes, go through different stages that involve changes in their chemical composition and pH until they become lysosomes, which feature low pH, high ionic strength and proteolytic enzymes that affect the stability of NPs and their cargo 75 , 160 . Materials that change in response to acidic conditions and have a proton sponge effect have been investigated to aid in endosomal escape, enabling NPs to avoid degradation 40 , 169 , 173 . LNPs, which include cationic and ionizable materials, are good examples of these intracellularly triggered delivery mechanisms and are often used to carry nucleic acids into cells 3 , 174 , 175 , 176 , 177 , 178 , 179 . Materials can respond to the acidic endosomal pH, but NPs have also been designed to react to the reductive endosomal environment 180 , 181 . As the redox potential of the endosome increases, cleavable linkers incorporated into the NP design may allow the NP to degrade, disrupt the endosomal membrane and release its cargo intracellularly 180 , 181 . In addition to responsive NPs, complex shapes, such as nanostars, have also been shown to improve the intracellular delivery of genetic material as they can efficiently enter cells and escape endosomes 172 .

Once in the cytosol, the cargo may still need to reach certain intracellular environments 75 , 160 , 161 . Because cells are highly compartmentalized, reaching these organelles may require crossing additional intracellular membranes 161 . For example, the nuclear membrane is a barrier for genome editing or DNA delivery 75 , 149 . NPs targeting the mitochondria for specific cancers or as neurogenerative or cardiovascular therapies face similar barriers 75 , 182 ; to overcome this challenge, pH-responsive NP systems could aid in precise delivery to the mitochondrial environment 183 .

Cellular heterogeneity

In addition to the general cellular barriers described above, cells form heterogeneous populations both within a patient and across a patient population. Many cellular variations occur based on the characteristics of an individual. For example, in human fibroblast cells from fetal lungs and epithelial cells from fetal colons, younger cells took up more NPs than old cells, and younger cells were less susceptible to toxicity 184 . Additionally, a study found that cell sex altered the uptake of AuNPs in human amniotic stem cells and fibroblasts isolated from saliva, demonstrating yet another factor to consider in NP delivery 185 .

Drug-resistant cells contribute to the cellular heterogeneity that challenges NP delivery 186 . For example, resistance to platinum (II)-based drugs, such as oxaliplatin and cisplatin, which distort DNA structure to induce apoptosis, can occur if cancer cells overexpress efflux pumps or increase their rate of DNA repair. Thus, smart NP platforms must be engineered to overcome these barriers. For example, micelles deliver NPs more effectively to the nucleus, and thus the cell has fewer opportunities to acquire drug resistance 187 , 188 . Thus, both cell type and acquired phenotypes that lead to a heterogeneous cell population create diverse barriers to NP delivery, but new developments in NP design may help overcome these obstacles.

To account for the vast heterogeneity of biological barriers and disease states within and across patient populations, methods must be developed to deliver therapeutics in a manner that is highly modular and customizable. This section details the effects of various NP properties on delivery, with a focus on how individual NP design choices (such as architecture, material properties, targeting and responsiveness; Fig.  5 ) can overcome barriers specific to individual diseases and patients.

Surface and material properties, architecture, targeting moieties and responsiveness are all attributes of nanoparticles (NPs) that can be altered in intelligent designs to tailor the platform to a specific application. Different combinations of these four properties allows for seemingly endless permutations of NP features and platforms. PEG, poly(ethylene glycol).

NPs for cancer therapy

Cancer remains the second leading cause of death worldwide 189 . Cancer is heterogeneous, and the development of effective cancer therapies is very challenging partially because of this complexity. However, precision medicine has emerged as a promising approach, and targeted chemotherapeutics have been developed that can treat patients who express specific biomarkers. The first drug of this type, imatinib (Gleevec; Novartis), is given to patients with chronic myeloid leukaemia who express the BCR–ABL fusion protein from the Philadelphia chromosome 190 . FDA approval of imatinib opened the field for many other successful targeted chemotherapeutics 190 , 191 , 192 . However, these therapies and others could be more effective if delivery is improved. For example, imatinib has also been delivered using a NP system, which enhanced tumour accumulation and regression in vivo, improving the survival ratio to 40% after 60 days in a melanoma mouse model 193 . Improvements in delivery could overcome some limitations of therapeutics that have failed to make it to the clinic, including small-molecule drugs with limited water solubility or antibodies with low stability 194 . Similarly, many chemotherapeutics have off-target toxicity and induce adaptive resistance, which limit efficacy. Furthermore, there are many biological barriers associated with cancer, specifically at the tumour site. Improved delivery techniques could offset many of these concerns. In order to best leverage our knowledge and treatment of individual cancer patients, both therapeutics and their delivery systems can be personalized for a given patient.

Adapting to the tumour microenvironment

The tumour microenvironment heavily influences patient prognosis, as it affects chemotherapeutic efficacy 195 . Although the EPR effect and FDA approval of early NP systems has given hope for NP-based delivery, these early systems do not improve overall patient survival, and there is still significant work to be done using smart NP designs to improve cargo delivery or remodel microenvironments and thus increase the efficacy of existing therapies 69 .

For example, incorporating cell membranes into NPs can improve their accumulation in cancerous tissue. NPs wrapped with membranes that are harvested from a patient’s own cancer cells homotypically adhere to patient-derived cancer cell lines; mismatch between the donor and host results in weak targeting 196 , 197 . NPs wrapped with macrophage or leukocyte membranes recognize tumours, and hybrid membranes, such as erythrocyte–cancer cell hybrids, can further increase specificity 197 , 198 , 199 . NPs that utilize these membranes show a twofold to threefold increase in drug activity over the free drug 198 . In a similar fashion, material properties can cause NPs to preferentially distribute to certain tissues. For example, a poly(β-amino-ester) (PBAE) ter-polymer/PEG lipid conjugate was optimized for lung localization, achieving efficacy two orders of magnitude above the pre-optimized form both in vitro and in vivo 179 . Other PBAE polymers have been developed that preferentially target glioblastoma cells over healthy cells in vitro 200 . Even AuNPs can be optimized to passively target triple-negative breast cancer cells, which notoriously lack traditional cell surface targets 201 . Designs like these, as well as the more generalizable trends for NP size and shape, are being used to improve the percentage of chemotherapeutic dose that makes it to the solid tumour site.

Within the tumour microenvironment, responsive particles can improve tumour penetration, overcoming the high interstitial pressure and dense ECM that typically prevent NP permeation 150 , 202 . Endogenous triggers — such as the acidic and hypoxic environment of the tumour — can be used to induce NP degradation and drug release 147 , 150 , 203 , 204 . High enzyme levels of matrix metalloproteinases (MMPs) and other extracellular proteases can serve as triggers 10 , 205 , 206 , 207 , and the Warburg effect, a metabolic shift towards anaerobic glycolysis 195 , can be exploited as well 208 . Exogenous triggers — such as light, sound waves, radio frequencies and magnetic fields — can also be used and tightly controlled from outside the body 206 . A non-invasive existing clinical technique, ultrasound, can trigger local release from a systemically administered particle 68 , 209 . Near-infrared light, another exogenous trigger, has low absorption by natural tissues and therefore good biocompatibility 210 , 211 . Regardless of the trigger type, chemotherapeutics delivered locally in this responsive fashion have fewer off-target toxicities and other negative systemic effects.

One example of smart NP design, iCluster, is a stimuli-responsive clustered NP system that breaks down into smaller and smaller pieces as it overcomes biological barriers in the tumour environment 204 . The initial size of ~100 nm favours extended circulation in the bloodstream and capitalizes on the EPR effect as the NP extravasates through the tumour vasculature 204 . At the tumour site, the low pH triggers breakdown of the system into much smaller (~5 nm) dendrimers, which have improved tissue penetration and thus deliver more of the platinum chemotherapeutic cisplatin to cancer cells 204 . This system is a vast improvement over the traditional intravenous administration of free cisplatin: administration of the free drug inhibits tumour growth by 10%, whereas the iCluster system inhibits growth by up to 95% in in vivo studies 204 . Additionally, free cisplatin commonly causes irritation and cytotoxicity, especially in the kidney. Size-switching is not a unique property of this system, and has been achieved using various other triggers and materials 10 , 207 , 212 , 213 . NP systems such as these have great potential to improve therapeutic efficacy; their design is versatile and can be tailored specifically to the tumour microenvironment.

Another example of optimally designed delivery is a poly(acrylamide-co-methacrylic acid) nanogel, which can be modified with bioactive moieties for numerous applications including local pH response, cell targeting, transduction of visible light for photothermal therapy or degradation in the intracellular environment 11 . This platform was able to maintain the function of multiple modifications, allowing for each added small molecule, peptide or protein to contribute new responsive or recognitive properties 11 . Nanomaterials that utilize a similar, modular approach could be rapidly designed to deliver multiple therapeutic agents intracellularly or respond to sequential biological stimuli.

Active targeting to cancer cells

Existing chemotherapeutics have various mechanisms and sites of action. Some disrupt DNA within the nucleus (doxorubicin, platinum drugs), and others work within the cytosol or affect organelles such as the mitochondria 214 . Each drug must be delivered to its site of action at therapeutic levels in order to work properly, indicating a need for NP trafficking to these sites.

Antibodies, carbohydrates and other ligands on the NP surface can induce specific and efficient NP uptake 124 . Examples of tumour cell targeting moieties include antibodies 121 , peptides 126 , integrin ligands 124 , glucose 122 , transferrin 34 , 215 and folic acid 123 (Fig.  5 ). As these technologies advance, some systems now incorporate multiple targeting modalities in a single NP 195 . Whereas some of these targeting schemes are generalizable, such as folic acid (folate receptors are overexpressed on >40% of human cancers) 216 , most require tumour profiling to establish receptor or ligand overexpression. Additionally, not all receptor targeting improves specificity. Some receptors overexpressed in tumour cell lines are also expressed in healthy tissues, limiting efficacy.

There is also often a trade-off between residence time in the circulation and cellular uptake. Recently, NPs have been developed with detachable stealth corona systems and charge-reversal systems (negative or neutral charge for circulation, positive charge for uptake), in an attempt to optimize both properties 217 , 218 . One such system utilizes an MMP-degradable linker to attach PEG to the surface of the NP: in the tumour microenvironment, the PEG coating is degraded, exposing a cell-penetrating peptide 219 . In this way, systems can be developed that change a given property to optimize for the delivery barrier they currently face.

NPs for immunotherapy

Although immune checkpoint inhibitors have shown significant promise for cancer treatment 220 , there are still challenges with efficacy, patient variability and off-target effects when immunomodulators are used 221 . Some immunotherapeutics, such as proteins, have limited delivery potential when administered freely, and thus NPs have the potential to significantly improve delivery by protecting immunotherapeutics and enhancing their interaction with immune cells 222 .

Immune activation

The immune system is trained to eliminate cancerous cells from the body, but certain genetic traits can allow cancerous cells to evade and suppress immune cells. To resensitize these cells, cancer vaccines aim to train the body to recognize cancerous cells by using antigens either from the patient or from allogenic tumour cells. For example, Sipuleucel-T, an FDA-approved cancer vaccine (albeit with limited efficacy) 221 , utilizes recombinant antigens specific to the tumour type. Although the drugs are not yet in the clinic, other groups have also developed synthetic peptides and tumour lysates with the ultimate goal of patient personalization 223 , 224 , 225 . NPs can protect these antigens from degradation, improve the likelihood that they are presented to target immune cells and reduce off-target effects. Antigen-presenting cells (APCs) that take up these NP systems present the antigen cargo to T cells to prime and activate them. NPs used in these systems can be polymeric (PLGA) 226 , lipid-based (liposomes, LNPs) 227 , 228 , inorganic (gold, silica) 229 , 230 or biologically derived (cell-membrane vesicles) 231 , 232 . NP-based cancer vaccines are currently being used in clinical trials 233 . Recently, NPs have been extensively explored in vaccines against SARS-CoV-2 (which causes COVID-19), with multiple successful late-stage clinical trials. Companies such as Moderna and BioNTech use LNPs to encapsulate mRNA that encodes for a COVID-19 antigen. As of 30 November 2020, Moderna and BioNTech/Pfizer have met their primary efficacy end points in phase III trials and have applied for Emergency Use Authorization. As with other applications, NP architecture, material properties and active targeting can affect cellular uptake, antigen presentation and the strength of the immune response 234 .

Macrophages, B cells and dendritic cells are all APCs and can be targeted by NPs to improve the specificity of immune activation. Passive targeting includes optimizing size, shape ratios and using positively charged particles to interact with the negatively charged cell membranes 235 , 236 . APCs also express numerous carbohydrate-recognizing lectin receptors for endocytosis, and these have been exploited for cell-specific active targeting 178 . Some of these lectin receptors are expressed at high levels in certain APCs, such as the C-type lectin receptors lymphocyte antigen 75 (also known as DEC-205) and C-type lectin domain family 9 member A (CLEC9A), which can be used to target dendritic cells 237 . Mannose is commonly used to target macrophages and tumour-associated macrophages 238 , 239 , 240 , but can target dendritic cells as well 241 . Particles coated with galactose, dextran or sialoadhesin can deliver to macrophages 198 , 240 , 242 . CD19-targeting NPs can be used to actively target B cells 243 , and NPs with lipoprotein surfaces can activate the scavenger receptor class B1 (SRB1) receptor on dendritic cells 244 . More generally, NP properties can be optimized for accumulation at tolerogenic organs, such as the liver and spleen, where immunological antigens are naturally produced 245 . Immune-recruiting systems, such as polymeric hydrogels and scaffolds, could also be used to optimize interactions with APCs. These systems work with APC-targeted NPs, allowing them to recruit and reprogramme APCs 246 . All of these methods aim to increase the likelihood that an antigen will interact with an APC, improving the efficacy of antigen-based therapies and lowering the dosage needed to reach therapeutic levels.

The stimulator of interferon genes (STING) pathway also leads to immune cell activation and antitumour effects, and can be activated by cytosolic double-stranded DNA (which typically comes from pathogens). STING agonists, typically cyclic dinucleotides, show promising antitumour activity, but are unstable and highly polar, which reduces cellular uptake 247 . NPs improve the delivery of STING agonists 248 , 249 , 250 ; a single STING NP dose of one formulation increased survival for at least 80 days in mice 249 . Additionally, some NPs with cyclic structures (cyclic lipids) that mimic double-stranded DNA can stimulate STING regardless of their cargo 174 .

Other immunotherapy approaches target T cells directly. Numerous targeting schemes have been used to target NPs to T cells. Examples include NPs targeting PD1 (ref. 251 ), CD3 (ref. 134 ) and THY1 (also known as CD90) 135 . The tLyp1 peptide, typically used for tumour targeting, has been used to target regulatory T cells, an immunosuppressive T cell subtype 193 . Checkpoint inhibitors, an anticancer immune-boosting strategy, are typically monoclonal antibodies that target PD1, PDL1 or CTLA4. As for other applications, the usage of free antibodies is limited by stability concerns. Additionally, less than a third of patients who receive these checkpoint inhibitors see a robust response 249 . In an attempt to improve these therapies by enhancing efficacy and reducing side effects, NPs have been formulated for monoclonal antibody (anti-PD1) delivery 252 , 253 , and other NP formulations disrupt immune checkpoints through the delivery of small interfering RNAs (siRNAs) 254 .

Genetically modified T cells have also shown promise in the treatment of metastatic and blood cancers. These T cells are constructed to express transgenic T cell receptors (TCRs) or CARs, which allow for T cells to specifically target and eliminate cancerous cells 255 . These T cells are extracted from patients before in vitro expansion using artificial APCs, and new NP formulations may allow for translation of this process in vivo 255 , 256 . Artificial APC design is similar to that of traditional NPs in the sense that their architecture, materials and targeting influence T cell activation 257 . Alternative methods of CAR T production could reduce the complexity of antigen delivery to T cells using NPs, including the delivery of CAR-encoding DNA in vivo and the delivery of CAR-encoding mRNA to produce transiently modified T cells 258 , 259 .

Immune suppression

Diseases such as rheumatoid arthritis and systemic lupus erythematosus also result from incorrect immune regulation: hyperactivation. In these autoimmune diseases, T cells and B cells are sensitized to self-antigens 260 . Autoimmune diseases are typically treated with general immunosuppressants, which can cause serious side effects. Conditions caused by immune overactivation could benefit from more targeted immunotherapies.

Cellular targets for immune suppression include APCs 261 , autoreactive T cells and B cells 262 , and regulatory T cells and B cells 263 , 264 . Antigen-specific immunotherapy aims to reprogramme or reduce reactive cells or impart them with tolerance to certain antigens 260 . By targeting a subset of immune cells, antigen-specific immunotherapy has potential to modulate the immune system without compromising systemic immunity. Passive and active targeting schemes similar to those used in immune-activating therapies are used for immune inhibition. For example, NPs coated with anti-CD2/CD4 antibodies target T cells and can be used to increase the number of regulatory T cells in circulation, whereas non-coated NPs at equivalent doses could not 263 . Similarly, sialic acid-binding immunoglobulin-like lectins (Siglecs) can be used to target and induce tolerance in B cells 262 .

Immune tolerance can also be induced through the delivery of immunosuppressant agents. NPs that deliver IL-2 and TGFβ can expand the number of regulatory T cells in vivo, suppressing the symptoms of lupus 263 . The active form of vitamin D 3 has immunosuppressive effects because it modulates dendritic cell function 261 . Active vitamin D 3 can cause hypercalcaemia when administered systemically, so NP delivery is a promising alternate strategy. PLGA NPs have been used extensively to deliver immunomodulators and prevent allograft rejection 265 ; PLGA NPs anchored to a hydrogel allow for local and sustained (28-day) delivery of tacrolimus, a common immunosuppressant 266 . For more long-term effects, genetic engineering — reprogramming immune cells at the genomic level — could be effective 267 .

NPs for genome editing

Recent advances in CRISPR, transcription activator-like effector nuclease (TALEN) and zinc-finger nuclease (ZFN) technologies are making it increasingly easy to engineer the genome for widespread use in biomedical research, drug development and discovery, and gene therapy 268 . This is important in the context of precision medicine, as over 3,000 human genes have been associated with Mendelian diseases but less than 5% of rare diseases have effective treatments 268 . Advances in genome editing are now making it possible to correct many of these rare diseases. However, efficient and safe delivery is still needed for genome-editing systems to effectively target and enter tissues and cells of interest, while also minimizing toxicity 269 . Delivery of genome-editing systems is challenging because these systems are multicomponent, hold sensitive cargo and need to overcome several extracellular and intracellular biological barriers to reach the genome of target cells. Lipid-based and polymer-based NPs have delivered a range of nucleic acids in vivo, and are in various stages of clinical development 24 , 44 , 104 , 183 . For example, a LNP siRNA drug termed Onpattro (patisiran) was recently approved by the FDA for the treatment of amyloidosis 270 . In the context of genome editing, NPs have the potential to be less toxic and immunogenic than viral vectors, which have a history of safety concerns 271 , 272 .

Intracellular targeting

Most NP-based systems for genome editing are formulated by electrostatic complexation of nucleic acids with cationic materials, which are delivered intracellularly through mechanisms including receptor-mediated endocytosis and phagocytosis 273 . Cationic materials both complex with nucleic acids and impart responsive properties to NPs that aid in endosomal escape. Charged materials currently used for nucleic acid delivery include lipids (lipofectamine, rationally designed lipids, combinatorial libraries of ionizable lipid-like materials) 21 , 22 , 274 and polymers (polyethylene imine (PEI), jetPEI, poly(amido amine) (PAA), polylysine (PLL), cyclodextrins and poly(β-amino esters)) 275 , 276 , 277 , 278 . These systems are responsive to the intracellular environment, and can be optimized to incorporate passive and active targeting elements to ensure endocytic uptake.

The final destination of the cargo for RNA interference is the cytosol 214 . However, gene editing requires access to DNA. Strategies for nuclear targeting generally fall into two categories: using particles that are small enough to pass through the nuclear pore complex, or incorporating functionality that is used after endosomal escape 214 , 279 . NP properties can passively influence intracellular trafficking and the final destination 179 , but particles can also be actively targeted to specific intracellular sites and organelles, such as the mitochondria 27 . Defects within the mitochondrial DNA can also play a significant role in disease onset. However, success with mitochondrial DNA genetic engineering is currently limited to highly controlled in vitro settings 214 .

Applications of genome engineering

Cystic fibrosis is caused by genetic defects in the gene that encodes cystic fibrosis transmembrane conductance regulator (CFTR) protein. There is currently no cure for this life-threatening disorder, but it is a monogenic disorder and therefore amenable to gene therapy. In vitro, the CFTR  gene can be replaced or precisely repaired 280 . However, gene therapy for cystic fibrosis has been largely unsuccessful in vivo due to issues with gene expression and delivery 281 , 282 . However, NPs could aid in overcoming these delivery barriers 283 , 284 , 285 . Multiple inhalable NP formulations have been developed, and some have shown successful delivery of genetic material such as mRNA 248 , 286 .

Cystic fibrosis affects cells that produce mucus, making the mucus extra thick. This is the main symptom of the disease but is also a significant barrier to delivery. NPs have been developed with improved muco-penetrating properties for use in lung delivery for cystic fibrosis and for oral delivery. NPs smaller than the mucus mesh pores have improved penetration, as do systems with inert hydrophilic coatings (such as PEG or polyethylene oxide) 283 , 285 . PEGylation has been shown to improve penetration through cystic fibrosis mucus ex vivo 287 . However, mucus can be highly variable between patients, and existing murine models may not accurately mimic the thickened airway mucus produced by patients with cystic fibrosis 283 . Other existing methods for improving mucosal delivery include the incorporation of muco-penetrating lipoplexes 176 , mucolytic proteins 149 , thiolated hyaluronic acid coatings 48 and N -acetylcysteine 137 . All of these methods attempt to improve the transversal of mucosal barriers by altering NP surface properties 118 .

Even though it has been estimated that restoration of 10–35% of CFTR protein function would substantially improve the manifestations of the disease 284 , a higher percentage would be needed for a genuine cure. For genetic diseases such as these, the fetal stage is the most effective time for gene editing, as genetic defects are present in a small number of cells. However, fetal delivery is a challenge.

The biological barriers to in utero delivery are actually fewer than may be expected. NP therapies can be injected directly into an umbilical vessel, the amniotic fluid or specific fetal tissue 288 . The limitations in fetal delivery come from concerns regarding the interaction between fetus and mother. Viral vectors have successfully delivered DNA editing machinery in utero in a mouse model 289 . However, these vectors have more toxicity concerns than NP systems do. Although the usage of NPs for this type of delivery is not widespread, there have been early successes with in utero NP delivery of peptide nucleic acids, resulting in a level of gene editing sufficient to alter the disease to manageable levels 290 .

Looking to the future, NPs have the potential to improve genome editing by exerting more precise control and reducing safety concerns. Several companies, including CRISPR Therapeutics, Intellia Therapeutics and Editas Medicine, are currently developing CRISPR–Cas9 therapeutics. Intellia Therapeutics is currently developing LNPs for in vivo delivery to treat several liver diseases, including amyloidosis, α1-antitrypsin deficiency and hepatitis B virus infection. With precision NP design, gene editing holds promise to cure diseases and significantly improve patient lives.

Conclusions

This Review has discussed numerous NP designs optimized for therapeutic delivery and engineered to overcome the heterogeneous biological barriers found across patient populations and diseases. These barriers to delivery are complicated by patient comorbidities, varying stages of disease progression and unique physiologies. This diverse array of needs can be met using NPs designed for different patient populations or pathologies, or intersections of the two. NP platforms offer an assortment of modifiable features such as size, shape, charge, surface properties and responsiveness that can be selected to optimize delivery for a specific application, therapeutic and patient population. This customization can be utilized synergistically with precision medicine therapies to improve patient stratification methods when screening NP platforms, widen the accessibility of precision therapeutics by allowing new patients to qualify for existing therapies with newly enhanced delivery mechanisms and, ultimately, increase the overall therapeutic efficacy of both precision medicines and NP delivery platforms.

Of these NP characteristics, size and shape have been extensively studied across numerous biological states, and, in some cases, trends have been identified that can be used for intelligent NP design. For example, NP charge is of particular importance in muco-penetrating applications and intracellular applications that require endosomal escape, whereas targeting surface markers takes precedent in applications where specific cell types must uptake NPs, as in many cancer and immunotherapy applications. However, as the design considerations become more complicated, so do efforts to generalize trends across large populations — sacrificing the accuracy of the findings within a small population in the hope of generating an all-encompassing principle of delivery. Therefore, investigations of NP design and the resulting interactions within the human body need to be more thoroughly analysed to improve the specificity of these claims, especially as we move towards stratifying patient populations to determine the most suitable NP platforms for these subgroups. Through the continued exploration of NP technologies in laboratory settings, researchers have the opportunity to collect data and analyse outcomes to add to the ever-growing library of known design–function relationship trends in nanomedicine. However, it is imperative that the trends observed in research settings be contextualized before attempting to generalize findings broadly, as seemingly minor differences in NP composition, animal models and pathology may greatly alter the performance of NPs and must be considered when moving NP technology towards clinical translation.

Current clinical successes with NPs in precision medicine have been largely diagnostic, such as the ability to recognize early stages of a disease by specific ligand–receptor interactions or the use of biomarkers to identify which therapeutics might be best for a particular patient. For example, determining the level of the EPR effect that a cancer patient exhibits can inform how effectively a NP therapy would accumulate at the solid tumour site 152 . However, this Review has focused on potential therapeutic applications of NPs — specifically, their use in the precision medicine fields of oncology, immunotherapy and genome engineering — as these platforms have immense potential to improve efficacy of precision medicine therapies but have yet to see the clinical progress achieved by diagnostic applications. This lack of clinical progress is likely because NP platforms are screened for efficacy in broad populations, in which the vast heterogeneity of biological barriers in patients could mask the potential for successfully treating slightly smaller subgroups. As mentioned previously, the introduction of stratified patient populations may also accelerate clinical progress, as stratified populations will likely respond more uniformly to NP treatment. However, as ongoing clinical trials are generally unstratified, we are currently unable to predict which NP platforms will be most useful for precision applications, and more stratified trials are necessary.

Of course, moving to screen NPs through a precision lens — thus, limiting the number of patients that are eligible to receive a medication — will reduce the potential market size of each NP-based therapeutic. This reduction may raise concern when considering the high cost of development for advanced NP designs and, thus, the increased financial risks associated with the potentially failed clinical translation of a NP formulation. However, NP platforms that are found to work well in specific patient populations may have applications in the delivery of numerous therapeutics, both precision-based and generic. Thus, the development of one highly effective NP platform for a stratified group could lead to multiple successful clinical applications. Furthermore, precision NP designs may allow for greater therapeutic efficacy compared with NPs developed for broad populations, and significant improvements in survival, quality of life and even dosing could justify the higher price point of these precision delivery systems.

As more advanced NP designs are explored, this research could influence the future rational design of drug carriers for various therapeutics, both personalized and generic, thereby benefitting an array of cargos including small molecules, nucleotides and proteins. Furthermore, the approval of more NP platforms could ease the path to the clinic for novel applications of these NPs for the delivery of previously approved therapeutic cargos. By working to develop more NP platforms and precision medicines for FDA approval and clinical use, we are taking steps towards more modular patient therapy designs and creating the potential for future prescriptions to not just include the optimal therapy but also pair it with the optimized delivery platform. The concept of offering multiple delivery platforms for a single therapeutic is not new in the clinical marketplace. Commonly prescribed drugs such as those used for birth control are already offered in multiple forms (oral pills, injections and implants) to fit the patient’s lifestyle. The expansion of both precision medicine and advanced NP platforms will contribute to the continued clinical progress of personalized medicines, allowing for seemingly niche markets to grow.

In precision medicine-relevant applications, the usage of NPs allows for improved cellular targeting, fewer off-target effects and more tailored therapies such as multidrug treatments. All of this can be achieved by engineering NPs for the application at hand, improving accumulation at the site of interest and introducing responsivity for on-demand drug release, to minimize unwanted toxicities and enable a new range of dosages or combinatorial treatments. By optimizing this specificity and local activity of NP delivery systems, the effects of precision medicine therapeutics can be improved as well, widening the populations they benefit and improving patient outcome overall. As the work described in this Review shows, intelligent NP design can improve precision medicine as a whole and the insight provided by precision medicine — such as patient stratification and genetic profiling — can inform the rational selection of a NP platform to, ultimately, generate the ideal NP-based precision therapy.

The White House Office of the Press Secretary. National Nanotechnology Initiative: leading to the next industrial revolution. White House https://clintonwhitehouse4.archives.gov/WH/New/html/20000121_4.html (2000).

Kou, L. et al. Transporter-guided delivery of nanoparticles to improve drug permeation across cellular barriers and drug exposure to selective cell types. Front. Pharmacol. 9 , 1–16 (2018).

Google Scholar  

Blanco, E., Shen, H. & Ferrari, M. Principles of nanoparticle design for overcoming biological barriers to drug delivery. Nat. Biotechnol. 33 , 941–951 (2015).

CAS   PubMed   PubMed Central   Google Scholar  

Mitragotri, S. et al. Drug delivery research for the future: expanding the nano horizons and beyond. J. Control. Release 246 , 183–184 (2017).

CAS   PubMed   Google Scholar  

Wechsler, M. E., Vela Ramirez, J. E. & Peppas, N. A. 110th anniversary: nanoparticle mediated drug delivery for the treatment of Alzheimer’s disease: crossing the blood–brain barrier. Ind. Eng. Chem. Res. 58 , 15079–15087 (2019).

Hua, S., de Matos, M. B. C., Metselaar, J. M. & Storm, G. Current trends and challenges in the clinical translation of nanoparticulate nanomedicines: pathways for translational development and commercialization. Front. Pharmacol. 9 , 790 (2018).

PubMed   PubMed Central   Google Scholar  

Anselmo, A. C. & Mitragotri, S. Nanoparticles in the clinic: an update. Bioeng. Transl. Med. 4 , 1–16 (2019). This review comprehensively examines recent NP systems that are currently clinically approved or under clinical trials .

Wagner, A. M., Gran, M. P. & Peppas, N. A. Designing the new generation of intelligent biocompatible carriers for protein and peptide delivery. Acta Pharm. Sin. B 8 , 147–164 (2018).

Culver, H. R., Clegg, J. R. & Peppas, N. A. Analyte-responsive hydrogels: intelligent materials for biosensing and drug delivery. Acc. Chem. Res. 50 , 170–178 (2017).

Ruan, S. et al. Matrix metalloproteinase-sensitive size-shrinkable nanoparticles for deep tumor penetration and pH triggered doxorubicin release. Biomaterials 60 , 100–110 (2015).

Clegg, J. R. et al. Synthetic networks with tunable responsiveness, biodegradation, and molecular recognition for precision medicine applications. Sci. Adv. 5 , eaax7946 (2019).

Cheng, Q. et al. Selective organ targeting (SORT) nanoparticles for tissue-specific mRNA delivery and CRISPR–Cas gene editing. Nat. Nanotechnol. 15 , 313–320 (2020).

Collins, F. S. & Varmus, H. A new initiative on precision medicine. N. Engl. J. Med. 372 , 793–795 (2015).

Fonseca-Santos, B., Gremião, M. P. D. & Chorilli, M. Nanotechnology-based drug delivery systems for the treatment of Alzheimer’s disease. Int. J. Nanomed. 10 , 4981–5003 (2015).

CAS   Google Scholar  

Sercombe, L. et al. Advances and challenges of liposome assisted drug delivery. Front. Pharmacol . 6 , 286 (2015).

Fenton, O. S., Olafson, K. N., Pillai, P. S., Mitchell, M. J. & Langer, R. Advances in biomaterials for drug delivery. Adv. Mater. 30 , 1705328 (2018).

Sarfraz, M. et al. Development of dual drug loaded nanosized liposomal formulation by a reengineered ethanolic injection method and its pre-clinical pharmacokinetic studies. Pharmaceutics 10 , 1–22 (2018).

Sedighi, M. et al. Rapid optimization of liposome characteristics using a combined microfluidics and design-of-experiment approach. Drug Deliv. Transl. Res. 9 , 404–413 (2019).

Alyautdin, R., Khalin, I., Nafeeza, M. I., Haron, M. H. & Kuznetsov, D. Nanoscale drug delivery systems and the blood–brain barrier. Int. J. Nanomed. 9 , 795–811 (2014).

Leung, A. K. K., Tam, Y. Y. C., Chen, S., Hafez, I. M. & Cullis, P. R. Microfluidic mixing: a general method for encapsulating macromolecules in lipid nanoparticle systems. J. Phys. Chem. B 119 , 8698–8706 (2015).

Cheng, X. & Lee, R. J. The role of helper lipids in lipid nanoparticles (LNPs) designed for oligonucleotide delivery. Adv. Drug Deliv. Rev. 99 , 129–137 (2016).

Kulkarni, J. A., Witzigmann, D., Leung, J., Tam, Y. Y. C. & Cullis, P. R. On the role of helper lipids in lipid nanoparticle formulations of siRNA. Nanoscale 11 , 21733–21739 (2019).

Berraondo, P., Martini, P. G. V., Avila, M. A. & Fontanellas, A. Messenger RNA therapy for rare genetic metabolic diseases. Gut 68 , 1323–1330 (2019).

Vhora, I., Lalani, R., Bhatt, P., Patil, S. & Misra, A. Lipid-nucleic acid nanoparticles of novel ionizable lipids for systemic BMP-9 gene delivery to bone-marrow mesenchymal stem cells for osteoinduction. Int. J. Pharm. 563 , 324–336 (2019).

Patel, S., Ryals, R. C., Weller, K. K., Pennesi, M. E. & Sahay, G. Lipid nanoparticles for delivery of messenger RNA to the back of the eye. J. Control. Rel. 303 , 91–100 (2019).

Brown, S. B., Wang, L., Jungels, R. R. & Sharma, B. Effects of cartilage-targeting moieties on nanoparticle biodistribution in healthy and osteoarthritic joints. Acta Biomater. 101 , 469–483 (2019).

Zhang, C. X. et al. Mitochondria-targeted cyclosporin A delivery system to treat myocardial ischemia reperfusion injury of rats. J. Nanobiotechnol. 17 , 1–16 (2019).

Le, Z. et al. Hydrogen-bonded tannic acid-based anticancer nanoparticle for enhancement of oral chemotherapy. ACS Appl. Mater. Interfaces 10 , 42186–42197 (2018).

He, C. et al. siRNA release kinetics from polymeric nanoparticles correlate with RNAi efficiency and inflammation therapy via oral delivery. Acta Biomater. 103 , 213–222 (2020).

Zhang, L. et al. Microfluidic-assisted polymer-protein assembly to fabricate homogeneous functional nanoparticles. Mater. Sci. Eng. C 111 , 110768 (2020).

Caldorera-Moore, M., Vela Ramirez, J. E. & Peppas, N. A. Transport and delivery of interferon-α through epithelial tight junctions via pH-responsive poly(methacrylic acid-grafted-ethylene glycol) nanoparticles. J. Drug Target. 27 , 582–589 (2019).

Knight, F. C. et al. Mucosal immunization with a pH-responsive nanoparticle vaccine induces protective CD8 + lung-resident memory T cells. ACS Nano 13 , 10939–10960 (2019).

Strand, M. S. et al. Precision delivery of RAS-inhibiting siRNA to KRAS driven cancer via peptide-based nanoparticles. Oncotarget 10 , 4761–4775 (2019).

Jose, S. et al. Transferrin-conjugated docetaxel–PLGA nanoparticles for tumor targeting: influence on MCF-7 cell cycle. Polymers 11 , 1905 (2019).

CAS   PubMed Central   Google Scholar  

Liu, X. et al. Glucose and H 2 O 2 dual-responsive polymeric micelles for the self-regulated release of insulin. ACS Appl. Bio Mater. 3 , 1598–1606 (2020).

Afsharzadeh, M., Hashemi, M., Mokhtarzadeh, A., Abnous, K. & Ramezani, M. Recent advances in co-delivery systems based on polymeric nanoparticle for cancer treatment. Artif. Cells Nanomed. Biotechnol. 46 , 1095–1110 (2018).

Patra, J. K. et al. Nano based drug delivery systems: recent developments and future prospects. J. Nanobiotechnol. 16 , 71 (2018).

Volpatti, L. R. et al. Glucose-responsive nanoparticles for rapid and extended self-regulated insulin delivery. ACS Nano 14 , 488–497 (2020).

Rideau, E., Dimova, R., Schwille, P., Wurm, F. R. & Landfester, K. Liposomes and polymersomes: a comparative review towards cell mimicking. Chem. Soc. Rev. 47 , 8572–8610 (2018).

Shae, D. et al. Endosomolytic polymersomes increase the activity of cyclic dinucleotide STING agonists to enhance cancer immunotherapy. Nat. Nanotechnol. 14 , 269–278 (2019).

Zelmer, C. et al. Organelle-specific targeting of polymersomes into the cell nucleus. Proc. Natl Acad. Sci. USA 117 , 2770–2778 (2020).

Lee, S.-W. et al. An open-label, randomized, parallel, phase II trial to evaluate the efficacy and safety of a cremophor-free polymeric micelle formulation of paclitaxel as first-line treatment for ovarian cancer: a Korean Gynecologic Oncology Group Study (KGOG-3021). Cancer Res. Treat. 50 , 195–203 (2018).

Xu, L., Zhang, H. & Wu, Y. Dendrimer advances for the central nervous system delivery of therapeutics. ACS Chem. Neurosci. 5 , 2–13 (2014).

Mendes, L. P., Pan, J. & Torchilin, V. P. Dendrimers as nanocarriers for nucleic acid and drug delivery in cancer therapy. Molecules 22 , 1401 (2017).

Kannan, R. M., Nance, E., Kannan, S. & Tomalia, D. A. Emerging concepts in dendrimer-based nanomedicine: from design principles to clinical applications. J. Intern. Med. 276 , 579–617 (2014).

Menjoge, A. R., Kannan, R. M. & Tomalia, D. A. Dendrimer-based drug and imaging conjugates: design considerations for nanomedical applications. Drug Discov. Today 15 , 171–185 (2010).

Cao, S.-J. et al. Nanoparticles: oral delivery for protein and peptide drugs. AAPS PharmSciTech 20 , 190 (2019).

Tian, H. et al. Uniform core-shell nanoparticles with thiolated hyaluronic acid coating to enhance oral delivery of insulin. Adv. Healthc. Mater. 7 , 1–12 (2018).

Valcourt, D. M., Dang, M. N., Scully, M. A. & Day, E. S. Nanoparticle-mediated co-delivery of Notch-1 antibodies and ABT-737 as a potent treatment strategy for triple-negative breast cancer. ACS Nano 14 , 3378–3388 (2020).

Yang, W., Liang, H., Ma, S., Wang, D. & Huang, J. Gold nanoparticle based photothermal therapy: development and application for effective cancer treatment. Sustain. Mater. Technol. 22 , e00109 (2019).

Wang, J., Potocny, A. M., Rosenthal, J. & Day, E. S. Gold nanoshell-linear tetrapyrrole conjugates for near infrared-activated dual photodynamic and photothermal therapies. ACS Omega 5 , 926–940 (2020).

Bobo, D., Robinson, K. J., Islam, J., Thurecht, K. J. & Corrie, S. R. Nanoparticle-based medicines: a review of FDA-approved materials and clinical trials to date. Pharm. Res. 33 , 2373–2387 (2016).

Arias, L. S. et al. Iron oxide nanoparticles for biomedical applications: a perspective on synthesis, drugs, antimicrobial activity, and toxicity. Antibiotics 7 , 46 (2018).

PubMed Central   Google Scholar  

Huang, K.-W. et al. Highly efficient and tumor-selective nanoparticles for dual-targeted immunogene therapy against cancer. Sci. Adv. 6 , eaax5032 (2020).

Xu, C., Nam, J., Hong, H., Xu, Y. & Moon, J. J. Positron emission tomography-guided photodynamic therapy with biodegradable mesoporous silica nanoparticles for personalized cancer immunotherapy. ACS Nano 13 , 12148–12161 (2019).

Wagner, A. M., Knipe, J. M., Orive, G. & Peppas, N. A. Quantum dots in biomedical applications. Acta Biomater. 94 , 44–63 (2019).

Zhang, Y., Meng, S., Ding, J., Peng, Q. & Yu, Y. Transition metal-coordinated graphitic carbon nitride dots as a sensitive and facile fluorescent probe for β-amyloid peptide detection. Analyst 144 , 504–511 (2019).

Manshian, B. B., Jiménez, J., Himmelreich, U. & Soenen, S. J. Personalized medicine and follow-up of therapeutic delivery through exploitation of quantum dot toxicity. Biomaterials 127 , 1–12 (2017).

Aguado, B. A., Grim, J. C., Rosales, A. M., Watson-Capps, J. J. & Anseth, K. S. Engineering precision biomaterials for personalized medicine. Sci. Transl. Med. 10 , 8645 (2018).

van der Meel, R. et al. Smart cancer nanomedicine. Nat. Nanotechnol. 14 , 1007–1017 (2019).

Scheetz, L. et al. Engineering patient-specific cancer immunotherapies. Nat. Biomed. Eng. 3 , 768–782 (2019).

Papi, M. et al. Converting the personalized biomolecular corona of graphene oxide nanoflakes into a high-throughput diagnostic test for early cancer detection. Nanoscale 11 , 15339–15346 (2019).

Shaban, M., Hasanzadeh, M. & Solhi, E. An Fe 3 O 4 /PEDOT:PSS nanocomposite as an advanced electroconductive material for the biosensing of the prostate-specific antigen in unprocessed human plasma samples. Anal. Methods 11 , 5661–5672 (2019).

Zhang, T. et al. The synergistic effect of Au-COF nanosheets and artificial peroxidase Au@ZIF-8(NiPd) rhombic dodecahedra for signal amplification for biomarker detection. Nanoscale 11 , 20221–20227 (2019).

Kaur, A., Shimoni, O. & Wallach, M. Novel screening test for celiac disease using peptide functionalised gold nanoparticles. World J. Gastroenterol. 24 , 5379–5390 (2018).

Tan, T. et al. Bioinspired lipoproteins-mediated photothermia remodels tumor stroma to improve cancer cell accessibility of second nanoparticles. Nat. Commun. 10 , 3322 (2019).

Wilson, R. A. et al. MicroRNA regulation of endothelial TREX1 reprograms the tumour microenvironment. Nat. Commun. 7 , 13597 (2016).

Papa, A. L. et al. Ultrasound-sensitive nanoparticle aggregates for targeted drug delivery. Biomaterials 139 , 187–194 (2017).

Chen, Q. et al. Photothermal therapy promotes tumor infiltration and antitumor activity of CAR T cells. Adv. Mater. 31 , e1900192 (2019).

Patel, R. B. et al. Development of an in situ cancer vaccine via combinational radiation and bacterial-membrane-coated nanoparticles. Adv. Mater. 31 , e1902626 (2019).

Wilhelm, S. et al. Analysis of nanoparticle delivery to tumours. Nat. Rev. Mater. 1 , 16014 (2016). This comprehensive analysis of NP delivery surveys literature from 2005 to 2015 and finds that, on average, only 0.7% of the administered NP dose is delivered to a solid tumour, and, further, explores possible reasons for this inefficiency .

Ensign, L. M., Cone, R. & Hanes, J. Oral drug delivery with polymeric nanoparticles: the gastrointestinal mucus barriers. Adv. Drug Deliv. Rev. 64 , 557–570 (2012). This review thoroughly examines barriers of the gastrointestinal tract and more deeply explores their effects on orally administered NPs .

Nag, O. K. & Delehanty, J. B. Active cellular and subcellular targeting of nanoparticles for drug delivery. Pharmaceutics 11 , 543 (2019).

Oliva, N. et al. Regulation of dendrimer/dextran material performance by altered tissue microenvironment in inflammation and neoplasia. Sci. Transl. Med. 7 , 272ra11 (2015). This study looks at the impact of disease on the tissue microenvironment and discovers that in states of carcinogenesis and inflammation the material–tissue interactions change, even when established biomaterials are used .

Gehr, P. & Zellner, R. Biological Responsese to Nanoscale Particles: Molecular and Cellular Aspects and Methodological Approaches (Springer, 2019).

Hoshyar, N., Gray, S., Han, H. & Bao, G. The effect of nanoparticle size on in vivo pharmacokinetics and cellular interaction. Nanomedicine 11 , 673–692 (2016). This review presents the influence of NP size on cellular uptake, biodistribution and pharmacokinetics, and also introduces models for interpretation of this result .

Wang, X. Y., Ishida, T. & Kiwada, H. Anti-PEG IgM elicited by injection of liposomes is involved in the enhanced blood clearance of a subsequent dose of PEGylated liposomes. J. Control. Rel. 119 , 236–244 (2007).

McSweeney, M. D. et al. A minimal physiologically based pharmacokinetic model that predicts anti-PEG IgG-mediated clearance of PEGylated drugs in human and mouse. J. Control. Rel. 284 , 171–178 (2018).

Yang, Q. et al. Analysis of pre-existing IgG and IgM antibodies against polyethylene glycol (PEG) in the general population. Anal. Chem. 88 , 11804–11812 (2016).

Chen, B. M. et al. Measurement of pre-existing IgG and IgM antibodies against polyethylene glycol in healthy individuals. Anal. Chem. 88 , 10661–10666 (2016).

Hu, C. M. J. et al. Nanoparticle biointerfacing by platelet membrane cloaking. Nature 526 , 118–121 (2015).

Wibroe, P. P. et al. Bypassing adverse injection reactions to nanoparticles through shape modification and attachment to erythrocytes. Nat. Nanotechnol. 12 , 589–594 (2017).

Anselmo, A. C. et al. Platelet-like nanoparticles: mimicking shape, flexibility, and surface biology of platelets to target vascular injuries. ACS Nano 8 , 11243–11253 (2014).

Palanikumar, L. et al. pH-responsive high stability polymeric nanoparticles for targeted delivery of anticancer therapeutics. Commun. Biol. 3 , 95 (2020).

Tanaka, R. et al. Furry nanoparticles: synthesis and characterization of nanoemulsion-mediated core crosslinked nanoparticles and their robust stability in vivo. Polym. Chem. 11 , 4408–4416 (2020).

Jarvis, M. et al. Detachment of ligands from nanoparticle surface under flow and endothelial cell contact: assessment using microfluidic devices. Bioeng. Transl. Med. 3 , 148–155 (2018).

Hosta-Rigau, L. & Städler, B. Shear stress and its effect on the interaction of myoblast cells with nanosized drug delivery vehicles. Mol. Pharm. 10 , 2707–2712 (2013).

Khor, S. Y. et al. Elucidating the influences of size, surface chemistry, and dynamic flow on cellular association of nanoparticles made by polymerization-induced self-assembly. Small 14 , 1–13 (2018).

Cooley, M. et al. Influence of particle size and shape on their margination and wall-adhesion: implications in drug delivery vehicle design across nano-to-micro scale. Nanoscale 10 , 15350–15364 (2018).

Pieter, S. & Mahmoudi, M. Drug Delivery Systems (World Scientific, 2017).

Da Silva-Candal, A. et al. Shape effect in active targeting of nanoparticles to inflamed cerebral endothelium under static and flow conditions. J. Control. Rel. 309 , 94–105 (2019). This work studies the effect of shape on NP targeting efficiency in the bloodstream, with rod-shaped particles showing better accumulation than spherical .

Uhl, C. G., Gao, Y., Zhou, S. & Liu, Y. The shape effect on polymer nanoparticle transport in a blood vessel. RSC Adv. 8 , 8089–8100 (2018).

Malota, Z., Glowacki, J., Sadowski, W. & Kostur, M. Numerical analysis of the impact of flow rate, heart rate, vessel geometry, and degree of stenosis on coronary hemodynamic indices. BMC Cardiovasc. Disord. 18 , 1–16 (2018).

Chacón-Lozsán, F., et al. Hemodynamic management of high blood pressure. Ther. Adv. Cardiol. 1 , 73–79 (2017).

Tenzer, S. et al. Rapid formation of plasma protein corona critically affects nanoparticle pathophysiology. Nat. Nanotechnol. 8 , 772–781 (2013).

Cox, A. et al. Evolution of nanoparticle protein corona across the blood–brain barrier. ACS Nano 12 , 7292–7300 (2018).

Schwartz, S. Unmet needs in developing nanoparticles for precision medicine. Nanomedicine 12 , 271–274 (2017).

von Roemeling, C., Jiang, W., Chan, C. K., Weissman, I. L. & Kim, B. Y. S. Breaking down the barriers to precision cancer nanomedicine. Trends Biotechnol. 35 , 159–171 (2017).

Chen, D., Parayath, N., Ganesh, S., Wang, W. & Amiji, M. The role of apolipoprotein- and vitronectin-enriched protein corona on lipid nanoparticles for vivo targeted delivery and transfection of oligonucleotides in murine tumor models. Nanoscale 11 , 18806–18824 (2019). This study analyses the impact of NP make-up on protein corona and resulting tumour-specific delivery, showing that PEG content can influence outcome and highlighting the ability of NP design features to dictate biodistribution .

Fornaguera, C. et al. In vivo retargeting of poly(β-aminoester) (OM-PBAE) nanoparticles is influenced by protein corona. Adv. Healthc. Mater. 8 , 1–11 (2019).

Bertrand, N. et al. Mechanistic understanding of in vivo protein corona formation on polymeric nanoparticles and impact on pharmacokinetics. Nat. Commun. 8 , 777 (2017).

Yeo, E. L. L., Cheah, J. U. J., Thong, P. S. P., Soo, K. C. & Kah, J. C. Y. Gold nanorods coated with apolipoprotein E protein corona for drug delivery. ACS Appl. Nano Mater. 2 , 6220–6229 (2019).

Dal Magro, R. et al. ApoE-modified solid lipid nanoparticles: a feasible strategy to cross the blood–brain barrier. J. Control. Rel. 249 , 103–110 (2017).

Dong, Y. et al. Lipopeptide nanoparticles for potent and selective siRNA delivery in rodents and nonhuman primates. Proc. Natl Acad. Sci. USA 111 , 3955–3960 (2014).

Key, J. et al. Soft discoidal polymeric nanoconstructs resist macrophage uptake and enhance vascular targeting in tumors. ACS Nano 9 , 11628–11641 (2015).

Hui, Y. et al. Role of nanoparticle mechanical properties in cancer drug delivery. ACS Nano 13 , 7410–7424 (2019).

Nizzero, S., Ziemys, A. & Ferrari, M. Transport barriers and oncophysics in cancer treatment. Trends Cancer 4 , 277–280 (2018).

Kumar, S., Anselmo, A. C., Banerjee, A., Zakrewsky, M. & Mitragotri, S. Shape and size-dependent immune response to antigen-carrying nanoparticles. J. Control. Rel. 220 , 141–148 (2015).

Xie, X., Liao, J., Shao, X., Li, Q. & Lin, Y. The effect of shape on cellular uptake of gold nanoparticles in the forms of stars, rods, and triangles. Sci. Rep. 7 , 1–9 (2017).

Wang, W., Gaus, K., Tilley, R. D. & Gooding, J. J. The impact of nanoparticle shape on cellular internalisation and transport: what do the different analysis methods tell us? Mater. Horiz. 6 , 1538–1547 (2019).

Garapaty, A. & Champion, J. A. Shape of ligand immobilized particles dominates and amplifies the macrophage cytokine response to ligands. PLoS One 14 , 12–14 (2019).

Ganson, N. J. et al. Pre-existing anti-polyethylene glycol antibody linked to first-exposure allergic reactions to pegnivacogin, a PEGylated RNA aptamer. J. Allergy Clin. Immunol. 137 , 1610–1613.e7 (2016).

Povsic, T. J. et al. Pre-existing anti-PEG antibodies are associated with severe immediate allergic reactions to pegnivacogin, a PEGylated aptamer. J. Allergy Clin. Immunol. 138 , 1712–1715 (2016).

Zhong, Q., Merkel, O. M., Reineke, J. J. & da Rocha, S. R. P. Effect of the route of administration and PEGylation of poly(amidoamine) dendrimers on their systemic and lung cellular biodistribution. Mol. Pharm. 13 , 1866–1878 (2016).

Battaglia, L. et al. Lipid nanoparticles for intranasal administration: application to nose-to-brain delivery. Expert Opin. Drug Deliv. 15 , 369–378 (2018).

Dölen, Y. et al. Nanovaccine administration route is critical to obtain pertinent iNKt cell help for robust anti-tumor T and B cell responses. Oncoimmunology 9 , 1738813 (2020).

McLennan, D. N., Porter, C. J. H. & Charman, S. A. Subcutaneous drug delivery and the role of the lymphatics. Drug Discov. Today Technol. 2 , 89–96 (2005).

Dong, W. et al. Comparative study of mucoadhesive and mucus-penetrative nanoparticles based on phospholipid complex to overcome the mucus barrier for inhaled delivery of baicalein. Acta Pharm. Sin. B 10 , 1576–1585 (2019).

Cone, R. A. Barrier properties of mucus. Adv. Drug Deliv. Rev. 61 , 75–85 (2009).

Wu, L. et al. Quantitative comparison of three widely-used pulmonary administration methods in vivo with radiolabeled inhalable nanoparticles. Eur. J. Pharm. Biopharm. 152 , 108–115 (2020). This study investigates three widely used routes of pulmonary administration in mouse models, revealing that the administration route affects polymeric NP biodistribution and highlighting another means of heterogenizing NP delivery .

Pant, K. et al. Active targeting of dendritic polyglycerols for diagnostic cancer imaging. Small 16 , 1905013. (2019).

Yi, Y. et al. Glucose-linked sub-50-nm unimer polyion complex-assembled gold nanoparticles for targeted siRNA delivery to glucose transporter 1-overexpressing breast cancer stem-like cells. J. Control. Release 295 , 268–277 (2019).

Yan, Y. et al. Dually active targeting nanomedicines based on a direct conjugate of two purely natural ligands for potent chemotherapy of ovarian tumors. ACS Appl. Mater. Interfaces 11 , 46548–46557 (2019).

Ahmad, A., Khan, F., Mishra, R. K. & Khan, R. Precision cancer nanotherapy: evolving role of multifunctional nanoparticles for cancer active targeting. J. Med. Chem. 62 , 10475–10496 (2019).

Du, Y., Liu, X., Liang, Q., Liang, X. J. & Tian, J. Optimization and design of magnetic ferrite nanoparticles with uniform tumor distribution for highly sensitive MRI/MPI performance and improved magnetic hyperthermia therapy. Nano Lett. 19 , 3618–3626 (2019).

Hong, S. T. et al. Improving the anticancer effect of afatinib and microRNA by using lipid polymeric nanoparticles conjugated with dual pH-responsive and targeting peptides. J. Nanobiotechnol. 17 , 89 (2019).

Saraiva, C. et al. Nanoparticle-mediated brain drug delivery: overcoming blood–brain barrier to treat neurodegenerative diseases. J. Control. Release 235 , 34–47 (2016).

Zhou, Q. et al. Enzyme-activatable polymer–drug conjugate augments tumour penetration and treatment efficacy. Nat. Nanotechnol. 14 , 799–809 (2019).

Sindhwani, S. et al. The entry of nanoparticles into solid tumours. Nat. Mater. 19 , 566–575 (2020). This groundbreaking study shows that inter-endothelial gaps are not responsible for NP entry into solid tumours as was previously thought, but instead finds that up to 97% of NPs actively enter tumours through endothelial cells .

Johnsen, K. B. et al. Modulating the antibody density changes the uptake and transport at the blood–brain barrier of both transferrin receptor-targeted gold nanoparticles and liposomal cargo. J. Control. Release 295 , 237–249 (2019).

Sousa, F., Dhaliwal, H. K., Gattacceca, F., Sarmento, B. & Amiji, M. M. Enhanced anti-angiogenic effects of bevacizumab in glioblastoma treatment upon intranasal administration in polymeric nanoparticles. J. Control. Release 309 , 37–47 (2019).

Musumeci, T., Bonaccorso, A. & Puglisi, G. Epilepsy disease and nose-to-brain delivery of polymeric nanoparticles: an overview. Pharmaceutics 11 , 118 (2019).

Bruinsmann, F. A. et al. Nasal drug delivery of anticancer drugs for the treatment of glioblastoma: preclinical and clinical trials. Molecules 24 , 4312 (2019).

Lamson, N. G., Berger, A., Fein, K. C. & Whitehead, K. A. Anionic nanoparticles enable the oral delivery of proteins by enhancing intestinal permeability. Nat. Biomed. Eng. 4 , 84–96 (2020). This work screens inorganic NPs for the oral delivery of proteins and finds the properties needed to enhance intestinal permeation via opening tight junctions, which include negative charge and small size .

Banerjee, A., Qi, J., Gogoi, R., Wong, J. & Mitragotri, S. Role of nanoparticle size, shape and surface chemistry in oral drug delivery. J. Control. Release 238 , 176–185 (2016).

Yong, J. M., Mantaj, J., Cheng, Y. & Vllasaliu, D. Delivery of nanoparticles across the intestinal epithelium via the transferrin transport pathway. Pharmaceutics 11 , 298 (2019).

Zhang, S. et al. BSA nanoparticles modified with N -acetylcysteine for improving the stability and mucoadhesion of curcumin in the gastrointestinal tract. J. Agric. Food Chem. 67 , 9371–9381 (2019).

Zheng, N. et al. Mesoporous silica nanorods for improved oral drug absorption. Artif. Cells Nanomed. Biotechnol. 46 , 1132–1140 (2018).

Li, D. et al. Influence of particle geometry on gastrointestinal transit and absorption following oral administration. ACS Appl. Mater. Interfaces 9 , 42492–42502 (2017).

Zhuang, J. et al. The influence of nanoparticle shape on bilateral exocytosis from Caco-2 cells. Chin. Chem. Lett. 29 , 1815–1818 (2018).

Berardi, A. & Baldelli Bombelli, F. Oral delivery of nanoparticles — let’s not forget about the protein corona. Expert Opin. Drug Deliv. 16 , 563–566 (2019).

PubMed   Google Scholar  

Durán-Lobato, M., Niu, Z. & Alonso, M. J. Oral delivery of biologics for precision medicine. Adv. Mater. 32 , e1901935 (2019).

Kruse, C. R. et al. The effect of pH on cell viability, cell migration, cell proliferation, wound closure, and wound reepithelialization: In vitro and in vivo study. Wound Repair Regen. 25 , 260–269 (2017).

Hathaway, H. et al. Thermally triggered release of the bacteriophage endolysin CHAPK and the bacteriocin lysostaphin for the control of methicillin resistant Staphylococcus aureus (MRSA). J. Control. Release 245 , 108–115 (2017).

Holme, M. N. et al. Shear-stress sensitive lenticular vesicles for targeted drug delivery. Nat. Nanotechnol. 7 , 536–543 (2012).

Sykes, E. A. et al. Tailoring nanoparticle designs to target cancer based on tumor pathophysiology. Proc. Natl Acad. Sci. USA 113 , E1142–E1151 (2016).

Zan, Y., Dai, Z., Liang, L., Deng, Y. & Dong, L. Co-delivery of plantamajoside and sorafenib by a multi-functional nanoparticle to combat the drug resistance of hepatocellular carcinoma through reprograming the tumor hypoxic microenvironment. Drug Deliv. 26 , 1080–1091 (2019). This work focuses on a NP system that can reprogramme the tumour microenvironment, working to reverse drug resistance of tumour cells and resensitize patients to existing first-line therapies .

Dai, Q. et al. Quantifying the ligand-coated nanoparticle delivery to cancer cells in solid tumors. ACS Nano 12 , 8423–8435 (2018).

Witten, J. & Ribbeck, K. The particle in the spider’s web: transport through biological hydrogels. Nanoscale 9 , 8080–8095 (2017).

Nagel, G., Sousa-Herves, A., Wedepohl, S. & Calderón, M. Matrix metalloproteinase-sensitive multistage nanogels promote drug transport in 3D tumor model. Theranostics 10 , 91–108 (2020).

Price, L. S. L., Stern, S. T., Deal, A. M., Kabanov, A. V. & Zamboni, W. C. A reanalysis of nanoparticle tumor delivery using classical pharmacokinetic metrics. Sci. Adv. 6 , eaay9249 (2020).

Lee, H. et al. 64 Cu-MM-302 positron emission tomography quantifies variability of enhanced permeability and retention of nanoparticles in relation to treatment response in patients with metastatic breast cancer. Clin. Cancer Res. 23 , 4190–4202 (2017).

Ding, J. et al. Engineered nanomedicines with enhanced tumor penetration. Nano Today 29 , 100800 (2019).

Wadajkar, A. S. et al. Decreased non-specific adhesivity, receptor targeted (DART) nanoparticles exhibit improved dispersion, cellular uptake, and tumor retention in invasive gliomas. J. Control. Release 267 , 144–153 (2017).

Nance, E. A. et al. A dense poly(ethylene glycol) coating improves penetration of large polymeric nanoparticles within brain tissue. Sci. Transl. Med. 4 , 149ra119 (2012).

Stephen, Z. R. et al. Time-resolved MRI assessment of convection-enhanced delivery by targeted and nontargeted nanoparticles in a human glioblastoma mouse model. Cancer Res. 79 , 4776–4786 (2019).

Fahy, J. V. & Dickey, B. F. Airway mucus function and dysfunction. N. Engl. J. Med. 363 , 2233–2247 (2010).

Livraghi-Butrico, A. et al. Contribution of mucus concentration and secreted mucins Muc5ac and Muc5b to the pathogenesis of muco-obstructive lung disease. Mucosal Immunol. 10 , 395–407 (2017).

Ho, L. W. C., Liu, Y., Han, R., Bai, Q. & Choi, C. H. J. Nano-cell interactions of non-cationic bionanomaterials. Acc. Chem. Res. 52 , 1519–1530 (2019).

Behzadi, S. et al. Cellular uptake of nanoparticles: journey inside the cell. Chem. Soc. Rev. 46 , 4218–4244 (2017).

Trimble, W. S. & Grinstein, S. Barriers to the free diffusion of proteins and lipids in the plasma membrane. J. Cell Biol. 208 , 259–271 (2015).

Foroozandeh, P. & Aziz, A. A. Insight into cellular uptake and intracellular trafficking of nanoparticles. Nanoscale Res. Lett. 13 , 339 (2018).

Zhang, L. et al. Shape effect of nanoparticles on tumor penetration in monolayers versus spheroids. Mol. Pharm. 16 , 2902–2911 (2019).

Lunnoo, T., Assawakhajornsak, J. & Puangmali, T. In silico study of gold nanoparticle uptake into a mammalian cell: interplay of size, shape, surface charge, and aggregation. J. Phys. Chem. C. 123 , 3801–3810 (2019).

Zhang, L. et al. Tumor chemo-radiotherapy with rod-shaped and spherical gold nano probes: shape and active targeting both matter. Theranostics 9 , 1893–1908 (2019).

Dzuricky, M., Xiong, S., Weber, P. & Chilkoti, A. Avidity and cell uptake of integrin-targeting polypeptide micelles is strongly shape-dependent. Nano Lett. 19 , 6124–6132 (2019).

Shang, L., Nienhaus, K. & Nienhaus, G. U. Engineered nanoparticles interacting with cells: size matters. J. Nanobiotechnol. 12 , 5 (2014).

Zhang, L., Chen, H., Xie, J., Becton, M. & Wang, X. Interplay of nanoparticle rigidity and its translocation ability through cell membrane. J. Phys. Chem. B 123 , 8923–8930 (2019).

Hocking, K. M. et al. Nanotechnology enabled modulation of signaling pathways affects physiologic responses in intact vascular tissue. Tissue Eng. Part A 25 , 416–426 (2019).

Shahin, M., Safaei-Nikouei, N. & Lavasanifar, A. Polymeric micelles for pH-responsive delivery of cisplatin. J. Drug Target 22 , 629–637 (2014).

Chen, L. Q. et al. Nanotoxicity of silver nanoparticles to red blood cells: size dependent adsorption, uptake, and hemolytic activity. Chem. Res. Toxicol. 28 , 501–509 (2015).

Yue, J., Feliciano, T. J., Li, W., Lee, A. & Odom, T. W. Gold nanoparticle size and shape effects on cellular uptake and intracellular distribution of siRNA nanoconstructs. Bioconjug. Chem. 28 , 1791–1800 (2017).

Chen, G. et al. A biodegradable nanocapsule delivers a Cas9 ribonucleoprotein complex for in vivo genome editing. Nat. Nanotechnol. 14 , 974–980 (2019).

Miao, L. et al. Delivery of mRNA vaccines with heterocyclic lipids increases anti-tumor efficacy by STING-mediated immune cell activation. Nat. Biotechnol. 37 , 1174–1185 (2019).

Maugeri, M. et al. Linkage between endosomal escape of LNP-mRNA and loading into EVs for transport to other cells. Nat. Commun. 10 , 4333 (2019).

Almeida, A. P. B. et al. Mucopenetrating lipoplexes modified with PEG and hyaluronic acid for CD44-targeted local siRNA delivery to the lungs. J. Biomater. Appl. 34 , 617–630 (2019).

Whitehead, K. A. et al. Degradable lipid nanoparticles with predictable in vivo siRNA delivery activity. Nat. Commun. 5 , 4277 (2014).

Oberli, M. A. et al. Lipid nanoparticle assisted mRNA delivery for potent cancer immunotherapy. Nano Lett. 17 , 1326–1335 (2017).

Kaczmarek, J. C. et al. Optimization of a degradable polymer–lipid nanoparticle for potent systemic delivery of mRNA to the lung endothelium and immune cells. Nano Lett. 18 , 6449–6454 (2018).

Wei, L., Zhao, Y., Hu, X. & Tang, L. Redox-responsive polycondensate neoepitope for enhanced personalized cancer vaccine. ACS Cent. Sci. 6 , 404–412 (2020).

Gao, Y. et al. pH/redox dual-responsive polyplex with effective endosomal escape for codelivery of siRNA and doxorubicin against drug-resistant cancer cells. ACS Appl. Mater. Interfaces 11 , 16296–16310 (2019).

Yuan, P. et al. Mitochondria-targeting, intracellular delivery of native proteins using biodegradable silica nanoparticles. Angew. Chem. 131 , 7739–7743 (2019).

Yamada, Y., Fukuda, Y., Sasaki, D., Maruyama, M. & Harashima, H. Development of a nanoparticle that releases nucleic acids in response to a mitochondrial environment. Mitochondrion 52 , 67–74 (2020).

Foroozandeh, P., Aziz, A. A. & Mahmoudi, M. Effect of cell age on uptake and toxicity of nanoparticles: the overlooked factor at the nanobio interface. ACS Appl. Mater. Interfaces 11 , 39672–39687 (2019).

Serpooshan, V. et al. Effect of cell sex on uptake of nanoparticles: the overlooked factor at the nanobio interface. ACS Nano 12 , 2253–2266 (2018).

Dong, Y. et al. Targeted blocking of miR328 lysosomal degradation with alkalized exosomes sensitizes the chronic leukemia cells to imatinib. Appl. Microbiol. Biotechnol. 103 , 9569–9582 (2019).

Yang, X. et al. Delivery of platinum (II) drugs with bulky ligands in trans-geometry for overcoming cisplatin drug resistance. Mater. Sci. Eng. C. 96 , 96–104 (2019).

Murakami, M. et al. Improving drug potency and efficacy by nanocarrier-mediated subcellular targeting. Sci. Transl. Med. 3 , 64ra2 (2011).

American Cancer Society. Cancer facts & figures 2019 (ACS, 2019).

Dalby, M., Cree, I. A., Challoner, B. R., Ghosh, S. & Thurston, D. E. The precision medicine approach to cancer therapy: part 1 — solid tumours. Pharm. J. https://doi.org/10.1211/PJ.2019.20207119 (2019).

Article   Google Scholar  

Hauschild, A. et al. Long-term outcomes in patients with BRAF V600-mutant metastatic melanoma receiving dabrafenib monotherapy: analysis from phase 2 and 3 clinical trials. Eur. J. Cancer 125 , 114–120 (2020).

Chapman, P. B. et al. Improved survival with vemurafenib in melanoma with BRAF V600E mutation. N. Engl. J. Med. 364 , 2507–2516 (2011).

Ou, W. et al. Regulatory T cell-targeted hybrid nanoparticles combined with immuno-checkpoint blockage for cancer immunotherapy. J. Control. Rel. 281 , 84–96 (2018).

Lyon, R. P. et al. Self-hydrolyzing maleimides improve the stability and pharmacological properties of antibody–drug conjugates. Nat. Biotechnol. 32 , 1059–1062 (2014).

Roma-Rodrigues, C., Mendes, R., Baptista, P. V. & Fernandes, A. R. Targeting tumor microenvironment for cancer therapy. Int. J. Mol. Sci. 20 , 840 (2019).

Rao, L. et al. Cancer cell membrane-coated nanoparticles for personalized therapy in patient-derived xenograft models. Adv. Funct. Mater. 29 , 1905671 (2019).

He, Z., Zhang, Y. & Feng, N. Cell membrane-coated nanosized active targeted drug delivery systems homing to tumor cells: a review. Mater. Sci. Eng. C. 106 , 110298 (2020).

Wang, Y., Luan, Z., Zhao, C., Bai, C. & Yang, K. Target delivery selective CSF-1R inhibitor to tumor-associated macrophages via erythrocyte-cancer cell hybrid membrane camouflaged pH-responsive copolymer micelle for cancer immunotherapy. Eur. J. Pharm. Sci. 142 , 105136 (2020).

Li, R., He, Y., Zhang, S., Qin, J. & Wang, J. Cell membrane-based nanoparticles: a new biomimetic platform for tumor diagnosis and treatment. Acta Pharm. Sin. B 8 , 14–22 (2018).

Karlsson, J. et al. Engineered nanoparticles for systemic siRNA delivery to malignant brain tumours. Nanoscale 11 , 20045–20057 (2019).

Ramchandani, D. et al. Nanoparticle delivery of miR-708 mimetic impairs breast cancer metastasis. Mol. Cancer Ther. 18 , 579–591 (2019).

Zhang, Y. R. et al. Strategies to improve tumor penetration of nanomedicines through nanoparticle design. Wiley Interdiscip. Rev. Nanomed. Nanobiotechnol. 11 , 1–12 (2019).

Xu, C. et al. Biodegradable nanotheranostics with hyperthermia-induced bubble ability for ultrasound imaging-guided chemo-photothermal therapy. Int. J. Nanomed. 14 , 7141–7153 (2019).

Li, H. J. et al. Stimuli-responsive clustered nanoparticles for improved tumor penetration and therapeutic efficacy. Proc. Natl Acad. Sci. USA 113 , 4164–4169 (2016).

Wang, D. et al. Engineering nanoparticles to locally activate T cells in the tumor microenvironment. Sci. Immunol. 4 , 1–14 (2019).

Kim, C. S., Duncan, B., Creran, B. & Rotello, V. M. Triggered nanoparticles as therapeutics. Nano Today 8 , 439–447 (2013).

Wong, C. et al. Multistage nanoparticle delivery system for deep penetration into tumor tissue. Proc. Natl Acad. Sci. USA 108 , 2426–2431 (2011).

Lim, W. Q., Phua, S. Z. F. & Zhao, Y. Redox-responsive polymeric nanocomplex for delivery of cytotoxic protein and chemotherapeutics. ACS Appl. Mater. Interfaces 11 , 31638–31648 (2019).

Mannaris, C. et al. Acoustically responsive polydopamine nanodroplets: a novel theranostic agent. Ultrason. Sonochem. 60 , 104782 (2020).

Riley, R. S. et al. Evaluating the mechanisms of light-triggered siRNA release from nanoshells for temporal control over gene regulation. Nano Lett. 18 , 3565–3570 (2018).

Dariva, C. G., Coelho, J. F. J. & Serra, A. C. Near infrared light-triggered nanoparticles using singlet oxygen photocleavage for drug delivery systems. J. Control. Rel. 294 , 337–354 (2019).

Shi, H. et al. Programmed co-delivery of platinum nanodrugs and gemcitabine by a clustered nanocarrier for precision chemotherapy for NSCLC tumors. J. Mater. Chem. B 8 , 332–342 (2020).

Li, H. J. et al. Smart superstructures with ultrahigh pH-sensitivity for targeting acidic tumor microenvironment: instantaneous size switching and improved tumor penetration. ACS Nano 10 , 6753–6761 (2016). This study demonstrates the usage of pH-triggered particle size-switching in the tumour microenvironment to facilitate NP extravasation and accumulation .

Chen, W. H., Luo, G. F. & Zhang, X. Z. Recent advances in subcellular targeted cancer therapy based on functional materials. Adv. Mater. 31 , 1–39 (2019).

Goswami, U. et al. Transferrin–copper nanocluster–doxorubicin nanoparticles as targeted theranostic cancer nanodrug. ACS Appl. Mater. Interfaces 10 , 3282–3294 (2018).

Qiao, Y. et al. Stimuli-responsive nanotherapeutics for precision drug delivery and cancer therapy. Wiley Interdiscip. Rev. Nanomed. Nanobiotechnol. 11 , 1–20 (2019).

Jin, Q., Deng, Y., Chen, X. & Ji, J. Rational design of cancer nanomedicine for simultaneous stealth surface and enhanced cellular uptake. ACS Nano 13 , 954–977 (2019).

Fang, Y. et al. Cleavable PEGylation: a strategy for overcoming the “PEG dilemma” in efficient drug delivery. Drug Deliv. 24 , 22–32 (2017).

Guo, F. et al. Matrix metalloprotein-triggered, cell penetrating peptide-modified star-shaped nanoparticles for tumor targeting and cancer therapy. J. Nanobiotechnol. 18 , 48 (2020).

Eno, J. Immunotherapy through the years. J. Adv. Pract. Oncol. 8 , 747–753 (2017).

Marshall, H. T. & Djamgoz, M. B. A. Immuno-oncology: emerging targets and combination therapies. Front. Oncol. 8 , 1–29 (2018).

Urbanavicius, D., Alvarez, T., Such, G. K., Johnston, A. P. R. & Mintern, J. D. The potential of nanoparticle vaccines as a treatment for cancer. Mol. Immunol. 98 , 2–7 (2018).

Truex, N. L. et al. Automated flow synthesis of tumor neoantigen peptides for personalized immunotherapy. Sci. Rep. 10 , 723 (2020).

Guo, Y., Lei, K. & Tang, L. Neoantigen vaccine delivery for personalized anticancer immunotherapy. Front. Immunol. 9 , 1499 (2018).

Dombroski, J. A., Jyotsana, N., Crews, D. W., Zhang, Z. & King, M. R. Fabrication and characterization of tumor nano-lysate as a preventative vaccine for breast cancer. Langmuir 36 , 6531–6539 (2020).

Gu, P. et al. Rational design of PLGA nanoparticle vaccine delivery systems to improve immune responses. Mol. Pharm. 16 , 5000–5012 (2019).

Firdessa-Fite, R. & Creusot, R. J. Nanoparticles versus dendritic cells as vehicles to deliver mRNA encoding multiple epitopes for immunotherapy. Mol. Ther. Methods Clin. Dev. 16 , 50–62 (2019).

Soema, P. C., Willems, G. J., Jiskoot, W., Amorij, J. P. & Kersten, G. F. Predicting the influence of liposomal lipid composition on liposome size, zeta potential and liposome-induced dendritic cell maturation using a design of experiments approach. Eur. J. Pharm. Biopharm. 94 , 427–435 (2015).

Cao, F., Yan, M., Liu, Y., Liu, L. & Ma, G. Photothermally controlled MHC class I restricted CD8+ T-cell responses elicited by hyaluronic acid decorated gold nanoparticles as a vaccine for cancer immunotherapy. Adv. Healthc. Mater. 7 , 1–12 (2018).

An, M., Li, M., Xi, J. & Liu, H. Silica nanoparticle as a lymph node targeting platform for vaccine delivery. ACS Appl. Mater. Interfaces 9 , 23466–23475 (2017).

Fontana, F., Liu, D., Hirvonen, J. & Santos, H. A. Delivery of therapeutics with nanoparticles: what’s new in cancer immunotherapy? Wiley Interdiscip. Rev. Nanomed. Nanobiotechnol. 9 , 1–26 (2017).

Tran, T. H. et al. Nanoparticles for dendritic cell-based immunotherapy. Int. J. Pharm. 542 , 253–265 (2018).

Grippin, A. J., Sayour, E. J. & Mitchell, D. A. Translational nanoparticle engineering for cancer vaccines. Oncoimmunology 6 , e1290036 (2017).

Liu, J., Miao, L., Sui, J., Hao, Y. & Huang, G. Nanoparticle cancer vaccines: design considerations and recent advances. Asian J. Pharm. Sci. 15 , 576–590 (2019).

Tostanoski, L. H., Gosselin, E. A. & Jewell, C. M. Engineering tolerance using biomaterials to target and control antigen presenting cells. Discov. Med. 21 , 403–410 (2016).

Zupan&#x010D;i&#x010D;, E. et al. Rational design of nanoparticles towards targeting antigen-presenting cells and improved T cell priming. J. Control. Rel. 258 , 182–195 (2017).

Macri, C., Dumont, C., Johnston, A. P. R. & Mintern, J. D. Targeting dendritic cells: a promising strategy to improve vaccine effectiveness. Clin. Transl. Immunol. 5 , e66–e68 (2016).

Zhang, F. et al. Genetic programming of macrophages to perform anti-tumor functions using targeted mRNA nanocarriers. Nat. Commun. 10 , 3974 (2019).

Li, K. et al. Polarization of tumor-associated macrophage phenotype via porous hollow iron nanoparticles for tumor immunotherapy in vivo. Nanoscale 12 , 130–144 (2020).

Chavez-Santoscoy, A. V. et al. Tailoring the immune response by targeting C-type lectin receptors on alveolar macrophages using ‘pathogen-like’ amphiphilic polyanhydride nanoparticles. Biomaterials 33 , 4762–4772 (2012).

Zhang, C. et al. Targeted antigen delivery to dendritic cell via functionalized alginate nanoparticles for cancer immunotherapy. J. Control. Release 256 , 170–181 (2017).

Chen, W. C. et al. Antigen delivery to macrophages using liposomal nanoparticles targeting sialoadhesin/CD169. PLoS One 7 , 1–9 (2012).

Tatar, A. S. et al. CD19-targeted, Raman tagged gold nanourchins as theranostic agents against acute lymphoblastic leukemia. Colloids Surf. B Biointerfaces 184 , 110478 (2019).

Qian, Y. et al. Targeting dendritic cells in lymph node with an antigen peptide-based nanovaccine for cancer immunotherapy. Biomaterials 98 , 171–183 (2016).

Pearson, R. M., Casey, L. M., Hughes, K. R., Miller, S. D. & Shea, L. D. In vivo reprogramming of immune cells: technologies for induction of antigen-specific tolerance. Adv. Drug Deliv. Rev. 114 , 240–255 (2017).

Singh, A. & Peppas, N. A. Hydrogels and scaffolds for immunomodulation. Adv. Mater. 26 , 6530–6541 (2014).

Marloye, M., Lawler, S. E. & Berger, G. Current patent and clinical status of stimulator of interferon genes (STING) agonists for cancer immunotherapy. Pharm. Pat. Anal. 8 , 87–90 (2019).

Liu, Y. et al. An inhalable nanoparticulate STING agonist synergizes with radiotherapy to confer long-term control of lung metastases. Nat. Commun. 10 , 1–15 (2019).

Cheng, N. et al. A nanoparticle-incorporated STING activator enhances antitumor immunity in PD-L1-insensitive models of triple-negative breast cancer. JCI Insight 3 , 1–20 (2018).

Wilson, D. R. et al. Biodegradable STING agonist nanoparticles for enhanced cancer immunotherapy. Nanomed. Nanotechnol. Biol. Med. 14 , 237–246 (2018).

Schmid, D. et al. T cell-targeting nanoparticles focus delivery of immunotherapy to improve antitumor immunity. Nat. Commun. 8 , 1–11 (2017).

Wang, C., Ye, Y., Hochu, G. M., Sadeghifar, H. & Gu, Z. Enhanced cancer immunotherapy by microneedle patch-assisted delivery of anti-PD1 antibody. Nano Lett. 16 , 2334–2340 (2016).

Wang, C., Sun, W., Wright, G., Wang, A. Z. & Gu, Z. Inflammation-triggered cancer immunotherapy by programmed delivery of CpG and Anti-PD1 antibody. Adv. Mater. 28 , 8912–8920 (2016).

Deng, H. & Zhang, Z. The application of nanotechnology in immune checkpoint blockade for cancer treatment. J. Control. Rel. 290 , 28–45 (2018).

Dunn, Z. S., Mac, J. & Wang, P. T cell immunotherapy enhanced by designer biomaterials. Biomaterials 217 , 119265 (2019).

Hickey, J. W., Vicente, F. P., Howard, G. P., Mao, H. Q. & Schneck, J. P. Biologically inspired design of nanoparticle artificial antigen-presenting cells for immunomodulation. Nano Lett. 17 , 7045–7054 (2017).

Sunshine, J. C., Perica, K., Schneck, J. P. & Green, J. J. Particle shape dependence of CD8 + T cell activation by artificial antigen presenting cells. Biomaterials 35 , 269–277 (2014).

Billingsley, M. et al. Ionizable lipid nanoparticle mediated mRNA delivery for human CAR T cell engineering. Nano Lett. 20 , 1578–1589 (2020).

Smith, T. T. et al. In situ programming of leukaemia-specific T cells using synthetic DNA nanocarriers. Nat. Nanotechnol. 12 , 813–822 (2017).

Pozsgay, J., Szekanecz, Z. & Sármay, G. Antigen-specific immunotherapies in rheumatic diseases. Nat. Rev. Rheumatol. 13 , 525–537 (2017).

Jung, H. H. et al. Polymeric nanoparticles containing both antigen and vitamin D3 induce antigen-specific immune suppression. Immune Netw. 19 , 1–12 (2019).

Pang, L., Macauley, M. S., Arlian, B. M., Nycholat, C. M. & Paulson, J. C. Encapsulating an immunosuppressant enhances tolerance induction by Siglec-engaging tolerogenic liposomes. ChemBioChem 18 , 1226–1233 (2017).

Horwitz, D. A., Bickerton, S., Koss, M., Fahmy, T. M. & La Cava, A. Suppression of murine lupus by CD4 + and CD8 + T reg cells induced by T cell–targeted nanoparticles loaded with interleukin-2 and transforming growth factor β. Arthritis Rheumatol. 71 , 632–640 (2019).

Hu, T. et al. pVAX1-A20 alleviates colitis in mice by promoting regulatory T cells. Dig. Liver Dis. 51 , 790–797 (2019).

Shahzad, K. A. et al. Design and optimization of PLGA particles to deliver immunomodulatory drugs for the prevention of skin allograft rejection. Immunol. Invest. 49 , 840–857 (2019).

Li, R. et al. Sustained release of immunosuppressant by nanoparticle-anchoring hydrogel scaffold improved the survival of transplanted stem cells and tissue regeneration. Theranostics 8 , 878–893 (2018).

Zhang, Y. et al. In situ repurposing of dendritic cells with CRISPR/Cas9-based nanomedicine to induce transplant tolerance. Biomaterials 217 , 119302 (2019).

Yin, H., Kauffman, K. J. & Anderson, D. G. Delivery technologies for genome editing. Nat. Rev. Drug Discov. 16 , 387–399 (2017).

Yin, H. et al. Non-viral vectors for gene-based therapy. Nat. Rev. Genet. 15 , 541–555 (2014).

Riley, R. S., June, C. H., Langer, R. & Mitchell, M. J. Delivery technologies for cancer immunotherapy. Nat. Rev. Drug Discov. 18 , 175–196 (2019).

Cavazzana-Calvo, M. et al. Gene therapy of human severe combined immunodeficiency (SCID)-X1 disease. Science 288 , 669–672 (2000).

Hacein-Bey-Abina, S. et al. LMO2-associated clonal T cell proliferation in two patients after gene therapy for SCID-X1. Science 302 , 1181–1185 (2003).

Whitehead, K. A., Langer, R. & Anderson, D. G. Knocking down barriers: advances in siRNA delivery. Nat. Rev. Drug Discov. 8 , 129–138 (2009).

Semple, S. C. et al. Rational design of cationic lipids for siRNA delivery. Nat. Biotechnol. 28 , 172–176 (2010).

Davis, M. E. et al. Evidence of RNAi in humans from systemically administered siRNA via targeted nanoparticles. Nature 464 , 1067–1070 (2010).

Mangraviti, A. et al. Polymeric nanoparticles for nonviral gene therapy extend brain tumor survival in vivo. ACS Nano 9 , 1236–1249 (2015).

Sarett, S. M. et al. Lipophilic siRNA targets albumin in situ and promotes bioavailability, tumor penetration, and carrier-free gene silencing. Proc. Natl Acad. Sci. USA 114 , E6490–E6497 (2017).

He, Y. et al. Polyethyleneimine/DNA polyplexes with reduction-sensitive hyaluronic acid derivatives shielding for targeted gene delivery. Biomaterials 34 , 1235–1245 (2013).

Xu, C. F. et al. Rational designs of in vivo CRISPR–Cas delivery systems. Adv. Drug Deliv. Rev. https://doi.org/10.1016/j.addr.2019.11.005 (2019).

Article   PubMed   Google Scholar  

Maule, G. et al. Allele specific repair of splicing mutations in cystic fibrosis through AsCas12a genome editing. Nat. Commun. 10 , 3556 (2019).

Moss, R. B. et al. Repeated aerosolized AAV-CFTR for treatment of cystic fibrosis: a randomized placebo-controlled phase 2B trial. Hum. Gene Ther. 18 , 726–732 (2007).

Alton, E. W. F. W. et al. Repeated nebulisation of non-viral CFTR gene therapy in patients with cystic fibrosis: a randomised, double-blind, placebo-controlled, phase 2b trial. Lancet Respir. Med. 3 , 684–691 (2015).

Guan, S. et al. Self-assembled peptide–poloxamine nanoparticles enable in vitro and in vivo genome restoration for cystic fibrosis. Nat. Nanotechnol. 14 , 287–297 (2019).

Robinson, E. et al. Lipid nanoparticle-delivered chemically modified mRNA restores chloride secretion in cystic fibrosis. Mol. Ther. 26 , 2034–2046 (2018).

Witten, J., Samad, T. & Ribbeck, K. Selective permeability of mucus barriers. Curr. Opin. Biotechnol. 52 , 124–133 (2018).

Patel, A. K. et al. Inhaled nanoformulated mRNA polyplexes for protein production in lung epithelium. Adv. Mater. 31 , 1805116 (2019).

Suk, J. S. et al. Lung gene therapy with highly compacted DNA nanoparticles that overcome the mucus barrier. J. Control. Rel. 178 , 8–17 (2014).

Deweerdt, S. The fix is in utero. Nature 564 , 6–8 (2018).

Massaro, G. et al. Fetal gene therapy for neurodegenerative disease of infants. Nat. Med. 24 , 1317–1323 (2018).

Ricciardi, A. S. et al. In utero nanoparticle delivery for site-specific genome editing. Nat. Commun. 9 , 1–11 (2018).

Caster, J. M., Patel, A. N., Zhang, T. & Wang, A. Investigational nanomedicines in 2016: a review of nanotherapeutics currently undergoing clinical trials. Wiley Interdiscip. Rev. Nanomed. Nanobiotechnol. 9 , 1416 (2017).

Wu, S. et al. Rapid label-free isolation of circulating tumor cells from patients’ peripheral blood using electrically charged Fe 3 O 4 nanoparticles. ACS Appl. Mater. Interfaces 12 , 4193–4203 (2020).

León-Janampa, N. et al. Synthesis, characterization and bio-functionalization of magnetic nanoparticles to improve the diagnosis of tuberculosis. Nanotechnology 31 , 175101 (2020).

Muhammad, P. et al. SERS-based nanostrategy for rapid anemia diagnosis. Nanoscale 12 , 1948–1957 (2020).

Download references

Acknowledgements

M.J.M. acknowledges support from a Burroughs Wellcome Fund Career Award at the Scientific Interface (CASI), a US National Institutes of Health (NIH) Director’s New Innovator Award (DP2 TR002776), a grant from the American Cancer Society (129784-IRG-16-188-38-IRG), the NIH (NCI R01 CA241661, NCI R37 CA244911 and NIDDK R01 DK123049), an Abramson Cancer Center (ACC)–School of Engineering and Applied Sciences (SEAS) Discovery Grant (P30 CA016520) and a 2018 American Association for Cancer Research (AACR)–Bayer Innovation and Discovery Grant (Grant Number 18-80-44-MITC). N.A.P. acknowledges support from the UT–Portugal Collaborative Research Program (CoLAB), the NIH (R01-EB022025-4 and R01-EB-00246-21), the National Science Foundation (Grant 1033746), the Pratt Foundation, the Cockrell Family Regents Chair and NSF Graduate Research Fellowships. R.M.H. was supported by a National Science Foundation (NSF) Graduate Research Fellowship (DGE 1845298). M.M.B. was supported by an NIH Training in HIV Pathogenesis T32 Program (T32 AI007632).

Author information

Authors and affiliations.

Department of Bioengineering, University of Pennsylvania, Philadelphia, PA, USA

Michael J. Mitchell, Margaret M. Billingsley & Rebecca M. Haley

Abramson Cancer Center, Perelman School of Medicine, University of Pennsylvania, Philadelphia, PA, USA

Michael J. Mitchell

Institute for Immunology, Perelman School of Medicine, University of Pennsylvania, Philadelphia, PA, USA

Cardiovascular Institute, Perelman School of Medicine, University of Pennsylvania, Philadelphia, PA, USA

Institute for Regenerative Medicine, Perelman School of Medicine, University of Pennsylvania, Philadelphia, PA, USA

Department of Biomedical Engineering, The University of Texas at Austin, Austin, TX, USA

Marissa E. Wechsler & Nicholas A. Peppas

Department of Chemical Engineering, The University of Texas at Austin, Austin, TX, USA

Nicholas A. Peppas

Department of Pediatrics, The University of Texas at Austin, Austin, TX, USA

Department of Surgery and Perioperative Care, The University of Texas at Austin, Austin, TX, USA

Department of Molecular Pharmaceutics and Drug Delivery, The University of Texas at Austin, Austin, TX, USA

Department of Chemical Engineering and Koch Institute for Integrative Cancer Research, Massachusetts Institute of Technology, Cambridge, MA, USA

Robert Langer

You can also search for this author in PubMed   Google Scholar

Contributions

M.M.B. and R.M.H. conducted the initial literature search and outlined the general manuscript format. M.M.B., R.M.H. and M.E.W. wrote the initial manuscript draft, with contributions from M.J.M., N.A.P. and R.L. All authors reviewed and critically revised previous versions of the manuscript. All authors read and approved the final manuscripts.

Corresponding authors

Correspondence to Michael J. Mitchell , Nicholas A. Peppas or Robert Langer .

Ethics declarations

Competing interests.

A list of entities with which R.L. is involved, compensated or uncompensated, can be found in the Supplementary information. M.J.M., M.M.B., R.M.H., M.E.W. and N.A.P. declare no competing interests.

Additional information

Publisher’s note.

Springer Nature remains neutral with regard to jurisdictional claims in published maps and institutional affiliations.

Supplementary information

A system of specialized cells that clear foreign bodies from blood circulation.

Having polyethylene glycol (PEG) polymer chains attached, typically to molecules or macrostructures.

A group of distinct plasma proteins that induce inflammation and aid in the clearance of foreign bodies or damaged cells by enhancing antibody and phagocytic cell activity.

The movement or leakage of something (cells, blood, nanoparticles and so on) from a blood vessel into the tissue around it.

Numerical comparisons of a nanoparticle’s height and width.

(BBB). A biological filter made of endothelial cells that restricts the movement of substances from the body to the brain.

(MPS). The phagocytic cell population of the immune system.

The metabolism of a drug within the liver and enterocytes before the drug reaches the systemic circulation.

The narrowing of a bodily passage (such as a blood vessel).

Rights and permissions

Reprints and permissions

About this article

Cite this article.

Mitchell, M.J., Billingsley, M.M., Haley, R.M. et al. Engineering precision nanoparticles for drug delivery. Nat Rev Drug Discov 20 , 101–124 (2021). https://doi.org/10.1038/s41573-020-0090-8

Download citation

Accepted : 16 October 2020

Published : 04 December 2020

Issue Date : February 2021

DOI : https://doi.org/10.1038/s41573-020-0090-8

Share this article

Anyone you share the following link with will be able to read this content:

Sorry, a shareable link is not currently available for this article.

Provided by the Springer Nature SharedIt content-sharing initiative

This article is cited by

Nanomaterial-encapsulated sting agonists for immune modulation in cancer therapy.

• Yuanliang Yan

Biomarker Research (2024)

Synergistic vesicle-vector systems for targeted delivery

• Christine Ardelle Marquez
• Ji-Young Ahn

Journal of Nanobiotechnology (2024)

Flash nanoprecipitation allows easy fabrication of pH-responsive acetalated dextran nanoparticles for intracellular release of payloads

• Krystal A. Hughes
• Bishal Misra
• Sharan Bobbala

Discover Nano (2024)

Intranasal administration enhances size-dependent pulmonary phagocytic uptake of poly(lactic-co-glycolic acid) nanoparticles

• Seung Ho Baek
• Eun-Ha Hwang
• Jung Joo Hong

EJNMMI Radiopharmacy and Chemistry (2024)

Development of hybrid biomicroparticles: cellulose exposing functionalized fusion proteins

• Joanna Żebrowska
• Piotr Mucha
• Piotr M. Skowron

Microbial Cell Factories (2024)

Quick links

• Explore articles by subject
• Guide to authors
• Editorial policies

Sign up for the Nature Briefing: Translational Research newsletter — top stories in biotechnology, drug discovery and pharma.

• Open access
• Published: 28 July 2021

Recent advances in transdermal drug delivery systems: a review

• Woo Yeup Jeong 1   na1 ,
• Mina Kwon 1   na1 ,
• Hye Eun Choi 1   na1 &
• Ki Su Kim   ORCID: orcid.org/0000-0002-6289-9467 1  

Biomaterials Research volume  25 , Article number:  24 ( 2021 ) Cite this article

63k Accesses

191 Citations

17 Altmetric

Metrics details

Various non-invasive administrations have recently emerged as an alternative to conventional needle injections. A transdermal drug delivery system (TDDS) represents the most attractive method among these because of its low rejection rate, excellent ease of administration, and superb convenience and persistence among patients. TDDS could be applicable in not only pharmaceuticals but also in the skin care industry, including cosmetics. Because this method mainly involves local administration, it can prevent local buildup in drug concentration and nonspecific delivery to tissues not targeted by the drug. However, the physicochemical properties of the skin translate to multiple obstacles and restrictions in transdermal delivery, with numerous investigations conducted to overcome these bottlenecks. In this review, we describe the different types of available TDDS methods, along with a critical discussion of the specific advantages and disadvantages, characterization methods, and potential of each method. Progress in research on these alternative methods has established the high efficiency inherent to TDDS, which is expected to find applications in a wide range of fields.

Introduction

Drug delivery system (DDS) is a generic term for a series of physicochemical technologies that can control delivery and release of pharmacologically active substances into cells, tissues and organs, such that these active substances could exert optimal effects [ 1 , 2 ]. In other words, DDS covers the routes of administration and drug formulations that efficiently deliver the drug to maximize therapeutic efficacy while minimizing any side effect [ 3 , 4 , 5 ]. Depending on the delivery route, there are many types of administration modalities, such as oral administration, transdermal administration, lung inhalation, mucosal administration, and intravenous injection. Among them, the transdermal drug delivery system (TDDS) represents an attractive approach.

TDDS has become one of the most widely investigated routes of noninvasive drug delivery into the body through the skin, unlike conventionally used direct administration routes that make use of needle-based injections. TDDS has significantly influenced the delivery of various therapeutic agents, especially in pain management, hormonal therapy, and treatment of diseases of the cardiovascular and central nervous systems [ 6 , 7 , 8 , 9 ]. TDDS does not involve passage through the gastrointestinal tract; therefore, there is no loss due to first-pass metabolism, and drugs can be delivered without interference from pH, enzymes, and intestinal bacteria. In addition, TDDS can be used to control drug release according to usage restrictions, thereby contributing to the high persistence of this method. Most importantly, because TDDS is a noninvasive administration method and involves minimal pain and burden on the patient, drugs can be safely and conveniently administered to children or the elderly [ 10 , 11 , 12 ].

However, it still does not utilize its full potential due to the innate skin barrier. The skin is the outermost organ with a multi-layered structure, and the role of the skin is to protect our body by blocking environmental hazards such as chemicals, heat, and toxins [ 13 , 14 ]. (Fig.  1 ). Such skin can be divided into the epidermis, which has the protective function, and the dermis, where blood vessels are located, and produces skin cells, and each layer has elements that interfere with transdermal delivery.

The structure of skin

First, the skin barrier effect of the epidermis occurs in the stratum corneum, the outermost layer, and is a property of blocking external substances. The barrier effect is very significant in the transport of substances having a large molecular weight. In TDDS, it is generally accepted that the delivery of substances with small molecular weights utilizes the intracellular pathway. However, for substances having a large molecular weight, methods and various mechanisms using the intracellular pathway in addition to the intercellular pathway are introduced and used [ 15 , 16 , 17 ]. This is due to the structure of the skin because the part called lipid containing both cells and hydrophilic substances and hydrophobic substances does not have a perfectly regular position but exists with regularity [ 18 ]. These structural features can be explained by the principles of physicochemical properties that are attempted to enhance drug delivery through the skin. Next, the vascular system in the dermal layer can inhibit transdermal delivery. A one-cell-thick layer of endothelial cells terminating in the papillary loops of the superficial arteriovenous plexus near the dermal-epidermal junction in the upper dermis represents the interface between the tissues surrounding the skin and the human vasculature. The role of the endothelium in the skin is like that of the whole body. It actively responds to pressure, shear, osmotic pressure, heat, chemokines, and cytokines by modulating permeability and inducing vasodilation or constriction [ 19 ]. Therefore, the biggest issue of TDDS is to resolve the barrier effect of the stratum corneum, deliver the drug to the skin tissue, and pass through the cellular and vascular tissue to reach the target tissue. The problem is that only a small amount of the drug can be delivered through the skin tissue [ 20 , 21 ].

To solve this problem, various novel TDDS techniques have been intensively developed and have emerged as attractive administration methods. In addition, such development could represent a competitive advantage over other drug administration methods in terms of the delivered dose, cost-effectiveness, and therapeutic efficacy [ 21 , 22 , 23 , 24 ].

Here, we review various transdermal drug delivery techniques (Table  1 ). We summarize the characteristics of active/passive transdermal delivery and characterization methods. In addition, we discuss future perspectives in the field of TDDS.

Enhancement of transdermal delivery by equipment (active delivery)

External stimuli, such as electrical, mechanical, or physical stimuli, are known to enhance skin permeability of drugs and biomolecules, as compared to the delivery of drugs by topical application on the skin [ 73 ]. TDDS supplemented by appropriate equipment is termed as active transdermal delivery, which is known to deliver drugs quickly and reliably into the skin. In addition, this mode of enhanced TDDS can accelerate the therapeutic efficacy of delivered drugs (Fig.  2 ) [ 75 , 76 , 77 ].

A Experimental set up for skin permeation test using iontophoresis. B In vitro drug release profiles of drug-loaded AuNP oleogels (d-AuNP) on skin. C Fluorescence spectroscopy images obtained from skin permeation experiment after 1 h of application. Arrows mark the top surface of the skin segment treated with d-AuNP. D A schematic illustration of sonophoresis-assisted transdermal drug delivery. E Penetration pathways of LaNO3 after treatment with low frequency sonophoresis, and TEM images of SC after treatment with low frequency sonophoresis using RuO4 fixation in the absence of low frequency sonophoresis (left) and after 5 min (middle) and 10 min (right) of treatment with low frequency sonophoresis. F Schematic Illustration Showing the Fabrication Process of the MTX-Loaded HA-Based Dissolving MN Patch. G Quantitative analysis of epidermal thickness. H Therapeutic effects of MTX-loaded MNs and oral administration of the same dose and a double dose of MTX on IMQ-induced psoriasis-like skin inflammation. Representative photographs of left ear lesions and skin sections stained with H&E and Ki67 on day 7. A , B , C Reproduced from [ 29 ], copyright permission by American Chemical Society 2020. D Reproduced from [ 74 ], copyright permission by Springer Nature 2021. E Reproduced from [ 40 ], copyright permission by Elsevier 2010. F , G , H Reproduced from [ 54 ], copyright permission by American Chemical Society 2019

Iontophoresis

Iontophoresis promotes the movement of ions across the membrane under the influence of a small externally applied potential difference (less than 0.5 mA/cm 2 ), which has been proven to enhance skin penetration and increase release rate of several drugs with poor absorption/permeation profiles. This technique has been utilized in the in vivo transport of ionic or nonionic drugs by the application of an electrochemical potential gradient [ 25 ]. The efficacy of iontophoresis depends on the polarity, valency, and mobility of the drug molecule, the nature of the applied electrical cycle, and the formulation containing the drug. In particular, the dependence on current makes drug absorption through iontophoresis less dependent on biological parameters, unlike most other drug delivery systems (Fig. 2 A, B) [ 26 ]. This modality could additionally include electronic means of reminding patients to change dosages, if desired, to increase patient compliance [ 27 , 28 ].

Sonophoresis

The desired range of ultrasound frequencies generated by an ultrasound device can improve transdermal drug delivery [ 30 , 31 ]. Low-frequency ultrasound is more effective, because it facilitates drug movement by creating an aqueous path in the perturbed bilayer through cavitation (Fig. 2 C) [ 32 ]. The drug under consideration is mixed with a specific coupler, such as a gel or a cream, which transmits ultrasonic waves to the skin and disturbs the skin layers, thereby creating an aqueous path through which the drug can be injected. Drugs typically pass through passages created by the application of ultrasonic waves with energy values between 20 kHz and 16 MHz. Ultrasound also increases the local temperature of the skin area and creates a thermal effect, which further promotes drug penetration. Several drugs of different classes have been delivered by this method regardless of their solubility, dissociation and ionization constants, and electrical properties (including hydrophilicity), such as mannitol and high molecular weight (MW) drugs such as insulin. However, the exact mechanism of drug penetration through this method is not yet completely understood, and problems with device availability, optimization of duration of exposure and treatment cycles for delivery, and undesirable side effects including burns persist.

Electroporation

This method uses the application of high voltage electric pulses ranging from 5 to 500 V for short exposure times (~ms) to the skin, which leads to the formation of small pores in the SC that improve permeability and aid drug diffusion [ 34 , 35 ]. For safe and painless drug administration, electric pulses are introduced using closely positioned electrodes. This is a very safe and painless procedure involving permeabilization of the skin and has been used to demonstrate the successful delivery of not only low MW drugs, such as doxorubicin, mannitol, or calcein, but also high MW ones such as antiangiogenic peptides, oligonucleotides, and the negatively charged anticoagulant heparin. However, this method has the disadvantages of small delivery loads, massive cellular perturbation sometimes including cell death, heating-induced drug damage, and denaturation of protein and other biomacromolecular therapeutics.

Photomechanical waves

Photodynamic waves transmitted to the skin can penetrate the SC, allowing the drug to pass through the transiently created channel [ 37 , 39 ]. The incident wave produces limited ablation, which is achieved by low radiation exposure of approximately 5–7 J/cm 2 to increase the depth to 50–400 μm for successful transmission. This limited ablation showed a longer increase and duration as compared to that in other direct ablation techniques, which made it necessary to control properties of the photodynamic waves to ensure delivery of the product to the intended depth in the skin. The wave generated by a single laser pulse also showed increased skin permeability within minutes, allowing macromolecules to diffuse into the skin. Dextran macromolecules of 40 kDa weight and 20 nm latex particles could be delivered by a single photodynamic laser pulse of a 23-ns duration.

Microneedle

The microneedle drug delivery system is a novel drug delivery system, in which drugs are delivered to the circulatory system through a needle [ 41 ]. This represents one of the most popular methods for transdermal drug delivery and is an active area of current research. This involves a system in which micron-sized needles pierce the superficial layer of the skin, resulting in drug diffusion across the epidermal layer. Because these microneedles are short and thin, these deliver drugs directly to the blood capillary area for active absorption, which helps in avoiding pain [ 42 ]. Scientists have attempted to use multiple techniques for appropriate optimization and geometric measurements required for effective insertion of microneedles into human skin, which also represents the broad objective of research on microneedles.

The fabrication of microneedle system has been widely investigated with considering the objective, drug type and dose, and targets for use [ 43 ]. Up to now, the microneedle can be fabricated with laser-mediated techniques and photolithography. The laser-mediated fabrication techniques are used for manufacturing metal or polymer microneedle. The 3D structure of a microneedle is generated through cutting or ablating on a flat metal/polymer surface using a laser [ 44 , 45 ]. Photolithography is known as the method of elaborately fabricating microneedle and has the advantage of being able to manufacture needles of various shapes using various materials. This method is mainly used to manufacture dissolving/hydrogel microneedles or silicon microneedles via making an inverse mold based on the microneedle structure through etching of photoresist [ 46 ]. In addition, 3D printing [ 47 ], Microstereolithography [ 48 ], and Two-photon polymerization [ 49 ] are also investigated for preparing various microneedle system.

The prepared microneedles could be of several types, such as solid microneedles that simply make a physical path through which drugs can be absorbed, drug-coated microneedles which facilitate delivery of drugs coated on the surfaces of the needles as the latter enter the skin, dissolving microneedles made of drug formulations that dissolve in the body, naturally delivered melting needles which involve drug storage in hollow needles followed by administration (such as a specific injection type), and microneedle patches combined with diverse patch types (Fig. 2 D, E) [ 50 , 51 , 52 , 53 , 54 ].

Thermal ablation

Thermal ablation, also known as thermophoresis, is a promising technique for selectively disrupting the stratum corneum structure by localized heat which provides enhanced drug delivery through microchannels created in the skin [ 55 ]. To ablate the stratum corneum by thermal ablation, a high temperature above 100 °C is required and this leads to heating and vaporization of keratin. Additionally, the degree of alteration of the stratum corneum structure is proportional to the locally elevated temperature, indicating that it is an ideal technique for precise control of drug delivery. The thermal exposure should be short within microseconds to create a high enough temperature gradient across the skin for selective ablation of the stratum corneum without damaging the viable epidermis. Micron-scale defects created from thermal ablation are small enough (50–100 μm in diameter) to avoid the potential to cause pain, bleeding, irritation, and infection. Therefore, the patient is well tolerated if there is no damage to the cells of the deeper tissues. In addition, thermal ablation has better control and reproducibility than other approaches such as mechanical abrasion, chemical treatment, or tape-stripping. And it offers effective delivery of small molecules as well as high molecular weight compounds. However, the structural changes in the skin must be evaluated, especially when using higher energy for enhancing the diffusion rate of drug molecules.

Thermal ablation can usually be induced by laser and radiofrequency methods depending on the different sources of thermal energy [ 56 , 57 ]. Laser thermal ablation methodologies utilize a laser to induce micropore structure of skin as well as the increase of the skin temperature which increases skin diffusivity. Laser light energy is absorbed by water and pigments of the skin and transforms to thermal energy leading to water excitation and explosive evaporation from the epidermis. The degree of the ablated skin depth can be precisely controlled upon tuning many parameters such as wavelength, pulse length, energy, number and repetition rate, tissue thickness, absorption coefficient, and duration time of laser exposure. Laser thermal ablation, especially when using Er:YAG laser, makes it possible to increase the penetration of drugs by more than 100 times and enhance the delivery of both lipophilic and hydrophilic drugs including biomacromolecules such as peptides, proteins, vaccines, and DNAs [ 56 , 57 , 58 ].

Radiofrequency thermal ablation involves the placement of an array of needle-like metallic microelectrodes directly onto skin and application of high frequency electric current into the skin in radiofrequency range (100–500 kHz) which produce micron-scale pathways in stratum corneum. Exposure of the skin to a high radiofrequency causes ionic vibrations within the tissue leading to the generation of localized heat in specific areas of the skin. And thus, induced heat cause water evaporation and ablates the cells of the epidermis under each filament creating microchannels up to 50 μm in depth. The process is completed within a few seconds and microchannels are filled with intestinal fluid through which hydrophilic molecules can permeate. The rate of drug delivery is proportional to the degree of the ablated skin depth which is controlled by the size and density of the microchannels. Radiofrequency thermal ablation can sustain the drug release and enhance the delivery of a wide range of drugs with hydrophilic nature including macromolecules using a low-cost, disposable device [ 59 ].

TDDS using chemical enhancers (passive delivery)

To achieve enhanced transdermal delivery and therapeutic efficacy, drugs should have low MW (less than 1 kDa), an affinity for lipophilic and hydrophilic phases, short half-life, and a lack of skin irritability [ 64 ]. Many factors affect drug penetration through the skin, such as species differences, skin age and site, skin temperature, state of the skin, area of application, duration of exposure, moisture content of the skin, pretreatment methods, and physical characteristics of the penetrant.

Recent studies that have focused on aspects of transdermal drug delivery technologies ranging from the development of chemical enhancers that increase the spread of drugs across the skin or increase the solubility of drugs in the skin to novel innovative approaches that extend this concept to the design of super-strong formulations, microemulsions, and vesicles [ 65 , 66 ] (Fig.  3 ). Penetration enhancers can be used alone or in combination with chemical penetration enhancers with proven superior skin penetration as compared to that of individual chemicals. These synergistic systems include eutectic mixtures and nanoparticle composite self-assembled vesicles. Therefore, research in recent years have focused on the application of suitable molecular simulation methodologies in understanding the skin lipid barrier, mechanisms regulating penetration of molecules across the skin and transport of penetration enhancers, and perturbations in the skin barrier function.

A SEM images of the prepared ALA-ES gels. B TEM images of ALA-ES in human HS tissue dermis. ALA-ES is indicated using black arrows. C TEM image of OA-UCNP. The nanoparticles show near-spherical shape with an average diameter around 25 nm. D Microscopy images of a section of a sample pig ear skin under 980 nm excitation laser. E Schematic illustration of W/O/W emulsification of HA-PLGA. F Fluorescence microscopic images of histological sections of rat skin at 4 and 12 h after topical application of Rho B-encapsulated HA-PLGA NPs. Scale bar, 100 μm. A , B Reproduced from [ 63 ], copyright permission by American Chemical Society 2018. C , D Reproduced from [ 67 ], copyright permission by IOP Publishing Ltd. 2020. E , F , G Reproduced from [ 68 ], copyright permission by BioMed Central Ltd. 2019

Vesicles are colloidal particles filled with water and consist of amphiphilic molecules in a bilayer arrangement. Under conditions of excess water, these amphiphilic molecules form concentric bilayers with one or more shells (multilayer vesicles). Vesicles can carry water-soluble and fat-soluble drugs to achieve transdermal absorption. When utilized for topical applications, vesicles can be used to achieve sustained release of stored drugs. It is also possible to employ vesicles in TDDS to control the absorption rate through a multilayered structure. Owing to the presence of different components, vesicle systems can be divided into several types, such as liposomes, transfersomes, and ethosomes, depending on the properties of the constituent substances [ 60 ].

Liposomes are circular soft vesicles formed by one or more bilayer membranes that separate an aqueous medium from another. Their main components are usually phospholipids, with or without cholesterol. Phospholipid molecules are mainly composed of different polar head groups and two hydrophobic hydrocarbon chains. Polar groups can be either positively or negatively charged. Hydrocarbon chain molecules have different lengths and different degrees of unsaturation. The formation of liposomes occurs spontaneously upon reconstitution of a dry lipid film in an aqueous solution. This unique structure allows liposomes to be both hydrophilic and hydrophobic and affords encapsulation of both water-soluble and fat-soluble substances. However, some studies have shown that liposomes can only remain on the surface of the skin and cannot pass through the granular layer of the epidermis, thereby minimizing the amount of drug absorbed into the blood circulation. This property increases the retention of drugs that stay on the skin, prolong their activity at the site of the lesion, and allow long-term sustained release. Therefore, liposomes are the preferred system of choice for the topical treatment of skin diseases [ 61 ].

Transfersomes are also called deformable liposomes, or elastic or highly flexible liposomes. The most important feature of these vesicles is the elasticity that results from the addition of single-chain surfactants. These surfactants make the phospholipid bilayer fluid and vesicles highly deformable, thereby rendering these into first-generation transfersomes. Over time, second-generation transfersomes have emerged, consisting of at least one basic bilayer building block (typically fluid-phase phosphatidylcholine lipids) and at least two or more polar lipophilic substances. Third-generation transfersomes are a combination of amphiphilic surfactants, with or without phospholipids. The possibility of deformation has facilitated the design of transfersomes to those capable of penetrating skin pores 5 to 10 times smaller than their size to enable delivery of skin-penetrating drugs with MW up to 1000 kDa. In addition, TDDS using transfersomes allows the administration of macromolecular drugs such as peptides or proteins [ 62 ].

Ethosomes are composed of phospholipids, alcohols, and water. Compared with liposomes, ethosomes have higher alcohol concentrations. Ethosomes promote the percutaneous penetration of drugs, with phospholipids also contributing to the process [ 63 ] (Fig. 3 A, B). The flexibility and fluidity of ethosomes increase as water molecules near the lipid headgroup are replaced by alcohol. Ethosomes have the characteristic size of small particles, a stable structure, and a high capture efficiency that can delay drug release; therefore, compared with regular liposomes, ethosomes can transport drugs with deep penetration or directly through the skin. These formulations are also known to greatly improve drug release into the circulating blood and drug transdermal efficacy. Ethosomes are a type of multiphase dispersion system characterized by better stability and a longer retention period than those of transfersomes.

Polymeric nanoparticles

Nanoparticles (NPs) are nanocarriers with sizes ranging between 1 and 1000 nm and can be classified into several types according to their composition. Drug administration in the form of NPs leads to targeted and controlled release behavior, changes in in vivo dynamics of the drug, and extends the drug residence time in the blood, which further lead to improved drug bioavailability and reduced toxicity and side effects. NPs are conventionally generated by polymerization and crosslinking, and biodegradable polymeric materials such as gelatin and polylactic acid (PLA) are often used [ 67 , 69 , 70 ]. In the field of TDDS, polymeric NPs are gaining increased attention because they can overcome the limitations of other lipid-based systems, such as by conferring protection to unstable drugs against degradation and denaturation and achieving continuous drug release to reduce side effects. Increase in the concentration gradient improves transdermal penetration of the drug. Depending on the manufacturing method and structure, polymeric NPs can be classified as nanospheres, nanocapsules, and polymer micelles. Widely used polymers include polylactic acid, poly(D,L-lactide-co-glycolide) (PLGA), polycaprolactone, polyacrylic acid, and natural poly esters (including chitosan, gelatin, and alginate). These polymer chains can be synthesized by covalent linkage of two or more single polymeric units under specific conditions, such as the presence of a synthetic membrane that mimics the cellular lipid bilayer membrane. Although these polymers can form a complex structure, the polymer membrane is highly structured owing to the high MW polymer chains; for this reason, polymeric NPs, characterized by high mechanical strength and non-deformability cannot pass through pores with dimensions smaller or equal to their size. However, these NPs can be difficult to break down, which means drugs can be stored for a substantially long period, followed by its release from the NPs and diffusion into deeper layers of the skin (Fig. 3 C-G) [ 68 , 78 , 79 , 80 ].

Nanoemulsion

Nanoemulsions are a mixture characterized by low viscosity and isotropic, thermodynamic, and dynamic stability [ 71 ]. The mixture consists of transparent or translucent oil globules dispersed in an aqueous phase stabilized by an interfacial membrane formed by surfactant or co-surfactant molecules of extremely small droplet size. The particle size of commonly used nanoemulsions ranges from 100 to 1000 nm, although an upper limit to the particle size has been proposed on account of its nanoscale dimensions. Nanoemulsions are different from microemulsions; although nanoemulsions have almost the same droplet size range, composition, and appearance as microemulsions, they differ greatly in terms of structural aspects and long-term thermodynamic stability. The small particle size, large specific surface area, and low surface tension of nanoemulsions provide excellent wettability that ensures close contact with the skin. In addition, nanoemulsions offer many other benefits such as high solubilization capacity and physical stability, improved bioavailability, ease of preparation, production with less energy input, and long shelf life. Nanoemulsions exhibit a shorter transdermal time and better transdermal absorption than commonly used topical skin preparations. Depending on the composition, nanoemulsions can include oil-in-water (O/W: oil phase dispersed in a continuous aqueous phase), water-in-oil (W/O: aqueous phase dispersed in a continuous oil phase), and bicontinuous/multiphasic emulsion. Several studies have reported the increased use of O/W nanoemulsions as a delivery system for encapsulating lipophilic components in pharmaceuticals, highlighting the immense potential of nanoemulsions in contributing to novel TDDS-based advances in pharmaceutical applications [ 58 , 68 ].

Methods for characterizing TDDS

The evaluation of delivery efficiency and effectiveness is a very important process in TDDS. There are various methods used for this, depending on the type and purpose of the drug to be delivered. However, the three most common methods involve the use of diffusion cells, tape stripping, and microscopic and spectroscopic examination [ 81 , 82 ], in which each method makes use of a distinct analysis method. As the drug applied to the surface is absorbed, all these characterization methods are based on the principle of measuring the amount of the drug in each surface layer or storing an imaging material instead the drug to visually confirm the absorption behavior.

Diffusion cell method

Tests employing diffusion cells represent the gold standard in the evaluation of TDDS, with Franz diffusion cells being the most common used setup (Fig.  4 ) [ 84 , 85 ]. This technique determines important relationships among the skin, active pharmaceutical ingredients, and the nature of the formulation. The diffusion cell consists of a chamber for drug application, a membrane within which the drug may diffuse, and an acceptor media chamber from which samples may be investigated. Diffusion cells are categorized into two main classes, namely, static and flow-through cells. In static cells, as in the popular Franz diffusion cell, the donor, the membrane, and the acceptor modules could be placed either vertically or horizontally. There are Franz cells that open from above; therefore, the measurement runs under conditions of atmospheric pressure. However, most of these cells are closed from the top, leading to increased pressure, which translates to an overestimation of penetration values. Nowadays, “hand-sampler” Franz diffusion cells have been replaced by systems connected to an automated sampler. These automated sampling systems facilitate the work of researchers and reduce errors from manually conducted experiments.

A Schematic illustrations of the static Franz diffusion cell. B Permeation profiles of ketoprofen (KTP) for 24 h in different conditions of matrix, medium, pH, and type of membrane. A , B Reproduced from [ 83 ], copyright permission by MDPI 2018

Tape stripping

Tape stripping is a commonly used minimally invasive method to test the penetration of topically applied formulations through the SC, where a layer of the SC is removed with an adhesive tape followed by examination of the skin layer on the adhesive tape (Fig.  5 ) [ 74 , 83 , 86 , 87 ]. The tape stripping process is performed after an appropriate incubation time post topical application of the test composition. The composition may be removed or left on the skin to provide the original amount of components to be used during the measurement. The adhesive tape is placed on the skin surface and is always removed from the same selection. It is important that the adhesive tape is always flattened with the same force as the roller to eliminate the effect of creases and recesses on tape stripping. In addition, the removal rate is an important factor. The slower the adhesive tape removal rate, the higher the adhesion of the SC to the patch, which increases the amount of skin removed from the patch. The removed adhesive tape contains both the SC layer and the active ingredients of the composition used. Several methods can be used to test samples harvested using adhesive tape. High-performance liquid chromatography (HPLC) analysis produces quantitative results, whereas spectroscopic methods produce semiquantitative insights. During HPLC analysis, the test material on the adhesive tape is extracted and analyzed on chromatographic separation. It is also possible to detect active substances using atomic absorption spectroscopy. However, the most prevalent method used to characterize skin harvested by tape stripping is attenuated total reflectance-Fourier transform infrared spectroscopy (ATR-FTIR). These spectroscopic measurements are based on sample irradiation, and changes in oscillations and bonding angles between atoms due to the absorption or scattering of infrared rays. The change in radiation on passing through the sample is measured by plotting the transmitted radiation as a function of wavelength/wavenumber. This analysis yields a spectrum that could be analyzed for both qualitative and quantitative information. Therefore, the depth of penetration is determined by the wavelength of the infrared radiation, the refractive index of the ATR crystal, and the measured material and angle of reflection. Tape stripping combined with ATR-FTIR spectroscopy is suitable for detecting a variety of exogenous substances in specific layers of the SC. However, the difficulty with this method is that characteristic peaks of the substance to be detected often overlap with peaks specific to the skin.

A Diagram illustrating the process of skin tape stripping. B Imaging of a 4-mm volar skin surface area of a healthy forearm with optical coherence tomography (OCT). A Reproduced from [ 66 ], copyright permission by Springer Nature 2021. B Reproduced from [ 74 ], copyright permission by Frontiers Media 2019

Microscopic and spectroscopic methods

Microscopy-based techniques can also provide important information about the spatial distribution of the drug within different skin layers or shed light on the mechanism of penetration. The two most common modalities of microscopy are confocal laser scanning microscopy (CLSM) and two-photon fluorescence microscopy (2-PFM) (Fig.  6 ) [ 58 , 68 , 71 , 72 , 74 , 80 , 81 , 82 , 83 , 84 , 85 , 86 , 87 , 90 ].

A CLSM images (100× magnification) of skin samples treated with free C-6, C-6/NLC, and C-6/SLN. B Enlarged CLSM Fig. (200× magnification). C Fluorescence intensity in receptor fluid at various times. D Reconstructed two-photon images in XZ orthogonal and 3D views. E Averaged normalized FITC-EGF signal intensity along the z-axis from the surface to the dermal layer of human skin samples. F Penetration depth of FITC-EGF with different thresholds of fluorescence intensity (50, 20, 10, and 5%) measured at the skin surface. A , B , C Reproduced from [ 88 ], copyright permission by Springer Nature 2018. D , E , F Reproduced from [ 89 ], copyright permission by OSA Publishing 2018

CLSM is a non-invasive method developed for fluorescence microscopy [ 88 , 91 , 92 ]. In the past few years, CLSM has been widely adopted as a technique to visualize fluorescent model compounds in the skin. CLSM can be used to examine skin structure without destroying tissue samples and is widely employed to evaluate the effect of physical and chemical enhancers on skin permeability. This method can be adapted for use in both in vivo and in vitro conditions. CLSM is used to diagnose common skin dysfunction and identify malignant lesions, along with characterization of keratinization and pigmentation disorders. CLSM can be applied to probe the mechanism underlying the promotion of transdermal transport by nanoparticle formulations. Fluorescent markers (e.g., fluorescein, Nile red, and 5-bromodeoxyuridine) can be included in the encapsulated nanostructured formulations. The therapeutic effectiveness of these formulations can be examined by CLSM to determine the penetration profile of these fluorescent markers across skin tissue or skin appendages.

In addition, 2-PFM has become an important tool for imaging skin cells [ 89 ]. This setup commonly uses a Ti-sapphire laser as the excitation source. In single-photon fluorescence, a fluorescent photon is generated when a high-energy photon excites the fluorophore and increases the energy level of one of its electrons to an excited state. In two-photon excitation, the combined energy transfer of two low-energy photons is sufficient to raise the same electron to a high energy level. The setup of a two-photon microscope is very similar to that of a CLSM, with two major differences. The 2-PFM setup works with an adjustable Ti-sapphire high-frequency pulsed laser, which emits red and near-infrared rays in the wavelength range of 650–1100 nm. Another significant difference is that there is no pinhole in front of the detector. The most relevant advantage of 2-PFM is that the total energy delivered to the specimen is much lower than that of other techniques. In addition, the two-photon excitation phenomenon involves fluorescence excitation of the sample in very small focal volumes, thereby reducing the possibility of photobleaching and photodamage. Skin samples can be studied without cryofixation or sectioning. For imaging of UV-absorbing fluorophores, less scattering and less absorption make deep tissue imaging possible using infrared excitation. The limitations of 2-PFM include the fact that this setup requires relatively expensive lasers and complex cooling systems. It also has a lower lateral resolution than other technologies; however, in practice, the resolution difference is not significant.

Conclusion and future perspectives

The development of TDDS technology is widely recognized as the development of a mass delivery methodology, which makes it the preferred drug injection modality for transdermal delivery across skin types, while preventing first-pass metabolism and other sensitivities associated with various alternative drug administration routes. In various devices and TDDSs, drugs can be delivered through the skin to the systemic circulation. Drugs are generally reliably and safely delivered through TDDS and are safe and stable from biochemical modifications until they reach the target tissue. TDDS is noninvasive, nonallergenic, and has a set duration and dose delivery method, which allows for uniform distribution of drugs at prescribed and controlled rates. Many new and old formulations are in the process of improving the bioavailability of low-absorption drugs via easy routes of administration that allow large doses to be administered over a long period of time. Therefore, the TDDS technology is growing rapidly in the pharmaceutical field and has succeeded in capturing key value in the market for biomedical applications as a formulation system that can improve drug delivery through topical routes. However, despite extensive research over the past few decades, passive methods such as chemical enhancers have had limited success in increasing transdermal transport of small molecules and have only had a relatively poor ability to increase transport of macromolecules under potentially clinically acceptable conditions. Active transport methods using external devices have more extensively increased the transdermal delivery efficiency of drugs and macromolecules. However, the ability of these technologies to effectively deliver drugs is partially balanced by their reliance on electronic control devices that require energy sources, which limits their utility and cost. Methods of piercing micron-sized pores into the skin, such as microneedles can significantly increase the transdermal delivery of drugs, macromolecules, or particles, but more studies are needed to achieve more safety/low skin damage and cost-effectiveness.

In recent years, the scale of TDDS in the domestic and overseas drug delivery system market has increased, as confirmed through increasing research studies, patents, and commercially available products from many companies and research institutes. In addition, microneedles are attracting great attention even among TDDS modalities, which complement the limitations of the existing simple application type and patch type needles and combine the advantages of microneedles to obtain higher treatment efficiency and effects. For this, manufacturing and commercialization methods are being developed, with judicious implementation of latest technologies, such as 3D bioprinting. Advances in these TDDSs could provide the driving force for controlling prevalence of diseases of cardiovascular and central nervous systems, diabetes, neuromuscular diseases, genetic diseases, and infectious and localized infectious diseases, while spearheading advances in vaccination and supporting patient preference for self-administration of drugs for long-term treatment.

Availability of data and materials

Not applicable.

Abbreviations

Drug delivery system

Transdermal drug delivery system

Stratum corneum

Molecular weight

High-performance liquid chromatography

Attenuated total reflectance Fourier transform infrared spectroscopy

Confocal laser scanning microscope

Two-photon scanning fluorescence microscope

Poly(lactic-co-glycolic acid)

Vega-Vásquez P, Mosier NS, Irudayaraj J. Nanoscale drug delivery systems: from medicine to agriculture. Front Bioeng Biotechnol. 2020;8:79. https://doi.org/10.3389/fbioe.2020.00079 .

Article   Google Scholar  

Vargason AM, Anselmo AC, Mitragotri S. The evolution of commercial drug delivery technologies. Nat Biomed Eng. 2021. https://doi.org/10.1038/s41551-021-00698-w .

Mali AD, Bathe R, Patil M. An updated review on transdermal drug delivery systems. Int J Adv Sci Res. 2015;1(6):244–54. https://doi.org/10.7439/ijasr.v1i6.2243 .

Li C, Wang J, Wang Y, Gao H, Wei G, Huang Y, et al. Recent progress in drug delivery. Acta Pharm Sin B. 2019;9(6):1145–62. https://doi.org/10.1016/j.apsb.2019.08.003 .

Kumar JA, Pullakandam N, Prabu SL, Gopal V. Transdermal drug delivery system: An overview. Int J Pharm Sci Rev Res. 2010;3(2):49–54.

CAS   Google Scholar  

Roohnikan M, Laszlo E, Babity S, Brambilla DA. Snapshot of transdermal and tropical drug delivery research in Canada. Pharmaceutics. 2019;11(6):256. https://doi.org/10.3390/pharmaceutics11060256 .

Article   CAS   Google Scholar  

Peña-Juárez MC, Guadarrama-Escobar OR, Escobar-Chávez JJ. Transdermal delivery Systems for Biomolecules. J Pharm Innov. 2021;6:1–14.

Google Scholar  

Ali H. Transdermal drug delivery system & patient compliance. MOJ Bioequiv Availab. 2017;3(2):47–8.

Leppert W, Malec–Milewska M, Zajaczkowska R, Wordliczek J. Transdermal and Topical Drug Administration in the Treatment of Pain. Molecules. 2018;23(3):681.

Akhter N, Singh V, Yusuf M, Khan RA. Non-invasive drug delivery technology: development and current status of transdermal drug delivery devices, techniques and biomedical applications. Biomed Tech. 2020;65(3):243–72. https://doi.org/10.1515/bmt-2019-0019 .

Pires LR, Vinayakumar KB, Turos M, Miguel V, Gaspar J. A perspective on microneedle-based drug delivery and diagnostics in Paediatrics. J Pers Med. 2019;9(4):49. https://doi.org/10.3390/jpm9040049 .

Ruby PK, Pathak SM, Aggarwal D. Critical attributes of transdermal drug delivery system (TDDS) – a generic product development review. Drug Dev Ind Pharm. 2014;40(11):1421–8. https://doi.org/10.3109/03639045.2013.879720 .

Ali S, Shabbir M, Shahid N. The structure of skin and transdermal drug delivery system - a review. Res J Pharm Tech. 2015;8(2):103–9. https://doi.org/10.5958/0974-360X.2015.00019.0 .

Wang M, Luo Y, Wang T, Wan C, Pan L, Pan S, et al. Artificial skin perception. Adv Mater. 2020;33:e2003014.

Hutton AR, McCrudden MT, Larrañeta E, Donnelly RF. Influence of molecular weight on transdermal delivery of model macromolecules using hydrogel-forming microneedles: potential to enhance the administration of novel low molecular weight biotherapeutics. J Mater Chem B. 2020;8(19):4202–9. https://doi.org/10.1039/D0TB00021C .

Andrews SM, Jeong EH, Prausnitz MR. Transdermal delivery of molecules is limited by full epidermis, Not Just Stratum Corneum. Pharm Res. 2013;30(4):1099–109.

Chaulagain B, Jain A, Tiwari A, Verma A, Jain SK. Passive delivery of protein drugs through transdermal route. Artif Cells Nanomed Biotechnol. 2018;46(1):472–87. https://doi.org/10.1080/21691401.2018.1430695 .

Uchechi O, Ogbonna J, Attama AA. Nanoparticles for dermal and transdermal drug delivery. In: Application of nanotechnology in drug delivery. Sezer AD: InTech C; 2014. p. 193–235.

Zhou X, Hao Y, Yuan L, Pradhan S, Shrestha K, Pradhan O. Nano-formulations for transdermal drug delivery: a review. Chin Chem Lett. 2018;29(12):1713–24. https://doi.org/10.1016/j.cclet.2018.10.037 .

Kováčik A, Kopečná M, Vávrová K. Permeation enhancers in transdermal drug delivery: benefits and limitations. Expert Opin Drug Deliv. 2020;17(2):145–55. https://doi.org/10.1080/17425247.2020.1713087 .

Pawar PM, Solanki KP, Mandali VA. Recent advancements in transdermal drug delivery system. Int J Pharm Clin Res. 2018;10(3):65–73.

Mujoriya R, Dhamande KA. Review on transdermal drug delivery system. Res J Sci Tech. 2011;3(4):227–31.

Patel R, Patel A, Prajapati B, Shinde G, Dharamsi A. Transdermal drug delivery systems: A mini review. Int J Adv Res. 2018;6(5):891–900. https://doi.org/10.21474/IJAR01/7109 .

Kakar S, Singh R, Rani P. A review on transdermal drug delivery. Innoriginal Int J Sci. 2016;3(4):1–5.

Wang Y, Zeng L, Song W, Liu J. Influencing factors and drug application of iontophoresis in transdermal drug delivery: an overview of recent progress. Drug Deliv Transl Res. 2021. https://doi.org/10.1007/s13346-021-00898-6 .

Dhal S, Pal K, Giri S. Transdermal delivery of gold nanoparticles by a soybean oil-based oleogel under iontophoresis. ACS Appl Bio Mater. 2020;3(10):7029–39. https://doi.org/10.1021/acsabm.0c00893 .

Moarefian M, Davalos RV, Tafti DK, Acheniec LE, Jones CN. Modeling iontophoretic drug delivery in a microfluidic device. Lab Chip. 2020;20(18):3310–21. https://doi.org/10.1039/D0LC00602E .

Byrne JD, Yeh JJ, DeSimone JM. Use of iontophoresis for the treatment of cancer. J Control Release. 2018;284:144–51. https://doi.org/10.1016/j.jconrel.2018.06.020 .

Sloan JB, Soltani K. Iontophoresis in dermatology: a review. J Am Acad Dermatol. 1986;15(4):671–84. https://doi.org/10.1016/S0190-9622(86)70223-5 .

Park J, Lee H, Lim GS, Kim N, Kim D, Kim YC. Enhanced transdermal drug delivery by sonophoresis and simultaneous application of sonophoresis and iontophoresis. AAPS PharmSciTech. 2019;20(3):96. https://doi.org/10.1208/s12249-019-1309-z .

Seah BC, Teo BM. Recent advances in ultrasound-based transdermal drug delivery. Int J Nanomedicine. 2018;13:7749–63. https://doi.org/10.2147/IJN.S174759 .

Nguyen HX, Banga AK. Electrically and ultrasonically enhanced transdermal delivery of methotrexate. Pharmaceutics. 2018;10(3):117. https://doi.org/10.3390/pharmaceutics10030117 .

Escobar-Chávez JJ, Díaz-Torres R, Domínguez-Delgado CL, Rodríguez-Cruz IM, López-Arellano R, Hipólito EAM. Therapeutic applications of sonophoresis and sonophoretic devices. In: Percutaneous Penetration Enhancers Physical Methods in Penetration Enhancement. Springer-Verlag Berlin Heidelberg; 2017. p. 31–58.

Charoo NA, Rahman Z, Repka MA, Murthy SN. Electroporation: An avenue for transdermal drug delivery. Curr Drug Deliv. 2010;7(2):125–36. https://doi.org/10.2174/156720110791011765 .

Chen X, Zhu L, Li R, Pang L, Zhu S, Ma J, et al. Electroporation-enhanced transdermal drug delivery: effects of logP, pKa, solubility and penetration time. Eur J Pharm Sci. 2020;151:105410. https://doi.org/10.1016/j.ejps.2020.105410 .

Sokołowska E, Błachnio-Zabielska AU. A critical review of electroporation as a plasmid delivery system in mouse skeletal muscle. Int J Mol Sci. 2019;20(11):2776. https://doi.org/10.3390/ijms20112776 .

Dermol-Černe J, Pirc E, Miklavčič D. Mechanistic view of skin electroporation – models and dosimetry for successful applications: an expert review. Expert Opin Drug Deliv. 2020;17(5):689–704. https://doi.org/10.1080/17425247.2020.1745772 .

Escobar-Chávez JJ, Bonilla-Martínez D, Villegas-González MA, Revilla-Vázquez AL. Electroporation as an efficient physical enhancer for skin drug delivery. J Clin Pharmacol. 2009;49(11):1262–83. https://doi.org/10.1177/0091270009344984 .

Lin CH, Aljuffali IA, Fang JY. Lasers as an approach for promoting drug delivery via skin. Expert Opin Drug Deliv. 2014;11(4):599–614. https://doi.org/10.1517/17425247.2014.885501 .

Lee S, Kollias N, McAuliffe DJ, Flotte TJ, Doukas AG. Topical drug delivery in humans with a single photomechanical wave. Pharm Res. 1999;16(11):1717–21. https://doi.org/10.1023/A:1018954015441 .

Agrawal S, Gandhi SN, Gurgar P, Saraswathy N. Microneedles: An advancement to transdermal drug delivery system approach. J Appl Pharm Sci. 2020;10(3):149–59.

Zhao Z, Chen Y, Shi Y. Microneedles: a potential strategy in transdermal delivery and application in the management of psoriasis. RSC Adv. 2020;10(24):14040–9. https://doi.org/10.1039/D0RA00735H .

Jung JH, Jin SG. Microneedle for transdermal drug delivery: current trends and fabrication. J Pharm Investig. 2021. https://doi.org/10.1007/s40005-021-00512-4 .

Shakya AK, Ingrole RSJ, Joshi G, Uddin MJ, Anvari S, Davis CM, et al. Microneedles coated with peanut allergen enable desensitization of peanut sensitized mice. J Control Release. 2019;314:38–47. https://doi.org/10.1016/j.jconrel.2019.09.022 .

Lim J, Tahk D, Yu J, Min DH, Jeon NL. Design rules for a tunable merged-tip microneedle. Microsyst Nanoeng. 2018;4(1):1–10.

Dardano P, Caliò A, Palma VD, Bevilacqua MF, Matteo AD, Stefano LD. A photolithographic approach to polymeric microneedles array fabrication. Materials. 2015;8(12):8661–73. https://doi.org/10.3390/ma8125484 .

Han D, Morde RS, Mariani S, Mattina AAL, Vignali E, Yang C, et al. 4D printing of a bioinspired microneedle array with backward-facing barbs for enhanced tissue adhesion. Adv Funct Mater. 2020;30(11):1909197. https://doi.org/10.1002/adfm.201909197 .

Xiao Z, Li X, Zhang P, Wang Y. Research of polymeric microneedles for transdermal drug delivery. Progress Chem. 2017;29(12):1518.

Balmert SC, Carey CD, Falo GD, Sethi SK, Erdos G, Korkmaz E, et al. Dissolving undercut microneedle arrays for multicomponent cutaneous vaccination. J Control Release. 2020;317:336–46. https://doi.org/10.1016/j.jconrel.2019.11.023 .

Kim HM, Lim YY, An JH, Kim MN, Kim BJ. Transdermal drug delivery using disk microneedle rollers in a hairless rat model. Int J Dermatol. 2012;51(7):859–63. https://doi.org/10.1111/j.1365-4632.2011.05343.x .

Li K, Yoo KH, Byun HJ, Lim YY, Kim MN, Hong HK, et al. The microneedle roller is an effective device for enhancing transdermal drug delivery. Int J Dermatol. 2012;51(9):1137–9. https://doi.org/10.1111/j.1365-4632.2010.04703.x .

Bariya SH, Gohel MC, Mehta TA, Sharma OP. Microneedles: an emerging transdermal drug delivery system. J Pharm Pharmacol. 2012;64(1):11–29.

Lee YS, Kang TG, Cho HR, Lee GJ, Park OK, Kim SY, et al. Localized delivery of Theranostic nanoparticles and high-energy photons using microneedles-on-bioelectronics. Adv Mater. 2021;33(24):2100425. https://doi.org/10.1002/adma.202100425 .

Du H, Liu P, Zhu J, Li Y, Zhang L, Zzhu J, et al. Hyaluronic acid-based dissolving microneedle patch loaded with methotrexate for improved treatment of psoriasis. ACS Appl Mater Interfaces. 2019;11(46):43588–98. https://doi.org/10.1021/acsami.9b15668 .

Alkilani AZ, Mccrudden MTC, Donnelly RF. Transdermal drug delivery: innovative pharmaceutical developments based on disruption of the barrier properties of the stratum corneum. Pharmaceutics. 2015;7(4):438–70. https://doi.org/10.3390/pharmaceutics7040438 .

Amjadi M, Mostaghaci B, Sitti M. Recent advances in skin penetration enhancers for transdermal gene and drug delivery. Curr Gene Ther. 2017;17(2):139–46.

Li Y, Guo L, Lu W. Laser ablation-enhanced transdermal drug delivery. Photonics Lasers Med. 2013;2(4):315–22.

Szunerits S, Boukherroub R. Heat: a highly efficient skin enhancer for transdermal drug delivery. Front Bioeng Biotechnol. 2018;6:15. https://doi.org/10.3389/fbioe.2018.00015 .

Parhi R, Mandru A. Enhancement of skin permeability with thermal ablation techniques: concept to commercial products. Drug Delivery Translat Res. 2021;11:817-41.

Babaie S, Del Bakhshayesh AR, Ha JW, Hamishehkar H. Invasome: A novel nanocarrier for transdermal drug delivery. Nanomaterials. 2020;10(2):341. https://doi.org/10.3390/nano10020341 .

Palac Z, Engesland A, Flaten GE, Škalko-Basnet N, Filipović-Grčić J, Vanić Ž. Liposomes for (trans)dermal drug delivery: the skin-PVPA as a novel in vitro stratum corneum model in formulation development. J Liposome Res. 2014;24(4):313–22. https://doi.org/10.3109/08982104.2014.899368 .

Rajan R, Jose S, Mukund VP, Vasudevan DT. Transferosomes-a vesicular transdermal delivery system for enhanced drug permeation. J Adv Pharm Technol Res. 2011;2(3):138–43. https://doi.org/10.4103/2231-4040.85524 .

Zhang Z, Liu Y, Chen Y, Li L, Lan P, He D, et al. Transdermal delivery of 5-aminolevulinic acid by nanoethosome gels for photodynamic therapy of hypertrophic scars. ACS Appl Mater Interfaces. 2019;11(4):3704–14. https://doi.org/10.1021/acsami.8b17498 .

Stanekzai A, Sudhakar CK, Zhakfar AM, Karan VS. Recent approaches in transdermal drug delivery system. Res J Pharm Tech. 2019;12(9):4550–8. https://doi.org/10.5958/0974-360X.2019.00783.2 .

Hadgraft J, Lane ME. Passive transdermal drug delivery systems. Am J Drug Deliv. 2006;4(3):153–60. https://doi.org/10.2165/00137696-200604030-00003 .

Gandhi K, Dahiya A, Monika LT, Singh K. Transdermal drug delivery – a review. Int J Res Pharm Sci. 2012;3(3):379–88.

Dhal S, Gavara RR, Pal K, Banerjee I, Mishra M, Giri S. Facile transdermal delivery of upconversion nanoparticle by iontophoresis-responsive magneto-upconversion oleogel. Nano Express. 2020;1(1):010012. https://doi.org/10.1088/2632-959X/ab81e1 .

Jeong WY, Kim SD, Lee SY, Lee HS, Han DW, Yang SY, et al. Transdermal delivery of Minoxidil using HA-PLGA nanoparticles for the treatment in alopecia. Biomater Res. 2019;23(1):16. https://doi.org/10.1186/s40824-019-0164-z .

Lee SY, Kang MS, Jeong WY, Han DW, Kim KS. Hyaluronic acid-based theranostic nanomedicines for targeted cancer therapy. Cancers. 2020;12(4):940. https://doi.org/10.3390/cancers12040940 .

Jijie R, Barras A, Boukherroub R, Szunerits S. Nanomaterials for transdermal drug delivery: beyond the state of the art of liposomal structures. J Mater Chem B. 2017;5(44):8653–75. https://doi.org/10.1039/C7TB02529G .

Ali A, Ahmad U. Nanoemulsion as a vehicle in transdermal drug delivery. Insights Biomed. 2018;3:15.

Halnor V, Pande V, Borawake D, Nagare H. Nanoemulsion: A novel platform for drug delivery system. J Mat Sci Nanotechol. 2018;6(1):104.

Benson HA, Grice JE, Mohammed Y, Namjoshi S, Roberts MS. Topical and transdermal drug delivery: from simple potions to smart technologies. Curr Drug Deliv. 2019;16(5):444–60. https://doi.org/10.2174/1567201816666190201143457 .

Ramadon D, McCrudden MT, Courtenay AJ, Donnelly RF. Enhancement strategies for transdermal drug delivery systems: current trends and applications. Drug Deliv Transl Res. 2021;20:1–34.

Lee HJ, Song CY, Baik SM, Kim DY, Hyeon TG, Kim DH. Device-assisted transdermal drug delivery. Adv Drug Deliv Rev. 2018;127:35–45. https://doi.org/10.1016/j.addr.2017.08.009 .

Bird D, Ravindra NM. Transdermal drug delivery and patches - an overview. Med Devices Sens. 2020;3(6):e10069.

Hanbali OA, Khan HS, Sarfraz M, Arafat M, Ijaz S, Hameed A. Transdermal patches: design and current approaches to painless drug delivery. Acta Pharma. 2019;69(2):197–215. https://doi.org/10.2478/acph-2019-0016 .

Escobar-Chávez JJ, Díaz-Torres R, Rodríguez-Cruz IM, Domínguez-Delgado CL, Morales RS, Ángeles-Anguiano E, et al. Nanocarriers for transdermal drug delivery. Res Rep Transdermal Drug Delivery. 2012;1:3–17.

Patra JK, Das G, Fraceto LF, Campos EV, Del Pilar R-TM, Acosta-Torres LS, et al. Nano based drug delivery systems: recent developments and future prospects. J Nanobiotechnol. 2018;16(1):71. https://doi.org/10.1186/s12951-018-0392-8 .

Buhecha MD, Lansley AB, Somavarapu S, Pannala AS. Development and characterization of PLA nanoparticles for pulmonary drug delivery: co-encapsulation of theophylline and budesonide, a hydrophilic and lipophilic drug. J Drug Deliv Sci Technol. 2019;53:101128. https://doi.org/10.1016/j.jddst.2019.101128 .

Raza R, Mittai A, Kumar P, Alam S, Prakash S, Chauhan N. Approaches and evaluation of transdermal drug delivery system. Int J Drug Dev Res. 2015;7(1):222–33.

Sheth NS, Mistry RB. Formulation and evaluation of transdermal patches and to study permeation enhancement effect of eugenol. J. Appl. Pharm. Sci. 2011;1(3):96–101.

Salamanca CH, Barrera-Ocampo A, Lasso JC, Camacho N, Yarce CJ. Franz diffusion cell approach for pre-formulation characterisation of ketoprofen semi-solid dosage forms. Pharmaceutics. 2018;10(3):148. https://doi.org/10.3390/pharmaceutics10030148 .

Lee JD, Kim JY, Jang HJ, Lee BM, Kim KB. Percutaneous permeability of 1-phenoxy-2-propanol, a preservative in cosmetics. Regul Toxicol Pharmacol. 2019;103:56–62. https://doi.org/10.1016/j.yrtph.2019.01.002 .

Rajithaa P, Shammika P, Aiswarya S, Gopikrishnan A, Jayakumar R, Sabitha M. Chaulmoogra oil based methotrexate loaded topical nanoemulsion for the treatment of psoriasis. J Drug Deliv Sci Technol. 2019;49:463–76. https://doi.org/10.1016/j.jddst.2018.12.020 .

Escobar-Chávez JJ, Merino-Sanjuán V, López-Cervantes M, Urban-Morlan Z, Piñón-Segundo E, Quintanar-Guerrero D, et al. The tape-stripping technique as a method for drug quantification in skin. J Pharm Pharm Sci. 2008;11(1):104–30. https://doi.org/10.18433/J3201Z .

Russell LM, Guy RH. Novel imaging method to quantify stratum corneum in dermatopharmacokinetic studies: proof-of-concept with acyclovir formulations. Pharm Res. 2012;29(12):3362–72. https://doi.org/10.1007/s11095-012-0831-4 .

Gu Y, Yang M, Tang X. Lipid nanoparticles loading triptolide for transdermal delivery: mechanisms of penetration enhancement and transport properties. J Nanobiotechnol. 2018;16(1):1–14.

Ahn J, Kim KH, Choe K, Lim JH, Lee SK, Kim YS, et al. Quantitative two-photon microscopy imaging analysis of human skin to evaluate enhanced transdermal delivery by hybrid-type multi-lamellar nanostructure. Biomed Opt Express. 2018;9(8):3974–82. https://doi.org/10.1364/BOE.9.003974 .

Berekméri A, Tiganescu A, Alase AA, Vital E, Stacey M, Wittmann M. Non-invasive approaches for the diagnosis of autoimmune/autoinflammatory skin diseases-a focus on psoriasis and lupus erythematosus. Front Immunol. 2019;10:1931. https://doi.org/10.3389/fimmu.2019.01931 .

Ashtikar M, Matthäus C, Schmitt M, Krafft C, Fahr A, Popp J. Non-invasive depth profile imaging of the stratum corneum using confocal raman microscopy: first insights into the method. Eur J Pharm Sci. 2013;50(5):601–8. https://doi.org/10.1016/j.ejps.2013.05.030 .

Zhang LW, Monteiro-Riviere NA. Use of confocal microscopy for nanoparticle drug delivery through skin. J Biomed Opt. 2012;18(6):061214. https://doi.org/10.1117/1.JBO.18.6.061214 .

Download references

This work was supported by a 2-Year Research grant from the Pusan National University.

Author information

Woo Yeup Jeong, Mina Kwon and Hye Eun Choi contributed equally to this work.

Authors and Affiliations

School of Chemical Engineering, Pusan National University, 2 Busandaehak-ro 63 beon-gil, Geumjeong-gu, Busan, 46241, Republic of Korea

Woo Yeup Jeong, Mina Kwon, Hye Eun Choi & Ki Su Kim

You can also search for this author in PubMed   Google Scholar

Contributions

Conceptualization, K.S.K.; investigation, W.Y.J., M.K., and H.E.C.; manuscript preparation, W.Y.J., M.K., H.E.C., and K.S.K.; visualization, W.Y.J., M.K., and H.E.C; supervision, K.S.K.; funding acquisition, K.S.K. All authors have read and agreed to the version of the manuscript for publication.

Corresponding author

Correspondence to Ki Su Kim .

Ethics declarations

Ethics approval and consent to participate, consent for publication, competing interests.

The authors declare no conflict of interest.

Additional information

Publisher’s note.

Springer Nature remains neutral with regard to jurisdictional claims in published maps and institutional affiliations.

Rights and permissions

Open Access This article is licensed under a Creative Commons Attribution 4.0 International License, which permits use, sharing, adaptation, distribution and reproduction in any medium or format, as long as you give appropriate credit to the original author(s) and the source, provide a link to the Creative Commons licence, and indicate if changes were made. The images or other third party material in this article are included in the article's Creative Commons licence, unless indicated otherwise in a credit line to the material. If material is not included in the article's Creative Commons licence and your intended use is not permitted by statutory regulation or exceeds the permitted use, you will need to obtain permission directly from the copyright holder. To view a copy of this licence, visit http://creativecommons.org/licenses/by/4.0/ . The Creative Commons Public Domain Dedication waiver ( http://creativecommons.org/publicdomain/zero/1.0/ ) applies to the data made available in this article, unless otherwise stated in a credit line to the data.

Reprints and permissions

About this article

Cite this article.

Jeong, W.Y., Kwon, M., Choi, H.E. et al. Recent advances in transdermal drug delivery systems: a review. Biomater Res 25 , 24 (2021). https://doi.org/10.1186/s40824-021-00226-6

Download citation

Received : 11 May 2021

Accepted : 18 July 2021

Published : 28 July 2021

DOI : https://doi.org/10.1186/s40824-021-00226-6

Share this article

Anyone you share the following link with will be able to read this content:

Sorry, a shareable link is not currently available for this article.

Provided by the Springer Nature SharedIt content-sharing initiative

• Transdermal drug delivery
• Active/passive method
• Characterization

Biomaterials Research

ISSN: 2055-7124

• Submission enquiries: Access here and click Contact Us
• General enquiries: [email protected]

• Reference Manager
• Simple TEXT file

People also looked at

Review article, research progress in brain-targeted nasal drug delivery.

• 1 Department of Anesthesiology, The General Hospital of Western Theater Command, Chengdu, China
• 2 College of Medicine, Southwest Jiaotong University, Chengdu, China
• 3 Department of Neurosurgery, The General Hospital of Western Theater Command, Chengdu, China

The unique anatomical and physiological connections between the nasal cavity and brain provide a pathway for bypassing the blood–brain barrier to allow for direct brain-targeted drug delivery through nasal administration. There are several advantages of nasal administration compared with other routes; for example, the first-pass effect that leads to the metabolism of orally administered drugs can be bypassed, and the poor compliance associated with injections can be minimized. Nasal administration can also help maximize brain-targeted drug delivery, allowing for high pharmacological activity at lower drug dosages, thereby minimizing the likelihood of adverse effects and providing a highly promising drug delivery pathway for the treatment of central nervous system diseases. The aim of this review article was to briefly describe the physiological structures of the nasal cavity and brain, the pathways through which drugs can enter the brain through the nose, the factors affecting brain-targeted nasal drug delivery, methods to improve brain-targeted nasal drug delivery systems through the application of related biomaterials, common experimental methods used in intranasal drug delivery research, and the current limitations of such approaches, providing a solid foundation for further in-depth research on intranasal brain-targeted drug delivery systems (see Graphical Abstract ).

Graphical Abstract. According to the factors that affect the absorption of drugs into the brain through the nose, the bioavailability of drugs administered through the nasal brain pathway can be improved by changing the dosage form of the drug, adding mucosal adhesives, applying osmotic enhancers, adding vasoconstrictors, and using a new drug delivery system.

1 Introduction

In recent years, the incidence rate of brain diseases such as Alzheimer’s disease, Parkinson’s disease, stroke, and brain tumors has continued to increase, which has been hugely detrimental to human health ( Palmer, 2010 ). These changes have led to increased research activity to develop better therapeutic strategies; however, there are many obstacles affecting the development of new drugs for the treatment of such diseases, one of which is the existence of the blood–brain barrier (BBB) ( Palmer, 2010 ; Xue et al., 2023 ). The BBB is a complex multicellular structure with extremely low permeability, which limits the movement of molecules between the blood and the neural tissues comprising the central nervous system (CNS) to help maintain homeostasis ( Xue et al., 2023 ). However, the tight junctions of the BBB and the presence of efflux transporters and metabolic enzymes limit the entry of approximately 98% of chemical drugs and nearly 100% of macromolecular drugs (including peptides and genetic drugs) into the brain when administered orally or via injection, making it difficult to ensure such drugs reach the target neural tissues at the concentrations required to induce therapeutic effects; this problem has continued to be a bottleneck that negatively impacts drug delivery in the CNS ( Nguyen et al., 2021 ). Therefore, exploring novel drug delivery systems and pathways to overcome the limitations imposed by the BBB to allow for targeted drug distribution in the CNS has become a top priority for future research.

Nasal administration is a systemic route of administration that leads to rapid drug absorption and high bioavailability, while minimizing the likelihood of adverse effects; these factors, in combination with the ease of administration, have led to increasing research interest in intranasal drug delivery methods ( Landis et al., 2012 ). There is a direct anatomical pathway between the nasal cavity and brain, and studies have shown that some drugs can be administered intranasally, allowing the BBB to be bypassed to facilitate drug entry into the brain, providing a promising approach to brain-targeted drug delivery ( Pardeshi and Belgamwar, 2013 ).

The aim of this review article was to summarize what is currently known about the physiological structure of the nasal cavity and brain, the pathways through which drugs can enter the brain through the nose, factors that affect brain-targeted nasal drug delivery, methods to improve brain-targeted nasal drug delivery through the application of related biomaterials, common experimental methods used in intranasal drug delivery research, as well as current limitations to provide a reference for researchers in the field.

2 Physiological structure of the nasal cavity and brain

2.1 physiological structure of the nasal cavity.

The human nasal cavity system is a complex structure, with an approximate length of 12–15 cm ( Hu et al., 2023 ). It is composed of two cavities, each of which is approximately 7.5 mL in volume, separated by the nasal septum, with an inner surface covered by mucosa; the mucosa is composed of the epithelium and lamina propria and is approximately 150 cm 2 in total surface area ( Hu et al., 2023 ). It is generally believed that the average pH of nasal mucus ranges from 5.5 to 6.5 ( Hu et al., 2023 ). The nasal cavity can be divided into three distinct areas based on function and organizational structure, including the vestibular, respiratory, and olfactory areas ( Babu et al., 2023 ). The vestibular area is located in the most anterior part of the nasal cavity and is covered with squamous epithelial cells and contains nasal hair secretory glands that limit drug penetration into the tissue; thus, this area is not typically regarded as a site of drug absorption ( Xinchen et al., 2023 ). The epithelial cells of the respiratory area, which occupies 80%–90% of the inner surface of the nasal cavity, consist of goblet, ciliated, intermediate, and basal cells, and the area contains a large number of mucus-secreting glands and trigeminal nerve branches ( Boyuklieva et al., 2023 ). The olfactory area is located at the top of the nasal cavity, below the cribriform plate ( Wu et al., 2023 ). The epithelial cell layer of the olfactory area is composed of olfactory, supporting, and basal cells. Olfactory cells are bipolar neurons; at one end, there is a dendrite that expands into segments and divides into many cilia that extend into the mucosal layer to allow for communication with the external environment; at the opposite end, there is an axon that can pass through the basal membrane and become wrapped in olfactory sheath cells, ultimately merging into the olfactory nerve, passing through the ethmoid plate and entering the olfactory bulb of the CNS ( Awad et al., 2023 ). In addition, the interstitial fluid surrounding the olfactory nerve bundle is connected to the cerebrospinal fluid (CSF) in the subarachnoid space, and supporting and basal cells can differentiate into new olfactory neurons and other cell types. The lamina propria of the olfactory and respiratory areas contain an abundance of blood and lymphatic vessels that extend to the lymph nodes deep within the neck ( Shrewsbury, 2023 ).

2.2 Relevant physiological structure of the CNS

Three protective capsules surround the brain. The outermost layer is the dura mater, with dural sinuses that contain venous vessels, and the innermost layer is the pia mater. The arachnoid membrane exists between the inner and outer layers, with the CSF flowing in the interstices below it ( Kadry et al., 2020 ). CSF is produced by the choroid plexus and flows within the ventricles and subarachnoid space, eventually entering the dural sinuses and mixing with venous blood ( Kadry et al., 2020 ). When the CSF flows through the sieve plates, up to 5% can enter the lamina propria of the olfactory mucosa through the perivascular space near the olfactory filaments before infiltrating into the lymphatic vessels of the nose ( Engelhardt and Sorokin, 2009 ). CSF plays an important role in protecting the brain, maintaining intracranial pressure, and regulating brain metabolism. Although there is an abundant blood supply in brain tissue, a tight BBB is formed by the presence of atresia bands between the endothelial cells of the brain capillaries, an intact basement membrane, and an astrocyte membrane coating the exterior of the vessels. In contrast, the barrier between brain tissue and the CSF, which is composed of soft meninges and glial membranes, is less effective at minimizing the movement of molecules, providing a physiological basis for the pathway of drug delivery from the nose to the CNS via the CSF ( Illum, 2004 ).

3 The pathways through which drugs can enter the brain through the nose

At present, the precise pathways through which drugs are transported to the brain following absorption across the nasal mucosa remain unclear; however, four main approaches have been identified to date.

3.1 Olfactory nerve pathway

Some studies have shown that most neurophilic viruses (such as rabies, herpetic stomatitis, and equine encephalomyelitis viruses), steroid hormones, metal ions (such as cadmium and nickel), and proteins enter the brain through the olfactory nerve pathway ( Crowe et al., 2018 ). After crossing the olfactory mucosa, these substances are absorbed at the axon terminals of olfactory neurons through pinocytosis, endocytosis, or simple diffusion ( Chung et al., 2022 ). They subsequently flow through the axonal plasma of neurons and are directly transported through the sieve plates to the olfactory bulb, reaching the rhinencephalon ( Chung et al., 2022 ). The olfactory nerve pathway is considered to be the most important pathway for the entry of drugs into the brain from the nose, and it represents the most direct pathway for bypassing the BBB ( Ruigrok and de Lange, 2015 ). However, axonal transport is relatively slow, the rate of which can vary from 0.1∼4 mm/d to 20∼400 mm/d depending on the properties of the drug being administered. Thus, this pathway is characterized by slow absorption, preventing a drug from entering the brain quickly. At the fastest rate, a drug could enter the brain within 1–2 h, however, at the slower rates, some drugs may not enter the brain for up to 24 h, limiting the clinical applications ( Singh and Shukla, 2023 ).

3.2 Olfactory mucosal epithelial pathway

The mucosal epithelial pathway (i.e., the nose–brain pathway) is the route through which drugs reach the olfactory mucosal epithelium and enter the CNS directly via cytosolic action or diffusion. Most small-molecule drugs, such as lidocaine, dopamine, 5-fluorouracil, dihydroergotamine, and insulin, are absorbed into the brain via this pathway ( Illum, 2003 ). This pathway is further divided into two components, including a transcellular transport pathway ( Illum, 2003 ) and a paracellular pathway ( Wolburg et al., 2008 ); in the former, drug molecules are transported across mucosal epithelia other than olfactory neuron receptor cells via mechanisms involving carrier transport or cytosolic or passive diffusion into supporting and glandular cells surrounding the olfactory nerve. In the latter pathway, drug molecules enter the intercellular fluid through either the interstitial space of the supporting cells or the peripheral cleft between the supporting cells and the olfactory nerve. If the drug molecules transported to the basement membrane are close to the axons in the lamina propria, they will be transported to the CSF within the cells surrounding the neuron. Simultaneously, drugs in the lamina propria may enter the lymphatic and systemic circulation. If the drug molecules transported to the basement membrane pass through the lamina propria, they will enter the space around the olfactory nerve bundle before being transported into the CSF. This pathway is dependent on the anatomical connections between the olfactory submucosa and subarachnoid space ( Ladel et al., 2019 ). Compared to the rate of absorption via the olfactory nerve pathway, the olfactory mucosal epithelial pathway allows for more rapid drug absorption, facilitating drug entry into brain tissue and the CSF within just a few minutes after nasal administration ( Yokel, 2022 ).

3.3 Trigeminal nerve pathway

The ophthalmic and maxillary nerve branches of the trigeminal nerve can extend to epithelial cells in the olfactory and respiratory areas of the nasal cavity, with the opposite end entering the CNS at the pons and terminating at the spinal nucleus of the trigeminal nerve in the brainstem or passing through the ethmoid plate and terminating at the olfactory bulb area ( Schaefer et al., 2002 ). In a study conducted by Thorne et al. (2004) , strong radioactivity was observed in the trigeminal nerve, trigeminal branch ganglia, and olfactory bulb following intranasal administration of iodine-125-conjugated insulin-like growth factor 1, with a concentration 10 times higher in the trigeminal nerve than in the olfactory bulb, confirming that, in some cases, the trigeminal nerve may serve as a pathway through which drugs can enter the brain following intranasal administration. However, the transit time along the trigeminal nerve has been reported to be 17–56 h longer than that along the olfactory nerve ( Gänger and Schindowski, 2018 ).

3.4 Blood circulation pathway

Low-molecular-weight lipophilic drugs predominantly enter the brain following absorption into the general circulation through the rich capillary network in the lamina propria of the respiratory region. However, after entering the general circulation, drugs must cross the BBB to reach the CNS; thus, this pathway is a limiting factor in the therapeutic application of many drugs ( Illum, 2000 ).

Following nasal administration, a drug will eventually reach the CNS through one or more of the aforementioned pathways, with differences in drug properties, formulations, and routes of administration dictating the dominant pathway of a drug delivery system.

4 Factors affecting brain-targeted nasal drug delivery

After a drug is administered via the nose, some of it may enter the CNS directly, whereas some will be absorbed into the general circulation, and some will be removed by the cilia of the nasal mucosa. The amount of drug that can be absorbed through the nasal mucosa is mainly influenced by the physiological characteristics of the nasal cavity, the nature of the drug itself, and factors related to its formulation.

4.1 Physiological properties of the nasal cavity

4.1.1 mechanism of mucociliary clearance.

One of the main limitations associated with nasal absorption is the rapid removal of a drug from the nasal cavity by ciliated cells. The cylindrical cilia lining the surface of the nasal mucosa sway at a speed of 5∼6 mm/min in order to move mucus from the nasal cavity to the pharynx and to eliminate substances adhered to the surface of the nasal mucosa. Although this process is a protective mechanism for maintaining optimal respiratory system functionality, it can also shorten the contact time between a drug and the nasal mucosa, directly affecting the efficiency of drug absorption through the nose, as a drug molecule typically remains within the nasal cavity for a period of approximately 20–30 min after administration ( Du et al., 2023 ). To counteract the effects of this clearance mechanism, it is usually possible to extend the drug adhesion time and increase drug absorption by selecting appropriate dosage forms, using bioadhesive materials, or modifying the surface of drug carriers, such as nanoparticles ( Badran et al., 2023 ), liposomes ( Kannavou et al., 2023 ), and microemulsions ( Pailla et al., 2022 ) with specific ligands.

4.1.2 Expression of transport proteins

Efflux mechanisms involving transport proteins expressed in the nasal mucosa, such as multidrug resistance protein 1, breast cancer resistance protein, and P-glycoprotein (P-gp), can limit the absorption of certain drugs. Therefore, inhibiting the synthesis and expression of these proteins can reduce drug clearance ( Chen et al., 2018 ). For example, a previous study demonstrated that P-gp plays an important role in drug absorption and efflux in the nasal cavity, and treatment with PSC-833, a specific P-gp inhibitor, increased the amount of quinidine (a P-gp substrate used as a model drug) reaching the brain via the nasal nerves ( Bors et al., 2020 ).

4.1.3 Role of enzymes

Many enzyme systems catalyze the biological processes occurring within the nasal mucosa; these systems include cytochrome enzyme subunits, epoxidases, polymerases, peptidases, and proteases, all of which can lead to drug degradation in the nasal cavity, producing a “pseudo-first pass effect” that limits the amount of peptide- and protein-based drugs that might otherwise enter the brain ( Oliveira et al., 2016 ). Administration of enzyme inhibitors can reduce the expression or activity of enzymes, thereby increasing drug uptake ( Khatri et al., 2023 ).

4.1.4 Blood vessel absorption

The abundance of capillaries in the nasal cavity can increase the amount of a drug being absorbed into the bloodstream, resulting in high drug concentrations in the general circulation. However, this also reduces the probability of direct drug entry into the brain and increases the likelihood of side effects ( Kumar et al., 2016 ). The use of vasoconstrictors at the administration site can, to some extent, prevent excessive drug diffusion into the bloodstream, thereby increasing brain-targeted drug delivery.

4.2 The physical and chemical properties of the drug itself

4.2.1 drug molecular weight.

Under normal circumstances, a higher molecular weight will lower the efficiency of drug absorption ( Fortuna et al., 2014 ). For lipid-soluble drugs, the rate of absorption across the nasal mucosa is nearly 100%; however, these rates are also affected by the molecular weight of a drug, with the absorption efficiency being significantly reduced when the molecular weight exceeds 1 kDa. For macromolecular water-soluble drugs, the concentration of drugs entering the CSF decreases as the relative molecular weight increases. In addition, the absorption efficiency of polar drugs is affected by the molecular weight; for example, when the molecular weight of a polar drug is less than 300 Da, the absorption efficiency is mostly unaffected by the drug’s physical and chemical properties, whereas the absorption efficiency is highly affected by the physical and chemical properties of a drug whose molecular weight exceeds 300 Da ( Appu et al., 2016 ).

4.2.2 Drug solubility

Before a drug can be absorbed through the nasal mucosa, it must first dissolve in nasal secretions. In terms of composition, water accounts for 90% of the components of nasal secretions, with the remainder comprising small amounts of mucin, proteins, inorganic salts, and lipids, among other factors ( Hussain et al., 2023 ). Therefore, drugs with high water solubility will also be soluble in nasal secretions. However, for drugs with low water solubility, carriers such as cyclodextrins ( De Gaetano et al., 2022 ), microemulsions ( Pires et al., 2021 ), and nanoparticles ( Abla and Mehanna, 2023 ) can be selected as drug delivery vectors, allowing the carrier instead of the drug to come into direct contact with the nasal mucosa. The shells of these carriers have good hydrophilicity, which improves their solubility in mucus.

4.2.3 Liposolubility of drugs

The smooth and rapid absorption of drugs across the mucosal layer is dependent on their liposolubility. Usually, drugs with high lipid solubility have good compatibility with the nasal mucosa and are more likely to be absorbed. However, the high fat solubility of a drug can also be detrimental as it can facilitate entry into the general circulation after absorption through the nasal cavity, which is not conducive for drug entry into the brain ( Chen et al., 2022 ; Tanna et al., 2023 ).

4.2.4 Drug viscosity

In general, drugs with an appropriate viscosity should be more likely to remain in contact with the nasal mucosa for longer periods of time, thereby improving the absorption efficiency ( Bseiso et al., 2022 ). However, a previous study that explored the effect of viscosity on absorption efficiency of metoclopramide hydrochloride drugs reported that although the retention time of a drug in the nasal cavity was prolonged as the viscosity increased, the absorption efficiency actually decreased ( Zaki et al., 2007 ). Most new drug carriers, such as microspheres, gels, and nanoparticles, have been developed with an appropriate drug viscosity in mind, improving both the retention time of drugs in the nasal cavity as well as the absorption efficiency.

4.3 Preparation-related factors

4.3.1 ph value.

Both the pH of a nasal preparation and the composition of the surface of the nasal cavity affect the dissolution, absorption, and penetration of a drug, with the optimal absorption occurring when a drug is in a non-dissipative state ( Yu et al., 2011 ). The pH of nasal secretions is approximately 5.5–6.5, and preparations with a pH too far outside this range will irritate the nasal cavity ( England et al., 1999 ). Ideally, the pH of the nasal cavity should be maintained at approximately 4.5–6.5 for optimal buffering.

4.3.2 Osmotic pressure

A change in osmotic pressure not only causes the contraction or relaxation of nasal epithelial cells and affects drug absorption, but it may also exacerbate nasal mucosal edema and reduce the beat frequency of ciliated cells in the nasal cavity ( Lemoine et al., 2005 ). Any drugs administered via the nasal cavity should be isotonic with the nasal mucosa (i.e., equivalent to 0.9% sodium chloride solution).

4.3.3 Formulation

Common formulations currently used for the nasal administration of drugs include nasal drops, sprays, powders, gels, microspheres, membranes, emulsions, liposomes, nanoparticles, and micelles ( Khan et al., 2017 ). The physical properties of a drug delivery system can affect the dissolution and retention time of drugs in the nasal cavity. Nasal sprays generally result in significantly higher bioavailability of a drug compared with that of drugs administered via nasal drops, which are quickly cleared by nasal cilia. In contrast, drugs administered via nasal sprays are mainly deposited within the front portion of the nasal cavity, with only a small proportion being slowly cleared into the throat, thereby extending the retention time of drugs in the nasal cavity, facilitating absorption, and improving the bioavailability ( Kumar et al., 2017 ). Powder-based formulations tend to result in stronger and more prolonged contact with the mucosa than solution-based formulations, leading to higher concentration gradients on both sides of the mucosa, thereby increasing drug absorption and brain bioavailability. Many other new dosage forms, such as gel or microsphere preparations, emulsions, liposomes, nanoparticles, and micelles are also conducive to improving the nasal absorption of drugs ( Marcello and Chiono, 2023 ). In summary, an appropriate dosage form should be selected based on the aforementioned specifications on a case-by-case basis depending on the properties of the drug being administered.

4.3.4 Effects of preservatives

Some commonly used preservatives in nasal preparations include benzalkonium chloride, ethylenediaminetetraacetic acid, ethylparaben, and thiomersal. Improper selection of a preservative can affect the absorption of drugs across the nasal mucosa ( Li et al., 2018 ). Lipophilic preservatives such as ethylparaben can reversibly accelerate or reduce the frequency of movement of ciliated cells in the nasal cavity, whereas polar preservatives, such as benzalkonium chloride, tend to reduce the movement of cilia in the nasal cavity.

4.3.5 Dosage

Typically, a higher dosage of a drug will lead to greater absorption and improve the efficacy. However, higher drug concentrations may irritate the nasal mucosa. Moreover, the volume of the nasal cavity is limited, and a dosage that is too high can lead to overflow from the nostrils or entry into the pharynx, causing discomfort. According to a previous study, the appropriate nasal administration volume is 0.05–0.15 mL, and the maximum volume should not exceed 0.20 mL ( Mittal et al., 2014 ).

4.3.6 Administration method

For humans, the method of administration primarily depends on the dosage form ( Krishnan et al., 2017 ). When using nasal drops, for example, the head should be tilted back at an appropriate angle before the medication is administered. Sprays should be directly administered into the nasal cavity, and a plaster must be applied to the inner wall of the nasal cavity. In experimental animals, medications are usually administered directly.

5 Methods to improve brain-targeted nasal mucosal drug delivery through the application of related biomaterials

5.1 penetration enhancers.

Following intranasal drug administration, the actual amount of drug absorbed through the nasal mucosa is very limited; therefore, a key research goal is to identify ways to increase this absorption. One means of achieving this goal is through the use of penetration enhancers, which increase the likelihood that drug molecules will be subsequently transport into the brain.

5.1.1 Pz-peptidase

Pz-peptidase has been shown to promote drug permeability, and the enzyme can instantly and reversibly induce the opening of tight junctions ( Chikuma et al., 1993 ).

5.1.2 Cell-penetrating peptides

Cell-penetrating peptides are short peptides that can penetrate the cell and/or nuclear membrane to guide any connected peptides, proteins, or other bioactive molecules into cells ( Kamei, 2017 ). Although the precise transmembrane mechanism of cell-penetrating peptides is not fully understood, the effects are likely mediated via signal transduction pathways and endocytosis ( Kamei et al., 2021 ).

5.1.3 Chitosan

The cations present in the amino groups of chitosan can bind to anions in the mucosa, thereby improving the permeability of epithelial cell membranes and promoting the opening of the tight junctions between epithelial cells and enhancing drug absorption ( Shim and Yoo, 2020 ). Chitosan functions as a mucosal adhesive and has a good safety profile, which has led to it being extensively investigated in recent years for its potential therapeutic applications ( Hard et al., 2023 ).

5.1.4 Cyclodextrins

Cyclodextrins are cyclic compounds composed of D-glucose molecules connected by 1,4-glycosidic bonds; they are water-soluble, non-reducing, molecules that take the form of a white crystalline powder ( Rassu et al., 2021 ). Commonly used α-, β-, and γ-cyclodextrins are composed of six, seven, and eight glucose molecules, respectively ( Papakyriakopoulou et al., 2021 ). The unique spatial structure of cyclodextrins allows them to form inclusion complexes with many substances, especially those that are lipophilic. Thus, cyclodextrins can be used as nasal mucosal absorption enhancers, solubilizers, or stabilizers to promote drug absorption, either directly or indirectly ( Papakyriakopoulou et al., 2021 ).

5.2 Mucosal adhesives

A significant limitation of the nasal route of administration is the rapid clearance of drug molecules mediated by mucosal fibers within the nasal cavity. Thus, the use of mucosal adhesives can inhibit the clearance functions of mucosal cilia, increasing the retention time of drugs at the mucosa and effectively enhancing both drug absorption and the bioavailability of drugs in the brain.

5.2.1 Chitosan

As described in several studies, various types of chitosan and hydroxypropyl methylcellulose can also be used as mucosal adhesives to facilitate nasal drug administration ( Nižić et al., 2020 ; Kamali et al., 2023 ).

5.2.2 Receptor–ligand interactions

To increase the nose-to-brain delivery of nanoscale drugs, selecting appropriate ligands for surface modification of the formulation can increase the affinity between the drug delivery system and the mucosa, leading to enhanced mucosal adsorption. The most commonly used targeting ligands are proteins whose receptors are expressed in the olfactory region, namely lactoferrin ( Li et al., 2023 ) or certain glycoproteins ( Pernet et al., 2023 ). Several lectins, such as wheat germ agglutinin ( Su et al., 2020 ) and solanum tuberosum lectin ( Chen et al., 2012 ), have also been used to promote nose-to-brain drug delivery.

5.3 New drug delivery systems

The identification of novel drug delivery systems capable of increasing the absorption of drugs across the nasal mucosa and promoting drug transport to the brain have become a research hotspot in the field of targeted drug delivery in the CNS. Some of these systems can achieve their effects without irritating the nasal mucosa, and many have a good safety profile.

5.3.1 Liposomes

Liposomes are novel dosage forms of targeted drug delivery systems. Liposomes are enclosed bilayer membranes of phospholipids (such as lecithin and cholesterol) that contain hydrophilic cores and possess the characteristics and functions of biofilms ( Duong et al., 2023 ). Lipid-soluble compounds can be embedded in phospholipid bilayer membranes, whereas water-soluble compounds can be encapsulated in hydrophilic parts; therefore, liposomes can carry both lipid- and water-soluble drugs ( Duong et al., 2023 ). Owing to their surface charge, liposomes increase the contact time with the mucosa and improve the bioavailability of drugs while also protecting encapsulated biomolecules from degradation by enzymes in the nasal mucosa ( Semyachkina-Glushkovskaya et al., 2022 ). Liposomes cause only slight or no damage to the nasal mucosa, without inducing irritation or ciliary toxicity, and they are suitable for long-term administration ( Semyachkina-Glushkovskaya et al., 2022 ).

5.3.2 In situ gel preparations

An in situ gel, also known as an in vivo gel, is a polymer that can undergo a phase change at the drug delivery site into a liquid or semi-solid form ( Tang et al., 2011 ). Such preparations can effectively extend the retention time of drugs in the nasal cavity and increase drug concentrations in brain tissue ( He et al., 2008 ). Depending on the factor that induces the phase change, in situ gels can be classified as temperature-, pH-, and ionic-type gels ( He et al., 2008 ). These gels have a highly hydrophilic three-dimensional network structure, which increases the absorption of drugs across the nasal mucosa by increasing the water permeability. The gels can induce their effects without damaging the mucosal surface because the drugs are transported through the paracellular pathway along with the flow of water ( Agrawal et al., 2020 ). Therefore, in recent decades, researchers have been focusing on optimizing the properties of these gels and improving the understanding of the processes through which they exert their effects.

5.3.3 Microsphere preparations

Microspheres are a new dosage form developed in recent years involving a particle dispersion system formed by drug dispersion and absorption in a polymer matrix ( Rassu et al., 2018 ). Microspheres are a spherical drug delivery system composed of various materials, including but not limited to, albumin, gelatin, polylactides, and starches ( Rassu et al., 2018 ). Microspheres act as strong bio-adhesives and can extend the retention time of drugs in the nasal cavity to 4 h ( Pandey et al., 2020 ). In addition, they can protect drugs from enzymatic metabolism, thereby greatly enhancing their bioavailability ( Pandey et al., 2020 ).

5.3.4 Emulsions

An emulsion is a heterogeneous dispersion system composed of two or more immiscible or partially miscible liquids ( Gizurarson, 2012 ). The particle size of a microemulsion ranges from 10 to 100 nm, they have a transparent appearance, and they contain surfactants and cosurfactants in a thermodynamically and dynamically stable system ( Froelich et al., 2021 ). Microemulsions are capable of carrying large amounts of lipophilic and water-soluble drugs while simultaneously protecting them from degradation, hydrolysis, and oxidation ( Froelich et al., 2021 ).

5.3.5 Nanoparticles

Nanoparticles are drug delivery systems with a particle size of 10–100 nm that are generated using polymer materials as carriers to adsorb or encapsulate drugs within the carrier material ( Antunes et al., 2023 ). Nanoparticles can protect encapsulated drugs from degradation, preventing their unintended removal from the nasal cavity, extending the residence time of drugs in the nasal mucosa, and promoting the direct transport of certain drugs from the nasal cavity to the brain, without increasing drug concentrations within the general circulation; thus, nanoparticles can help facilitate brain-targeted drug delivery ( Ferreira et al., 2023 ). Although the precise mechanism remains unknown, the transportation of nanoparticles to the brain may involve receptor-mediated endocytosis into cerebral capillary endothelial cells ( Montegiove et al., 2022 ).

6 Common experimental methods used in intranasal drug delivery research

Currently, there is no suitable in vitro method for evaluating brain-targeted drug delivery following nasal administration, and most existing evaluation methods are based on pharmacokinetic or pharmacodynamic techniques.

6.1 Cerebellomedullary cistern puncture (single-point puncture method)

The procedure for cerebellomedullary cistern puncture is as follows: after a certain period of drug, administration, the skin on the dorsal side of the head and neck of the mouse is cut open, the foramen magnum is exposed, and a syringe is inserted into the cerebellomedullary cistern to extract the CSF, from which the drug content is quantified ( Šakić, 2019 ). However, due to insufficient CSF supplementation post-extraction, normal intracranial pressure is difficult to maintain, as it is affected by the CSF volume. This method can only be used to determine the CSF concentration of a single mouse at a certain time point after drug administration, and it cannot be used to obtain complete data to characterize the changes in drug concentrations in the CSF over time. It is also difficult to distinguish differences in drug distribution within brain tissue using this method. Because of the need for a large number of animals for experimental purposes, this method has been utilized less frequently than others in brain-targeted drug delivery research ( Zhang et al., 2023 ).

6.2 Brain tissue homogenization method

In the brain tissue homogenization method, the whole brain of an animal is collected post-drug administration according to a predefined experimental timeline, the meninges and blood stains are removed, different brain tissues (such as the olfactory bulb and cerebellum) are separated, and the drug content is measured after weighing, homogenization, and sample pretreatment ( Onaolapo et al., 2023 ). This method allows researchers to assess the distribution of drugs in brain tissue at specific time points after drug administration; however, to minimize the impact of individual variability in experimental animals, a large sample size is generally required. Despite this limitation, this method remains one of the most widely used in experimental research ( Ramos et al., 2023 ).

6.3 Radionuclide labeling method

The radionuclide labeling method uses isotope labeling to quantify drug content in tissue after administration, making it a suitable technique for studying the brain distribution of peptides or proteins after nasal administration ( Elsharkawy et al., 2023 ). This method allows for rapid detection with high sensitivity, and it does not require tedious drug extraction steps after tissue homogenization, reducing the risk of experimental errors. However, this method cannot distinguish between raw materials, degradation products, and conjugates, making it impossible to determine the true concentration of a drug from the measured total radioactivity ( Janssen et al., 2023 ).

6.4 Brain microdialysis method

The microdialysis method is a rapidly evolving in vitro or in vivo brain chemistry samplingtechnique developed in recent years; it offers good temporal and spatial resolution for determining the concentration of free drugs in the CNS. Compared to traditional research methods, it does not alter the total amount of CSF; thus, it can be performed without affecting the normal physiological function of experimental animals, and it allows for continuous sampling and quantification in a single animal ( Zhu et al., 2023 ). The technique can also be used to determine the concentration of drugs in different brain tissues as well as changes in the concentration of non-binding drugs in the brain’s extracellular fluid over time. This technology has become indispensable in the study of brain-targeting drugs and their active pharmacological effects in the CNS, and it is suitable for in vivo biochemical and deep brain region research ( Venturini et al., 2023 ). However, this method has high instrument requirements and costs, and it is not conducive to large-scale experiments.

6.5 Pharmacodynamic evaluation method

When a drug concentration is difficult to measure, its known pharmacological effects can be assessed to indirectly infer the degree of drug absorption into the brain through the nasal mucosa.

7 Brain-targeted nasal drug delivery for the prevention and treatment of CNS diseases

Brain-targeted nasal drug delivery has broad application prospects in the prevention and treatment of CNS diseases. By administering drugs through the nasal cavity, drugs can directly enter the brain, improve their efficacy, and reduce adverse reactions. Meanwhile, brain-targeted nasal drug delivery can also improve patients’ cognitive function, emotional state, and quality of life. Therefore, brain-targeted nasal drug delivery will become one of the important means for the prevention and treatment of CNS diseases in the future. The application of brain-targeted nasal drug delivery in different CNS diseases is summarized as follows (See Table 1 ).

Table 1. The application of brain-targeted nasal drug delivery in CNS diseases.

8 Discussion and conclusion

Many scholars have conducted studies on the brain-targeting effects of nasally administered drugs, with aims that have ranged from identifying the mechanism of action to improving drug dosage forms. However, it is still necessary to rationally assess nasal administration systems, as the following are some important limitations that have yet to be overcome: 1) Low targeting efficiency is currently the biggest and most common problem in brain-targeted drug delivery research. The efficiency of such targeted drug delivery systems is usually only several times higher than that of clinically administered drugs or conventional non-targeted drug delivery systems. 2) At present, brain targeting studies mainly focus on ways to improve drug delivery to the brain, with fewer studies focusing on how drugs are distributed after they enter the brain. The structure of brain tissue is complex, and the organ is central to the control of most bodily functions, making it difficult to investigate drug concentrations in various regions in humans. The ability to maximize the concentration of a therapeutic drug at the site of a specific lesion after entry into the brain is of great clinical significance. 3) The use of excipients, receptors, or carriers to enhance drug delivery to the brain can inadvertently affect the metabolism and distribution of endogenous substances, and long-term use of certain medications may cause serious toxic side effects. 4) There are significant anatomical differences between experimental animals and humans, and the results of studies based on animal models may differ significantly from those of studies involving humans. For example, in experimental studies, animals are often anesthetized, which results in low sensitivity to nasal mucosal stimulation and enhanced drug absorption. 5) The effectiveness of brain targeted nasal administration in treating central nervous system diseases is easily influenced by the experimental animal model and the age of the experimental animals. 6) The physiological condition of the nasal cavity has a significant impact on the absorption of drugs. Nevertheless, nasal administration of drugs in order to bypass the BBB and facilitate direct entry of a drug into the CSF or brain tissue is still a safe, convenient, and non-destructive method of targeted drug delivery to the CNS. To optimize the clinical utility of brain-targeted nasal drug administration strategies, it will be necessary to further increase the volume of drugs that can be transported through the nose into the brain, to reduce the significant variability in drug absorption that is caused by physiological changes in the nasal cavity, and to reduce the irritation and long-term toxicity of nasal preparations.

Author contributions

QH: Writing – original draft. XC: Writing – review and editing. SY: Methodology, Supervision, Writing – review and editing. GG: Writing – review and editing. HS: Funding acquisition, Writing – review and editing.

The author(s) declare financial support was received for the research, authorship, and/or publication of this article. This study was supported by the Joint key project (grant number 2019LH01), the Department of Science and Technology of Sichuan Province (grant numbers 2022JDRC0041 and 22CXRC0178), the Medical Innovation Project (grant number 21WQ040), and the Hospital Management Project of the General Hospital of the Western Theater Command (grant numbers 2021-XZYG-C25, 2021-XZYG-B22, and 2021-XZYG-B21).

Conflict of interest

The authors declare that the research was conducted in the absence of any commercial or financial relationships that could be construed as a potential conflict of interest.

Publisher’s note

All claims expressed in this article are solely those of the authors and do not necessarily represent those of their affiliated organizations, or those of the publisher, the editors and the reviewers. Any product that may be evaluated in this article, or claim that may be made by its manufacturer, is not guaranteed or endorsed by the publisher.

Abbreviations

BBB, blood–brain barrier; CNS, central nervous system; CSF, cerebrospinal fluid; P-gp, P-glycoprotein.

Abla, K. K., and Mehanna, M. M. (2023). The battle of lipid-based nanocarriers against blood-brain barrier: A critical review. J. Drug Target. 31, 832–857. doi: 10.1080/1061186X.2023.2247583

PubMed Abstract | Crossref Full Text | Google Scholar

Agrawal, M., Saraf, S., Saraf, S., Dubey, S. K., Puri, A., Gupta, U., et al. (2020). Stimuli-responsive In situ gelling system for nose-to-brain drug delivery. J. Control Release 327, 235–265. doi: 10.1016/j.jconrel.2020.07.044

Antunes, J. L., Amado, J., Veiga, F., Paiva-Santos, A. C., and Pires, P. C. (2023). Nanosystems, drug molecule functionalization and intranasal delivery: An update on the most promising strategies for increasing the therapeutic efficacy of antidepressant and anxiolytic drugs. Pharmaceutics 15:998. doi: 10.3390/pharmaceutics15030998

Appu, A. P., Arun, P., Krishnan, J. K., Moffett, J. R., and Namboodiri, A. M. (2016). Rapid intranasal delivery of chloramphenicol acetyltransferase in the active form to different brain regions as a model for enzyme therapy in the CNS. J. Neurosci. Methods 259, 129–134. doi: 10.1016/j.jneumeth.2015.11.027

Awad, R., Avital, A., and Sosnik, A. (2023). Polymeric nanocarriers for nose-to-brain drug delivery in neurodegenerative diseases and neurodevelopmental disorders. Acta Pharm. Sin. B 13, 1866–1886. doi: 10.1016/j.apsb.2022.07.003

Babu, S. R., Shekara, H. H., Sahoo, A. K., Harsha Vardhan, P. V., Thiruppathi, N., and Venkatesh, M. P. (2023). Intranasal nanoparticulate delivery systems for neurodegenerative disorders: A review. Ther. Deliv. 14, 571–594. doi: 10.4155/tde-2023-0019

Badran, M. M., Alanazi, A. E., Ibrahim, M. A., Alshora, D. H., Taha, E., and Alomrani, A. H. (2023). Optimization of bromocriptine-mesylate-loaded polycaprolactone nanoparticles coated with chitosan for nose-to-brain delivery: In vitro and in vivo studies. Polymers 15:3890. doi: 10.3390/polym15193890

Bonaccorso, A., Gigliobianco, M. R., Lombardo, R., Pellitteri, R., Di Martino, P., Mancuso, A., et al. (2023). Nanonized carbamazepine for nose-to-brain delivery: Pharmaceutical formulation development. Pharm. Dev. Technol. 28, 248–263. doi: 10.1080/10837450.2023.2177673

Bors, L. A., Bajza, Á, Mándoki, M., Tasi, B. J., Cserey, G., Imre, T., et al. (2020). Modulation of nose-to-brain delivery of a P-glycoprotein (MDR1) substrate model drug (quinidine) in rats. Brain Res. Bull. 160, 65–73. doi: 10.1016/j.brainresbull

Crossref Full Text | Google Scholar

Boyuklieva, R., Zagorchev, P., and Pilicheva, B. (2023). Computational, in vitro, and in vivo models for nose-to-brain drug delivery studies. Biomedicines 11:2198. doi: 10.3390/biomedicines11082198

Bseiso, E. A., AbdEl-Aal, S. A., Nasr, M., Sammour, O. A., and El Gawad, N. A. (2022). Nose to brain delivery of melatonin lipidic nanocapsules as a promising post-ischemic neuroprotective therapeutic modality. Drug Deliv. 29, 2469–2480. doi: 10.1080/10717544.2022.2104405

Chen, J., Zhang, C., Liu, Q., Shao, X., Feng, C., Shen, Y., et al. (2012). Solanum tuberosum lectin-conjugated PLGA nanoparticles for nose-to-brain delivery: In vivo and in vitro evaluations. J. Drug Target 20, 174–184. doi: 10.3109/1061186X.2011.622396

Chen, S., Xie, Q., Yang, M., Shi, Y., Shi, J., and Zeng, X. (2022). Scutellaria baicalensis extract-phospholipid complex: Preparation and initial pharmacodynamics research in rats. Curr. Pharm. Biotechnol. 23, 847–860. doi: 10.2174/1389201022666210729142257

Chen, S., Zhang, J., Wu, L., Wu, H., and Dai, M. (2018). Paeonol nanoemulsion for enhanced oral bioavailability: Optimization and mechanism. Nanomedicine 13, 269–282. doi: 10.2217/nnm-2017-0277

Chikuma, T., Tanaka, W., Yamada, K., Ishii, Y., Tanaka, A., and Kato, T. (1993). High-performance liquid chromatographic determination of PZ-peptidase activity. J. Chromatogr. 635, 81–87. doi: 10.1016/0021-9673(93)83117-b

Chung, S., Peters, J. M., Detyniecki, K., Tatum, W., Rabinowicz, A. L., and Carrazana, E. (2022). The nose has it: Opportunities and challenges for intranasal drug administration for neurologic conditions including seizure clusters. Epilepsy Behav. Rep. 21:100581. doi: 10.1016/j.ebr.2022.100581

Crowe, T. P., Greenlee, M. H., Kanthasamy, A. G., and Hsu, W. H. (2018). Mechanism of intranasal drug delivery directly to the brain. Life Sci. 195, 44–52. doi: 10.1016/j.lfs.2017.12.025

De Gaetano, F., d’Avanzo, N., Mancuso, A., De Gaetano, A., Paladini, G., Caridi, F., et al. (2022). Chitosan/cyclodextrin nanospheres for potential nose-to-brain targeting of idebenone. Pharmaceuticals 15:1206. doi: 10.3390/ph15101206

Du, L., Chen, L., Liu, F., Wang, W., and Huang, H. (2023). Nose-to-brain drug delivery for the treatment of CNS disease: New development and strategies. Int. Rev. Neurobiol. 171, 255–297. doi: 10.1016/bs.irn.2023.05.014

Duong, V. A., Nguyen, T. T., and Maeng, H. J. (2023). Recent advances in intranasal liposomes for drug, gene, and vaccine delivery. Pharmaceutics 15:207. doi: 10.3390/pharmaceutics15010207

Elkomy, M. H., Abo El-Ela, F., Zaki, R. M., Alsaidan, O. A., Elmowafy, M., Zafar, A., et al. (2023). Intranasal nanotransferosomal gel for quercetin brain targeting: II. Antidepressant effect in an experimental animal model. Pharmaceutics 15:2095. doi: 10.3390/pharmaceutics15082095

ElShagea, H. N., Makar, R. R., Salama, A. H., Elkasabgy, N. A., and Basalious, E. B. (2023). Investigating the targeting power to brain tissues of intranasal rasagiline mesylate-loaded transferosomal in situ gel for efficient treatment of Parkinson’s disease. Pharmaceutics 15:533. doi: 10.3390/pharmaceutics15020533

Elsharkawy, F. M., Amin, M. M., Shamsel-Din, H. A., Ibrahim, W., Ibrahim, A. B., and Sayed, S. (2023). Self-assembling lecithin-based mixed polymeric micelles for nose to brain delivery of clozapine: In-vivo assessment of drug efficacy via radiobiological evaluation. Int. J. Nanomed. 18, 1577–1595. doi: 10.2147/IJN.S403707

Engelhardt, B., and Sorokin, L. (2009). The blood-brain and the blood-cerebrospinal fluid barriers: Function and dysfunction. Semin. Immunopathol. 31, 497–511. doi: 10.1007/s00281-009-0177-0

England, R. J., Homer, J. J., Knight, L. C., and Ell, S. R. (1999). Nasal pH measurement: A reliable and repeatable parameter. Clin. Otolaryngol. Allied Sci. 24, 67–68. doi: 10.1046/j.1365-2273.1999.00223.x

Ferreira, M. D., Duarte, J., Veiga, F., Paiva-Santos, A. C., and Pires, P. C. (2023). Nanosystems for brain targeting of antipsychotic drugs: An update on the most promising nanocarriers for increased bioavailability and therapeutic efficacy. Pharmaceutics 15:678. doi: 10.3390/pharmaceutics15020678

Fonseca-Santos, B., Cazarin, C. A., da Silva, P. B., Dos Santos, K. P., da Rocha, M. C., Báo, S. N., et al. (2023). Intranasal in situ gelling liquid crystal for delivery of resveratrol ameliorates memory and neuroinflammation in Alzheimer’s disease. Nanomedicine 51:102689. doi: 10.1016/j.nano.2023.102689

Fortuna, A., Alves, G., Serralheiro, A., Sousa, J., and Falcão, A. (2014). Intranasal delivery of systemic-acting drugs: Small-molecules and biomacromolecules. Eur. J. Pharm. Biopharm. 88, 8–27. doi: 10.1016/j.ejpb.2014.03.004

Froelich, A., Osmałek, T., Jadach, B., Puri, V., and Michniak-Kohn, B. (2021). Microemulsion-based media in nose-to-brain drug delivery. Pharmaceutics 13:201. doi: 10.3390/pharmaceutics13020201

Fuchigami, T., Itokazu, Y., Morgan, J. C., and Yu, R. K. (2023). Restoration of adult neurogenesis by intranasal administration of gangliosides GD3 and GM1 in The olfactory bulb of A53T alpha-synuclein-expressing Parkinson’s-disease model mice. Mol. Neurobiol. 60, 3329–3344. doi: 10.1007/s12035-023-03282-2

Gänger, S., and Schindowski, K. (2018). Tailoring formulations for intranasal nose-to-brain delivery: A review on architecture, physico-chemical characteristics and mucociliary clearance of the nasal olfactory mucosa. Pharmaceutics 10:116. doi: 10.3390/pharmaceutics10030116

Gizurarson, S. (2012). Anatomical and histological factors affecting intranasal drug and vaccine delivery. Curr. Drug Deliv. 9, 566–582. doi: 10.2174/156720112803529828

Hard, S. A., Shivakumar, H. N., and Redhwan, M. A. (2023). Development and optimization of in-situ gel containing chitosan nanoparticles for possible nose-to-brain delivery of vinpocetine. Int. J. Biol. Macromol. 253:127217. doi: 10.1016/j.ijbiomac.2023.127217

He, C., Kim, S. W., and Lee, D. S. (2008). In situ gelling stimuli-sensitive block copolymer hydrogels for drug delivery. J. Control Release 127, 189–207. doi: 10.1016/j.jconrel.2008.01.005

Hope, J., Copolov, D., Tiller, J., Galbally, M., Hopwood, M., Newton, R., et al. (2023). What clinicians need to know about intranasal esketamine for treatment-resistant depression? Australas Psychiatry 31, 841–845. doi: 10.1177/10398562231211171

Hu, X., Yue, X., Wu, C., and Zhang, X. (2023). Factors affecting nasal drug delivery and design strategies for intranasal drug delivery. Zhejiang Da Xue Xue Bao Yi Xue Ban. 52, 328–337. doi: 10.3724/zdxbyxb-2023-0069

Hussain, A., Altamimi, M. A., Ramzan, M., Mirza, M. A., and Khuroo, T. (2023). GastroPlus- and HSPiP-oriented predictive parameters as the basis of valproic acid-loaded mucoadhesive cationic nanoemulsion gel for improved nose-to-brain delivery to control convulsion in humans. Gels 9:603. doi: 10.3390/gels9080603

Illum, L. (2000). Transport of drugs from the nasal cavity to the central nervous system. Eur. J. Pharm. Sci. 11, 1–18. doi: 10.1016/s0928-0987(00)00087-7

Illum, L. (2003). Nasal drug delivery–possibilities, problems and solutions. J. Control Release 87, 187–198. doi: 10.1016/s0168-3659(02)00363-2

Illum, L. (2004). Is nose-to-brain transport of drugs in man a reality? J. Pharm. Pharmacol. 56, 3–17. doi: 10.1211/0022357022539

Janssen, B., Tian, G., Lengyel-Zhand, Z., Hsieh, C. J., Lougee, M. G., Riad, A., et al. (2023). Identification of a putative α-synuclein radioligand using an in silico similarity Search. Mol. Imaging Biol. 25, 704–719. doi: 10.1007/s11307-023-01814-9

Kadry, H., Noorani, B., and Cucullo, L. (2020). A blood-brain barrier overview on structure, function, impairment, and biomarkers of integrity. Fluids Barriers CNS 17:69. doi: 10.1186/s12987-020-00230-3

Kamali, H., Tafaghodi, M., Eisvand, F., Ahmadi-Soleimani, S., Khajouee, M., Ghazizadeh, H., et al. (2023). Preparation and evaluation of the in situ gel-forming chitosan hydrogels for nasal delivery of morphine in a single unit dose in rats to enhance the analgesic responses. Curr. Drug Deliv. doi: 10.2174/1567201820666230724161205 [Epub ahead of print].

Kamei, N. (2017). [Nose-to-brain delivery of peptide drugs enhanced by coadministration of cell-penetrating peptides: Therapeutic potential for dementia]. Yakugaku Zasshi 137, 1247–1253. doi: 10.1248/yakushi.17-00138

Kamei, N., Suwabe, S., Arime, K., Bando, H., Murata, K., Yamaguchi, M., et al. (2021). Investigation of the transport pathways associated with enhanced brain delivery of peptide drugs by intranasal coadministration with penetratin. Pharmaceutics 13:1745. doi: 10.3390/pharmaceutics13111745

Kannavou, M., Karali, K., Katsila, T., Siapi, E., Marazioti, A., Klepetsanis, P., et al. (2023). Development and comparative in vitro and in vivo study of BNN27 mucoadhesive liposomes and nanoemulsions for nose-to-brain delivery. Pharmaceutics 15:419. doi: 10.3390/pharmaceutics15020419

Khan, A. R., Liu, M., Khan, M. W., and Zhai, G. (2017). Progress in brain targeting drug delivery system by nasal route. J. Control Release 268, 364–389. doi: 10.1016/j.jconrel.2017.09.001

Khatri, D. K., Preeti, K., Tonape, S., Bhattacharjee, S., Patel, M., Shah, S., et al. (2023). Nanotechnological advances for nose to brain delivery of therapeutics to improve the Parkinson therapy. Curr. Neuropharmacol. 21, 493–516. doi: 10.2174/1570159X20666220507022701

Krishnan, J. K., Arun, P., Chembukave, B., Appu, A. P., Vijayakumar, N., Moffett, J. R., et al. (2017). Effect of administration method, animal weight and age on the intranasal delivery of drugs to the brain. J. Neurosci. Methods 286, 16–21. doi: 10.1016/j.jneumeth

Kumar, H., Mishra, G., Sharma, A. K., Gothwal, A., Kesharwani, P., and Gupta, U. (2017). Intranasal drug delivery: A non-invasive approach for the better delivery of neurotherapeutics. Pharm. Nanotechnol. 5, 203–214. doi: 10.2174/2211738505666170515113936

Kumar, N. N., Gautam, M., Lochhead, J. J., Wolak, D. J., Ithapu, V., Singh, V., et al. (2016). Relative vascular permeability and vascularity across different regions of the rat nasal mucosa: Implications for nasal physiology and drug delivery. Sci. Rep. 6:31732. doi: 10.1038/srep31732

Ladel, S., Schlossbauer, P., Flamm, J., Luksch, H., Mizaikoff, B., and Schindowski, K. (2019). Improved in vitro model for intranasal mucosal drug delivery: Primary olfactory and respiratory epithelial cells compared with the permanent nasal cell line RPMI 2650. Pharmaceutics 11:367. doi: 10.3390/pharmaceutics11080367

Landis, M. S., Boyden, T., and Pegg, S. (2012). Nasal-to-CNS drug delivery: Where are we now and where are we heading? An industrial perspective. Ther. Deliv. 3, 195–208. doi: 10.4155/tde.11.149

Larijani, S. F., Hassanzadeh, G., Zahmatkesh, M., Radfar, F., and Farahmandfar, M. (2023). Intranasal insulin intake and exercise improve memory function in amyloid-β induced Alzheimer’s-like disease in rats: Involvement of hippocampal BDNF-TrkB receptor. Behav Brain Res. 460:114814. doi: 10.1016/j.bbr.2023.114814

Lemoine, J. L., Farley, R., and Huang, L. (2005). Mechanism of efficient transfection of the nasal airway epithelium by hypotonic shock. Gene Ther. 12, 1275–1282. doi: 10.1038/sj.gt.3302548

Li, L., Tan, L., Zhang, Q., Cheng, Y., Liu, Y., Li, R., et al. (2023). Nose-to-brain delivery of self-assembled curcumin-lactoferrin nanoparticles: Characterization, neuroprotective effect and in vivo pharmacokinetic study. Front. Bioeng. Biotechnol. 11:1168408. doi: 10.3389/fbioe.2023.1168408

Li, Y., Fan, X., Li, W., Yang, P., Zhang, H., Tang, D., et al. (2018). Metoclopramide nasal spray in vitro evaluation and in vivo pharmacokinetic studies in dogs. Pharm. Dev. Technol. 23, 275–281. doi: 10.1080/10837450.2017.1316734

Marcello, E., and Chiono, V. (2023). Biomaterials-enhanced intranasal delivery of drugs as a direct route for brain targeting. Int. J. Mol. Sci. 24:3390. doi: 10.3390/ijms24043390

Meirinho, S., Rodrigues, M., Santos, A. O., Falcão, A., and Alves, G. (2023). Nose-to-brain delivery of perampanel formulated in a self-microemulsifying drug delivery system improves anticonvulsant and anxiolytic activity in mice. Int. J. Pharm. 642:123145. doi: 10.1016/j.ijpharm.2023.123145

Michaelian, J. C., McCade, D., Hoyos, C. M., Brodaty, H., Harrison, F., Henry, J. D., et al. (2023). Pilot randomized, double-blind, placebo-controlled crossover trial evaluating the feasibility of an intranasal oxytocin in improving social cognition in individuals living with Alzheimer’s disease. J. Alzheimers Dis. Rep. 7, 715–729. doi: 10.3233/ADR-230013

Misilimu, D., Li, W., Chen, D., Wei, P., Huang, Y., Li, S., et al. (2022). Intranasal salvinorin A improves long-term neurological function via immunomodulation in a mouse ischemic stroke model. J. Neuroimmune Pharmacol. 17, 350–366. doi: 10.1007/s11481-021-10025-4

Mittal, D., Ali, A., Md, S., Baboota, S., Sahni, J. K., and Ali, J. (2014). Insights into direct nose to brain delivery: Current status and future perspective. Drug Deliv. 21, 75–86. doi: 10.3109/10717544.2013.838713

Mohammad, Khan, U., Saifi, Z., Bora, J., Warsi, M. H., Abourehab, M. A., et al. (2023). Intranasal inorganic cerium oxide nanoparticles ameliorate oxidative stress induced motor manifestations in haloperidol-induced parkinsonism. Inflammopharmacology 31, 2571–2585. doi: 10.1007/s10787-023-01274-1

Montegiove, N., Calzoni, E., Emiliani, C., and Cesaretti, A. (2022). Biopolymer nanoparticles for nose-to-brain drug delivery: A new promising approach for the treatment of neurological diseases. J. Funct. Biomater. 13:125. doi: 10.3390/jfb13030125

Nguyen, T. T., Dung Nguyen, T. T., Vo, T. K., Tran, N. M., Nguyen, M. K., Van Vo, T., et al. (2021). Nanotechnology-based drug delivery for central nervous system disorders. Biomed. Pharmacother. 143, 112–117. doi: 10.1016/j.biopha.2021.112117

Nižić, L., Potaś, J., Winnicka, K., Szekalska, M., Erak, I., Gretić, M., et al. (2020). Development, characterisation and nasal deposition of melatonin-loaded pectin/hypromellose microspheres. Eur. J. Pharm. Sci. 141:105115. doi: 10.1016/j.ejps.2019.105115

Oliveira, P., Fortuna, A., Alves, G., and Falcao, A. (2016). Drug-metabolizing enzymes and efflux transporters in nasal epithelium: Influence on the bioavailability of intranasally administered drugs. Curr. Drug Metab. 17, 628–647. doi: 10.2174/1389200217666160406120509

Onaolapo, O. J., Olofinnade, A. T., Ojo, F. O., Falade, J., and Onaolapo, A. (2023). Prepubertal continuous dietary folate fortification enhances the brain function of adult mice by modulating antioxidant status, inflammation, and brain neurotransmitter levels. Antiinflamm. Antiallergy Agents Med. Chem. 22, 198–209. doi: 10.2174/0118715230249814230925060325

Pailla, S. R., Sampathi, S., Junnuthula, V., Maddukuri, S., Dodoala, S., and Dyawanapelly, S. (2022). Brain-targeted intranasal delivery of zotepine microemulsion: Pharmacokinetics and pharmacodynamics. Pharmaceutics 14:978. doi: 10.3390/pharmaceutics14050978

Palmer, A. M. (2010). The role of the blood-CNS barrier in CNS disorders and their treatment. Neurobiol. Dis. 37, 3–12. doi: 10.1016/j.nbd.2009.07.029

Pandey, J., Shankar, R., Kumar, M., Shukla, K., and Kumari, B. (2020). Development of nasal mucoadhesive microspheres of granisetron: A potential drug. Drug Res. 70:367. doi: 10.1055/a-1193-4781

Papakyriakopoulou, P., Manta, K., Kostantini, C., Kikionis, S., Banella, S., Ioannou, E., et al. (2021). Nasal powders of quercetin-β-cyclodextrin derivatives complexes with mannitol/lecithin microparticles for Nose-to-Brain delivery: In vitro and ex vivo evaluation. Int. J. Pharm. 607:121016. doi: 10.1016/j.ijpharm.2021.121016

Pardeshi, C. V., and Belgamwar, V. S. (2013). Direct nose to brain drug delivery via integrated nerve pathways bypassing the blood-brain barrier: An excellent platform for brain targeting. Expert Opin. Drug Deliv. 10, 957–972. doi: 10.1517/17425247.2013.790887

Pernet, V., Joly, S., Spiegel, S., Meli, I., Idriss, S., Maigler, F., et al. (2023). Nogo-A antibody delivery through the olfactory mucosa mitigates experimental autoimmune encephalomyelitis in the mouse CNS. Cell Death Discov. 9:290. doi: 10.1038/s41420-023-01588-7

Pires, P. C., Fazendeiro, A. C., Rodrigues, M., Alves, G., and Santos, A. O. (2021). Nose-to-brain delivery of phenytoin and its hydrophilic prodrug fosphenytoin combined in a microemulsion - formulation development and in vivo pharmacokinetics. Eur. J. Pharm. Sci. 164:105918. doi: 10.1016/j.ejps.2021.105918

Ramos, A. C., Rezende, K. R., Teixeira, C. M., Fernandes, A. R., Santos, H., Machado, R. D., et al. (2023). Bioanalytical LC-QTOF/MS Method for a N-phenylpiperazine Derivate (LQFM05): An anxiolytic- and antidepressant-like prototype drug applied to pharmacokinetic and biodistribution studies. Pharmaceuticals 16:930. doi: 10.3390/ph16070930

Rassu, G., Ferraro, L., Pavan, B., Giunchedi, P., Gavini, E., and Dalpiaz, A. (2018). The role of combined penetration enhancers in nasal microspheres on in vivo drug bioavailability. Pharmaceutics 10:206. doi: 10.3390/pharmaceutics10040206

Rassu, G., Sorrenti, M., Catenacci, L., Pavan, B., Ferraro, L., Gavini, E., et al. (2021). Versatile nasal application of cyclodextrins: Excipients and/or actives? Pharmaceutics 13:1180. doi: 10.3390/pharmaceutics13081180

Ruigrok, M. J., and de Lange, E. C. (2015). Emerging insights for translational pharmacokinetic and pharmacokinetic-pharmacodynamic studies: Towards prediction of nose-to-brain transport in humans. AAPS J. 17, 493–505. doi: 10.1208/s12248-015-9724-x

Sabry, S. A., Abd El Razek, A. M., Nabil, M., Khedr, S. M., El-Nahas, H. M., and Eissa, N. G. (2023). Brain-targeted delivery of Valsartan using solid lipid nanoparticles labeled with Rhodamine B; a promising technique for mitigating the negative effects of stroke. Drug Deliv. 30:2179127. doi: 10.1080/10717544.2023.2179127

Saha, P., Singh, P., Kathuria, H., Chitkara, D., and Pandey, M. M. (2023). Self-assembled lecithin-chitosan nanoparticles improved rotigotine nose-to-brain delivery and brain targeting efficiency. Pharmaceutics 5:851. doi: 10.3390/pharmaceutics15030851

Šakić, B. (2019). Cerebrospinal fluid collection in laboratory mice: Literature review and modified cisternal puncture method. J. Neurosci. Methods 311, 402–407. doi: 10.1016/j.jneumeth.2018.09.025

Schaefer, M. L., Böttger, B., Silver, W. L., and Finger, T. E. (2002). Trigeminal collaterals in the nasal epithelium and olfactory bulb: A potential route for direct modulation of olfactory information by trigeminal stimuli. J. Comp. Neurol. 444, 221–226. doi: 10.1002/cne.10143

Semyachkina-Glushkovskaya, O., Shirokov, A., Blokhina, I., Telnova, V., Vodovozova, E., Alekseeva, A., et al. (2022). Intranasal delivery of liposomes to glioblastoma by photostimulation of the lymphatic system. Pharmaceutics 15:36. doi: 10.3390/pharmaceutics15010036

Shaikh, R. G., Ramanujan, B., Singh, R. K., Vibha, D., Mehta, S., Appukuttan, R., et al. (2022). The IN-MIDAZ study - Intranasal midazolam in aborting seizures - An epilepsy monitoring unit based randomized controlled trial for efficacy. Epilepsy Res. 188:107037. doi: 10.1016/j.eplepsyres.2022

Shim, S., and Yoo, H. S. (2020). The application of mucoadhesive chitosan nanoparticles in nasal drug delivery. Mar. Drugs 18:605. doi: 10.3390/md18120605

Shrewsbury, S. B. (2023). The upper nasal space: Option for systemic drug delivery, mucosal vaccines and “Nose-to-Brain”. Pharmaceutics 15:1720. doi: 10.3390/pharmaceutics15061720

Silva, S., Fonseca, C., Bicker, J., Falcão, A., and Fortuna, A. (2023). Intranasal administration of sertraline ensures sustained brain delivery and antidepressant effect in a mouse model of depression. Eur. J. Pharm. Biopharm. doi: 10.1016/j.ejpb.2023.12.002 [Epub ahead of print].

Singh, S., and Shukla, R. (2023). Nanovesicular-mediated intranasal drug therapy for neurodegenerative disease. AAPS PharmSciTech 24:179. doi: 10.1208/s12249-023-02625-5

Su, Y., Sun, B., Gao, X., Dong, X., Fu, L., Zhang, Y., et al. (2020). Intranasal delivery of targeted nanoparticles loaded with miR-132 to brain for the treatment of neurodegenerative diseases. Front. Pharmacol. 11:1165. doi: 10.3389/fphar.2020.01165

Tang, Y., Heaysman, C. L., Willis, S., and Lewis, A. L. (2011). Physical hydrogels with self-assembled nanostructures as drug delivery systems. Expert Opin. Drug Deliv. 8, 1141–1159. doi: 10.1517/17425247.2011.588205

Tanna, V., Sawarkar, S. P., and Ravikumar, P. (2023). Exploring nose to brain nano delivery for effective management of migraine. Curr. Drug Deliv. 20, 144–157. doi: 10.2174/1567201819666220401091632

Thorne, R. G., Pronk, G. J., Padmanabhan, V., and Frey, W. II (2004). Delivery of insulin-like growth factor-I to the rat brain and spinal cord along olfactory and trigeminal pathways following intranasal administration. Neuroscience 127, 481–496. doi: 10.1016/j.neuroscience.2004.05.029

Venturini, S., Bhatti, F., Timofeev, I., Carpenter, K. L., Hutchinson, P. J., Guilfoyle, M. R., et al. (2023). Microdialysis-based classifications of abnormal metabolic states after traumatic brain injury: A systematic review of the literature. J. Neurotrauma 40, 195–209. doi: 10.1089/neu.2021.0502

Wolburg, H., Wolburg-Buchholz, K., Sam, H., Horvát, S., Deli, M. A., and Mack, A. F. (2008). Epithelial and endothelial barriers in the olfactory region of the nasal cavity of the rat. Histochem. Cell Biol. 130, 127–140. doi: 10.1007/s00418-008-0410-2

Wu, D. D., Salah, Y. A., Ngowi, E. E., Zhang, Y. X., Khattak, S., Khan, N. H., et al. (2023). Nanotechnology prospects in brain therapeutics concerning gene-targeting and nose-to-brain administration. iScience 26:107321. doi: 10.1016/j.isci.2023.107321

Xinchen, Y., Jing, T., and Jiaoqiong, G. (2023). Lipid-based nanoparticles via nose-to-brain delivery: A mini review. Front. Cell Dev. Biol. 11:1214450. doi: 10.3389/fcell.2023.1214450

Xue, S., Zhou, X., Yang, Z. H., Si, X. K., and Sun, X. (2023). Stroke-induced damage on the blood-brain barrier. Front. Neurol. 14:1248970. doi: 10.3389/fneur.2023.1248970

Yokel, R. A. (2022). Direct nose to the brain nanomedicine delivery presents a formidable challenge. Wiley Interdiscip. Rev. Nanomed. Nanobiotechnol. 14:e1767. doi: 10.1002/wnan.1767

Yu, C., Gu, P., Zhang, W., Qi, N., Cai, C., He, H., et al. (2011). Preparation and evaluation of zolmitriptan submicron emulsion for rapid and effective nasal absorption in beagle dogs. Drug Dev. Ind. Pharm. 37, 1509–1516. doi: 10.3109/03639045.2011.587432

Zaki, N. M., Awad, G. A., Mortada, N. D., and Abd Elhady, S. S. (2007). Enhanced bioavailability of metoclopramide HCl by intranasal administration of a mucoadhesive in situ gel with modulated rheological and mucociliary transport properties. Eur. J. Pharm. Sci. 32, 296–307. doi: 10.1016/j.ejps.2007.08.006

Zhang, L. L., Yuan, Y., Liang, M. Y., Liu, M. X., Wang, D. X., Xie, J. X., et al. (2023). [A glass micropipette vacuum technique of cerebrospinal fluid sampling in C57BL/6 mice]. Sheng Li Xue Bao 75, 197–204.

Google Scholar

Zhou, X., Deng, X., Liu, M., He, M., Long, W., Xu, Z., et al. (2023). Intranasal delivery of BDNF-loaded small extracellular vesicles for cerebral ischemia therapy. J. Control Release 357, 1–19. doi: 10.1016/j.jconrel

Zhu, M., Liu, G., Chen, H., Ma, W., Cui, L., Nimmu, N. V., et al. (2023). Analytical strategies in neurotransmitter measurements: A mini literature review. Biomed. Chromatogr. 37:e5521. doi: 10.1002/bmc.5521

Keywords : nasal administration, brain targeting, drug delivery, blood–brain barrier, biomaterials, research progress

Citation: Huang Q, Chen X, Yu S, Gong G and Shu H (2024) Research progress in brain-targeted nasal drug delivery. Front. Aging Neurosci. 15:1341295. doi: 10.3389/fnagi.2023.1341295

Received: 20 November 2023; Accepted: 22 December 2023; Published: 17 January 2024.

Reviewed by:

Copyright © 2024 Huang, Chen, Yu, Gong and Shu. This is an open-access article distributed under the terms of the Creative Commons Attribution License (CC BY) . The use, distribution or reproduction in other forums is permitted, provided the original author(s) and the copyright owner(s) are credited and that the original publication in this journal is cited, in accordance with accepted academic practice. No use, distribution or reproduction is permitted which does not comply with these terms.

*Correspondence: Gu Gong, [email protected] ; Haifeng Shu, [email protected]

This article is part of the Research Topic

Progress in New Drug Development and Drug Delivery Strategies for Central Nervous System Diseases

• Search Menu
• Advance Articles
• Collections
• Author Guidelines
• Submission Site
• Open Access Options
• Self-Archiving Policy
• About Regenerative Biomaterials
• About the Chinese Society for Biomaterials
• Editorial Board
• Advertising and Corporate Services
• Journals Career Network
• Journal disclaimer
• Journals on Oxford Academic
• Books on Oxford Academic

Article Contents

Introduction, materials and methods, result and discussion, supplementary data.

• < Previous

ROS-responsive hydrogels with spatiotemporally sequential delivery of antibacterial and anti-inflammatory drugs for the repair of MRSA-infected wounds

• Article contents
• Figures & tables
• Supplementary Data

Bowen Qiao, Jiaxin Wang, Lipeng Qiao, Aziz Maleki, Yongping Liang, Baolin Guo, ROS-responsive hydrogels with spatiotemporally sequential delivery of antibacterial and anti-inflammatory drugs for the repair of MRSA-infected wounds, Regenerative Biomaterials , Volume 11, 2024, rbad110, https://doi.org/10.1093/rb/rbad110

• Permissions Icon Permissions

For the treatment of MRSA-infected wounds, the spatiotemporally sequential delivery of antibacterial and anti-inflammatory drugs is a promising strategy. In this study, ROS-responsive HA-PBA/PVA (HPA) hydrogel was prepared by phenylborate ester bond cross-linking between hyaluronic acid-grafted 3-amino phenylboronic acid (HA-PBA) and polyvinyl alcohol (PVA) to achieve spatiotemporally controlled release of two kinds of drug to treat MRSA-infected wound. The hydrophilic antibiotic moxifloxacin (M) was directly loaded in the hydrogel. And hydrophobic curcumin (Cur) with anti-inflammatory function was first mixed with Pluronic F127 (PF) to form Cur-encapsulated PF micelles (Cur-PF), and then loaded into the HPA hydrogel. Due to the different hydrophilic and hydrophobic nature of moxifloxacin and Cur and their different existing forms in the HPA hydrogel, the final HPA/M&Cur-PF hydrogel can achieve different spatiotemporally sequential delivery of the two drugs. In addition, the swelling, degradation, self-healing, antibacterial, anti-inflammatory, antioxidant property, and biocompatibility of hydrogels were tested. Finally, in the MRSA-infected mouse skin wound, the hydrogel-treated group showed faster wound closure, less inflammation and more collagen deposition. Immunofluorescence experiments further confirmed that the hydrogel promoted better repair by reducing inflammation (TNF-α) and promoting vascular (VEGF) regeneration. In conclusion, this HPA/M&Cur-PF hydrogel that can spatiotemporally sequential deliver antibacterial and anti-inflammatory drugs showed great potential for the repair of MRSA-infected skin wounds.

The skin separates the internal environment of the human body from the external environment and plays a crucial role in the human body [ 1 , 2 ]. However, skin is very vulnerable to injury, which includes external factors (surgery, pressure, burns and cuts, etc.) and pathological factors (diabetes or vascular disease, etc.) [ 3 ]. These injuries can cause gaps in the protective barrier of the skin, allowing pathogens such as bacteria to attack the human body through gaps. Research showed that methicillin-resistant Staphylococcus aureus (MRSA) can exist in 7–30% of wounds [ 4 ], and MRSA may spread into the blood, even endanger life. Because MRSA is an antibiotic-resistant pathogen that can cause multiple serious infections, research by the World Health Organization have shown that the mortality rate of MRSA-infected patients is 64% higher than that of other infected patients [ 5 ]. Therefore, effective treatment strategies for MRSA infection are particularly important for people’s lives and health.

At present, there are many strategies for the treatment of MRSA-infected wounds, such as antibiotics, bacteriophages and nanomedicine platforms [ 6 , 7 ]. Among numerous treatment methods, the use of antibiotics that can kill drug-resistant bacteria can directly and effectively treat MRSA infection, but the unreasonable use of drug doses gradually reduces the therapeutic effect of antibiotics, and even forms the antibiotic resistance [ 8 , 9 ]. Therefore, designing and developing better biocarrier to control antibiotics release is a mainstream direction for the treatment of MRSA infection [ 10–12 ]. Among the current medical wound dressings (gauze, adhesive bandage, foam and hydrogel, etc.) [ 13–15 ], hydrogel is a good platform for antibiotic controlled release [ 16 ]. At the same time, hydrogel offers the advantages that other materials do not have, such as moisturizing, self-healing, and on-demand functional designs [ 17 , 18 ]. Therefore, it is of great practical significance to develop an antibacterial hydrogel-based antibiotic delivery system for MRSA-infected wounds.

After the formation of the wound, the locally recruited inflammatory cells immediately migrate to the wound site, making the wound repair enter the inflammatory phase [ 19 , 20 ]. However, excessive oxidative stress at the wound site can produce excessive ROS [ 21 ], which in turn triggers chronic inflammation [ 22 ], creating a vicious cycle. On the other hand, the colonization of MRSA in the wound bed undoubtedly leads to severe infection [ 23 ], exacerbating the inflammatory response and causing tissue damage [ 24 ]. Infection also weakens immune cells, making it difficult to reverse pathological changes, prolonging the inflammatory phase and hindering further wound repair [ 25 , 26 ]. Therefore, in order to treat infected wounds, it is necessary to remove bacteria [ 27 , 28 ], reduce ROS and inflammation [ 29–31 ], so as to restore balance in the microenvironment [ 32 , 33 ] of the wound site and facilitate orderly subsequent repair.

There have been studies on hydrogels for the treatment of MRSA infection from the perspective of antibacterial, anti-inflammatory and antioxidant, but the load of drugs is mostly reflected in the controlled release of a single drug. For example, due to the presence of carboxyl groups in pectin and gelatin, curcumin (Cur)-loaded photocrosslinked hydrogels composed of methacrylated gelatin and methacrylated pectin can release more Cur under alkaline conditions, showing great advantages for the treatment of infected wounds [ 34 ]. Singh et al. prepared a hydrogel system by chitosan and poly(N-isopropylacrylamide-co-methacrylic acid) (PNIPAM-co-MAA) microgels. Due to the temperature-responsiveness of PNIPAM and the pH-responsiveness of the carboxylic acid groups in MAA, the release of moxifloxacin in the hydrogel can achieve dual-responsive control of temperature and pH [ 35 ]. The above studies have shown that a single controlled release of a drug can only achieve a single purpose, which cannot meet the multiple needs of infected wounds for antibacterial and anti-inflammatory etc., and also cannot meet the spatiotemporally sequential delivery of multiple needs. Therefore, design of hydrogel dressings that can deliver different drugs at different times to treat MRSA-infected wounds is a promising strategy.

Moxifloxacin is a broad-spectrum fluoroquinolone antibiotic [ 36 ], which is an appropriate choice for the treatment of skin bacterial infections. Curcumin is a kind of diketone polyphenol compound, which has many functions such as antibacterial, anti-inflammatory, and antioxidant [ 37–39 ]. The safety of Cur has been certified by the World Health Organization and U.S. Food and Drug Administration [ 40 ]. Besides, the industrial production of Cur is relatively mature, and the product is cheap and economical [ 41 ]. Based on the hydrophilicity of moxifloxacin hydrochloride and the hydrophobicity of Cur, they can be designed to exist in two different forms in hydrogels, that is, moxifloxacin hydrochloride can be loaded in the hydrogel by directly mixed with the hydrogel precursor solution. And for Cur, it can be firstly mixed with Pluronic F127 (PF) to form Cur-encapsulated PF micelles (Cur-PF), and then loaded into the hydrogel to achieve sustained release of Cur [ 42 ]. So, the release of Cur is characterized by sustained release, while the release of moxifloxacin directly loaded in the hydrogel is characterized by high efficiency and fast release. This is consistent with the treatment characteristics of MRSA-infected wounds, and has not been reported.

In this study, ROS-responsive HPA hydrogel loaded with antibiotic moxifloxacin (M) and anti-inflammatory ingredient Cur-PF was prepared by cross-linking of phenylboronic acid ester between hyaluronic acid-grafted 3-amino phenylboronic acid (HA-PBA) and polyvinyl alcohol (PVA) to treat MRSA-infected wound healing. In this hydrogel, the phenylboronic acid ester bond formed by HA-PBA and PVA has ROS responsiveness, which can realize the responsive release of drugs. Moxifloxacin and Cur exist in different forms in the hydrogel, which makes them spatially different from each other. Besides, the spatial difference between the two drugs results in a difference in their release rate, which further results in a time difference. Therefore, the hydrogel designed above can the hydrogel prepared in this study showed rapidly release of moxifloxacin to endow antibacterial property, and exhibited sustained release of Cur for anti-inflammatory under ROS-responsive conditions. Those differences meet the different needs of MRSA-infected wounds at different treatment periods based on spatiotemporally sequential delivery time-space sequential release of the two drugs. The swelling, degradation, self-healing, biocompatibility, responsive sequential release, antibacterial, anti-inflammatory and antioxidant properties of the hydrogel were tested, and their effectiveness in repairing in full-thickness skin was verified in a MRSA-infected mouse skin wound model. This is the first time to realize the spatiotemporally sequential delivery of antibacterial and anti-inflammatory drugs on hyaluronic acid (HA)-based hydrogel for repairing the MRSA-infected skin wound of mouse.

Hyaluronic acid (Mn = 800 000 Da), 3-amino phenylboronic acid (PBA), polyvinyl alcohol type 224 (PVA), 1-(3-dimethylaminopropyl)-3-ethyl carbodiimide hydrochloride (EDC), N-hydroxysuccinimide (NHS), and 2,2-diphenyl-1-(2,4,6-trinitrophenyl) hydrazide (DPPH) were purchased from Macklin; Cur, PF, and 2′,7′-dichlorofluorescein diacetate (DCFH-DA) were purchased from Sigma Aldrich. All reagents were used directly without special purification.

Preparation of hyaluronic acid grafted 3-amino phenylboronic acid (HA-PBA)

HA (1 g, 2.5 mmol) was firstly dissolved in 100 ml of deionized water, then the EDC (575 mg, 3 mmol) and NHS (345 mg, 3 mmol) were added. The pH of the solution was controlled at 5 ∼ 6. After the above mixed solution was stirred for 20 min, PBA (410.65 mg, 3 mmol) was added to the mixture, and 1 M HCl was used to control the pH of the solution at 5 ∼ 6. The reaction lasted overnight at room temperature. Then, the mixture was dialyzed (3500 KD) for 3 days, and freeze-dried.

Preparation of curcumin encapsulated Pluronic F127 micelles (Cur-PF)

According to the references [ 43 ], a one-step solid dispersion method was used to synthesize Cur-PF micelle. The feed ratio of Cur and PF polymer was 2:98.

Preparation of hydrogels

The synthesized HA-PBA polymer was dissolved in distilled water at a concentration of 3 wt%; PVA was dissolved in distilled water at a concentration of 10 wt%. Then, the HA-PBA and PVA were mixed in a certain volume ratio and vortex rapidly to obtain HPA hydrogels with final HA-PBA concentration of 1.5 wt% and the final PVA concentration of 1, 2 and 3 wt%, and these hydrogels were named as HPA1, HPA2, and HPA3, respectively.

To prepare the hydrogels loaded with Cur and moxifloxacin, the Cur-PF was dissolved in distilled water at a concentration of 30 wt%. Then, Cur-PF and HA-PBA were mixed in advance with a certain volume ratio, and the final concentration of Cur-PF was 5 wt%. The mass content of moxifloxacin and Cur in Cur-PF was the same, and they were premixed into HA-PBA solution. Then, the above mixed solution was mixed with PVA solution, and the hydrogels obtained were named as HPA1/M&Cur-PF, HPA2/M&Cur-PF and HPA3/M&Cur-PF, respectively.

The characterization of hydrogels

The tests of nuclear magnetic resonance ( 1 H-NMR) [ 44 ], Fourier transform infrared spectroscopy (FT-IR) [ 45 ], field emission scanning electron microscopy (SEM) [ 46 ], transmission electron microscope (TEM) [ 43 ], swelling [ 47 ], degradation [ 48 ], self-healing, rheological and mechanical properties [ 49 ], DPPH scavenging [ 50 ], ROS scavenging [ 51 ] and biocompatibility were all carried out according to the literature [ 52 ]. And the operational details can be found in SI. All animal experiments were conducted in accordance with the current guidelines for experimental animal care, and were approved by the Professional Committee of Xi’an Jiaotong University.

In vitro drug release assay

The drug release characteristics of HPA/M&Cur-PF hydrogels were tested in PBS or 1 mM H 2 O 2 for moxifloxacin and Cur. The drug released from the hydrogel was analyzed by UV-Vis spectrophotometer at 420 nm (Cur) and 288.57 nm (Moxifloxacin), respectively [ 53 ]. The details can be found in SI.

Antibacterial property test of the hydrogels

To test the antibacterial properties of the released drugs from the HPA/M&Cur-PF hydrogels, the samples were placed on solid medium (nutrient agar) in contact with the bacteria and the zones of inhibition around each sample were measured to record the antibacterial effect of HPA/M hydrogel loaded with moxifloxacin, and HPA/M&Cur-PF hydrogel loaded with moxifloxacin and Cur [ 44 ]. The details can be found in SI.

Anti-inflammatory experiments of HPA hydrogels

Macrophage polarization was induced by lipopolysaccharide (LPS). Hydrogel leachate was used instead of culture medium, and after incubation for 48 h, total RNA of macrophages was isolated and reverse-transcribed and amplified for further analysis of related gene expression [ 54 ].

Wound healing in an in vivo MRSA infection model

To further evaluate the promoting effect of HPA/M&Cur-PF hydrogel on wound healing, a wound healing model of MRSA-infected mouse back skin was established. The details can be found in SI.

Histological and immunohistochemical evaluation

Collect wound specimens on the Days 3, 7, and 14 after treatment. Then, hematoxylin–eosin (HE) staining was performed to evaluate the epidermal regeneration and inflammation of the wound. Masson staining was used to evaluate collagen deposition in wound beds. On the other hand, immunofluorescence staining was performed using TNF-α and VEGF antibodies, respectively.

Statistical analysis

All experimental data were statistically analyzed, and the results were expressed as mean ± SD. Statistical differences were determined by one-way ANOVA and a Student t -test. In all cases, if P  <   0.05, there is a significant difference.

Ethics approval

All protocols about animal experiments were approved by the animal research committee of Xi’an Jiaotong University (approval number: 2023-1469).

Synthesis of hydrogel

In this study, based on the dynamic phenylboronic acid ester bond between HA-PBA and PVA, and Cur-PF and antibiotic moxifloxacin (M), a series of hydrogel dressings with good antibacterial, anti-inflammatory and antioxidant effects, and stimulus-responsive drug release in different spatiotemporal sequences were prepared. Figure 1 showed the overall strategy to prepare HPA/M&Cur-PF hydrogels for MRSA-infected skin wounds healing. Firstly, PBA was grafted onto HA through amidation reaction, forming HA-PBA ( Figure 1A ). Secondly, Cur was encapsulated in PF by taking advantage of the self-assembly characteristics of PF to form Cur-PF ( Figure 1B ). Figure 1C is the structural diagram of PVA. Figure 1D showed the specific preparation procedure of HPA/M&Cur-PF hydrogel, namely Cur-PF and moxifloxacin were first mixed with HA-PBA precursor solution, and then HA-PBA in the mixed solution formed phenylboronic acid ester dynamic bond through the combination of phenylboronic acid group with the diol group structure on PVA. The hydrogel was named as HA-PBA/PVA1 (HPA1), HA-PBA/PVA2 (HPA2), and HA-PBA/PVA3 (HPA3) according to the final concentration of PVA in the hydrogel varying from 10, 20–30 mg/ml. Figure 1E showed the application of HPA/M&Cur-PF hydrogel in the MRSA-infected skin wound of mice. Based on the response of phenylboronic acid ester dynamic bond to ROS, and the different loading forms of moxifloxacin and Cur in hydrogel, two drugs in the HPA/M&Cur-PF hydrogel achieved stimulus-responsive release in different spatiotemporal sequences. When the HPA/M&Cur-PF hydrogel was applied to the MRSA-infected skin wounds of mice, it can achieve responsive anti-inflammatory and antioxidant property on the basis of rapid antibacterial action, and synergistically promote the wound repair.

Schematic diagram of preparation and application of HPA/M&Cur-PF hydrogel. ( A ) Hyaluronic acid-grafted 3-amino phenylboronic acid (HA-PBA). ( B ) polyvinyl alcohol (PVA). ( C ) preparation of curcumin-encapsulated pluronic F127 micelles (Cur-PF). ( D ) structure diagram of HPA/M&Cur-PF hydrogel. ( E ) Application of hydrogel in MRSA-infected skin wound healing of mouse.

As shown in Figure 2A , the peak of HA-PBA at 7–8 ppm in the 1 H-NMR spectrum comes from hydrogen on the benzene ring, demonstrating the successful grafting of PBA. The chemical shifts ( δ , ppm) of the peaks were assigned as below: 7.58 (m, 4H, A), 1.89 (s, 3H, B). It could be seen that phenylboronic acid was successfully grafted. Through integral calculation, the grafting ratio of PBA was 14.2%. Meanwhile, as shown in Figure 2B , the changes in the peaks at 1459 and 1517 cm −1 in the FT-IR spectrum came from benzene ring, and the peak at 1340 cm −1 is attributed to the stretching vibration of the B–O, once again proving the successful grafting of PBA. The TEM image of Cur-PF in Figure 2C confirmed the formation of Cur-PF micelles with a diameter of around 300 nm, which is consistent with the results of previous studies [ 43 ]. As shown in the Supplementary Figure S1 , the diameter of Cur-PF micelles was tested by using dynamic light scattering, and the experimental results showed that it is distributed in the range of 220–458 nm, which is consistent with the TEM results. Figure 2D showed the state of HPA hydrogel before and after gelation. Both HA-PBA and PVA are in the liquid state with fluidity. After mixing and shaking them in a certain proportion within 2 min, the gelatinized HPA hydrogel without flowing state can be observed. The test tube inversion method was used to measure the gelation time at constant temperature of 25°C. Supplementary Table S1 showed the average gelation time of the HPA2 hydrogel was 50.4 ± 2.4 s, while the gelation time of the HPA2/PF hydrogel was a little longer, about 69.3 ± 2.1 s, which may be due to the surfactant of PF [ 55 , 56 ]. Figure 2E showed the SEM images of all hydrogels. With the increase of PVA concentration in HPA1, HPA2 and HPA3 hydrogel, the crosslinking-density of hydrogels was increased, and the pore size also became smaller. When the PF micelles were added, it can be seen that the pore size of HPA2/PF hydrogel was more uniform than that of HPA2, which may be caused by the nature of non-ionic surfactant of PF. Figure 2F showed the statistics of the pore diameter of all hydrogels, which more intuitively showed that with increasing of PVA concentration, the pore size of the hydrogel gradually decreased from 191.8 ± 49.3 µm of HPA1 hydrogel to 92.6 ± 30.7, 68.3 ± 24.4 and 66.2 ± 10.3 µm for HPA2, HPA3 and HPA2/PF hydrogels, respectively. And the pore size of HPA2/PF hydrogel was more uniform compared to HPA2 hydrogel.

( A ) 1 H-NMR spectrum of HA-PBA, A represents the hydrogen on the benzene ring and B represents the hydrogen on the methyl group. ( B ) FT-IR spectra of HA, PBA, and HA-PBA. ( C ) The TEM image of Cur-PF micelle. ( D ) Gelation display of HPA hydrogel. ( E ) HPA1, HPA2, HPA3 and of HPA2/PF hydrogels’ SEM image (scale bar: 200 μm) and ( F ) pore size diameter statistics.

Mechanical properties, swelling, degradation and self-healing of hydrogels

With the introduction of the wet healing theory [ 57 ], maintaining a certain level of humidity at the wound site is beneficial for the repair of skin wounds [ 58 ]. Figure 3A showed the equilibrium swelling ratio of HPA hydrogels, all hydrogels can absorb more than 60 times water of their own mass. The swelling ratio of HPA hydrogels decreased with increasing of PVA concentration, and their swelling ratios were 8878.8 ± 470.1%, 7360.0 ± 372.5% and 5971.1 ± 322.1% for HPA1, HPA2 and HPA3, respectively. This should be due to the increase of PVA content, which provides more crosslinkable sites and increases the crosslinking density within a certain range. Besides, the swelling ratio of HPA2/PF was 6560.5 ± 107.3%, slightly lower than that of HPA2 hydrogel, because of hydrogen bonding between PF micelles and HPA hydrogel network. The good swelling properties of the HPA hydrogels make it natural and great advantages in the management of wound exudate, and it can absorb the exudate well while maintaining the wound moist.

( A ) Swelling behavior of HPA hydrogels. ( B ) Degradation behavior of HPA hydrogels. ( C ) Rheological behavior of HPA hydrogels. ( D ) The strain–stress curves of HPA hydrogels during the compression test. ( E ) Self-healing display of HPA hydrogels. ( F ) The rheological properties of HPA2/PF hydrogel when alternate step strain switched from 1% to 1500%. ( G ) ROS-responsive properties of hydrogels. The release of ( H ) moxifloxacin and ( I ) curcumin from HPA2/M&Cur-PF hydrogels in PBS or H 2 O 2 .

Biodegradability is a crucial criterion for measuring medical materials [ 59 ]. Therefore, the degradation performance of HPA hydrogels was further evaluated. Figure 3B showed that all hydrogels have good degradability. With increasing of PVA concentration, the crosslinked network of hydrogels was closer, so the degradation rate of HPA3 hydrogels was the slowest. Specifically, after 3 days of testing, the remaining weight of HPA1, HPA2 and HPA3 hydrogels groups remained 54.6 ± 1.8%, 72.0 ± 4.0%, and 83.0 ± 1.7%, respectively. Since the addition of PF micelles maked the crosslinked network of hydrogel more closely, HPA2/PF degraded more slowly than HPA2, showed 88.0 ± 1.7% remaining weight after 3 days of testing. Importantly, the remaining weight of all hydrogels was less than 30% after the 12 days test. Experiments verified that all hydrogel dressings prepared in this study had reasonable degradation properties.

Appropriate modulus is very important for hydrogel dressings. It can provide a good mechanical matching between hydrogel dressings and skin tissues, ensure a comfortable sense of wear during use, and reduce physical damage to damaged tissues. Therefore, rheological tests were used to evaluate the properties of HPA hydrogels and Figure 3C showed that with increasing of PVA concentration, the modulus of the hydrogel showed a trend of gradual increase. The storage modulus of HPA1 was 40.6 Pa, that of HPA2 was 139.1 Pa, and that of HPA3 was 424.7 Pa. Due to the addition of PF micelles, the modulus of HPA2/PF hydrogel was also increased to 227.4 Pa compared with HPA2 hydrogel. The thermal stability of the hydrogels in the range of 25–40°C was tested. As shown in the Supplementary Figure S2 , the modulus of HPA2 and HPA2/PF hydrogels did not change much. In addition, we can clearly observe that the HPA2/PF hydrogel can still maintain the gel-forming state under the high temperature condition of 40°C. The hydrogel has good thermal stability between 25°C and 40°C, so when applied to the skin surface, the hydrogel can still maintain its own stability even under body temperature conditions.

Skin’s inevitable stretching during human activities is very easy to damage the hydrogel dressings. Therefore, higher requirements are put forward to the mechanical properties of hydrogels. As shown in Figure 3D , it can be seen from the stress–strain test that all hydrogels showed good compressibility. When the strain was 80%, with the increase of PVA concentration, the stress of the hydrogels was 2161.7, 2445.9 and 3547.0 Pa, respectively, and all hydrogels were unbroken. After the addition of PF micelles, the stress of HPA2/PF hydrogel increased to 4998.7 Pa at the strain of 80%. The above experimental results showed that the mechanical properties of the hydrogel gradually increase with increasing of PVA concentration, and it can be effectively improved by adding PF micelles. The results of the above tests showed that the properties of HPA hydrogels, including pore size, swelling, degradation, modulus, and mechanical properties, can be easily regulated by adjusting the ratio between components. This tunable property provides a broader selectivity for the hydrogel to adapt to different wound states and repair stages.

The damage of hydrogel dressing will lose the protective effect to wound, which puts forward requirements for the self-healing performance of skin wound dressing. As shown in Figure 3E , after the hydrogel was cut into two halves, it can quickly self-heal and merge within 5 min. Figure 3F showed that when the strain was higher than 1500%, G″ of hydrogel was higher than G′, which means that the hydrogel network was damaged. Therefore, by constantly switching low strain (γ = 1%) and high strain (γ = 1500%), the quantitative self-healing function was tested. At the beginning of the test, at the first low strain (γ = 1%), G′ and G″ was 241.3 and 138.4 Pa, respectively, and G′ > G″. After switching to the high strain (γ = 1500%), G′ changed to 27.6 Pa and G″ changed to 44.7 Pa, G″ > G′, which means that the hydrogel network crashed. In the following tests, when γ was cyclically transformed, G′ and G″ can recover to the initial value, with no significant difference in modulus compared to the first test. In conclusion, due to the existence of phenylboronic acid ester dynamic bond, the HPA hydrogel prepared in this study has good self-healing performance, which provides the quickly restore of structural integrity when it is ruptured under external force, not only ensures the physical barrier effect, but also avoids the possible bacterial invasion after rupture. As shown in the Supplementary Figure S3 , shear rate scanning measurements indicate that all hydrogels in this experiment exhibit shear-thinning behavior, where the viscosity of the material depends on the shear force, and thus the hydrogels in this work are injectable.

ROS-responsive drug release of hydrogels

The biggest problem of using antibiotics is bacterial resistance [ 60 ]. Loading drugs in hydrogels can greatly avoid the repeated use of drugs and reduce the generation of bacterial resistance. In particular, the ROS response caused by the presence of phenylboronic acid dynamic ester bond in HPA hydrogel can realize intelligent on-demand drug delivery in the hydrogel system [ 61 ]. As shown in Figure 3G , when 200 μl of 10 mM H 2 O 2 was added to the prepared 500 μl HPA2/M&Cur-PF hydrogel, a certain fluidity of the hydrogel can be observed just within 20 min, and the hydrogel network collapsed completely at 4 h. This is mainly due to the destruction of the hydrogel network structure caused by the reaction of the phenylboronic acid ester structure with H 2 O 2 . As shown in Supplementary Figure S4 , the modulus of HPA2 hydrogels and HPA2/PF hydrogels were tested under 1 mM H 2 O 2 or PBS for the same time, respectively. At 10 min, the modulus of the hydrogel in PBS did not change much compared to the initial value, but the modulus of the hydrogel in H 2 O 2 decreased dramatically by about 200 Pa. At 20 min, the modulus of the hydrogel in PBS decreased due to the swelling and absorbing of water, and at this time, the hydrogel in H 2 O 2 was close to the state of non-gelation. In this study, two drugs were loaded into the hydrogel dressings. One was moxifloxacin with antibacterial effect directly mixed in the hydrogel, and the other was Cur with anti-inflammatory and antioxidant functions encapsulated in PF micelles. As shown in Figure 3H , the release rate of moxifloxacin in PBS is relatively slow, and the release time can be more than 250 h. While in 1 mM H 2 O 2 release solution, the release amount of moxifloxacin in 36 h can reach more than 80%. The above results showed that hydrogel can deliver antibiotics quickly and effectively after being applied to wounds due to the ROS responsiveness of phenylboronic acid ester bonds. These results mean that when bacterial infection is serious, leading to excessive inflammation and the production of a large amount of ROS, the ROS response can be quickly activated, so that the antibiotic can be released rapidly to prevent the continuation of severe infection in a short time. The release amount of Cur in PBS and H 2 0 2 was 14% and 38% at 24 h. In summary, moxifloxacin can be rapidly released more than 80% within 36 h, and Cur can be continuously released within 288 h under the action of 1 mM H 2 O 2 . In a word, the spatiotemporally sequential release of these two drugs allows for a rapid antimicrobial treatment and then sustained anti-inflammatory and antioxidant effect at the site of the infected wound.

Antibacterial properties of hydrogels

The antibacterial effect of the antibiotics released from the hydrogel was verified by the inhibition zone test. The control group is just the holes punched on the agar plate by the punch, so the diameter size and shape of the control group do not change with time and environmental factors, and at the same time to ensure that the initial diameter is the same between every group, and to confirm whether the experimental group formed a ring of inhibition around the holes. As shown in Figure 4A , at the 12 h, the diameter of the inhibition zone for HPA2/M hydrogel and HPA2/M&Cur-PF hydrogel against E. coli were 2.20 ± 0.02 and 2.26 ± 0.02 cm, respectively, and the diameter of the inhibition zone against MRSA were 2.08 ± 0.01 and 2.05 ± 0.01 cm, respectively. This suggested that the antibacterial effect of HPA2/M&Cur-PF hydrogels is better than that of HPA2/M hydrogels. Until the 60 h, there was still obvious inhibition zone. At this time, the diameter of HPA2/M hydrogel and HPA2/M&Cur-PF hydrogel against E. coli was 1.30 ± 0.09 and 1.50 ± 0.02 cm, respectively, and the diameter of the inhibition zone against MRSA was 1.24 ± 0.02 and 1.33 ± 0.01 cm, respectively. The appearance of inhibition zone indicates that moxifloxacin and Cur-PF loaded in the hydrogel diffuse to the surrounding environment, thus killing the bacteria in a certain area. In addition, the inhibition zone diameter of HPA2/M&Cur-PF hydrogel for MRSA was statistically larger than HPA2/M hydrogel, and showed significant different ( P  < 0.05). This is because Cur also has antibacterial effect [ 62 ], which is synergistic with moxifloxacin. More obviously, at the 84 h, the inhibition zone of HPA2/M hydrogel had disappeared, while the inhibition zone of HPA2/M&Cur-PF hydrogel was still slightly larger than 7 mm, suggesting that HPA2/M&Cur-PF hydrogels had better antibacterial effect than HPA2/M hydrogels. In conclusion, the hydrogel dressings prepared in this study had good sustained antibacterial effect.

( A ) The inhibition zone of hydrogel for E. coli and MRSA within 96 h. In the figure, 1, 2 and 3 represent the control group, the HPA2/M group and the HPA2/M&Cur-PF group, respectively. Statistics of inhibition zone of hydrogel for ( B ) E. coli and ( C ) MRSA within 96 h (* P  < 0.05).

Biocompatibility of hydrogel

Good biocompatibility is an essential prerequisite for biomedical materials [ 63 ]. In this study, the prepared hydrogels were tested from the aspects of blood compatibility and cell compatibility. As shown in Figure 5A , in the blood compatibility test, the materials in experimental group showed comparable to or even lower hemolysis ratio than that of the PBS group, with hemolysis ratio of below 5%, which was considered to be a good range of blood compatibility. The test results indicated that they will not cause significant hemolysis when applied to biological tissues. Figure 5B showed the cell compatibility of the hydrogels. Because of the good adhesion and proliferation effect of HA on cells [ 64 ], the cell viability of HPA2 hydrogel group was higher than that of the control group after co-culture. The cell viability of the HPA2/M&Cur-PF group compared to the control group was 80.3%, 82.1%, and 94.4% in the first 3 days, respectively. Although the cell viability of the experimental group was slightly decreased in the first 2 days of testing compared to the control group, the significant differences were all P  < 0.05, indicating that the material was not toxic. Figure 5C exhibited the Live/Dead staining images of L929 cells after co-cultured with the hydrogel leachate for one day, which is consistent with the quantitative statistical results of cell viability. In general, the hydrogel dressings prepared in this study had good biocompatibility.

( A ) Hemolysis ratio of hydrogel. ( B ) Cell compatibility of hydrogel on L929 cells within 3 days test. ( C ) Live/dead straining of L929 cells on the first day. ( D ) DPPH scavenging statistics of hydrogel. ( E ) Representative images of ROS scavenging experiments and ( F ) statistical results for fluorescent areas. Statistic of expression results of ( G ) TNF-α and ( H ) IL-1β after application of hydrogels on macrophages (* P  < 0.05, ** P  < 0.01, *** P  < 0.001).

Antioxidation and anti-inflammation of hydrogel

If MRSA-infected wounds are not treated in a timely and appropriate manner, it can easily change to severe infection, causing severe inflammation at the wound site, and producing a large amount of ROS through oxidative stress [ 65 , 66 ]. Therefore, we first used the scavenging experiment of stable free radical DPPH• to evaluate the antioxidant properties of the hydrogels. Figure 5D showed that the presence of PBA group brings little DPPH• scavenging capacity to the hydrogels. Furthermore, due to the good antioxidant capacity of Cur, 500 μl of HPA2/Cur-PF hydrogel can scavenge more than 90% of DPPH. In addition, Figure 5E further showed the ROS scavenging ability of the hydrogels by DCFH-DA staining. LPS-induced macrophage (RAW 264.7) produced a large amount of ROS, while normal macrophage (RAW 264.7) cells in the control group only produced a small amount of ROS and the HPA2/Cur-PF hydrogel group showed similar or even lower ROS intensity than the control group. The statistical results in Figure 5F showed that the ROS production in the experimental group was not significantly different from that in the control group. As is well known, high concentrations of ROS can cause DNA damage and cell death, while low concentrations of ROS play a crucial role in signal transduction and wound repair [ 67 ]. Therefore, the experimental group didn’t show a significant different amount of ROS compared to the control group, which is very beneficial for wound repair. As shown in Figure 5G and H , the significantly increased expression levels of TNF-α and IL-1β in the LPS group indicate that macrophage (RAW 264.7) cells were successfully induced into an inflammatory state, while the expression levels of both inflammatory factors were significantly reduced under the action of the hydrogel group, indicating that the hydrogel has significant anti-inflammatory effects. In summary, the HPA2/Cur-PF hydrogels have suitable antioxidant and anti-inflammatory effects, which lays a strong foundation for its use as a dressing for wound with MRSA infection.

Promoting effect of HPA/M&Cur-PF hydrogel on healing of MRSA-infected skin wounds

A series of in vitro experiments have verified the good antibacterial, anti-inflammatory and antioxidant effects of HPA/M&Cur-PF hydrogel. We further established a MRSA-infected mouse back skin wound model, and evaluated the repair-promoting effect of the hydrogel prepared in this study. Figure 6A showed the entire time course from modeling to wound repair. First, a circular wound with a diameter of 8 mm was created on the back skin of mouse, and 10 μl of 10 8 CFU/ml MRSA were injected. On the 3rd, 7th, and 14th day, the skin at the wound site was observed and sectioned for study. The commercially available Tegaderm™ dressing was selected as the control group [ 68 ]. Based on the previous characterization tests, HPA2 hydrogel was selected as the representative and applied in the subsequent animal experiments. The experimental groups were HPA2 hydrogel, HPA2/M hydrogel loaded with moxifloxacin (the concentration of moxifloxacin was 2 mg/ml), HPA2/Cur-PF hydrogel loaded with Cur-encapsulated PF (the concentration of Cur was 2 mg/ml), and HPA2/M&Cur-PF hydrogel loaded with moxifloxacin and Cur-encapsulated PF (the concentration of moxifloxacin and Cur were both 2 mg/ml). In Figure 6B , the wounds in each group showed different phenomena over time after surgery. Control group showed significant suppuration on the third day after treatment, which even lasted until the seventh day. This is consistent with the clinical phenomenon that MRSA-infected wounds have severe bacterial infection with a lot of inflammation and even difficulty in healing. The HPA2 hydrogel group did not have a good anti-infection effect compared to the other three experimental groups, and no significant sepsis was found. In Figure 6C , the wound areas of each group were plotted over time. In Figure 6D , the wound closure ratio of each group was analyzed. Compared with the control group, after 7 days of treatment, there was a significant difference ( P  < 0.01) in the wound closure of HPA2, HPA2/M, HPA2/Cur-PF, and HPA2/M&Cur-PF group, respectively. Besides, there was a significant difference in the wound closure of HPA2/M&Cur-PF and HPA2 ( P  < 0.05), which indicated that the hydrogel loaded with the two drugs promoted better wound repair. On the 14th day, there still was a significant difference ( P  < 0.01) between HPA2/M&Cur-PF and HPA2/M and HPA2/Cur-PF, which proved that moxifloxacin and Cur promoted MRSA-infected wound healing, and the synergistic effect of the two drugs was more favorable for wound repair. In conclusion, on the 14th day, all hydrogel groups showed good healing effect. The HPA2/M&Cur-PF hydrogel group showed the best healing effect, with almost complete wound closure. But the control group still had an obvious wound, with 34.3% of the remaining wound area, which was significantly different from the other hydrogel groups ( P  < 0.001).

( A ) Schematic diagram of the in vivo wound healing experimental program in an infected full-thickness skin defect model. ( B ) The pictures of wounds on the 3rd, 7th, and 14th day were divided into five groups: Tegaderm™ film dressing (control), HPA2 hydrogel, HPA2/M hydrogel, HPA2/Cur-PF hydrogel and HPA2/M&Cur-PF hydrogel, and ( C ) plotting of wound area over time. ( D ) Statistics on changes in wound closure ratio. ( E ) HE staining of the wound site on the 3rd, 7th, and 14th day. ( F ) Masson staining of the wound site on the seventh day. ( G ) Statistics of relative inflammation at the wound site on the third day. ( H ) Thickness of regenerated epidermis at the wound site on the seventh day. ( I ) Statistics of relative collagen deposition at the wound site on the seventh day (* P  < 0.05, ** P  < 0.01, *** P  < 0.001).

The effect of each group on wound repair was further evaluated by HE staining. In Figure 6E , on the third day, a large amount of inflammation can be seen in the control group, while the inflammation in other hydrogel groups is relatively low. Figure 6G showed the statistics of inflammation. Due to the effects of moxifloxacin and Cur, the inflammation of HPA2/M, HPA2/Cur-PF and HPA2/M&Cur-PF hydrogel groups is significantly lower than that of the control group and HPA2 hydrogel group ( P  < 0.01). The dual effect of antibacterial and anti-inflammatory effects makes the relative amount of inflammation between HPA2/M&Cur-PF hydrogel group and HPA2/M and HPA2/Cur-PF have a significant difference ( P  < 0.01). In Figure 6E , on the seventh day, the epidermal regeneration in the control group was discontinuous, accompanied by blood scabs, which was mainly because the wound site was infected by MRSA and the bacteria were not removed in time, and the infection also made it difficult for wound healing to continue the next step. However, obvious continuous epidermal regeneration was seen in HPA2/M, HPA2/Cur-PF and HPA2/M&Cur-PF hydrogel groups. In Figure 6H , statistics showed that the thickness of regenerated epidermis in HPA2/M&Cur-PF hydrogel group was the thickest, which was significantly different from the other four groups ( P  < 0.01). The metabolism of collagen participates in the whole process of wound repair and the regenerated collagen constitutes an important part of the repaired wound, so its importance is obvious. In Figure 6F , masson staining showed the deposition of collagen in each group on the seventh day. And in Figure 6I , a statistical analysis of the masson staining was performed. The relative collagen deposition of the four test groups was significantly higher than that of the control group, and the difference was significant compared with the control group ( P  < 0.05). Among the experimental groups, the HPA2/M&Cur-PF group, which was more effective for wound repair, had 3.4 times more collagen deposition than the control group. In summary, HPA2/M&Cur-PF hydrogel dressings can promote wound closure, reduce inflammation and promote collagen deposition in MRSA-infected mouse skin wound healing.

Immunofluorescence staining analysis

Wound healing is a complex process, and the expression of relevant cytokines can reflect some specific situations during the wound healing process [ 69 , 70 ]. The third day after wound formation is considered to be the inflammatory period. At this time, a proinflammatory cytokine, tumor necrosis factor (TNF-α) [ 71 ], was selected to evaluate the effect of hydrogel on inflammation control. From the immunofluorescence staining analysis of TNF-α in Figure 7A and the statistics in Figure 7B , it can be seen that the colonization of bacteria at the wound leads to the aggravation of inflammation, and there were significant differences between control and HPA2/M, HPA2/Cur-PF and HPA2/M&Cur-PF hydrogel groups, respectively. Specifically, inflammation in HPA2/M and HPA2/Cur-PF hydrogel groups was significantly reduced. The reduction in inflammation in the HPA2/M hydrogel was due to the antibacterial effect of moxifloxacin, while the reduction in inflammation in HPA2/Cur-PF hydrogel was due to the antibacterial and anti-inflammatory effects of Cur. Most importantly, HPA2/M&Cur-PF hydrogel group has the synergistic effect of two drugs, so it showed the lowest inflammation. The generation of blood vessels is an essential physiological process in wound healing, which can rebuild normal blood flow for wound tissue, provide nutrition and oxygen for the tissue, and accelerate the process of wound repair [ 72 , 73 ]. On the seventh day, vascular endothelial growth factor (VEGF) [ 74 ] was used to evaluate the neovascularization at the wound. Figure 7A showed the immunofluorescence staining analysis of VEGF and Figure 7C showed the statistical results. Significant differences ( P  < 0.01) were found between the control group and the HPA2/M hydrogel group, the HPA2/Cur-PF hydrogel group and the HPA2/M&Cur-PF hydrogel group, respectively. In HPA2/M and HPA2/Cur-PF hydrogel groups, due to the effect of two drugs to avoid bacterial infection and inflammation in the early stage, the hydrogel promoted the formation of new blood vessels. In conclusion, based on the expression of TNF-α and VEGF, HPA2/M&Cur-PF hydrogel showed better therapeutic effect in MRSA-infected skin wounds.

( A ) Immunofluorescence staining images of TNF-α on the third day and VEGF on the seventh day, with green indicating TNF-α expression and red indicating VEGF. ( B ) Quantitative analysis of the relative area percentage ( n  = 3) of TNF-α and ( C ) VEGF. For all quantitative analyses, the commercial film group data for TNF-α on the third day and VEGF on the seventh day were set as 100% (* P  < 0.05, ** P  < 0.01).

In this study, HA-based HPA/M&Cur-PF hydrogel dressing with spatiotemporally sequential delivery of antibacterial and anti-inflammatory drugs was constructed based on the dynamic bond of phenylboronic acid ester for the first time, and was used to repair MRSA-infected skin wounds. The rheological properties, self-healing, biocompatibility, responsiveness, spatiotemporally sequential drug delivery, antibacterial, antioxidant and anti-inflammatory properties of hydrogels were verified. Under ROS conditions, hydrogels can release more than 80% of moxifloxacin within 36 h to perform quickly anti-infection effect, and release Cur up to 288 h to perform a sustained anti-inflammation effect. Finally, in the MRSA-infected mouse skin wound healing, the hydrogel-treated group exhibited faster wound closure, reduced inflammation, and promoted epidermal growth and collagen deposition. Immunofluorescence staining results also demonstrated the hydrogels’ ability to reduce inflammation while also promoting angiogenesis. In conclusion, HPA/M&Cur-PF hydrogel dressing has a significant effect on the repair of skin wounds infected by MRSA, providing ideas for responsive spatiotemporally sequential drug delivery strategies.

Supplementary data are available at Regenerative Biomaterials online.

This work was jointly supported by the National Natural Science Foundation of China (grant numbers 51973172, 52273149), supported by 111 Project 2.0 (grant number BPO618008), the Natural Science Foundation of Shaanxi Province (grant number 2020JC-03), and China Postdoctoral Science Foundation (grant number 2022M712498), State Key Laboratory for Mechanical Behavior of Materials, and the World-Class Universities (Disciplines) and the Characteristic Development Guidance Funds for the Central Universities.

Conflicts of interest statement . None declared.

Byrd AL , Belkaid Y , Segre JA. The human skin microbiome . Nat Rev Microbiol 2018 ; 16 : 143 – 55 .

Google Scholar

Murphree RW. Impairments in skin integrity . Nurs Clin North Am 2017 ; 52 : 405 – 17 .

Whitney JD. Overview: acute and chronic wounds . Nurs Clin North Am 2005 ; 40 : 191 – 205 .

Gurusamy KS , Koti R , Toon CD , Wilson P , Davidson BR. Antibiotic therapy for the treatment of methicillin‐resistant Staphylococcus aureus (MRSA) in non surgical wounds . Cochrane Database Syst Rev 2013 ; 11 : CD010427 .

Nandhini P , Kumar P , Mickymaray S , Alothaim AS , Somasundaram J , Rajan M. Recent developments in methicillin-resistant Staphylococcus aureus (MRSA) treatment: a review . Antibiotics 2022 ; 11 : 606 .

Duan YY , Xu PY , Ge PY , Chen LF , Chen Y , Kankala RK , Wang SB , Chen AZ. NIR-responsive carrier-free nanoparticles based on berberine hydrochloride and indocyanine green for synergistic antibacterial therapy and promoting infected wound healing . Regen Biomater 2023 ; 10 : rbad076 .

Le T , Tran DL , Hoang Thi TT , Lee Y , Park KD. Injectable reactive oxygen and nitrogen species-controlling hydrogels for tissue regeneration: current status and future perspectives . Regen Biomater 2022 ; 9 : rbac069 .

Livermore DM. Antibiotic resistance in staphylococci . Int J Antimicrob. Agents 2000 ; 16 : 3 – 10 .

Frieri M , Kumar K , Boutin A. Antibiotic resistance . J Infect Public Health 2017 ; 10 : 369 – 78 .

Sharifi S , Hajipour MJ , Gould L , Mahmoudi M. Nanomedicine in healing chronic wounds: opportunities and challenges . Mol Pharm 2020 ; 18 : 550 – 75 .

Zheng Z , Yang X , Fang M , Tian J , Zhang S , Lu L , Zhou C , Xu C , Qi Y , Li L. Photothermal effective CeO2NPs combined in thermosensitive hydrogels with enhanced antibacterial, antioxidant, and vascularization performance to accelerate infected diabetic wound healing . Regen Biomater 2023 ; 10 : rbad072 .

Ren X , Hu Y , Chang L , Xu S , Mei X , Chen Z. Electrospinning of antibacterial and anti-inflammatory Ag@ hesperidin core-shell nanoparticles into nanofibers used for promoting infected wound healing . Regen Biomater 2022 ; 9 : rbac012 .

Dhivya S , Padma VV , Santhini E. Wound dressings–a review . Biomedicine (Taipei) 2015 ; 5 : 22 – 8 .

Rezvani Ghomi E , Khalili S , Nouri Khorasani S , Esmaeely Neisiany R , Ramakrishna S. Wound dressings: current advances and future directions . J Appl Polym Sci 2019 ; 136 : 47738 .

Cao D , Ding J. Recent advances in regenerative biomaterials . Regen Biomater 2022 ; 9 : rbac098 .

Op’t Veld RC , Walboomers XF , Jansen JA , Wagener FADTG. Design considerations for hydrogel wound dressings: strategic and molecular advances . Tissue Eng Part B Rev 2020 ; 26 : 230 – 48 .

Brumberg V , Astrelina T , Malivanova T , Samoilov A. Modern wound dressings: hydrogel dressings . Biomedicines 2021 ; 9 : 1235 .

Zeng Z , Zhu M , Chen L , Zhang Y , Lu T , Deng Y , Ma W , Xu J , Huang C , Xiong R. Design the molecule structures to achieve functional advantages of hydrogel wound dressings: advances and strategies . Compos B Eng 2022 ; 247 : 110313 .

Tottoli EM , Dorati R , Genta I , Chiesa E , Pisani S , Conti B. Skin wound healing process and new emerging technologies for skin wound care and regeneration . Pharmaceutics 2020 ; 12 : 735 .

Kolaczkowska E , Kubes P. Neutrophil recruitment and function in health and inflammation . Nat Rev Immunol 2013 ; 13 : 159 – 75 .

Schäfer M , Werner S. Oxidative stress in normal and impaired wound repair . Pharmacol Res 2008 ; 58 : 165 – 71 .

Bryan N , Ahswin H , Smart N , Bayon Y , Wohlert S , Hunt JA. Reactive oxygen species (ROS)–a family of fate deciding molecules pivotal in constructive inflammation and wound healing . Eur Cell Mater 2012 ; 24 : 249 – 65 .

Andersson H , Lindholm C , Fossum B. MRSA–global threat and personal disaster: patients’ experiences . Int Nurs Rev 2011 ; 58 : 47 – 53 .

Scalise A , Bianchi A , Tartaglione C , Bolletta E , Pierangeli M , Torresetti M , Marazzi M , Di Benedetto G. Microenvironment and microbiology of skin wounds: the role of bacterial biofilms and related factors . Semin Vasc Surg 2015 ; 28 : 151 – 9 .

Stadelmann WK , Digenis AG , Tobin GR. Impediments to wound healing . Am J Surg 1998 ; 176 : 39S – 47S .

Negut I , Grumezescu V , Grumezescu AM. Treatment strategies for infected wounds . Molecules 2018 ; 23 : 2392 .

Gurusamy KS , Koti R , Toon CD , Wilson P , Davidson BR. Antibiotic therapy for the treatment of methicillin‐resistant Staphylococcus aureus (MRSA) infections in surgical wounds . Cochrane Database of Syst Rev 2013 ; 8 : CD009726 .

Selders GS , Fetz AE , Radic MZ , Bowlin GL. An overview of the role of neutrophils in innate immunity, inflammation and host-biomaterial integration . Regen Biomater 2017 ; 4 : 55 – 68 .

Landén NX , Li D , Ståhle M. Transition from inflammation to proliferation: a critical step during wound healing . Cell Mol Life Sci 2016 ; 73 : 3861 – 85 .

Gong X , Luo M , Wang M , Niu W , Wang Y , Lei B. Injectable self-healing ceria-based nanocomposite hydrogel with ROS-scavenging activity for skin wound repair . Regen Biomater 2022 ; 9 : rbab074 .

Chen Y , Cai J , Liu D , Liu S , Lei D , Zheng L , Wei Q , Gao M. Zinc-based metal organic framework with antibacterial and anti-inflammatory properties for promoting wound healing . Regen Biomater 2022 ; 9 : rbac019 .

Yang Y , Wang L , Zhou Y , He Y , Lin S , Zeng Y , Zhou Y , Li W , He Z , Zhao Q. Antioxidant-enriched autologous biogel promoted diabetic wound healing by remodeling inherent posttraumatic inflammatory patterning and restoring compromised microenvironment homeostasis . Regen Biomater 2022 ; 9 : rbac023 .

Kirchner S , Lei V , MacLeod AS. The cutaneous wound innate immunological microenvironment . Int J Mol Sci 2020 ; 21 : 8748 .

Bostancı NS , Büyüksungur S , Hasirci N , Tezcaner A. pH responsive release of curcumin from photocrosslinked pectin/gelatin hydrogel wound dressings . Biomater Adv 2022 ; 134 : 112717 .

Singh N , Aery S , Juneja S , Kumari L , Lone MS , Dar AA , Pawar SV , Mehta SK , Dan A. Chitosan hydrogels with embedded thermo-and pH-responsive microgels as a potential carrier for controlled release of drugs . ACS Appl Bio Mater 2022 ; 5 : 3487 – 99 .

Balfour JAB , Wiseman LR. Moxifloxacin . Drugs 1999 ; 57 : 363 – 73; discussion 374 .

Stanić Z. Curcumin, a compound from natural sources, a true scientific challenge–a review . Plant Foods Hum Nutr 2017 ; 72 : 1 – 12 .

Barzegar A , Moosavi-Movahedi AA. Intracellular ROS protection efficiency and free radical-scavenging activity of curcumin . PLoS One 2011 ; 6 : e26012 .

Zheng D , Huang C , Huang H , Zhao Y , Khan MRU , Zhao H , Huang L. Antibacterial mechanism of curcumin: a review . Chem Biodivers 2020 ; 17 : e2000171 .

Ali A , Jaber N. Pharmacological aspects of curcumin: review article . Int J Pharmacogn 2018 ; 5 : 313 – 26 .

Yi XL , Qaria MA , Sivasamy S , Jian ZS , Dao CZ. Curcumin production and bioavailability: a comprehensive review of curcumin extraction, synthesis, biotransformation and delivery systems . Ind Crops Prod 2021 ; 172 : 114050 .

Kaur J , Singla P , Kaur I. Labrasol mediated enhanced solubilization of natural hydrophobic drugs in pluronic micelles: physicochemical and in vitro release studies . J Mol Liq 2022 ; 361 : 119596 .

Qu J , Zhao X , Liang Y , Zhang T , Ma PX , Guo B. Antibacterial adhesive injectable hydrogels with rapid self-healing, extensibility and compressibility as wound dressing for joints skin wound healing . Biomaterials 2018 ; 183 : 185 – 99 .

Liang Y , Zhao X , Hu T , Han Y , Guo B. Mussel-inspired, antibacterial, conductive, antioxidant, injectable composite hydrogel wound dressing to promote the regeneration of infected skin . J Colloid Interface Sci 2019 ; 556 : 514 – 28 .

He J , Shi M , Liang Y , Guo B. Conductive adhesive self-healing nanocomposite hydrogel wound dressing for photothermal therapy of infected full-thickness skin wounds . Chem Eng J 2020 ; 394 : 124888 .

Yang Y , Li M , Pan G , Chen J , Guo B. Multiple stimuli‐responsive nanozyme‐based cryogels with controlled NO release as self‐adaptive wound dressing for infected wound healing . Adv Funct Mater 2023 ; 33 : 2214089 .

Zhao X , Dong R , Guo B , Ma PX. Dopamine-incorporated dual bioactive electroactive shape memory polyurethane elastomers with physiological shape recovery temperature, high stretchability, and enhanced C2C12 myogenic differentiation . ACS Appl Mater Interfaces 2017 ; 9 : 29595 – 611 .

Qiao L , Liang Y , Chen J , Huang Y , Alsareii SA , Alamri AM , Harraz FA , Guo B. Antibacterial conductive self-healing hydrogel wound dressing with dual dynamic bonds promotes infected wound healing . Bioact Mater 2023 ; 30 : 129 – 41 .

Wang C , Liang Y , Huang Y , Li M , Guo B. Porous photothermal antibacterial antioxidant dual–crosslinked cryogel based on hyaluronic acid/polydopamine for non-compressible hemostasis and infectious wound repair . J Mater Sci Technol 2022 ; 121 : 207 – 19 .

Li M , Pan G , Yang Y , Guo B. Smart aligned multi-layered conductive cryogels with hemostasis and breathability for coagulopathy epistaxis, nasal mucosal repair and bleeding monitoring . Nano Today 2023 ; 48 : 101720 .

Wang X , Wu B , Zhang Y , Dou X , Zhao C , Feng C. Polydopamine-doped supramolecular chiral hydrogels for postoperative tumor recurrence inhibition and simultaneously enhanced wound repair . Acta Biomater 2022 ; 153 : 204 – 15 .

He J , Liang Y , Shi M , Guo B. Anti-oxidant electroactive and antibacterial nanofibrous wound dressings based on poly(ε-caprolactone)/quaternized chitosan-graft-polyaniline for full-thickness skin wound healing . Chem Eng J 2019 ; 385 : 123464 .

Liang Y , Li M , Yang Y , Qiao L , Xu H , Guo B. pH/glucose dual responsive metformin release hydrogel dressings with adhesion and self-healing via dual-dynamic bonding for athletic diabetic foot wound healing . ACS Nano 2022 ; 16 : 3194 – 207 .

Sunnapu O , Khader R , Dhanka M , Kumar Vemula P , Karuppannan S. Enzyme‐responsive hydrogel for delivery of the anti‐inflammatory agent zingerone . ChemNanoMat 2022 ; 8 : e202200334 .

Cortés H , Hernández-Parra H , Bernal-Chávez SA , Prado-Audelo MLD , Caballero-Florán IH , Borbolla-Jiménez FV , González-Torres M , Magaña JJ , Leyva-Gómez G. Non-ionic surfactants for stabilization of polymeric nanoparticles for biomedical uses . Materials (Basel) 2021 ; 14 : 3197 .

Pluronic F. Applications of thermo-reversible pluronic F-127 gels in pharmaceutical formulations . J Pharm Pharm Sci 2006 ; 9 : 339 – 58 .

Xian C , Qing YW , Yan PL , Xiao FW , Feixiang C , Xiang YX , Ying ML. Clinical research on chronic refractory wounds treated by the wound ostomy specialist group using the theory of wet healing . Int Med Health Guid News 2022 ; 28 : 1997 .

Harries RL , Bosanquet DC , Harding KG. Wound bed preparation: TIME for an update . Int Wound J 2016 ; 13 Suppl 3 : 8 – 14 .

Daniel S. Biodegradable polymeric materials for medicinal applications . In: Thomas S, Balakrishnan P, editors. Green composites. Singapore: Springer Singapore, 2021 ; 351 – 72 .

Urban-Chmiel R , Marek A , Stępień-Pyśniak D , Wieczorek K , Dec M , Nowaczek A , Osek J. Antibiotic resistance in bacteria—a review . Antibiotics 2022 ; 11 : 1079 .

Chen W , Zhen X , Wu W , Jiang X. Responsive boron biomaterials and their biomedical applications . Sci China Chem 2020 ; 63 : 648 – 64 .

Peng Y , Ao M , Dong B , Jiang Y , Yu L , Chen Z , Hu C , Xu R. Anti-inflammatory effects of curcumin in the inflammatory diseases: status, limitations and countermeasures . Drug Des Devel Ther 2021 ; 15 : 4503 – 25 .

Marin E , Boschetto F , Pezzotti G. Biomaterials and biocompatibility: an historical overview . J Biomed Mater Res A 2020 ; 108 : 1617 – 33 .

Cai S , Wu C , Yang W , Liang W , Yu H , Liu L. Recent advance in surface modification for regulating cell adhesion and behaviors . Nanotechnol Rev 2020 ; 9 : 971 – 89 .

Wang G , Yang F , Zhou W , Xiao N , Luo M , Tang Z. The initiation of oxidative stress and therapeutic strategies in wound healing . Biomed Pharmacother 2023 ; 157 : 114004 .

Medzhitov R. Origin and physiological roles of inflammation . Nature 2008 ; 454 : 428 – 35 .

Deng Z , Shi F , Zhou Z , Sun F , Sun M-H , Sun Q , Chen L , Li D , Jiang C-Y , Zhao R-Z , Cui D , Wang X-J , Jing Y-F , Xia S-J , Han B-M. M1 macrophage mediated increased reactive oxygen species (ROS) influence wound healing via the MAPK signaling in vitro and in vivo . Toxicol Appl Pharmacol 2019 ; 366 : 83 – 95 .

Guo B , Liang Y , Dong R. Physical dynamic double-network hydrogels as dressings to facilitate tissue repair . Nat Protoc 2023 ; 18 : 3322 – 54 .

Singh S , Young A , McNaught CE. The physiology of wound healing . Surgery (Oxford) 2017 ; 35 : 473 – 7 .

Tsirogianni AK , Moutsopoulos NM , Moutsopoulos HM. Wound healing: immunological aspects . Injury 2006 ; 37 Suppl 1 : S5 – 12 .

Zelová H , Hošek J. TNF-α signalling and inflammation: interactions between old acquaintances . Inflamm Res 2013 ; 62 : 641 – 51 .

Guerra A , Belinha J , Jorge RN. Modelling skin wound healing angiogenesis: a review . J Theor Biol 2018 ; 459 : 1 – 17 .

Johnson KE , Wilgus TA. Vascular endothelial growth factor and angiogenesis in the regulation of cutaneous wound repair . Adv Wound Care (New Rochelle) 2014 ; 3 : 647 – 61 .

Dvorak HF. VPF/VEGF and the angiogenic response . Semin Perinatol 2000 ; 24 : 75 – 8 .

• anti-inflammatory agents
• inflammation
• moxifloxacin
• anti-bacterial agents
• methicillin-resistant staphylococcus aureus

Email alerts

Citing articles via, affiliations.

• Online ISSN 2056-3426
• Print ISSN 2056-3418
• Copyright © 2024 Oxford University Press and Chinese Society for Biomaterials
• About Oxford Academic
• Publish journals with us
• University press partners
• What we publish
• New features  
• Open access
• Institutional account management
• Rights and permissions
• Get help with access
• Accessibility
• Advertising
• Media enquiries
• Oxford University Press
• Oxford Languages
• University of Oxford

Oxford University Press is a department of the University of Oxford. It furthers the University's objective of excellence in research, scholarship, and education by publishing worldwide

• Copyright © 2024 Oxford University Press
• Cookie settings
• Cookie policy
• Privacy policy
• Legal notice

This Feature Is Available To Subscribers Only

Sign In or Create an Account

This PDF is available to Subscribers Only

For full access to this pdf, sign in to an existing account, or purchase an annual subscription.