Block Copolymer Micelles in Nanomedicine Applications - Chemical

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Cite This: Chem. Rev. 2018, 118, 6844−6892

Block Copolymer Micelles in Nanomedicine Applications Horacio Cabral,† Kanjiro Miyata,‡ Kensuke Osada,† and Kazunori Kataoka*,§,∥ Department of Bioengineering, Graduate School of Engineering, and ‡Department of Materials Engineering, Graduate School of Engineering, The University of Tokyo, 7-3-1 Hongo, Bunkyo-ku, Tokyo 113-8656, Japan § Innovation Center of NanoMedicine, Kawasaki Institute of Industrial Promotion, 3-25-14, Tonomachi, Kawasaki-ku, Kawasaki 210-0821, Japan ∥ Policy Alternatives Research Institute, The University of Tokyo, 7-3-1 Hongo, Bunkyo-ku, Tokyo 113-0033, Japan Downloaded via 5.189.202.108 on August 9, 2018 at 15:09:48 (UTC). See https://pubs.acs.org/sharingguidelines for options on how to legitimately share published articles.

†

ABSTRACT: Polymeric micelles are demonstrating high potential as nanomedicines capable of controlling the distribution and function of loaded bioactive agents in the body, effectively overcoming biological barriers, and various formulations are engaged in intensive preclinical and clinical testing. This Review focuses on polymeric micelles assembled through multimolecular interactions between block copolymers and the loaded drugs, proteins, or nucleic acids as translationable nanomedicines. The aspects involved in the design of successful micellar carriers are described in detail on the basis of the type of polymer/payload interaction, as well as the interplay of micelles with the biological interface, emphasizing on the chemistry and engineering of the block copolymers. By shaping these features, polymeric micelles have been propitious for delivering a wide range of therapeutics through effective sensing of targets in the body and adjustment of their properties in response to particular stimuli, modulating the activity of the loaded drugs at the targeted sites, even at the subcellular level. Finally, the future perspectives and imminent challenges for polymeric micelles as nanomedicines are discussed, anticipating to spur further innovations.

CONTENTS 1. Introduction 2. Block Copolymer Self-Assemblies for Drug Delivery 3. Design Criteria of Block Copolymers for SelfAssembly of Polymeric Micelles 3.1. Shell-Forming Segments 3.2. Core-Forming Segments 3.3. Stimuli Sensitivity 3.4. Modification of the Surface of Polymeric Micelles with Ligand Molecules 4. Polymeric Micelles as Carriers of Small Drugs 4.1. Polymeric Micelles Physically Incorporating Hydrophobic Drugs 4.2. Polymeric Micelles Reversibly Conjugating Hydrophobic Drugs to the Core-Forming Segments 4.2.1. Drug Loading in Polymeric Micelles via Hydrolytically Labile Bonds 4.2.2. Drug Loading in Polymeric Micelles via Redox Labile Bonds 4.3. Self-Assembly of Polymeric Micelles via Polymer−Metal Complexation 5. Polymeric Micelles Assembled via Electrostatic Interactions: Polyion Complex (PIC) Micelles 5.1. PIC Micelles Loaded with Organic Chemicals 5.2. PIC Micelles Loaded with Proteins 5.3. PIC Micelles Loaded with Oligonucleotides

© 2018 American Chemical Society

5.3.1. Biological and Structural Characteristics of Oligonucleotides 5.3.2. Design Criteria of Block Copolymers for Oligonucleotide Delivery 5.3.3. Therapeutic Potential of Oligonucleotide-Loaded Micelles 5.4. PIC Micelles Loaded with pDNA 5.4.1. Biological and Structural Characteristics of pDNA for Encapsulation 5.4.2. Structure and Functionality of Polyplex Micelles for Managing Each Step in Systemic Delivery 5.4.3. Design Criteria of Block Copolymers toward Systemic Gene Therapy 5.4.4. Removal of Free Polymers for Safe Delivery System 5.4.5. pDNA Packaging into Rod-, Globule-, or Toroid-Shapes 5.5. PIC Micelles Loaded with mRNA 6. Future Perspectives Author Information Corresponding Author ORCID Notes Biographies Acknowledgments

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Received: March 28, 2018 Published: June 29, 2018 6844

DOI: 10.1021/acs.chemrev.8b00199 Chem. Rev. 2018, 118, 6844−6892

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1. INTRODUCTION To elicit a given therapeutic response, the indispensable amount of active drug should reach its site of action, while the effective dose must be maintained at the target site for a definite time. However, for most drugs, such a process poses several inexorable barriers, including their fast degradation in the harsh in vivo environment, inadequate pharmacokinetics, lack of selectivity for the targeted tissues, and widespread biodistribution after systemic administration, which is a potential cause of toxicity. For example, most anticancer drugs are water-insoluble molecules with low-molecular weight (MW), which distribute to the whole body after intravenous injection, critically limiting the dosage due to the underlying risk of side effects, while proteins and genes designed as therapeutic agents present unfavorable pharmacokinetics, poor internalization by targeted cells, and are rapidly degraded. Thus, by using targeted nanomedicines, therapeutic strategies capable of controlling the spatiotemporal distribution of drugs, as well as avoiding side effects, could be designed as effective approaches to satisfy unmet medical demands, or constitute the foundation for unprecedented or revolutionary treatments.1−5 Successful nanomedicines should stably circulate in the blood compartment, avoiding unspecific interactions with blood components, as well as with the reticuloendothelial systems (RES),6 and selectively extravasating at the diseased site, which is followed by the uptake by target cells and the intracellular release of the cargo. Nanomedicines with macromolecular or particulate nature show low volume of distribution, indicating the limited extravasation of these carriers in healthy tissues because of the continuous and tight lining of blood vessels in these tissues1−6 (Figure 1).

Figure 2. (A) Low-molecular-weight drugs can extravasate in healthy tissues, which leads to undesirable side effects. Conversely, nanomedicines do not pass to healthy tissues due to their relatively large size. (B) Nanomedicines selectively accumulate in tumors by the enhanced permeability and retention (EPR) effect, which is based on the leaky vasculature of tumors with poorly aligned endothelial cells and large fenestrations, allowing nanomedicines to leak extensively into tumor tissues (upper panel). Moreover, slow venous return and poor lymphatic clearance in tumor tissues cause the retention of nanomedicines. In tumors with low permeability (lower panel), vascular barriers, such as pericytes, and interstitial barriers, such as fibrosis, can limit the effective size of nanomedicines for extravasation and penetration into tumors. (C) Nanomedicines can also extravasate by vascular bursts spontaneously forming in tumor microvessels, providing an increased pressure-driven accumulation and retention of sub-100 nm nanomedicines, even in poorly permeable tumors. The openings are defined as dynamic vents, and the transient extravasation of nanomedicines are called eruptions. (D) Ligand-installed nanomedicines targeting vascular receptors can promote the extravasation of nanomedicines by transcytosis into poorly permeable tumors or even nonpermeable diseased tissues.

medicines with biocompatible materials, such as poly(ethylene glycol) (PEG), is essential to avoid adsorption of opsonin proteins, which facilitate their recognition and elimination from the bloodstream by the RES, principally in liver and spleen, resulting in lowered unspecific biodistribution and impaired pharmacokinetic properties.6,9,10 Besides the effective shielding of the surface of nanomedicines with biocompatible polymers, other approaches for extending the blood life of nanomedicines have been recently reported, including the selective interaction of nanomedicines with specific plasma proteins, such as apolipoproteins, to form a protective protein corona11,12 or the installation of “don’t eat me” signal (CD47) on the surface of nanomedicines to avoid the uptake by macrophages.13,14 In addition, the size and the charge of the nanomedicines affect their biodistribution. Nanomedicines smaller than ∼150 nm tend to accumulate in the liver, while at sizes larger than ∼150 nm, they are retained in the spleen.15,16 Positively charged systems are rapidly sequestered in the liver, spleen, and lungs, whereas neutral or slightly negative charged nanomedicines, which have much lower opsonization rate, present prolonged circulation in the bloodstream.15 Beyond stable circulation in the body, nanomedicines must access the target tissues. In this way, the preferential accumulation of macromolecular nanomedicines in solid

Figure 1. Nanomedicines encounter several barriers en route to their therapeutic targets. The design parameters of nanomedicines affect their distribution in the body and allow them to overcome these barriers.

Thus, stable nanomedicine formulations can effectively eliminate accidental extravasation, while avoiding leakage, as well as degradation, of the loaded drugs in blood (Figure 2A). The macromolecular character of the nanomedicines also substantially reduces the renal clearance of the loaded drugs, as the threshold for glomerular filtration in kidney is approximately 50 kDa,7 or 6 nm,8 which results in the extension of the drug’s half-life. Moreover, the modification of the nano6845

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tumors is a significant advantage for developing targeted antitumor therapies. This prominent accumulation of drug carriers in tumors is mediated by the so-called enhanced permeability and retention (EPR) effect, which is based on the high permeability of the malignant vasculature to macromolecules, and the impaired lymphatic drainage in tumor tissues, which increases the retention of the macromolecules (Figure 2B).17 The transport of nanomedicines through the blood vessels of tumors has been indicated to occur accross interendothelial gaps and transendothelial channels.17,18 Moreover, tumor vessels show fenestrations and vesicular vacuolar organelles with diameters ranging from 50 to 100 nm,19−23 which could also facilitate the transvascular transport of smallsized nanomedicines into tumors. The efficiency of the EPR effect for selective tumor targeting is highly dependent on the pathophysiological characteristics, including the degree of angiogenesis and lymphangiogenesis of tumors, the density of cells and stromal components, and the intratumoral pressure.17,24,25 Because of the heterogeneous vasculature and permeability of tumors, the perfusion and distribution of nanosized carriers within the tumor mass could be rendered uneven. Therefore, to properly exploit the EPR effect for tumor targeting and homogeneously distribute adequate dosages of drugs, the carriers should have a prolonged blood half-life and be sufficiently small for successfully extravasating and deeply penetrating from the blood compartment into tumor tissues.1−5 This effective size has been reported to be between 10 and 100 nm in diameter, although for tumors with low permeability, such as pancreatic and gastric cancer, nanomedicines with sizes below 50 nm should be considered (Figure 2B).26−28 Nanomedicines in the sub-100 nm size can also take advantage of dynamic vascular bursts in the tumor vasculature mediated by intratumoral and vascular pressure gradients, which facilitate the access of nanomedicines to tumor interstitium (Figure 2C).29 Moreover, nanomedicines can be equipped with ligands targeting vascular receptors, which can facilitate the extravasation from the bloodstream into poorly permeable tumors or even nonpermeable diseased tissues (Figure 2D).30−32 In addition, tumors present specific features, such as acidic intratumoral environment ranging from pH 6.5 to pH 7.0,33 or expression of specific biomarkers,34,35 which can be used for triggering specific functions of nanomedicines, increasing their selectivity toward cancerous tissues. Following accumulation in the targeted tissues, nanomedicines are internalized by the cells (Figure 3), where they can be activated for achieving precise actions upon sensing specific stimuli. For example, as nanosized materials are mainly internalized by endocytosis, the pH drop inside these endosomal compartments (pH 5−5.5) serves as a specific trigger for programmed functions of the nanomedicines, such as drug release35−38 or disruption of endosomal membranes.39−41 Moreover, because endosomes ultimately merge with lysosomes, where lysosomal enzymes may degrade the drugs or genes, nanomedicines controlling the intracellular trafficking and the subcellular targeting of their cargo could substantially improve the efficiency of delivery and the therapeutic response. Such subcellular delivery is particularly important for constructing nanomedicines capable of overcoming drug resistance by circumventing the defense mechanisms of cancer cells,42,43 as well as for nanomedicines delivering nucleic acids to the cytoplasm or the nucleus,41,44 which otherwise would not be able to escape from endosomal

Figure 3. After accumulation in tissues, the nanomedicines are internalized by the cells, where they can elicit further subcellular targeting. The surface of nanomedicines can be functionalized with ligand moieties directed to specific receptors on the cells. Moreover, upon endocytosis, the constituents of the nanomedicines can be engineered to allow endosomal escape of the cargo and subcellular delivery.

compartments to reach their targets. Additionally, the cellular recognition and intracellular delivery of nanomedicines can be further improved by including ligand molecules directed to specific cell surface receptors.45 Various nanomedicine strategies, including liposomes, polymer−drug conjugates, nanoparticles, and polymeric micelles, have been considered for selective drug delivery to tumors,1−5 while improving solubility of drugs and tailoring pharmacokinetics with the premise of reducing toxicity, and the clinical translation of several of these nanomedicines further supports the approach as an effective therapeutic modality.46−48 Among clinical translationable nanomedicines, polymeric micelles have demonstrated unique advantages for incorporating a wide range of bioactive molecules and overcoming any biological barriers en route to the therapeutic target.

2. BLOCK COPOLYMER SELF-ASSEMBLIES FOR DRUG DELIVERY Over recent decades, block copolymers comprising immiscible blocks and their self-assemblies have been extensively studied from basic to applied science. Such interest has been driven by the ability of block copolymers featuring a narrow distribution of MW and compartmentalized segments for distinct phase separation to form highly ordered multimolecular architectures, including micelles and vesicles. The self-assembling behavior of amphiphilic block copolymers can be well described from the thermodynamic standpoint, as illustrated in Figure 4.3 Amphiphilic block copolymers assemble with each other to decrease the interfacial area of insoluble blocks for lowered interfacial free energy. On the other hand, the increase in the assembling number of block copolymers is accompanied by the increase in insoluble core size, leading to the stretching of the blocks forming the core, as their connection with the blocks forming the shell should be aligned at the interface to circumvent thermodynamically unfavorable phase mixing. In addition, the association of the block copolymers also increases the density of shell-forming segments directed toward a stretched conformation. The stretching of both shell- and core-forming segments, as well as 6846

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entropy loss determines the thermodynamically stable size of polymeric micelles. Once micellization happens, the segregated (or concentrated) core-forming segments can acquire additional proximity forces, such as van der Waals forces and hydrogen bonding, for further core stabilization. In addition to the above fundamental understanding, additional aspects can be considered for the assembly of polymeric micelles, particularly highlighting their chemical structure. For example, block ionomers, that is, block copolymers combining neutral and charged segments, which are completely soluble in aqueous milieu, can also form multimolecular self-assemblies, polyion complex (PIC) micelles (Figure 5), when they are mixed with oppositely charged ion-containing block copolymers (or block ionomers).49,50 The PIC formation is mainly driven by the release of counterions triggered by ion-pairing between oppositely charged ionomer segments, apparently similar to the exclusion of poor solvent molecules from insoluble segments upon the self-assembly of

Figure 4. Crucial factors for determining the size of polymeric micelles.

their connection alignment at the interface upon multimolecular assembly, apparently decrease the conformational entropy of block copolymers, which is compensated by the decreased interfacial free energy. Thus, the balance between the decreased interfacial free energy and the conformational

Figure 5. Self-assembled polymeric micelles of block copolymers embody a versatile platform for loading bioactive molecules by controlling the interaction of the payloads with the segments forming the core (upper panel). The hydrophilic shell, high loading efficiency, ability to introduce ligands, and their relatively small size are substantial advantages of polymeric micelles for acting at the biological interface (upper panel). The considerations for the development of medicinal products based on polymeric micelles are described in the lower panel.53 6847

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proteins in blood, reducing their immunogenicity, and protecting them from proteolytic degradation,54 is widely chosen for developing polymeric micelles because of its hydrophilicity (three water molecules are bound per oxyethylene unit), linearity, chain flexibility, lack of charge, and availability in a wide range of MWs with narrow MW distribution. Indeed, the ring-opening anionic polymerization of ethylene oxide offers well-defined PEG with MW distribution (Mw/Mn) around 1.01, which facilitates constructing narrowly distributed block copolymers capable of assembling monodisperse polymeric micelles with low polydispersity index. The end-groups of PEG can be further controlled by choosing the appropriate initiator and ω-endcapping moiety, which is a significant benefit for preparing block copolymers or branched polymers, or even installing ligand molecules on the PEG shell of polymeric micelles. PEG prevents aggregation of polymeric micelles in vivo, as well as during storage, due to the steric repulsive effect. The formation of an effective shell hindering the surface of particles depends on the MW of the PEG, the surface density, and the conformation of the PEG chains on the surface.10 As a consequence, the surface for recognition by the immune system through opsonization is reduced. While PEGs with MW of 20−50 kDa are used for conjugating small molecules, oligonucleotides, or siRNA, to avoid their renal clearance,54,55 PEGs with 1−20 kDa are used for conjugating larger biomolecules, such as antibodies, for coating nanoparticles, or for constructing self-assembled nanostructures.54,55 The steric hindrance of PEG also allows shielding the charge of the core of polymeric micelles, which results in the neutralization of their zeta-potential, reducing charge-related interactions in the body.10,54,55 Because PEG is nonbiodegradable, the effective clearance of the polymer is preferable for avoiding effects from long-term exposure. The MW of PEG determines its clearance from the body, with PEGs larger than 60 kDa being slowly excreted predominantly through the liver, and PEGs with MW below 60 kDa showing faster excretion through the kidneys.54,55 A concern about PEG retention in the body is the prolonged accumulation of PEG within lysosomes following chronic administration, as some preclinical studies have indicated the formation of PEG-containing intracellular vesicles in the kidneys of animal models.55 Moreover, several studies have indicated that specific antibody responses can be generated against PEG, which can reduce the half-life of PEGylated drugs, so-called accelerated blood clearance (ABC) phenomena,56 and, in rare cases, trigger allergic responses in particular individuals.55 The ABC phenomena have been reported for a wide-range of formulations, including PEGylated proteins and liposomes, in both animal models and humans.57,58 Nevertheless, various polymeric micelle formulations equipped with PEG shielding do not present ABC phenomena.59−63 Improved understanding of the means by which anti-PEG antibodies (IgG and IgM) bind to PEGylated materials suggests that these antibodies do not bind to the main chain of PEG.63 Instead, the affinity of anti-PEG IgM decreases at high PEG concentrations, and the binding of anti-PEG IgM was shown to depend on the length of the hydrophobic block conjugated to the PEG chain.62,63 Thus, nanomedicines with highly dense PEG shielding, such as polymeric micelles, may avoid capture by anti-PEG IgM. Moreover, because the number of anti-PEG IgM in the body is limited, for example, approximately 1012 in mouse serum, the ABC phenomena would be negligible when

amphiphilic block copolymers. Further details regarding PIC micelles can be found in the section related to oligonucleotideloaded PIC micelles (section 5.3). An important criterion for constructing self-assembled systems for drug delivery is their stability against dissociation under highly diluted conditions in body fluids. The static stability of self-assemblies can be well documented by the critical association (or micelle) concentration (CAC or CMC).51 As indicated by lower CAC values,51,52 polymeric micelles are generally more stable than low-MW surfactant micelles because of the greater interfacial free energy derived from the larger insoluble segments. Furthermore, the segregation of the core-forming segments in the micellar core can generate a variety of intermolecular forces to lower the CAC of the micelles. The significant contribution of these forces to micelle stabilization causes the hysteresis in CAC of polymeric micelles. It should also be noted that the drug payloads can additionally stabilize the micellar core through their interactions with the core-forming segments, where lowMW hydrophobic drugs molecularly dispersed in the core can act as filler molecules to strengthen the intermolecular association, and some platinum drugs, for example, cisplatin, make coordinate bonding with the core-forming segments to apparently fix the core structure. The core-forming segment can be further designed to trigger the release of the payloads in response to the specific signals, such as pH and redox potential, in the body fluids. Thus, the chemical structure of coreforming segments should be carefully designed according to the characteristics of target drug molecules for both their stable encapsulation out of the target site and triggered release in the target site (Figure 5).

3. DESIGN CRITERIA OF BLOCK COPOLYMERS FOR SELF-ASSEMBLY OF POLYMERIC MICELLES Polymeric micelles can be constructed by using both biocompatible synthetic polymers and natural macromolecules, which can combine the processability and adaptability of the former with the ability of the latter to program assemblage mechanisms and control the structure and function.3,53 The segments of the block copolymers control the assembly process, the arrangement of the polymers and the cargo, and the stability of the nanoassemblies, as well as the performance in biological environments.3,53 However, the selection of the polymers should not only consider the structural and functional roles in the final micellar assembly, but also the safety of these segments, which become significant when repeated administration of micelles is needed. Thus, besides being biocompatible and nontoxic, as indicated by the FDA guidelines for biomedical polymers, it is preferable to reduce the amount of polymers in the body after the drug delivery is accomplished to prevent unwanted side effects, such as activation of immune responses.53 Therefore, block copolymers used for constructing risk-free micelles should be designed to be biodegradable, which allows the complete disintegration of the polymers into the forming monomers, and to be safely excreted from the body without accumulation, avoiding any long-term toxicity.3,53 3.1. Shell-Forming Segments

Several hydrophilic and flexible polymers have been used as shell-forming blocks to achieve effective steric stabilization of micelles. PEG, which was originally introduced in the pharmaceutical field with the aim of extending the half-life of 6848

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Table 1. Hydrophilic Polymers Commonly Used for Constructing the Shell of Polymeric Micelles polymer poly(ethylene glycol)

polyvinyl alcohol

poly(N-vinyl-2-pyrrolidone) poly[N-(2-hydroxypropyl) methacrylamide] poly(oxazolines)

zwitterionic polymers

polysaccharides

poly(amino acid)s

advantages

disadvantages

clinically approved reduced immunogenicity reduced opsonization precise control of end-groups for chemical modification availability in a wide range of narrowly distributed molecular weights clinically approved biocompatible reduced immunogenicity clinically approved (plasma expander) biocompatible clinically approved (plasma expander) biocompatible easy modification of pendant groups biocompatible reduced immunogenicity precise control of structure and chemistry temperature responsive biocompatible reduced immunogenicity precise control of structure and chemistry clinically approved biodegradable reduced immunogenicity nontoxic reduced opsonization (dextran, heparan sulfate) work as ligand molecules (chitosan, hyaluronic acid) biodegradable antifouling properties

refs

nonbiodegradable allergic reaction in sensitive individuals

10, 54, 55

nonbiodegradable in physiological conditions

69

potential immunogenicity nonbiodegradable nonbiodegradable

68, 69

not clinically approved

55, 72

71

not clinically approved 75−77, 93−96 long-term effects after systemic administration are not known high variability in molecular weight and structure 78, 97−99 toxic impurities difficulty in control of conjugation site

charged (poly(glutamic acid), poly(aspartic acid)) 64, 78, 101−106 potentially immunogenic

tion of the corresponding monomers. PVP has been used as a plasma expander, that is, for providing volume to the circulatory system, due to its high water solubility and low toxicity.67 However, PVP lack biodegradability in physiological conditions, and various cases of allergic reactions to PVP have been reported.85 PVA also shows remarkably low toxicity, and it is used in several biomedical applications; particularly, it has been approved as an ophthalmic demulcent.86 While preclinical observations demonstrated that PVA does not have any carcinogenicity and teratogenicity, as well as no impairment in ophthalmological, hematological, clinical chemistry, and motor activity parameters, single or several intravenous injections of PVA in rats showed the development of renal lesions, with subsequent deposition of immunoglobulins and complement activation.87 Poly(N-(2-hydroxypropyl)methacrylamide)’s (PHPMA) biocompatibility, nonimmunogenicity, and possibility for functionalization promoted its broad application in the pharmaceutical and biomedical field.71 PHPMA was first used as a plasma expander, and then as anticancer drug delivery vehicle, with the HPMA copolymer conjugating the anticancer drug doxorubicin (DOX) being the first polymer−drug conjugate entering clinical trials (PK1).71 PHPMA can be used to form the shell of polymeric micelles, while it can also be modified with hydrophobic groups and be applied as the core-forming segment, in which hydrophobic drugs can be incorporated.71 Moreover, PHPMA can be readily used for preparing block copolymers via RAFT by introducing additional monomers in the polymer solution.88 PHPMA is not degraded in the body, and the effects of long-term exposure of PHPMA in humans

injecting a significantly larger number of PEGylated nanoparticles.63 This assumption may also reasonably explain why polymeric micelles, which are smaller and produce more particles per weight of material than PEGylated liposomes, did not show the ABC phenomena. In view of the above concerns on PEG, several hydrophilic polymers are being evaluated as operative alternatives to PEG for modification of biomolecules, as well as shell-forming segments of polymeric micelles.55,64 Non-PEG approaches typically involve the use of biocompatible hydrophilic polymers, including poly(glycerol) (PG),65,66 poly(N-vinyl-2-pyrrolidone) (PVP),67,68 poly(vinyl alcohol) (PVA),69 poly(acrylamide) (PAAm),70 poly[N-(2-hydroxypropyl) methacrylamide] (PHPMA),71 poly(oxazolines) (POxs),72 poly(acrylic acid),73 poly(malic acid),74 zwitterionic polymers,75−77 polysaccharides,78 and poly(amino acids) (PAAs)64,79 (Table 1). Linear and hyperbranched PGs are nontoxic and nonimmunogenic, although they are nonbiodegradable in vivo.65,66 Low-MW PGs are predominantly excreted by kidneys, while PGs with high MW show accumulation in liver and spleen. Because of the similarity with PEG chemical structure, PGs might show similar issues regarding prolonged exposure in the body. The vinyl polymers PVP and PVA are mainly obtained by radical polymerization in solution, and block copolymers having PVP or PVA segments can be synthesized through various polymerization techniques, including atom transfer radical polymerization (ATRP),80,81 reversible addition− fragmentation chain-transfer polymerization (RAFT),82,83 or organostibine-mediated polymerization,84 by sequential addi6849

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pure PAAs with low degree of polydispersity can be prepared by ring-opening polymerization of N-carboxyanhydrides of amino acids, which can in turn be initiated by macroinitiators, resulting in macroinitiator-PAA block copolymers and eliminating the need for conjugation of large macromolecules.3 Poly(glutamic acid) (P(Glu)),101 poly(aspartic acid) (P(Asp)), 1 0 2 poly(hydroxyethyl- L -glutamine), 1 0 3 poly(hydroxyethyl-L-asparagine),104 and poly(sarcosine)105,106 have been considered as potentially safe candidates for substituting PEG with the possibility to be degraded by endogenous proteinases in vivo to their corresponding amino acids. PAAs have been used for assembling the shell of various nanostructures incorporating bioactive molecules. Nevertheless, PAAs pose particular drawbacks as shell-forming materials, including their charge as for P(Glu) or P(Asp), which can induce nonspecific interactions with the biomolecules that should be incorporated in the core of micelles, or with blood components once injected, with the possibility of immunogenicity. Besides, PAAs are not soluble in some organic solvents, which could complicate synthetic protocols for constructing block copolymers or assembling nanostructures. A head-to-head comparison of PEGylated liposomes and liposomes modified with hydrophilic PEG alternatives, that is, PHPMA, poly(N,N-dimethyl acrylamide), poly(N-acryloyl morpholine), PVP, and poly(2-methyl-2-oxazoline) (PMOx), showed that each polymer increased the half-life of liposomes as compared to unmodified liposomes, although the PEGylated liposomes and the PMOx-modified liposomes presented much longer circulation than the liposomes coated with the other polymers.107 However, a second injection of the PEGylated liposomes and the PMOx-modified liposomes induced ABC phenomena, with a 20-fold drop in the blood circulation halflives and high accumulation in liver and spleen, whereas the blood circulation of the liposomes equipped with the other polymeric coatings remained unaffected.107 These observations support the importance of developing alternatives to PEG that can be readily translated for reformulating nanomedicine platforms. Moreover, elucidating the mechanisms that allow some polymers, as well as some nanostructures, like polymeric micelles, to avoid the ABC phenomena would facilitate the rational design of nanomedicines for producing improved formulations in the future.

remain unknown. In a zebrafish embryo assay, which allows in vivo toxicity screening, PHPMA was demonstrated to be nontoxic, although prolonged incubation (72 h) at high concentrations (0.5 mg/mL) showed signs of developmental defects.89 It is worth nothing that such high exposure conditions will not be reached in vivo. Importantly, backbone-degradable PHPMA copolymers can be synthesized by alternating synthetic and degradable blocks by using bifunctional chain transfer agents containing biodegradable units, such as enzyme cleavable peptides.90 This strategy allows preparing copolymers with molecular weights larger than the threshold for renal excretion for achieving prolonged blood circulation, while it facilitates the elimination of the copolymer by the kidneys after it is degraded into smaller molecular weight species.90 POx’s have demonstrated the ability to reduce nonspecific protein binding, and the safe administration of POx-based polymer therapeutics has been recently demonstrated in nonhuman primates.55,72 POx’s are synthesized via living cationic ring-opening polymerization of 2-oxazolines, and are remarkable candidates for drug delivery because of their highly structural and chemical variability, which can facilitate the design of polymer therapeutics, including polymeric micelles.91 Moreover, block copolymers having POx segments can be readily obtained by using the polymer as a macroinitiator. However, POx’s have not entered clinical trials so far, and poly(2-ethyl-2-oxazoline) (PEtOx) has been only approved as a food additive by the U.S.’s Food and Drug Administration. Moreover, PEtOx was found to be degraded to PEI by proteinase K, a nonhuman enzyme,92 although the biodegradation of POxs in humans remains unkown. Zwitterionic polymers, such as poly(2-methacryloyloxyethyl phosphorylcholine) (PMPC)75 and betaine-based polymers,76,77 have also been considered as a compelling alternative to PEG. The zwitterionic polymers achieve high antifouling properties due to their strong hydration effect, which is attributed to electrostatic interaction,76,77 and various nanoparticle-based systems functionalized with zwitterionic polymers have shown improved stability in the in vitro75,93−95 and in vivo conditions.96 Unfortunately, the long-term effects of these polymers after systemic administration, particularly following repeated and cumulative dosing, remain to be studied. For using biopolymers as shell-forming materials, important properties should be considered, including their availability, purity, stability, toxicity, or immunogenicity. Polysaccharides offer availability and well-established modification schemes, as well as biocompatibility and biodegradability. Some polysaccharides, such as dextran and heparin, can inhibit opsonization and complement activation,97 while other polysaccharides, such as chitosan,98 hyaluronic acid,99 and chondroitin sulfate,100 have shown capability as ligand molecules, increasing cellular uptake on target cells. Nevertheless, for using these polymers for constructing micellar nanostructures, there are still several challenges, such as the high variability of MW and structure, which directly affect the nanostructure of the micelles and their pharmacokinetics/ pharmacodynamics, the contamination with bioactive impurities, and the difficulty to control the conjugating position and introduction rate of core-forming segments or therapeutic molecules. PAAs from natural sources also present large variability of MW and structure, and the potential of impurities. However,

3.2. Core-Forming Segments

The segregation of the core from aqueous milieu, which can be elicited through various interactions, including hydrophobic interaction,108,109 metal complexation,110 hydrogen bonding,111 and electrostatic interaction,49,50 is the essential driving force for the formation of micelles and the incorporation of payloads. Moreover, through strong cohesive forces in the core, stable micellar systems with low CAC can be achieved for retaining their integrity during circulation at extremely diluted conditions upon intravenous injection. The cohesive forces between the polymers forming the core and the cargo have been correlated with lower CACs.3,51 Thus, the stability can be affected by different physicochemical aspects of the coreforming blocks, including their stiffness,3,51 crystallinity,3,51,112 and the prearranged chemical cross-linking between polymer segments in the core,3,51 as well as the affinity between these blocks with the drugs mediated by noncovalent interactions, such as π−π stacking,3,109,113 hydrogen bonding,111 or electrostatic interaction.49,50,114 In this way, the localization 6850

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Table 2. Polymers Commonly Used as Core-Forming Segments of Polymeric Micelles Based on the Payload cargo hydrophobic drugs

polymer

limitations

polyesters

biodegradable clinically approved (poly(D,Llactide)) high loading biodegradable synthesis of narrowly distributed poly(amino acid)s with controlled degree of polymerization could be intricated clinically approved (poly(glutamic acid)) high loading capacity high affinity with the cargo tunable release rate effective endosomal escape nonbiodegradable-toxic

118−127

biodegradable controllable endosomal escape

3, 49, 50, 128, 129, 131−134, 136, 334

poly(ethylene imine) poly(amino acid)s

high CAC low affinity with drug molecules low loading efficiency low affinity with drug molecules risk of burst release

refs

block copolymers (e.g., poloxamers) are commercially available

polyamino acids

negatively charged biomacromolecules

advantages

polyethers

complex synthetic processes

108, 115−117

3, 37, 109, 113

148, 152, 153

improve the stability and decrease the CAC of micelles, may also contribute to the segregation of the drugs in the core as the strong interpolymer interactions overwhelm the polymer− drug interactions.124−126 Thus, even though polyester-based micelles are more stable than polyether micelles, they frequently show an initial burst release of the incorporated drug.126,127 PAAs are advantageous backbones for constructing biocompatible and biodegradable cores of micelles, as the side groups of PAAs can be readily modified for appropriate assembly, loading, and release of the payload.3 PEG-b-PAA copolymers forming polymeric micelles can be prepared by ring-opening polymerization of the N-carboxyanhydride of amino acids, which allows precise control of the degree of polymerization (DP) and MW distribution without racemization at the chiral centers.3 Commonly used PAA blocks include P(Asp), P(Glu), poly(L-lysine) (P(Lys)), and poly(histidine) (P(His)) (Table 2). These polypeptide blocks usually consist of monomer units with identical configuration, which may allow the adoption of secondary structures under specific conditions, regulating the structure of polymeric micelles in a comparable hierarchical way to proteins. Assemblies built from secondary structures may in turn modulate the physicochemical properties of micelles, such as their association number, CAC, size, and stability, which may affect their pharmaceutical properties. Moreover, hydrogen bonding between the PAA segments in the core can further increase the micelle stability, concurrently with improvement of the affinity with the incorporated drugs. In addition, depending on the type of amino acid, the PAA block may bear positive or negative charge at their side groups, which can be exploited for additional stabilization and affinity enhancement with the cargo through ion complexation. Micelles incorporating negatively charged biomacromolecules, such as negatively charged photosensitizers,128 proteins,129 polysaccharides,130 small interfering ribonucleic acid (siRNA),131 antisense oligonucleotides,132 plasmid DNA (pDNA),133 and mRNA (mRNA),134 can be self-assembled through electrostatic interaction by pairing oppositely charged block copolymers, neutralizing the charge and inducing the required amphiphilicity for micelle assembly.49,50 Moreover, such micelles should be provided with the ability to release their contents from the endocytotic compartments into the

of the drug in the core structure and the rate of drug release may also determine the micelle stability. For constructing cores containing hydrophobic drugs, the most commonly used polymers are polyethers, polyesters, and PAAs (Table 2), which show different features for assembling the core of micelles. The first report using polyethers as the core segment of drug-loaded polymeric micelles showed the capability of poly(ethylene oxide)-poly(propylene oxide)poly(ethylene oxide) copolymers for solubilizing haloperidol.108 Other polyethers, such as poly(butylene oxide) or poly(styrene oxide), have also been investigated for constructing the core of micelles, incorporating hydrophobic molecules.115,116 Micelles having a polyether core have shown stability superior to those made from low-MW surfactants, although their stability upon systemic injection is low due to their relatively high CAC and limited affinity of the drugs with the polyether segment.117 In addition, polyethers are not degraded in the body, which could pose risks after long-term exposure. Polyesters have shown to be safe for use in humans as biodegradable surgical sutures, tissue engineering scaffolds, and controlled drug delivery systems.118 Particularly, poly(εcaprolactone) (PCL), poly(glycolic acid) (PGA), and poly(D,L-lactic acid) (PDLLA) are among the most widely considered polyesters for building the core of polymeric micelles.119−121 The rate of hydrolysis of these polyesters depends on their MW with more hydrophilic (PGA > PDLLA > PCL) and shorter chains degrading faster.122 Moreover, polyesters forming crystal structures in the core of micelles hydrolyze at a slower rate than amorphous cores.122 Polyester cores have shown increased loading of hydrophobic anticancer agents when compared to polyethers. For example, micelles of PEG-b-PCL copolymers incorporating paclitaxel (PTX) have shown approximately 25% w/w loading efficiency,122 while those from poly(ethylene oxide)-poly(propylene oxide)-poly(ethylene oxide) copolymers loaded a maximum of 1.46% w/w PTX.123 Nevertheless, the affinity between polyesters and drugs is relatively low, and the loaded drugs tend to spontaneously partition in the hydrophobic polyester cores depending on the drug/polymer fractions; for example, the PTX in the core of micelles prepared for PEG-b-PCL copolymers starts to partition at PTX:PCL higher than 5%.124 Besides, the crystallization of the polyester, which can 6851

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riers.41,136,145 These approaches are described in detail in later sections of this Review (sections 5.3 and 5.4).

cytoplasm of cells after they are endocytosed, as prolonged exposure in endosomes and lysosomes presenting nucleases and acidic pH could decrease the activity of the confined micelles. The endosomal escape is a key step for the effective delivery of nucleic acids into cytoplasm or nuclei (see sections 5.3−5.5). For effective endosomal escape, two mechanisms have been proposed, as follows: (i) the proton sponge effect, which is based on the buffering ability of the polymer for boosting the influx of protons and chloride ions during the acidification of endosomes and lysosomes, thus increasing the osmotic pressure and disrupting the endosomal vesicle;135 and (ii) the destabilization of endosomal/lysosomal membranes, which is mediated by electrostatic interaction between the cationic polymers and the anionic phospholipids forming the membrane of the intracellular vesicles, which ultimately disturbs the vesicular structure and allows the leakage into the cytoplasm.136 However, the validity of the proton sponge effect is strongly argued, and critical reports have shown that the buffering capacity of the carriers alone may be insufficient for achieving endosomal escape,137 that the premised proton influx into endosomes/lysosomes may not occur after using carriers with buffering capacity,138 and that the theoretical osmotic pressure achieved during the process may not be sufficient for lysing the endosomes.139 On the other hand, the endosome escape of macromolecules may happen after the formation of pores in the endosomal membrane rather than the proposed endosomal bursting of the proton sponge effect.136,140 Several polycations have been considered as core-forming backbones, including poly(ethylenimine) (PEI),141,142 P(Lys),49,50,128,132,133,143 cationic polyaspartamides (P(Asp(R))),131,134,144,145 polyamidoamine dendrimers (PAMAM),146 and poly(2-(N,N-dimethylamino)ethyl methacrylate),147 with varying abilities for endosomal escape (Table 2). PEIs have become one of the standards for nonviral gene delivery, offering significant transfection and protection of nucleic acids against nuclease degradation.148 At pH 7, PEI is only protonated on every third- or fourth-amine,142 which confers siginificant buffering capacity over a wide pH range, and the possibility to disrupt endosomal membranes. However, the nonbiodegradability of PEIs associated with their high toxicity is a limiting factor for their application in vivo.149−151 Thus, considerable efforts to prepare biodegradable PEIs for reducing the toxicity have been dedicated. In particular, the construction of PEI chains that degrade in the intracellular reductive environment has shown lower toxicity than nonbiodegradable PEIs.152,153 On the other hand, PAAs are readily biodegradable, thus presenting a safer option than PEI. P(Lys) was among the first polymers to be applied in nonviral gene delivery.44,154 P(Lys) is biodegradable, which is a great advantage for in vivo applications, although it still presents moderate cytotoxicity. While polyplexes from P(Lys) are internalized by cells as efficiently as PEI-based polyplexes, their transfection efficiencies are much lower than those of the PEIbased polyplexes, which may be attributed to the absence of amino groups with pKa between 5 and 7, thereby limiting the endosomal escape. In this way, the application of biodegradable P(Asp(R)) having side chains of aminoethylene repeats with buffering capacity in the pH range of endosomes and lysosomes has shown a significant increase in the transfection ability as compared to P(Lys), and has received much attention for constructing safe and efficient gene car-

3.3. Stimuli Sensitivity

Polymeric micelles can be designed to respond to particular stimuli for achieving spatiotemporal control of their functions, such as reporting the conditions of their surroundings, releasing their cargo, and exerting therapeutic effects (Figure 6). Such signals can be endogenously present in the body, and

Figure 6. Design of polymeric micelles respoding to various endogenous and exogenous stimuli.

intensified or distinctly expressed in diseased tissues. For example, the microenvironment of tumors presents unique stimuli as compared to healthy tissues for selectively activating the micelles, including acidic interstitial pH between pH 6.5 and pH 7.2 due to the aerobic glycolysis and lactate production, overexpression of particular biomolecules, extracellular reactive oxygen species (ROS), and altered redox potential.33−35,38,155,156 Moreover, intracellular spurs, such as endosomal/lysosomal pH (pH 6.5−4.5), enzymes, adenosine triphosphate (ATP), intracellular ROS, and redox potential, can be further exploited for controlling the action of micelles inside the cells.35−37,157−159 By merging such stimuliresponsive approaches with polymeric micelles having promoted accumulation in tumor tissues, it is possible to develop effective anticancer therapies with high selectivity. In addition, the application of exogenous triggers, such as temperature, 160,161 light irradiation, 161,162 and ultrasound,163−165 to targeted tissues has been used for developing pinpoint therapies based on polymeric micelles. Polymeric micelles can incorporate various functions to sense the subtle changes in the surroundings by fine-tuning their multicomponent and compartmentalized structure. Thus, polymeric micelles designed to respond to pH changes are usually equipped with ionizable groups or pH-cleavable linkages that are neutral or stable at pH 7.4, but rapidly protonate or hydrolyze at mild acidic pH, respectively.35−38,159,166 Such pH-sensitivity can be applied for developing systems reporting intratumoral pH, releasing their cargo intratumorally or intracellularly. Besides the pH variations, the differential enzymatic activity in healthy and diseased tissues can be used for developing systems with high activation selectivity.167 The enzymatic sensitivity can be provided by introducing moieties in the main chain of the building blocks, or in their pendant groups, which are 6852

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triggering the release of drugs from polymeric micelles through mechanical effects, which may destabilize the micelles and improve the intracellular drug delivery.163−165 In addition, due to the versatile structure of polymeric micelles, it is possible to combine both endogenous and exogenous stimuli-responsive strategies within single platforms eliciting cooperative synergistic effects toward therapies with enhanced efficacy.

selectively recognized and degraded by enzymes overexpressed in interstitial or intracellular environments.167 Nevertheless, the core−shell structure of polymeric micelles may limit the access of bulky enzymes to the enzyme-sensitive links within the core, reducing the drug release rate. The distinct distribution of ATP in intracellular and extracellular environments can also be used for triggering the function of micelles. Thus, while the ATP in plasma is less than 1 μM,168 the intracellular ATP concentrations range from 1 to 10 mM,169 providing a stimulus for micelle activation. Polymeric micelles can be equipped for sensing the changes on ATP concentration by, for example, cross-linking the core with boronate esters that can be tailored for ligand exchange with ATP and micelle dissociation at relevant ATP levels.159,166,170,171 Besides, the reductive environments of tumors145 or the elevated glutathione in the cytosol (2−10 mM),172 which is 1000-fold higher than the levels at the extracellular space, can be applied for triggering micelle activation. Reduction-sensitive polymeric micelles can be modified with disulfide bonds for drug conjugation or crosslinking, which are cleaved under tumoral or cytosolic reductive conditions.41,158 The disulfide approach is potentially safer, as disulfide bonds are used in nature for stabilizing protein structures. Diselenide bonds have also been applied for constructing reduction-sensitive systems,173 because they are cleaved more easily than disulfide bonds under reductive conditions due to their lower bond energy.174 Selenium- and sulfide-containing polymers, such as diselenide-bearing poly(urethane), poly(propylene sulfide), and poly(thioether ketal), can be also used as building blocks that respond to ROS by changing their water solubility.175−177 In addition, ROScleavable moieties, like boronic esters or thioketals, have been used for controlling the release from various nanoparticle systems.178,179 The prevalence of ROS not only in pathologies, but also in healthy biological signals, indicates the potential of ROS-responsive polymeric micelles for developing complex systems interacting with biological events beyond targeted diagnosis and therapy. External stimuli are useful for directing the action of polymeric micelles to specific sites, although their application in disseminated diseases might be restricted. In the clinic, external stimuli, such as heating, light irradiation, and ultrasound, are being used as locoregional treatments. Heating is usually applied in cancer therapy as hyperthermia by raising the temperature of tumor tissues up to 42 °C, which leads to cellular damage. Thus, temperature-sensitive polymeric micelles should be designed to be stable at healthy body temperature (37 °C), while promptly collapsing at locally heated tissues. This is usually achieved by taking advantage of blocks that undergo sharp changes in water solubility due to temperature, such as poly(N-isopropylacrylamide) (PNIPAM), with a lower critical solution temperature (LCST) of 32 °C160 that can be tuned within a desired range, that is, around 42 °C, by random copolymerization with other monomers.161 Light is used in the clinic for the local treatment of various superficial malignancies due to the low penetration depth of light into tissues, particularly in combination with photosensitizer agents. Light-responsive polymeric micelles rely on the light-mediated structural changes of polymer-conjugated chromophores, such as azobenzene, cinnamoyl, nitrobenzyl, pyrene, or spirobenzopyran groups, which lead to the disassembly of the micelles and the release of their cargo upon illumination.161,162 As compared to light, ultrasound can reach deeper tissues for

3.4. Modification of the Surface of Polymeric Micelles with Ligand Molecules

The hydrophilic protective shell of polymeric micelles strongly reduces the contact with blood components, allowing the extension of the half-life of nanomedicines in the bloodstream. However, the shell concurrently limits the internalization of such nanomedicines into the target cells. The poor cell uptake is a crucial problem for applications where the payload loaded in the micelles should be delivered to particular intracellular targets. This effect has been termed “PEG dilemma”, as most reports associate this phenomenon with the PEGylated surface of nanomedicines,180 although it is also observed for other hydrophilic shell-forming segments.181 Advantageously, to overcome this problem, the surface of nanomedicines, including polymeric micelles, can be readily modified with ligand molecules for specific interaction and increased internalization in particular cell populations (Figure 5).45 Moreover, by providing these nanomedicines with the ability for selectively delivering the drugs to their subcellular targets, significant enhancements in therapeutic efficacies could be attained. The efficiency of ligand-installed micelles depends on design aspects, such as the density of ligands on the surface of micelles31,182 and the flexibility of the spacer between the ligand and the micelles, which should provide high mobility to the ligands for better accessibility to the receptor.183 Moreover, the characteristics of the selected ligand−receptor system include the binding affinities,30,154,184 which may be enhanced due to the multivalent effect of several ligands on the carriers,185−188 the receptor internalization,189−191 the biodistribution,192,193 and availability of the receptors.192,193 The variable expression of the receptors194 also determines the efficiency of the targeting ligand strategy. For targeting tumor cells, the aspects involving the binding affinity between ligands and receptors, the density of receptors on the cancer cells, and the internalization rate may in turn affect the penetration and overall accumulation of ligand-installed nanomedicines,195−197 with high binding affinities and receptor densities, and low internalization rates leading to unfavorable penetration profiles. Besides, the intratumoral distribution of the receptors, and the variation of the receptor expression with the stage of the tumors, may also affect the targeting efficiency of ligandinstalled nanomedicines,198−201 and ligand alternatives capable of achieving effective targeting of cancer cells require suitable recognition of the versatile receptor landscapes of tumors. For example, the folate receptor has been considered as a suitable marker for ligand-mediated targeting of tumors, as its expression increases as the tumor stage aggravates, suggesting improved targeting efficiency in advanced stages of cancer.199−201 In fact, the tumor selectivity of folate has been demonstrated in humans, with folate−drug conjugates reaching phase III clinical trials.202 Unfortunately, the therapeutic benefits of the folate−drug conjugate were not sufficient to warrant approval.203 Installing folate on nanomedicines may overcome the limitations of the small molecule conjugates, 6853

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medicines can be readily optimized in murine tumor models by directly comparing nanomedicines with different ligand densities, such comparisons are not plausible in human clinical trials as currently outlined. This is particularly important as the results obtained in animal models may not be readily extrapolated to human settings due to cross-species differences in the expression of receptors or in the binding affinities.217,218 Ligand-installed micelles also offer the possibility for developing selective therapeutic strategies to other disorders besides cancer. Accordingly, several sugar-installed polymeric micelles, including galactose, lactose, and mannose, have been reported for delivering bioactive agents, particularly to immune cells,188,219−221 which could be effectively used for immunotherapies and gene therapy.222,223

such as fast clearance and off-target side effects, while providing higher affinity through multivalent binding. Various ligands, including antibodies,204 antibody fragments,205,206 aptamers,207 peptides,31,208 and small molecules,209 installed on nanomedicines have been confirmed to improve the targeting efficiency to specific cell populations. Cancer-specific antibodies are an effective type of tumordirected ligands for modification of nanomedicines due to their high specificity and affinity for the targeted antigens.204 Polymeric micelles installed with antibodies and antibody fragments have demonstrated enhanced intracellular delivery of cytotoxic agents, improving the efficacy against several tumor models, although the installation of antibody fragments may be more suitable for modification of micelles, as the final size of the constructs will be less affected.204 As compared to antibody−drug conjugates, which can only load 3−4 drug molecules per antibody,210 antibody- or antibody fragmentinstalled polymeric micelles represent a substantial improvement on the deliverable payload, as hundreds of therapeutic agents can be delivered by a single micelle. Despite the clear enhancements of cell recognition, uptake, and intracellular delivery, the mechanisms of tumor accumulation of systemically administered nanomedicines having ligands for targeting specific epitopes in cancerous tissues are still based on the EPR effect (Figure 2B), and the process is controlled by the blood circulation, extravasation, penetration, and retention of the nanomedicines within tumors. Indeed, several reports have indicated that the presence of tumordirected ligands on nanomedicines did not necessarily enhance the drug levels in tumors,192,211−213 suggesting that the installation of ligands does not always correlate with effective accumulation. Nevertheless, ligands can also be used for overcoming several biological barriers, including extravasation and tissue penetration in otherwise impermeable tissues (Figure 2D). We have recently demonstrated that polymeric micelles installed with cyclic Arg-Gly-Asp (cRGD) peptides, which can target αvβ3- and αvβ5-integrins overexpressed in neovasculature and cancer cells,214 not only enhanced the cellular uptake of micelles by cancer cells, but also achieved efficient drug delivery in a mouse model of glioblastoma,31,215 which is notorious for its poor permeability due to the presence of the blood−brain tumor barrier.216 These cRGDinstalled micelles rapidly penetrated and accumulated within brain tumor tissues,31,215 whereas the micelles modified with a nontargeting control ligand, that is, cyclic Arg-Ala-Asp (cRAD), showed significantly lower extravasation and accumulation within the same tumors, suggesting an active extravasation pathway for cRGD-installed micelles, most likely transcytosis. It is worth noting that nanomedicines with too strong binding through multivalent attachment may not be appropriate for this type of transport and would be prone to remain bound to the endothelial lining, thus reducing the transport into the interstitium. The modification of the surface of nanomedicines with ligands may affect their pharmacokinetic and biodistribution, and even the toxicity profiles of nanomedicines, as the surface features of the carriers determine the interaction with blood constituents and cells. Thus, the density of the ligand, the adsorption of proteins to the ligand-installed nanomedicines, and the charge of the nanomedicines after installation of the ligands should be precisely controlled for avoiding unfavorable pharmacokinetics and biodistribution. Whereas the pharmacokinetics and biodistribution profiles of ligand-installed nano-

4. POLYMERIC MICELLES AS CARRIERS OF SMALL DRUGS Most polymeric micelles incorporating small hydrophobic molecules have been designed for the delivery of hydrophobic anticancer drugs, which frequently have to be injected with surfactants and organic solvents. Upon systemic administration, such low-MW anticancer agents distribute to the whole body, reducing the effective dose in the targeted tissues and inducing toxicity. Moreover, the rapid clearance of anticancer drugs from the body results in repeated administrations to maintain an effective drug concentration in tumors, which can further potentiate chronic toxicities and even lead to acquired drug resistance. Thus, polymeric micelles are significantly advantageous for stabilizing the drugs in aqueous conditions, protecting these agents within their core from outer environments, stably circulating in the bloodstream, and selectively accumulating in solid tumors, where they can release the loaded drugs in a programmed manner. The drugs can be incorporated into the core of micelles through physical interactions, that is, by taking advantage of the interaction of the drug with the hydrophobic core-forming segment, or through conjugation of the drugs to the core-forming backbone via labile bonds, which can be cleaved at specific conditions to recover the active drug. In the following sections, we will review the particularities of each approach. 4.1. Polymeric Micelles Physically Incorporating Hydrophobic Drugs

Our polymeric micelles loading DOX are the first example of micelles physically incorporating hydrophobic anticancer agents (Figure 7).109,113,224 These micelles are prepared by mixing free DOX and PEG-b-P(Asp) copolymer conjugated with DOX to the carboxylate moieties of the amino acid via amide bonds, which resulted in the physical entrapment of free DOX in the core of micelles via π−π stacking.109,225 These physically entrapped DOX molecules further stabilize the micelles by serving as an agglomerant, thus reducing the CAC and maintaining the micelle structure upon injection. The optimized version of these DOX-loaded micelles226 was the first micelle formulation to proceed into clinical trials with the developing name NK911 (Nippon Kayaku, Co.) in 2001, setting a valuable benchmark for developing drug-loaded polymeric micelles for translation into human application. Moreover, on the basis of the development of the polymeric micelles incorporating DOX, it is noted that engineering the cohesive forces between core-forming segments and loaded drugs is essential for constructing micelles that can tolerate dilution after intravenous injection. Thus, besides the ability of 6854

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systemic and intracellular barriers to exert effective delivery as described in section 3.3. Thus, selective drug release from polymeric micelles could improve drug targeting and enhance the efficacy of the delivered therapeutics. For this, the intrinsic stimuli in the body, which are given by the unique pathways of micelles or pathological features of the malignancies, could serve as specific triggers for the discharge of drugs from polymeric micelles, for example, the endosomal/lysosomal pH (pH 5−6) determined by the endocytosis of micelles or the mild acidic pH of solid tumors and inflammatory tissues (pH 7.0−6.5).33 In this way, polymeric micelles prepared from polyesters present an increase in drug release rate at acidic pH, as polyester hydrolysis is accelerated.230,231 However, such systems do not release their cargo synchronously to the rapid changes in pH. On the other hand, polymeric micelles prepared from polymers that rapidly change polarity in response to pH can result in localized burst releases of their cargo, such as DOX-loaded micelles constructed from PEG-bP(His),37,232 which show a boost in the drug release rate when pH decreases due to the protonation of imidazole moieties in the P(His) block forming the core.

Figure 7. Self-assembled doxorubicin-loaded micelles constructed from poly(ethylene glycol)-b-poly(α,β-aspartic acid(doxorubicin)) and free doxorubicin. Free doxorubicin is physically loaded in the micelle core by π−π stacking with the doxorubicin molecules conjugated to the polymer, thereby serving as an agglomerant and stabilizer of the micelle structure.

4.2. Polymeric Micelles Reversibly Conjugating Hydrophobic Drugs to the Core-Forming Segments

the PAA block for stabilizing the core of micelles through hydrogen bonds with the cargo, as well as with surrounding PAA chains, the pendant groups of the PAA segment can be readily modified with moieties having high affinity with the drugs to promote physical incorporation. This was confirmed by using PEG-b-poly(β-benzyl-L-aspartate) (PEG-b-PBLA) copolymer to prepare DOX-loaded polymeric micelles, where the poor water solubility of the PBLA block and π−π stacking of benzyl residues with DOX lead to micelles with remarkably high stability in blood.227 Small-angle X-ray scattering of micelles prepared from PEG-b-PBLA copolymer loading hydrophobic drugs, that is, a retinoid antagonist (LE540), showed that the drug could be uniformly dissolved within the core of these micelles, indicating the importance for controlling the affinity of drugs with the core-forming segment for stable drug-loading.228 A comparable approach was used to construct polymeric micelles incorporating PTX by carefully tailoring the payload-core affinity through the introduction of 4-phenyl-1butanol in the side groups of the P(Asp) segment.229 These PTX-loaded micelles stably circulated in the bloodstream, enhanced drug accumulation in tumor tissues, and achieved both high efficacy and reduced side effects in preclinical tumor models, which promoted their clinical translation (NK105; Nippon Kayaku, Co.). In these clinical studies, NK105 demonstrated a safe profile and therapeutic improvement, with a 15-fold higher area under the time−concentration curve (AUC) in plasma than that of conventional PTX dosage, indicating the stable blood circulation of NK105.47 Interestingly, this AUC value is also 33-fold larger than that of micelles prepared from PEG-b-PDLLA copolymer physically incorporating PTX in their hydrophobic core (Genexol-PM; Samyang Biopharm), whereas the maximum concentration in plasma of NK105 was much higher than that of Genexol-PM, and the total plasma clearance of NK105 was more than 70-fold slower than Genexol-PM. These observations indicate the importance of controlling the affinity of the drug cargo with the coreforming blocks for avoiding drug leakage and micelle disruption during circulation in blood. NK105 has almost completed Phase III clinical trials.47 Polymeric micelles physically incorporating drugs could also be provided with stimuli-responsive properties for overcoming

The first example of polymeric micelles having hydrophobic drugs conjugated to the core-forming segment consisted of PEG-b-poly(L-lysine) copolymer and the anticancer drug cyclophosphamide to the poly(L-lysine) block.233 Our initial studies on DOX-loaded micelles prepared by PEG-b-poly(α,βaspartatic acid) copolymer conjugated with DOX through amide bonds also indicated that amphiphilic block copolymers conjugating hydrophobic drugs in the core-forming segment can form polymeric micelles in aqueous environment, even without adding free drug molecules.225 Thus, by using labile bonds, polymeric micelles can be programmed to release the drugs upon sensing specific endogenous and/or exogenous stimuli. Such ability permits reducing drug leakage from the micelles while circulating in the body, and enables designing systems capable of spatiotemporal control of the therapeutic action of drugs. 4.2.1. Drug Loading in Polymeric Micelles via Hydrolytically Labile Bonds. In this section, we focus on polymeric micelles conjugating the drugs with linkages that are sensitive to hydrolysis either at neutral pH or at acidic pH, rather than enzymatic hydrolysis. Accordingly, because of their relative sensitivity to hydrolysis, esters have been widely used for developing polymeric micelles with controlled release property. However, the rate of hydrolysis of ester linkages could be excessively slow or fast depending on the type of ester and the particular purpose of the drug. For delivering anticancer drugs, ester bonds can be designed to be gradually hydrolyzed at neutral pH, or to rapidly release the contents upon sensing the acidic pH of tumors or endosomes. For example, polymeric micelles encapsulating the active metabolite of the topoisomerase I inhibitor irinotecan, that is, 7-ethyl10-hydroxy-CPT (SN-38), were prepared by esterification of the phenol group of SN-38 and the carboxylate moieties of the P(Glu) segment of PEG-b-P(Glu) copolymer.234 These SN38-loaded micelles were effectively stabilized in physiological conditions by the extremely low solubility of SN-38 in water and the strong tendency of the drug for π−π stacking.235 The ester bond within the core of these micelles can be cleaved in phosphate buffered saline, gradually releasing approximately 6855

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route may affect the efficiency of reduction-sensitive systems, as some endo/lysosomal pathways have shown differing redox potentials, for example, reductive environment for folatemediated endocytosis pathway243 or oxidizing potential for HER-2-mediated uptake.244,245 Interestingly, further selectivity can be provided to reduction-sensitive systems through combination with methods that allow the escape from endosomal compartments into the thiol-rich cytosolic environment, such as buffering polycations245 or photochemical internalization (PCI).240

57% of the conjugated SN-38 from the micelles after 24 h incubation at 37 °C.234 Although the release is not inhibited at neutral pH, such release profile can reduce the leakage of drug while the micelles circulate in the body. Thus, these micelles have demonstrated safety and antitumor activity in Phase I clinical trials, which have stimulated several ongoing Phase II trials.47 As polymeric micelles are internalized by endocytosis, the conditions within the endocytic vesicles can be used to trigger drug release from the micelles for delivering their cargo intracellularly and enhancing drug efficiency. To exploit this pathway of micelles, pH-sensitive DOX-loaded micelles were constructed by conjugating DOX to PEG-b-P(Asp) via a hydrazone bond.36 These micelles were shown to selectively release the drug at the acidic pH of endosomes, while at extracellular pH, that is, pH 7.4, the release was negligible. The prevention of drug release at pH 7.4 and the selective accumulation of these micelles in tumor tissues increased the maximum tolerated dose in mice 4-fold (40 mg kg−1 injected intravenously 3 times every 4 days) as compared to free DOX (10 mg kg−1, similar schedule). Moreover, the enhanced delivery of DOX with these micelles to tumors significantly suppressed the growth of subcutaneous C26 tumors.236 This system is now being studied in the clinical stage, although DOX has been changed to the 4′-epimer of DOX, epirubicin (EPI), as it has lower cardiotoxicity than DOX, but similar efficacy.237 A Phase I study of these pH-sensitive EPI-loaded micelles (Figure 8; NC-6300; Nanocarrier Co.) has just been completed to confirm the safety, tolerability, and recommended dosage.238

4.3. Self-Assembly of Polymeric Micelles via Polymer−Metal Complexation

Metal complexation has also been used for self-assembling polymeric micelles. Several combinations of metal ions, including naturally available ions, such as iron,246 zinc,247 and copper,248 have been combined with block copolymers having a core-forming segment bearing side groups with the ability to coordinate with metal ions. PAAs represent an attractive core-forming segment for constructing polymeric micelles from metal complexation. Indeed, approximately onehalf of all proteins contain metals, and about one-third of all proteins are proposed to require metals to perform their functions.249 Binding of metal ions to proteins or peptides is based on the interaction between an electron-donating group present on a protein surface and a metal ion presenting one or more accessible coordination sites,249 which can be subsequently reversed by changes in pH or concentration of nucleophiles, such as Cl− ions. The possibility to reversibly complex metals to PAAs and trigger their release upon specific environmental changes can be exploited for assembling polymeric micelles loading metal-containing drugs. Platinum anticancer drugs, cisplatin and oxaliplatin (oxalato(trans-l-1,2-diaminocyclohexane)platinum(II)), are widely used for the treatment of various cancers,250,251 although they are usually associated with severe dose-related side effects and the occurrence of inherent or developed resistance. Liposomes and polymer−drug conjugates were initially considered for delivering these drugs, with cisplatin-loaded liposomes (Lipoplatin) reaching clinical application.252 However, several limitations have been observed during the development of these formulations, including the relative large size of liposomes, loading efficiency, and leakage of drugs during storage and in the bloodstream. Particularly, for polymer−platinum drug conjugates, the systems lose solubility, cross-link, and aggregate at high loading ratios. For the preparation of platinum drug-loaded polymeric micelles, such an increase in hydrophobicity at high platinum drug substitution ratios can be exploited for assembling uniformly distributed nanostructures. Thus, our polymeric micelles incorporating cisplatin253 or DACHPt254 can spontaneously self-assemble in water through polymer−metal complexation of the drugs with the carboxylate moieties in the PAA segment of block copolymers. When these micelles are exposed to physiological concentration of chloride ions, the carboxylate− platinum complex can be dissociated, showing a sustained drug release in physiological saline and a slow disassembly of the micelles into unimers.255 Polymeric micelles incorporating cisplatin were initially constructed from PEG-b-P(Asp).110,256 Although these cisplatin-loaded micelles increased the blood circulation and tumor accumulation of the drug, they were rapidly degraded in the bloodstream, resulting in high accumulation in spleen and liver.110 For improving the

Figure 8. pH-sensitive epirubicin-loaded micelles (NC-6300) prepared from poly(ethylene glycol)-b-poly(aspartate-hydrazide-epirubicin). The block copolymer contains 8−10 epirubicin molecules per chain. The hydrazone bond between the drug and the polymer is stable at extracellular pH (7.4), which avoids drug leakage during blood circulation, but it is readily cleaved at the mild acidic pH of tumors (6.5) or endosomes (5.5), allowing tumor selective drug release. These micelles are currently under clinical evaluation.

4.2.2. Drug Loading in Polymeric Micelles via Redox Labile Bonds. Besides relying on pH changes to cleave acidlabile bonds, selective drug release of polymeric micelles could be attained by either oxidation or reduction. As described in section 3.3, disulfide bonds responding to reductive environments can be readily introduced to the polymeric micelle architectures as moieties for drug loading or cross-linking of the core-forming segments,239−241 or in the main block copolymer chain to induce the cleavage of the block copolymer and micelle disruption.242 Accordingly, micelles bearing disulfide bonds have shown selective drug release and micelle disruption under reductive cytosolic or endo/lysosomal conditions.239−242 It is worth noting that the intracellular 6856

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5. POLYMERIC MICELLES ASSEMBLED VIA ELECTROSTATIC INTERACTIONS: POLYION COMPLEX (PIC) MICELLES As described in section 3.2, charged macromolecules, such as proteins and nucleic acids, can also be encapsulated into micelles through electrostatic interactions (or charge neutralization) with oppositely charged block ionomers. The multivalent complexation between the charged groups on the core-forming segment of the block copolymer and multiple charged moieties on the payloads facilitates the formation of stable PIC micelles capable of enduring physiological salt concentrations even in diluted conditions. This section highlights the chemical approaches for stable encapsulation of charged macromolecules, as well as triggered payload release.

bloodstream stability of these micelles, the block copolymer was changed to PEG-b-P(Glu) (Figure 9).253 The cisplatin-

Figure 9. Cisplatin-loaded polymeric micelles are self-assembled by mixing poly(ethylene glycol)-b-poly(L-glutamic acid) copolymer and cisplatin in water. The poly(L-glutamic acid) block forms α-helical bundles in the core of the micelles after complexation with cisplatin, which stabilizes the micelles under the harsh in vivo conditions.

5.1. PIC Micelles Loaded with Organic Chemicals

PIC micelles have been used for loading organic chemicals having multiple charged groups in their structure, such as adenosine 5′-triphosphate,260 fluorescent probes,261 and photosensitizer molecules (PSs).262 The latter micelles have been effectively applied for photodynamic therapy (PDT), which is based on the selective activation of PSs in a target tissue upon light irradiation at the appropriate wavelength to generate ROS, damaging target cells via apoptosis and necrosis.263 While physical incorporation of hydrophobic PSs into polymeric micelles can be achieved by hydrophobic interaction between PSs and the hydrophobic core-forming segment of the copolymer, 262 most hydrophobic PSs incorporated into micelles easily form aggregates due to π−π interactions,264,265 which severely decrease the generation of singlet oxygen due to self-quenching of the excited states.264,265 To solve this problem with conventional PSs, we developed ionic dendrimeric photosensitizers in which a core of porphyrin (DPor) or phthalocyanine (DPc) is surrounded by negatively charged dendritic wedges, thereby sterically preventing the aggregation of the center dye molecules.114,128,266 The multiple anionic groups on the dendrimer periphery allow their stable incorporation into PIC micelles through the electrostatic interaction with oppositely charged PEG-b-polyelectrolyte copolymers.267−269 Such micelles exhibit long blood-circulation times and tumor selectivity based on the EPR effect, decreasing unfavorable biodistribution to healthy tissues and reducing adverse effects, such as skin photosensitivity.267−269 This enhanced tumor accumulation of dendrimeric PS-loaded polymeric micelles improved therapeutic efficacy without damaging healthy tissues, as demonstrated in various subcutaneous tumors, as well as in orthotopic models of bladder cancer. Moreover, dendrimeric PS-loaded polymeric micelles also showed selective accumulation in lesions of age-related macular degeneration (AMD), which are characterized by a permeable vasculature, providing effective targeted treatment in a rat model of AMD.268 PDT has been recently used for enhancing the endosomal escape of macromolecular compounds or nanocarriers by the photochemical disruption of the endosomal membranes, allowing their delivery into the cytoplasm in a light-inducible manner, that is, PCI.270 This technology has been demonstrated to be useful for the in vitro and in vivo delivery of pDNA,271 siRNA,272 and immunotoxins.273 Moreover, PCI has been used for complementing the intracellular action of camptothecin-loaded polymeric micelles, which release the drug in response to the reductive cytosolic environment, as an

loaded micelles from PEG-b-P(Glu) showed a slower dissociation rate than the micelles prepared from PEG-bP(Asp), with a gradual release of cisplatin, increasing the blood circulation of the micelles, which allowed attaining high distribution in solid tumors and low accumulation of kidney, liver, and spleen.253−256 This high tumor selectivity of the micelles prepared from PEG-b-P(Glu) promoted superior antitumor effects and lower side effects than free cisplatin.253,255 This stability enhancement and superior biological performance of the micelles from PEG-b-P(Glu) were related to the ordered arrangement of poly(L-glutamic acid-cisplatin) blocks into α-helical bundle structures in their core.255 These bundles prevented the undesirable disintegration of the micelles during circulation, allowing a gradual erosion-like process. Cisplatin-loaded micelles are currently under clinical evaluation (NC-6004; Nanocarrier Co.), showing only mild toxicities and enhanced efficacy against various malignancies, and Phase III studies are being performed in patients with advanced pancreatic cancer.47,257,258 DACHPt-loaded micelles are also prepared from PEG-bP(Glu), which procure them with high stability in physiological conditions, extended blood circulation, and high tumor accumulation. Moreover, by following the stability of the micelles in the bloodstream and tumor tissues in real-time through a dual-fluorescent labeling that reported the position of micelles and their kinetic stability, DACHPt-loaded micelles were found to keep their micelle form during blood circulation, extravasate from the blood vessels into the tumor, and dissociate within the cells in tumors, which enhanced the delivery of DACHPt to the nucleus of cells.43 This subcellular drug targeting is a significant benefit of DACHPt-loaded micelles for evading detoxification mechanisms in the cytoplasm and improving the efficacy of the loaded drug, allowing the micelles to overcome oxaliplatin-resistance in an in vivo model of human colon cancer.43 Moreover, DACHPtloaded micelles have demonstrated higher efficacy than oxaliplatin in various tumor models and reduced the side effects, particularly oxaliplatin’s dose limiting neurotoxicity.259 Phase I clinical studies of DACHPt-loaded micelles are underway (NC-4016; Nanocarrier Co.) for treating patients with refractory advanced or metastatic solid tumors or lymphoma.47 6857

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effective approach for light-activated chemotherapy.240 Thus, the PCI-mediated activation of the camptothecin-loaded micelles allowed specific cytotoxic effects only at the illuminated tumor tissues, thereby reducing side effects and improving the therapeutic efficacy.240 It is also worth noting that the light dose necessary for PCI is much lower than that required for PDT, ensuring the application in deeper tissues, as well as low photocytotoxicity necessary for some applications, such as gene delivery. In addition, for applications requiring minimal cytotoxicity, it is necessary that the PCI is achieved by selective photodamage to the endosomal/lysosomal compartments. Thus, PS-loaded polymeric micelles can offer a safe approach for PCI, as they are internalized through the endocytosis and their photodamage is mainly to endosomal/ lysosomal membranes. The application of PCI for achieving light-controlled gene delivery by using polymeric micelles will be described in sections 5.4.2.3 and 5.4.3.

Figure 10. A representative scheme on chemical “charging” of proteins by installing charged moieties through reversible bonds.

original antigen recognition.279 Thus, polymeric micelles loading charge-converted antinuclear pore complex antibodies were engineered to effectively target the nucleus of cancer cells by fine-tuning the charge-conversional rate of the antibodies at endosomal pH, as well as the rate of endosome escape of the micelles, to recover the antibody affinity before reaching the cytosol.280 These polymeric micelles for intracellular delivery of antibodies have potential for developing therapies capable of modulating, inhibiting, or defining the functions of a wide range of target antigens at post-translational level.

5.2. PIC Micelles Loaded with Proteins

Protein drugs, including cytokines, enzymes, and antibodies, are one of the mainstreams in recent pharmaceuticals. In fact, one-half of the top 10 pharmaceutical products in the 2014− 2015 sales were protein drugs, that is, antibodies.274 The primary hurdles in protein therapeutics are the metabolic stability and the immunogenicity. As the target molecules of such proteins are outside of the cells or on the cellular membrane, the modification of these proteins with PEG has been widely considered for enhancing stability, as well as reducing immunogenicity. Indeed, several formulations of PEGylated proteins are clinically available, for example, PEGα-interferon 2a/b for treatment of hepatitis C, and various formulations are under clinical trials.275 Meanwhile, when the target molecule of proteins exists in the cytosol or the nucleus, more functionalities beyond simple PEGylation are required for overcoming the delivery barriers, that is, the cellular membrane, the endosomal membrane, and/or the nuclear membrane (Figure 3). Strongly charged proteins, such as lysozyme, can be readily encapsulated into PIC micelles using PEG-aniomers, such as PEG-b-P(Asp), through significant electrostatic interaction.49 Interestingly, lysozymes encapsulated in PIC micelles show apparently enhanced enzymatic activity against a substrate, pnitrophenyl penta-N-acetyl-β-chitopentaoside, as reflected by the decreased Michaelis−Menten constant, possibly as a result of the corona layer of PIC micelles acting as a reservoir of the substrate.276 On the other hand, many proteins have a relatively low charge density, which may not be enough for stable PIC formation. To overcome the weakly charged nature of proteins, approaches for chemical “charging” have been exploited by installing charged moieties (primary amines or carboxylic acids) to proteins through reversible bonds. For instance, immunoglobulin G (IgG) as a representative antibody has relatively weak charges as indicated by the middling isoelectric point of 5.0−9.5, which is obviously lower than that of lysozyme (10.5−11.0).277 The cationic primary amines of lysine residues in IgG can be modified with citraconic anhydride and converted to anionic sites through maleic acid amide formation (Figure 10). This maleic acid amide is stable at neutral pH, whereas it undergoes hydrolysis at the acidic pH of endosomes (pH ≈ 5).278 The citraconic anhydride-charged IgG can be stably encapsulated into PIC micelles using PEG-catiomers at pH 7.4, while it is selectively released from the PIC micelle at acidic pH for eliciting the

5.3. PIC Micelles Loaded with Oligonucleotides

Oligonucleotides are also interesting molecules for delivery because of their potential for developing versatile therapeutic modalities with high target specificity, which can be generated by the simple yet diverse combination of four varying nucleobases, that is, guanine (G), cytosine (C), adenine (A), and thymine (T) in case of DNA or uracil (U) in case of RNA. As illustrated in Figure 11, the so-called central dogma in

Figure 11. Oligonucleotide therapeutics for targeting and silencing the central dogma of molecular biology. Antigene winds to genomic DNA to make a triple helix for inhibiting the transcription. Antisense complementary makes the double strand with mRNA in the nucleus and/or cytoplasm for inhibiting the translation. siRNA makes a protein complex in the cytoplasm, termed RISC, where its doublestranded RNA structure is processed to single-stranded “guide” RNA for recognizing and degrading the complementary mRNA.

molecular biology can be precisely modulated at each stage by oligonucleotides, for example, single-stranded DNA, termed “antigene” for binding to genomic DNA, and “antisense” for binding mRNA, and double-stranded RNA, termed small interfering RNA (siRNA) for binding mRNA.281 6858

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5.3.1. Biological and Structural Characteristics of Oligonucleotides. Despite the strong potential of oligonucleotide therapeutics, only six pharmaceutical formulations, Fomivirsen (Vitravene in trade name), Mipomersen (Kynamro in trade name), Eteplirsen (Exondys 51 in trade name), Nusinersen (Spinraza in trade name), Pegaptanib (Macugen in trade name), and Defibrotide (Defitelio in trade name), have been approved by the U.S. FDA for the treatment of cytomegalovirus retinitis in immunocompromised patients, familial hypercholesterolemia, Duchenne muscular dystrophy, spinal muscular atrophy, AMD of the retina, and hepatic venoocclusive disease, respectively, while several tens formulations are undergoing clinical trials.281,282 The limited approvals in oligonucleotide therapeutics are mainly due to the poor bioavailability of oligonucleotides. Naturally occurring oligonucleotides are readily degraded by nucleases in biological fluids, particularly in the bloodstream. Moreover, the size of oligonucleotides, typically 5−15 kDa, facilitates their rapid excretion from the kidney.283 Overall, the arrival rate of biologically active (or intact) oligonucleotides to their target sites of action, that is, the cytoplasm or nucleus, is substantially low. Thus, the needs for delivery methodologies have arisen to improve the poor bioavailability of oligonucleotides. Apart from the biological activity, it is interesting to note the unique structure of oligonucleotides. Oligonucleotides generally consist of repeating units of phosphate and a (deoxy)ribose backbone associated with a nucleobase (i.e., nucleotide), having appreciably high negative charge density under physiological conditions derived from these phosphate groups. The negatively charged macromolecular structures generate considerable electrostatic repulsion against the negatively charged cellular membrane, resulting in their poor cellular uptake. Thus, the cellular uptake of nucleic acids can be dramatically improved by canceling their negative charges (or charging to the positive) through PIC formation with catiomers. It should be also noted that double-stranded nucleic acids have a persistent length of approximately 60 nm, which is much longer than that of single-stranded nucleic acids.284 Accordingly, short double-stranded nucleic acids, such as siRNA with a length of around 6 nm,285 are considered as a rigid cylindrical architecture (Figure 12). Recently, the effect of

Figure 13. Comparison in PIC formation between flexible ssRNA and rigid siRNA with PEG-b-P(Lys). (A) ssRNA shows a two-step complexation mode, the unit PIC formation from oppositely charged ionomers, and the micelle (or multimolecular assembly) formation from unit PICs above a CAC. siRNA selectively forms unit PICs with PEG-b-P(Lys) without multimolecular assemblies, presumably due to the rigid structure generating substantially limited entropy gain upon multimolecular assembly. (B) Size of PICs prepared from ssRNA (▲) and siRNA (●) with PEG-b-P(Lys) plotted against ionomer concentration. The hydrodynamic diameter of PICs was determined by fluorescence correlation spectroscopy using fluorescently labeled ssRNA and siRNA.286

prepared from ssRNA and PEG-b-P(Lys) plotted against ionomer concentration displayed three distinct phases: (i) a smaller size-constant region at low concentration derived from the selective formation of unit PICs, (ii) a size-increasing region at intermediate concentration from multimolecular association of unit PICs above a CAC, and (iii) a larger sizeconstant region in higher concentration from the formation of PIC micelles as a major portion (Figure 13B). In contrast, it was found that siRNA significantly hampered such micelle formation (or multimolecular assembly) with PEG-b-P(Lys), as their PIC size plotted against ionomer concentration displayed only one phase of smaller size-constant region in a wide range of concentrations (Figure 13B). This is presumably due to the siRNA rigidity that considerably limited the conformational and/or positional entropy gain accompanied by the phase separation of insoluble PIC core upon multimolecular assemblies.286 Instead of mutimolecular assemblies, siRNA and PEG-b-P(Lys) selectively constructed unit PICs without forming multimolecular assemblies even at a fairly high ionomer concentration reaching the overlapping concentration (C*) of PEG-b-P(Lys) (Figure 13B).286 Thereby, additional attractive forces, such as hydrophobic interaction and hydrogen bonding, may be required for the multimolecular assembly between siRNA and PEG-b-P(Lys). 5.3.2. Design Criteria of Block Copolymers for Oligonucleotide Delivery. The primary bottleneck of nucleic acids for pharmaceutical applications is their susceptibility to enzymatic degradation in biological fluids. When mixed with plasma, naturally occurring nucleic acids are immediately digested within a few minutes. Thus, stabilization strategies of nucleic acids have been long sought since the beginning of oligonucleotide therapeutics, and can be divided into two major approaches. One is the chemical modification of nucleotide backbone. A variety of chemically modified nucleotide backbones, including phosphorothioate, 2′-Omethylated ribose, and “locked (or bridged)” nucleotides, have been synthesized for enhanced stability of nucleic acids, and some of them are used in the clinically approved formulations, as well as the clinically tested ones.287 The details in chemical modification approaches are described elsewhere.287 The other is the encapsulation of nucleic acids

Figure 12. A typical structure of 21mer/21mer siRNA. siRNA is a double-stranded RNA, which has 2 nucleic base overhangs at both 3′ ends.

siRNA rigidity on micelle formation with a block catiomer, PEG-b-P(Lys), was investigated in comparison with a flexible single-stranded RNA (ssRNA) control. A combination of ssRNA (21-mer) and PEG-b-P(Lys) (a DP of P(Lys) segment: ∼40) clearly showed a two-step complexation behavior directed toward micelle formation, as follows: (i) the formation of primary assemblies that are a minimal ionomer pair for charge-neutralization, termed unit (or unimer) PIC, and (ii) the secondary or multimolecular association of unit PICs above a CAC (Figure 13A). Indeed, the size of PICs 6859

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within nanoparticulate formulations, such as PIC micelles.132 In this way, nucleic acid payloads can be sequestered from external milieu, as well as enzymatic attacks, leading to their longevity in biological fluids. Indeed, numerous cationic lipids and polymers have been developed for the preparation of PICs, termed lipoplex and polyplex, respectively, with negatively charged nucleic acids. Particularly, PIC micelles have been considered as one of the most promising platforms for systemic oligonucleotide delivery because of the aforementioned properties suitable for stable circulation in the bloodstream and efficient penetration in the target tissue. Oligonucleotide delivery needs careful design criteria for block copolymers and their assemblies because of the fragility and inefficient cellular uptake efficiency (or poor membrane permeability) of naked oligonucleotides. As illustrated in Figure 3, successful oligonucleotide delivery needs apparently conflicting functions, that is, stable encapsulation of nucleic acids outside of target cells versus selective release of nucleic acids in the cytoplasm or nucleus of target cells, and stealthiness in the bloodstream versus endosome membrane disruptivity for translocation from endosome to cytoplasm. A sophisticated approach for merging these conflicting requirements is the creation of smart PIC micelles that exert the desired function in response to specific biological signals or environments. The following sections describe the prominent strategies for developing such smart PIC micelles directed to successful oligonucleotide delivery. 5.3.2.1. Stable Encapsulation and Selective Release of Nucleic Acid Payloads. PEG-b-catiomers can encapsulate oligonucleotides into PIC micelles through their chargeneutralization, and be utilized as a platform material for further functionalization or chemical modification. As described in the previous section, P(Lys) and PEI are the most widely explored catiomer segments as they have primary amines, which can be readily modified to obtain additional functionality. It should be also noted that PEI contains low pKa amines that can protonate in acidic endosomal compartments for the endosome disruption through increased osmotic pressure and/or direct membrane binding.41,148 The PIC micelles prepared with oligonucleotides and PEG-b-catiomers are generally stable in the absence of competitively charged macromolecules, but more likely dissociate in the presence of charged biomacromolecules, such as anionic proteoglycans and endogenous RNA species, through their counter polyanion exchange. Thus, the stabilization of micellar structures in extracellular milieu is the first step for systemic oligonucleotide delivery. One of the simplest stabilization concepts is increasing the apparent charge number (or MW) of oligonucleotides. More charges per oligonucleotide can generate more ion-pairing sites with PEG-b-catiomers, rendering PIC micelles more resistant against the counter polyanion exchange.288 To increase the apparent negative charge, oligonucleotides can be conjugated with each other or into a polyanion backbone through hydrogen bonding or intracellularly cleavable bonds, such as disulfide bond. For example, “sticky” siRNAs were prepared by extending the length of overhangs to make intermolecular hydrogen bonds, or introducing thiol moieties at the end of siRNA strands to produce intermolecular disulfide bonds (Figure 14).288,289 These sticky siRNAs were converted to a polymerized siRNA for the construction of more stable PICs through more ion-pairing sites. However, the issue on this sticky approach is the uncontrollable polymerization process,

Figure 14. Various approaches for increasing the apparent charge numbers (or MWs) of oligonucleotides.

associated with the difficulty in quality control. In contrast, the grafting of oligonucleotides to a polymer backbone is more manageable, as the grafting sites are limited on the backbone. When thiolated siRNAs were grafted into the polyanion backbone via disulfide bond, the siRNA-grafted polymer significantly stabilized its PICs in serum-containing media, similar to the sticky approach. Notably, the enhanced PIC stability was dramatically reduced after treatment under a cytoplasm-mimicking reductive condition, accelerating the release of siRNA payloads.288 It should be noted that this grafting approach can simultaneously provide additional functions, such as endosome disruptivity, by using functional polymer backbones.290 Indeed, siRNA was grafted to the side chains of an endosome-disrupting polymer through an acidlabile maleic acid amide bond for dual functions of the acidic pH-triggered siRNA release and endosome disruption. The resulting siRNA-grafted polymer, thus, constructed more stable PICs as compared to monomeric siRNA, while inducing acidic pH-responsive destabilization of PICs for enhanced endosomal escape of siRNA.290 Nevertheless, these approaches for increasing the charge number of oligonucleotides need relatively complicated reaction processes to obtain polymerized oligonucleotides or oligonucleotide-grafted polymers. Another stabilization approach is providing additional attractive forces to PIC micelles by installing stabilizing units, such as hydrophobic groups and hydrogen-bonding groups, into catiomer segments. The catiomer modification with hydrophobic groups, for example, cholesterol and long alkyl chains, was demonstrated to appreciably stabilize the micellar core through hydrophobic interaction, similar to the amphiphilic polymeric micelles.291 In this approach, the introduction rate of hydrophobic groups needs to be optimized to take a balance between extracellular micelle stability and intracellular payload release.291 It should be noted that a single cholesterol introduction to the ω-end of cationic P(Asp(R)) derivative segment significantly stabilized the micellar structure, possibly due to a more distinct phase separation between cholesterol-derived hydrophobic phase and catiomer/ oligonucleotide-derived PIC phase.292 This fact motivated us to further develop a triblock copolymer comprising PEG, catiomer, and hydrophobic segments for the construction of a PIC micelle equipped with spatially ordered compartments, which are the hydrophobic core, intermediate PIC layer, and hydrophilic PEG shell (Figure 15).293 The resulting triblock copolymer micelles showed higher stability, as compared to the diblock copolymer micelles derived from the catiomer segment with cholesterol-modified ω-end, under serum-containing media. 6860

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Figure 15. A triblock copolymer micelle featuring the hydrophobic core, intermediate PIC layer, and hydrophilic PEG shell. Reprinted with permission from ref 293. Copyright 2014 Elsevier.

The reversible cross-linking of the micellar core can also be used for stabilizing the systems. To this end, disulfide crosslinking has been one of the most widely explored examples because it can be preferentially cleaved in the reductive cytoplasm containing a high concentration of glutathione in the reduced form.158 Thus, disulfide cross-linked PIC micelles loading siRNA showed a certain level of stability in serumcontaining media and induced a significantly enhanced gene silencing in cultured cells.294 Nevertheless, their blood half-life after systemic administration was not significantly prolonged as compared to non-cross-linked control micelles, possibly due to the leakage of siRNA payloads from the cross-linking network in the harsh bloodstream including abundant charged biomacromolecules. Thus, the cross-linking approach was combined with the aforementioned hydrophobic stabilization using cholesterol-modified siRNA (Chol-siRNA), that is, CholsiRNA-loaded/disulfide cross-linked PIC micelles, which significantly enhanced the blood circulation.295 Meanwhile, these results suggest that direct covalent conjugation of siRNA with catiomer segment via reversible bonds should also stabilize the PIC micelle by suppressing the payload leakage. Disulfide bond is available for this purpose by installing thiol functionality to both siRNA and catiomer segment (Figure 14). On the other hand, a more sophisticated approach is using the natural siRNA structure for covalent conjugation with catiomer segments. To this end, we focused on the cis-diol at 3′-ends of siRNA as it can make a reversible ester bond with tetravalent phenylboronic acid (PBA).159 Accordingly, natural siRNA loaded in these micelles works as a macromolecular cross-linker between PBA-functionalized catiomer segments (Figure 16).159 Body fluids contain several cis-diol compounds, which could be exploited as potential triggers for siRNA release from the PBA-functionalized PIC micelles (PBA-micelles) through ligand exchange reaction. Thus, while PBA-micelles were stable at blood levels of glucose (5 mM) and ATP (0.5 mM), they released siRNA payloads in the presence of intracellular levels of ATP (5 mM), thereby allowing selective cytosolic release of siRNA.159 These results also suggest that the negative charges derived from triphosphate in ATP may be crucial for micelle destabilization and triggering siRNA release from PBA-micelles. 5.3.2.2. Targeted Oligonucleotide Delivery to Specific Cells. Following micelle stabilization, the next step is to selectively deliver the oligonucleotides to specific tissues and cells. As described in the preceding section, ligand molecules directed to specific receptors on the target cellular surface have been installed onto the surface of micelles for improving the site-specific delivery. In oligonucleotide delivery, liver and cancer tissues have been mainly targeted by ligand-installed nanomedicines. For liver targeting, sugar molecules, such as lactose221 and N-acetylgalactosamine,296 have been tested as

Figure 16. ATP-responsive PIC micelles for selective release of nucleic acid payloads. The boronate ester between the polymer and the siRNA is stable in physiological concentrations of ATP, but readily exchanged at intracellular levels of ATP. The triphosphate group in ATP further interferes with the PIC formation leading to the release of siRNA.

the ligands targeting asialoglycoprotein receptors expressed on the hepatocyte surface. It should be noted that asialoglycoprotein receptors are composed of three parts, each of which is bound to a single sugar molecule, in a triangle formation.297 Thus, trimeric sugar ligands were demonstrated to recognize more effectively the hepatocyte surface through trivalent binding between the ligand and receptor.296,298 Similarly, the multivalent binding manner is critical for other ligandmediated active targeting. The aforementioned cRGD peptide ligand was reported to show 1 order of magnitude greater binding constant to αvβ3 integrin receptor in a tetrameric form, as compared to a monomeric control.299 Indeed, cRGDinstalled PIC micelles demonstrated the importance of cRGD ligand density for active siRNA delivery to cancer cells, where significant gene silencing activity was observed for siRNAloaded PIC micelles prepared with cRGD-installed PEG-bcatiomers, but not for control micelles with a 1:1 mixture of cRGD-installed PEG-b-catiomers and nonmodified PEG-bcatiomers.300 This result indicates the existence of a critical ligand density for significant binding to the target cells. 5.3.2.3. Safe Endosomal Escape. After endocytosis by cells, macromolecular drugs, including PIC micelles, are subjected to endosomal transport from early endosome to late endosome, followed by lysosomal digestion. This fact suggests that the endosomal escape functionality should be crucial for successful translocation of oligonucleotides into the cytoplasm. PEI with a MW of 25 kDa is the golden standard catiomer used for endosomal escape of nucleic acid.148 Nevertheless, the significant cytotoxic effect of PEI, which is mainly attributed to the cytoplasmic and mitochondrial membrane damage,301,302 has hampered its pharmaceutical use, generating a demand for less toxic endosome-disrupting catiomers. A major approach is the preparation of biodegradable PEI by conjugating low-MW PEIs (or oligoethylenimines, oligo-EIs) 6861

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via intracellularly cleavable bonds.152,303−306 This approach is based on the fact that the catiomer cytotoxicity is generally reduced as the MW (or cationic charge density/number) decreases, while high-MW PEIs permit more stable PIC formation with nucleic acids through more ion-pairing sites. For example, PEIs with a low MW (oligo-EIs) were conjugated to each other through disulfide bonds to increase their apparent MW.305 In this way, chain-extended oligo-EI showed delivery efficacy comparable to 25 kDa PEI, although its cytotoxicity was appreciably lower. Nevertheless, the control of the MW of chain-extended oligo-EI is difficult, similar to the case of polymerized siRNA, which limits the possibilities for assembling well-defined nanostructures. As briefly discussed in section 3.2, the underlying mechanism of endosomal escape of PEI relies on its buffering capacity, that is, the change in the protonation degree between extracellular neutral pH 7.4 and endosomal acidic pH around 5.5 (Δα7.4−5.5). Larger Δα7.4−5.5 indicates that the catiomers contain more protonatable amines in response to the endosomal acidification. A larger amount of protonated amines in the endosomes increases the positive charge density of catiomers for their stronger binding to the endosomal membrane, allowing the acidic pH-accelerated membrane disruption.136 By using biodegradable polyaspartamides bearing protonatable amines in their pendant groups as the polycation blocks, it was indicated to design safe polymeric micelles with optimal Δα7.4−5.5 for achieving efficient endosomal escape.41 The details for the design of such nontoxic polycation segments for effective endosomal escape, and the undelying mechanisms, are described in section 5.4.2.3 for the cytosolic delivery of pDNA. Even though endosome-disrupting catiomers are successfully synthesized to possess a large Δα7.4−5.5, the PEG palisade may compromise their membrane activity in the micellar formulation. This PEG dilemma can be overcome by rendering the PEG segment detachable from the catiomer segment through acid-labile bonds. For example, pH-detachable micelles loading siRNA were prepared by using 2-propionic3-methylmaleic anhydride as the linker between the PEG and the polycation segments.307 These micelles were stable at pH 7.4, although at pH 6.5, their PEG shielding was removed, promoting cellular uptake and gene silencing. Moreover, siRNA-loaded micelles with hypoxia-sensitive PEG detachment were constructed by using PEG-b-PEI-1,2-dioleoyl-sn-glycero3-phosphoethanolamine copolymer with an azobenzene linker between the PEG and the PEI blocks.308 Disulfide bonds can also be used for achieving effective detachment of PEG shell under intracellular reductive conditions, thereby increasing cytosolic delivery and transfection ability.309 Another solution to the PEG dilemma is the reversible masking of cationic sites in the catiomer segment via the acid-labile bond to generate biologically inert ionomers.310,311 Thus, following a strategy comparable to that previously described for charge conversion of proteins, catiomers, for example, PAsp(DET), can be readily masked by reaction of primary amines with maleic acid anhydrides, for example, carboxylated dimethyl maleic acid (CDM) anhydride, and converted to a zwitterionomer via the maleic acid amide bond (Figure 17A).312 This ionomer is stable at neutral pH, while readily reverted back to the parent catiomer at acidic pH.312 Thus, when this zwitterionomer is installed onto the micellar surface similar to targeting ligand molecules, the obtained PIC micelle can become membrane active preferentially at acidic pH, as illustrated in Figure 17B.

Figure 17. (A) Activation (or demasking) scheme of PAsp(DETCDM) to PAsp(DET) by acid hydrolysis. Only the α-isoform of Asp(DET) unit is represented for simplicity. (B) PAsp(DET-CDM)functionalized PIC micelles can become membrane active preferentially at acidic pH through the PAsp(DET-CDM) demasking.

Indeed, the PAsp(DET-CDM) covalently conjugated on the micellar surface via copper free click chemistry dramatically enhanced the gene silencing activity of siRNA-loaded PIC micelles in cultured cancer cells.312 Of note, PAsp(DETCDM) could be directly conjugated to siRNA by reacting dibenzylcyclooctyne-functionalized CDM moieties with azidefunctionalized siRNA through copper-free click conjugation, as the aforementioned siRNA-grafted endosome-disrupting polymer (Figure 14). In this way, the obtained siRNA-grafted PAsp(DET-CDM) induced acidic pH-selective membrane destabilization and siRNA liberation through the single demasking reaction of CDM moieties.312 5.3.3. Therapeutic Potential of OligonucleotideLoaded Micelles. This section describes the therapeutic potential of oligonucleotide-loaded PIC micelles particularly for RNAi-based cancer therapy, which is one of the representative therapeutic targets in RNAi therapeutics.313 As compared to conventional chemotherapy, the inherent advantage of RNAi-based cancer therapy is the appreciably lower side effects because of the target specificity in RNAi therapeutics. Nevertheless, the significant gene silencing in nontarget genes, so-called off-target effect, cannot be completely avoided so far, and naturally occurring oligonucleotides are a potential immunostimulator despite their negligible chemical toxicity. Thus, cancer-targeted oligonucleotide delivery is an urgent demand for successful RNAi-based cancer therapy. RNAi-based cancer therapy has a certain complexity in design, that is, the sequence design of oligonucleotides according to the target gene (or mRNA) and the binding 6862

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Chemical Reviews

Review

position on the mRNA strand, directed toward higher RNAi activity with lower off-target effects. Vascular endothelial growth factor (VEGF), which is known as a stimulator of angiogenesis, is one of the most widely explored target genes in RNAi-based cancer therapy and has been clinically tested in human cancer patients.314 This is because the VEGF knockdown suppresses the angiogenesis in cancer, which can reduce the vascular density of tumors and disturb the blood supply, allowing the cancer growth inhibition, termed antiangiogenic cancer therapy.315 A great advantage of the antiangiogenic cancer therapy is that siRNA does not directly kill the cancer cells, and thus it should be unnecessary to deliver siRNA to all of the cancer cells. It should be also noted that the delivery target in antiangiogenic cancer therapy may not only be the cancer cells but also the cancer-related endothelial cells when the VEGF receptors are selected as target genes. Indeed, we tested siRNAs targeting VEGF (siVEGF) and its receptor 2 (siVEGFR2) as therapeutic candidates for antiangiogenic cancer therapy against a subcutaneous cervical cancer (HeLa cell) model using cRGD-installed PIC micelles.300 The obtained results clearly showed the significant anticancer activity of the targeted PIC micelles by delivering siVEGF and siVEGFR2. In addition, the critical role of the targeting ligand is also evident, as nontargeted control micelles exerted no therapeutic activity. The delivery efficacy of cRGD-installed PIC micelles was further validated for subcutaneous cervical cancer (SiHa cell)316 and lung cancer (A549 cell) models.292 The targeted micelles showed the significant anticancer activity for both cancer models, demonstrating the strong potential for RNAibased cancer therapy. In the case of treatment of the SiHa model, the siRNA targeting human papillomavirus (HPV)derived E6/E7 oncogene (siE6/E7) was selected as a therapeutic candidate because more than 70% of cervical cancer patient cases are caused by HPV infection, that is, the expression of E6/E7 oncogene.317 The E6/E7 knockdown can induce the apoptosis of cervical cancer cells by recovering the expression of p53 as a tumor suppressor gene. Of note, the therapeutic strategy silencing such “exogenous” oncogenes is highly promising from the safety standpoint because normal cells do not express exogeneous oncogenes. Thus, the risk for “on-target” side effects in normal cells should be negligible. In the case of treatment of the A549 tumors, the siRNA targeting human polo-like kinase 1 (siPLK1) was used as a therapeutic siRNA. PLK1 is a cell cycle regulator, and the PLK1 knockdown can trigger the apoptosis of cancer cells. Moreover, siRNA targeting of PLK1 has been clinically tested in human cancer patients.318 The systemic delivery of siPLK1 by the targeted PIC micelles also elicited a significant anticancer effect in a lung cancer model, comparable to the siE6/E7 delivery to the cervical cancer model. These results suggest that the targeted PIC micelles may deliver siRNA to a wide range of tumors. More recently, cRGD-installed PIC micelles were further tested for antisense oligonucleotide (ASO) delivery.319 As illustrated in Figure 13B, the lower CAC of single-strand RNA PICs for micellization (or secondary association) suggests the stable micelle formation with ASO as compared to doublestrand RNA (or siRNA). This enhanced stability of ASOloaded micelles, as well as their relatively small size (