nanohydrogels for delivering macromolecular therapeutics.

COREL-07187; No of Pages 12 Journal of Controlled Release xxx (2014) xxx–xxx

Contents lists available at ScienceDirect

Journal of Controlled Release journal homepage: www.elsevier.com/locate/jconrel

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Kuntal Ganguly, Kiran Chaturvedi, Uttam A. More, Mallikarjuna N. Nadagouda, Tejraj M. Aminabhavi ⁎

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Department of Pharmaceutical Engineering and Chemistry, SET’s College of Pharmacy, S.R. Nagar, Dharwad 580 002, India

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Article history: Received 11 March 2014 Accepted 7 May 2014 Available online xxxx

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Keywords: Macromolecules Peptides/proteins Microspheres Nanoparticles Genes Hormones Vaccines

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1. Introduction

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The development of recombinant macromolecular therapeutics has grown quite rapidly over the past decade due to the advent of peptide and protein drugs [1]. In recent years, macromolecular therapeutics such as proteins, peptides, small interfering RNA (siRNA), vaccines and hormones have emerged as a significant class of medicine used in the treatment of various deadly diseases. With more than 130 FDA approved products available today in the market and many more in the pipeline, such drugs are gaining a significant importance in almost every discipline, such as cancer therapy, inflammatory disease, vaccines, and as diagnostics. These drugs have numerous advantages over the small-molecule generic drugs, since they are highly specific and exhibit a complex set of functions such as biochemical reactions, protein-based membrane receptors and channels, cellular or organ transport of molecules and transcellular scaffolding support for which small synthetic molecules can hardly mimic [2]. Macromolecular drugs do not easily cross the mucosal surfaces and biological membranes, since these are susceptible to loss of native structure through cleavage of peptide bonds and destruction of amino acid residues (e.g., proteolysis, oxidation, deamination, and elimination) and conformational changes due to the disruption of non-covalent interactions such as aggregation, precipitation, and adsorption. Specialized uptake mechanisms like transmucosal M-cell uptake in Peyer’s patches and other lymphoid tissues may be necessary to transport

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Increased interest in developing novel micro/nanohydrogel based formulations for delivering macromolecular therapeutics has led to multiple choices of biodegradable and biocompatible natural polymers. This interest is largely due to the availability of large number of highly pure recombinant proteins and peptides with tunable properties as well as RNA interference technology that are used in treating some of the deadly diseases that were difficult by the conventional approaches. The majority of marketed drugs that are now available are in the form of injectables that pose limited patient compliance and convenience. On the other hand, micro/nanotechnology based macromolecular delivery formulations offer many alternative routes of administration and advantages with improved patient compliance and efficient or targeted delivery of intracellular therapeutics to the site of action. This review outlines and critically evaluates the research findings on micro and nano-carrier polymeric hydrogels for the delivery of macromolecular therapeutics. © 2014 Published by Elsevier B.V.

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Polysaccharide-based micro/nanohydrogels for delivering macromolecular therapeutics

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⁎ Corresponding author. Tel.: +91 836 2448540; fax: +91 836 2467190. E-mail address: [email protected] (T.M. Aminabhavi).

such water-soluble macromolecules through mucosal surface to systemic circulation, since these are prone to rapid clearance in liver as well as other body tissues and may require accurate dosing [3]. Polymeric (especially those of polysaccharide-based)-based delivery systems will diminish the inherent instability of these drugs to improve their bioavailability after administering through oral, nasal, pulmonary and other routes [4]. Presently, protein drugs and antigens are administered parenterally i.e., by subcutaneous (sc) or intramuscular injections as well as intravenous (iv) infusions, but these pose problems of oscillating drug concentrations [5]. Drugs like growth hormone, insulin, oxytocin, parathyroid hormone, and vasopressin have short half-lives of b25 min [6], which necessitate multiple injections per week causing the compliance issues, especially when long-term treatment is required as in the treatment of diabetes mellitus by insulin. These drawbacks impose immense challenges and opportunities for developing delivery vehicles using biopolymeric hydrogels. Among the various approaches, researchers have developed needlefree administration routes with high bioavailability such as pulmonary, oral, and nasal delivery [7–9]. Other approaches include extending circulation time and masking immunogenicity of protein drugs through conjugation with other biopolymers as well as developing injectable or transmucosal controlled release (CR) systems including liposomes, polymeric micro/nanoparticles, and hydrogels [4]. Therefore, development of efficient micro/nanocarrier-based delivery systems provides tremendous opportunities for improving the patient compliance and pharmaco-economic benefits. This review compiles the literature on such materials since 2000 until now. The current status and future

http://dx.doi.org/10.1016/j.jconrel.2014.05.014 0168-3659/© 2014 Published by Elsevier B.V.

Please cite this article as: K. Ganguly, et al., Polysaccharide-based micro/nanohydrogels for delivering macromolecular therapeutics, J. Control. Release (2014), http://dx.doi.org/10.1016/j.jconrel.2014.05.014

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Hydrogels are the cross-linked 3D network structures prepared from hydrophilic polymers that are capable of retaining a large amount of water and remain insoluble due to their physical and/or chemical cross-linking. Since their early development in the 1960s [10], innumerable hydrogels are available for a wide range of pharmaceutical applications [11]. These hydrogels can be prepared from both natural and synthetic polymers composing homopolymers, copolymers, and interpenetrating polymer networks (IPNs) by selecting proper building blocks as well as using relevant cross-linking approaches [12–14]. Hydrogels possess high water content and are soft networks, resembling those of natural extracellular matrices that minimize the tissue irritation or cell adherence [15]. The high loads of water-soluble therapeutically active proteins, peptides, siRNA, DNA, vaccines, etc., can be encapsulated into their 3D networks, due to their porous structure along with retaining high water content. Unlike other delivery systems (microparticles, emulsions, etc.), where preparation conditions are sometimes detrimental to proteins (i.e., use of organic solvents and protein denaturating processes like homogenization, exposure to interfaces, and shear forces), hydrogel preparation procedures are beneficial to preserve the protein stability as mild conditions such as aqueous environment and room temperature are commonly employed. Their unique properties have created increasing interest in developing the CR systems for proteins/peptides to maintain therapeutic plasma concentrations in the surrounding tissues or in circulation for longer time. Hydrogels can be prepared from natural as well as synthetic polymers. Chemically cross-linked networks have permanent junctions, while physical networks have transient junctions that arise from either polymer chain entanglements or physical interactions such as ionic interactions, H-bonds or hydrophobic interactions. The physical appearance as matrix, film or microsphere depends on the polymerization method used to prepare hydrogels. Hydrogel networks are also based on the network electrical charge, called nonionic (neutral), ionic (including anionic or cationic), and amphoteric electrolyte (ampholytic) containing both acidic and basic groups. Hydrogel-forming natural polymers include proteins such as collagen, gelatin and polysaccharides like starch, sodium alginate (NaAlg), chitosan (CS), and agarose. Hydrogels based on homopolymers consists of a single monomer with a cross-linked skeletal structure, while copolymer-based hydrogels are formed from two or more different types of monomers with at least one hydrophilic component arranged in a random, block or alternating configuration along the polymer backbone [16]. On the other hand, IPNs are made from two independently cross-linked synthetic and/or natural polymers in a network, while in a semi-IPN hydrogel, one polymer component is cross-linked, and the other is not. Various such structures are depicted in Fig. 1.

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The stimuli-responsive hydrogels that can undergo volume transitions in response to various physical stimuli such as temperature, electric or magnetic field, light, pressure, and sound, as well as chemical stimuli like pH, solvent composition, ionic strength, and molecular species, are widely employed in drug delivery area [12,17]. Upon removing the external stimuli, swollen hydrogel contracts to the unswollen state. Stimuli-responsive hydrogels offer remarkable prospects for the delivery of macromolecules and genes as the carriers are active contributors to optimize the therapy instead of a passive delivery vehicle. Recent interest in cell mechanics and effects of substrate elasticity on cell structure as well as its function together with the ability of synthesizing the novel polymers that approximate the material property of biological tissues has motivated research on different materials for use in wound healing and tissue engineering [18]. Synthetic gels prepared from polyisocyanopeptides grafted with oligo(ethylene glycol) side chains reportedly mimic gels prepared from intermediate filaments in almost all aspects. These responsive polymers have a stiff and helical architecture with a tunable thermal transition where the chains bundle together to generate transparent gels at extremely low concentrations. These materials show a very fast sol–gel phase transition that can be completely reversible. However, the ease of modification of these materials provides avenues for the preparation of functional biomimetic materials required in biomedical applications [19].

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Naturally occurring polysaccharides are frequently used in the delivery of macromolecular therapeutics as these are highly biodegradable and biocompatible, and can be prepared as conjugates or complexes with proteins, peptides and other biomacromolecules. Among the widely investigated polysaccharides, NaAlg, chondroitin sulfate, CS, and hyaluronic acid (HA) are the prime candidates. Specifically, these polysaccharides in combination with other polymers offer the desired chemical and/or biological advantages. Some representative members of this class are discussed briefly here, but details can be found elsewhere [20].

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CS, a copolymer of glucosamine and N-acetylglucosamine, has been widely used in drug delivery area [21]; it is a nontoxic, mucoadhesive, biodegradable, non-allergic and easily absorbable polymer, whose properties can be tailored to suit to specific applications in the form micro and nanoparticles or hydrogels [2,8,22–25]. Innumerable studies have been reported on colon specificity of CS [26,27], but its intestinal delivery to colon is insufficient because of its deswelling nature in alkaline media. In this pursuit, polyelectrolyte complex of CS (such as CS-pectin and CS-NaAlg) with water-soluble polyionic species that are swollen in neutral pH was developed [28–30]. High insulin association efficiency of 81% was reported for CS-NaAlg NPs (size, 850 nm)

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prospects of this emerging field of micro/nanotechnology will be the focus of our discussion that surrounds the delivery of macromolecular therapeutics, with a special emphasis on biopolymeric hydrogels.

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Fig. 1. Types of hydrogel network structures used in macromolecular drug delivery.


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Dextran (DEX) is a complex polysaccharide consisting of α-1,6linked D-glucopyranoses with some degree of 1,3-branching. Being hydrophilic, it can be crosslinked by various methods to be effective as a hydrogel carrier system [37]. Van Tomme et al. [38] published an extensive review on DEX-based hydrogels for protein delivery. Peptide crosslinking resulted in enzyme-dependent degradation controlled by cellsecreting enzymes, thus mimicking the degradation of the natural extra cellular matrix [39]. Taking these advantages, self-assembled NPs of quaternized CS (N-(2-hydroxyl) propyl-3-trimethyl ammonium CS chloride (HTCC)) and DEX sulfate were prepared by ionic-gelation that showed internalization into Caco-2 cells without loss of cell viability. This system demonstrated fast release of therapeutics in phosphate buffer solution (pH 7.4), but with a slow release in acidic media (pH 1.4) [40]. A novel method was reported [41] to prepare uniform spherical bovine serum albumin (BSA)-Zn2+ NPs of 50–350 nm size under mild condition by adding BSA and zinc acetate to a solution of PEG followed by freeze-drying. Here, no protein aggregation was observed and due to freeze-induced phase separation, the BSA–Zn2 + complex was squeezed into dispersed particle giving spherical NPs. The same protocol was followed to prepare recombinant human growth hormone (rhGH)–Zn2+− DEX NPs [42] by vortexing a mixture of rhGH and zinc acetate added to DEX-PEG solution followed by freeze-drying for overnight at −20 °C. The lyophilized powder was washed with dichloromethane and centrifuged to remove PEG to obtain NPs of size 20– 170 nm with N90% encapsulation efficiency (EE) and these showed Nb2-11 cell proliferation activity. An almost identical procedure was adopted for encapsulating granulocyte-macrophage colony-stimulating factor (GM-CSF), granulocyte colony-stimulating factor (G-CSF), β-galactosidase, myoglobin (MYO), and BSA into DEX NPs by aqueous–aqueous freezing-induced phase separation [43] as shown in Fig. 2 to produce NPs of 200– 500 nm size for encapsulating BSA that showed N98% EE with no protein aggregation or loss of bioactivity. Oral targeted insulin NPs (150–300 nm size) using vitamins B12 (VB12) were reported by Chalasani et al. [44]. The authors used aminoalkyl VB12 derivatives synthesized at 5′-hydroxy ribose and e-propionamide sites via carbamate and ester/amide linkages coupled to succinic acid modified DEX NPs with varying extent of cross-linking that showed 70–75% reduction in blood glucose level (BGL) to maintain anti-diabetic effects up to 54 h in STZ-induced diabetic rats. NPs with low levels of cross-linking were superior and more effective with VB12 derivatives of carbamate linkage. The pharmacological availability relative to sc insulin was around 29% that was superior (1.5-fold) to NPs conjugate of ester linked VB12. The NPs demonstrated a similar oral insulin efficacy in congenital diabetic mice (60% BGL reduction in 20 h). The microcapsules of thermo-responsive poly(N-isopropylacrylamide) (PNIPAAm) were prepared [45] and used for the CR of stromal cell-

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produced by an ionotropic gelation method with no major conformational changes of insulin in terms of α-helix and β-sheet content [31]. Polyelectrolyte complexes of CS with other biopolymers [30,32], folic acid-conjugated NPs of CS for targeted delivery of 5-aminolevulinic acid (5-ALA) to colon [33] as well as CS-tamarind kernel powder interpolymer complex films have been used for colon delivery [34]. Chemical cross-linking of CS with polyethylene glycol (PEG) produced a swellable delivery system in both acidic and alkaline pH media [35], but crosslinking of CS with glutaraldehyde (GA) has the disadvantage of toxicity. Alternatively, genipin, a natural aglycone, was found to have minimum toxic effects [36]. Among the many CR systems of CS, pHsensitive CS-Eudragit L100-55 NPs prepared by a coacervation method using high viscosity HPMC was useful for the delivery of insulin through gut mucosa. See recent reviews on CS hydrogel delivery systems [2,8].

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Fig. 2. Formulation strategy used for preparing protein-loaded dextran NPs. Reprinted from Ref. [43], © 2014, with permission from Creative Commons Attribution 2.0 Generic.

derived factor (SDF)-1α, an important chemokine for stem cell recruitment/homing. Here, the double-phase emulsified condensation polymerization was used to prepare the interconnected porous glycidyl methacrylated dextran (Dex-GMA)/gelatin microcapsules by adopting plasma-graft pore-filling polymerization to graft PNIPAAm onto the surface pores of microcapsules. The in vitro results of these microcapsules exhibited thermo-responsive release of drug in response to temperature variations. The sc implantation of these microcapsules resulted in a sustained and long-term release of SDF-1α compared to those formulations without PNIPAAm-grafting. A new multicomponent delivery system based on DEX, protamine (Prot), and solid lipid nanoparticles (SLN) containing pCMS-EGFP plasmid were investigated for clathrin/caveolae dependant transfection efficiency [46]. These surface modified SLNs were tested for erythrocyte interaction and potential agglutination. The insulin-loaded NPs of 500 nm produced from complexation of DEX and CS [47] showed good stability with zeta potential value of −15 mV under optimal composition of DS:CS (mass ratio of 1.5:1) at pH 4.8 with almost no insulin release at pH b 5.2 up to 24 h, but insulin release occurred in pH 6.8, suitable for oral delivery. In a continuing study, the authors [48] used these systems to deliver insulin in a pH dependent manner to achieve association efficiency of 70% with 24-h sustained release in rat models with a bioavailability of 5.6% at 50 IU/kg dose compared to insulin solution. In another study [49], cationic character was imparted to DEX by conjugating spermine to oxidized DEX by reductive amination. Selfassembly of cationized DEX and CXCR4-siRNAs produced 55 nm size NPs with a zeta potential of 40 mV, which significantly downregulated CXCR4 expression (tested in colorectal cancer metastasis in Balb/c mice) through CXCR4 silencing. .


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NaAlg is an unbranched polysaccharide consisting of 1-4′-linked β- D -mannuronic acid (M) and α-L -guluronic acid (G) moieties in varying compositions [59], which is capable of forming strong hydrogels by simple addition of metal ions to its aqueous solution. Such a mild method of preparation makes the hydrogels very suitable to encapsulate living cells for investigating their CR properties. NaAlg hydrogels were also used for tissue regeneration, while simultaneously releasing growth factors or cytokines [60,61]. Some representative novel micro/nanocarriers of NaAlg will be discussed. NaAlg-based NPs developed as pH-responsive delivery vehicles for insulin prepared by spray-drying, ionic crosslinking by electrohydrodynamic spraying and solvent diffusion methods have shown the EE values ranging from 38 to 90% [62,63]. Their pH dependent CR properties protected the integrity of insulin at higher temperatures during spray-drying process. Sharma et al. [64] developed biodegradable PVA and NaAlg electrospun composite nanofiber-based transmucosal patches for sublingual insulin delivery. Their high water holding capacity resulted in high insulin loading that improved the mucoadhesivity with EE of 99%. NaAlg-based microcapsules added with poly-L-lysine and poly-L-ornithine crosslinked under UV radiation were also used for connexin-43 carboxyl-terminus mimetic peptide (CT1) delivery to the wound site. A study [65] on NaAlg-based micro-devices for BSA (66 kDa) showed the dependence of release on MW and particle size distribution. Mahidhara et al. [66] developed novel NaAlg-CS coated ceramic nano-carrier loaded with multi-functional anti-cancer bovine lactoferrin (Lf), a natural milk based protein, to improve its intestinal absorption. These systems showed the size dependent endocytosis and transcytosis of NPs in colon cancer cell-lines.

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Hydrogel-based micro/nanoparticles are promising for macromolecular drug delivery as these show particular advantages in oral delivery. In oral vaccine delivery, these devices, if prepared in the size range of 1– 10 μm, are more suitable than the nanocarriers for proper cell mediated immune responses and high drug loading. As noted before, such micro/ nanohydrogel systems can be prepared by many different techniques including solvent evaporation, spontaneous emulsification/solvent diffusion, salting out/emulsification-diffusion, supercritical fluid technology, spray-drying, ionic gelation, micelle, and reverse micelle formation [78]. Self-assembled nanohydrogels are particularly attractive, since these are easy to prepare, are affordable and can effectively incorporate biopharmaceuticals like biosimilars, proteins and peptides. The release of cargos from such systems can be fine-tuned by tailoring

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Hyaluronic acid (HA) is a naturally occurring anionic linear glycosaminoglycan made up of repeating disaccharide units of D-glucuronic acid and N-acetylglucosamine having MW up to 107 Da. It is a major component of synovial fluid, and is present primarily in extracellular matrix (ECM) due to its viscoelastic properties that helps to reduce the friction between bones [50]. Delivery of proteins and peptides has been attempted with photopolymerized HA and HA-tyramine conjugates cross-linked by disulfide bond formation [51–53]. The negative charge of HA hinders protein release rate due to the charge of protein, but enzymatic degradation of HA by hyaluronidase alleviates this problem. Even though there are no commercially available delivery products using HA, except Declage, it is considered to have a greater potential to develop as conjugate or even as physically or chemically crosslinked hydrogel depot system [54]. Microparticle formulation using Hyaff11p50, where 50% of carboxyl groups of HA are esterified with benzyl alcohol, was prepared by solvent evaporation and spray-drying method for a 7-day release of hGH [55]. Another formulation of hGH/ HA/lecithin prepared by spray-drying method resulted in small particles to be injected through a 26-gauge needle, and showed improved patient compliance [56]. Novel hybrid hyaluronan-based nanogels (average size b160 nm) [57] were suggested for spontaneous encapsulation of proteins and peptides that showed the CR profiles without altering chemical stability of macromolecular therapeutics. HA and ferric oxide were ionically gelated to form hybrid nanoparticles for peptide delivery. Hybrid materials such as cyclodextrin (used as porogen) coupled porous microparticles of HA ionically complexed with lysozyme were also developed for protein delivery that showed improved EE and protective nature [58]. Thus, HA-based systems appear to be more suitable than poly(lactic acid-co-glycolic acid) (PLGA), as it is more biocompatible and hydrophilic in nature

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Recently, affinity-based growth factor delivery system was developed by incorporating heparin into photocrosslinkable NaAlg hydrogels (HP-NaAlg) that prolonged the release of therapeutic proteins [67]. Heparin modification showed minimal biodegradation and the release of growth factors sustained up to 3 weeks with no initial burst release. On the other hand, implantation of bone morphogenetic protein-2 (BMP-2)-loaded HP-NaAlg hydrogels induced a moderate bone formation around the implant periphery. Importantly, BMP-2-loaded HPNaAlg hydrogels induced significantly more osteogenesis than BMP-2-loaded unmodified HP-NaAlg hydrogels with 1.9-fold greater peripheral bone formation and 1.3-fold greater calcium content in BMP2-loaded HP-NaAlg hydrogels than the BMP-2-loaded unmodified HP-NaAlg hydrogels even after 8 weeks of implantation. El-Sherbiny et al. [68] developed water-soluble copolymer of sodium acrylate grafted onto carboxymethyl cellulose (CMC). When these were used along with NaAlg as a pH-sensitive IPN hydrogel prepared by ionotropic gelation in the presence of Ca+2 showed b 18% and 68% after 8 h in pH 7.4 for BSA in 2 h in pH 2 media. On the other hand, NaAlg-CS microspheres prepared [69] by membrane emulsification technique in the presence of Ca+2 and CS solidification achieved an EE of 57% and their in vitro release under blood pH showed stable and sustained release up to 14 days and these formulations protected the chemical stability of insulin released. In vivo studies showed a reduction of BGL of diabetic rats stable up to 60 h after oral administration of insulin-loaded microspheres. However, pH-responsive, biodegradable microparticle carrier system based on ionotropically cross-linked mixture of NaAlg and chemically modified carboxymethyl CS coated [70] through polyelectrolyte complexation with CS grafted PEG showed the release ranging from 32 to 83.1%. Many types of modified NaAlg matrices have [71] been mixed with other polymers to obtain either IPNs or covalently linked mixed polymer systems to improve the stability of hydrogels along with the CR of encapsulated proteins [72]. An elegant example is simple sulfonation of uronic acids in NaAlg that specifically binds heparin-binding proteins, including growth factors to allow the CR of these proteins and thus provide enhanced therapeutic activity as demonstrated in a murine hindlimb ischemia model in rats [73]. Goycoolea et al. [74] developed the NPs for transmucosal delivery of macromolecules by ionic gelation of CS hydrochloride with pentasodium tripolyphosphate (TPP), complexed with NaAlg that exhibited 41–52% of EE. These showed enhanced systemic absorption of insulin after nasal administration to conscious rabbits. However, some novel composite NaAlg/PLGA microparticles [75] could deliver bovine insulin exhibiting reproducible EE values for insulin, releasing up to 4 months, while NaAlg nanohydrogels loaded with insulin prepared [76] by a reverse emulsification-diffusion method showed the CR profile for tissue or organ targeting for DNA/gene delivery. The NPs of PLGA [77] imparted pH sensitivity when prepared using hypromellose phthalate by a modified emulsion solvent diffusion method and showed relative bioavailability of 7%.

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Despite the significant advances in macromolecular drug delivery through different approaches, oral route continues to be the most extensively studied approach for therapeutic delivery mainly due to the convenience and patient compliance [85]. The large MW, 3D structure and high aqueous solubility of macromolecules pose unique challenges for absorption through transmucosal route. A number of delivery systems including pH-sensitive hydrogels and intestinal M cell targeted devices that utilize paracellular and transcellular absorption pathways have been developed for protein/peptide drugs in order to prevent the deleterious effect of GIT [86]. However, restricted and delayed absorption as

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Transmucosal delivery of vaccines using hydrogel-based nanocarrier systems have been explored due to their advantages such as efficient encapsulation, protecting the drug from pH and enzymatic degradation, mucosal adjuvant incorporation to enhance immune response and targeted delivery to mucosal inductive sites [91–93]. Even though adverse pH and enzymatic conditions of GIT pose problems for oral delivery of vaccines, these systems are quite useful, since they are prepared by spray-drying, double emulsion, ionic gelation/polyelectrolyte complexation, and phase separation methods that are more suitable for vaccine delivery as these methods protect the 3D structure of vaccines and antigens. Physicochemical properties of hydrogel-based nanocarriers are also critical for effective vaccine delivery. For orally administered NPs, a mean size of around 500 nm showed better uptake, since the M-cell membrane is negatively charged, so a positive zeta potential is beneficial for M-cell transport [93]. The most widely used hydrogel-based systems are chitosan, dextran, alginate, and such other polymers. Of these, CS has inherent mucoadhesive property due to ionic interactions between its quaternary amino (NH3+) and negative sialic acid group of mucin; thus, it can open the tight junctions of intestinal epithelium. Various derivatives of CS (e.g. trimethyl CS (TMCS), carboxymethyl CS (CMCS), and thiolated CS) have been prepared to impart more mucoadhesivity and higher penetration enhancement capability [94]. Hybrid materials such as HA (polysaccharide) and oligosaccharide (e.g. cyclodextrin and cyclodextrin derivatives)-based NPs showed better biocompatibility, in vivo stability and pharmacological efficacy. Various self-hydrolyzing hydrogels based on Michael-type addition of PEG-diester-dithiol or non-degradable PEGdithiol cross-linkers onto 4-arm PEG-vinyl sulfone have also been used as tunable delivery systems for antigens like immunoglobulins. These systems had a gel mesh size of 13–35 nm and showed controlled diffusivity based on factors such as molecular weight of drug, polymer and chemical structure of the cross-linker. The CS and CS-based NPs [95–102] showed both systemic and mucosal immune responses when delivered through the nasal route, but mucoadhesivity of CS-based NPs leads to prolonged residence time and increased M cell uptake. Alginate coated CS and TMCS NPs were developed for nasal [103,104] and oral [104–106] delivery to

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well as variable intestinal motility and gastric emptying time would create a major hurdle to achieve the desired pharmacodynamic efficacy. Among other routes of administration, pulmonary [87], nasal [88], and buccal [89] routes have been suggested for macromolecular drug delivery. However, hydrogel-based parenteral delivery of macromolecular therapeutics have some unique advantages such as controlled release, depot therapy, biocompatible, biodegradable and targeting capability using ligands, aptamers, etc. The drawbacks such as decreased patient compliance, temperature sensitiveness and tissue damage caused by parenteral therapy pose problems for chronic therapy [90].

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cross-link density of the matrix. Other strategies to tailor drug release from hydrogels rely upon reversible protein–polymer interaction or encapsulation of protein in a second delivery system (e.g., micro/ nanoparticles and liposomes) dispersed in a hydrogel matrix [79]. The raspberry-like nanohydrogels (A-CHPNG) with a narrow size distribution (40–120 nm size) [80] were employed in protein (interleukin-12 (IL-12)) delivery by cross-linking acrylate group-modified cholesterol-bearing pullulan nanohydrogel (CHPANG) with thiol group-modified PEG (PEGSH) through Michael addition (see Fig. 3). These nanohydrogels showed the EE value of 96% for IL-12, with a steady plasma IL-12 level in mice up to 72 h following the sc administration. Surface characteristics of the NPs play a predominant role in their uptake by liver, spleen and RES. The NPs containing more hydrophobic surfaces are preferentially taken up by liver, followed by spleen and lungs [78,81]. On the other hand, hydrophilic NPs (35 nm) prepared from poly(vinyl pyrrolidone) (PVP) displayed only b1% uptake by spleen and liver and about 5–10% of these could circulate in bloodstream even after 8 h of iv injection; however, the NPs (45–126 nm) prepared using 50% vinyl pyrrolidone and 50% N-isopropyl acrylamide showed preferential uptake by the liver [82]. Ligand-mediated active targeting has emerged as a novel strategy to support the effectiveness of nanocarriers to improve the delivery of therapeutics. Among different receptors expressed over gastrointestinal tract (GIT), folic acid (FA) receptors were reported to be present in sufficient quantity to improve the uptake and transport of bioactive or vesicular systems across the GIT [83]. These FA receptors are also over-expressed in various tumor types and thereby target to various tumors [84]. Aptamers that are oligonucleic acids or peptides are also used to deliver macromolecular drugs (siRNA, proteins and peptides), but these are used in conjunction with ligand-based target molecules such as transferrin and FA to acquire better therapeutic efficacy. Hydrogel-based nanocarriers of CS, due to their positive charges, are specifically suitable for ligand/aptamer appending as well as encapsulating negatively charged siRNA that are critical to increase the cargo loading and EE.

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Fig. 3. Schematics of the formation of cholesterol-bearing pullulan nanohydrogel structure (A-CHPNG) assemblies. Reprinted from Ref. [80], © 2014, with permission from Elsevier.


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While delivering the hormone, it becomes necessary to ensure safety and therapeutic efficacy to reduce the frequency of administration and to enhance the patient compliance and comfort. Various novel delivery systems have been explored for human growth hormone (hGH) delivery. Even though hGH itself is secreted in a pulsatile manner, continuous infusion of hGH via pump may not provide the required clinical efficacy as it elevates insulin-like growth factor-1 (IGF-1) levels similar to that of daily injections, indicating that pulsatile hGH delivery is not suitable [108]. However, the most commonly used formulations for hGH are oral tablets and transdermal patches for sex steroid hormones as well as intramuscular or sc injections [109]. Other approaches like crystal formulation, PEGylation, and loading of hydrogels have also been tried for the CR of hGH [110,111]. Recently, CriticalSorb, a proprietary absorption promoter used in nasal delivery, was investigated for investigating hGH absorption efficacy in conscious rat model. Here, absorption-enhancing component is Solutol HS15, which consists of polyglycol mono- and di-esters of 12-hydroxystearic acid combined with free PEG, which showed 49% bioavailability in 2 h for the ratio of Solutol HS15 to hGH of 4:1 (mg/mg) at which no nasal mucosa toxicity was found even up to 6 months [112]. Even though PLGA-based microsphere formulation, viz., Nutropin depot, was the first marketed CR product of hGH, its high manufacturing costs led to its withdrawal from the market. This system showed sustained release of hGH up to 1 month [113], but suffered from low EE, high initial burst release, protein aggregation, denaturation, and

505 506

O

509 510

504

R O

7. Delivery of hormones

502 503

P

508

500 501

inflammation due to the use of organic solvents. Moreover, acidic degradation products [114] were observed with this system. To overcome these drawbacks, zinc was incorporated into these microspheres to stabilize hGH [108,111], since Zn+ 2 would induce dimerization of hGH, and hence, the reversible complex would be more stable than monomeric hGH. However, irreversible aggregation of hGH in the PLGA microsphere was effectively reduced by the zinc–hGH complex and the CR of hGH was observed compared to daily injection. The only commercially available CR formulation of hGH is hyaluronate-based microparticle formulation launched as a once-aweek injection formulation (Declage) in Korea by LG Life Science in 2007. This system showed improved drug loading (ca. 20%) and bioavailability compared to PLGA microspheres, which maintained serum hGH level for 30 h in cynomolgus monkeys [56,115]. However, injectable, biodegradable, and thermosensitive hydrogel-based micro/nanocarriers seem to be the better choices for hGH delivery, since their high water content and temperature dependent gelation properties without the use of organic solvents or chemical cross-linkers made these systems as efficient hGH delivery carriers [116]. Park et al. [117] developed dual ionic interaction nanohydrogel system composed of a positively charged polyelectrolyte complex (PEC) containing hGH and anionic thermosensitive poly(organophosphazene) hydrogel to enhance the CR of hGH (see Fig. 4). These nanodevices of 500 nm size and zeta potential of +8 mV suppressed the initial burst release showing better in vitro and in vivo correlation with 13-fold increase in AUC compared to pristine hGH solution and enhanced bioavailability in hypophysectomized rat model. Pyo et al. [118] used an entirely different approach, i.e., solution enhanced dispersion by supercritical fluids (SEDS), to produce nano-sized recombinant hGH using ethanol to help supercritical CO2 to extract water from the aqueous protein solution. Various sizes of hGH NPs were prepared by this method with a narrow particle size distribution from the aqueous ethanol solution without using any additive. However, these novel delivery systems require rigorous in vivo pre-clinical, clinical safety and efficacy testing data prior to further development as useful therapeutic products. Tang et al. [119] used an emulsion cross-linking technique to prepare CMCS and HA conjugate hydrogel microspheres for delivering

T

507

impart pH responsiveness as well as to control burst release of antigen from the NPs. For instance, nasal delivery of alginate MPs with tetanus toxoid [107] showed a strong systemic and mucosal immune response. Despite the promising in vivo results of various NPs for vaccine delivery, their clinical uses are more challenging. The crossing of mucosal barrier and those related to mucosal immune stimulation as well as their correlation with NP characteristics could result in improved NP delivery. However, targeted delivery approach to cellular and subcellular specific delivery would be more beneficial as it leads to an optimum therapeutic NP system for vaccine delivery.

D

498 499


E

6

Fig. 4. Dual ionic interaction model based on hGH loaded polyelectrolyte complex and anionic hydrogel for sustained delivery of hGH. Reprinted from Ref. [117], © 2014, with permission from Elsevier.


534 535 536 537 538 539 540 541 542 543 544 545 546 547 548 549 550 551 552 553 554 555 556 557 558 559 560 561 562 563 564 565 566 567 568 569 570 571


593 594 595 596 597 598 599 600 601 602 603 604 605 606 607 608 609 610 611 612 613 614 615 616 617 618 619 620 621 622 623 624 625 626 627 628 629 630 631 632 633 634 635 636 637

t1:3

PEC hGH-PS in poly(organophosphazene) Recombinant hGH-NP by SEDS method HA-carboxymethyl CS conjugate PLGA-CS LEC-CS LEC-CS NP transdermal suspension CS NPs CS-coated tripalmitin NPs PEG-coated tripalmitin NPs PEG-coated tripalmitin/Miglyol NPs HA–CS nanocomplexes CS NP (spray-dried)

500/+8

hGH

[117]

t1:4

50

hGH

[118]

t1:5

4100–5900/−17 to −24 430–590 122–347/8–33 113−129/−9 to +13 267/+35 538/+29 226/−35 207/−37 163–193/−32 to −75 215/+28

Catalase sCT Melatonin Melatonin sCT sCT sCT sCT sCT sCT

[119] [122] [123] [124] [125] [126] [126] [126] [127] [128]

t1:6 t1:7 t1:8 t1:9 t1:10 t1:11 t1:12 t1:13 t1:14 t1:15

O

F

Ref.

hGH, human growth hormone; PLGA, poly(lactic-co-glycolic acid); CS, chitosan; PEG, polyethylene glycol; NPs, nanoparticles; sCT, salmon calcitonin; HA, hyaluronic acid; PEC, polyelectrolyte complex; SEDS, solution-enhanced dispersion by supercritical fluids; LEC, lecithin; PS, protamine sulfate.

R O

591 592

Drug

t1:16 t1:17 t1:18 t1:19

8. Gene delivery

638

Efficient in vivo delivery of genes remains the biggest challenge of nucleic acid-based biopharmaceuticals, mainly because of their large molecular size, hydrophilicity, negative charge, acid labile nature and ease of degradation by nuclease, preventing an effective transport across the intestinal epithelium causing low transfection efficiency via oral route. As a result, parenteral route has been preferred for delivering nucleic acids using non-viral vectors containing polymers, complexes of cationic lipids or peptides with plasmid DNA [129]. Cationic polymerbased non-viral vectors have been commonly used as carriers for nucleic material delivery [2], and self-assembly of cationic polymer with DNA (polyplex) that interacts with negatively charged cellular membrane was also employed [130]. Polyplexes containing macromolecular drugs must escape from endosomes for successful transport of nucleic material, since DNA is prone to degradation by lysosomal enzymes. After endosomal escape, polyplexes get into cytoplasm where they unpack DNA and deliver it either near the nucleus or in the nucleus. The translocated DNA in the nucleus then causes gene expression [131]. To achieve this, the chosen carrier polymer should have a good DNA binding capability such that it can condense DNA into polyplexes and high buffering capacity to enable endosomal escape as well as good intracellular vector unpacking to release DNA. In this regard, CS can effectively bind with DNA, protecting it from nuclease degradation and these polyplexes are capable of adhering to intestinal epithelium and M cells targeted transport across the mucosal boundary that can transfect epithelial and immune cells associated with the gut lymphoid tissue [132]. Among others, CS has a lower cytotoxicity than the commonly used synthetic polyethyleneimine (PEI) [133] and its endosomal disruptive property, as well as proton sponge-type mechanism, ensures entry into the cell. Upon internalization within the endosome and due to the change in pH inside the cell, CS gets protonated and eventually causes the rupture of the membrane to release of the genetic material into the cell [2]. CS derivatives such as glycol, o-carboxymethyl, trimethylated, thiolated and 6-N,N,N-trimethyltriazole CS and its hydrophobic modifications like deoxycholic acid, 5-β-cholanic acid, and N-acylated chitosan have all been used to overcome the limited CS solubility to improve transfection efficiency. A recent study on CS-based NPs for nucleic acid delivery showed that transfection efficiency depends on MW of CS, its deacetylation degree, complex formulation, and pH of the environment [134]. Chen et al. [132] studied oral gene therapy using CS/DNA NPs carrying murine erythropoietin (Epo) gene. Over 25% increase in hematocrit levels was observed by delivering Epo genes (responsible for stimulating red

639 640

P

589 590

Size (nm)/zeta potential (mV)

D

587 588

System

E

585 586

t1:1 t1:2

Table 1 Micro/ nanohydrogels for hormone delivery.

T

583 584

C

581 582

E

579 580

R

578

R

576 577

N C O

574 575

antioxidant enzyme catalase through oral route to prevent enzymatic degradation. Cross-linking reduced the swelling to increase the resistance to hyaluronidase digestion of these systems. The encapsulated catalase exhibited superior stability over wide pH ranges (pH 2, 6 and 8) compared to the native enzyme. These catalaseloaded microspheres, in contrast to native catalase, effectively decreased the intracellular H2O2 level and protected the HT-29 colonic epithelial cells against H2O2-induced oxidative damage to preserve the cell viability. Salmon calcitonin (sCT), a therapeutic analogue of calcitonin, a cyclic polypeptide hormone, is another important drug used to prevent osteoclastic bone resorption (potent hypocalcemic). It is mainly available as injection and nasal spray formulation in the market, but these have met with several disadvantages. The oral bioavailability of sCT is b0.1% due to extensive proteolytic degradation in the GIT and poor permeation across intestinal epithelial cells [120]. The hydrogel-based NPs using thiomer derivative of glycol chitosan (GCS) synthesized by coupling with thioglycolic acid (TGA) have been evaluated for pulmonary delivery of peptides [121]. This study reported the NPs (200–300 nm) prepared from GCS and GCS-TGA by ionic gelation with a net positive surface charge that have shown high calcitonin encapsulation. On the other hand, CS surface modified sCT loaded PLGA NPs (430–590 nm) prepared [122] by w/o/w emulsification and solid dipping methods showed the EE of 50% with improved sustained release, showing short-period hypocalcemic effect. The mucoadhesive NPs [123] prepared from lecithin/CS investigated for transmucosal delivery of melatonin showed no cell membrane damage and were non-cytotoxic. Transdermal delivery of melatonin was also studied [124] using the NPs with size differing in lecithin type (Lipoid S45 and S100) and CS content that ranged from 114 to 332 nm with zeta potential of 4.6–31.2 mV. The study indicated 7.2% of melatonin loading with 1.3- to 2.3-fold higher flux compared to melatonin solution; the highest flux of 9.0 μg/cm2/h was achieved with lecithin/CS for a weight ratio of 20:1 and these NPs were safe to use at concentration up to 200 g/mL for skin application. A similar study by Prego et al. [125] highlighted the concentration dependent improved cellular uptake and internalization of NPs loaded with sCT. The in vivo data suggested that both CS-coated systems (neat NPs and CS-coated NPs) showed enhanced intestinal absorption of sCT, thus enhancing the therapeutic efficacy in rat model. More efforts have been made on developing the CR formulations of sCT, of which sCT loaded solid lipid NPs coated with PEG or CS were investigated to study the effect of composition of core shell on the delivery of sCT [126]; according to this study, the nature of coating affected the surface association of the peptide. The NPs covered by positively charged CS reduced the burst release that was more pronounced for NPs coated with PEG than those coated with CS. After the initial burst release, these systems showed continuous and slow release of peptide, independently of the nature of coating. The slow release was due to affinity of peptide for lipids and the absence of degradation of lipid matrix under in vitro release conditions. The intra-articular sCT-HA sustained release system using CS as electrostatic-based nanocomplex reduced the inflammation and inflammatory gene expression when injected through intraarticular route in a mice model [127] that showed a remarkable anti-inflammatory effect by reducing the NR4A2 mRNA expression in vitro. However, inhalable co-spray dried powders of sCT-loaded CS nanoparticles (sCT-CS NPs) of 200 nm size prepared with mannitol showed improved pulmonary absorption in rats [128]; the NPs were formed spontaneously after adding sCT in CS solution containing TPP. For inhalation, NPs were co-spray dried with mannitol as a carrier, which displayed better aerodynamic properties for pulmonary delivery, showing higher protein absorption than native sCT. Table 1 summarizes various micro/nanoparticle-based platforms used in hormone delivery.

U

572 573

7


641 642 643 644 645 646 647 648 649 650 651 652 653 654 655 656 657 658 659 660 661 662 663 664 665 666 667 668 669 670 671 672 673 674 675 676 677 678 679 680 681

705 706 707 708 709 710 711 712 713 714 715 716 717 718 719 720 721 722 723 724 725

F

O

R O

703 704

P

701 702

D

699 700

T

697 698

C

695 696

E

693 694

R

691 692

R

689 690

O

688

C

686 687

Recently, siRNA has emerged as a powerful therapeutic agent for the treatment of diseases as well as repair of faulty genes, resulting in the production of faulty proteins [2,5,8]. The siRNAs are more stable to nuclease degradation than unmodified antisense oligonucleotides, since they are highly sequence specific and are required in relatively small doses, but their poor serum stability in systemic circulation and poor cellular uptake makes the therapy a big challenge. Moreover, naked siRNAs are quickly excreted out of the kidney as their biological half life is b1 h, but their size, negative charge and inability to cross cellular membrane in unmodified form make gene knock-down all the more difficult [5]. Various synthetic (poly(l-lysine) or polyethylenimine (PEI) and their analogs) and natural polymers (CS, collagen, gelatin and their derivatives) have been used for their delivery. Mimi et al. [141] prepared PEI-based nanohydrogels with gelatin core (200 nm) to encapsulate siRNA by a two stage process involving the preparation of gelatin NPs followed by conjugation with branched PEI to impart cationic property (see Fig. 5). These NPs showed increase in transfection efficiency to HeLa cells from 41 to 84%. The delivered siRNA inhibited 70% of human argininosuccinate synthetase 1 (ASS1) gene expression. Another study [142] involving cationic DEX nanohydrogels for encapsulating siRNA and photochemical internalization has shown the potential for gene silencing using intracellular vesicles as depots for siRNA. This technique destabilized the endosomal vesicles to prolong the knockdown of target protein. According to Salva et al. [143], local delivery of CS/vascular endothelial growth factor gene (VEGF)siRNA nanoplex in a rat breast cancer model using CS (75 kDa, 75–85% DDA)-VEGF siRNA nanocomplex remarkably suppressed the VEGF expression and tumor volume. Crosslinked CS NPs of b150 nm for encapsulating siRNA were prepared by ionic crosslinking (CS to siRNA mass ratios of 10:1, 30:1 and 50:1) to explore their potential to deliver siRNA to lungs via a jet nebulizer [144] to observe high EE and non-aggregation at the pH of the airway. Complete binding of siRNA to CS was possible at a ratio 50:1 and high cell viability (N85%) was observed even at the highest CS concentration of 83 μg/mL. In another study, interpolyelectrolyte complexes of CS/siRNA showed rapid uptake (1 h) of Cy5-labeled NPs into NIH 3T3 cells, followed by accumulation for over 24 h [145]. These showed knockdown of the endogenous enhanced green fluorescent protein (EGFP) in both H1299 human lung carcinoma cells and murine peritoneal macrophages (77.9% and 89.3% reduction in EGFP fluorescence, respectively). Nasal administration of these NPs resulted in effective in vivo RNA interference in bronchiole epithelial cells of transgenic EGFP mice (37% and 43% reduction compared to mismatch and untreated control, respectively).

N

684 685

blood cell production) to intestinal epithelium. Bowman et al. [135] used the NPs of CS as a gene carrier for oral delivery of factor VIII DNA in a mouse model with hemophilia A (a disorder of the blood coagulation cascade caused by defective factor VIII). These NPs led to higher plasmid copy numbers in Peyer's patch tissue compared to the naked DNA delivery. Zheng et al. [136] studied the in vitro and in vivo transfection efficiency of three types of CS NPs, i.e., quaternized CS-60% trimethylated CS oligomer (TMCSO-60%), CS (43–45 kDa, 87%), and CS (230 kDa, 90%) for encapsulating pDNA encoding green fluorescent protein (GFP) prepared by complex coacervation technique. In vitro results of these indicated the highest transfection efficiency for NPs of TMCSO60% followed by CS (43–45 kDa, 87%) and CS (230 kDa, 90%). In vivo studies indicated the most prominent GPF expression in gastric and upper intestinal mucosa. TMCSO-60% was the most efficient with better activity and minimal toxicity making it an efficient carrier for oral delivery at optimal CS/pDNA ratio of 3.2:1. The pullulan NPs (45 nm) encapsulated with nucleic acid (pBUDLacZ plasmid) showed encouraging results with gene expression comparable to commercial Lipofectamine 2000 [137]. Zeng et al. [138] prepared PLGA-CS NPs by spontaneous emulsion diffusion method for the delivery of pDNA that showed much higher EE and higher cellular uptake as well as higher hepatitis B virus (HBV) gene-silencing efficiency than plain-PLGA NPs and naked pDNA. In order to achieve prolonged delivery and high transfection efficiency of cationic PLA/DNA complexes, Chen et al. [139] used the copolymer of methoxy polyethylene glycol-PLA (MePEG-PLA) to prepare MePEGPLA-CS NPs as well as PLA-CS NPs by diafiltration method. The NPs of MePEG-PLA-CS had a zeta potential of 15.7 mV with a size of N100 nm, while PLA-CS NPs had the zeta potential of 4.5 mV at pH 7.4. The transfection efficiency of MePEG-PLA-CS/DNA complexes was better than PLA-CS/DNA and Lipofectamine/DNA complexes, which mediated higher gene expression in stomach and intestine of BALB/C mice compared to PLA-CS/DNA and Lipofectamine/DNA complexes, making these as the efficient non-viral vectors for gene delivery. Superior mucoadhesive character and prolonged pharmacological effect compared to drug solution and plain-PLGA were studied [140] by developing CS-modified PLGA NPs in which nuclear factor kappa B (NF-kB) decoy oligonucleotide (ODN) was loaded to treat DEX sulfate induced experimental colitis. The positive zeta potential of these NPs enabled greater cellular uptake and was stable during incubation with nuclease (DNase I) and in simulated gastric fluid. Ionic complex formation between ODN and CS on the NP surface prevented ODN degradation due to acidic conditions and nuclease.

U

682 683


E

8

Fig. 5. Steps involved in the preparation of gelatin-PEI core-shell nanohydrogels. Reprinted from Ref. [141], © 2014, with permission from Elsevier.


726 727 728 729 730 731 732 733 734 735 736 737 738 739 740 741 742 743 744 745 746 747 748 749 750 751 752 753 754 755 756 757 758 759 760 761 762 763 764 765 766 767 768 769


787 788 789 790 791 792 793 794 795 796 797 798 799 800 801 802 803 804

t2:1 t2:2

Table 2 Micro/nanohydrogel systems for gene delivery.

t2:3

System

Size (nm)/zeta potential (mV)

t2:4 t2:5 t2:6 t2:7 t2:8 t2:9 t2:10 t2:11 t2:12 t2:13 t2:14

Quaternized CS 60% TMCSO-60% CS-modified PLGA PLA and CS MePEG-PLA-CS PLGA/CS Ge core and PEI shell DEX-HEMA-co-TMAEMA CS CS HTCSC TMCS

t2:15 t2:16 t2:17

TMCS-MAA Guanidinylated CS CS (MW 90, 50–190, 190–310/ 250,173 kDa) CS (MW 190–310/250 kDa)/ PEGylated CS CS (MW 50–190; 190–310/250; 173 kDa)/PEGylated CS/PEI PEGylated CS/PEI/HA

t2:18 t2:19 t2:20 t2:21 t2:22 t2:23 t2:24 t2:25 t2:26 t2:27

F

O

785 786

Ref.

b200 59, 100/+13 65/+5 94/+13 367/13 200/+40 195–267/17–30 b150/+35 to +41 40–600/+19 to +31 143/+32 112, 124, 129, 166/ +5 to 19 112–243/+5 to +15 100/+15 135–294/+20 to +28

pDNA pDNA pDNA pDNA Nf-kB ODN siRNA siRNA siRNA siRNA siRNA siRNA

[136] [138] [139] [139] [140] [141] [142] [144] [145] [146] [147]

siRNA siRNA siRNA

[147] [149] [150]

124/+22

siRNA

[150]

181, 191, 188/+29, +32/+25 181/−16

siRNA

[150]

siRNA

[150]

U

N C O

R

Drug

PLGA, poly(lactic-co-glycolic acid); PEI, polyethyleneimine; DEX, HEMA-co-TMAEMAdextran hydroxyethyl methacrylate-co-[2-(methacryloyloxy)-ethyl]trimethylammonium chloride; PLA, poly(lactic acid); MePEG, PLA-CS-methoxypolyethyleneglycol-PLA-chitosan; HA, hyaluronic acid; pDNA, plasmid DNA; Ge, gelatin; TMCS, trimethyl chitosan; TMCS, trimethylated chitosan oligomer; HTCSC, N-((2-hydroxy-3-trimethylammonium) propyl) chitosan chloride; Nf-kB ODN, nuclear factor kappa B decoy oligonucleotide; TMCS, MAA is trimethyl chitosan and methacrylic acid copolymer.

806 807

R O

783 784

Polymeric micro/nanohydrogels have been widely explored in recent years with renewed interest for increasing the macromolecular therapeutic outcome and patient compliance. Such systems are particularly advantageous to deliver macromolecular therapeutics compared to other systems due to their lesser cost, 3D network structures and high water retaining capacity, in addition to biocompatible and biodegradable nature. The use of mild preparation conditions is particularly well suited for preserving the structural integrity, delicate nature and stability of these molecules, while their 3D structures are critical for biological efficacy. However, relatively rapid release of macromolecules from hydrogel matrix, burst release, low mechanical strength and less duration of action pose some problems. The fast diffusion of macromolecules from such matrices can be circumvented by increasing the drug/ polymer ratio, cross-link density and by manipulating physicochemical cross-linking methods. Even if these associated problems are resolved, micro/nanoparticle-based systems are difficult to use due to their scaleup, high manufacturing cost and limited sales potential. A better understanding of the patient and medical fraternity needs as well as public– private partnerships would help in developing novel therapies for public healthcare requirements, which would offer better pharmaco-economic benefit. In conclusion, micro/nanohydrogel-based delivery of macromolecules offers unique opportunities and challenges. However, their further preclinical and clinical studies are necessary for assessing the safety and efficacy of these delivery systems prior to marketing.

P

781 782

805

D

779 780

9. Conclusions

Acknowledgements

808 809 810 811 812 813 814 815 816 817 818 819 820 821 822 823 824 825 826 827 828 829 830

Professor Tejraj M. Aminabhavi thanks the All India Council for Technical Education (AICTE), New Delhi, India [1-51/RIFD/EF(13)/ 2011-12] for Emeritus Fellowship. We also acknowledge Dr. S. Rame Gowda Research Institute of Science and Technology, Dharwad, for a partial support of this study in the form of student fellowships and instrumental facilities. We thank Mr. Shrikant A. Tiwari for technical assistance.

831 832

References

838

E

777 778

T

776

C

774 775

E

772 773

Wei et al. [146] developed NPs of N-((2-hydroxy-3trimethylammonium) propyl) CS chloride (HTCSC) for encapsulating telomerase reverse transcriptase siRNA for oral delivery. Here, the polymer coating protected siRNA from enzymatic degradation to permeate to the intestine. Based on these data, authors developed a “two-in-one” nano-complex with both paclitaxel and siRNA encapsulated into one system, which simultaneously transported siRNA and paclitaxel (PTX) to tumor cells and increased drug concentration, leading to better tumor suppression. Similarly, pH-responsive nanocarriers of trimethylchitosan (TMC) and methacrylic acid (MAA) copolymer were reported for the oral delivery of siRNA [147], where incorporation of MAA into polyplex was increased. A significant decrease in zeta potential for MAA-TMCsiRNA complex with greater transfection efficiency (in L929 cells) than TMC-siRNA complex was observed. Guanidinylated chitosan (GCS) formed stable complexes with plasmid DNA under physiological pH that showed lower cytotoxicity with higher transfection efficiency than CS and 8-fold increase in cellular uptake [148,149]. The authors tested these for siRNA delivery to lungs, since GCS was able to condense siRNA at a weight ratio 40:1, forming NPs of diameter ~ 100 nm. Guanidinylation helped to enhance siRNA gene silencing activity than the pristine CS due to better cellular internalization. These NPs were further coupled with salbutamol, a β2adrenoceptor agonist, which improved the targeting specificity of green fluorescent protein (GFP)-siRNA carrier to lung cells having β2adrenergic receptor. Results indicated enhancement in gene silencing activity both in vitro and in vivo (using aerosol treatment in the lung of enhanced green fluorescent protein (EGFP) transgenic mice). A study [150] on NPs of CS used for iv administration of siRNA that suggests the presence of PEGylated CS and PEI is important to achieve high levels of gene silencing in vitro. Since the stability in blood and plasma is significant to achieve the desired result, usual procedure of combining TPP with CS may not be sufficient. Stable NPs were produced by increasing the amount of PEG and inclusion of anionic polymer, viz., HA. Some representative micro/nanohydrogel delivery systems for gene delivery are summarized in Table 2.

R

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9

[1] H. He, W. Dong, J. Gong, J. Wang, V.C. Yang, Developing macromolecular therapeutics: the future drug of choice, Front. Chem. Eng. Chin. 4 (1) (2010) 10–17. [2] W.E. Rudzinski, T.M. Aminabhavi, Chitosan as a carrier for targeted delivery of small interfering RNA, Int. J. Pharm. 399 (2010) 1–11. [3] Z. Antosova, M. Mackova, V. Kral, T. Macek, Therapeutic application of peptides and proteins: parenteral forever? Trends Biotechnol. 27 (11) (2009) 628–635. [4] R.C. Mundargi, V.R. Babu, V. Rangaswamy, P. Patel, T.M. Aminabhavi, Nano/micro technologies for delivering macromolecular therapeutics using poly(d, llactide-co-glycolide) and its derivatives, J. Control. Release 125 (2008) 193–209. [5] K. Chaturvedi, K. Ganguly, A.R. Kulkarni, V.H. Kulkarni, M.N. Nadagouda, W.E. Rudzinski, T.M. Aminabhavi, Cyclodextrin-based siRNA delivery nanocarriers: A state-of the-art review, Expert Opin. Drug Deliv. 8 (2011) 1455–1468. [6] R. Singh, S. Singh, J.W. Lillard, Past, present, and future technologies for oral delivery of therapeutic proteins, J. Pharm. Sci. 97 (7) (2008) 2497–2523. [7] M. Werle, A. Makhlof, H. Takeuchi, Oral protein delivery: a patent review of academic and industrial approaches, Recent Pat. Drug Deliv. Formul. 3 (2) (2009) 94–104. [8] K. Chaturvedi, K. Ganguly, M.N. Nadagouda, T.M. Aminabhavi, Polymeric hydrogels for oral insulin delivery, J. Control. Release 165 (2013) 129–138. [9] V.R. Babu, P. Patel, R.C. Mundargi, V. Rangaswamy, T.M. Aminabhavi, Developments in polymeric devices for oral insulin delivery, Expert Opin. Drug Deliv. 5 (2008) 403–415. [10] O. Wichterle, D. Lim, Hydrophilic gels for biological use, Nature 185 (1960) 117–118. [11] J.I. Kim, B.S. Lee, C. Chun, J.K. Cho, S.Y. Kim, S.C. Song, Long-term theranostic hydrogel system for solid tumors, Biomaterials 33 (7) (2012) 2251–2259. [12] M. Hamidi, A. Azadi, P. Rafiei, Hydrogel nanoparticles in drug delivery, Adv. Drug Deliv. Rev. 60 (2008) 1638–1649. [13] M.D. Moya, C. Alvarez, Cyclodextrin-based nanogels for pharmaceutical and biomedical applications, Int. J. Pharm. 428 (2012) 152–163. [14] K. Ganguly, T.M. Aminabhavi, A.R. Kulkarni, Colon targeting of 5-Fluorouracil using polyethylene glycol crosslinked chitosan microspheres enteric coated with cellulose acetate phthalate, Ind. Eng. Chem. Res. 50 (21) (2011) 11797–11807. [15] N. Kashyap, N. Kumar, M.N.V. Ravi Kumar, Hydrogels for pharmaceutical and biomedical applications, Crit. Rev. Ther. Drug Carrier Syst. 22 (2) (2005) 107–150.


833 834 835 836 837

839 840 841 842 843 844 845 846 847 848 849 850 851 852 853 854 855 856 857 858 859 860 861 862 863 864 865 866 867 868 869 870 871 872 873

D

P

R O

O

F

[46] D. Delgado, A.R. Gascón, A. Del Pozo-Rodríguez, E. Echevarría, A.P. Ruiz de Garibay, J.M. Rodríguez, M.A. Solinís, Dextran-protamine-solid lipid nanoparticles as a nonviral vector for gene therapy: in vitro characterization and in vivo transfection after intravenous administration to mice, Int. J. Pharm. 425 (2012) 35–43. [47] B. Sarmento, A. Ribeiro, F. Veiga, D. Ferreira, Development and characterization of new insulin containing polysaccharide nanoparticles, Colloids Surf. B: Biointerfaces 53 (2006) 193–202. [48] B. Sarmento, A. Ribeiro, F. Veiga, D. Ferreira, R. Neufeld, Oral bioavailability of insulin contained in polysaccharide nanoparticles, Biomacromolecules 8 (2007) 3054–3060. [49] F. Abedini, H. Hosseinkhani, M. Ismail, A.J. Domb, A.R. Omar, P.P. Chong, P.D. Hong, D.S. Yu, I.Y. Farber, Cationized dextran nanoparticle-encapsulated CXCR4-siRNA enhanced correlation between CXCR4 expression and serum alkaline phosphatase in a mouse model of colorectal cancer, Int. J. Nanomedicine 7 (2012) 4159–4168. [50] G. Kogan, L. Soltés, R. Stern, P. Gemeiner, Hyaluronic acid: a natural biopolymer with a broad range of biomedical and industrial applications, Biotechnol. Lett. 29 (2007) 17–25. [51] F. Lee, J.E. Chung, M. Kurisawa, An injectable hyaluronic acid-tyramine hydrogel system for protein delivery, J. Control. Release 134 (2009) 186–193. [52] S.K. Hahn, J.S. Kim, T. Shimobouji, Injectable hyaluronic acid microhydrogels for controlled release formulation of erythropoietin, J. Biomed. Mater. Res. A 80 (2007) 916–924. [53] S.K. Hahn, E.J. Oh, H. Miyamoto, T. Shimobouji, Sustained release formulation of erythropoietin using hyaluronic acid hydrogels crosslinked by Michael addition, Int. J. Pharm. 322 (2006) 44–51. [54] E.J. Oh, K. Park, K.S. Kim, J. Kim, J.A. Yang, J.H. Kong, M.Y. Lee, A.S. Hoffman, S.K. Hahn, Target specific and long-acting delivery of protein, peptide, and nucleotide therapeutics using hyaluronic acid derivatives, J. Control. Release 141 (2010) 2–12. [55] E. Esposito, E. Menegatti, R. Cortesi, Hyaluronan-based microspheres as tools for drug delivery: a comparative study, Int. J. Pharm. 288 (2005) 35–49. [56] S.J. Kim, S.K. Hahn, M.J. Kim, D.H. Kim, Y.P. Lee, Development of a novel sustained release formulation of recombinant human growth hormone using sodium hyaluronate microparticles, J. Control. Release 104 (2005) 323–335. [57] A. Kumar, B. Sahoo, A. Montpetit, S. Behera, R. Lockey, S. Mohapatra, Development of hyaluronic acid-Fe2O3 hybrid magnetic nanoparticles for targeted delivery of peptides, Nanomedicine: NBM 3 (2007) 132–137. [58] E. Lee, M. Kwon, K. Na, J. Bae, Protein release behavior from porous microparticle with lysozyme/hyaluronate ionic complex, Colloids Surf. B: Biointerfaces 55 (2007) 125–130. [59] H.H. Tønnesen, J. Karlsen, Alginate in drug delivery systems, Drug Dev. Ind. Pharm. 28 (2002) 621–630. [60] M.C. Peters, B.C. Isenberg, J.A. Rowley, D.J. Mooney, Release from alginate enhances the biological activity of vascular endothelial growth factor, J. Biomater. Sci. Polym. Ed. 9 (1998) 1267–1278. [61] S.Y. Rabbany, J. Pastore, M. Yamamoto, T. Miller, S. Rafii, R. Aras, M. Penn, Continuous delivery of stromal cell-derived factor-1 from alginate scaffolds accelerates wound healing, Cell Transplant. 19 (2010) 399–408. [62] K. Bowey, B.E. Swift, L.E. Flynn, R.J. Neufeld, Characterization of biologically active insulin-loaded alginate microparticles prepared by spray drying, Drug Dev. Ind. Pharm. 39 (2013) 457–465. [63] T. Suksamran, P. Opanasopit, T. Rojanarata, T. Ngawhirunpat, U. Ruktanonchai, P. Supaphol, Biodegradable alginate microparticles developed by electrohydrodynamic spraying techniques for oral delivery of protein, J. Microencapsulation 26 (2009) 563–570. [64] A. Sharma, A. Gupta, G. Rath, A. Goyal, R.B. Mathur, S.R. Dhakate, Electrospun composite nanofiber-based transmucosal patch for anti-diabetic drug delivery, J. Mater. Chem. B 1 (27) (2013) 3410–3418. [65] K. Moore, J. Amos, J. Davis, R. Gourdie, J.D. Potts, Characterization of polymeric microcapsules containing a low molecular weight peptide for controlled release, Microsc. Microanal. 19 (2013) 213–226. [66] G. Mahidhara, R.K. Kanwar, J.R. Kanwar, Novel nanoplatform for oral delivery of anti-cancer biomacromolecules, Int. J. Nanotechnol. 9 (2012) 942–960. [67] O. Jeon, C. Powell, L.D. Solorio, M.D. Krebs, E. Alsberg, Affinity-based growth factor delivery using biodegradable, photocrosslinked heparin-alginate hydrogels, J. Control. Release 154 (2011) 258–266. [68] I.M. El-Sherbiny, A. Salama, A.A. Sarhan, Ionotropically cross-linked pH-sensitive IPN hydrogel matrices as potential carriers for intestine-specific oral delivery of protein drugs, Drug Dev. Ind. Pharm. 37 (2011) 121–130. [69] Y. Zhang, W. Wei, P. Lv, L. Wang, G. Ma, Preparation and evaluation of alginatechitosan microspheres for oral delivery of insulin, Eur. J. Pharm. Biopharm. 77 (2011) 11–19. [70] I.M. El-Sherbiny, Enhanced pH-responsive carrier system based on alginate and chemically modified carboxymethyl chitosan for oral delivery of protein drugs: Preparation and in vitro assessment, Carbohydr. Polym. 80 (2010) 1125–1136. [71] A.W. Chan, R.J. Neufeld, Tuneable semi-synthetic network alginate for absorptive encapsulation and controlled release of protein therapeutics, Biomaterials 31 (2010) 9040–9047. [72] L. Pescosolido, S. Miatto, C. Di Meo, C. Cencetti, T. Coviello, F. Alhaique, P. Matricardi, Injectable and in situ gelling hydrogels for modified protein release, Eur. Biophys. J. 39 (2010) 903–909. [73] E. Ruvinov, J. Leor, S. Cohen, The effects of controlled HGF delivery from an affinitybinding alginate biomaterial on angiogenesis and blood perfusion in a hindlimb ischemia model, Biomaterials 31 (2010) 4573–4582. [74] F.M. Goycoolea, G. Lollo, C. Remuñán-López, F. Quaglia, M.J. Alonso, Chitosanalginate blended nanoparticles as carriers for the transmucosal delivery of macromolecules, Biomacromolecules 10 (2009) 1736–1743.

N

C

O

R

R

E

C

T

[16] P. Munk, T.M. Aminabhavi, Introduction to macromolecular science, 2nd edn. John Wiley & Sons, New York, 2002. [17] W. Zhao, X. Jin, Y. Cong, Y. Liu, J. Fu, Degradable natural polymer hydrogels for articular cartilage tissue engineering, J. Chem. Technol. Biotechnol. 88 (2013) 327–339. [18] I. Levental, P.C. Georges, P.A. Janmey, Soft biological materials and their impact on cell function, Soft Matter 2 (2006) 1–9. [19] P.H. Kouwer, M. Koepf, V.A. Le Sage, M. Jaspers, A.M. van Buul, Z.H. Eksteen-Akeroyd, T. Woltinge, E. Schwartz, H.J. Kitto, R. Hoogenboom, S.J. Picken, R.J. Nolte, E. Mendes, A.E. Rowan, Responsive biomimetic networks from polyisocyanopeptide hydrogels, Nature 493 (7434) (2013) 651–655. [20] J. Kadokawa, Precision polysaccharide synthesis catalyzed by enzymes, Chem. Rev. 111 (2011) 4308–4345. [21] M.D. Buschmann, A. Merzouki, M. Lavertu, M. Thibault, M. Jean, V. Darras, Chitosans for delivery of nucleic acids, Adv. Drug Deliv. Rev. 65 (2013) 1234–1270. [22] S.A. Agnihotri, N.N. Mallikarjuna, T.M. Aminabhavi, Recent advances on chitosanbased micro-and nanoparticles in drug delivery, J. Control. Release 100 (1) (2004) 5–28. [23] S. Al-Qadi, A. Grenha, D. Carrión-Recio, B. Seijo, C. Remuñán-López, Microencapsulated chitosan nanoparticles for pulmonary protein delivery: in vivo evaluation of insulin-loaded formulations, J. Control. Release 157 (2012) 383–390. [24] P. Mukhopadhyay, K. Sarkar, M. Chakraborty, S. Bhattacharya, R. Mishra, P.P. Kundu, Oral insulin delivery by self-assembled chitosan nanoparticles: in vitro and in vivo studies in diabetic animal model, Mater. Sci. Eng. C 33 (2013) 376–382. [25] C. Vauthier, C. Zandanel, A.L. Ramon, Chitosan-based nanoparticles for in vivo delivery of interfering agents including siRNA, Curr. Opin. Colloid Interface Sci. 18 (2013) 406–418. [26] H. Zhang, I.A. Alsarra, S.H. Neau, An in vitro evaluation of a chitosan containing multiparticulate system for macromolecule delivery to the colon, Int. J. Pharm. 239 (2002) 197–205. [27] H. Tozaki, T. Odoriba, N. Okada, T. Fujita, A. Terabe, T. Suzuki, S. Okabe, S. Muranishi, A. Yamamato, Chitosan capsules for colon specific drug delivery: enhanced localization of 5-amino salicylic acid in the large intestine accelerates healing of TNBSinduced colitis in rats, J. Control. Release 82 (2002) 51–61. [28] Y. Xu, C. Zhan, L. Fan, L. Wang, H. Zheng, Preparation of dual crosslinked alginate chitosan blend gel beads and in vitro controlled release in oral site- specific drug delivery system, Int. J. Pharm. 336 (2007) 329–337. [29] F. Bigucci, B. Luppi, T. Cerchiara, M. Sorrenti, G. Bettinetti, L. Rodriquez, V. Zecchi, Chitosan/pectin polyelectrolyte complexes: Selection of suitable preparative conditions for colon-specific delivery of vancomycin, Eur. J. Pharm. Sci. 35 (2008) 435–441. [30] E. Assaad, Y.J. Wang, X.X. Zhu, M.A. Mateescu, Polyelectrolyte complex of carboxymethyl starch and chitosan as drug carrier for oral administration, Carbohydr. Polym. 84 (2011) 1399–1407. [31] B. Sarmento, D.C. Ferreira, L. Jorgensen, M. Van de Weert, Probing insulin's secondary structure after entrapment into alginate/chitosan nanoparticles, Eur. J. Pharm. Biopharm. 65 (2007) 10–17. [32] J. Du, J. Dai, J. Liu, T. Dankovich, Novel pH-sensitive polyelectrolyte carboxymethyl konjac glucomannan-chitosan beads as drug carriers, React. Funct. Polym. 66 (2006) 1055–1061. [33] P. Li, Y. Wang, F. Zeng, L. Chen, Z. Peng, L.X. Kong, Synthesis and characterization of folate conjugated chitosan and cellular uptake of its nanoparticles in HT-29 cells, Carbohydr. Res. 346 (2011) 801–806. [34] G. Kaur, S. Jain, A.K. Tiwary, Chitosan-carboxymethyl tamarind kernel powder interpolymer complexation: investigations for colon drug delivery, Sci. Pharm. 78 (2010) 57–78. [35] A.R. Kulkarni, V.I. Hukkeri, H.W. Sung, H.F. Liang, A novel method for the synthesis of the PEG-crosslinked chitosan with a pH independent swelling behavior, Macromol. Biosci. 5 (2005) 925–928. [36] S.C. Chen, Y.C. Wu, F.L. Mi, Y.H. Lin, L.C. Yu, H.W. Sung, A novel pH-sensitive hydrogel composed of N, O-carboxymethyl chitosan and alginate cross-linked by genipin for protein drug delivery, J. Control. Release 96 (2004) 285–300. [37] Y.Z. Du, Q. Weng, H. Yuan, F.Q. Hu, Synthesis and antitumor activity of stearateg-dextran micelles for intracellular doxorubicin delivery, ACS Nano 4 (2010) 6894–6902. [38] S.R. Van Tomme, W.E. Hennink, Biodegradable dextran hydrogels for protein delivery applications, Expert Rev. Med. Dev. 4 (2007) 147–164. [39] S.G. Lévesque, M.S. Shoichet, Synthesis of enzyme-degradable, peptide-crosslinked dextran hydrogels, Bioconjug. Chem. 18 (2007) 874–885. [40] S. Shu, X. Zhang, Z. Wu, Z. Wang, C. Li, Delivery of protein drugs using nanoparticles self-assembled from dextran sulfate and quaternized chitosan, J. Control. Release 152 (2011) e170–e172. [41] W. Yuan, F. Wu, Y. Geng, S. Xu, T. Jin, An effective approach to prepare uniform protein-Zn2+ nanoparticles under mild conditions, Nanotechnology 18 (2007) 145601–145608. [42] W. Yuan, Z. Hu, J. Su, F. Wu, Z. Liu, T. Jin, Preparation and characterization of recombinant human growth hormone-Zn2 + dextran nanoparticles using aqueous phase-aqueous phase emulsion, Nanomedicine: NBM 8 (2012) 424–427. [43] F. Wu, Z. Zhou, J. Su, L. Wei, W. Yuan, T. Jin, Development of dextran nanoparticles for stabilizing delicate proteins, Nanoscale Res. Lett. 8 (2013) 197. [44] K.B. Chalasani, G.J. Russell-Jones, A.K. Jain, P.V. Diwan, S.K. Jain, Effective oral delivery of insulin in animal models using vitamin B12-coated dextran nanoparticles, J. Control. Release 122 (2007) 141–150. [45] F.M. Chen, H. Lu, L.A. Wu, L.N. Gao, Y. An, J. Zhang, Surface-engineering of glycidyl methacrylated dextran/gelatin microcapsules with thermo-responsive poly (N-isopropylacrylamide) gates for controlled delivery of stromal cell-derived factor-1α, Biomaterials 34 (2013) 6515–6527.

U

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960 961 962 963 964 965 966 967 968 969 970 971 972 973 974 975 976 977 978 979 980 981 982 983 984 985 986 987 988 989 990 991 992 993 994 995 996 997 998 999 1000 1001 1002 1003 1004 1005 1006 1007 1008 1009 1010 1011 1012 1013 1014 1015 1016 1017 1018 1019 1020 1021 1022 1023 1024 1025 1026 1027 1028 1029 1030 1031 1032 1033 1034 1035 1036 1037 1038 1039 1040 1041 1042 1043 1044 1045


[110]

[111]

[112]

[113]

[114]

[115]

D

[116]

C

E

R

R

N C O

F

[109]

O

[108]

R O

[107]

of alginate coated chitosan nanoparticles for mucosal vaccination, J. Control. Release 114 (2006) 348–358. F. Chen, Z.-R. Zhang, F. Yuan, X. Qin, M. Wang, Y. Huang, In vitro and in vivo study of N-trimethyl chitosan nanoparticles for oral protein delivery, Int. J. Pharm. 349 (2008) 226–233. M. Tafaghodi, S.A. Sajadi Tabassi, M.R. Jaafari, Induction of systemic and mucosal immune responses by intranasal administration of alginate microspheres encapsulated with tetanus toxoid and CpG-ODN, Int. J. Pharm. 319 (2006) 37–43. H.J. Lee, G. Riley, O. Johnson, J.L. Cleland, N. Kim, M. Charnis, L. Bailey, E. Duenas, A. Shahzamani, M. Marian, A.J.S. Jones, S.D. Putney, In vivo characterization of sustained-release formulations of human growth hormone, J. Pharmacol. Exp. Ther. 281 (1997) 1431–1439. J. Cázares-Delgadillo, A. Ganem-Rondero, Y.N. Kalia, Human growth hormone: New delivery systems, alternative routes of administration, and their pharmacological relevance, Eur. J. Pharm. Biopharm. 78 (2011) 278–288. H.J. Chung, Y. Lee, T.G. Park, Thermo-sensitive and biodegradable hydrogels based on stereocomplexed Pluronic multi-block copolymers for controlled protein delivery, J. Control. Release 127 (2008) 22–30. G. Tae, J.A. Kornfield, J.A. Hubbell, Sustained release of human growth hormone from in situ forming hydrogels using self-assembly of fluoroalkyl-ended poly (ethylene glycol), Biomaterials 26 (2005) 5259–5266. L. Illum, F. Jordan, A.L. Lewis, CriticalSorb(tm): A novel efficient nasal delivery system for human growth hormone based on Solutol HS15, J. Control. Release 162 (2012) 194–200. O.L. Johnson, W. Jaworowicz, J.L. Cleland, L. Bailey, M. Charnis, E. Duenas, C. Wu, D. Shepard, S. Magil, T. Last, A.J.S. Jones, S.D. Putney, The stabilization and encapsulation of human growth hormone into biodegradable microspheres, Pharm. Res. 14 (1997) 730–735. K. Fu, D.W. Pack, A.M. Klibanov, R. Langer, Visual evidence of acidic environment within degrading poly(lactic-co-glycolic acid) (PLGA) microspheres, Pharm. Res. 17 (2000) 100–106. H.H. Kwak, W.S. Shim, M.K. Choi, M.K. Son, Y.J. Kim, H.C. Yang, T.H. Kim, G.I. Lee, B.M. Kim, S.H. Kang, C.K. Shim, Development of a sustained-release recombinant human growth hormone formulation, J. Control. Release 137 (2009) 160–165. K.Y. Lee, S.H. Yuk, Polymeric protein delivery systems, Prog. Polym. Sci. 32 (7) (2007) 669–697. M.R. Park, B.B. Seo, S.C. Song, Dual ionic interaction system based on polyelectrolyte complex and ionic, injectable, and thermosensitive hydrogel for sustained release of human growth hormone, Biomaterials 34 (2013) 1327–1336. D. Pyo, C. Lim, D. Cho, D. Oh, Size prediction of recombinant human growth hormone nanoparticles produced by supercritical fluid precipitation, Anal. Bioanal. Chem. 387 (2007) 901–907. D.W. Tang, S.H. Yu, W.S. Wu, H.Y. Hsieh, Y.C. Tsai, F.L. Mi, Hydrogel microspheres for stabilization of an antioxidant enzyme: effect of emulsion cross-linking of a dual polysaccharide system on the protection of enzyme activity, Colloids Surf. B: Biointerfaces 113 (2014) 59–68. H.E. Lee, M.J. Lee, C.R. Park, A.Y. Kim, K.H. Chun, H.J. Hwang, D.H. Oh, S.O. Jeon, J.S. Kang, T.S. Jung, G.J. Choi, S. Lee, Preparation and characterization of salmon calcitonin-sodium triphosphate ionic complex for oral delivery, J. Control. Release 143 (2010) 251–257. A. Makhlof, M. Werle, Y. Tozuka, H. Takeuchi, Nanoparticles of glycol chitosan and its thiolated derivative significantly improved the pulmonary delivery of calcitonin, Int. J. Pharm. 397 (2010) 92–95. M.J. Lee, D.Y. Seo, H.E. Lee, G.J. Choi, Therapeutic effect of chitosan modification on salmon-calcitonin-loaded PLGA nanoparticles, Korean J. Chem. Eng. 28 (6) (2011) 1406–1411. A. Hafner, J. Lovrić, D. Voinovich, H. Filipović-Grčić, Melatonin-loaded lecithin/ chitosan nanoparticles: physicochemical characterisation and permeability through Caco-2 cell monolayers, Int. J. Pharm. 381 (2009) 205–213. A. Hafner, J. Lovric, I. Pepic, J.F. Grcic, Lecithin/chitosan nanoparticles for transdermal delivery of melatonin, J. Microencapsulation 28 (2011) 807–815. C. Prego, M. Garcıá, D. Torres, M.J. Alonso, Transmucosal macromolecular drug delivery, J. Control. Release 101 (2005) 151–162. M. Garcia-Fuentes, D. Torres, M.J. Alonso, New surface-modified lipid nanoparticles as delivery vehicles for salmon calcitonin, Int. J. Pharm. 296 (2005) 122–132. S.M. Ryan, J. McMorrow, A. Umerska, H.B. Patel, K.N. Kornerup, L. Tajber, E.P. Murphy, M. Perretti, O.I. Corrigan, D.J. Brayden, An intra-articular salmon calcitonin-based nanocomplex reduces experimental inflammatory arthritis, J. Control. Release 167 (2013) 120–129. C. Sinsuebpol, J. Chatchawalsaisin, P. Kulvanich, Preparation and in vivo absorption evaluation of spray dried powders containing salmon calcitonin loaded chitosan nanoparticles for pulmonary delivery, Drug Des. Devel. Ther. 7 (2013) 861–873. I.A. Khalil, K. Kogure, H. Akita, H. Harashima, Uptake pathways and subsequent intracellular trafficking in nonviral gene delivery, Pharmacol. Rev. 58 (2006) 32–45. J.M. Garza, N. Jessel, G. Ladam, V. Dupray, S. Muller, J.F. Stoltz, P. Schaaf, J.C. Voegel, P. Lavalle, Polyelectrolyte multilayers and degradable polymer layers as multicompartment films, Langmuir 21 (2005) 12372–12377. C.R. Middaugh, C.M. Wiethoff, Barriers to nonviral gene delivery, J. Pharm. Sci. 92 (2003) 203–217. J. Chen, W.L. Yang, G. Li, J. Qian, J.L. Xue, S.K. Fu, D.R. Lu, Transfection of mEpo gene to intestinal epithelium in vivo mediated by oral delivery of chitosan-DNA nanoparticles, World J. Gastroenterol. 10 (2004) 112–116. M. Koping-Hoggard, I. Tubulekas, H. Guan, K. Edwards, M. Nilsson, K.M. Varum, P. Artursson, Chitosan as a nonviral gene delivery system. Structure-property relationships and characteristics compared with polyethylenimine in vitro and after lung administration in vivo, Gene Ther. 8 (2001) 1108–1121.

P

[106]

E

[117]

T

[75] A. Schoubben, P. Blasi, S. Giovagnoli, L. Perioli, C. Rossi, M. Ricci, Novel composite microparticles for protein stabilization and delivery, Eur. J. Pharm. Sci. 36 (2009) 226–234. [76] S.M. Lee, H.Y. Lee, Y.M. Kim, M.K. Kim, S.O. Sohn, H.D. Ghim, Preparation and characterization of release behavior for hydrophilic dung loaded alginate nanohydrogels, NSTI Nanotechnology Conference and Trade Show - NSTI Nanotech, Technical Proceedings, 4, 2007, pp. 250–253. [77] F.-D. Cui, A.-J. Tao, D.-M. Cun, L.-Q. Zhang, K. Shi, Preparation of insulin loaded PLGA-Hp55 nanoparticles for oral delivery, J. Pharm. Sci. 96 (2) (2007) 421–427. [78] K.S. Soppimath, T.M. Aminabhavi, A.R. Kulkarni, W.E. Rudzinski, Biodegradable polymeric nanoparticles as drug delivery devices, J. Control. Release 70 (2001) 1–20. [79] A.K. Gaharwar, N.A. Peppas, A. Khademhosseini, Nanocomposite hydrogels for biomedical applications, Biotechnol. Bioeng. 111 (2014) 441–453. [80] U. Hasegawa, S.I. Sawada, T. Shimizu, T. Kishida, E. Otsuji, O. Mazda, K. Akiyoshi, Raspberry-like assembly of cross-linked nanogels for protein delivery, J. Control. Release 140 (2009) 312–317. [81] I. Brigger, C. Dubernet, P. Couvreur, Nanoparticles in cancer therapy and diagnosis, Adv. Drug Deliv. Rev. 54 (2002) 631–651. [82] U. Gaur, S.K. Sahoo, T.K. De, P.C. Ghosh, A. Maitra, P.K. Ghosh, Biodistribution of fluoresceinated dextran using novel nanoparticles evading reticuloendothelial system, Int. J. Pharm. 202 (2000) 1–10. [83] A. Balasubramaniem, Z.M. Mohammed, N.D. Vaziri, H.M. Said, Effect of folate oversupplementation on folate uptake by human intestinal and renal epithelial cells, Am. J. Clin. Nutr. 86 (1) (2007) 159–166. [84] Y. Lu, P.S. Low, Folate-mediated delivery of macromolecular anticancer therapeutic agents, Adv. Drug Deliv. Rev. 64 (2012) 342–352. [85] M. Goldberg, I. Gomez-Orellana, Challenges for the oral delivery of macromolecules, Nat. Rev. 2 (2003) 289–295. [86] M.C. Chen, K. Sonaje, K.J. Chen, H.-W. Sung, A review of the prospects for polymeric nanoparticle platforms in oral insulin delivery, Biomaterials 32 (2011) 9826–9838. [87] J.S. Patton, Mechanisms of macromolecule absorption by the lungs, Adv. Drug Deliv. Rev. 19 (1996) 3–36. [88] M.I. Ugwoke, R.M. Agu, N. Verbeke, R. Kinget, Nasal mucoadhesive drug delivery: background, applications, trends and future perspectives, Adv. Drug Deliv. Rev. 57 (2005) 1640–1665. [89] P. Shakya, N.V.S. Madhav, A.K. Shakya, K. Singh, Palatal mucosa as a route for systemic drug delivery: a review, J. Control. Release 151 (2011) 2–9. [90] M. Oak, R. Mandke, J. Singh, Smart polymers for peptide and protein parenteral sustained delivery, Drug Discov. Today: Technol. 9 (2) (2012) e131–e140. [91] J. Holmgren, C. Czerkinsky, Mucosal immunity and vaccines, Nat. Med. Suppl. 11 (4) (2005) 545–553. [92] B. Slütter, N. Hagenaars, W. Jiskoot, Rational design of nasal vaccines, J. Drug Target. 16 (1) (2008) 1–17. [93] S. Sharma, T.K.S. Mukkur, H.A.E. Benson, Y. Chen, Pharmaceutical aspects of intranasal delivery of vaccines using particulate systems, J. Pharm. Sci. 98 (3) (2009) 812–843. [94] A. Bernkop-Schnürch, M. Hornof, D. Guggi, Thiolated chitosans, Eur. J. Pharm. Biopharm. 57 (2004) 9–17. [95] B.D. Zhu, Y.Q. Qie, J.L. Wang, Y. Zhang, Q.Z. Wang, Y. Xu, H.H. Wang, Chitosan microspheres enhance the immunogenicity of an Ag85B-based fusion protein containing multiple T-cell epitopes of mycobacterium tuberculosis, Eur. J. Pharm. Biopharm. 66 (2007) 318–326. [96] K. Khatri, A.K. Goyal, P.N. Gupta, N. Mishra, S.P. Vyas, Plasmid DNA loaded chitosan nanoparticles for nasal mucosal immunization against Hepatitis B, Int. J. Pharm. 354 (2008) 235–241. [97] M.L. Kang, S.G. Kang, H.L. Jiang, S.W. Shin, D.Y. Lee, J.M. Ahn, N. Rayamahji, I.K. Park, S.J. Shin, C.S. Cho, H.S. Yoo, In vivo induction of mucosal immune responses by intranasal administration of chitosan microspheres containing Bordetella bronchiseptica DNT, Eur. J. Pharm. Biopharm. 63 (2006) 215–220. [98] J.-L. Huang, Y.-X. Yin, Z.-M. Pan, G. Zhang, A.-P. Zhu, X.-F. Liu, X.A. Jiao, Intranasal immunization with chitosan/pCAGGS-flaA nanoparticles inhibits campylobacter jejuni in a white leghorn model, J. Biomed. Biotechnol. (2010) 8, http://dx.doi. org/10.1155/2010/589476 (Article ID 589476). [99] B. Sayin, S. Somavarapu, X.W. Li, M. Thanou, D. Sesardic, H.O. Alpar, S. Senel, Mono– n-carboxymethyl chitosan (MCC) and N-trimethyl chitosan (TMC) nanoparticles for non-invasive vaccine delivery, Int. J. Pharm. 363 (2008) 139–148. [100] M. Amidi, S.G. Romeijn, J.C. Verhoef, H.E. Junginger, L. Bungener, A. Huckriede, D.J. Crommelin, W. Jiskoot, N-trimethyl chitosan (TMC) nanoparticles loaded with influenza subunit antigen for intranasal vaccination: biological properties and immunogenicity in a mouse model, Vaccine 25 (2007) 144–153. [101] M. Amidi, S.G. Romeijn, G. Borchard, H.E. Junginger, W.E. Hennink, W. Jiskoot, Preparation and characterization of protein-loaded N-trimethyl chitosan nanoparticles as nasal delivery system, J. Control. Release 111 (2006) 107–116. [102] B. Sayin, S. Somavarapu, X.W. Li, D. Sesardic, S. Şenel, O.H. Alpar, TMC-MCC (N-trimethyl chitosan-mono-N-carboxymethyl chitosan) nanocomplexes for mucosal delivery of vaccines, Eur. J. Pharm. Sci. 38 (2009) 362–369. [103] O. Borges, A. Cordeiro-da-Silva, J. Tavares, N. Santarem, A. de Sousa, G. Borchard, H. E. Junginger, Immune response by nasal delivery of Hepatitis B surface antigen and codelivery of a CpG ODN in alginate coated chitosan nanoparticles, Eur. J. Pharm. Biopharm. 69 (2008) 405–416. [104] X.Y. Li, X.Y. Kong, S.A. Shi, X.L. Zheng, G. Guo, Y.Q. Wei, Z.Y. Qian, Preparation of alginate coated chitosan microparticles for vaccine delivery, BMC Biotechnol. 8 (2008) (article 89). [105] O. Borges, A. Cordeiro-da-Silva, S.G. Romeijn, M. Amidi, A. De Sousa, G. Borchard, H. E. Junginger, Uptake studies in rat Peyer's patches, cytotoxicity and release studies

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[118]

[119]

[120]

[121]

[122]

[123]

[124] [125] [126] [127]

[128]

[129] [130]

[131] [132]

[133]


1132 1133 1134 1135 1136 1137 1138 1139 1140 1141 1142 1143 1144 1145 1146 1147 1148 1149 1150 1151 1152 1153 1154 1155 1156 1157 1158 1159 1160 1161 1162 1163 1164 1165 1166 1167 1168 1169 1170 1171 1172 1173 1174 1175 1176 1177 1178 1179 1180 1181 1182 1183 1184 1185 1186 1187 1188 1189 1190 1191 1192 1193 1194 1195 1196 1197 1198 1199 1200 1201 1202 1203 1204 1205 1206 1207 1208 1209 1210 1211 1212 1213 1214 1215 1216 1217

[143] E. Şalva, L. Kabasakal, F. Eren, N. Özkan, F. Çakalağaoğlu, J. Akbuğa, Local delivery of chitosan/VEGF siRNA nanoplexes reduces angiogenesis and growth of breast cancer in vivo, Nucleic Acid Ther. 22 (2012) 40–48. [144] K. Sharma, S. Somavarapua, A. Colombani, N. Govind, K.M.G. Taylor, Nebulised siRNA encapsulated crosslinked chitosan nanoparticles for pulmonary delivery, Int. J. Pharm. 455 (2013) 241–247. [145] K.A. Howard, U.L. Rahbek, X. Liu, C.K. Damgaard, S.Z. Glud, M.Ø. Andersen, M.B. Hovgaard, A. Schmitz, J.R. Nyengaard, F. Basenbacher, J. Kjems, RNA interference in vitro and in vivo using a chitosan/siRNA nanoparticle system, Mol. Ther. 14 (4) (2006) 476–484. [146] W. Wei, P.P. Lv, X.M. Chen, Z.G. Yue, Q. Fu, S.Y. Liu, H. Yue, G.H. Ma, Codelivery of mTERT siRNA and paclitaxel by chitosan-based nanoparticles promoted synergistic tumor suppression, Biomaterials 34 (2013) 3912–3923. [147] V. Dehousse, N. Garbacki, A. Colige, B. Evrard, Development of pH-responsive nanocarriers using trimethylchitosans and methacrylic acid copolymer for siRNA delivery, Biomaterials 31 (2010) 1839–1849. [148] X. Zhai, P. Sun, Y. Luo, C. Ma, J. Xu, W. Liu, Guanidinylation: a simple way to fabricate cell penetrating peptide analogue-modified chitosan vector for enhanced gene delivery, J. Appl. Polym. Sci. 121 (2011) 3569–3578. [149] Y. Luo, X. Zhai, C. Ma, P. Sun, Z. Fu, W. Liu, J. Xu, An inhalable β2-adrenoceptor ligand-directed guanidinylated chitosan carrier for targeted delivery of siRNA to lung, J. Control. Release 162 (2012) 28–36. [150] H. Ragelle, R. Riva, G. Vandermeulen, B. Naeye, V. Pourcelle, C.S. Le Duff, C. D'Haese, B. Nysten, K. Braeckmans, S.C. De Smedt, C. Jérôme, V. Préat, Chitosan nanoparticles for siRNA delivery: optimizing formulation to increase stability and efficiency, J. Control. Release 176 (2014) 54–63.

F

[134] S. Mao, W. Sun, T. Kissel, Chitosan-based formulations for delivery of DNA and siRNA, Adv. Drug Deliv. Rev. 62 (2010) 12–27. [135] K. Bowman, R. Sarkar, S. Raut, K.W. Leong, Gene transfer to hemophilia A mice via oral delivery of FVIII-chitosan nanoparticles, J. Control. Release 132 (2008) 252–259. [136] F. Zheng, X.W. Shi, G.F. Yang, L.L. Gong, H.Y. Yuan, Y.J. Cui, Y. Wang, Y.M. Du, Y. Li, Chitosan nanoparticle as gene therapy vector via gastrointestinal mucosa administration: results of an in vitro and in vivo study, Life Sci. 80 (2007) 388–396. [137] M. Gupta, A.K. Gupta, Hydrogel pullulan nanoparticles encapsulating pBUDLacZ plasmid as an efficient gene delivery carrier, J. Control. Release 99 (2004) 157–166. [138] P. Zeng, Y. Xu, C. Zeng, H. Ren, M. Peng, Chitosan-modified poly(d, l-lactide-coglycolide) nanospheres for plasmid DNA delivery and HBV gene-silencing, Int. J. Pharm. 415 (2011) 259–266. [139] J. Chen, B. Tian, X. Yin, Y. Zhang, D. Hu, Z. Hu, M. Liu, Y. Pan, J. Zhao, H. Li, C. Hou, J. Wang, Y. Zhang, Preparation, characterization and transfection efficiency of cationic PEGylated PLA nanoparticles as gene delivery systems, J. Biotechnol. 130 (2) (2007) 107–113. [140] K. Tahara, S. Samura, K. Tsuji, H. Yamamoto, Y. Tsukada, Y. Bando, H. Tsujimoto, R. Morishita, Y. Kawashima, Oral nuclear factor-κB decoy oligonucleotides delivery system with chitosan modified poly(D, L-lactide-co-glycolide) nanospheres for inflammatory bowel disease, Biomaterials 32 (2011) 870–878. [141] H. Mimi, K.M. Ho, Y.S. Siu, A. Wu, P. Li, Polyethyleneimine-based core-shell nanogels: a promising siRNA carrier for argininosuccinate synthetase mRNA knockdown in HeLa cells, J. Control. Release 158 (2012) 123–130. [142] K. Raemdonck, B. Naeye, A. Høgset, J. Demeester, S.C. De Smedt, Prolonged gene silencing by combining siRNA nanogels and photochemical internalization, J. Control. Release 145 (2010) 281–288.

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Macromolecular therapeutics.

Enhancing Endosomal Escape for Intracellular Delivery of Macromolecular Biologic Therapeutics.

Immunogenicity of coiled-coil based drug-free macromolecular therapeutics.

Jotham Musinguzi: delivering for Africa.

Which model for delivering care?

Clinical therapeutics for phenylketonuria.

New therapeutics for osteoporosis.

Oxygen Delivering Biomaterials for Tissue Engineering.

gum acacia bipolymeric nanohydrogels--promising carrier for zinc oxide nanoparticles.

Enzyme therapeutics for systemic detoxification.

Splicing therapeutics for Alzheimer's disease.

Newer therapeutics for hepatitis C.

Reference materials for cellular therapeutics.

Next-Generation Therapeutics for IBD.

Emerging therapeutics for Alzheimer's disease.

Therapeutics: Bacterial treatment for cancer.

Delivery materials for siRNA therapeutics.

RNA-targeted Therapeutics for ALS.

Therapeutics.

Methods for delivering and evaluating the efficacy of cognitive enhancement.

The challenges of delivering validated personas for medical equipment design.

Delivering bad news to patients.

Therapeutics.

Delivering protein drugs orally.