research paper on targeted drug delivery

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• Cancers (Basel)

Targeted Delivery Methods for Anticancer Drugs

Valery v. veselov.

1 Center of Bioanalytical Investigation and Molecular Design, Sechenov First Moscow State Medical University, 8 Trubetskaya ul, 119991 Moscow, Russia; ur.xednay@gnulebindercas (V.V.V.); ur.liam@0502kirer (A.E.N.)

Alexander E. Nosyrev

László jicsinszky.

2 Department of Drug Science and Technology, University of Turin, Via P. Giuria 9, 10125 Turin, Italy; moc.liamg@ykzsniscijl

Renad N. Alyautdin

3 Department of Pharmacology, Sechenov First Moscow State Medical University, 119991 Moscow, Russia; ur.dempxe@nidtuayla

Giancarlo Cravotto

4 World-Class Research Center “Digital Biodesign and Personalized Healthcare”, Sechenov First Moscow State Medical University, 8 Trubetskaya ul, 119991 Moscow, Russia

Simple Summary

The current main technological strategies for the delivery of anticancer drugs are discussed herein. This comprehensive review may help researchers design suitable delivery systems.

Several drug-delivery systems have been reported on and often successfully applied in cancer therapy. Cell-targeted delivery can reduce the overall toxicity of cytotoxic drugs and increase their effectiveness and selectivity. Besides traditional liposomal and micellar formulations, various nanocarrier systems have recently become the focus of developmental interest. This review discusses the preparation and targeting techniques as well as the properties of several liposome-, micelle-, solid-lipid nanoparticle-, dendrimer-, gold-, and magnetic-nanoparticle-based delivery systems. Approaches for targeted drug delivery and systems for drug release under a range of stimuli are also discussed.

1. Introduction

Over the past 30 years, the number of successful cancer treatments has significantly increased, predominantly driven by our improved understanding of carcinogenesis processes, cell biology, and the tumor microenvironment [ 1 , 2 ]. However, many cancers are still fatal despite the sustained effort being invested in preclinical and clinical research. One of the ways to improve the survival rate of cancer patients is the targeted delivery of anticancer drugs. Advances in biomedical science and biotechnology have led to the discovery and development of effective drug carriers such as liposomes, dendrimers, and gold and magnetic nanoparticles [ 3 , 4 , 5 , 6 ]. The principal difference between these new types of formulation and classical ones is their suitability for the potential development of technologies for targeted drug delivery to specific tissues, cells, and even intracellular organelles. The essence of targeted delivery lies in the surface of a drug container (carrier) bearing a modified drug or molecule with a functional group that can be recognized by the target cell receptors. Folic acid modification is a classic example as it is actively taken up by tumor cells [ 7 , 8 , 9 ]. Antibodies and aptamers are universal molecules that recognize the surface of a target cell [ 10 , 11 , 12 ]. Thanks to advances in basic biomedical research, the antigenic portraits of cells are becoming more and more detailed, allowing us to distinguish one cell from another based on their surface characteristics. Orally or parenterally administered medicines are distributed throughout the body, with only a small portion reaching the target area. Targeted delivery methods, therefore, make it possible to reduce the dosage of an administered drug and minimize its effect on other cells, which is very important in chemotherapy as drugs are highly toxic. The presence of recognizing or recognizable molecules on the surface of a delivery system allows it to concentrate on the desired area. It is also vital that the delivery system penetrates the cell and that the drug is then delivered to the nucleus, mitochondria, endoplasmic reticulum, and other organelles. In fact, the concept of intracellular drug delivery is under active development. Knowledge of the signaling pathways involving proteins that lead to different cellular structures is essential to achieve efficient intracellular transport [ 13 ]. Equally important is the need for more knowledge of the motor proteins of cells, which directionally move loads over long distances inside cells. It is also necessary to understand the mechanisms by which drugs are released from delivery systems, including diffusion, degradation, swelling, and other processes that can control the release of drugs [ 14 , 15 , 16 ].

2. Types of Containers and Carriers

A drug-delivery system may be considered suitable for clinical practice if it is non-toxic, biocompatible, stable in blood, non-immunogenic, non-thrombogenic, and biodegradable [ 17 ].

The enhanced permeability and retention (EPR) effect is the principle behind the passive targeting used in all containers and carriers. This principle and term were proposed in 1986 [ 18 ]. Rapid tumor growth is accompanied by neovascularization with wide fenestrations and the suppression of lymphatic drainage [ 19 ]. Traditionally, the EPR concept presumes that small molecules enter via diffusion and leave the interstitial space of the tumor, whereas macromolecules (containers, carriers) are no longer able to do so after extravasation [ 20 ]. In addition to the traditional explanation, other theories concerning how the pathophysiological characteristics of tumor growth shape the EPR effect have been put forward [ 21 , 22 ]. Thus, the EPR effect has been accepted as a universal principle incorporated into the design of anticancer drug-delivery systems [ 23 ]. However, there are currently serious disputes about the effectiveness of the EPR effect when using nanoparticles [ 24 , 25 ]. At the same time, it is important to note that the current understanding of the EPR effect is based on results obtained in animal models, meaning that the results of EPR-effect studies in patients must be collated if delivery systems that fully exploit the EPR effect are to be successfully designed [ 26 ].

2.1. Liposomes

Liposomes are spherical vesicles consisting of one or more lipid bilayers. A liposome has a hollow structure that is usually filled with a solvent and can deliver a variety of substances. Its hydrophobic membrane allows it to merge with cell membranes and transport its contents inside cells. Liposomes are most often composed of phospholipids and cholesterol, but may also include other lipids to improve endocytosis and tissue compatibility. Many methods have been developed to produce a range of liposomal compositions [ 27 ], and all described liposome fabrication methods combine lipids with the aqueous phase in some way [ 28 , 29 ]. The thin-layer hydration method, also known as the Bangham method, is one of the first and still most commonly used methods for the preparation of liposomes [ 30 ]. This method involves lipids being dissolved in the organic phase and removing the organic solvent, usually by evaporation, to form a lipid film. The lipid film is then dispersed in an aqueous medium that contains the drug under vigorous stirring to form the sealed spherical structures; liposomes. The short elimination half-life of liposomes, caused by their opsonization principally in the liver and spleen, is a crucial weak point in their use [ 31 ]. The modification of liposome surfaces with various functional ligands, such as polyethylene glycol (PEG) coating, reduces the interaction between the surface and blood components, thus ensuring that the liposomes have a longer residence time in the bloodstream [ 32 ]. PEG can be attached to liposome surfaces in a variety of ways:

• Physical adsorption onto the surface of liposomes.
• Covalent attachment using reactive groups on the surface of preformed liposomes.
• Inclusion of a PEG-lipid conjugate in liposome preparations.

The most common method anchors the polymer in the membrane using a cross-linked lipid (e.g., PEG-distearoylphosphatidylethanolamine) [ 33 ]. The presence of PEG on liposome surfaces reduces their aggregation [ 34 ]. To ensure targeted delivery, PEG is also often covalently bound to proteins (transport, signaling) so that, while the mechanism of action of the proteins does not change, there is a change in protein pharmacokinetics; PEG-asparaginase (used in the treatment of leukemia), PEG-aldesleukin (an antineoplastic agent), PEG-filgrastim (for the treatment of chemotherapy-induced febrile neutropenia), and PEG-epoetin-β (for the treatment of anemia) are commonly used in the treatment of cancer [ 35 , 36 ]. The liposomal delivery of anticancer drugs has been successfully used in cancer therapy for several decades [ 37 ].

2.2. Micelles

Micelles are particles of tens of nanometers in size with a hydrophobic core and a hydrophilic surface and are commonly used as carriers of hydrophobic drugs. Like liposomes, they can be delivered directly into the bloodstream through the respiratory tract or skin. In recent years, amphiphilic block copolymers, which spontaneously form micellar structures, have attracted much attention because of their use in the delivery of cytostatic drugs [ 38 , 39 ]. Amphiphilic block copolymers are usually assembled from two or three blocks, with PEG being the most common hydrophilic block in the copolymer structure. Other hydrophilic block-forming polymers include chitosan, polyvinylpyrrolidone, and poly(N-isopropyl acrylamide) [ 40 ]. Polymers of various compound classes are used as hydrophobic polymer blocks for micellar core creation: polyethers (poly(propylene oxide)) polyesters (polylactide), polycarboxylic acids (poly(aspartic acid)) and lipids (distearoylphosphatidyl ethanolamine) [ 40 ]. Micelles that contain functional groups (-NH2, -COOH) in their core can transfer drugs by chemical modification and not just by physical encapsulation [ 41 ], and various cytostatic drug micelles (doxorubicin, paclitaxel) have shown significant results in several in-vitro and in-vivo studies [ 42 ]. Paclitaxel encapsulated in micelles has been tested in clinical trials in patients with malignant tumors with a resulting reduction in toxicity and no change in the antitumor activity compared to free paclitaxel [ 43 ].

2.3. Solid-Lipid Nanoparticles

Solid-lipid nanoparticles are colloidal nanoparticles stabilized by surfactants and composed of mono-, di- and triglycerides, solid fats, and waxes. They have been developed as an alternative to liposome technologies to increase stability, modulate the release of encapsulated drugs, reduce costs, and simplify manufacturing [ 44 ]. Unlike liposomes, which are usually injected into the body intravenously, intraperitoneally, subcutaneously, and orally, solid-lipid nanoparticles can be administered via different routes, via inhalation, intranasally, and intravesically [ 45 ], thus ensuring the local targeting of the drug. Recent in-vitro and in-vivo experiments have shown that solid-lipid nanoparticles that contain cytostatic drugs appear to be superior to conventional drug solutions and are comparable to other encapsulated systems in many aspects, such as efficacy, pharmacokinetics, and bioavailability [ 46 ]. However, clinical studies have not yet been conducted in this area.

2.4. Gold Nanoparticles

Gold nanoparticles (AuNP) can boast a combination of unique physical and chemical properties relative to other biomedical nanotechnologies and can selectively deliver cytostatic drugs [ 47 , 48 ]. AuNPs offer significant potential for new approaches to cancer treatment as they are easy to produce, have low toxicity, and display antiangiogenic properties [ 49 ]. AuNPs are up to 100 nm in size, have a pronounced EPR-effect, and, as a result, preferentially accumulate in tumors.

AuNP-based supports are most often synthesized using colloidal methods; gold salts (e.g., hydrogen tetrachloroaurate (III)) are reduced in the presence of surface stabilizers that prevent the aggregation of the resulting solution [ 50 ]. Spherical AuNPs are principally used to create delivery systems because they can be synthesized on a large scale with high monodispersity. The other forms of AuNP include nanorods, nanoshells, and nano cells [ 51 ]. AuNPs can undergo surface modification thanks to their covalent and non-covalent bond-forming properties [ 51 ]. A stabilizing agent (e.g., citric acid) is responsible for the overall charge of the AuNP surface. The correct choice of a stabilizing agent allows various biomolecules (DNA, antibodies, polypeptides) to be conjugated to the AuNP surface via electrostatic interactions, whereas covalent attachment to AuNPs is usually achieved via the interaction between gold and thiol, amine, and carboxylate functional groups [ 52 ]. Unlike liposomes and micelles, the drug is conjugated directly to the AuNP surface using various linkers [ 52 , 53 , 54 ]. It is worth noting that the overwhelming majority of studies on AuNP-based directional transport are based on spherical AuNPs, and this is, at least in part, because they undergo surface modification and penetrate cells more easily than more complex AuNPs. A drug conjugated to AuNPs has shown increased antitumor potential compared to the free drug in in-vitro and in-vivo studies [ 47 , 55 , 56 , 57 , 58 ].

2.5. Magnetic Nanoparticles

Magnetic targeting is of great interest in the treatment of malign tumors as the technique not only provides targeted drug delivery but also makes it possible to monitor the accumulation of magnetic nanoparticles (MNP) in tumors using magnetic resonance imaging (MRI) [ 59 , 60 ]. MNPs that carry a drug are first accumulated in the target tissue using an external magnetic field, and the drug is then released from the MNPs in a controlled manner [ 61 ].

MNPs are magnetic materials with small particle sizes (from 10 to 100 nm), a large specific surface area, magnetic response, and superparamagnetism [ 62 ]. This superparamagnetism means that MNPs are in a single-domain state, as they are uniformly magnetized throughout the entire volume [ 63 ], and that the orientation of their magnetic moment changes with temperature [ 63 ]. Iron oxides, for example magnetite (Fe 3 O 4 or FeO.Fe 2 O 3 ) and maghemite (γ-Fe 2 O 3 ), are usually used for MNP production [ 60 , 64 ]. The MNP core, which consists of magnetite, maghemite, or a mixture of the two, is usually obtained via the precipitation of Fe 2+ and Fe 3+ iron salts from an aqueous solution [ 65 , 66 ]. Moreover, it is possible to regulate the size of the resulting nanoparticles by adding various iron salts (chloride, sulfate, nitrate, etc.) and by changing the ratio of Fe 2+ and Fe 3+ , the pH, and the ionic strength in the solution [ 62 , 67 ]. Reactions are carried out in an inert atmosphere to prevent the oxidation of the formed nanoparticles [ 68 ]. The formed MNPs have a hydrophobic surface and are coated with synthetic and natural polymers to reduce nanoparticle agglomeration [ 60 ] and further modify the surface to conjugate drugs and biomolecules [ 69 ]. The most commonly used polymers are PEG, dextran, polyvinylpyrrolidone, polyaniline, alginate various fatty acids, and chitosan [ 70 , 71 ]. In general, the conjugates of MNPs with various cytostatics show decreased overall toxicity, and the concentration of cytostatic agents is required to achieve a therapeutic effect [ 72 , 73 , 74 , 75 , 76 , 77 ]. The ability of MNPs to accumulate in tumors has also been confirmed by MRI [ 59 , 74 , 78 , 79 , 80 ].

2.6. Dendrimers

Dendrimers are three-dimensional, monomolecular, highly branched monodisperse macromolecules [ 81 ] that usually have rotational symmetry and often take on a spherical shape. In general, dendrimers have a hydrophobic core from which they branch, ending in terminal functional groups responsible for their solubility in water [ 82 ]. These dendrimers can retain hydrophobic drugs and increase their concentration in water. Biocompatibility, easy excretion from the body, and a significantly improved EPR effect are the most remarkable advantages of dendrimers. However, dendrimers have one significant drawback; they are cytotoxic for normal cells due to the physiological stability of cationic groups on their surfaces [ 83 ]. The problem of dendrimer cytotoxicity is usually solved by modifying their surface using biocompatible polymers, for example, PEG. The PEG-modified dendrimer surface provides the necessary screening of the cationic surface charge, which leads to a biologically safe carrier [ 84 ].

Dendrimer synthesis is a rather laborious process. There are two principal approaches to the synthesis of dendrimers; divergent and convergent methods [ 85 ]. In the divergent version, a base reagent (a molecule that is protected at its end groups, if necessary) is attached to the original branching center (which has several end groups). The protecting groups are removed, and a 1st generation dendrimer is formed. Subsequently, dendrimers of higher generations are obtained by attaching either the original branching center or the base reagent, followed by deprotection [ 86 ]. In the convergent method, the arms of the dendrimer are synthesized first and then connected [ 86 ], and this method produces more monodisperse dendrimers than the divergent version. However, the size of dendrimers obtained using the convergent method is limited due to steric hindrance, whereas dendrimers of a wider variety of sizes can be obtained using the divergent method [ 85 , 86 ]. The most widely used dendrimers are currently the commercially available poly(amidoamine) (PAMAM) dendrimers [ 87 , 88 , 89 ]. Delivery systems based on poly(propylene imine) [ 90 ], polylysine [ 91 ], carbosilane [ 92 ], and phosphorus dendrimers [ 93 ] have also been developed. Numerous studies have shown the effectiveness of using different dendrimers for targeted transport in cancer therapy [ 94 , 95 , 96 , 97 , 98 , 99 ], and several clinical trials using various dendrimers as targeted delivery systems are underway [ 100 ].

2.7. Albumin-Based Nanoparticles

Albumin is the most abundant plasma protein in human blood, with a molecular weight of about 67 kDa. Due to its endogenous origin, it is non-toxic, non-immunogenic, biocompatible, and biodegradable [ 101 ]. Human serum albumin (HSA) and the cheaper bovine serum albumin (BSA) and ovalbumin (OVA) have been used to create delivery systems [ 102 ]. HSA has several ligand binding sites that can be used for transfer via both hydrophobic and electrostatic interactions [ 103 , 104 ], and the presence of a free cysteine residue on its surface means that albumin easily conjugates with a variety of ligands [ 105 , 106 ]. Receptors, such as albondin (Gp60) and secreted protein acidic and rich in cysteine (SPARC), have been shown to overexpress in some cancers [ 107 ] and can mediate albumin transcytosis [ 108 ], while the Gp30 and Gp18 receptors, the megalin/cubilin complex and the neonatal Fc receptor (FcRn) are also involved in albumin transport [ 106 ]. Albumin-based delivery systems can therefore accumulate in tumors via mechanisms beyond the EPR effect. Albumin-based nanoparticles are obtained by various methods, including emulsification, self-assembly of thermal gelation, desolvation, and nanospray drying [ 109 , 110 ]. The patented nanoparticle albumin-bound (NAB) technology, which consists of the evaporation of an emulsion with the creation of cross-links between albumin units, is the best-known preparation method for albumin-based nanoparticles [ 105 , 106 ]. In addition to Abraxane ® , which is created with the help of NAB technology and has been successfully used in clinical practice [ 111 , 112 , 113 ], work is also underway to create a range of albumin-based nanoparticles [ 104 , 105 , 106 , 107 , 108 , 109 , 110 , 111 , 112 , 113 , 114 , 115 , 116 , 117 , 118 ].

2.8. Porous Materials

Zeolites are hydrated crystalline aluminosilicates consisting of tetrahedral groups, [SiO 4 ] 4− and [AlO 4 ] 5− , united by common vertices into a three-dimensional framework. The open frame-cavity structure of zeolites has a negative charge, which is compensated for by counterions [ 119 ]. Zeolites have a porous structure that can absorb various substances, making zeolites an ideal material for drug-delivery systems [ 120 ]. To prevent the untimely release of a drug, either a zeolite with an optimal pore size is selected [ 121 ], or its surface is modified with various ligands [ 122 , 123 ]. In general, zeolites are promising carriers for creating systems for the delivery of cytotoxic substances [ 124 , 125 , 126 , 127 ].

Mesoporous silica particles (MSP) are another porous material used for drug delivery [ 128 ]. Their pore size can be adjusted from 2 to 50 nm, as in the case of zeolites, to tune them for a specific drug [ 129 , 130 , 131 ]. The surfaces of MSPs are rich in reactive silanol groups, which can be used for conjugation with various substances [ 132 ], and MSPs have been developed with several structures. The morphology and size of both the particles themselves and their pores can be controlled via the choice of a synthetic method [ 133 , 134 ]. MSP-based delivery systems have shown high drug-loading capacity, successfully controlled release, and increased antitumor activity [ 79 , 135 , 136 , 137 ].

2.9. Carbon Nanoparticles

Carbon has many allotropic modifications, including carbon nanotubes, fullerenes, and nanodiamonds, which have found applications as carriers for drug delivery [ 138 ]. Carbon-based quantum dots are also used (see the Section 2.10 ). Carbon nanoparticles have a high specific surface area and hydrophobicity. Carbon nanotubes and fullerenes have cavities in their structure and can encapsulate active substances [ 138 , 139 ]. However, unlike fullerenes and carbon nanotubes, the surface of nanodiamonds is rougher, which increases adhesion with drugs [ 140 ]. Under the action of acidic oxidation, carboxyl groups are formed on the surface of carbon nanoparticles and are used for surface modification, as well as for the covalent attachment of anticancer drugs [ 141 , 142 , 143 ]. Carbon nanoparticles with the desired properties can be obtained by correctly choosing and adapting the synthesis method [ 144 , 145 , 146 ].

While carbon nanoparticles are currently widely used for drug delivery, their toxic properties are concentration-dependent [ 147 , 148 , 149 ]. Attention should therefore be paid to the delivery method when developing carbon-nanoparticle supports. For example, it has been shown that the absorption of fullerene by the respiratory and digestive tracts is low [ 150 ]. In addition, as has been demonstrated, inhaled carbon nanotubes can act on the body similarly to asbestos [ 151 ].

2.10. Quantum Dots

Quantum dots are inorganic semiconductor nanocrystals and are typically up to 10 nm in size. Quantum dots have fluorescent, optical and electronic properties [ 152 ], with cadmium-compound-based and carbon quantum dots being the most widespread [ 153 , 154 ]. In addition to drug delivery, quantum dots can visualize cancer cells due to their unique optical properties, which derive from quantum and other effects [ 155 ]. The quantum dots used in biomedicine typically consist of a core and a coating with the core imparting optical properties to the system and the coating performing a protective function, which enables the surface to be functionalized with various ligands and is responsible for water solubility [ 156 ]. The quantum-dot core may be composed of cadmium compounds, such as cadmium selenide (CdSe), cadmium sulfide (CdS), and cadmium telluride (CdTe), and these quantum dots have shown notable results as drug carriers [ 157 , 158 , 159 ]. It is important to mention that these quantum dots are not biodegradable and are not cell and environmentally friendly due to the toxicity of cadmium compounds [ 160 ].

Carbon-based quantum dots, which can be classified as either carbon quantum dots or graphene quantum dots, are widely used in various fields of biomedicine [ 161 ]. They possess low toxicity, high specific surface area, high photostability and are easily modified [ 162 ]. Carbon-based quantum dots are excellent carriers for anticancer drugs due to their biocompatibility, ease of manufacture, and lower environmental impact [ 163 , 164 , 165 , 166 ].

2.11. Calcium Phosphate

Calcium phosphate (CaP)-based nanoparticles are crystalline formations of predominantly carbonate apatite capable of transporting a drug both on their surface and within their structure [ 167 , 168 ]. Minerals based on CaPs are the main inorganic components of the bones and teeth of vertebrates and humans [ 169 ]. CaP-based nanoparticles have several peculiar properties that make them attractive for delivering anticancer drugs. CaPs are fully biodegradable, release non-toxic calcium and phosphate ions upon degradation, and decompose faster than other inorganic nanoparticles (zeolites, mesoporous silica particles, carbon nanoparticles, and quantum dots) [ 168 , 170 ]. Moreover, CaP-based nanoparticles have pH-sensitive solubility; they are insoluble at the physiological pH of blood plasma (7.4) but quickly dissolve in acidic biological media (pH < 5), for example, in endosomes and lysosomes, where they rapidly release encapsulated substances [ 170 , 171 , 172 ]. There are currently many approaches for synthesizing CaP-based nanoparticles, and the careful selection of synthesis conditions makes it possible to control the size and morphology of the resulting particles [ 173 , 174 , 175 , 176 ]. Although nanoparticles with different morphologies are used for delivery in cancer therapy, including rod shapes [ 169 , 177 ], porous structures [ 178 , 179 ], and core-shell shapes [ 180 , 181 ], spherical nanoparticles are the most commonly used since they are more thermodynamically stable [ 182 , 183 ].

2.12. Oligo- and Polysaccharide-Based Drug-Delivery Systems

2.12.1. chitosan.

Chitosan is a type of amino polysaccharide polymer (see Figure 1 ) produced via the deacetylation of chitin, and is the second most common biopolymer in nature after cellulose. Chitin is the main component of the exoskeleton of arthropods and many other invertebrates and is also part of the cell walls of fungi [ 184 ]. Chitosan has amino functionalities that are useful for biopolymer modification [ 185 ]. Its biodegradability, biocompatibility, low immunogenicity, and non-toxicity mean that chitosan is used in delivery systems for various chemotherapeutic drugs [ 186 ]. Chitosan and its derivatives, such as carboxymethyl chitosan, sulfated chitosan, sulfated benzaldehyde chitosan, and polypyrrole-chitosan, have been shown to have anticancer activity in and of themselves [ 187 , 188 , 189 , 190 ]. This property is assumed to be related to the antioxidant properties of chitosan and its derivatives, which are capable of trapping cancer-causing free radicals [ 191 ].

Structural formulas of chitosan and chitin.

As mentioned earlier, chitosan can be a hydrophilic moiety in an amphiphilic block copolymer [ 192 ]. The presence of free amino groups in the chitosan backbone grants it a unique polycationic character that ensures that negatively charged drugs, such as doxorubicin, are properly encapsulated [ 193 ]. Some chitosan-based hydrogels that contain a significant amount of water and retain a self-organized three-dimensional structure have been developed and can be used for the encapsulation and delivery of anticancer drugs [ 194 , 195 ]. Various forms of delivery systems, such as microspheres, film capsules, etc., have been obtained using water-insoluble species of chitosan [ 196 ], meaning that the properties of chitosan-based delivery systems are easy to modulate. Depending on the preparation method selected, it is possible to regulate the particle size, toxicity, thermal and chemical stability, and release kinetics [ 197 ].

2.12.2. Cyclodextrins

Cyclodextrins (CDs) are a family of cyclic oligosaccharides that consist of glucose subunits obtained by enzymatic means from starch [ 198 ]. The most commonly used CD types are α-, β- and γ-CDs ( Figure 2 a), named according to the number of glucose residues they possess [ 199 , 200 ].

( a ) Structural formula of cyclodextrins; ( b ) 3D structure of cyclodextrins.

CDs have a hydrophilic outer surface and a significantly less hydrophilic cavity. They take the form of a truncated cone with a cylindrical-torus cavity inside ( Figure 2 b) [ 201 ]. CDs can form various complexes with hydrophilic, lipophilic, and amphiphilic substances. CDs, therefore, can often increase the solubility and bioavailability of many anticancer drugs [ 202 ], and various cyclodextrin derivatives are widely used to create drug-delivery systems [ 203 ].

CDs are often combined with other nanoparticles to create delivery systems [ 201 ]. It has been shown that the loading of liposomes with anticancer drugs in combination with cyclodextrin increases their half-life, reduces toxicity, and increases liposome loading [ 204 , 205 , 206 , 207 ]. Several CD-based polymers have also been developed and used successfully in drug transport [ 208 ].

The cross-linking of CDs results in unique particles, namely CD polymers and nanosponges. Nanosponges are a type of nanoparticle that has a porous structure with a pore size of several nanometers. Due to their unique structure, CD polymers and nanosponges can encapsulate various substances in their pores and act as drug transporters [ 209 ]. One of the advantages of natural CD-based delivery systems is the creation of effective oral, mucosal and transdermal drug formulations [ 210 ]. Cyclodextrin-based macromolecules can transfer oligonucleotides, siRNAs, or their fragments into the cells. Though the promising results reported by these associations, experimental trials are still in progress [ 211 , 212 , 213 , 214 ].

2.12.3. Pectins

Pectins are polysaccharides that are mainly formed from residues of galacturonic acid. Pectins are extracted in different ways from higher plants, mainly from their fruits. Consequently, the structures of pectins can be very diverse, although they can be classified into three types based on their general characteristics: homogalacturonan, rhamnogalacturonan-I, and substituted galacturonans [ 215 ]. Pectin and its various modifications have anticancer activities [ 216 , 217 , 218 ], and the majority of studies on natural and modified pectins, and their delivery systems, have mainly focused on colon cancer [ 219 , 220 , 221 ]. This is primarily because pectin is not digested in the gastrointestinal tract until it reaches the colon, where it is fermented and breaks down to release encapsulated active ingredients [ 222 , 223 ]. Pectin-based microgranules and microspheres have been developed to encapsulate anticancer drugs and release them directly into the colon [ 224 , 225 ]. Another use of pectins is as a drug carrier in the preparation of various hydrogels [ 226 , 227 ]. As pectins contain carboxyl groups, it is possible to use them to create negatively charged particles that retain drugs thanks to electrostatic interactions [ 228 ]. Moreover, pectin has been used to create self-organizing polymer nanoparticles to deliver ursolic acid [ 229 ].

3. The Targeting Methods of Delivery Systems

Active targeting is used to increase the concentration of cytostatics in the desired organ or tissues to achieve higher and more selective therapeutic activity. The surface of the container or carrier is modified with various recognizable or recognition molecules, such as monoclonal antibodies or their fragments, aptamers, proteins, peptides, and low molecular weight compounds, to grant active-targeting properties [ 230 ].

3.1. Antibodies and Aptamers

Numerous monoclonal antibodies (mAb) have recently been developed against various epitopes of cancer cells and are used as therapeutic agents in and of themselves [ 231 ]. mAb-conjugated containers or carriers specifically bind to a target cell (receptor, protein, etc.) in the desired areas and then release the encapsulated drug. The surface modification of the carrier systems with mAbs can either be achieved via non-covalent physical interactions or the formation of covalent bonds [ 232 ], with non-covalent bonding being faster than covalent. However, the antigen-binding domains of mAbs are arranged chaotically in non-covalent conjugation, which can lead to disruption in mAb functionality [ 233 ]. The most commonly used method for non-covalent conjugation is the streptavidin-biotin method [ 234 ]. It consists of the preliminary non-covalent conjugation of the surface of the delivery systems with streptavidin, which has a high affinity for biotin, which, in turn, is covalently linked to the mAbs.

The most commonly used cancer targets for mAbs are:

• Epidermal growth factor receptor (EGFR) [ 235 , 236 , 237 ].
• Human epidermal growth factor receptor 2 (HEP2) [ 238 , 239 , 240 ].
• B-lymphocyte antigen CD19 [ 241 , 242 , 243 ].
• Guanine deaminase (GAH) [ 244 , 245 ].
• Receptor cluster of differentiation 47 (CD47) [ 246 , 247 ].

Conjugation with aptamers is a newer approach to targeted delivery. Aptamers are oligonucleotides (DNA, RNA aptamers) or peptide molecules that specifically bind to specific target molecules and can be considered analogs of monoclonal antibodies. However, they have many advantages over antibodies. Their production is much easier, cheaper, and faster than monoclonal antibodies. They have a much smaller size and, therefore, more easily penetrate tissues and cells, as well as having higher affinity and specificity [ 248 ]. The potential for the in-vivo targeting of RNA aptamers in cancer therapy was demonstrated for the first time in 2006 [ 249 ]. More than 20 different systems for targeted transport are currently being developed using oligonucleotide aptamers [ 250 ].

Peptide aptamers consist of a short (10–20 amino acid), conformationally limited peptide sequence that is inserted into a scaffold protein (most often the bacterial protein thioredoxin A) [ 251 , 252 ]. A unique feature of peptide aptamers is that their variable region has a double limitation as both ends are connected to a framework (protein), unlike oligonucleotide aptamers and antibodies [ 252 ]. For this reason, peptide aptamers have limited conformations and require less energy to bind to the target, which, in turn, increases their affinity.

3.2. Proteins and Peptides

The most commonly used protein for targeting delivery systems is transferrin, a serum glycoprotein that transports iron into cells by binding to transferrin receptors on the cell surface [ 253 ]. The transferrin receptor is present in malignant tumors at levels that are hundreds of times higher than in normal cells ( Figure 3 ) [ 254 ]. Containers and carriers, their surface modified by transferrin molecules, can therefore penetrate cancer cells and accumulate in them [ 230 ]. The following proteins and peptides are used to target other receptors that are overexpressed in cancer cells:

• A designed ankyrin repeat protein (DARPin) can target the Epithelial cell adhesion molecule (EpCAM) [ 255 , 256 ].
• The peptide K237 targets the kinase insert domain receptor (KDR) [ 257 , 258 ].
• The peptide bombesin targets the gastrin-releasing peptide receptor [ 259 , 260 ].
• The peptide octreotide targets the somatostatin receptor type 2 [ 261 , 262 ].

Schematic representation of the penetration of transferrin into a cell via receptor-mediated endocytosis. Original diagram inspired by [ 263 ].

3.3. Low Molecular Weight Compounds

Modifying the surfaces of containers and carriers with folic acid is currently a commonly used methodology to ensure the delivery of an encapsulated drug into cancer cells [ 264 ]. Folic acid is a low molecular weight compound, a vitamin required by eukaryotic cells for the biosynthesis of purines and pyrimidines [ 265 ]. Folate uptake by cells occurs via two mechanisms: via the low-affinity-reduced folate carrier (RFC), found in almost all cells, and via the high-affinity glycosylphosphatidylinositol-linked folate receptor (FOLR), which has limited distribution [ 266 ]. FOLR can transport conjugated folate into cells, unlike RFC [ 266 , 267 ], and FOLR is also significantly expressed in different types of human tumors but is minimally expressed in most normal tissues [ 268 ]. Consequently, FOLR is a target for the selective delivery of anticancer molecules. A wide range of containers and carriers, such as liposomes, micelles, gold nanoparticles, dendrimers, and magnetic nanoparticles, have been targeted to cancerous tumors using folic acid as the targeting component [ 269 ].

Other small molecule compounds are used to target specific cancers. For example, the asialoglycoprotein receptor (ASGPR) is overexpressed on the surfaces of hepatocytes in hepatocellular carcinoma [ 270 ]. Studies have shown that modifying the surface of a container or carrier with D-galactose residues or N-acetylgalactosamine effectively targets the delivery system to hepatocytes via ASGPR [ 271 , 272 ]. Surface modification with lactose is also used to target hepatocytes via ASGPR [ 273 , 274 ]. It has been shown that the cells of some cancers, such as brain cancer, colon cancer, melanoma, and breast cancer, overexpress sigma receptors [ 275 , 276 , 277 , 278 ]. The conjugation of containers and carriers with anisamide, which has a high affinity for sigma receptors, has been proposed as a means of targeting sigma receptors [ 279 ]. More than ten different delivery systems that use the sigma receptor as a target have already been developed [ 280 ].

3.4. Small Molecule-Drug Conjugates

Small molecule-drug conjugates (SMDC) are a drug delivery system without using nanocontainers and nanocarriers [ 281 ]. Typically, SMDCs consist of an anticancer agent coupled to a targeting ligand via a linker capable of being cleaved under various stimuli (See Section 4 for details) [ 282 ]. Besides the binding ability to the cellular target, the spacer also increases the hydrophilicity of the conjugate [ 281 , 283 ]. Antibodies [ 284 , 285 ]) and aptamers [ 286 ], peptides [ 287 , 288 ] and low molecular weight compounds [ 289 , 290 ] can act as a targeting ligand. Although SMDCs do not exhibit an EPR effect and, therefore, do not passively accumulate in solid tumors, they nevertheless passively perfuse the cancer mass more thoroughly and faster than nanoparticles [ 282 ]. When creating SMDCs, it should be taken into account that they have a short half-life compared to nanoparticles [ 291 ]. Various SMDCs have been developed and are being used successfully in cancer therapy [ 292 ].

Various small interfering RNA (siRNA) conjugates are also used for cancer gene therapy [ 293 , 294 ]. Chemical modification of siRNAs (at the 2′ position, at the ribose ring, or using nucleotide phosphorothioate) improves their stability, increases cell specificity, and reduces off-target effects [ 295 ]. siRNA conjugates show efficient RNA interference both in vitro and in vivo [ 296 ].

4. Stimuli-Responsive Drug Release

To be delivered to the desired area of the body, an active substance is either encapsulated in the delivery system or covalently associated with it. There are two main mechanisms of drug release: firstly, as a result of endocytosis or fusion with the cell membrane (in the case of lipid delivery systems), and, secondly, under the influence of stimuli [ 297 ]. These stimuli can be internal and thus inherent to the affected area of the body, such as changes in enzyme levels, pH, and temperature; or external, such as a magnetic field, ultrasound, and light [ 298 ].

4.1. Enzyme-Sensitive Release

The expression pattern of enzymatic proteins in the tumor may be altered in some types of cancer [ 299 ]. There are two main approaches to controlling the release of a drug from delivery systems under the action of enzymes [ 300 ]:

• The drug is conjugated to the delivery system with a linker cleaved by an enzyme that is overexpressed in the tumor environment.
• Enzyme cleavage sites are embedded into the envelope of the scaffolds, thereby destroying the envelope near or inside the tumor and releasing the encapsulated drug.

Several materials sensitive to various enzymes have been obtained to date [ 300 ]. For example, an octapeptide sensitive to metalloproteinase has been developed and used as a linker [ 301 ]. Other enzymes that have been used for drug release include phospholipase [ 302 ], α-amylase [ 303 ], glucose oxidase [ 304 ], and cancer-associated proteases [ 305 ].

4.2. pH-Sensitive Release

Due to changes in the metabolic environment, the extracellular pH is usually lower in tumors (≈6.5) than in blood and normal tissues (≈7.4) [ 306 ]. The pH level in tumor tissue is not uniform; intracellular pH is similar in tumor and normal tissues, and extracellular pH is more acidic [ 307 ]. This difference in pH means that a cellular transmembrane gradient is formed between normal tissue and tumor tissue. Exploiting this gradient allows drugs to be directly delivered into the cytosol of cancer cells, which are weak electrolytes with the corresponding pKa [ 307 ]. A weakly acidic drug in protonated form can freely penetrate through a cell membrane, reach a region with a more basic pH, and then become trapped inside the cell, leading to a significant difference in drug concentration between normal and tumor tissues.

There are two main approaches to using pH as a stimulus for drug release. The first approach is to introduce various chemical bonds, which are hydrolyzed and destroyed under conditions of acidic pH, into the delivery system. Most often, bonds are introduced into the delivery system of drugs, as presented in Table 1 .

List of pH-labile bonds.

The second approach exploits the ability of different polymers to be protonated/deprotonated at different pH levels. At physiological pH, such polymers remain deprotonated/deionized, but under acidic conditions, the polymers are protonated or change their charge, causing structural transformation or disintegration in the delivery system and the subsequent release of the encapsulated drugs [ 16 ]. Conjugation of urocanic acid with various polymers makes it possible to give them pH-dependent properties [ 308 , 309 ].

4.3. Temperature-Sensitive Release

Mild hyperthermia plays a pivotal role in changing the tumor microenvironment by increasing blood-flow velocity, oxygenation, and vascular permeability [ 319 ]. It has been shown that most delivery systems, up to 400 nm in diameter, can extravasate from the tumor environment into tumor cells when heated to 42 °C in vitro [ 320 ]. Moreover, the inclusion of thermosensitive fragments in a delivery system changes their properties in areas with elevated temperatures, leading to the release of the encapsulated drug. At a specific temperature, lipid carrier systems that contain lysolipids or oligoglycerol undergo a gel-liquid phase transition involving the release of the active substance [ 321 ]. Several thermosensitive polymers have been developed with a lower critical solution temperature (LCST) of about 40 °C. Below this temperature, the polymers are soluble in water but become insoluble in water above this temperature. Such polymers are used in anticancer-molecule delivery systems [ 320 ]. Table 2 lists some characteristic polymers used for temperature-sensitive release from delivery systems and indicates their LCST [ 322 ]. Hyperthermia in the environment of a tumor can also be caused externally, for example, by applying an alternating magnetic field around the tumor, causing magnetic nanoparticles to heat up and creating hyperthermia in the area. Furthermore, in an alternating magnetic field, the magnetic nanoparticles themselves have a strong cytotoxic effect on cancer cells [ 239 , 323 ]. The creation of hyperthermia in the desired area can also be achieved using a laser; photothermal inducing agents can be included in the structure of delivery systems and absorb emitted light and convert it into local heat [ 324 ]. The most commonly used photothermal material is gold nanoparticles [ 325 , 326 ].

The LCST of polymers in an aqueous solution.

4.4. Other Stimuli

The redox environment of tumor cells is changed by an increased level of glutathione (GSH) usually 4 times higher than in normal cells [ 331 ]. Glutathione regulates the reducing environment of the cell by forming and destroying disulfide bonds via reaction with the excess of reactive oxygen species (ROS) [ 332 , 333 ]. Redox-sensitive delivery systems usually contain disulfide, diselenide, or succinimide-thioether bonds [ 334 , 335 , 336 , 337 ]. Under the influence of glutathione, these bonds are reduced and destroyed ( Table 3 ).

List of labile groups sensitive to GSH and ROS.

Some studies have shown, indicating oncogenic transformation compared to normal cells, that cancer cells constantly generate high levels of intracellular ROS, such as hydrogen peroxide, hydroxyl radical, and superoxide anion [ 338 ]. Some ROS-sensitive transport systems have also been developed to exploit this abnormal biochemical change [ 339 , 340 ], and all of the delivery systems that have been developed contain ROS-sensitive linkers. In essence, the linkers are based on the organic compounds of chalcogens (S, Se, Te), such as thioesters [ 341 , 342 ], thioketals [ 343 ], diselenides [ 344 ], monoselenides [ 345 ], and tellurides [ 346 ]. Under the action of ROS, the two-stage oxidation of thioesters, monoselenides, and tellurides occurs, first to the oxidation state of +4 and then to +6. Accordingly, delivery systems that contain these groups undergo a phase transition from hydrophobic compounds to more hydrophilic ones [ 347 ]. Linkers that contain other ROS-sensitive groups are oxidized with bond cleavage ( Table 3 ). As in the case of pH-sensitive release, delivery systems with ROS-sensitive bonds release their drugs either via a phase transition or the breaking of a chemical bond.

It has been shown that monosulfides and monoselenides are only ROS-sensitive linkers despite disulfides and diselenides being redox- and ROS-sensitive linkers [ 348 , 349 ]. Arylboronic ethers are widely used as the ROS-sensitive linker as, under the action of hydrogen peroxide, arylboronic esters are oxidized to boronic acid and phenol, and the bond in the para-position of the aryl ring is destroyed [ 350 ].

5. Conclusions

Over recent decades, tremendous progress has been made in the field of targeted delivery in cancer therapy. Several targeted-delivery drugs have been approved and included in clinical practice. Delivery systems can target different parts of a tumor using specific targeting fragments and avoid the problems associated with multidrug resistance. With detailed studies of the physiological differences between normal and diseased tissues, it is possible to develop target-specific drug delivery systems able to respond to local stimuli. However, some aspects require a more detailed study. In fact, a deeper understanding of the EPR effect, of the interactions between nanoparticles and cells, of tumor targeting, and of the metastatic microenvironment is certainly needed. Moreover, further insights into the biodistribution, pharmacokinetics, toxicity, and role of delivery systems in therapeutic protocols are essential if they are to become part of standard-treatment algorithms. Adverse immunological reactions also require careful consideration when using targeted delivery. Only once studies into these factors are complete will it be possible to unleash the full potential of cytostatic drug-delivery systems in cancer therapy.

Acknowledgments

The Sechenov First Moscow State Medical University and the University of Turin are warmly acknowledged.

Author Contributions

Writing—original draft preparation, V.V.V. and R.N.A.; writing—review and editing, L.J. and G.C.; supervision, A.E.N. and G.C. All authors have read and agreed to the published version of the manuscript.

This research was supported by the Ministry of Science and Higher Education of the Russian Federation: World-Class Research Center, Sechenov First Moscow State Medical University, and the University of Turin.

Conflicts of Interest

The authors declare no conflict of interest.

Publisher’s Note: MDPI stays neutral with regard to jurisdictional claims in published maps and institutional affiliations.

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• 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

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• 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.

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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.

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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).

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Michael J. Mitchell, Margaret M. Billingsley & Rebecca M. Haley

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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.

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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).

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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

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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

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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.

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The authors declare that they have no competing interests.

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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).

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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

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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

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Received : 19 July 2018

Accepted : 25 August 2018

Published : 19 September 2018

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

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Revolutionary molecular device unleashes potential for targeted drug delivery and self-healing materials

In a new breakthrough that could revolutionise medical and material engineering, scientists have developed a first-of-its-kind molecular device that controls the release of multiple small molecules using force.

The researchers from The University of Manchester describe a force-controlled release system that harnesses natural forces to trigger targeted release of molecules, which could significantly advance medical treatment and smart materials.

The discovery, published today in the journal Nature , uses a novel technique using a type of interlocked molecule known as rotaxane. Under the influence of mechanical force -- such as that observed at an injured or damaged site -- this component triggers the release of functional molecules, like medicines or healing agents, to precisely target the area in need. For example, the site of a tumour.

It also holds promise for self-healing materials that can repair themselves in situ when damaged, prolonging the lifespan of these materials. For example, a scratch on a phone screen.

Guillaume De Bo, Professor of Organic Chemistry at The University of Manchester, said: "Forces are ubiquitous in nature and play pivotal roles in various processes. Our aim was to exploit these forces for transformative applications, particularly in material durability and drug delivery.

"Although this is only a proof-of-concept design, we believe that our rotaxane-based approach holds immense potential with far reaching applications -- we're on the brink of some truly remarkable advancements in healthcare and technology."

Traditionally, the controlled release of molecules with force has presented challenges in releasing more than one molecule at once, usually operating through a molecular "tug of war" game where two polymers pull at either side to release a single molecule.

The new approach involves two polymer chains attached to a central ring-like structure that slide along an axle supporting the cargo, effectively releasing multiple cargo molecules in response to force application. The scientists demonstrated the release of up to five molecules simultaneously with the possibility of releasing more, overcoming previous limitations.

The breakthrough marks the first time scientists have been able to demonstrate the ability to release more than one component, making it one of the most efficient release systems to date.

The researchers also show versatility of the model by using different types of molecules, including drug compounds, fluorescent markers, catalyst and monomers, revealing the potential for a wealth of future applications.

Looking ahead, the researchers aim to delve deeper into self-healing applications, exploring whether two different types of molecules can be released at the same time. For example, the integration of monomers and catalysts could enable polymerization at the site of damage, creating an integrated self-healing system within materials.

They will also look to expand the sort of molecules that can be released.

Prof De Bo said: "We've barely scratched the surface of what this technology can achieve. The possibilities are limitless, and we're excited to explore further."

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• Lei Chen, Robert Nixon, Guillaume De Bo. Force-controlled release of small molecules with a rotaxane actuator . Nature , 2024; 628 (8007): 320 DOI: 10.1038/s41586-024-07154-0

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ORIGINAL RESEARCH article

Development and therapeutic evaluation of 5d3(cc-mln8237)3.2 antibody-theranostic conjugates for psma-positive prostate cancer therapy.

• 1 Department of Radiology and Radiological Science, The Johns Hopkins University School of Medicine, Baltimore, United States
• 2 Department of Ophthalmology, The Johns Hopkins University School of Medicine, Baltimore, United States
• 3 The Solomon H Snyder Department of Neuroscience, Johns Hopkins University School of Medicine, Baltimore, United States
• 4 Laboratory of Structural Biology, Institute of Biotechnology of the Czech Academy of Sciences, Vestec, Czechia
• 5 Department of Molecular and Comparative Pathobiology, School of Medicine, Johns Hopkins Medicine, Baltimore, United States
• 6 Department of Oncology, The Sidney Kimmel Comprehensive Cancer Center, The Johns Hopkins University School of Medicine, Baltimore, United States
• 7 Department of Pathology, University Medical Center Utrecht, Utrecht, Netherlands, Netherlands
• 8 Department of Pharmacology and Molecular Sciences, School of Medicine, Johns Hopkins Medicine, Baltimore, United States

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Prostate cancer (PC) is an aggressive cancer that can progress rapidly and eventually become castrateresistant prostate cancer (CRPC). Stage IV metastatic castrate-resistant prostate cancer (mCRPC) is an incurable late-stage cancer type with a low 5-year overall survival rate. Targeted therapeutics such as antibody-drug conjugates (ADCs) based on high-affinity monoclonal antibodies and potent drugs conjugated via smart linkers are being developed for PC management. Conjugating further with in vitro or in vivo imaging agents, ADCs can be used as antibody-theranostic conjugates (ATCs) for diagnostic and image-guided drug delivery. In this study, we have developed a novel ATC for PSMA(+) PC therapy utilizing (a) anti-PSMA 5D3 mAb, (b) Aurora A kinase inhibitor, MLN8237, and (c) for the first time using tetrazine (Tz) and trans-cyclooctene (TCO) click chemistry-based conjugation linker (CC linker) in ADC development. The resulting 5D3(CC-MLN8237)3.2 was labeled with suitable fluorophores for in vitro and in vivo imaging. The products were characterized by SDS-PAGE, MALDI-TOF, and DLS and evaluated in vitro by optical imaging, flow cytometry, and WST-8 assay for cytotoxicity in PSMA(+/-) cells. Therapeutic efficacy was determined in human PC xenograft mouse models following a designed treatment schedule. After the treatment study animals were euthanized, and toxicological studies, complete blood count (CBC), blood clinical chemistry analysis, and H&E staining of vital organs were conducted to determine side effects and systemic toxicities. The IC50 values of 5D3(CC-MLN8237)3.2-AF488 in PSMA(+) PC3-PIP and PMSA(-) PC3-Flu cells are 8.17 nM and 161.9 nM, respectively. Pure MLN8237 shows 736.9 nM and 873.4 nM IC50 values for PC3-PIP and PC3-Flu cells, respectively. In vivo study in human xenograft mouse models confirmed high therapeutic efficacy of 5D3(CC-MLN8237)3.2-CF750 with significant control of PSMA(+) tumor growth with minimal systemic toxicity in the treated group compared to PSMA(-) treated and untreated groups. Approximately 70% of PSMA(+) PC3-PIP tumors did not exceed the threshold of the tumor size in the surrogate Kaplan-Meyer analysis. The novel ATC successfully controlled the growth of PSMA(+) tumors in preclinical settings with minimal systemic toxicities. The therapeutic efficacy and favorable safety profile of novel 5D3(CC-MLN8237)3.2 ATC demonstrates their potential use as a theranostic against aggressive PC.

Keywords: prostate cancer, PSMA -prostate-specific membrane antigen, Image-guided drug delivery, targeted therapy, theranostics, Aurora A kinase inhibition

Received: 13 Feb 2024; Accepted: 15 Apr 2024.

Copyright: © 2024 Liatsou, Assefa, Liyanage, Surasinghe, Nováková, Bařinka, Gabrielson, Raman, Artemov and Hapuarachchige. 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) or licensor 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: Sudath Hapuarachchige, Department of Radiology and Radiological Science, The Johns Hopkins University School of Medicine, Baltimore, United States

Disclaimer: 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.

Revolutionizing the Treatment of Idiopathic Pulmonary Fibrosis: From Conventional Therapies to Advanced Drug Delivery Systems

• Review Article
• Published: 08 April 2024
• Volume 25 , article number  78 , ( 2024 )

Cite this article

• Sanskriti Singh 1 &
• Sarika Wairkar   ORCID: orcid.org/0000-0002-0124-1741 1  

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Idiopathic pulmonary fibrosis (IPF) is a chronic and progressive interstitial lung disease that has been well-reported in the medical literature. Its incidence has risen, particularly in light of the recent COVID-19 pandemic. Conventionally, IPF is treated with antifibrotic drugs—pirfenidone and nintedanib—along with other drugs for symptomatic treatments, including corticosteroids, immunosuppressants, and bronchodilators based on individual requirements. Several drugs and biologicals such as fluorofenidone, thymoquinone, amikacin, paclitaxel nifuroxazide, STAT3, and siRNA have recently been evaluated for IPF treatment that reduces collagen formation and cell proliferation in the lung. There has been a great deal of research into various treatment options for pulmonary fibrosis using advanced delivery systems such as liposomal-based nanocarriers, chitosan nanoparticles, PLGA nanoparticles, solid lipid nanocarriers, and other nanoformulations such as metal nanoparticles, nanocrystals, cubosomes, magnetic nanospheres, and polymeric micelles. Several clinical trials are also ongoing for advanced IPF treatments. This article elaborates on the pathophysiology of IPF, its risk factors, and different advanced drug delivery systems for treating IPF. Although extensive preclinical data is available for these delivery systems, the clinical performance and scale-up studies would decide their commercial translation.

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Pulmonary Fibrosis: Unveiling the Pathogenesis, Exploring Therapeutic Targets, and Advancements in Drug Delivery Strategies

Kirti Aggarwal, Sandeep Arora & Kalpana Nagpal

Idiopathic Pulmonary Fibrosis: New and Emerging Treatment Options

Richard J. Hewitt & Toby M. Maher

Targeted Therapy for Idiopathic Pulmonary Fibrosis: Where To Now?

Sunad Rangarajan, Morgan L. Locy, … Victor J. Thannickal

Abbreviations

Idiopathic pulmonary fibrosis

Gastroesophageal reflux disease

Telomerase reverse transcriptase

Coronavirus disease 2019

Food and Drug Administration

Zeta potential

Truncated basic fibroblast growth factor

Spliceosome-associated factor

Small interfering ribonucleic acid

Polylactic-co-glycolic acid

Polylactic acid

Polyvinyl alcohol

Signal transducer and activator of transcription-3

Alcaligenes xylosoxidans exopolysaccharide

(3‑Aminopropyl) triethoxysilane

Self-microemulsifying drug delivery system

Signal transducer and activator of transcription 3

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Singh, S., Wairkar, S. Revolutionizing the Treatment of Idiopathic Pulmonary Fibrosis: From Conventional Therapies to Advanced Drug Delivery Systems. AAPS PharmSciTech 25 , 78 (2024). https://doi.org/10.1208/s12249-024-02793-y

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Received : 14 December 2023

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DOI : https://doi.org/10.1208/s12249-024-02793-y

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