Recent advances in the design, development, and targeting mechanisms of polymeric micelles for delivery of siRNA in cancer therapy

Progress in Polymer Science 64 (2017) 154–181

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Progress in Polymer Science journal homepage: www.elsevier.com/locate/ppolysci

Recent advances in the design, development, and targeting mechanisms of polymeric micelles for delivery of siRNA in cancer therapy Muhammad Wahab Amjad a,1 , Prashant Kesharwani b,c,1 , Mohd Cairul Iqbal Mohd Amin a,∗∗ , Arun K. Iyer c,d,∗ a Centre for Drug Delivery Research, Faculty of Pharmacy, Universiti Kebangsaan Malaysia, Jalan Raja Muda Abdul Aziz, 50300 Kuala Lumpur, Malaysia b Department of Pharmaceutical Technology, School of Pharmacy, The International Medical University, Jalan Jalil Perkasa 19, Kuala Lumpur 57000, Malaysia c Use-inspired Biomaterials & Integrated Nano Delivery (U-BiND) Systems Laboratory, Department of Pharmaceutical Sciences, Eugene Applebaum College of Pharmacy and Health Sciences, Wayne State University, Detroit, MI 48201, USA d Molecular Therapeutics Program, Barbara Ann Karmanos Cancer Institute, Wayne State University, School of Medicine, Detroit, MI 48201, USA

Abbreviations: AFM, atomic force microscopy; AMD, age-related macular degeneration; ANG, Angiopep-2; apoB, apolipoprotein B; ATP, adenosine triphosphate; AURKA, Aurora A kinase; BMA, butylmethacrylate; BPEI, branched PEI; BSA, bovine serum albumin; CCP, charge-conversional polymer; Chol, cholesteryl; Chol-siRNA, cholesterol-modified siRNA; CMC, critical micelle concentration; cRGD, cyclo-Arg-Gly-Asp; DA, 1-octadecylamine; DMAEMA, dimethylaminoethyl methacrylate; DN, dimethoxy nitrobenzyl; DOX, doxorubicin; DP, N,N-dimethyldipropylenetriamine; DTX, docetaxel; ECM, extracellular matrix; EGFR, epidermal growth factor receptor; EPR, enhanced permeability and retention; FITC, fluorescein isothiocyanate; Gal-MNP, N-acetylgalactosamine functionalized mixed micellar nanoparticles; Gal-PEG, galactosylated PEG; HDP, low molecular hyaluronic acid-1octadecylamine-spermine; HLP, low molecular hyaluronic acid-1-laurylamine-spermine; HOP, low molecular hyaluronic acid-1-octanamine-spermine; HSC, hepatic stellate cells; HUVEC, human umbilical vein endothelial cell; 2IT, 2-iminothiolane; IV, intravenous; LA, 1-laurylamine; LCST, lower critical solution temperature; LHRH, luteinizing hormone-releasing hormone; LMHA, low molecular weight hyaluronic acid; mAb-SA, streptavidinconjugated monoclonal antibody mRNA messenger ribonucleic acid; MAL-PEG-NHS, alpha-maleimide-omega-N-hydroxysuccinimide ester polyethylene glycol; MePEG-b-PVL, methoxy PEG-b-poly(␦-valerolactone); M6P, mannose 6-phosphate; MPEG, methoxy poly(ethylene glycol)MPS mononuclear phagocyte system; mPEG-b-PCL-b-PPEEA, monomethoxy PEG-b-PCL-b-poly(2-aminoethyl ethylene phosphate); MTT, (3-(4,5-dimethylthiazol-2-yl)2,5-diphenyltetrazolium bromide); NB, nanobubble; NGR, Asn-Gly-Arg; NIR, near-infrared; NSC-PLL-PA, N-succinyl chitosan-poly-l-lysine-palmitic acid; OA, 1-octanamine; ODN, oligonucleotide; PAA, polyacrylic acid; pDNA, plasmid deoxyribonucleic acid; PEO, poly(ethylene oxide); PEG, poly(ethylene glycol) (PEG); PPO, poly(propylene oxide); Pgp, P-glycoprotein; PAA, poly(l-amino acid); PCL, poly(␧-caprolactone); PDLLA, poly(d,llactide); PGA, poly(glycolide); PIC, polyion complex; PEI, polyethylenimine; PHDCA, PEGylated cyanoacrylate-co-n-hexadecyl cyanoacrylate; PLL, poly(l-lysine); PPA, polyphosphoramidate; PLys, poly(l-lysine); PAsp(DET), poly{N-[N-(2-aminoethyl)-2-aminoethyl]aspartamide}; PDMA-b-PDPA, poly(2-(dimethylamino)ethyl methacrylate)-block-poly(2-(diisopropylamino)ethyl methacrylate); PEC, polyelectrolyte complex; polyHPMA-co-PDSMA, poly[N-(2-hydroxypropyl)methacrylamide-co-N-(2-(pyridin-2-yldisulfanyl)ethyl)methacrylamide); PEI-C-AuNPs, polyethyleneimine (PEI)-coated gold nanoparticles; PLG*LAG, matrix metalloproteinase 2 (MMP-2)-degradable peptide; PAsp, polyaspartamide derivative; PPEEA, poly(2-aminoethyl ethylene phosphate); PECbD, mPEG-PCL-b-PDMAEMA; PECgD, mPEG-PCL-g-PDMAEMA; PEC, polyelectrolyte complex; PBS, phosphate buffered saline; PLL, poly(l-leucine); Plk 1, polo-like kinase 1; PIPAAm, poly(N isopropylacrylamide); PSpMA, poly(spiropyran-methacrylate); PPyMA, poly(1-pyrenylmethyl methacrylate); PNBMA, poly(2-nitrobenzylmethyl methacrylate); PDT, photodynamic therapy; R-, arginine; R9, cell penetrating peptide; RAFT, reversible addition fragmentation chain transfer; RES, reticuloendothelial system; RHAMM, receptor for hyaluronan-mediated motility; RISC, RNA-induced silencing complex; RNAi, RNA interference; RVG, rabies virus glycoprotein; ScFvs, single chain-fragmented antibodies; siAC, siRNA-loaded anticeramidase; siDNMTs, DNA methyltransferases 1 and/or 3b siRNA; siRNA, small interfering ribonucleic acid; SP, spermine; SP, spiropyran; SS, disulphide linkage; TEP, tetraethylenepentamine; TNF-␣, tumor necrosis factor alpha; TP, tetraethylenepentamine; TPP, sodium triphosphate; UCST, upper critical solution temperature; VEGF, vascular endothelial growth factor; V-, valine. ∗ Corresponding author at: Use-inspired Biomaterials & Integrated Nano Delivery (U-BiND) Systems Laboratory, Department of Pharmaceutical Sciences, Eugene Applebaum College of Pharmacy and Health Sciences, Wayne State University, Detroit, MI 48201, USA. Fax: +1 313 577 2033. ∗∗ Corresponding author. Fax: +603 2698 3271. E-mail addresses: [email protected] (M.C.I. Mohd Amin), [email protected] (A.K. Iyer). 1 These authors contributed equally to this work. http://dx.doi.org/10.1016/j.progpolymsci.2016.09.008 0079-6700/© 2016 Elsevier Ltd. All rights reserved.

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Article history: Received 1 November 2015 Received in revised form 13 August 2016 Accepted 9 September 2016 Available online 19 September 2016 Keywords: Polymeric micelles siRNA Active targeting Passive targeting Multidrug resistance Proton sponge effect

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a b s t r a c t Small interfering RNA (siRNA) is a relatively novel nucleic acid-based therapy to treat diseases such as cancer. Nevertheless, substantial obstacles to its clinical applications have been reported, such as low cellular uptake, immunogenicity, off-target effects, and instability in physiological environments. The design of appropriate delivery vehicles capable of transporting siRNA to target cells has been pursued. Nanoparticles are extensively studied for the delivery of siRNA. Among the various nanocarriers, polymeric micelles have recently gained strong interest. Polymeric micelles of average nanometer size are straightforward to design and modify. Hydrophilic groups incorporated in the polymeric micelles can extend in vivo half-life of siRNA to ensure adequate accumulation in tumors, be exchanged for cations that electrostatically interact with siRNA, and be coupled to various ligands for cell-specific targeting. The polymeric micelle core provides stability and serves as a loading dock for drugs. In this review, the different types of polymers used, the design and characterization of polymeric micelles for siRNA delivery, and the established polymeric micelle targeting mechanisms are discussed. © 2016 Elsevier Ltd. All rights reserved.

Contents 1. 2. 3. 4. 5. 6. 7. 8. 9.

Introduction . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 155 Challenges facing the delivery of siRNA . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 156 siRNA–drug combinations . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 156 siRNA delivery and activity in cells . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 156 Types of polymers used for fabricating polymeric micelles . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 158 Design of polymeric micelles for siRNA delivery . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 159 Effect of size, charge, and surface functionality on the biodistribution of polymeric micelles . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 159 Fate of polymeric micelles and pharmacokinetics in the body . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 160 Targeting mechanisms of polymeric micelles . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 166 9.1. Passive targeting (the EPR-effect) . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 166 9.2. Active targeting . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 167 9.2.1. Ligand-coupled micelles . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 167 9.2.2. Stimuli-responsive micelles . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 171 9.2.3. Photodynamic therapy . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 173 10. siRNA-based therapy . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 173 10.1. Targeted delivery of siRNA to the liver . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 173 10.2. Targeted delivery of siRNA to tumors . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 174 10.3. Miscellaneous routes for administration of siRNA nanoparticles . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 175 11. Conclusions and future directions . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 175 Acknowledgments . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 176 References . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 176

1. Introduction The mechanism of RNA interference (RNAi) was discovered in plants, while the mechanism of gene silencing was later found in Caenorhabditis elegans. Gene expression may be inhibited by small interfering RNA (siRNA) via the well-controlled enzyme-mediated gene silencing mechanism [1]. RNA–protein interactions are categorized into four main phases: association of the RNA-induced silencing complex (RISC) and siRNA, stimulation of RISC, and recognition and cleavage of the target gene. Intracellularly, siRNA is integrated into the RISC, which separates the RNA duplex strands and removes the sense strand, while activated RISC uses the antisense strand to guide the cleavage of messenger RNA (mRNA) [2,3]. The mRNA cleavage of RNAi is mediated by an endonuclease, Argonaute 2, within the RISC [4,5].

siRNA blocks the expression of target genes in numerous cells. Apart from drug development and biological research, siRNA possesses remarkable therapeutic properties that may be applied to the treatment of, for example, macular degeneration and cancer by inhibiting the overexpression of angiogenic growth factors or oncogenes. The Nobel Prize (medicine) in 2006 reiterated the commitment of global research to RNAi [6]. Ten years after the discovery of RNAi, some therapeutic siRNAs are now undergoing clinical trials [7,8], the majority of which involve local administration of siRNA via the intranasal or intravitreal route, for example Refs. [9,10]. The first therapeutic siRNA, bevasiranib, designed to treat wet neovascular age-related macular degeneration (AMD) by targeting vascular endothelial growth factor (VEGF) [11,12], has reached phase III of clinical testing. Another VEGF-targeting therapeutic siRNA, AGN-745, was the sec-

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ond to reach clinical trials and the initial data on its safety and efficacy have shown promise [12]. Most therapeutic siRNAs require carriers for systemic or intravenous (IV) administration. Tekmira Pharmaceuticals Corporation has used a lipid nano-carrier for the development of RNAi products. TKM-ApoB was developed using siRNA against the mRNA of ApoB to treat hypercholesterolemia. In July 2009, the company started a phase I trial for TKM-ApoB but ended the trial at the beginning of 2010, as the initial results were unsatisfactory. Calando Pharmaceuticals designed siRNA-complexed cyclodextrin-derived nano-carrier (CALAA-01) targeting ribonucleotide reductase (RRM2) [13], which is currently in phase I of clinical testing for the treatment of solid tumors. The delivery of nucleic acids (mostly siRNAs) offers promising prospects for cancer treatment. Following IV administration, siRNA is rapidly degraded by nucleases in the blood stream [14,15]. A solution to this problem is to condense siRNA with a polycationic vector. Since siRNA is 21–23 base pairs (bp) in size and exhibits minor electrostatic interactions with polycations, it contributes little to the total size of the delivery complex. The polymeric micelle-mediated intracellular transport of siRNA may aid in overcoming the cellular barriers encountered by siRNA by facilitating cytoplasmic siRNA release, endosomal escape, and intracellular uptake, for example [3,16,17]. Upon reaching the cytosol, siRNA can exert its inhibitory effects (Fig. 1). There is a need to develop nanocarriers capable of safely and efficiently delivering siRNA to the site(s) of interest. A model delivery vehicle for the in vivo administration of siRNA should be non-immunogenic and non-toxic, efficient in condensing siRNA and sustaining its integrity prior to reaching the target, able to evade rapid clearance to reach the target site, and internalized and disassociated intracellularly to release its siRNA load, thereby exposing the siRNA to the mRNA [18–20]. Therefore, nanomedicine-derived carriers have been extensively explored for siRNA delivery [3,21]. Polymeric micelles have gained increased attention for their role in drug delivery, as they possess several bioand physicochemical benefits over those of other vehicles. These include their high stability, biocompatibility, ability to solubilize poorly soluble agents, and capability to accumulate at poorly vascularized sites [22–26]. This review focuses on the most recently used polymers to develop polymeric micelles as siRNA carriers, their design, characterization, and targeting mechanisms. 2. Challenges facing the delivery of siRNA Successful siRNA delivery is the main challenge in understanding the utility and potential of siRNA. Although the local targets are easily accessible, difficulties may be faced when siRNA is intended to be transported to organs/tissues that are only accessible via systemic siRNA administration [27–29]. The systemic delivery of siRNA is hindered by several obstacles. Naked siRNA is less stable after in vivo IV administration because of rapid degradation by nucleases. Rapid excretion by the kidneys and uptake by the mononuclear phagocyte system (MPS) contribute to siRNA’s short half-life. Furthermore, their relatively

large size (about 13 kDa), negative charge, and hydrophilic nature hamper crossing of the plasma membrane [30–32]. Their stimulation of the immune system and ability to produce off-target effects are some other challenges associated with the use of siRNA [4]. Endogenous miRNAs can compete with siRNAs by saturating the shared RNAi machinery, thus hindering normal miRNA-mediated gene regulation [33–35]. Several strategies have been recommended to address these challenges. Chemical alteration of the bases, sugars, or phosphate linkages of siRNA may lessen immune stimulation and enhance the stability of siRNA. Alterations of the unlocked/locked nucleic acids and of the 2 sugar position may decrease immunostimulatory activity and enhance resistance against endonucleases, while increased stability and resistance against exonucleases may be achieved via the induction of a phosphorothioate bond in the siRNA backbone at the 3 end [4]. Phosphate or sugar linkages are more common than base modifications. The binding potency/specificity and C5-methylation of pyrimidines may be increased by using pseudouracil or 2-thiouracil. Improving the structure and sequence of siRNA and restriction of the amount of exogenous RNA may help avoid the off-target effects of siRNA and the saturation of RNAi machinery [33,36,37]. The use of nanocarriers or the chemical modification of siRNA could help to address the issues related to the rapid clearance, short half-life, and systemic instability of naked siRNA [33,36,38–41]. The electrostatic complexation of siRNA with cationic carriers may mask the anionic charge of siRNA, thereby enhancing cellular internalization [42]. 3. siRNA–drug combinations By virtue of its pathophysiology, tumors form new blood capillaries with the help of angiogenic molecules to meet growth nutrient and oxygen requirements for proliferation. For instance, the inhibition of vascular endothelial growth factor (VEGF) activity leads to the suppression of various tumor-promoting factors, such as the proliferation of endothelial cells, angiogenesis, and tumor progression [43]. siRNA in combination with other chemotherapeutic drugs has been extensively investigated in a number of clinical trials, with appreciable results. It was established from these investigations that such combinations can significantly enhance the sensitivity of the chemotherapeutic agent by suppressing genes involved in the development of chemotherapeutic resistance (Fig. 2). Investigators exploring combinations of VEGF siRNA and anticancer drugs have reported superior antitumor activity compared with administration of either compound alone. Similarly, several studies have reported VEGF siRNA–chemotherapy resistant gene combinations to effectively inhibit angiogenesis [44], hence mandating the exploration of siRNA chemotherapeutic drug combinations. 4. siRNA delivery and activity in cells siRNA is a hydrophilic polyanion that cannot readily cross the lipid bilayers of the cell membrane [45]. In comparison with other nucleic acids, such as plas-

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Fig. 1. Intracellular transport of siRNA by polymeric micelles. Micelles can be internalized via various membrane constituents, such as lipid rafts, caveolae, or clathrin. Endocytosis may be receptor-dependent or independent. In all instances, micelles enter through endosomal vesicles, which fuse with lysosomes to result in the degradation of siRNA. To evade this degradation pathway, micelles need to be released from the endosomal vesicles. Upon reaching the cytosol, siRNA can exert its inhibitory effect.

Fig. 2. Multidrug resistance (MDR). MDR is one of the main challenges faced by chemotherapy and other therapeutic treatments. The target cells become therapy-insensitive after prolonged exposure to permeability glycoprotein (Pgp)-binding therapeutics, as Pgp, a highly efficient drug efflux pump, effluxes a large number of therapeutics out of the cell.

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mid DNA (pDNA), siRNA seems to be transported into cells through some form of endocytosis, and is then sequestered in endosomal/lysosomal vesicles where it may undergo degradation by nucleases [46]. Both endosomes and lysosomes possess acidic interiors (pH 5–6.5) [47], and lysosomes contain hydrolases such as ribonuclease, deoxyribonuclease, acid phosphatases, phosphodiesterases, and pyrophosphatase, which together can lead to degradation of endocytosed siRNA and subsequent prevention of RNAi [48]. Moreover, synthetic siRNA was found to be localized to the perinuclear regions and incapable of entering the nucleus following transfection by liposomes, even after an extended incubation period [49]. In comparison with siRNA introduced directly into the cytosol (bypassing the endocytic pathway), it could move quickly into the nucleus, the additional site of antisense activity other than cytoplasm. Based on this finding, siRNA was administered into mammalian cells either as synthetic or intracellularly expressed siRNA, after introduction of encoding genetic information by plasmid DNA (pDNA) or viral siRNA [46,50]. The difference between these approaches is that only that pDNA requires nuclear rather than cytosolic delivery, as transcription of the encoded DNA construct occurs in the nucleus. Efficient in vivo delivery of siRNA is more difficult to achieve than in vitro delivery, owing to limitations associated with target selectivity and homeostasis [51]. Nonetheless, the cellular membrane remains the primary obstacle for siRNA to be efficiently transported to the target site, even with in vitro delivery. A number of approaches have been established to deliver siRNA into cells: integration of siRNA into cations such as cationic liposomes or cationic polymers, and incorporating siRNA into viral vectors and distorting cell membrane integrity using various physical stimuli, including transfer of siRNA into cells by “gene gun” or magnetofection, or by lowering the cell membrane barrier (e.g., by electroporation or ultrasound) [25,51–56]. However, the use of physical stimuli such as electroporation has been shown to potentially decrease cell viability to less than 60%, despite high cellular uptake. Furthermore, transfection of siRNA into mammalian cells has been demonstrated to be dependent on cellular features such as cell type, confluency, and passage number. Likewise, when using cationic vectors as delivery systems, their compatibility in the growth medium, cytotoxicity, and physical characteristics of the cationic particles also significantly contribute to siRNA transfection efficiency [51,57,58]. For instance, even where cationic particles are sufficiently small enough to be taken up by cells via endocytosis, they should also be capable of escaping from endolysosomal vesicles to allow RNAi to occur. Several available strategies for endosomal escape are discussed in detail in Section 6. 5. Types of polymers used for fabricating polymeric micelles The synthesis of polymeric micelles mainly involves the use of amphiphilic diblock copolymers, although graft and triblock copolymers are also used. These three copolymer types all have specific benefits for the delivery of

drugs, such as prolongation of drug circulation time, the potential to alter the drug-release profile, or the addition of targeting ligands. The outer hydrophilic portion can consist of polyethers including poly(ethylene oxide) (PEO) and poly(ethylene glycol) (PEG). Additional hydrophilic portions comprise polymers including poly(trimethylene carbonate), poly(acryloylmorpholine), and poly(vinylpyrrolidone) [59,60]. Occasionally, polymers including polyelectrolyte and PEO are combined to constitute the hydrophilic part of the polymeric micelles. These polymers provide stealth properties to polymeric micelles, facilitating the escape from the reticuloendothelial system (RES), a vital property for attaining an extended circulation half-life. The length of PEG chains is generally 1–15 kDa; the use of longer PEG chains results in a heavier hydrophilic corona, thus enhancing the in vivo circulation half-life and stealth properties of the polymeric micelles. The properties of core-forming blocks can be improved by using block copolymers such as PEO-poly(l-amino acids) that offer functional groups. For example, the carboxyl functionality of the poly(l-aspartate) block in PEO-poly(l-aspartate) improves the formation of a chemical linkage with the drug [61], whereas poloxamers (Pluronics® , BASF Corporation), polyesters, and poly(l-amino acid) usually constitute the hydrophobic core. Pluronics® are the triblock copolymers of PEOm/2-b and poly(propylene oxide) PPOn-b-PEOm/2, where n and m represent the PPO and PEO repeating unit count, respectively, and b represents the ‘block’. The partitioning of the hydrophobic moieties and CMC in the micelles is affected by the size of the PPO block [62]. An exceptional characteristic of Pluronics® is their ability to constrain efflux via permeability glycoprotein (Pgp), which is associated with multiple drug resistance (MDR), by altering the intracellular ATP levels in endothelial cell monolayers and by increasing membrane fluidization [63]. Indeed, Pluronics® significantly reduce ATP levels in MDR cells [64]. Since Pgp functionality largely depends on sustained ATP levels, ATP depletion may be a mechanism underlying the inhibition of Pgp by Pluronics® . As shown in Fig. 2, MDR is a major limitation to the successful delivery of biomolecules. However, poly(l-glutamate/aspartates) are the most frequently used poly(l-amino acids) (PAAs), while the most frequently used polyesters for drug delivery include poly(␧-caprolactone) (PCL), poly(d,l-lactide) (PDLLA) and poly(glycolide) (PGA). PAAs are non-toxic, economical, and biocompatible with thiol, amino, carboxyl, and hydroxyl functional groups. The alterations in the core of micelles facilitating drug conjugation rely on the versatility of these functional groups. Regular PAAs are linked via peptide bonds between the ␣-carboxylic and ␣-amine groups of l-amino acids, and may either act as hydrophilic or hydrophobic blocks in copolymers that form the micelles. The hydrophilic parts frequently include poly(l-glutamate/lysine/aspartate). A random derivatization of the amino or carboxyl groups of PAAs yields a hydrophobic block, thus rendering them similar to the graft-type copolymer. Polyesters such as poly(␧caprolactone/glycolide/d,l-lactide) (PCL, PGA, and PDLLA) are among the most frequently used polymers for the synthesis of nanoparticles and micelles. The molecular weight

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and polymer crystallinity influence the degree of hydrolysis of the ester bonds, with smaller and more hydrophilic (PGA > PDLLA > PCL) chains degrading more rapidly [65]. Polyesters are often used for the targeted delivery of small hydrophilic molecules and nucleic acids [66]. Hence, polymeric micelles possess the ability to suppress MDR, which is facilitated mainly by monomers. The distribution and temporal control of polymeric micelles may be achieved by making use of the above-mentioned block copolymers. The distribution control affects the release and, hence, delivery of the loaded drug at the target site [67]. The temporal control refers to the ability to activate drug release at a specific time point in the course of the treatment. The modification of the polymeric micelles comprises surface functionalization, the addition of supporting moieties, and cross-linking in the micelle shell/core [68]. 6. Design of polymeric micelles for siRNA delivery Thus far, the design of polymeric micelles for siRNA delivery relies on two key approaches. The first approach includes the direct conjugation of hydrophobic or hydrophilic moieties to the siRNA through non-degradable or degradable bonds, followed by condensation with polycationic ions to form micelles referred to as polyion complex (PIC) micelles. The polyions in PIC micelles generally consist of PAAs including polyethylenimine (PEI) and poly(l-lysine/aspartate) [69]. The second approach involves the complexation of siRNA with polycation containing an amphipathic block copolymer after micellization [18,70]. Nanoparticles, along with polymeric micelles, are internalized in the cells via endocytosis. A major intracellular barrier faced by many siRNA delivery carriers is endosomal degradation, which can be bypassed using so-called endosomal escape, as alluded to in Section 1 and as shown in Fig. 1. Endocytosis is followed by the fusion of siRNA-complexed nano-carriers with the early, mildly acidic endosomes, which transform into late endosomes with a pH of 5.5. Lysosomes (pH 4.8) are the final organelles involved in the endosomal degradation pathway, in which nucleic acids and proteins are hydrolyzed because of the lysosomes’ more acidic environment. Therefore, the endosomal escape of siRNA is necessary to avoid lysosomal degradation, thereby enabling cytosolic siRNA release and the subsequent interaction of siRNA with RNAi components. Polymeric micelles may also be formed using PEI, which can act as a ‘proton sponge’ to rupture endosomes causing subsequent cytosolic siRNA release (Fig. 3). In addition, polymeric micelles may be designed using pH-responsive polymers to release siRNA at the endosomal pH. Lastly, photosensitizers, cellpenetrating peptides, and fusogenic lipids possessing high buffering capacity may also be used to design polymeric micelles [71]. 7. Effect of size, charge, and surface functionality on the biodistribution of polymeric micelles Factors such as particle size and surface composition are important for the biodistribution of polymeric micelles and therefore for attaining therapeutic efficacy [72–75].

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In an in vivo study carried out in polystyrene micelles of different particle sizes (50–500 nm), a higher accumulation of large particle-sized micelles was observed in the liver [76,77]. It was proposed that opsonization due to surface absorption of proteins mediated the hepatic uptake. At 4 ◦ C, an unexpectedly rapid uptake of the 50-nm particle sized polystyrene micelles was observed. Likewise, protein absorption was found to rely on the size of the micelles. A correlation between protein absorption and particle size was observed by studying PEGylated cyanoacrylate-con-hexadecyl cyanoacrylate (PHDCA) micelles of different particle sizes, ranging from <100 nm to >200 nm, which were incubated for 2 h in serum [78]. The absorption of proteins by small-sized (80 nm) micelles was 6%, whereas large-sized micelles (171 and 243 nm) exhibited protein absorption of 23% and 34%, respectively. The determination of micelle uptake in plasma clearance kinetics studies and by murine macrophages confirmed the existence of a relationship between micelle size and protein absorption. Larger micelles exhibited a blood clearance twice as fast as that of smaller micelles. The presence of a larger number of PEG units on the surface of smaller micelles may be a key factor behind this phenomenon. The biodistribution of micelles can be considerably affected by polydispersity and size, as stated by Rijcken et al. [79]. Even after 72 h, the cross-linked micelles in PBS (at pH 7.4 and 37 ◦ C) did not exhibit a considerable increase in polydispersity and size; however, after 10 h, the polydispersity of non-crosslinked micelles increased to 0.5. Slow hepatic uptake was the cause of the relatively long half-life (almost 8 h) of the micelles in systemic circulation. The micelles accumulated in the skin to an extent similar to that of small liposomes (<100 nm). Hence, it may be concluded that PEGylated micelles smaller than 100 nm exhibit decreased surface adsorption of plasma proteins in addition to reduced hepatic filtration. Additionally, these micelles have a high extravasation rate into permeable tissues and an extended blood residence time, indicating the significance of tunable surface composition and particle size for attaining efficient, targeted delivery. The uptake of polymeric micelles by the cells of the phagocytic system may be affected by the surface charge and functional groups of micelles. Indeed, in contrast to microparticles having carboxyl, hydroxyl, and sulfate groups, polystyrene microparticles containing a primary amine on their surface underwent considerably more phagocytosis. Therefore, as compared with negatively charged or neutral micelles, it is well established that cationic micelles have a faster cell uptake rate. Moreover, cationic micelles have a short systemic circulation half-life and high nonspecific internalization rate. A noticeable decrease in the cellular uptake rate was observed for anionic charged micelles. A change toward a more negative zeta potential was observed for bovine serum albumin (BSA)-absorbed micelles. Nonetheless, cellular uptake was not enhanced by BSA absorption. Interestingly, Tyr-Glu (neutral and negatively charged, respectively) PEG/PDLLA micelles showed no significant variance in clearance from the blood in an in vivo mouse biodistribution study [80]. The pharmacokinetic parameters showed that all micelles were dispersed primarily in the extracellular space of the

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Fig. 3. The proton sponge effect of cationic polymers. Polyethylenimine is the best example of a cationic polymer exhibiting the proton sponge effect, which enables the avoidance of endosomal/lysosomal degradation and, hence, the successful cytosolic release of siRNA.

spleen and liver. However, lower circulation times (10 times lower) of anionic Tyr-Glu PEG/PDLLA micelles in the spleen and liver were observed 4 h after injection. This relatively low accumulation may be attributed to electrostatic repulsion and synergic steric effects, which reduce cellular micelle uptake. Additionally, urinary excretion analysis confirmed the slow hepatic uptake of the micelles. Thiolated micelles (∼250 nm, −5 mV) [81] were found to accumulate to a larger extent in tumors than that observed for non-thiolated micelles, and had a short plasma half-life (∼3 h), indicating high tumor accumulation or uptake. Additionally, thiolated micelles displayed improved splenic uptake. The reaction with thiolated molecules or the formation of disulfide bonds in the blood resulted in the relatively high cellular uptake of thiolated micelles. Hence, neutral or anionic charged micelles may exhibit low, non-specific cellular uptake and less plasma protein absorption. Therefore, the surface functionality of long-circulating micelles should be thoroughly considered in the design and development of micelles. 8. Fate of polymeric micelles and pharmacokinetics in the body Most copolymers used in the design of polymeric micelles are biocompatible and biodegradable, resulting in their straightforward clearance from the body within a precise time interval without causing side effects. Nonetheless, it is important that polymeric micelles accumulate at the site of action before being cleared from the systemic circulation and that the release of siRNA at the site of action does not induce non-specific cytotoxicity. Polymeric micelles should be >42–50 kDa in size and have a molecular weight >70,000 g/mol to remain in the systemic circulation. The volume of distribution and rate of elimination of the copolymers from the body affect the plasma concentration of the copolymers. Protein–siRNA or protein–polymer

interactions may occur after micelle administration, altering the siRNA release rate [75,82,83]. The potential entry routes of micelles need to be studied in detail to understand the potential risks related to micelle exposure. As mentioned above, the main routes of entry are facilitated through IV, intradermal, lung, intestinal tract, and peritoneal exposure. IV administration is an important delivery route for micelles. Once administered, micelles accumulate in various organ systems including the thymus, testes, lungs, kidneys, brain, and heart. After IV administration, the micelles adhere to the capillary wall, extravasate into tissues, distribute at vessel junctions, and migrate toward the target (tumor) cells. The micelles may also be inhaled as aerosols into the respiratory tract, or be administered intratracheally, oropharyngeally, or intrapharyngeally. Once inhaled, the electrostatic force of the air carries the micelles from the upper to the lower respiratory tract. Subsequently, the micelles’ migration from pulmonary sites to the systemic circulation is via the lymphatic system. Micelles administered orally are usually absorbed via gut-associated lymphoid tissue (GALT) and via the epithelial cells of Peyer’s patches in the gut’s enterocytes. The distribution and penetration of micelles administered dermally are less significant and are restricted to the areas near the hair follicles and the topmost layers of the stratum corneum. Ophthalmic micelle-mediated delivery is possible via intravitreal administration, and micelle-mediated delivery to the auditory pathway is achieved via auricular administration [82,84]. Successful drug delivery by polymeric micelles inspired the development of engineered block copolymer micelles for delivering nucleic acids. Block copolymers, comprising polycations and PEG, may be used to prepare PIC micelles by mixing siRNA or DNA in an aqueous solution [85]. These micelles prevent the in vivo enzymatic degradation of siRNA and enhance the internalization of siRNA by tumor

M.W. Amjad et al. / Progress in Polymer Science 64 (2017) 154–181

cells. The polycation segments in PIC micelles are usually PAAs such as poly(l-lysine/aspartic acid) [69]. To obtain these polymers, ring-opening polymerization of ␣-amino acid N-carboxyanhydrides is initiated by amino-modified PEG under an anaerobic and highly anhydrous environment. Since siRNA is present in the micelle core, cytosolic siRNA release is perhaps one of the main challenges for drug delivery. Amphipathic copolymers comprising cationic blocks have been designed to deliver siRNA as a substitute of PIC micelles. In contrast to PIC micelles, these copolymer nanocarriers possess unique benefits. siRNA loading after the formation of positively charged nanoparticles avoids the formation of large-sized particles. This procedure is advantageous for the design of monodispersed and size-controllable siRNA-loaded nanoparticles, and shows specific benefits in vivo. Additionally, this technique may be appropriate to produce the large amounts of micelles needed for therapeutic use [86]. Polymeric micelles made from monomethoxy PEG-bPCL-b-poly(2-aminoethyl ethylene phosphate) (mPEG-bPCL-b-PPEEA) triblock copolymers were prepared by Sun et al. [59]. These polymeric micelles were spherical in morphology, and exhibited a CMC and zeta potential of 2.7 × 10−3 mg/mL and 45 mV, respectively. The average diameter of micelles after siRNA binding ranged from 98 to 125 nm, depending on N/P ratio. The N/P ratios used were 1/1 and 5/1, respectively. The micelles were stable because of hydrophobic–hydrophobic interactions in the PCL core. Furthermore, PPEEA functions as the binding site of siRNA, while PEG covers the hydrophobic core to shield the nanoparticles from clearance in the systemic circulation. These micelles permit the loading of siRNA after the formation of nanoparticles without altering the nanoparticle consistency. The siRNA-loaded micelles could efficiently enter and release siRNA inside the target cells, leading to substantial gene knockdown as shown by the targeting and subsequent efficient silencing of green fluorescence protein (GFP) by two siRNAs in HEK293 cells [86]. An MTT assay was used to demonstrate that neither micelles themselves nor siRNA-loaded micelles showed cytotoxicity, even at high concentrations. siRNA-loaded anticeramidase nanoparticles (micelleplex siAC) significantly silenced the acid ceramidase (AC) gene in vitro, and significantly inhibited tumor growth in a BT474 xenograft murine model after intracaudal administration (Fig. 4). Optimal micelleplex gene silencing conditions were measured at an N/P ratio of 10:1 and an siRNA dose of 100 nM, and these conditions were used for further BT474 cell transfection. Additionally, the micelleplex siAC decreased in vivo proliferation and induced apoptosis in breast tumor cells. However, it did not instigate an immune response or exhibit cytotoxicity [86,87]. The most recent studies using polymeric micelles for siRNA delivery are summarized in Table 1. Recently, polycationic PEG block copolymers with modifiable functional hydrophobic polyesters have been developed. For instance, Xiong et al. developed polymeric micelles to deliver siRNA using conventional degradable grafted block copolymers such as PEO-b-PCL with grafted spermine (SP)/N,N-dimethyldipropylenetriamine

161

(DP)/tetraethylenepentamine (TP) (PEO-b-P(CL-g SP/DP/TP)) [60]. Complexes were prepared at a polymer:siRNA mass ratio of 32:1 to ensure complete binding of siRNA by the polymers. Compared with PEI (IC50 , 6.58 mg/mL), PEO-b-P(CL-gSP), PEO-b-P(CL-g-TP), and PEO-b-P(CL-g-DP) showed significant lower cytotoxicity against MDA435/LCC6 cells, with IC50 values of 452, 165, and 130 mg/mL, respectively. Compared with PEO-b-PCL (IC50 > 500 mg/mL), PEO-b-P(CL-g-TP) and PEO-b-P(CL-b-DP) were more cytotoxic, whereas grafting of the endogenous polyamine SP to the PCL block did not increase the cytotoxicity of PEO-b-PCL significantly. These polymeric micelles effectively bound siRNA, prevented degradation of siRNA by nucleases, and effectively delivered MDR-1 siRNA to knockdown the expression of Pgp in resistant MDA435/LCC6 cancer cells [60]. To control the steric effect of the PEO corona, a novel siRNA carrier was also developed by incorporating the cell-penetrating ligand (TAT) and/or integrin ␣v␤3 ligand (RGD4C) onto the PEO shell of PEO-bP(CL-g-SP), thus enhancing the cellular internalization of micelles (Fig. 5). The composition of siRNA-complexed micelles was as follows; NON-micelle/siRNA, I:II:siRNA (8:8:1); RGD-micelle/siRNA, I:II:III:siRNA (8:4:4:1); TATmicelle/siRNA, I:II:IV:siRNA (8:4:4:1); and RGD/TATmicelle/siRNA, I:III:IV:siRNA (8:4:4:1), where I, II, III, and IV were various polymers. No cytotoxicity was reported for the polymeric micelles. In contrast to unmodified micelles, peptide-functionalized (RGD/TAT) micelles exhibited improved cell internalization and an efficient endosomal escape of siRNA in MDA435/LCC6 cells. The transfection of MDR siRNA-loaded, peptide-functionalized (RGD/TAT) micelles silenced Pgp mRNA and decreased its protein levels [86,88]. Another biodegradable triblock copolymer consisting of PDLLA, polyarginine, and monomethoxy PEG (mPEG 2000PDLLA 3000-b-R 15) was newly synthesized and found to self-assemble into cationic micelles [89]. The arginine moiety possessing the ability to penetrate the cell enhanced the transport of the polyplex micelles to the cytosol [90]. At pH 7.4, micelleplexes showed only slight hemolytic activity, which increased to 48.3% ± 9.2% and 24.5% ± 5.3% at pH 5.60 and 6.64, respectively. Endosome disruption and direct endosomal membrane penetration by protonation permitted epidermal growth factor receptor (EGFR)-targeted siRNA micelleplexes to escape endosomes, thus reducing (by 65%) the expression of EGFR in MCF-7 cells, similar to the reduction observed using Lipofectamine 2000 [89]. Various micelleplexes were prepared at different N/P ratios, ranging from 0.5/1 to 150/1 (siRNA concentration was fixed at 1 mM). For most studies, micelleplexes were prepared at an N/P ratio of 50/1 (200 nM siRNA). In vitro cytotoxicity and hemolysis assays demonstrated that polymeric nanomicelles showed greater cell viability and hemocompatibility than polyethyleneimine (PEI) or R15 peptide nanomicelles. Furthermore, no positive activation of the innate immune responses and no significant body weight loss were observed during treatment, suggesting that this polymeric micelle delivery system is non-toxic. Qi et al. synthesized a positively charged copolymer of mPEG-b-PCL-b-poly(l-lysine) (PLL) with a variable length

siRNA against the gene of interest

N/P ratio

GL3

0, 0.2, 0.4, 0.6, 1, 2, and 4 mPEG-b-PCL-b-PLL PEG-b-PPA/TPP

3–5

PEG-PLys-PAsp(DET-DN)

0–4

cRGD/Chol-siRNA

5–50

PAA

2:1

LMHA-OA-SP (HOP), LMHA-DA-SP (HDP), LMHA-LA-SP (HLP) Iminothiolane-modified PEG-b-PLL (PEG-b-PLL)-(2IT)-(cRGD)

1.2, 7.6 1.2, 7.6

2:1 4, 8 2

PEG-b-PLL(MPA) and PEG-b-PLL(N2IM-IM) Lactosylated-PEG–siRNA conjugate (through acid labile ␤-thiopropionate bond) HOP/HLP/HDP micelles PEG-b-PPA/siRNA/TPP PDMA-b-PDPA/siRNA

16

1

GAPD

Chemistry involved in synthesis

Cell line/animal model used

Outcomes

Ref.

Ring-opening polymerization and click reactions Ring-opening polymerization

Human liver cancer Huh-7 Luc cells HeLa and D407 cells

[91]

Sequential ring-opening polymerization, aminolysis reaction Ring-opening polymerization

HeLa-Luc cells

Improved gene silencing and siRNA delivery efficiency Enhanced siRNA transfection and gene knockdown efficiency Successful siRNA delivery

Ring-opening reaction

Huh7 cells

Efficient intracellular siRNA delivery

Ring-opening polymerization

HeLa-luc cells

Ring-opening copolymerization

B16F10-Luc cell line, BALB/c nude mice HuH-7 cells

Enhanced cellular uptake, wider subcellular [144] distribution, and increased gene silencing A 3-fold longer circulation half-life than that of [228] naked siRNA A 100-fold higher gene silencing efficiency [69] compared with that of free conjugates

[93] [225]

HeLa-Luc cells, BALB/c nude mice Significant gene-silencing activity [143] with HeLa-Luc tumor xenograft Caco-2 cells 35% cellular uptake, effective endosomal [226] escape of siRNA, and significant knockdown of the luciferase gene Promising siRNA carriers [102] Carbodiimide-mediated coupling Not available

Michael addition

[227]

Carbodiimide-mediated coupling Not available Not available HeLa and D407 cells Atom transfer radical A549 and A549-Luc cells polymerization

Efficient siRNA carriers Enhanced stability and knockdown efficiency Improved intracellular siRNA delivery

[102] [93] [229]

VEGF siRNA–PEG/PEI PEC

Not available

Enhanced siRNA accumulation in tumors

[109]

16

PEC-siRNA

Not available

PC-3 cells, female nude mice (nu/nu) with PC-3 tumor xenograft Multicellular layers of DLD-1

1–20

siRNA–PEG/KALA

Not available

PC-3 cells

1–30

MPEG–PCL-SS-Tat

S-180 sarcoma cells, male ICR mice

14:1

Cholesteryl peptides

12, 16 1–30

PEG-LHRH-PEI MPEG–PCL-SS-Tat

Fmoc-solid-phase peptide synthesis method, disulfide linking Fmoc-solid-phase peptide synthesis method EDC chemistry Fmoc-solid-phase peptide synthesis

2:1

(DMAEMA-co-BMA-co-PAA)

RAFT polymerization

1, 2, 4, 8

DMAEMA-(DMAEMA-BMA-PAA) RAFT polymerization

DoHH2 lymphoma and transduced HeLa-R cells Human cervical carcinoma and HeLa cells

PC-3, MCF-7, SK-BR-3 and LNCaP cells SKOV-3 and A2780 cells S-180 cells, male ICR mice bearing a S-180 tumor xenograft

Enhanced penetration of multicellular layers of [230] tumors and effective gene silencing N/P ratio-dependent gene silencing. The higher [231] the N/P ratio, the larger the gene silencing and vice versa High anti-tumor effect [232]

Efficient siRNA transfection

[233]

Enhanced silencing of VEGF gene expression Enhanced cellular uptake ability and high anti-tumor effect in vivo

[141] [232]

70% reduction of the gene expression

[151]

Uptake of siRNA in 90% of the cells and a 3-fold [234] increase in the intracellular siRNA

M.W. Amjad et al. / Progress in Polymer Science 64 (2017) 154–181

4 and 8

0–7

VEGF

Polymer combination

162

Table 1 Summary of the recent studies on polymeric micelles used as siRNA carriers.

GFP

Plk1

Poly (HPMA-co-PDSMA)-b-(PAAco-DMAEMA-co-BMA)

RAFT polymerization

HeLa cells

65% protein and 90% mRNA knockdown at 48 h after transfection with slight toxicity

[235]

16

PEI-C-AuNPs

EDC chemistry

Excellent gene silencing effect

[236]

10, 20, 40

siRNA–PLGA/LPEI

Carbodiimide-mediated coupling

Human breast cancer MDA-MB435 cells expressing GFP MDA-MB-435-GFP cells

[237]

200, 500, 750

siRNA-S-S-PE/PEG-PE

Not available

GFP-expressing C166 cells

Not available

Dendritic cells, C57/BL6 eGFP transgenic and Balb/C mice

0–16

Double-hydrophilic block copolymer and poly-l-lysine-based micelles 6PEG-siRNA-Hph1/cl-KALA

Enhanced gene silencing and improved cellular uptake A 50-fold downregulation of the GFP production over that of free siRNA eGFP gene silencing in C57/BL6 eGFP transgenic mice

Disulfide linking

MDA-MB-435-GFP

Enhanced GFP gene silencing efficiency

[240]

1–30

MPEG–PCL–Tat

Ring-opening copolymerization

Male Sprague-Dawley rats

[150]

1–30

MPEG–PCL–Tat

Ring-opening copolymerization

Male Sprague-Dawley rats

Accelerated transport of siRNA along the olfactory and trigeminal nerve pathway Improved brain delivery compared with that of intravenously delivered dextran

3:1

PEG77 –XPLG*LAGr9 X–PCL17

Not available

Enhancement of the inhibition of breast tumor growth

[241]

cRGD-PEG-PAsp(TEP)Chol/siRNA

Ring-opening polymerization

MDA-MB-231, HT-1080 and HEK293 cells, BALB/nu-nu nude mice with an MDA-MB-231 tumor xenograft A549 cells, BALB/c nude mice bearing an A549 tumor xenograft

Significant sequence-specific gene silencing of Plk1

[242]

PEG-CCP)/calcium phosphate (CaP) hybrid micelles

Ring-opening polymerization

HeLa-Luc cells, FVB/NJc1 female and EL1-Luc/EL1-SV40 male mice

[243]

PEG-CCP

Amide bond formation

SKOV3-Luc cells

Reduction in the luciferase-based luminescent signal from tumors in an siRNA sequence-specific manner Significantly higher gene silencing efficiency

Hyaluronic acid-spermine conjugates (HHSCs) siRNA-NBs

Carbodiimide-mediated coupling

GES-1 and SGC-7901 cells

[103]

Ring-opening polymerization

Rat C6 glioma cells, BALB/c nude mice with Rat C6 tumor xenograft A2780, SKOV3, SKOV3-tr and MDA-MB231 cell lines 293T and SKNO cells

Intracellular delivery of HHSC/siRNA complex enabled by HA-receptor-mediated endocytosis Highly effective RNA interference for tumor treatment Superior therapeutic activity

[245]

Efficient knockdown of a fusion oncogene

[246]

Effective silencing of GFP expression

[86]

Successful downregulation of apoB expression Enhanced SHP-1 gene silencing

[155] [247]

Effective intracellular delivery of siRNA to reverse MDR Efficient transfection, inhibition of HIF-1␣ expression, suppression of cell migration, and impaired angiogenesis

[88]

siLuc

COX-2

0.5–10

SIRT2

0, 0.5, 1, 2, 4, and 8

Survivin siGL3

1–12

EGFP, GL3

25, 50, 75, 100

siRNA-S-S-PE incorporated PEG2000 -PE PCL-b-PDMAEMA and PCL-b-PEG mixed micelles mPEG45 -b-PCL100 -b-PPEEA12

Not available

apoB SHP-1

0–10 0–6

Gal-MNP/siRNA PEI1.8 -DA/SHP-1 siRNA

Not available Carbodiimide-mediated coupling

P-gp

16:1

Not available

HIF-1␣

1:1–1:10

RGD/TAT decorated PEO-b-PCL micelles PCL29 -b-PPEEA21 and PCL40 -b-PEG45

Not available Ring-opening polymerization

Not available

40–70% GFP-expressed HEK293 cells NIH3T3 cells, Male BALB/c mice H9C2 cells, male Sprague-Dawley rats MDA435/LCC6 cells PC3 cells, male athymic (nu/nu) mice

[238] [239]

[150]

[244]

[196]

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6-FAM

10:1

[216]

163

164

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Fig. 4. (A) Micelleplex siAC inhibited the tumor growth of BT474 xenografts. (B) Anticeramidase (AC), 5-bromo-2-deoxyuridine (BrdU), terminal deoxynucleotidyl transferase-mediated dUTP-biotin nick-end labeling (TUNEL), and hematoxylin & eosin analyses of tumor tissue after 528 h of treatment with various formulations. (C) The serum level of interleukin-6 (IL-6), tumor-necrosis factor alpha (TNF-␣), interferon gamma (IFN-␥), IFN-␤, and IFN-␣ in mice bearing the BT474 xenograft at 192 h (A) and 528 h (B) after the intravenous (IV) administration of different formulations [223]. Copyright 2011. Reproduced with permission from WILEY-VCH Verlag GmbH & Co. KGaA.

Fig. 5. (A) The formation of TAT and/or RGD4C micellar siRNA complexes. (B) The expression of permeability glycoprotein (Pgp) as determined by flow cytometry. The P-gp-related fluorescence intensity after 2 days of incubation was normalized to that of the untreated controls. (C) Decrease in doxorubicin resistance in MDA435/LCC6-resistant cells after transfection with different multiple drug resistance (MDR) siRNA formulations. The cellular distribution and accumulation of doxorubicin after transfection was assessed by fluorescence microscopy [223]. Copyright 2011. Reproduced with permission from WILEY-VCH Verlag GmbH & Co. KGaA.

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PCL block via click chemistry [91]. In contrast to mPEG-bPLL micelles, the mPEG-b-PCL-b-PLL micelles were smaller because of the core-stabilizing potential of PCL. The average particle size of mPEG-b-PLL micelles was in the range of 130–300 nm, depending on the N/P ratio. The siRNA binding ability of the copolymers was evaluated using agarose gel electrophoresis. The complexes were prepared at N/P molar ratios of 0, 0.2, 0.4, 0.6, 1, 2, and 4. For anion displacement and nuclease resistance assays, complexes at an N/P ratio of 2 were used. For gene silencing and intracellular translocation, cells were transfected with complexes containing GL3-siRNA or control siRNA at various N/P ratios. The final concentration of siRNA in each well was 50 nM. A polyplex of GL3-siRNA with PEI-25 kDa at N/P = 10 was also used as a positive control. All three copolymers were significantly less toxic than PEI-25 kDa. At a concentration of 64 ␮g/mL, cell viabilities with the three copolymers were maintained at over 80%, while that for PEI-25 kDa decreased to 10%. In Huh-7 Luc cells, mPEG-b-PCL-b-PLL micelles (with an N/P ratio of 60 or 75) exhibited 50% gene knockdown at 48 and 72 h after treatment. These findings were similar to those obtained using Lipofectamine 2000 and were superior to those of 25-kDa branched PEI. At any N/P ratio, the mPEG-b-PLL complex showed little gene knockdown [91]. Lin et al. synthesized grafted PEG-PCL-g-poly(2dimethylaminoethyl methacrylate) (PDMAEMA) and block PEG-PCL-b-PDMAEMA copolymers [92]. The PEG-PCL-gPDMAEMA copolymers exhibited a higher zeta potential and higher positive surface charge than linear PEG-PCL-bPDMAEMA copolymers, resulting in improved endosome escape ability and cellular internalization. An N/P ratio of 5:1 was used for different experiments with PECbD1/siRNA complexes. Cytotoxicity of polymer NPs/siRNA complexes was dependent on the N/P ratio. Because of cell membrane damage caused by the high molecular weight and charge density, polymer NPs/siRNA complexes were highly cytotoxicity to cells at high N/P ratios. Overall, t-test results showed that there was no significant difference in cell viability between PECgD NPs/siRNA complexes and PECbD NPs/siRNA complexes at the same DP of DMAEMA. Furthermore, these micelles improved the tumor-targeting capability in vivo [92]. An efficient stabilization technique is important for successful polycationic nanoparticle-mediated delivery of siRNA since physiological conditions reduce the stability of polyelectrolyte complex nanoparticles. To stabilize the siRNA-containing nanoparticles, Nakanishi et al. used sodium triphosphate (TPP) for ionic crosslinking via cocondensation [93]. Co-encapsulation of siRNA and TPP into PEG-b-polyphosphoramidate (PEG-b-PPA) yielded PEG-bPPA/siRNA/TPP ternary nanoparticles [93]. In contrast to PEG-b-PPA/siRNA nanoparticles, the ternary nanoparticles displayed a great consistency with a small average particle size (80–100 nm) and improved stability in serumcontaining medium because of ionic crosslink stabilization between the cationic PPA segments and the negatively charged TPP. PEGb-PPA solutions were incubated with a mixture of siRNA and TPP at N/P ratios of 4 and 8. The TPP-crosslinked nanoparticles exhibited improved gene silencing and transfection efficiency compared with PEG-b-

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PPA/siRNA complexes. Cell viability was unaffected, either in the absence or presence of TPP in the complexes [93]. PEI has been used for the in vitro and in vivo delivery of nucleic acids for several years [94–96]. High-molecular weight PEI is toxic but offers high transfection efficiency, whereas low-molecular weight PEI is non-toxic yet ineffective. Navarro et al. synthesized micelle-like nanoparticles (MNPs) derived from low-molecular weight PEI (1.8 kDa) and phospholipids, with an outer layer of PEG-stabilized liposomes that bypassed enzymatic degradation and enhanced cellular siRNA uptake [97]. The polymer/siRNA ratio was expressed as the N/P ratio and calculated by assuming that 43 g/mol corresponds to each repeating unit of PEI containing one amine, and that 316 g/mol corresponds to each repeating unit of siRNA containing one phosphate. For gel retardation studies, complexes containing 750 ng of siRNA with varying amounts of PLPEI or PEI 1.8 kDa in PBS were used. Different studies were performed by incubating PLPEI and PEI complexes at an N/P ratio of 10. These MNPs had an average particle diameter of 200 nm with a neutral zeta potential, and exhibited gene-silencing and siRNA-delivery capacity with enhanced biocompatibility [97]. The MNPs were further used in MCF-7 cells for silencing Pgp to overcome MDR. MNPs showed no toxicity toward NIH/3T3 and B16F10 cells in vitro over a concentration range of 1–7 ␮g/mL. Moreover, the viability of cells was higher than 80%, even at high concentrations (up to 50–100 ␮g/mL). Drug efflux was successfully inhibited after treatment of the cells with the MDR siRNA-complexed MNPs. The amount of doxorubicin (DOX) in MDR-treated cells was double that of the control, and resulted in a two-fold reduction in cell viability [97]. The cytotoxicity of positively charged polymers may be attributed to their high charge densities; therefore, cationic liposomes, peptides, and low-molecular weight carriers with a reduced positive charge may be clinically safer [98]. Low-molecular weight polycations are developed to deliver siRNA. Ryu et al., for example, assessed valine (V)and arginine (R)-containing amphipathic peptides for their use as novel siRNA carriers. The peptides had a low charge density and contained 1–4 R and six V blocks (R(1–4)V6). The V blocks formed the hydrophobic core whereas the positively charged shell consisted of R in an aqueous solution. In previous studies, the efficiency of plasmid DNA (pDNA) delivery by RV peptides was less than that of PEI (25 kDa), since RV peptides were considerably smaller than pDNA [99]. The molecular weight of RV peptides and siRNA is similar, which may augment the charge density in an aqueous solution during the formation of micelles. Compared with siRNA-PEI (25 kDa), siRNA in complexes formed by RV peptides was more stable; the R3V6 peptide silenced the reporter genes more effectively than the R4V6, R2V6, and R1V6 peptides. R3V6 (at a 1:20 weight ratio) exhibited a maximum siRNA delivery efficiency, and the R3V6/siVEGF complex downregulated the VEGF expression by 35% compared with untreated control. An siRluc/PEI25k complex was prepared at a 1:1 weight ratio, and the formation of the siRNA/RV peptide complex was verified using a gel retardation assay. A fixed amount of siRNA (0.7 ␮g) was mixed with increasing amounts of RV peptides in distilled water. The stabilities of the siRNA/RV peptide and siRNA/PEI25k

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complexes were evaluated through a heparin competition assay. The siRNA/R3V6 peptide and siRNA/PEI25k complexes were formed at 1:20 and 1:1 weight ratios, respectively. A cytotoxicity assay showed that RV peptides did not cause any cytotoxicity. Additionally, the suppression efficacy of the R3V6/siVEGF complex was comparable with that of the siVEGF/PEI (25 kDa) complex [100]. Researchers have also used natural polymers instead of synthetic peptides or polymers. For example, spermine, a tetraamine formed during cellular metabolism, is found in all eukaryotes. Hyaluronic acid, a natural non-sulfated glycosaminoglycan polysaccharide consisting of d-glucuronic acid and N-acetyl-d-glucosamine, is a major component of the extracellular matrix (ECM). Hyaluronic acid receptors, including receptor for hyaluronan-mediated motility (RHAMM) and CD44, are upregulated in tumor cells [101]. Shen et al. developed hydrophobic hyaluronic acid spermine conjugates (HHSCs) for use as siRNA carriers [102]. The polymers efficiently self-assembled into micelles, attached to the siRNA, shielded the siRNA from enzymatic degradation, and released the siRNA effectively at low levels of polyanionic heparin. The cytotoxicity of siRNA/HHSC complexes in SGC-7901 cells was lower than that of siRNA/PEI 25k and Lipofectamine 2000 complexes as shown by MTT assay. The weight ratio of HHSCs to siRNA varied from 0.5 to 10. For the heparin polyanion competition assay, complexes were prepared at a weight ratio (HHSCs:siRNA) of 5:1 to ensure complete binding of siRNA by the conjugates. To evaluate the stability of the complexes in serum conditions, siRNA/HHSC complexes were prepared at different polymer-to-siRNA weight ratios ranging from 4:1 to 32:1. For cell uptake studies, FITC-siRNA/HSCCs complexes and FITC-siRNA/PEI (25k) complexes had a weight ratio of 1:5. For gene silencing, siRNA/PEI 25k, siRNA/HHSC-1, siRNA/HHSC-2, and siRNA/HHSC-3 complexes with a weight ratio 1:5 were used. In contrast to PEI (25 kDa) and Lipofectamine 2000, HHSCs exhibited enhanced transfection efficiency in HA receptor-(CD44) overexpressing SGC-7901 cells as a result of the HA receptor-mediated endocytosis of HHSCs [103]. 9. Targeting mechanisms of polymeric micelles Two main types of tumor targeting routes, passive and active targeting, have been reported (Fig. 6). 9.1. Passive targeting (the EPR-effect) Systemic siRNA delivery can be categorized into active and passive delivery approaches. One of the most promising approaches for delivering the drugs and genes into tumor sites is based on the phenomenon of passive targeting by exploiting the anatomical and pathophysiological abnormalities of tumor vasculature and utilizing the enhanced permeability and retention (EPR) effect [104–106]. Particles of nano-dimension such as dendrimers, polymeric micelles, and liposomes, as well as macromolecules above the renal excretion threshold (typically above 40 kDa) have a tendency to accumulate more in the tumor tissues than normal tissues. This is because tumor blood vessels are poorly aligned, defective, and

highly leaky. The tumor blood vessels thus allow extravasation of plasma components as well as macromolecules and nanoparticles into the tumor interstitum. Furthermore, tumor tissues have poor lymphatic clearance meaning that accumulated nanoparticles cannot be cleared out of the tumor tissues. Defective lymphatic drainage thus allows nanoparticles to reside in the tumor microenvironment for long periods of time, allowing sustained release of the drug/gene cargo into the tumor tissues [75,106]. Using cationic micelle-forming block polymers, negatively charged pDNA and siRNA can be entrapped within the micellar core via electrostatic interaction and form stable complexes (polyplex micelles) under physiological pH [8,107,108]. In a reported study, Kim et al. developed stable siRNA-PEG/PEI polyplex micelles (polyelectrolyte complex) by conjugating VEGF siRNA with PEG for siRNA delivery. The PEC micelles were prepared by adding polyethylenimine (branched PEI) dissolved in PBS solution into the siRNA–PEG conjugate solution (N/P ratio = 16/1). After administration of these micelles to prostate cancer PC-3 xenograft-bearing mice, downregulation of VEGF gene expression was observed at both the transcriptional and protein level in the tumor, leading to reduced tumor microvessel density and potent tumor growth suppression compared with the naked siRNA and siRNA/PEI complex. The enhanced gene silencing by siRNA-PEG/PEI micelles was due to extended circulation time of siRNA and enhanced accumulation in the tumor via the EPR effect [109]. In another report, atelocollagen (a positively charged peptide) was complexed with negatively charged siRNA and these particles were shown to passively accumulate into targeted tumor owing to the EPR effect [8,110]. In another study, Zhang et al. synthesized triblock copolymer (N-succinyl chitosan-poly-l-lysine-palmitic acid, NSC-PLL-PA)-based micelles to co-deliver siRNA-Pglycoprotein (P-gp) and DOX. The micelles exhibited an average particle size of around 170 nm. siRNA strongly bound to the micelles at an N/P ratio of 20/1 as demonstrated by agarose gel electrophoresis. The micelles were more unstable in pH 5.3 medium than in pH 7.4 medium, which supported the in vitro fast release of siRNA and DOX observed in acidic environments. No polymeric toxicity was reported. The in vitro antitumor efficiency of the micelles increased considerably, particularly in HepG2/ADM cells, as a result of the downregulation of P-gp. Additionally, most of the micelles accumulated in the tumor area beyond 24 h post-injection, and the co-delivery system significantly inhibited tumor growth with synergistic effects in vivo [111]. Zheng et al. synthesized triblock copolymer glycol)-b-poly(l-lysine)-bpoly(l-leucine) poly(ethylene (PEGePLLePLLeu)-based polypeptide micelle NPs to systemically co-deliver siRNA-Bcl-2 and docetaxel (DTX). Strong siRNA binding occurred at a 10/1 N/P ratio between the carrier and siRNA as shown using a gel retardation assay. The blank micelle NPs did not significantly affect the viability of MCF-7 cells at concentrations from 0.5 mg/mL to 80 mg/mL, and cell viability was 100.01% when NP concentration was 0.5 mg/mL. The hydrophobic PLLeu core entrapped DTX, while the PLL polypeptide cationic backbone allowed for electrostatic interaction with the

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Fig. 6. The mechanisms underlying drug delivery by micelles. The extravasation of micelles via ineffective lymphatic drainage, the enhanced permeability and retention (EPR) effect, and the increased permeability of the tumor vasculature is the driving force behind passive tissue targeting. Functionalizing the surface of micelles with ligands that enhance cell-specific recognition and binding is referred to as active targeting. The micelles can release the drugs in close vicinity to the target sites, attach to the cell membrane to act as a sustained extracellular drug release depot, or enter the cell.

siRNA. The resultant micelle NP showed remarkably steady and passive targeting and strong biocompatible properties. The micelle complexes containing siRNA-Bcl-2 efficiently silenced the expression of Bcl-2 mRNA and protein. Furthermore, the co-delivery system of siRNA-Bcl-2 and DTX (DTX-siRNA-NPs) noticeably downregulated the anti-apoptotic Bcl-2 gene and improved antitumor activity with a lesser dose of DTX, leading to significant tumor growth inhibition in an MCF-7 xenograft murine model in comparison with siRNA or DTX treatment alone. The findings thus reported well-defined PEG-PLL-PLLeu polypeptide cationic micelles with an excellent synergistic effect between DTX and siRNA-Bcl-2 in combined cancer therapy [112]. Sun et al. demonstrated an inventive ‘two-in-one’ micelleplex method based on micelles of a triblock copolymer poly(ethylene glycol)-b-poly(␧-caprolactone)b-poly(2-aminoethylethylene phosphate) to co-deliver chemotherapeutic drug and siRNA. Effective binding of siRNA took place at an N/P ratio of 5/1. The micelleplex displayed compact and spherical morphology with a mean diameter of 50 nm. The researchers showed that the micelleplex was proficient in concurrently co-delivering paclitaxel and siRNA to the same tumor cells in vitro and in vivo. They also demonstrated that IV administra-

tion of a micelleplex containing paclitaxel and polo-like kinase 1 (Plk1)-specific siRNA could instigate synergistic tumor suppression in an MDA-MB-435s xenograft mouse model. The system required 1000-fold less paclitaxel than that required for paclitaxel monotherapy, and micelleplex delivery was independent of carrier-associated toxicity and activation of the innate immune response [113]. 9.2. Active targeting Active targeting enhances the interaction of polymeric micelles with tumor cells, extending the half-life and permitting micelles to effectively permeate into cells via receptor-mediated endocytosis [75]. The development of ligand-coupled and stimuli-responsive micelles for active targeting is described in the following sections. 9.2.1. Ligand-coupled micelles Ligand-coupled polymeric micelles are designed by conjugation of polymeric micelles or nanoparticles with antibodies and antibody fragments (F(ab)/2, F(ab/), and scFv) [114]; proteins including transferrin, ankyrin repeat proteins, and affibodies [115–117]; peptides such as CGNKRTRGC (LyP-1), F3 peptide, iRGD, KLWVLPKGGGC, KLWVLPK, and aptides [118–125]; nucleic acid-based lig-

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ands such as the A10 aptamer and A9 CGA aptamer [126,127]; and small molecules including folic acid, TPP, and ACUPA [128–131]. 9.2.1.1. Folate-coupled micelles. Folate-coupled micelles are primarily used to transport siRNA into cells. Folate has a strong affinity for its binding proteins and receptors, which are selectively overexpressed on cancer cells [132]. For example, the brain, mammary gland, ovary, prostate, lung, and colon epithelial tumors overexpress folate receptors. Therefore, folate-conjugated micelles of appropriate design can be targeted to and internalized in tumor cells by receptor-mediated endocytosis. Moreover, folate-conjugated micelles may revert to diffusion through drug efflux pumps expressed by cancer cells [133], remain present in recycled endosomes (pH 5–6), or diffuse into the cytosol. These features are important in improving the cellular uptake of drugs/siRNA and for evading potential lysosomal uptake of the micelles. Human pharyngeal cancer cells were exposed to fluorescent folate-conjugated and unconjugated pH-sensitive micelles for 3 and 24 h, resulting in intracellular accumulation of fluorescence and indicating that folate conjugation significantly improved the uptake of micelles by tumor cells. Another study used micelles mixed with folate-conjugated pH-sensitive PEO-b-PDLLA and folate-PEG-b-poly(l-histidine) block copolymers. These polymeric micelles also exhibited effective targeting of tumor cells in vitro [134]. Recently, Jang et al. developed a novel siRNA carrier, RNAtr NPs, in which numerous tandem copies of RNA hairpins arising from rolling circle transcription (RCT) can be easily adapted to tumor-targeted and systemic siRNA delivery. RNAtr NPs showed considerably higher melting temperature (Tm = 13 ◦ C), a characteristic closely linked with steadily hybridized structures. At a weight ratio of 1:0.2, the majority of RNAtr/DNA-Chol hybrids did not transfer down and remain in the culture wells. RNAtr/DNA-Chol/FA-DNA hybrids (1:0.2:0.2, w/w/w) exhibited a particle size of 190.1 ± 37.2 nm. RNAtr NPs offer a way to condense large amounts of multimeric RNA transcripts into compact NPs, particularly without the help of polycationic moieties, therefore lessening the possibility of cytotoxicity and immunogenicity by evading the use of synthetic polycationic agents. Flow cytometry data showed that RNAtr NPs had about 82% cell binding efficiency in SKOV3 cells, 3 h post treatment. However, folate-free RNAtr NPs displayed only 11.6% cell binding efficiency because of a lack of folate ligands, indicating that RNAtr NPs can be internalized into cancer cells in a folate receptor-specific manner. RNAtr NPs did not exhibit substantial cytotoxicity in cell viability assays, up to a concentration of 50 ␮g/mL. This approach permits the design of a platform for systemic delivery of siRNA to tumor sites, as the RCT reaction, which enzymatically produces RNA polymers in manifold copy numbers at low cost, can result in directly accessible routes for targeted and systemic delivery [135]. 9.2.1.2. Hyaluronic acid-coupled micelles. Amiji and coworkers assessed the potential of hyaluronic acid-

PEI/hyaluronic acid-PEG for the delivery of MDR siRNA as well as the efficacy of siRNA and paclitaxel co-loaded hyaluronic acid-PEI/hyaluronic acid-PEG nanoparticles to suppress ovarian cancer growth [136,137]. siRNA was encapsulated in HA-PEI/HA-PEG complexes at a mass ratio of 54:1 (polymer:siRNA). A cytotoxicity assay showed that 80% of cells remained viable after treatment with HA-PEI/HA-PEG/MDR1 siRNA for 6 h at a concentration of 600 ␮g/mL, while, the highest concentration of HA-PEI/HAPEG used for transfection was 170 ␮g/mL. This finding indicates an apparent lack of HA-PEI/HA-PEG/MDR1 siRNA cytotoxicity during transfection, as cells were shown to tolerate the complexes. The researchers found that hyaluronic acid-PEI/hyaluronic acid-PEG nanoparticles efficiently delivered MDR siRNA into multidrug-resistant ovarian cancer cells, leading to down-regulation of the MDR gene and expression of Pgp. Administration of siRNA-complexed hyaluronic acid-PEI/hyaluronic acidPEG nanoparticles followed by treatment with paclitaxel produced significant tumor growth inhibition, reduced expression of Pgp, and enhanced apoptosis in a multidrugresistant ovarian cancer mouse model. Their results propose that CD44-targeted siRNA-complexed hyaluronic acid-PEI/hyaluronic acid-PEG nanoparticles could serve as a therapeutic tool to evade multidrug resistance in ovarian cancer. In another study, Ganesh et al. designed and screened a range of CD44-targeting hyaluronic acid-based selfassembling nanoparticulate systems for the targeted delivery of siRNA [138]. To assess encapsulation of siRNA, hyaluronic acid polymer was functionalized with lipids of varying nitrogen content/carbon chain lengths and polyamines. Screening by dynamic light scattering and gel retardation assays exhibited that numerous hyaluronic acid-derivatives could efficiently complex siRNAs and form self-assembled nanoparticulate systems. Several derivatives of hyaluronic acid could transfect siRNAs into CD44 receptor overexpressing cancer cells. The nanoassemblies were prepared at various polymer-to-siRNA weight ratios, ranging from 9:1 to 450:1, to evaluate the effect of polymer concentration on siRNA encapsulation. Interestingly, the blocking of CD44 receptors on cells using free soluble hyaluronic acid before incubation of cy3-labeled siRNA-loaded hyaluronic acid nanoparticles resulted in more than 90% inhibition of receptor-mediated uptake, affirming target specificity. Moreover, SSB/PLK1 siRNA-complexed hyaluronic acid-PEI/PEG nanosystems showed target-specific and dose-dependent gene knockdown in both resistant and sensitive A549 lung cancer cells over-expressing CD44 receptors. Furthermore, these siRNA-encapsulated nanosystems exhibited in vivo tumor selective uptake and target specific gene knock down in solid and metastatic tumors. Yin et al. developed hyaluronic acid-based amphiphilic conjugate (HSOP) redox-sensitive micelles for tumortargeted co-delivery of Aurora A kinase (AURKA)-specific siRNA (si-AURKA) and paclitaxel [139]. The N/P ratios of HOP/siRNA and HSOP/siRNA ranged from 10 to 40. No toxicity of the polymeric micelles was reported. HSOP showed efficient loading capacities for both siRNA and paclitaxel, with desirable redox-sensitivity and adjustable dosing

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ratios confirmed by morphological changes of micelles together with the in vitro release of both drugs in different reducing environments. Furthermore, confocal microscopy and flow cytometry analyses confirmed that HSOP micelles were able to simultaneously co-deliver siRNA and paclitaxel into MDA-MB-231 breast cancer cells by hyaluronic acid receptor-mediated endocytosis, followed by quick transport of siRNA and paclitaxel into cytoplasm. The successful transport and delivery enhanced the synergistic effects of the drugs, resulting in better antitumor efficacy compared with non-sensitive co-loaded micelles and single drug-loaded micelles. An in vivo study showed that HSOP micelles could successfully accumulate in tumors and showed superior antitumor efficacy to redox-sensitive single-drug controls and a non-sensitive co-delivery control. 9.2.1.3. Transferrin-coupled micelles. Since certain cancer cells overexpress the transferrin receptor (depending on the degree of malignancy), tumor targeting may also be accomplished by transferrin-conjugated polymeric micelles. Ren et al. prepared transferrin-functionalized PEG-PLA micelles for potential targeted in vivo delivery to brain glioma [140]. Flow cytometry showed the in vitro targeting potential of the micelles to tumor cells, and fluorescence microscopy imaging of brain sections from C6 glioma tumor-bearing rats showed that transferrinfunctionalized PEG-PLA micelles were able to penetrate tumors in vivo. 9.2.1.4. Luteinizing hormone-releasing hormone (LHRH)coupled micelles. Micelles can also be conjugated to LHRH for tumor targeting. LHRH receptors are prevalent on the plasma membrane of normal cells but are overexpressed in cancer cells, such as ovarian cancer cells. Kim et al. modified polyelectrolyte complex (PEC) micelles with cancer cell targeting moieties for intracellular delivery of VEGF siRNA [141]. LHRH peptide was coupled to the distal end of PEG–siRNA conjugate as a targeting ligand. For A2780 cancer cells, the PEC micelles with LHRH showed increased cellular uptake in comparison with those lacking LHRH, leading to enhanced VEGF gene silencing efficiency through receptor-mediated endocytosis. 9.2.1.5. Epidermal growth factor (EGF)-coupled micelles. An alternative strategy for tumor targeting is the conjugation of polymeric micelles to EGF. Zeng et al. [142] synthesized methoxy PEG-b-poly(␦-valerolactone) MePEG-b-PVL and EGF-conjugated EGF-PEG-b-PVL micelles that selectively targeted the overexpressed EGF receptors (EGFR) on the plasma membrane of breast cancer cells. CMDiI, a hydrophobic fluorescent probe, was loaded into both MePEG-b-PVL and EGF-PEG-b-PVL micelles. Confocal laser scanning microscopy revealed that the CMDiI-functionalized EGF-PEG-b-PVL micelles accumulated intracellularly in MDA-MB-468 breast cancer cells after 2 h of incubation, whereas no noticeable cellular uptake was observed for the CM-DiI-functionalized MePEG-b-PVL micelles. The findings from confocal imaging were verified by calculating the intracellular CM-DiI fluorescence in cell lysate. Moreover, the presence of free EGF was

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found to reduce the uptake of EGF-PEG-b-PVL micelles. Cellular nuclear staining with Hoechst 33258 showed that EGF-PEG-b-PVL micelles primarily accumulated in the perinuclear area while some gathered in the nucleus. 9.2.1.6. Oligopeptide-based targeted micelles. For siRNA cancer therapy, Oe et al. developed PIC micelles to enhance siRNA accumulation in tumors and increase cancer cell uptake after systemic administration [143]. The polymer was designed to add disulfide cross-linking and cyclic RGD peptide ligands for cancer cell targeting, whereas cholesterol-modified siRNA (Chol-siRNA) gave further hydrophobic stabilization to micelles. Neither micelle formulation induced significant cytotoxicity, even at a high siRNA concentration (1000 nM), indicating negligible cytotoxic effects of siRNA micelles at concentrations used for gene silencing studies. The functionalization of cRGD ligands on the micelles efficiently facilitated siRNA accumulation in an in vivo subcutaneous cervical cancer model. Conclusively, the cRGD/Chol-siRNA micelles showed significant in vivo gene silencing, possibly owing to their targeting capability together with better stability conferred through both disulfide cross-linking and hydrophobic interactions of cholesterol [143]. Similarly, Christie et al. developed stable multifunctional micelles specifically containing poly(ethylene glycol)-block-poly(l-lysine) (PEG-b-PLL), and consisting of lysine amines engineered with 2-iminothiolane (2IT) and with a cyclo-Arg-Gly-Asp (cRGD) peptide on the PEG terminus. Formation of micelles occurred spontaneously upon mixing of block copolymer with siRNA, and the theoretical configuration of the 15–20 nm particle displaying siRNA in the core, a PEG shell, and the cRGD peptide on the micelle surface was revealed using atomic force microscopy (AFM). Introduction of the cRGD peptide resulted in enhanced biological activity; however, addition of siRNA optimized gene silencing ability, cell uptake, and subcellular distribution in vitro, and increased accumulation in tumors following IV injection in mice. More specifically, the authors investigated the subcellular distribution of fluorescent-labeled siRNA contained within cRGD-2IT-95 or 2IT-95 micelles, and found that cRGD-2IT-95 micelles displayed a broader cell distribution, suggesting diminished lysosomal accumulation of pDNA polyplexes containing the cRGD peptide. However, 2IT-95 micelles were mostly confined to lysosome compartments. Conclusively, cRGD-2IT-95 micelles were able to co-accumulate with 2IT-95 micelles within lysosomes, but 2IT-95 micelles were unable to colocalize with cRGD-2IT-95 micelles in nonlysosomal regions. This observation suggested that the cRGD peptide influenced the subcellular fate of micelles, resulting in broader subcellular distribution in nonlysosomal regions [144] (Fig. 7). Xiong and Lavasanifar developed polymeric micelles that could incorporate multiple tasks within a single system, such as the ability to house a mixture of therapeutic moieties with varying physicochemical properties (i.e., DOX and siRNA), pH-triggered drug release, cell membrane translocation, and passive and active cancer targeting. Micelles were prepared from degradable poly(ethylene oxide)-block-poly(␧-caprolactone) (PEO-bPCL) block copolymers, equipped with functional groups

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Fig. 7. Micelle structure and properties. (i) Schematic representation of polymeric micelle structure and key components. (ii) Atomic force microscopy image of cRGD-2IT-95 micelles showing spherical structures. (iii) Subcellular distribution of 2IT-95 and cRGD-2IT-95 micelles following coincubation in HeLa-luc cells for 4 h (500 nM siRNA for each micelle formulation). Lysosomes were stained with LysoTracker Green and micelles were prepared with Cy3or Cy5-labeled siRNA for cRGD-2IT-95 and 2IT-95 formulations, respectively. (A–C) Cell images showing individual fluorescence signals. (D) Overlay of all fluorescent signals onto the transmitted image [144]. Copyright 2012. Reproduced with permission from the American Chemical Society.

on both blocks. The PCL block functional group was used to integrate short polyamines for complexation with siRNA or to chemically attach DOX using a pH-sensitive hydrazone linkage. A virus mimetic shell was obtained by attaching two ligands, a cell-penetrating peptide TAT for membrane activity and an integrin Rv␤3-specific ligand (RGD4C) for cancer targeting. The micelles had average diameters ranging from 93 to 108 nm and were labeled with near-infrared fluorescent imaging probes to facilitate monitoring of micelle location in vivo upon IV administration. Dy677labeled siRNA was also used to evaluate the in vivo stability of the siRNA carrier. These micelles were shown to be capable of siRNA and DOX delivery to their intracellular targets, resulting in the inhibition of P-gp-mediated DOX resistance in vitro and targeting of Rv␤3-positive tumors in vivo [145]. Asn-Gly-Arg (NGR)-containing peptides can be used as ligands for targeted delivery of polymeric micelles to activate blood vessels in tumors. Son et al. synthesized a polymer based on branched PEI (BPEI) thiolated with propylene sulfide and combined it with alpha-maleimideomega-N-hydroxysuccinimide ester polyethylene glycol (MAL-PEG-NHS, MW: 5000) and cyclic NGR (cNGR) peptide [146]. The gene nanocarrier exhibited effective tumor targeting using the cNGR peptide. Wang et al. evaluated the anti-tumor efficacy of NGR peptide-linked PEG-b-PLA polymeric micelles [147]. HT1080 cells were selected as a model of positive tumor cells whereas HUVECs were used as a model of tumor endothelial cells. The findings exhibited that actively targeted polymeric micelles showed better uptake and stronger adhesion than undecorated polymeric micelles. Angiopep-2 (ANG) was used to prepare ANGpoly(lactic-co-glycolic acid) (PLGA) functionalized nanocarriers that encapsulated both EGFR siRNA and DOX

[148]. The N/P ratio of ANG/PLGA/DOX to siRNA ranged from 15:1 to 90:1. Results indicated that ANG/PLGA NPs exhibited no significant cytotoxicity to U87MG cells, even at a concentration of 100 mg/mL, suggesting biocompatibility of the drug carrier itself. The nanocarrier efficiently delivered siRNA and DOX into U87MG cells, resulting in significant EGFR silencing, apoptosis, and cell inhibition in vitro. Subsequent in vivo experiments using a brain orthotopic U87MG glioma xenograft model showed that the nanocarriers not only prolonged the lifespan of glioma-bearing mice but also produced obvious cell apoptosis in glioma tissue. Huo et al. prepared PIC micelles composed of rabies virus glycoprotein (RVG) peptide-tagged PEGylated polyasparthydrazide (PAHy) derivatives for siRNA delivery to the brain [149]. The PIC micelles were formed by electrostatic attraction between the polymer and siRNA. Using PEG, the siRNA-complexed PIC micelles were linked with RVG. The COOH-PEGg-PAHy-GTA/siRNA micelles were prepared at weight ratios of 10:1, 25:1, 50:1, 75:1, 100:1, and 150:1 polymer/siRNA. The cytotoxicity of siRNA-loaded PIC micelles was evaluated in Neuro 2a and HeLa cells by MTT assay. The results revealed that the micelles of COOH-PEGg-PAHy-GTA/siRNA and RVG-PEG-g-PAHy-GTA/siRNA did not exhibit significant cytotoxicity, even at the highest siRNA concentration of 8 ␮g/mL. Furthermore, no obvious difference in cytotoxicity was observed between the two types of micelles. Both types exhibited good biocompatibility in the Neuro 2a cell line. In comparison with non-functionalized micelles, RVG-functionalized micelles were easily internalized by the Neuro 2a cells and efficiently silenced gene expression. The in vivo study also verified the brain targeting ability of RVG-functionalized micelles.

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Kanazawa et al. designed nose-to-brain siRNA delivery micelles consisting of PEG-PCL copolymers conjugated to cell penetrating peptide (CPP), TAT (MPEG-PCL-TAT) [150]. Dextran (Mw: 10,000 Da) was used as a model siRNA in this study. As compared with IV delivery of dextran with or without MPEG-PCL-TAT, the intranasal dextran delivery with MPEG-PCL-TAT improved delivery to the brain. Moreover, the MPEG-PCL-TAT enhanced the transport along the trigeminal and olfactory nerve pathway owing to its high permeation across the nasal mucosa. 9.2.1.7. Antibody-coupled micelles. Palanca-Wessels et al. designed a delivery system consisting of (i) a streptavidinconjugated monoclonal antibody (mAb-SA) against CD22 and (ii) a diblock copolymer consisting of both a pHresponsive block to facilitate endosome release and a positively charged siRNA condensing block [151]. A minimum polymer-to-siRNA molar ratio of 2:1 was required for complete complexation, and the polymer was found to be non-toxic to normal cell lines. Improved uptake of siRNA was exhibited in transduced HeLa-R and DoHH2 lymphoma cells expressing CD22 but not in CD22 negative HeLaR cells. Compared with non-targeted polymeric micelles, gene knockdown was significantly enhanced with CD22targeted micelles. The CD22-targeted polymeric micelles containing 15 nmol/L siRNA showed 70% reduction of gene expression in DoHH2 cells. Dou et al. designed a nanocarrier consisting of a fusion protein of an anti-Her2 antibody fragment with a positively charged protamine (F5-P) for the delivery of DNA methyltransferases 1 and/or 3b siRNA (siDNMTs) into Her2-expressing breast tumor cells [152]. The F5-P/FAMsiRNA complexes had a molar ratio of 1:5. The carrier F5-P efficiently bound siRNA and delivered it to Her2-expressing BT474 breast cancer cells but not to Her2 non-expressing MDA-MB-231 breast cancer cells. Furthermore, siDNMTs delivery to BT474 cells efficiently silenced DNMT expression and enhanced the demethylation of RASSF1A tumor suppressor gene promoter, resulting in tumor cell proliferation suppression. 9.2.1.8. Carbohydrate-coupled micelles. Zhu et al. conjugated galactosylated PEG (Gal-PEG) to oligonucleotide (ODN) by a beta-thiopropionate acid-labile ester linkage [153]. The ODN with 3 -thiol functional group and Gal-PEG-acrylate were used at a molar ratio of 1:100. After intracaudal injection into rats, Gal-PEG-33P-ODN was quickly cleared from the circulation, and 60.2% of the dose was found to be accumulated in the liver 30 min after injection, more than that deposited after injection of 33P-ODNs. A biphasic plasma concentration versus time profile of Gal-PEG-33P-ODN was observed, with halflife (t1/2 ) of distribution and elimination as 4.38 ± 0.36 and 118.61 ± 22.06 min respectively. Pre-administration of excessive Gal-bovine serum albumin (BSA) reduced the hepatic uptake of Gal-PEG-33P-ODN from 60.2% to 35.9%, suggesting that galactose initiates the asialoglycoprotein receptor-mediated endocytosis of Gal-PEG-33P-ODN by hepatocytes. Zhu and Mahato also conjugated the antisense strands of siRNA to mannose 6-phosphate PEG (M6P-PEG) and

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Gal-PEG for siRNA targeted delivery to hepatic stellate cells (HSCs) and hepatocytes, respectively [154]. The N/P ratio of 3 -sulfhydryl siRNA to GalPEG-OPSS (or M6PPEG-OPSS) was 1:100. The polymer GalPEG-OPSS (or M6P-PEG-OPSS) was found to be non-toxic to normal cells. Without transfection reagents, both M6P-PEG–siRNA and Gal-PEG–siRNA conjugates silenced luciferase gene expression to about 40%, whereas 98% gene silencing was observed using cationic nanocarriers at a similar dose. The conjugation of M6P-PEG and Gal-PEG to TGF-␤1 siRNA also silenced the endogenous TGF-ˇ1 gene expression. Wang et al. designed N-acetylgalactosaminefunctionalized mixed micellar nanoparticles (Gal-MNP), capable of delivering siRNA to hepatocytes and silencing the target gene expression upon systemic administration [155]. MNP/siRNA complexes were prepared at N/P ratios ranging from 0 to 10 (siRNA concentration was fixed at 1 ␮M). Importantly, this delivery system did not activate the innate immune response or induce hepatotoxicity, a promising outcome for systemic delivery of siRNAs for liver disease therapy. The hepatocyte-targeting effect of Gal-MNP was exhibited by efficient accumulation of fluorescent siRNA in primary hepatocytes in vitro and in vivo. After IV administration of Gal-MNP/siapoB to BALB/c mice, efficient down-regulation of apolipoprotein B (apoB) expression was achieved in the liver of mice, at both transcriptional and protein level. Innate immunity or positive hepatotoxicity was not induced by Gal-MNP/siRNA systemic delivery.

9.2.2. Stimuli-responsive micelles 9.2.2.1. pH-responsive polymeric micelles. The extracellular pH (pHe) of most solid tumors is distinctly different from the physiological pH (7.4) of the adjacent normal tissue [156]. The average pH of tumors is around 7.06, commonly ranging from pH 5.7 to 7.8. Non-invasive imaging of in vivo tumor pH has been enabled by magnetic resonance spectroscopy on the basis of the chemical change in the pH-dependent resonance frequency. The local pH of the endosomal and lysosomal compartments ranges from 5 to 6 in the majority of cell types. Therefore, a weak acidity is believed to be one of the functional triggers for the specific release of chemotherapeutics/siRNA at tumor tissue sites and/or inside tumor cells [157,158]. Titratable moieties such as weak acids or weak bases are extensively applied to micelle-forming copolymers to cause disruption or interior structural alterations of polymeric micelles in acidic biological environments [159–161]. Particularly, poly(l-histidine) (P(His)) is generally used as a pH-responsive element in polymeric micelles [162]. The imidazole of the histidine residue is freely protonated, yielding positively charged micelles. Several studies have reported pH-responsive polymeric micelles composed of PEG-b-P(His) block copolymers [163]. Lee et al. designed block copolymer micelles from PEG-b-P(His) and PEG-b-PDLLA to improve the micelle stability at a physiological pH [163]. Additionally, pHresponsive, folate-conjugated block copolymer micelles [folate-PEG-b-P(His)/PEG-b-PDLLA] were designed and shown to exhibit efficient intracellular localization.

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Kim et al. reported a method to stabilize pH-responsive P(His)-based cores by the addition of hydrophobic lphenylalanine to the P(His) main chains (P(His-co-Phe)) [164]. Using a carboxylic acid-based system, Leroux et al. designed pH-responsive polymeric micelles containing a ternary arbitrary coblock of methacrylic acid, octadecylacrylate, and N-isopropylacrylamide as a pH-responsive unit [165]. The micelles showed a pH-dependent phase transition around pH 5.8, as evidenced by the altered structure of the micelle inner core. Biodegradable polyesters have been used as hydrophobic core-forming polymers in the design of block copolymer micelles. Biodegradable polyesters are commonly used as hydrophobic blocks in the hydrolysis-induced disruption of polymeric micelles, although the rate of polyester hydrolysis is comparatively slow under slightly acidic conditions (pH 5–6), necessitating several days to attain thorough disruption of the micelles’ structure [166]. Heller et al. described polymeric micelles consisting of PEG-b-poly(ortho ester) (POE) block copolymers that are pH-sensitive [167]. These micelles had a nanoscopic size (50–80 nm) and low CMC value. Acid-labile chemical bonds such as acetyl [168], hydrazine, cis-aconityl, and oxime [169] moieties are commonly used either to modify the structure of micelle-forming polymers or to conjugate drugs to polymer backbones for the design of efficient drug delivery systems that are reactive at the acidic pH of endosomes/lysosomes. Gillies and Fréchet, designed PEG micelles with a pH-sensitive hydrophobic core disruption property using an acid-labile core-forming P(Asp) derivative to increase the pH-triggered drug release from the carriers [170]. In their innovative system, a PEGdendrimer hybrid was used for the design of micelles with a pH-sensitive drug release function. The addition of hydrophobic aromatic groups to neighboring sites that are involved in the core formation of the block was accomplished via cyclic acetyl linkage [171]. The features of the designed micelles (CMC, disruption rate, and size) were controllable by fine-tuning the dendrimer generation, chemical structure, corona-forming PEG length, and the type of hydrophobic acetyl linkage [172]. 9.2.2.2. Temperature-responsive polymeric micelles. The temperature change during cooling/heating processes can be used as a trigger signal for altering the functional and structural characteristics of stimuli-responsive polymeric micelles. Particularly, mild heating (up to 41–43 ◦ C) of the body may be used as a potential cancer treatment [173] because of the high-temperature sensitivity of cancer cells: mild heating specifically affects the biological processes within cancer cells, including altered receptor expression, microtubule disruption, and inhibited DNA repair and synthesis. Temperature-responsive polymers exhibit a characteristic conformational change in reaction to the environmental temperature. These polymers are divided into two types on the basis of the temperature dependency of their solubility. One type of polymer shrinks or precipitates below certain temperatures (the upper critical solution temperature (UCST)) [174], whereas the other type shrinks and becomes insoluble in water above a

specific temperature, known as the lower critical solution temperature (LCST). LCST-type polymers have been widely used in drug delivery systems and as biomaterials [175]. In particular, poly(N isopropylacrylamide) (PIPAAm) has interested researchers working on intelligent biomedical applications. The latest development in polymer chemistry facilitates the selection of several types of polymers that are temperature-responsive, such as poly(2-alkyl-2-oxazoline)- [176], poly(oligo(ethylene glycol) methacrylate)- [177], and polypeptide-based polymers (for example, elastin-like polypeptide) and others. Generally, the LCST control of poly(N-substituted acrylamide) derivatives is achieved by random radical copolymerization with different ratios of co-monomers possessing hydrophobic/hydrophilic properties. Recently, the outmost surface functionalization of temperature-responsive micelles has been designed using block copolymers such as end-functional thermoresponsive blocks to add exclusive features to polymeric micelles, including bio-imaging and active targeting functions [178]. Hydrophobic groups placed at the periphery of micelles significantly stimulated the dehydration of corona-forming PIPAAm derivatives and led to a significant LCST shift in the micelles to a lower temperature, in contrast with that of micelles possessing hydrophilic surface moieties. Additionally, the amplitude of the LCST shifts relied on the molecular weight of the temperature-responsive chains [178]. The findings of earlier experiments showed that hydrophobically terminated linear PIPAAm systems had a lower LCST value than non-modified pure PIPAAm because of increased dehydration of the proximal IPAAm units through the freely movable hydrophobic end-groups [179]. A PIPAAm-corona micelle formed by mixing hydroxyland phenyl-based block copolymers showed a sharp phase transition at a temperature between specific LCST values of the individual homogeneous micelles [180]. This distinctive feature in the phase transition of micelles indicates that micellar LCST values can be thoroughly controlled by modifying the surface chemistry with certain stimuli such as bio-related interactions, redox reactions, light, and pH. Additionally, biodegradable polyesters such as PCL, PLA, and their copolymers have been used to construct hydrophobic cores [181–183]. 9.2.2.3. Light-responsive polymeric micelles. Light (UV/visible or near-infrared (NIR) light) is a practical activation signal as it can be applied to the body externally with temporal and spatial control. Additionally, several factors, including light intensity and wavelength, can also be controlled during photo-induced reactions. Polymeric micelles with light-responsive functions are categorized into two types on the basis of micelle disruption mechanism, which triggers either light-induced cleavage of chemical bonds or light-switchable property changes such as photo-isomerization. Photochromic compounds are well-known molecules that display reversible isomeric transformations in response to individual light wavelengths. These compounds are widely used to prepare light-responsive materials [184]. In the past five decades, azobenzene

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derivatives have been one of the most extensively studied photochromic compounds [185]. Wang et al. reported polymeric micelles that altered their morphology via photo-isomerization of the azobenzene units present in the block copolymer [186]. These morphological changes are largely accredited to the adaptable polarity of the hydrophobic blocks when exposed to visible and UV light. Some years ago, the photo-isomerization reaction of spiropyran (SP) derivatives was used as a basis for the formation of core–corona micelles that exhibited reversible light-induced micelle formation. The SP compounds undergo reversible isomerization between the hydrophilic zwitterionic and hydrophobic SP merocyanine forms upon irradiation with visible and UV light [187]. Hence, the significant change in the polarity of the SP units introduced in the core-forming chain results in reversible block copolymer assembly. Polymeric micelles containing a block copolymer called PEG-b-poly(spiropyran-methacrylate) (PEG-b-PSpMA) are disrupted by UV radiation at a wavelength of 365 nm, and are recovered at the visible light wavelength of 620 nm [188]. Therefore, the light-induced disruption of the micelle structure allows the respective drug load to be released. Zhao [189] and Yan et al. [190] designed polymeric micelles exhibiting an irreversible light-responsive dissociation profile. Poly(1-pyrenylmethyl methacrylate) (PPyMA) was used for the hydrophobic core-forming chains in the first report on these polymeric micelles [191,192]. Micelles composed of poly(2-nitrobenzylmethyl methacrylate) (PNBMA) [192] are another example of light-responsive polymeric micelles. A light-triggered drug release from the PNBMA core has been shown, and increasing the intensity of the UV light accelerated the drug release rate. A light-induced drug release system including 2-nitrobenzyl can also be attained by NIR irradiation, although the drug release rate is far slower than that under UV irradiation because of the low effectiveness of two-photon absorption of the 2-nitrobenzyl unit. This issue was addressed by introducing 7-diethylamino-4-(hydroxymethyl)coumarinyl, showing two-photon absorption at 794 nm, as a substitute chromophore in the hydrophobic block of poly([7(diethylamino)coumarin-4-yl]methyl methacrylate) [193]. 9.2.2.4. Ultrasound-based tumor targeting. This approach uses focused ultrasound for payload release. Ultrasound activates drug release from the micelles along with disruption of the plasma membrane, thereby improving cellular drug uptake [194,195]. Using this approach, the transducer is positioned to interact with a layer of water or waterbased gel applied to the skin, thus avoiding the need for an invasive/surgical administration route. Ultrasound may improve the cellular uptake of both released and entrapped siRNA at the site of irradiation, enhance the drug release from the micelles into the tumor interstitium, increase the diffusion of the drug through the tumor interstitium, improve the accumulation of drug-loaded micelles at the tumor site, and produce thermal effects. All of these properties improve the accumulation of siRNA inside the tumor.

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The technique can be further improved by modifying several features, including the physicochemical properties of the micelles, time between the injection of the drug and the ultrasound application, sonication frequency, and type of ultrasound waves applied [196] (Fig. 8). The physicochemical properties of the micelles include the state of the micelles’ inner core and the micelles’ hydrophobicity, as these affect the accumulation of drug in the tumor [197,198]. 9.2.2.5. Polymeric immunomicelles. Micelles having specific antibodies attached to their surface are referred to as immunomicelles, and target the drug/siRNA by specific interactions; they may therefore be adapted to a variety of targets. It has been shown that specific monoclonal, antinuclear nucleosome-restricted autoantibodies identify the exterior of several tumors through surface-attached nucleosomes, but not the exterior of normal, healthy cells [199]. A modified design method that uses PEG-poly(ethylene) (PEG-PE) with a p-nitrophenylcarboxyl-activated-free PEG end to link monoclonal antibodies (mAbs) to PEG-PE micelles has been introduced [200]. This novel PEG-PE linker retains its stability at a pH lower than 6; once the pH is higher than 7.5, the linker reacts with the amino groups of several peptides and proteins, including mAbs, yielding a stable carbamate linkage. The immunomicelles are quantified by sodium dodecyl sulfate polyacrylamide gel electrophoresis or fluorescence-labeled antibody techniques [201,202]. The coupling of different antibodies to polymeric micelles did not significantly alter the size of micelles, as reported by several researchers [203]. 9.2.3. Photodynamic therapy Photodynamic therapy (PDT) is an alternative method for the treatment of visible tumors using photochemical irradiation of tumor cells with the aid of photosensitizing agents. This technique permits the use of low quantities of siRNA, thereby considerably decreasing the quantity of PDT siRNA accumulating in healthy cells. 10. siRNA-based therapy Although numerous obstacles to siRNA delivery prevail, some carriers have been developed and investigated for in vivo application. The most notable uses for the delivery of siRNA include cancer and hepatic therapies. Two further administration routes, oral and topical, also exhibit significant potential for the development of novel siRNA therapeutics. 10.1. Targeted delivery of siRNA to the liver The liver is an important target for siRNA delivery as it is the site of several metabolic disorders, cancer, and viral infections [36,204,205]. Liver cirrhosis has been treated using knockdown of rat protein gp46 with siRNA [206]. Liver-targeted gene knockdown was achieved by IV injection of vitamin A-bound nanoparticles. Collagen secretion decreased following treatment and fibrotic zones were considerably reduced, resulting in 100% survival rates at most recurrent dosing intervals compared with 0% survival

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Fig. 8. Illustrative principle of siRNA delivery to tumor tissue and tumor cells using siRNA-NBs and low-frequency US exposure. (A) The siRNA-NBs complexes were small enough to accumulate in tumor tissues after intravenous injection. (B) With the assistance of low-frequency US, the permeability of endothelium and cancer cell membranes was enhanced simultaneously due to the activated cavitation of siRNA-NBs. The released siRNA micelles were directly delivered into cancer cells by the enhanced permeability of cancer cell membranes, known as “sonoporation”. In addition, after low-frequency US exposure, cationic siRNA copolymer micelles, which were effectively accumulated in tumor tissue by cavitation, can enter the cancer cells by cell endocytosis. Abbreviations: US, ultrasound; siRNA-NBs, siRNA-loaded nanobubbles [196]. Copyright 2013. Reproduced with permission from Elsevier Ltd.

for controls. Another study evaluated polymeric nanoparticles consisting of PEG for shielding, N-acetylgalactosamine for liver targeting, and a poly(vinyl ether) for endosomal lysis and observed 80% knockdown of apoB following IV injection of 2.5 mg siRNA/kg body weight [207]. More than 90% of PEGylated nanoparticles [208] were found to accumulate in the liver after IV injection. Gene silencing lasting almost 1 month was achieved using a factor VII knockdown model, with 50% knockdown efficiency observed after 21 days. Furthermore, monthly dosing resulted in similar silencing intensities, with no occurrence of an immune response. Additional experiments using this strategy demonstrated that other nanoparticles could also achieve substantial knockdown of factor VII at doses of 0.1 mg/kg in mice, but for shorter intervals compared with the higher dose of 1 mg/kg [209–211]. Furthermore, five different hepatocyte siRNA sequences were combined into one nanoparticle formulation, resulting in more than 50% knockdown of the individual gene target at a dose as low as 0.1 mg/kg per siRNA [212]. Effective reduction of choles-

terol levels in both nonhuman and rodent primates was observed by a research group using similar nanoparticles. Using siRNA to target LDL receptors, the researchers reported a 60% reduction in rodent cholesterol levels and a 50% reduction in LDL cholesterol concentration in primates following IV administration at a dose of 5 mg/kg [212]. 10.2. Targeted delivery of siRNA to tumors Ren et al. [213] described a peptide-based nanoparticulate system targeting the ovarian cancer oncogene ID4. The nanoparticles were observed to selectively target and penetrate deep into the tumor tissue. After IV administration of 5 mg/kg siRNA every 3 days for 25 days, tumor growth was reduced by 82% and survival rate was improved in comparison with that of controls and nonsense RNA nanoparticles. In another experiment, a protein consisting of Her2 single chain-fragmented antibodies (ScFvs) attached to protamine was complexed with polo-like kinase 1 (PLK1) siRNA for breast cancer sup-

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pression [214]. The nanoparticles considerably reduced Her2+ tumor growth, decreased the number of metastases, and extended the survival period after twice-weekly IV administration of a 2 mg/kg dose for 28 days. These nanoparticles were also used to deliver a mixture of antitumor siRNA, resulting in improved outcomes compared with PLK1 siRNA alone. Leuschner et al. [215] demonstrated that nanoparticles could effectively deliver siRNA to inflammatory monocytes. When the nanoparticles were applied to a colorectal cancer model and treated with CCR2 siRNA, a 75% decrease in tumor-linked macrophages was observed. This study also indicated the potential efficacy of this delivery system in several other disease models, such as myocardial infarction and atherosclerosis, and for extending the survival of transplanted pancreatic islets. Liu et al. [216] prepared micelles that could target siRNA delivery to hypoxic tumors to suppress hypoxiainducible factor-1␣. Following in vivo evaluation in a mouse prostate cancer model, the IV administration of 2 mg/kg siRNA-complexed micelles was shown to inhibit tumor growth by almost 60% and also increased responsiveness to DOX treatment. The in vivo co-delivery of paclitaxel and PLK1 siRNA using these micelles in a mouse breast cancer xenograft model also exhibited substantial antitumor effects and augmented the chemotherapeutic efficacy of paclitaxel [113]. The co-delivery of gemcitabine monophosphate with siRNA against the c-Myc oncogene using lipid/calcium phosphate nanoparticles produced a noteworthy reduction in tumor growth in a nonsmall-cell lung cancer mouse model [217]. Using the same nanoparticles, vascular endothelial growth factor (VEGF) siRNA improved treatment response, as demonstrated by a reduction in tumor vascular density and growth [218]. For tumor-targeted delivery of siRNA using nanoparticles, it is vital to consider the tumor vasculature to capitalize on the EPR effect. Recently, Li et al. exhibited that improving the tumor vascularity by ectopic expression of VEGF resulted in better efficiency of siRNA delivery to the tumor interstitium [219]. This outcome was validated when equating a poorly vascularized and highly vascularized tumor model, as siRNA-mediated gene knockdown was 60% in the latter model versus 15% in the former. 10.3. Miscellaneous routes for administration of siRNA nanoparticles Oral delivery is a suitable administration route for numerous therapeutics. Recently, Yin et al. [220] used micelles to orally deliver tumor necrosis factor alpha (TNF␣) siRNA for the treatment of inflammation. The addition of mannose-targeting ligand directed particle uptake to macrophages and enterocytes in the small intestine. Cysteamine was also incorporated to form disulfide bonds with the mucosal glycoproteins, maintaining localization of nanoparticles. Using a hepatic injury model, treatment with siRNA nanoparticles resulted in an 80% reduction in TNF-␣ levels in the spleen, liver, and lung, signifying that the orally-administered nanoparticles were transported to other organs after absorption from intestine. Topical delivery represents another beneficial but complex delivery route. Zheng et al. used 13-nm gold

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nanocarriers to deliver surface-conjugated siRNA [221]. They found that topical delivery of siRNA targeting EGFR resulted in 40% decreased epidermal thickness, decreased downstream signaling activation, and reduced EGFR expression in keratinocytes. The treatment did not exhibit accumulation of gold nanocarriers in internal organs, signifying its potential for the treatment of genetic disorders, tumors, and skin lesions. 11. Conclusions and future directions During the past few decades, the development of polymeric micelles to deliver a range of entities from chemotherapeutics to oligonucleotides, antibodies, siRNA, and DNA, has undergone a rapid evolution. The development of these sophisticated and advanced micelles has been enabled by chemically modifying the assembly of the block copolymers that are involved in micelle formation. Currently, there is a change in the trend to develop micelles loaded with only one drug type to micelles loaded with multiple payload types, which can be used for active targeting. This approach may aid in the delivery of imaging mediators and agent sensitive to distinct signals that are produced either externally or within the tumor microenvironment, and therefore allows for the temporal and spatial regulation of payload release. Nucleic-acid aptamers designed from a pool of arbitrary sequences for binding to a particular molecule of interest have potential as targeted delivery siRNAs functioning as antibody substitutes. Aptamers possess multiple benefits, including low immunogenicity, process-compatible storability, ready-to-use chemical synthesis, and highly specific binding to receptors and other proteins. Yet, more DNA and RNA aptamers need to be established for particular diseases or as markers of cancers to increase the use of this therapeutic strategy. Another promising technique for improved siRNA delivery includes the covalent conjugation of siRNAs to protein transduction domains or CPPs. Nanocassettes generated from DNA-dependent siRNA expression are a novel strategy for targeted drug delivery and the in vivo expression of siRNAs. All of these techniques for the targeted delivery of siRNAs are in their initial development stages. Increased interest by researchers is required to further investigate the potential of these techniques. A number of examples have been discussed in this review, which has covered a variety of polymeric micelle alterations, from polymeric micelles that depend on the EPR effect to more sophisticated systems that are stimulisensitive and incorporate targeting ligands, and including systems incorporating several modifications. The future of polymeric micelles can be regarded with optimism, not least because of the simplicity of introducing structural changes to the polymeric micelles and these micelles’ inherent advantages. Nevertheless, siRNA delivery by polymeric micelles requires more intensive studies before they may be effectively applied as clinically suitable therapeutics. These studies should, for example, address the difficulties in siRNA delivery and the number of obstacles faced by siRNA to reach its final site of action, i.e., the cytoplasm. On the basis of previous and ongoing research, the design and development of polymeric micelles for siRNA

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delivery should follow specific standards. These standards include the use of targeting ligands to enhance cell-specific uptake, the incorporation of endosomolytic moieties that aid in the endosomal escape of siRNA, the use of polycations that efficiently condense siRNA and enhance transfection, PEGylation to avoid non-specific siRNA interactions, a fast systemic clearance, fine-tuning of the size of micelles to prevent the micelles’ renal filtration and phagocytosis, and, to prevent recognition by immune surveillance cells, chemical alterations of siRNA to avoid immunostimulation and enhance the stability against nuclease degradation [222]. It is important to optimize all of the above-mentioned parameters to develop effective micelles for siRNA delivery, such that augmenting one of these parameters does not negatively affect the others. For example, polycations condense the siRNA inside the micelles and enhance the siRNA’s transfection efficacy; nonetheless, the in vitro and in vivo safety profile of micelles may be affected by polycations, an issue that highlights the importance of using low molecular weight polycations (which are comparatively safe). Thus far, several oncogenes, such as those related to drug resistance, apoptosis, angiogenesis, and proliferation, have been targeted with siRNA-complexed formulations. However, the efficient and safe targeted delivery of siRNA remains challenging. New polymers should be developed that permit effective protection and loading of siRNA with fewer side effects. Future research should also focus on the development of more stable siRNA-complexed micelles with a longer shelf life. In vivo safety issues, including immune stimulation and off-target effects of both micelle-forming materials and siRNA, should also be given focused attention. Lastly, to achieve optimum siRNA dosing regimens, the biodistribution of siRNA following IV administrations should be thoroughly determined. Acknowledgments MCIMA would like to acknowledge Universiti Kebangsaan Malaysia (UKM) for a research grant (GPK007818) for funding this study. AKI would like to acknowledge funding support from National Cancer Institute’s (NCI’s) 5R21CA179652-02 grant and starup funding from Wayne State University. References [1] Liu G, Wong-Staal F, Li QX. Development of new RNAi therapeutics. Histol Histopathol 2007;22:211–7. [2] Gondi CS, Rao JS. Concepts in in vivo siRNA delivery for cancer therapy. J Cell Physiol 2009;220:285–91. [3] Lee SJ, Kim MJ, Kwon IC, Roberts TM. Delivery strategies and potential targets for siRNA in major cancer types. Adv Drug Deliv Rev 2016;104:2–15. [4] Bumcrot D, Manoharan M, Koteliansky V, Sah DWY. RNAi therapeutics: a potential new class of pharmaceutical drugs. Nat Chem Biol 2006;2:711–9. [5] Zhao Y, Wang W, Guo S, Wang Y, Miao L, Xiong Y, Huang L. PolyMetformin combines carrier and anticancer activities for in vivo siRNA delivery. Nat Commun 2016;7, 11822/1-9. [6] Leng Q, Woodle MC, Lu PY, Mixson AJ. Advances in systemic siRNA delivery. Drugs Future 2009;34:721–37. [7] Melnikova I. RNA-based therapies. Nat Rev Drug Discov 2007;6:863–4. [8] Mishra V, Kesharwani P, Jain NK. siRNA nanotherapeutics: a Trojan horse approach against HIV. Drug Discov Today 2014;19:1913–20.

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