Journal of Controlled Release 193 (2014) 162–173

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Journal of Controlled Release

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Polysaccharide-based micro/nanohydrogels for deliveringmacromolecular therapeutics

Kuntal Ganguly, Kiran Chaturvedi, Uttam A. More, Mallikarjuna N. Nadagouda, Tejraj M. Aminabhavi ⁎Department of Pharmaceutical Engineering and Chemistry, SET’s College of Pharmacy, S.R. Nagar, Dharwad 580 002, India

⁎ Corresponding author. Tel.: +91 836 2448540; fax: +E-mail address: aminabhavit@gmail.com (T.M. Amina

http://dx.doi.org/10.1016/j.jconrel.2014.05.0140168-3659/© 2014 Elsevier B.V. All rights reserved.

a b s t r a c t

a r t i c l e i n f o

Article history:Received 11 March 2014Accepted 7 May 2014Available online 17 May 2014

Keywords:MacromoleculesPeptides/proteinsMicrospheresNanohydrogelsGenesHormones

Increased interest in developing novel micro/nanohydrogel based formulations for delivering macromoleculartherapeutics has led to multiple choices of biodegradable and biocompatible natural polymers. This interest islargely due to the availability of large number of highly pure recombinant proteins and peptides with tunableproperties as well as RNA interference technology that are used in treating some of the deadly diseases thatwere difficult to be treated by the conventional approaches. The majority of marketed drugs that are now avail-able are in the form of injectables that pose limited patient compliance and convenience. On the other hand,micro/nanotechnology basedmacromolecular delivery formulations offermany alternative routes of administra-tion and advantages with improved patient compliance and efficient or targeted delivery of intracellular thera-peutics to the site of action. This review outlines and critically evaluates the research findings on micro andnano-carrier polymeric hydrogels for the delivery of macromolecular therapeutics.

© 2014 Elsevier B.V. All rights reserved.

1. Introduction

The development of recombinant macromolecular therapeutics hasgrown quite rapidly over the past decade due to the advent of peptideand protein drugs [1]. In recent years, macromolecular therapeuticssuch as proteins, peptides, small interfering RNA (siRNA), vaccinesand hormones have emerged as a significant class of medicine used inthe treatment of various deadly diseases. With more than 130 FDAapproved products available today in the market and many more inthe pipeline, such drugs are gaining a significant importance in almostevery discipline, such as cancer therapy, inflammatory disease, vaccines,and as diagnostics. These drugs have numerous advantages over thesmall-molecule generic drugs, since they are highly specific and exhibita complex set of functions such as biochemical reactions, protein-basedmembrane receptors and channels, cellular or organ transport of mole-cules and transcellular scaffolding support for which small syntheticmolecules can hardly mimic [2].

Macromolecular drugs do not easily cross the mucosal surfaces andbiologicalmembranes, since these are susceptible to loss of native struc-ture through cleavage of peptide bonds and destruction of amino acidresidues (e.g., proteolysis, oxidation, deamination, and elimination)and conformational changes due to the disruption of non-covalent in-teractions such as aggregation, precipitation, and adsorption. Special-ized uptake mechanisms like transmucosal M-cell uptake in Peyer’spatches and other lymphoid tissues may be necessary to transport

91 836 2467190.bhavi).

such water-soluble macromolecules through mucosal surface to sys-temic circulation, since these are prone to rapid clearance in liver aswell as other body tissues and may require accurate dosing [3]. Poly-meric (especially those of polysaccharide-based)-based deliverysystemswill diminish the inherent instability of these drugs to improvetheir bioavailability after administering through oral, nasal, pulmonaryand other routes [4].

Presently, protein drugs and antigens are administered parenterallyi.e., by subcutaneous (sc) or intramuscular injections as well as intrave-nous (iv) infusions, but these pose problems of oscillating drug concen-trations [5]. Drugs like growth hormone, insulin, oxytocin, parathyroidhormone, and vasopressin have short half-lives of b25 min [6], whichnecessitate multiple injections per week causing the compliance issues,especially when long-term treatment is required as in the treatmentof diabetes mellitus by insulin. These drawbacks impose immensechallenges and opportunities for developing delivery vehicles usingbiopolymeric hydrogels.

Among the various approaches, researchers have developed needle-free administration routes with high bioavailability such as pulmonary,oral, and nasal delivery [7–9]. Other approaches include extending cir-culation time and masking immunogenicity of protein drugs throughconjugation with other biopolymers as well as developing injectableor transmucosal controlled release (CR) systems including liposomes,polymeric micro/nanoparticles, and hydrogels [4]. Therefore, develop-ment of efficient micro/nanocarrier-based delivery systems providestremendous opportunities for improving the patient compliance andpharmaco-economic benefits. This review compiles the literature onsuch materials since 2000 until now. The current status and future

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

2. Hydrogels

Hydrogels are the cross-linked 3Dnetwork structures prepared fromhydrophilic polymers that are capable of retaining a large amount ofwater and remain insoluble due to their physical and/or chemicalcross-linking. Since their early development in the 1960s [10], innumera-ble hydrogels are available for a wide range of pharmaceutical applica-tions [11]. These hydrogels can be prepared from both natural andsynthetic polymers composing homopolymers, copolymers, and inter-penetrating polymer networks (IPNs) by selecting proper building blocksas well as using relevant cross-linking approaches [12–14].

Hydrogels possess highwater content and are soft networks, resem-bling those of natural extracellular matrices that minimize the tissueirritation or cell adherence [15]. The high loads of water-soluble thera-peutically active proteins, peptides, siRNA, DNA, vaccines, etc., can beencapsulated into their 3D networks, due to their porous structurealong with retaining high water content. Unlike other delivery systems(microparticles, emulsions, etc.), where preparation conditions aresometimes detrimental to proteins (i.e., use of organic solvents andprotein denaturing processes like homogenization, exposure to inter-faces, and shear forces), hydrogel preparation procedures are beneficialto preserve the protein stability as mild conditions such as aqueousenvironment and room temperature are commonly employed. Theirunique properties have created increasing interest in developing theCR systems for proteins/peptides to maintain therapeutic plasma con-centrations in the surrounding tissues or in circulation for longer time.

Hydrogels can be prepared from natural as well as synthetic poly-mers. Chemically cross-linked networks have permanent junctions,while physical networks have transient junctions that arise from eitherpolymer chain entanglements or physical interactions such as ionic in-teractions, H-bonds or hydrophobic interactions. The physical appear-ance as matrix, film or microsphere depends on the polymerizationmethod used to prepare hydrogels. Hydrogel networks are also basedon 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 poly-mers include proteins such as collagen, gelatin and polysaccharides likestarch, sodium alginate (NaAlg), chitosan (CS), and agarose.

Hydrogels based on homopolymers consists of a single monomerwith a cross-linked skeletal structure,while copolymer-based hydrogelsare formed from two or more different types of monomers with at leastone hydrophilic component arranged in a random, block or alternatingconfiguration along the polymer backbone [16]. On the other hand, IPNsaremade from two independently cross-linked synthetic and/or naturalpolymers in a network, while in a semi-IPN hydrogel, one polymer com-ponent is cross-linked, and the other is not. Various such structures aredepicted in Fig. 1.

Fig. 1. Types of hydrogel network structures

The stimuli-responsive hydrogels that can undergo volume transi-tions in response to various physical stimuli such as temperature, elec-tric or magnetic field, light, pressure, and sound, as well as chemicalstimuli like pH, solvent composition, ionic strength, and molecularspecies, are widely employed in drug delivery area [12,17]. Uponremoving the external stimuli, swollen hydrogel contracts to theunswollen state. Stimuli-responsive hydrogels offer remarkable pros-pects for the delivery of macromolecules and genes as the carriers areactive contributors to optimize the therapy instead of a passive deliveryvehicle.

Recent interest in cellmechanics and effects of substrate elasticity oncell structure as well as its function together with the ability of synthe-sizing the novel polymers that approximate the material property ofbiological tissues has motivated research on different materials for usein wound healing and tissue engineering [18]. Synthetic gels preparedfrom polyisocyanopeptides grafted with oligo(ethylene glycol) sidechains reportedly mimic gels prepared from intermediate filaments inalmost all aspects. These responsive polymers have a stiff and helicalarchitecture with a tunable thermal transition where the chains bundletogether to generate transparent gels at extremely low concentrations.These materials show a very fast sol–gel phase transition that can becompletely reversible. However, the ease ofmodification of thesemate-rials provides avenues for the preparation of functional biomimeticmaterials required in biomedical applications [19].

3. Polysaccharide-based hydrogels

Naturally occurring polysaccharides are frequently used in thedelivery of macromolecular therapeutics as these are highly biode-gradable and biocompatible, and can be prepared as conjugates orcomplexes with proteins, peptides and other biomacromolecules.Among the widely investigated polysaccharides, NaAlg, chondroitinsulfate, CS, and hyaluronic acid (HA) are the prime candidates. Specifi-cally, these polysaccharides in combination with other polymers offerthe desired chemical and/or biological advantages. Some representativemembers of this class are discussed briefly here, but details can be foundelsewhere [20].

3.1. Chitosan

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

used in macromolecular drug delivery.

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

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produced by an ionotropic gelation method with no major conforma-tional changes of insulin in terms of α-helix and β-sheet content [31].

Polyelectrolyte complexes of CS with other biopolymers [30,32], folicacid-conjugated NPs of CS for targeted delivery of 5-aminolevulinic acid(5-ALA) to colon [33] as well as CS-tamarind kernel powder inter-polymer complex films have been used for colon delivery [34]. Chemicalcross-linking of CS with polyethylene glycol (PEG) produced a swellabledelivery system in both acidic and alkaline pH media [35], but cross-linking of CS with glutaraldehyde (GA) has the disadvantage of toxicity.Alternatively, genipin, a natural aglycone, was found to have minimumtoxic effects [36]. Among the many CR systems of CS, pH-sensitive CS-Eudragit L100-55 NPs prepared by a coacervationmethod using high vis-cosity HPMC was useful for the delivery of insulin through gut mucosa.See recent reviews on CS hydrogel delivery systems [2,8].

3.2. Dextran

Dextran (DEX) is a complex polysaccharide consisting of α-1,6-linked D-glucopyranoses with some degree of 1,3-branching. Being hy-drophilic, it can be crosslinked by various methods to be effective as ahydrogel carrier system [37]. Van Tomme et al. [38] published an exten-sive review on DEX-based hydrogels for protein delivery. Peptide cross-linking resulted in enzyme-dependent degradation controlled by cell-secreting enzymes, thus mimicking the degradation of the naturalextra cellular matrix [39]. Taking these advantages, self-assembledNPs of quaternized CS (N-(2-hydroxyl) propyl-3-trimethyl ammoniumCS chloride (HTCC)) and DEX sulfate were prepared by ionic-gelationthat showed internalization into Caco-2 cells without loss of cell viabil-ity. This system demonstrated fast release of therapeutics in phosphatebuffer 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 sphericalbovine serum albumin (BSA)-Zn2+ NPs of 50–350 nm size under mildcondition by adding BSA and zinc acetate to a solution of PEG followedby freeze-drying. Here, no protein aggregation was observed and dueto freeze-induced phase separation, the BSA–Zn2+ complex wassqueezed into dispersed particle giving spherical NPs. The same proto-col was followed to prepare recombinant human growth hormone(rhGH)–Zn2+− DEX NPs [42] by vortexing a mixture of rhGH and zincacetate added to DEX-PEG solution followed by freeze-drying forovernight at−20 °C. The lyophilized powder was washed with dichlo-romethane and centrifuged to remove PEG to obtain NPs of size 20–170 nm with N90% encapsulation efficiency (EE) and these showedNb2-11 cell proliferation activity.

An almost identical procedure was adopted for encapsulatinggranulocyte-macrophage colony-stimulating factor (GM-CSF), granulo-cyte colony-stimulating factor (G-CSF), β-galactosidase, myoglobin(MYO), and BSA into DEX NPs by aqueous–aqueous freezing-inducedphase separation [43] as shown in Fig. 2 to produce NPs of 200–500 nmsize for encapsulatingBSA that showed N98% EEwith noproteinaggregation 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 usedaminoalkyl VB12 derivatives synthesized at 5′-hydroxy ribose ande-propionamide sites via carbamate and ester/amide linkagescoupled to succinic acid modified DEX NPs with varying extent ofcross-linking that showed 70–75% reduction in blood glucose level(BGL) tomaintain anti-diabetic effects up to 54 h in STZ-induced diabeticrats. NPs with low levels of cross-linking were superior and more effec-tive with VB12 derivatives of carbamate linkage. The pharmacologicalavailability relative to sc insulin was around 29% that was superior(1.5-fold) to NPs conjugate of ester linked VB12. The NPs demonstrateda similar oral insulin efficacy in congenital diabeticmice (60% BGL reduc-tion in 20 h).

The microcapsules of thermo-responsive poly(N-isopropylacrylamide)(PNIPAAm) were prepared [45] and used for the CR of stromal cell-

derived factor (SDF)-1α, an important chemokine for stem cellrecruitment/homing. Here, the double-phase emulsified condensationpolymerizationwas used to prepare the interconnected porous glycidylmethacrylated dextran (DEX-GMA)/gelatin microcapsules by adoptingplasma-graft pore-filling polymerization to graft PNIPAAm onto thesurface pores of microcapsules. The in vitro results of these microcap-sules exhibited thermo-responsive release of drug in response totemperature variations. The sc implantation of these microcapsulesresulted in a sustained and long-term release of SDF-1α compared tothose formulations without PNIPAAm-grafting.

A new multicomponent delivery system based on DEX, protamine(Prot), and solid lipid nanoparticles (SLN) containing pCMS-EGFPplasmid were investigated for clathrin/caveolae dependant transfectionefficiency [46]. These surface modified SLNs were tested for erythrocyteinteraction and potential agglutination. The insulin-loaded NPs of500 nm produced from complexation of DEX and CS [47] showedgood stability with zeta potential value of−15mV under optimal com-position of DEX:CS (mass ratio of 1.5:1) at pH 4.8with almost no insulinrelease at pHb5.2 up to 24h, but insulin release occurred in pH 6.8, suit-able for oral delivery.

In a continuing study, the authors [48] used these systems todeliver insulin in a pH dependent manner to achieve associationefficiency of 70% with 24-h sustained release in rat models with abioavailability of 5.6% at 50 IU/kg dose compared to insulin solution.In another study [49], cationic character was imparted to DEX byconjugating spermine to oxidized DEX by reductive amination. Self-assembly of cationized DEX and CXCR4-siRNAs produced 55 nm sizeNPs with a zeta potential of 40 mV, which significantly downregulatedCXCR4 expression (tested in colorectal cancer metastasis in Balb/cmice) through CXCR4 silencing.

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3.3. Hyaluronic acid

Hyaluronic acid (HA) is a naturally occurring anionic linear glycos-aminoglycan made up of repeating disaccharide units of D-glucuronicacid and N-acetylglucosamine having MW up to 107 Da. It is a majorcomponent of synovial fluid, and is present primarily in extracellularmatrix (ECM) due to its viscoelastic properties that helps to reducethe friction between bones [50]. Delivery of proteins and peptides hasbeen attempted with photopolymerized HA and HA-tyramine conju-gates cross-linked by disulfide bond formation [51–53]. The negativecharge of HA hinders protein release rate due to the charge of protein,but enzymatic degradation of HA by hyaluronidase alleviates thisproblem.

Even though there are no commercially available delivery productsusing HA, except Declage, it is considered to have a greater potentialto develop as conjugate or even as physically or chemically cross-linked hydrogel depot system [54]. Microparticle formulation usingHyaff11p50, where 50% of carboxyl groups of HA are esterified withbenzyl alcohol, was prepared by solvent evaporation and spray-dryingmethod for a 7-day release of hGH [55]. Another formulation of hGH/HA/lecithin prepared by spray-dryingmethod resulted in small particlesto be injected through a 26-gauge needle, and showed improved patientcompliance [56].

Novel hybrid hyaluronan-based nanogels (average size b160 nm)[57] were suggested for spontaneous encapsulation of proteins andpeptides that showed the CR profiles without altering chemical stabilityof macromolecular therapeutics. HA and ferric oxide were ionicallygelated to form hybrid nanoparticles for peptide delivery. Hybrid mate-rials such as cyclodextrin (used as porogen) coupled porous microparti-cles of HA ionically complexed with lysozyme were also developed forprotein delivery that showed improved EE and protective nature [58].Thus, HA-based systems appear to be more suitable than poly(lacticacid-co-glycolic acid) (PLGA), as it ismore biocompatible andhydrophilicin nature

3.4. Sodium alginate

NaAlg, an unbranched polysaccharide consisting of 1-4′-linkedβ-D-mannuronic acid (M) and α-L-guluronic acid (G) moieties invarying compositions [59], is capable of forming strong hydrogelsby simple addition of metal ions to its aqueous solution. Such a mildmethod of preparation makes the hydrogels very suitable to encapsu-late living cells for investigating their CR properties. NaAlg hydrogelswere also used for tissue regeneration, while simultaneously releasinggrowth factors or cytokines [60,61]. Some representative novel micro/nanocarriers of NaAlg will be discussed.

NaAlg-based NPs developed as pH-responsive delivery vehiclesfor insulin prepared by spray-drying, ionic crosslinking by electro-hydrodynamic spraying and solvent diffusion methods have shownthe EE values ranging from 38 to 90% [62,63]. Their pH dependent CRproperties protected the integrity of insulin at higher temperaturesduring spray-drying process. Sharma et al. [64] developed biodegradablePVA and NaAlg electrospun composite nanofiber-based transmucosalpatches for sublingual insulin delivery. Their highwater holding capacityresulted in high insulin loading that improved the mucoadhesivity withEE of 99%.

NaAlg-based microcapsules added with poly-L-lysine and poly-L-or-nithine crosslinked under UV radiation were also used for connexin-43carboxyl-terminus mimetic peptide (CT1) delivery to the wound site. Astudy [65] on NaAlg-based micro-devices for BSA (66 kDa) showed thedependence of release onMW and particle size distribution. Mahidharaet al. [66] developed novel NaAlg-CS coated ceramic nano-carrier loadedwith multi-functional anti-cancer bovine lactoferrin (Lf), a natural milkbased protein, to improve its intestinal absorption. These systemsshowed the size dependent endocytosis and transcytosis of NPs incolon cancer cell-lines.

Recently, affinity-based growth factor delivery system was devel-oped by incorporating heparin into photocrosslinkable NaAlg hydrogels(HP-NaAlg) that prolonged the release of therapeutic proteins [67].Heparin modification showed minimal biodegradation and the releaseof 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 forma-tion around the implant periphery. Importantly, BMP-2-loaded HP-NaAlg hydrogels induced significantly more osteogenesis thanBMP-2-loaded unmodified HP-NaAlg hydrogels with 1.9-fold greaterperipheral bone formation and 1.3-fold greater calcium content in BMP-2-loaded HP-NaAlg hydrogels than the BMP-2-loaded unmodifiedHP-NaAlg hydrogels even after 8 weeks of implantation.

El-Sherbiny et al. [68] developedwater-soluble copolymer of sodiumacrylate grafted onto carboxymethyl cellulose (CMC). When these wereused along with NaAlg as a pH-sensitive IPN hydrogel prepared byionotropic gelation in the presence of Ca+2 showed b18% 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 emulsificationtechnique in the presence of Ca+2 and CS solidification achieved an EEof 57% and their in vitro release under blood pH showed stable andsustained release up to 14 days and these formulations protected thechemical stability of insulin released. In vivo studies showed a reductionof BGL of diabetic rats stable up to 60 h after oral administration ofinsulin-loaded microspheres. However, pH-responsive, biodegradablemicroparticle carrier system based on ionotropically cross-linked mix-ture of NaAlg and chemically modified carboxymethyl CS coated [70]through polyelectrolyte complexation with CS grafted PEG showed therelease ranging from 32 to 83.1%.

Many types of modified NaAlg matrices have [71] been mixed withother polymers to obtain either IPNs or covalently linked mixed poly-mer systems to improve the stability of hydrogels along with the CR ofencapsulated proteins [72]. An elegant example is simple sulfonationof uronic acids in NaAlg that specifically binds heparin-binding proteins,including growth factors to allow the CR of these proteins and thusprovide enhanced therapeutic activity as demonstrated in a murinehindlimb ischemia model in rats [73].

Goycoolea et al. [74] developed the NPs for transmucosal deliveryof macromolecules by ionic gelation of CS hydrochloride withpentasodium tripolyphosphate (TPP), complexed with NaAlg thatexhibited 41–52% of EE. These showed enhanced systemic absorptionof insulin after nasal administration to conscious rabbits. However,some novel composite NaAlg/PLGA microparticles [75] could deliverbovine insulin exhibiting reproducible EE values for insulin, releasingup to 4 months, while NaAlg nanohydrogels loaded with insulinprepared [76] by a reverse emulsification-diffusion method showedthe CR profile for tissue or organ targeting for DNA/gene delivery.The NPs of PLGA [77] imparted pH sensitivity when prepared usinghypromellose phthalate by amodified emulsion solvent diffusionmethodand showed relative bioavailability of 7%.

4. Micro/nanohydrogels

Hydrogel-basedmicro/nanoparticles are promising formacromolec-ular 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 mediatedimmune responses and high drug loading. As noted before, such micro/nanohydrogel systems can be prepared by many different techniquesincluding solvent evaporation, spontaneous emulsification/solventdiffusion, salting out/emulsification-diffusion, supercritical fluid tech-nology, spray-drying, ionic gelation, micelle, and reversemicelle forma-tion [78]. Self-assembled nanohydrogels are particularly attractive,since these are easy to prepare, are affordable and can effectively incor-porate biopharmaceuticals like biosimilars, proteins and peptides. Therelease of cargos from such systems can be fine-tuned by tailoring

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cross-link density of the matrix. Other strategies to tailor drug releasefrom hydrogels rely upon reversible protein–polymer interaction orencapsulation 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 sizedistribution (40–120 nm size) [80]were employed in protein (interleu-kin-12 (IL-12)) delivery by cross-linking acrylate group-modifiedcholesterol-bearing pullulan nanohydrogel (CHPANG) with thiolgroup-modified PEG (PEGSH) through Michael addition (see Fig. 3).These nanohydrogels showed the EE value of 96% for IL-12, witha steady plasma IL-12 level in mice up to 72 h following the scadministration.

Surface characteristics of the NPs play a predominant role in theiruptake by liver, spleen and RES. The NPs containing more hydrophobicsurfaces are preferentially taken up by liver, followed by spleen andlungs [78,81]. On the other hand, hydrophilic NPs (35 nm) preparedfrom poly(vinyl pyrrolidone) (PVP) displayed only b1% uptake byspleen and liver and about 5–10% of these could circulate in blood-stream even after 8 h of iv injection; however, the NPs (45–126 nm)prepared using 50% vinyl pyrrolidone and 50% N-isopropyl acrylamideshowed preferential uptake by the liver [82].

Ligand-mediated active targeting has emerged as a novel strategy tosupport the effectiveness of nanocarriers to improve the delivery oftherapeutics. Among different receptors expressed over gastrointestinaltract (GIT), folic acid (FA) receptors were reported to be present insufficient quantity to improve the uptake and transport of bioactive orvesicular systems across the GIT [83]. These FA receptors are alsoover-expressed in various tumor types and thereby target to varioustumors [84]. Aptamers that are oligonucleic acids or peptides are alsoused to deliver macromolecular drugs (siRNA, proteins and peptides),but these are used in conjunction with ligand-based target moleculessuch as transferrin and FA to acquire better therapeutic efficacy.Hydrogel-based nanocarriers of CS, due to their positive charges, arespecifically suitable for ligand/aptamer appending as well as encapsu-lating negatively charged siRNA that are critical to increase the cargoloading and EE.

5. Delivery routes

Despite the significant advances in macromolecular drug deliverythrough different approaches, oral route continues to be themost exten-sively studied approach for therapeutic delivery mainly due to the con-venience and patient compliance [85]. The large MW, 3D structure andhigh aqueous solubility of macromolecules pose unique challenges forabsorption through transmucosal route. A number of delivery systemsincluding pH-sensitive hydrogels and intestinal M cell targeted devicesthat utilize paracellular and transcellular absorption pathways havebeen developed for protein/peptide drugs in order to prevent the delete-rious effect of GIT [86]. However, restricted and delayed absorption as

Fig. 3. Schematics of the formation of cholesterol-bearing pullulan nanohydrogel structure (

well as variable intestinal motility and gastric emptying time wouldcreate 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 drugdelivery. However, hydrogel-based parenteral delivery of macromolec-ular therapeutics have some unique advantages such as controlledrelease, depot therapy, biocompatible, biodegradable and targetingcapability using ligands, aptamers, etc. The drawbacks such as decreasedpatient compliance, temperature sensitiveness and tissue damage causedby parenteral therapy pose problems for chronic therapy [90].

6. Delivery of vaccines and immunomodulatory agents

Transmucosal delivery of vaccines using hydrogel-based nanocarriersystems have been explored due to their advantages such as efficientencapsulation, protecting the drug from pH and enzymatic degradation,mucosal adjuvant incorporation to enhance immune response andtargeted delivery to mucosal inductive sites [91–93]. Even thoughadverse pH and enzymatic conditions of GIT pose problems for oraldelivery of vaccines, these systems are quite useful, since they are pre-pared by spray-drying, double emulsion, ionic gelation/polyelectrolytecomplexation, and phase separation methods that are more suitablefor vaccine delivery as these methods protect the 3D structure ofvaccines and antigens. Physicochemical properties of hydrogel-basednanocarriers are also critical for effective vaccine delivery. For orallyadministered NPs, amean size of around 500 nm showed better uptake,since the M-cell membrane is negatively charged, so a positive zetapotential 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 muco-adhesive property due to ionic interactions between its quaternaryamino (NH3+) and negative sialic acid group of mucin; thus, it canopen the tight junctions of intestinal epithelium. Various derivatives ofCS (e.g. trimethyl CS (TMCS), carboxymethyl CS (CMCS), and thiolatedCS) have been prepared to impart more mucoadhesivity and higherpenetration enhancement capability [94]. Hybrid materials such as HA(polysaccharide) and oligosaccharide (e.g. cyclodextrin and cyclodextrinderivatives)-based NPs showed better biocompatibility, in vivo stabilityand pharmacological efficacy. Various self-hydrolyzing hydrogels basedon Michael-type addition of PEG-diester-dithiol or non-degradable PEG-dithiol cross-linkers onto 4-arm PEG-vinyl sulfone have also been usedas tunable delivery systems for antigens like immunoglobulins. These sys-tems had a gel mesh size of 13–35 nm and showed controlled diffusivitybased on factors such asmolecular weight of drug, polymer and chemicalstructure of the cross-linker.

The CS and CS-based NPs [95–102] showed both systemic andmucosal immune responses when delivered through the nasal route,but mucoadhesivity of CS-based NPs leads to prolonged residencetime and increased M cell uptake. Alginate coated CS and TMCS NPswere developed for nasal [103,104] and oral [104–106] delivery to

A-CHPNG) assemblies. Reprinted from Ref. [80], © 2014, with permission from Elsevier.

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impart pH responsiveness as well as to control burst release of antigenfrom the NPs. For instance, nasal delivery of alginate MPs with tetanustoxoid [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 barrierand those related tomucosal immune stimulation aswell as their corre-lation with NP characteristics could result in improved NP delivery.However, targeted delivery approach to cellular and subcellular specificdeliverywould bemore beneficial as it leads to an optimum therapeuticNP system for vaccine delivery.

7. Delivery of hormones

While delivering thehormone, it becomes necessary to ensure safetyand therapeutic efficacy to reduce the frequency of administration andto enhance the patient compliance and comfort. Various novel deliverysystems have been explored for human growth hormone (hGH) delivery.Even though hGH itself is secreted in a pulsatile manner, continuous in-fusion of hGH via pump may not provide the required clinical efficacyas it elevates insulin-like growth factor-1 (IGF-1) levels similar to thatof daily injections, indicating that pulsatile hGH delivery is not suitable[108]. However, the most commonly used formulations for hGH areoral tablets and transdermal patches for sex steroid hormones as wellas intramuscular or sc injections [109].

Other approaches like crystal formulation, PEGylation, and loadingof 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 consciousrat model. Here, absorption-enhancing component is Solutol HS15,which consists of polyglycol mono- and di-esters of 12-hydroxystearicacid combined with free PEG, which showed 49% bioavailability in 2 hfor the ratio of Solutol HS15 to hGH of 4:1 (mg/mg) at which no nasalmucosa toxicity was found even up to 6 months [112].

Even though PLGA-based microsphere formulation, viz., Nutropindepot,was thefirstmarketedCRproduct of hGH, its highmanufacturingcosts led to its withdrawal from the market. This system showedsustained release of hGH up to 1 month [113], but suffered from lowEE, high initial burst release, protein aggregation, denaturation, and

Fig. 4. Dual ionic interactionmodel based on hGH loaded polyelectrolyte complex and anionic hfrom Elsevier.

inflammation due to the use of organic solvents.Moreover, acidic degra-dation products [114] were observed with this system. To overcomethese drawbacks, zinc was incorporated into these microspheres tostabilize hGH [108,111], since Zn+2 would induce dimerization ofhGH, and hence, the reversible complex would be more stable thanmonomeric hGH. However, irreversible aggregation of hGH in thePLGA microsphere was effectively reduced by the zinc–hGH complexand the CR of hGH was observed compared to daily injection.

The only commercially available CR formulation of hGH ishyaluronate-based microparticle formulation launched as a once-a-week injection formulation (Declage) in Korea by LG Life Science in2007. This system showed improved drug loading (ca. 20%) and bioavail-ability compared to PLGA microspheres, which maintained serum hGHlevel for 30 h in cynomolgusmonkeys [56,115]. However, injectable, bio-degradable, and thermosensitive hydrogel-based micro/nanocarriersseem to be the better choices for hGH delivery, since their high watercontent and temperature dependent gelation properties without theuse of organic solvents or chemical cross-linkers made these systemsas efficient hGH delivery carriers [116].

Park et al. [117] developed dual ionic interaction nanohydrogel sys-temcomposed of a positively chargedpolyelectrolyte complex (PEC) con-taining hGH and anionic thermosensitive poly(organophosphazene)hydrogel to enhance the CR of hGH (see Fig. 4). These nanodevices of500 nm size and zeta potential of +8 mV suppressed the initial burstrelease showing better in vitro and in vivo correlation with 13-foldincrease in AUC compared to pristine hGH solution and enhancedbioavailability in hypophysectomized rat model.

Pyo et al. [118] used an entirely different approach, i.e., solution en-hanced dispersion by supercritical fluids (SEDS), to produce nano-sizedrecombinant hGH using ethanol to help supercritical CO2 to extractwater from the aqueous protein solution. Various sizes of hGH NPswere prepared by this method with a narrow particle size distributionfrom the aqueous ethanol solutionwithout using any additive. However,these novel delivery systems require rigorous in vivo pre-clinical, clinicalsafety and efficacy testing data prior to further development as usefultherapeutic products.

Tang et al. [119] used an emulsion cross-linking technique to pre-pare CMCS and HA conjugate hydrogel microspheres for delivering

ydrogel for sustained delivery of hGH. Reprinted from Ref. [117], © 2014, with permission

Table 1Micro/ nanohydrogels for hormone delivery.

System Size (nm)/zetapotential (mV)

Drug Ref.

PEC hGH-PS inpoly(organophosphazene)

500/+8 hGH [117]

Recombinant hGH-NP by SEDSmethod

50 hGH [118]

HA-carboxymethyl CS conjugate 4100–5900/−17 to −24 Catalase [119]PLGA-CS 430–590 sCT [122]LEC-CS 122–347/8–33 Melatonin [123]LEC-CS NP transdermal suspension 113−129/−9 to +13 Melatonin [124]CS NPs 267/+35 sCT [125]CS-coated tripalmitin NPs 538/+29 sCT [126]PEG-coated tripalmitin NPs 226/−35 sCT [126]PEG-coated tripalmitin/Miglyol NPs 207/−37 sCT [126]HA–CS nanocomplexes 163–193/−32 to −75 sCT [127]CS NP (spray-dried) 215/+28 sCT [128]

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.

168 K. Ganguly et al. / Journal of Controlled Release 193 (2014) 162–173

antioxidant enzyme catalase through oral route to prevent enzymaticdegradation. Cross-linking reduced the swelling to increase the re-sistance to hyaluronidase digestion of these systems. The encapsu-lated catalase exhibited superior stability over wide pH ranges(pH 2, 6 and 8) compared to the native enzyme. These catalase-loaded microspheres, in contrast to native catalase, effectively de-creased the intracellular H2O2 level and protected the HT-29 colonicepithelial cells against H2O2-induced oxidative damage to preservethe cell viability.

Salmon calcitonin (sCT), a therapeutic analogue of calcitonin, a cyclicpolypeptide hormone, is another important drug used to prevent oste-oclastic bone resorption (potent hypocalcemic). It is mainly availableas injection and nasal spray formulation in the market, but these havemet with several disadvantages. The oral bioavailability of sCT is b0.1%due to extensive proteolytic degradation in the GIT and poor perme-ation across intestinal epithelial cells [120]. The hydrogel-based NPsusing thiomer derivative of glycol chitosan (GCS) synthesized by cou-pling with thioglycolic acid (TGA) have been evaluated for pulmonarydelivery of peptides [121]. This study reported the NPs (200–300 nm)prepared from GCS and GCS-TGA by ionic gelation with a net positivesurface charge that have shown high calcitonin encapsulation. On theother hand, CS surface modified sCT loaded PLGA NPs (430–590 nm)prepared [122] by w/o/w emulsification and solid dipping methodsshowed the EE of 50% with improved sustained release, showingshort-period hypocalcemic effect.

Themucoadhesive NPs [123] prepared from lecithin/CS investigatedfor transmucosal delivery of melatonin showed no cell membranedamage and were non-cytotoxic. Transdermal delivery of melatoninwas also studied [124] using the NPs with size differing in lecithintype (Lipoid S45 and S100) and CS content that ranged from 114 to332 nm with zeta potential of 4.6–31.2 mV. The study indicated7.2% of melatonin loading with 1.3- to 2.3-fold higher flux comparedto melatonin solution; the highest flux of 9.0 μg/cm2/h was achievedwith lecithin/CS for a weight ratio of 20:1 and these NPs were safeto use at concentration up to 200 g/mL for skin application. Asimilar study by Prego et al. [125] highlighted the concentrationdependent improved cellular uptake and internalization of NPsloaded with sCT. The in vivo data suggested that both CS-coatedsystems (neat NPs and CS-coated NPs) showed enhanced intestinalabsorption of sCT, thus enhancing the therapeutic efficacy in ratmodel.

More efforts have been made on developing the CR formulations ofsCT, of which sCT loaded solid lipid NPs coated with PEG or CS were in-vestigated to study the effect of composition of core shell on the deliveryof sCT [126]; according to this study, the nature of coating affected thesurface association of the peptide. The NPs covered by positivelycharged CS reduced the burst release that was more pronounced forNPs coated with PEG than those coated with CS. After the initial burstrelease, these systems showed continuous and slow release of peptide,independently of the nature of coating. The slow release was due to af-finity of peptide for lipids and the absence of degradation of lipidmatrixunder in vitro release conditions.

The intra-articular sCT-HA sustained release system using CSas electrostatic-based nanocomplex reduced the inflammationand inflammatory gene expression when injected through intra-articular route in a mice model [127] that showed a remarkableanti-inflammatory effect by reducing the NR4A2 mRNA expressionin vitro. However, inhalable co-spray dried powders of sCT-loadedCS nanoparticles (sCT-CS NPs) of 200 nm size prepared with mannitolshowed improved pulmonary absorption in rats [128]; the NPswere formed spontaneously after adding sCT in CS solution contain-ing TPP. For inhalation, NPs were co-spray dried with mannitol as acarrier, which displayed better aerodynamic properties for pulmonarydelivery, showinghigher protein absorption thannative sCT. Table 1 sum-marizes various micro/nanoparticle-based platforms used in hormonedelivery.

8. Gene delivery

Efficient in vivo delivery of genes remains the biggest challenge ofnucleic acid-based biopharmaceuticals, mainly because of their largemolecular size, hydrophilicity, negative charge, acid labile nature andease of degradation by nuclease, preventing an effective transportacross the intestinal epithelium causing low transfection efficiency viaoral route. As a result, parenteral route has been preferred for deliveringnucleic acids using non-viral vectors containing polymers, complexes ofcationic lipids or peptides with plasmid DNA [129]. Cationic polymer-based non-viral vectors have been commonly used as carriers fornucleic material delivery [2], and self-assembly of cationic polymerwith DNA (polyplex) that interacts with negatively charged cellularmembrane was also employed [130].

Polyplexes containing macromolecular drugs must escape fromendosomes for successful transport of nucleic material, since DNA isprone to degradation by lysosomal enzymes. After endosomal escape,polyplexes get into cytoplasm where they unpack DNA and deliver iteither near the nucleus or in the nucleus. The translocated DNA in thenucleus then causes gene expression [131]. To achieve this, the chosencarrier polymer should have a good DNA binding capability such thatit can condense DNA into polyplexes and high buffering capacity toenable endosomal escape as well as good intracellular vector unpackingto release DNA. In this regard, CS can effectively bind with DNA,protecting it fromnuclease degradation and these polyplexes are capableof adhering to intestinal epithelium andM cells targeted transport acrossthe mucosal boundary that can transfect epithelial and immune cellsassociated with the gut lymphoid tissue [132].

Among others, CS has a lower cytotoxicity than the commonly usedsynthetic polyethyleneimine (PEI) [133] and its endosomal disruptiveproperty, as well as proton sponge-type mechanism, ensures entryinto the cell. Upon internalization within the endosome and due tothe change in pH inside the cell, CS gets protonated and eventuallycauses the rupture of the membrane to release of the genetic materialinto the cell [2]. CS derivatives such as glycol, o-carboxymethyl,trimethylated, thiolated and 6-N,N,N-trimethyltriazole CS and its hydro-phobic modifications like deoxycholic acid, 5-β-cholanic acid, andN-acylated chitosan have all been used to overcome the limited CSsolubility to improve transfection efficiency.

A recent study on CS-basedNPs for nucleic acid delivery showed thattransfection 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 murineerythropoietin (Epo) gene. Over 25% increase in hematocrit levels wasobserved by delivering Epo genes (responsible for stimulating red

169K. Ganguly et al. / Journal of Controlled Release 193 (2014) 162–173

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 DNAin a mouse model with hemophilia A (a disorder of the blood coagula-tion cascade caused by defective factor VIII). These NPs led to higherplasmid copy numbers in Peyer's patch tissue compared to the nakedDNA delivery.

Zheng et al. [136] studied the in vitro and in vivo transfection effi-ciency of three types of CS NPs, i.e., quaternized CS-60% trimethylatedCS 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 ofthese indicated the highest transfection efficiency for NPs of TMCSO-60% followed by CS (43–45 kDa, 87%) and CS (230 kDa, 90%). In vivostudies indicated the most prominent GPF expression in gastric andupper intestinal mucosa. TMCSO-60%was themost efficient with betteractivity andminimal toxicitymaking it an efficient carrier for oral deliveryat optimal CS/pDNA ratio of 3.2:1. The pullulanNPs (45 nm) encapsulat-ed with nucleic acid (pBUDLacZ plasmid) showed encouraging resultswith gene expression comparable to commercial Lipofectamine 2000[137].

Zeng et al. [138] prepared PLGA-CS NPs by spontaneous emulsiondiffusion method for the delivery of pDNA that showed much higherEE and higher cellular uptake as well as higher hepatitis B virus (HBV)gene-silencing efficiency than plain-PLGA NPs and naked pDNA. Inorder to achieve prolonged delivery and high transfection efficiency ofcationic PLA/DNA complexes, Chen et al. [139] used the copolymer ofmethoxy polyethylene glycol-PLA (MePEG-PLA) to prepare MePEG-PLA-CS NPs as well as PLA-CS NPs by diafiltration method. The NPsof MePEG-PLA-CS had a zeta potential of 15.7 mV with a size ofN100 nm, while PLA-CS NPs had the zeta potential of 4.5 mV atpH 7.4. The transfection efficiency of MePEG-PLA-CS/DNA complexeswas better than PLA-CS/DNA and Lipofectamine/DNA complexes,which mediated higher gene expression in stomach and intestine ofBALB/C mice compared to PLA-CS/DNA and Lipofectamine/DNA com-plexes, making these as the efficient non-viral vectors for gene delivery.

Superior mucoadhesive character and prolonged pharmacologicaleffect 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 sulfateinduced experimental colitis. The positive zeta potential of these NPsenabled greater cellular uptake and was stable during incubation withnuclease (DNase I) and in simulated gastric fluid. Ionic complex forma-tion between ODN and CS on the NP surface prevented ODN degrada-tion due to acidic conditions and nuclease.

Fig. 5. Steps involved in the preparation of gelatin-PEI core-shell nanohydro

Recently, siRNA has emerged as a powerful therapeutic agent for thetreatment of diseases as well as repair of faulty genes, resulting in theproduction of faulty proteins [2,5,8]. The siRNAs are more stable tonuclease degradation than unmodified antisense oligonucleotides,since they are highly sequence specific and are required in relativelysmall doses, but their poor serum stability in systemic circulation andpoor cellular uptake makes the therapy a big challenge. Moreover,naked siRNAs are quickly excreted out of the kidney as their biologicalhalf life is b1 h, but their size, negative charge and inability to crosscellular membrane in unmodified form make gene knock-down all themore difficult [5]. Various synthetic (poly(l-lysine) or polyethylenimine(PEI) and their analogs) and natural polymers (CS, collagen, gelatin andtheir derivatives) have been used for their delivery.

Mimi et al. [141] prepared PEI-based nanohydrogels with gelatincore (200 nm) to encapsulate siRNA by a two stage process involvingthe preparation of gelatin NPs followed by conjugation with branchedPEI to impart cationic property (see Fig. 5). These NPs showed increasein transfection efficiency to HeLa cells from 41 to 84%. The deliveredsiRNA inhibited 70% of human argininosuccinate synthetase 1 (ASS1)gene expression. Another study [142] involving cationic DEX nano-hydrogels for encapsulating siRNA and photochemical internalizationhas shown the potential for gene silencing using intracellular vesicles asdepots for siRNA. This technique destabilized the endosomal vesicles toprolong 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 ex-pression and tumor volume.

Crosslinked CS NPs of b150 nm for encapsulating siRNA wereprepared by ionic crosslinking (CS to siRNA mass ratios of 10:1, 30:1and 50:1) to explore their potential to deliver siRNA to lungs via a jetnebulizer [144] to observe high EE and non-aggregation at the pH ofthe airway. Complete binding of siRNA to CS was possible at a ratio50:1 and high cell viability (N85%) was observed even at the highestCS concentration of 83 μg/mL. In another study, interpolyelectrolytecomplexes of CS/siRNA showed rapid uptake (1 h) of Cy5-labeled NPsinto NIH 3T3 cells, followed by accumulation for over 24 h [145].These showed knockdown of the endogenous enhanced green fluores-cent protein (EGFP) in both H1299 human lung carcinoma cells andmurine peritoneal macrophages (77.9% and 89.3% reduction in EGFPfluorescence, respectively). Nasal administration of these NPs resultedin effective in vivo RNA interference in bronchiole epithelial cells oftransgenic EGFP mice (37% and 43% reduction compared to mismatchand untreated control, respectively).

gels. Reprinted from Ref. [141], © 2014, with permission from Elsevier.

170 K. Ganguly et al. / Journal of Controlled Release 193 (2014) 162–173

Wei et al. [146] developed NPs of N-((2-hydroxy-3-trimethylammonium) propyl) CS chloride (HTCSC) for encapsulatingtelomerase reverse transcriptase siRNA for oral delivery. Here, the poly-mer coating protected siRNA fromenzymatic degradation to permeate tothe intestine. Based on these data, authors developed a “two-in-one”nano-complexwith both paclitaxel and siRNAencapsulated into one sys-tem, which simultaneously transported siRNA and paclitaxel (PTX) totumor cells and increased drug concentration, leading to better tumorsuppression. Similarly, pH-responsive nanocarriers of trimethylchitosan(TMC) and methacrylic acid (MAA) copolymer were reported for theoral delivery of siRNA [147], where incorporation of MAA into polyplexwas increased. A significant decrease in zeta potential for MAA-TMC-siRNA complex with greater transfection efficiency (in L929 cells) thanTMC-siRNA complex was observed.

Guanidinylated chitosan (GCS) formed stable complexes with plas-mid DNA under physiological pH that showed lower cytotoxicity withhigher transfection efficiency than CS and 8-fold increase in cellularuptake [148,149]. The authors tested these for siRNA delivery to lungs,since GCS was able to condense siRNA at a weight ratio 40:1, formingNPs of diameter ~100 nm. Guanidinylation helped to enhance siRNAgene silencing activity than the pristine CS due to better cellular inter-nalization. These NPs were further coupled with salbutamol, a β2-adrenoceptor agonist, which improved the targeting specificity ofgreen fluorescent protein (GFP)-siRNA carrier to lung cells having β2-adrenergic receptor. Results indicated enhancement in gene silencingactivity both in vitro and in vivo (using aerosol treatment in the lungof enhanced green fluorescent protein (EGFP) transgenic mice).

A study [150] on NPs of CS used for iv administration of siRNA thatsuggests the presence of PEGylated CS and PEI is important to achievehigh levels of gene silencing in vitro. Since the stability in blood andplas-ma is significant to achieve the desired result, usual procedure of com-bining TPP with CS may not be sufficient. Stable NPs were producedby increasing the amount of PEG and inclusion of anionic polymer,viz., HA. Some representative micro/nanohydrogel delivery systemsfor gene delivery are summarized in Table 2.

Table 2Micro/nanohydrogel systems for gene delivery.

System Size (nm)/zetapotential(mV)

Drug Ref.

Quaternized CS 60% TMCSO-60% b200 pDNA [136]CS-modified PLGA 59, 100/+13 pDNA [138]PLA and CS 65/+5 pDNA [139]MePEG-PLA-CS 94/+13 pDNA [139]PLGA/CS 367/13 Nf-kB ODN [140]Ge core and PEI shell 200/+40 siRNA [141]DEX-HEMA-co-TMAEMA 195–267/17–30 siRNA [142]CS b150/+35 to +41 siRNA [144]CS 40–600/+19 to +31 siRNA [145]HTCSC 143/+32 siRNA [146]TMCS 112, 124, 129, 166/

+5 to 19siRNA [147]

TMCS-MAA 112–243/+5 to +15 siRNA [147]Guanidinylated CS 100/+15 siRNA [149]CS (MW 90, 50–190, 190–310/250,173 kDa)

135–294/+20 to +28 siRNA [150]

CS (MW 190–310/250 kDa)/PEGylated CS

124/+22 siRNA [150]

CS (MW 50–190; 190–310/250;173 kDa)/PEGylated CS/PEI

181, 191, 188/+29,+32/+25

siRNA [150]

PEGylated CS/PEI/HA 181/−16 siRNA [150]

PLGA, poly(lactic-co-glycolic acid); PEI, polyethyleneimine; DEX, HEMA-co-TMAEMA-dextran hydroxyethyl methacrylate-co-[2-(methacryloyloxy)-ethyl]trimethylammoniumchloride; 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, MAAis trimethyl chitosan and methacrylic acid copolymer.

9. Conclusions

Polymeric micro/nanohydrogels have been widely explored inrecent years with renewed interest for increasing the macromoleculartherapeutic outcome and patient compliance. Such systems are particu-larly advantageous to deliver macromolecular therapeutics comparedto other systems due to their lesser cost, 3D network structures andhigh water retaining capacity, in addition to biocompatible and biode-gradable nature. The use of mild preparation conditions is particularlywell suited for preserving the structural integrity, delicate nature andstability of these molecules, while their 3D structures are critical forbiological efficacy. However, relatively rapid release of macromoleculesfrom hydrogel matrix, burst release, low mechanical strength and lessduration of action pose some problems. The fast diffusion of macromol-ecules from such matrices can be circumvented by increasing the drug/polymer ratio, cross-link density and by manipulating physicochemicalcross-linking methods. Even if these associated problems are resolved,micro/nanoparticle-based systems are difficult to use due to their scale-up, high manufacturing cost and limited sales potential. A better under-standing of the patient and medical fraternity needs as well as public–private partnerships would help in developing novel therapies for publichealthcare requirements, which would offer better pharmaco-economicbenefit. In conclusion, micro/nanohydrogel-based delivery of macromol-ecules offers unique opportunities and challenges. However, their furtherpreclinical and clinical studies are necessary for assessing the safety andefficacy of these delivery systems prior to marketing.

Acknowledgements

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

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