Chitosan: Its Applications in Drug Eluting
Devices
Mei-Chin Chena† · Fwu-Long Mib† · Zi-Xian Liaoc · Hsing-Wen Sungc*
a Department of Chemical Engineering, National Cheng Kung University, Tainan, Taiwan b Department of Biotechnology, Vanung University, Chungli, Taoyuan, Taiwan
c Department of Chemical Engineering, National Tsing Hua University, Hsinchu, Taiwan
(†Authors contributed equally; *corresponding to: [email protected])
Abstract Chitosan, a naturally occurring polysaccharide derived from chitin, has been widely applied in drug delivery, tissue regeneration, wound healing, blood coagulation and immunostimulation due to its well-known biocompatibility and biodegradability. Additionally, because of its unique cationic nature and the gel/film/matrix forming capabilities, chitosan has been considered as a promising material for the development of medical devices. The current concept for develop-ing medical devices often comprises the functionality of controlled release of bio-active agents such as drugs, proteins or growth factors in order to fulfill their clin-ical applications.However, in biological fluids, the hydrophilic chitosan matrices may swell and deform dramatically by hydration, thus resulting in a rapid loss of the encapsulated drugs from the delivery devices. Considerable efforts have there-fore been made in chemically modifying chitosan to improve its physical property and functionality. This review article is focused on the versatile modifications of chitosan matrices (ionic or chemical crosslinking) and their most recent research activities in drug eluting devices including vascular stents, artificial skin, bone grafts and nerve guidance conduits.
Keywords Drug-eluting stent, Crosslinking agent, Controlled release, Biode-gradable materials, Tissue Regeneration
Contents 1 Introduction. . . ….. 2 Modification/Crosslinking of Chitosan. . . 2.1 Chitosan Chemistry……… 2.1.1 Acylation……… 2.1.2 Quaternization……… Year of Publication: 2011 ISSN:0065-3195 Publisher:Springer Verlag. DOI Bookmark:10.1007/12_2011_116
2.1.3 Sugar-modification………. 2.1.4 Carboxymethylation……….. 2.1.5 Sulfation……… 2.1.6 Thiolation……… 2.2 Ionic Modification . . . 2.3 Modification by Aldehydes. . . 2.4 Modification by Genipin. . .
2.5 Modification by Other Crosslinking Agents. . .
3 Drug Eluting Stents . . .
3.1 Eluting Coating for Metallic Stent. . . 3.2 Eluting Membrane for Metallic Stent . . . 3.3 Polymeric Stent Crosslinked by an Epoxy Compound. . . . 3.4 Polymeric Stent Crosslinked by Genipin
4 Drug Eluting Medical Devices. . .
4.1 Wound Dressing/Artificial Skin. . . 4.2 Cartilage Graft. . .
4.3 Bone Graft. . .
4.4 Nerve Guidance Conduit. . . .. .
5 Conclusions . . . References . . . . . .
1 Introduction
Chitin [poly(N-acetyl-1,4-β-D-glucopyranosamine)] is an abundant mucopol-ysaccharide that is found in the exoskeleton of crustaceans, the cuticles of insects, and the cell walls of fungi [1,2]. It is structurally similar to cellulose, consisting of 2-acetamido-2-deoxy-β-D-glucose with a β-(1→4) linkage. Chitosan is obtained by the deacetylation of chitin, which is a copolymer of 2-acetamido-2-deoxy-D-galactopyranose (GlcNAc) and β-(1→4)-2-amino-2-deoxy-D-glucopyranose (GlcN) with a GlcN content of more than 50%. Applications of chitin are fewer than those of chitosan because chitin is chemically inert and soluble in neither wa-ter nor acid, whereas chitosan is relatively reactive and can be processed in vari-ous forms [3]. The solubility of chitosan depends on the distribution of free amino and N-acetyl groups. In dilute acids (pH < 6.5), the free amino groups are proto-nated and the molecule thus becomes soluble [4]. Accordingly, chitosan can be easily molded in various forms, such as powder, paste, film, fiber, porous scaffold and others, for various applications.
Owing to its well-known antibacterial nature, minimal foreign body reac-tions, biocompatibility and biodegradability, chitosan is regarded as a good candi-date for tissue engineering [4–6], wound healing [7–9], drug delivery [10–13] and gene delivery applications [14–17]. Functional modifications of chitosan by graft copolymerization or crosslinking are sometimes made to introduce desired proper-ties and enlarge the range of the potential applications. In recent years, chitosan-based medical devices have shown great potential for drug delivery, because of their controlled and sustained release properties. Many studies reported that tissue regeneration can be accelerated by the association of proteins or growth factors with controlled drug delivery devices. This review summarizes the most recent in-vestigations of versatile modifications of chitosan, chitosan-based drug-eluting devices and their formulations to fulfill various clinical requirements.
2 Modification/Crosslinking of Chitosan
2.1 Chitosan Chemistry
The presence of both reactive amino and hydroxyl groups in chitosan
pro-vides opportunities for chemical modification to impart useful physicochemical properties and distinctive biological functions [18–21]. Some of the potential chemical modification reactions have been proposed [22–24]. Specific reactions that involves the –NH2 group at the C-2 position and nonspecific reactions of –
OH groups at the C-3 and C-6 positions (especially esterification and
etherifica-tion) give rise to such chemical reactions as acylation,
tion, carboxymethlation, galactosylation, sulfation, and thiolation, which provide a variety of products with, for example, antibacterial, fungal, viral, anti-coagulation, and transmucosal properties [25,26]. The aforementioned chemical modified chitosan derivatives and their applications are discussed below.
2.1.1 Acylation
N-acyl chitosan derivatives are typically synthesized using anhydrides (linear or cyclic), acyl chlorides and lactones [27–32]. The degrees of acetylation (DA) of chitosan by a controlled N-acetylation reaction can be adjusted by mixing chitosan with acetic anhydride in a molar ratio of 0.1–2.0. Modifying the DA of chitosan significantly affects the material properties that are critical to tissue engineering applications, by increasing degradation times and enhancing cell adhesion [33]. Moreover, chitosan has been selectively N-acylated with linear aliphatic anhy-drides of different chain length, including acetic, propionic, butyric and hexanoic anhydrides, to prepare acetyl chitosan, propionyl chitosan, butyryl and N-hexanoyl chitosan, respectively [29]. The hydrophobic interaction of N-acylated chitosan that has been substituted with longer acyl chains is stronger. In contrast, the drug release rates of N-acylchitosan microspheres with longer acyl chains are less susceptible to enzymatic degradation than are their N-acetylchitosan counter-parts [34].
Since the amino group is more active than the hydroxyl groups, a new meth-od for the selective O-acylation of chitosan in methanesulfonic acid (MSA) has been developed [35]. The N,O-acylation of chitosan has also been conducted by dissolving the O-acyl chitosans that are synthesized via an MSA protection meth-od in N,N-dimethylacetamide and then performing N-acylations on the O-acyl chitosans [36,37].
Chitosan with aliphatic side chains such as stearoyl, palmitoyl and octanoyl groups has been hydrophobically modified by reacting carboxylic anhydride with chitosan through N-fatty acylation reactions [28,38]. N-palmitoyl chitosan (NPCS) has also been synthesized by reacting palmitic acid N-hydroxysuccinimide ester with chitosan. This NPCS was developed as a pH-triggered, in situ gelling system, which underwent a rapid nanostructural transformation within a narrow pH range (pH 6.5–7.0) [39,40]. In dilute aqueous media, NPCS can self-assemble into na-noparticles, owing to the hydrophobic interaction between their conjugated pal-mitoyl groups [41]. The extent of cellular uptake of NPCS nanoparticles increased with the degree of substitution (DS) of palmitoyl groups in NPCS. The internaliza-tion of NPCS nanoparticles was clearly related to the lipid raft-mediated routes. Figure 1 shows the reaction scheme for aforementioned acylation of chitosan.
2.1.2 Quaternization
N-trimethyl chitosan chloride (TMC) with introduced quaternary amino groups on the chitosan chain has been synthesized using iodomethane in an alka-line solution of N-methyl pyrrolidinone (NMP). Quaternization is based on the nucleophilic alkylation of the primary amino group at the C-2 position of chitosan [42]. TMC is the most important quaternized chitosan derivative, and has been proven effectively increase the paracellular permeation of hydrophilic macromo-lecular drugs across the mucosal epithelia by opening the tight junctions [43]. The degree of substitution of TMC crucially influences its ability to open up the tight junctions of intestinal mucosa. A degree of quaternization of approximately 40– 50% is reportedly optimal for transepithelial delivery of both low molecular weight compounds [44,45] and proteins [46–49].
Chitosan has been modified using glycidyl trimethyl ammonium chloride (GTMAC) to produce another quaternized chitosan derivative, N-(2-hydroxyl) propyl-3-trimethylammonium chitosan chloride (HTCC). When a primary amino group at the C-2 position of chitosan reacted with GTMAC, the chain of the qua-ternary ammonium group thus obtained was longer than that of the TMC. The pre-liminary protein release test from HTCC or modified HTCC nanoparticles in vitro reveals that they are functional carriers for protein delivery [50]. Figure 2 shows the reaction scheme for aforementioned quaternization of chitosan.
2.1.3 Sugar-modification
Carbohydrates such as lactose can be grafted onto the chitosan backbone at the C-2 position by reductive alkylation, or can be introduced at the C-6 position without opening of the ring [51]. These sugar-modified chitosan derivatives can be
used for drug targeting since the specific recognition of cells, viruses, and bacteria by sugars has been proven. D- and L-fucos-bound chitosan derivatives have been synthesized to interact specifically with lectin and cells [52].
Asialoglycoprotein receptors are known to be present on hepatocytes at a high density of 500,000 receptors per cell, and they are retained on several human hepatoma cell lines [53,54]. A lactosaminated N-succinyl-chitosan derivative has been synthesized as a liver-specific drug carrier in mice to target asialoglycopro-tein receptors [55]. Galactosylated chitosan (GC) can be synthesized by coupling chitosan with lactobionic acid via an active ester intermediate using 1-ethyl-3-(3-dimethylaminopropyl)-carbodiimide hydrochloride (EDC). GC has potential as a synthetic extracellular matrix for hepatocyte attachment [56]. Chitosan bearing single or antennary galactose residues has also been investigated to determine its specific targeting to HepG2 cells after it is formed into nanoparticles or complexes with DNA [57–59]. GC/DNA complex has a much higher transfection efficiency than the chitosan/DNA complex on HepG2 [59,60], indicating that a galactose lig-and that is attached to GC not only helps the GC/DNA complex bind to asialogly-coprotein receptors, but also promotes the internalization of GC/DNA complex within membrane-bound vesicles or endosomes [61]. Figure 3 shows the reaction scheme for aforementioned sugar-modification of chitosan.
2.1.4 Carboxymethylation
Both, N-carboxymethyl and O-carboxymethyl chitosan derivatives have been prepared under different reaction conditions using monochloroacetic acid in the presence of NaOH to achieve the N-versus-O selectivity [62,63]. The selective route for synthesizing N-carboxymethylation utilizes glyoxylic acid in a reductive amination sequence [64]. N-Carboxymethyl and 6-O-carboxymethyl chitosan (N-CC and 6-O-(N-CC) are both polyampholytic chitosan derivatives, whose molecules contain –COOH groups and –NH2 groups. These carboxymethyl chitosans are not
only soluble in water, but also have numerous outstanding chemical, physical and biological properties, including high viscosity, film formation, gel formation, non-toxicity, biodegradability, biocompatibility, and antibacterial and antifungal activi-ties, all of which make them attractive for used in biomedical, pharmaceutical and cosmetic applications [65,66].
N,O-Carboxymethyl chitosan (N,O-CC) is a chitosan derivative that has car-boxymethyl substituents at some of both the amine and the 6-hydroxyl sites of its glucosamine units. It can be easily prepared using chitosan, sodium hydroxide, isopropanol with chloroacetic acid [67]. N,O-CC as a wound dressing can stimu-late the extracellular lysozyme activity of fibroblasts and considerably promote the proliferation of skin fibroblasts. N,O-CC is also used in the development of
vari-ous functional hydrogels, including superporvari-ous or pH-sensitive hydrogels, which are used for protein drug delivery [68–71].
Another carboxymethyl chitosan derivative called N-succinyl chitosan is ob-tained by introducing succinyl groups at the N-terminal of the glucosamine units of chitosan [72]. Various macromolecule-antitumor drug conjugates could be pre-pared from the synthesis process of N-succinyl chitosan and drug conjugation [73]. For example, N-succinyl chitosan can directly react with an activated ester of glutaric mitomycin [74,75] or with mitomycin C using carbodiimide as a coupling reagent [76]. Figure 4 shows the reaction scheme for aforementioned carbox-ymethylation of chitosan.
2.1.5 Sulfation
N-, O- or N,O-sulfate derivatives of chitosan have been synthesized using numerous methods in the preparation of bioactive sulfated polysaccharides with anticoagulant activity, immunomodulating effects, and anti-virus activity [77]. Various reagents have been used in the sulfation of chitosan, including concentrat-ed sulfuric acid, oleum, sulfurtrioxide, sulfurtrioxide/pyridine, sulfurtriox-ide/trimethylamine, sulfur trioxide/sulfur dioxide, chlorosulfonic acid-sulfuric acid and chlorosulfonic acid, which is the most commonly used [78–80]. These synthe-ses have been performed under both homogeneous and heterogeneous conditions in such media as DMF, DMF-dichloroacetic acid, tetrahydrofuran, and formic acid [81]. Sulfated derivatives of chitosan exhibit different blood anticoagulant activi-ties depending on the degree and sites of substitution [82–84]. Increasing the sul-fur content in chitosan increased its anticoagulant activity. 6-O-carboxymethyl chitosan N-sulfate exhibited 23% of the activity of heparin [82], and the O-sulfated form of 6-O-carboxymethyl chitosan exhibited 45% of the activity of heparin in vitro [83]. N-carboxymethyl chitosan 3,6-disulfate exhibited anticoagu-lant activity similar to that of heparin [84].
N-sulfate, and 6-O-sulfated and 2-N,6-O-sulfated chitosans were synthesized and their effects on the bioactivity of bone morphogenetic protein-2 (BMP-2) were compared. The 6-O-sulfated substitution primarily increased the bioactivity of BMP-2, while 2-N-sulfate acted as a subsidiary group, which provided less ac-tivity. A low dose of 2-N, 6-O-sulfated chitosan markedly enhanced the activity of alkaline phosphatase (ALP) and the mineralization that was induced by BMP-2 [85]. N,O-SOCC that is substituted with 6-O-carboxymethyl groups and 2-N, 3-O, and 6-O sulfate groups may have well-balanced steric and electrostatic structures that can stably interact with bFGF to protect the growth factor against proteolytic inactivation. The sulfate groups appeared to bind predominantly to basic amino acid residues, while the carboxymethyl groups complemented sulfate groups,
sta-bilizing the binding structure [86]. Figure 5 shows the reaction scheme for afore-mentioned sulfation of chitosan.
2.1.6 Thiolation
Thiolated chitosans are synthesized by covalently coupling thiol groups to chitosan backbones. They exhibit in situ gel-forming properties owing to the for-mation of inter- and/or intramolecular disulfide bonds at pH values above 5 [87]. Sulfhydryl-bearing agents such as cysteine or thioglycolic acid can be covalently bound to chitosan. The thiolation reaction of chitosan involves amide bond for-mation between the carboxylic acid group of the thiol reagent and the primary amino group of chitosan, mediated by a water-soluble carbodiimide [87,88]. Thio-lated chitosans can also be prepared by modifying chitosan with 2-iminothiolane or isopropyl-S-acetylthioacetimidate, which is in situ cross-linkable. As well as a sulfhydryl group, a cationic moiety is also introduced in the form of an amidine substructure, resulting in a chitosan-4-thio-butylamidine conjugate (chitosan-TBA) or chitosan-thioethylamidine conjugate (chitosan-TEA) [89,90]. A new sys-tem of thiolated chitosan that is based on chitosan-thioglycolic acid (chitosan-TGA) in the presence of oxidizing agents such as hydrogen peroxide (H2O2),
so-dium periodate (NaIO4), ammonium persulfate ((NH4)2S2O8) and sodium
hypo-chlorite (NaOCl) has been examined to increase the dynamic viscosity and reduce the reverse volume phase time [91].
Chitosan has mucoadhesive properties [92], however the mucoadhesive properties of chitosan can be significantly improved by the immobilization of thiol groups on the polymer. The mucoadhesion is enhanced by the formation of disul-fide bonds with cysteine-rich subdomains of mucus glycoproteins, and these bonds are stronger than non-covalent bonds [92,93]. Chitosan-TBA conjugates provide 140 times stronger mucoadhesion than unmodified polymer [87,89]. These favorable results can be explained by the fact that chitosan-TBA combines the formation of disulfide bonds with improved ionic interactions between the ad-ditional cationic amidine groups of the modified chitosan and the anionic moieties that are provided by sialic acid and sulfonic acid substructures in the mucus layer. Figure 6 shows the reaction scheme for aforementioned thiolation of chitosan.
Thiolation of chitosan strongly improved the permeation-enhancing capabili-ties of chitosan, facilitating the paracellular transport of hydrophilic compounds. The immobilization of thiol groups can strongly improve the permeation-enhancing effect of chitosan. The uptake of rhodamine-123 in the presence of chi-tosan-TBA was three times that versus unmodified chitosans [94]. Permeation studies were also carried out using the fluorescent dye rhodamine-123 (Rho-123) as a P-gp (P-glycoprotein) substrate, indicating that chitosan-TBA potentially in-hibits the ATPase activity of P-gp in the intestine [95]. Furthermore,
chitosan-TBA seems to inhibit protein tyrosine phosphatase, opening tight junctions, to in-crease the permeation of hydrophilic compounds [96].
2.2 Ionic Modification
Chitosan is acid-soluble because its crystal structure can be broken down un-der acidic conditions. To preserve the stability of chitosan gel unun-der GI (gastroin-testinal) tract delivery or enzymatic degradation, the amine groups on its polymer-ic chains must be fixed by chempolymer-ical reagents. The ionpolymer-ic interactions between the positively charged amino groups on chitosan and either small anionic molecules, such as sulfates, citrates, and phosphates, or some metal ions, have been success-fully used in the preparation of chitosan beads or hydrogels [97,98].
Small anionic molecules bound with the protonated amino groups on chitosan via electrostatic attractions, but metal ions form coordinate-covalent bonds with chitosan rather than electrostatic interactions [99]. Tripolyphosphate (TPP) which can interact with chitosan by electrostatic attractions has been recognized as an acceptable food additive by the U.S. Food and Drug Administration [98]. In the ionotropic gelation method, chitosan dissolved in aqueous acidic solution is added dropwise with constant stirring to TPP solution to form spherical particles. Since the pKa of chitosan is in the range of 6.3–6.7, chitosan gel beads that are cured at pH values of less than 6 were really ionically crosslinked. However, chitosan gel beads that are cured at pH values of greater than 7 are coacervation-phase inver-sion-controlled, and depend slightly on crosslinking [100]. The ionic-crosslinking density of chitosan beads can be improved by changing the pH value of the curing agent, TPP, from basic to acidic [101].
The process by which chitosan is ionically cross-linked is simple and fre-quently performed under mild conditions without the use of organic solvents. Hence, several researchers have explored its potential pharmaceutical use [102,103]. Bodmeier et al. have developed a new approach for preparing chitosan nanoparticles that was based on the ionotropic gelation method [104], forming positively charged, stable CS nanoparticles with sizes in the range of 250–400 nm. Subsequently, ionically cross-linked chitosan nanoparticles have been extensively used to deliver various small molecular drugs and bioactive macromolecules. An-ti-cancer or protein drugs such as doxorubicin or insulin can be effectively en-trapped into the chitosan nanoparticles during ionotropic gelation process [105,106].
2.3 Modification by Aldehydes
Chitosan and its derivatives have been extensively developed for biomedical applications because of their biocompatibility and enzymatic degradability. Glu-taraldehyde, a bifunctional aldehyde, is traditionally utilized to cross-link chitosan. The cross-linking process involves formation of imine bonds between the amino groups on chitosan chains and bifunctional glutaraldehyde cross-linker [107]. However, glutaraldehyde is generally considered to be relatively toxic and the fate of this cross-linking reagent in human body has not been established. The glutar-aldehyde-crosslinked chitosan gels can be restructured into tablets, membranes, beads, and nano- or microspheres. Recently studies have demonstrated that cross-linking is one of the main factors that affect the physical and swelling properties of chitosan-based gels [108].
The degree of cross-linking of chitosan depends on the degree of deacetyla-tion but is found to be independent of the molecular weight of chitosan [108]. The concentration of glutaraldehyde and the reaction temperature are also used to con-trol the degree of cross-linking in chitosan-based microspheres [108–110]. Cross-linked chitosan microspheres has been investigated for potential use in a drug de-livery system for delivering anticancer drugs, such as 5-fluorouracil (5-FU), cis-platin (CCDP), oxantrazole (OXZ) and others [111–114]. Additionally, chitosan can be crosslinked by glutaraldehyde to form a network to promote the sustained release of bioactive reagents, including centchroman and progesterone [115,116]. Cross-linked chitosan microspheres reportedly released hydroquinone faster than do their uncross-linked counterparts [117]. The rate of release depends on the de-gree of cross-linking of chitosan microspheres. Gupta and Jabrail also reported an extended, constant release of centchroman, a selective estrogen receptor modula-tor, from cross-linked microspheres for 60–70 h, whereas this process continued for only 10 h from uncross-linked microshperes [116]. These studies demonstrated that drug release rates may be manipulated by controlling the degree of cross-linking of the chitosan hydrogels or microspheres.
2.4 Modification by Genipin
Genipin, a naturally occurring heterocyclic compound, can be obtained from its parent compound, geniposide, which is isolated from the fruits of Gardenia jasminoides ELLIS. Genipin and its related iridoid glucosides have been extensive-ly used as antiphlogistics and cholagogues in herbal medicine [118]. Comparing the cytotoxicity of genipin to that of glutaraldehyde in vitro using 3T3 fibroblasts via the MTT assay reveals that genipin is about 5000–10,000 times less cytotoxic than glutaraldehyde [119]. In an in vitro 3T3 cell culture or an in vivo rat model, a
chitosan membrane or injectable chitosan microspheres that are cross-linked with genipin histologically exhibited less cytotoxicity, better biocompatibility and slower degradation rate than their glutaraldehyde cross-linked counterparts [120,121]. Relevant studies strongly suggest that the compatibility of genipin is superior to glutaraldehyde.
The above results indicate that the crosslinking reactions were pH-dependent. Under neutral and acidic conditions, genipin reacted with chitosan via a nucleo-philic attack by the primary amino group, on the olefinic carbon atom at the C-3 position of deoxyloganin aglycone, followed by opening of the dihydropyran ring and attacked by the secondary amine group on the intermediate aldehyde group, forming heterocyclic amines [122]. The heterocyclic amines were further associat-ed to form cross-linkassociat-ed networks with short chains that comprisassociat-ed dimmer, trimer, and tetramer bridges. Under strongly basic conditions, genipin underwent a ring-opening polymerization prior to linking with chitosan [123]. The cross-linking bridges consisted of polymerized genipin macromers or oligomers. This ring-opening polymerization of genipin was initiated by extraction of a proton from the hydroxyl groups at the C-1 position of deoxyloganin aglycone, followed by opening of the dihydropyran ring to enable an aldol condensation [123].
Genipin has been adopted to cross-link chitosan in the preparation of hydro-gels, wound dressings, scaffolds, films, beads, and nano- or microspheres for wound healing, tissue engineering, and drug delivery [70,120,121,124–126]. Nov-el chitosan gNov-el beads have been synthesized by a coupled ionic and chemical co-crosslinking method using TPP and genipin as cross-linkers. The pH-dependent ionic/chemical co-crosslinking mechanism has an obvious effect on the swelling property and enzymatic degradation behavior of the prepared chitosan gel beads [127]. A pH-sensitive hydrogel of N,O-carboxymethyl chitosan (NOCC) and algi-nate, cross-linked by genipin, for protein drug delivery has been reported [70]. The amount of BSA that is released at pH 1.2 is relatively low (20%), while that released at pH 7.4 is significantly greater (80%). These results clearly suggest that the genipin-cross-linked NOCC/alginate hydrogel may be a suitable polymeric carrier for site-specific delivery of protein drug in the intestine.
2.5 Modification by Other Cross-linking Agents
In addition to glutaraldehde and genipin, numerous bifunctional reagents have been used to cross-link chitosan covalently, such as epichlorohydrin, diiso-cyanate or epoxy compound (4-butanediol diglycidyl ether or ethylene glycol di-glycidyl ether (EGDE) [128–132]. Figure 7 shows the reaction schemes for cross-linking chitosan with these bifunctional reagents, as well as TPP, glutaraldehyde and genipin. Among those bifunctional reagents, EGDE may be the most suitable
linker for reaction with chitosan to prepare flexible films. Observably, cross-linking with glutaraldehyde increases the tensile strength of 6-O-CC/water soluble polyurethane (WPU) composite membranes, but reduces their elongation. In con-trast, the elongation of 6-O-CC/WPU membrane increased upon reaction with EGDE [133]. Recently, a novel biodegradable stent, made of chi-tosan/poly(ethylene oxide) blend films that were cross-linked with EGDE and ex-hibit the shape-memory property, was developed for the sustained release of siro-limus [134]. Other studies, such as those involve the preparation of highly porous microspheres by cross-linking chitosan with EGDE to create an open porous sur-face structure have been performed with the goal of the sustained release of the antigen of the Newcastle disease vaccine [135]. Furthermore, quaternary ammoni-um, and aliphatic and aromatic acyl groups, can be introduced into the porous chi-tosan beads to adsorb an anti-inflammatory drug, indomethacin, via both electro-static and hydrophobic interactions [136].
Chitosan solutions that are physically mixed with β-glycerophosphate (β-GP), can be injected into the body in liquid form, forming a gel in situ at the body temperature. The rate of gelation depends on the degree of chitosan deacetylation, the concentration of β-GP, and the temperature and pH of the final solution [137]. The in situ gelation mechanism involves neutralization of the ammonium groups in chitosan, allowing strengthened hydrophobic and hydrogen bonding between the chitosan chains at elevated temperatures. A new, mild approach to in situ hy-drogel formation via a peroxidase-catalyzed cross-linking reaction that takes only seconds was recently developed. This approach has shown great potential for mak-ing injectable hydrogels for biomedical applications, such as tissue engineermak-ing and drug or protein delivery [138,139].
3 Drug Eluting Stents
Atherosclerosis is the primary cause of coronary heart disease, which is char-acterized by a narrowing (stenosis) of the arteries that supply blood to tissues of the heart [140–142]. Metallic coronary stents were originally developed to prevent abrupt artery closure and to reduce the likelihood of restenosis that is associated with percutaneous transluminal coronary angioplasty. They are, however, limited by the frequent occurrence of restenosis, which is caused by smooth muscle pro-liferation, and associated neointimal hyperplasia and target lesion revasculariza-tion [143].
Drug-eluting stents (DES) are a revolutionary technology because of their unique ability to provide both mechanical and biological solutions simultaneously to the target artery [144]. Generally, each DES comprises three components - the backbone stent, the active pharmacologic compound, and a drug carrier vehicle, which controls drug elution [145]. The biological effects of the pharmacological agents and the interaction of DES components with the arterial wall have been demonstrated to attenuate greatly in-stent restenosis, relative to that associated with a bare-metal stent [143].
Using polymers as drug delivery vehicles typically enables the best con-trolled and sustained drug release [146]. Current commercially available DES (such as Cypher, Taxus, Xience V, PromusTM and Endeavor) comprise a me-tallic stent with a coating of a non-erodable synthetic polymer that contains anti-proliferative agents [143]. The drugs can inhibit the proliferation and migration of vascular smooth muscle cells (VSMC), which are important factors in the devel-opment of neointima formation, preventing restenosis [147]. After the drug elutes from the polymer coatings, the residual synthetic polymer coatings remain in place. Eventually, the permanent presence of the nonresorbable polymer may lead to complications such as an exaggerated inflammatory response and neointimal hyperplasia at the implant site [148,149]. A principal target of current research is to develop the next generation of DES with a biodegradable and biocompatible coating to minimize these unfavorable effects.
Chitosan is a biodegradable, non-toxic and tissue-compatible polymer; it is regarded as an excellent candidate for drug delivery applications. Many studies have been conducted to explore the feasibility of using chitosan as a drug reservoir for DES, as described in the following sections.
3.1 Eluting Coating for Metallic Stent
Although DES have dramatically reduced in-stent restenosis, histopathologi-cal assessment in autopsy cases has demonstrated delayed healing and incomplete endothelialization. Additionally, evidence exists of hypersensitivity reactions in patients with stent thrombosis who had received DES [150]. A hypersensitivity re-action is caused by the nonerodable polymer. The mechanisms by which DES cause a prolongation of arterial healing and endothelialization, which may be sus-ceptible to late thrombosis, are poorly understood [150]. Recent investigations have found that antiproliferative drugs (sirolimus and paclitaxel) reduce neointi-mal formation by impeding VSMC proliferation and migration, and retard the normal healing of the injured arterial wall [150].
Bioactive surfaces and compounds that promote vascular healing are emerg-ing to improve stent performance and patient outcomes [143]. A biomacromolecu-lar layer-by-layer (LbL) coating of chitosan/heparin (CS/HP) onto a coronary stent has been designed to accelerate re-endothelialization and healing after coronary stent deployment [151]. Chitosan has promise as a bioactive material for implant devices because of its capacity to enhance wound healing and cell attachment [152]. Heparin is the most often used anticoagulant reagent in clinical use. It in-controvertibly suppresses sub-acute in-stent thrombus [153]. The results of in vitro culturing of porcine iliac artery endothelial cells, as well as a haemocompatibility evaluation, support the claim that the combination of chitosan and heparin can make the stent surface compatible with endothelial cells and haemocompatible.
A porcine coronary injury model and an arteriovenous shunt model were used to evaluate the application of such a surface-modified stainless steel stent in vivo. In the second week, neointimal layers were present on the surfaces of some bare-metal stent (BMS) samples, whereas the intimal tissues seemed to be more integrated on the surfaces of all of the CS/HP-coated stents. After four weeks of implantation, the healing on all BMS and CS/HP-coated stents was completed. No obvious inflammation response to any of the CS/HP-coated stents was observed, suggesting the good biocompatibility of the CS/HP coating and the safety of its in vivo application (Fig. 8) [151].
Endothelialization of the inner wall of the stented arteries was further exam-ined by SEM (Fig. 9) [151]. In the second week of development, cell attachment occurred on the BMS surface with a relatively low density. These cells appeared to be infantile and undeveloped, while on the CS/HP-coated stents, the luminal surface of the vessel wall and the stent struts had been covered with confluent shuttle-like endothelial cells. The endothelialization of the BMS surfaces was sta-tistically estimated to be approximately 30% as observed by SEM, whereas that of
the surfaces of the CS/HP-coated stents was almost 100%. After four weeks, the attached and grown shuttle-shaped endothelial cells completely covered the BMS and the CS/HP-coated stents. Therefore, the evidence conclusively established that the natural self-assembly coating accelerated the re-endothelialization process in the coronary stent system. This finding may provide a solution to the delayed heal-ing problem of DES, and thus the natural self-assembly coatheal-ing may have great potential in future use as drug reservoirs.
A separate attempt has also been made to use the LbL coating method to form self-assembling chitosan/hyaluronan (CS/HA) coatings for endovascular stents [154]. HA, a naturally occurring polyanion, plays an important structural and mechanical role in various tissues. The inhibitive effects of HA with respect to hyperplasia, observed after either systemic or local delivery, suggest that the anti-proliferative effects of HA may be related to its antiinflammatory properties [155]. Related results have demonstrated that CS/HA multilayer-coated NiTi disks has better antifouling properties than unmodified NiTi disks, as demonstrated by a de-crease in the adhesion of platelets in an in vitro assay (38% reduction). This result may be attributable to the hydration layer that surrounds HA molecules on the sur-face of the disks [156]. The hydrogel-like sursur-face that is formed by the multilayers exhibited good haemocompatibility.
To determine whether the CS/HA multilayers may be exploited as in situ drug delivery systems, the nitric-oxide-donor sodium nitroprusside (SNP) was in-corporated in the multilayers. SNP, a nitrous oxide donor that spontaneously de-composes in biological environments [157], is extensively used clinically to re-duce blood pressure and has emerged as a promising modality in the treatment of restenosis [158]. SNP-loaded multilayers have been shown to reduce platelet ad-hesion below that associated with multilayers devoid of the drug. The enhanced thromboresistance of the self-assembled multilayer, together with the anti-inflammatory and wound-healing properties of HA and CS, are expected to reduce neointimal hyperplasia that is associated with stent implantation.
The LbL self-assembly of chitosan and polyanions into multilayers has emerged as an efficient and versatile approach for forming biologically active sur-faces [154]. CS/HP LbL modification is simpler and less expensive than other methods of facilitating vascular healing (CD34 antibody-coated or RGD peptide-coated stent) [159,160]. An LbL coating increases stability during sterilization and storage. Additionally, bioactive molecules, such as DNA or proteins, may be in-corporated into the multilayer during the polyelectrolyte deposition process, form-ing drug-releasform-ing interfaces on various substrates [151].
3.2 Eluting Membrane for Metallic Stent
Membrane-covered stents have also been studied with a view to their use in the treatment of coronary diseases [161–163]. A stent that is covered with chi-tosan/polyethylene oxide (CS/PEO) membrane has been shown to have suitable properties for endoscopic implantation [164]. The results of mechanical tests indi-cate that a blend of chitosan with PEO with a high molecular weight can greatly improve the elasticity and strength of chitosan. CS/PEO membrane-coated stents have been demonstrated to sustain mechanical deformation during endoscopic ex-pansion and to resist physiological blood pressure. Additionally, the blended membrane has good hemocompatibility because it has a lower adsorption than a non-blended CS membrane. To study further the capacity of the CS/PEO mem-brane to act as a drug reservoir, SNP was used as a model drug and loaded in the membrane via an ionic interaction between the anionic drug and CS. The SNP-loaded membrane further reduced platelet adhesion. These results suggest that the CS/PEO membrane could also be easily loaded with other bioactive drugs for the treatment of various pathologies.
3.3 Polymeric Stent Crosslinked by an Epoxy Compound
Even though polymer-coated metallic DESs have revolutionized the treat-ment of obstructive coronary diseases, they do not provide an optimal solution [165]. FDA reports and autopsy findings suggest that metallic DESs may be a cause of systemic and intrastent hypersensitivity reactions that, in some cases, have been associated with late thrombosis and death [166]. Other reports have fur-ther suggested that the most likely cause of the hypersensitivity reaction is the non-erodable polymer coating of the DES [148,149,167]. Additionally, in-stent restenosis may be associated with allergic reactions to the nickel (stent compo-nent) and molybdenum (stent impurity) in metallic stents [149,168].
Stents are superior to angioplasty as they provide scaffolding of the vessel and prevent elastic recoil and detrimental remodeling following revascularization [143]. However, whether the presence of a permanent stent is favorable, or wheth-er it would be more advantageous that the stent was degraded and absorbed by the body once its task was done, is unclear [143]. Recently, biodegradable polymeric stents have attracted much interest as an alternative to metallic stents [169−171].
A self-expandable polymeric stent, made of chitosan-based films that are crosslinked with an epoxy compound (ethylene glycol diglycidyl ether), and ex-hibiting shape memory has been developed [169
–
171]. The flexibility of chitosan films can be significantly improved by blending with glycerol and PEO. Since co-valent crosslinks form in the stent matrix, a chitosan-based stent has shape memory (Fig. 10) [170]. It can expand (~150 s) from its crimped (temporary) toits fully expanded (permanent) states upon hydration (Fig. 11) [170], markedly faster than can polymeric stents that are made of poly-L-lactic acid (PLLA) (~20 min at 37°C for the Igaki-Tamai stent) [172] or PLLA/poly-D-L-lactide-glycolide (~8 min at 37°C for the stent that was developed by Venkatraman et al.) [173]. Rapid self-expandability of the stent is advantageous, as it helps prevent migration of the stent during its in vivo deployment. In a preliminary animal study, an im-planted chitosan-based stent was found to be intact and no thrombus was formed in the stent-implanted vessel.
Sustained delivery of antiproliferative drugs to prevent in-stent restenosis is critical for an ideal biodegradable polymeric stent. However, the drugs (sirolimus and paclitaxel) that are most commonly used in currently available DES systems are both lipophilic. The hydrophilic nature of chitosan matrices makes them una-ble to entrap poorly soluuna-ble therapeutic agents and greatly limits their range of ap-plications as drug-delivery systems. A consistent limitation of these hydrophilic systems appears to be the rapid loss of therapeutic agents because of that lack suf-ficiently strong interactions between the lipophilic drug and the hydrophilic poly-mer. Additionally, large aggregates may be formed by the interaction among lipo-philic pharmaceuticals in the drug loading process, resulting in a high local concentration and, thus, toxicity at the sites of the aggregates [174].
A nanoscale drug entrapment strategy has been proposed for the controlled and sustained release of sirolimus in chitosan-based stents [171]. Sirolimus was first entrapped in the core of self-assembled Pluronic block copolymer L121 (PEO5-PPO68-PEO5) micelles; the hydrophilic outer shell of L121 micelles then
maintained their uniform dispersion and stability in the hydrophilic chitosan ma-trix (Fig. 12) [171]. The SEM results indicated that large aggregates of sirolimus were present at random locations in the control group (sirolimus/chitosan-blended stents, blue arrow, Fig. 13a) [171], whereas spherical micelles were uniformly dispersed in the experimental stents with sirolimus-loaded micelles (Fig. 13b) [171].
In an in vitro drug release study, an initial burst release of more than 40% of sirolimus from the control stents was observed on the first day of the experiment, whereas no apparent burst release from the experimental stents occurred (Fig. 14) [171]. After ten days, the cumulative percentages of drug that were released from the low-dose and high-dose control stents were approximately 70% and 80%, re-spectively. These were followed by continuous release for about two months. In contrast, the cumulative percentage of the drug that was released in ten days from the experimental group was greatly lower, being 35% (low-dose stent) or 28% (high-dose stent); this was followed by a prolonged and sustained drug release for up to three (low-dose stent) or six months (high-dose stent). These results indicate that a nanoscale drug entrapment strategy can prevent the drug from aggregating and beneficially reduce its initial burst release, greatly extending the duration of drug release.
Abnormal VSMC proliferation isinvolved in restenosis following percutane-ous transluminal angioplasty [175,176]. Previpercutane-ous studies have demonstrated that macrolide antibiotic sirolimus can inhibit VSMC proliferation by blocking cell cy-cle progression at the G1/S transition [177]. To evaluate the activity of the
re-leased sirolimus, rat thoracic aorta smooth muscle cells (RASMC) were coincu-bated with test stents for three days. The cell viability of RASMC was significantly reduced when they were coincubated with either the control or exper-imental stent, because of the released sirolimus; the anti-proliferative activities of sirolimus that was released from these two stents on RASMC were similar to each other (Fig.15) [171]. Cell cycle analysis revealed that the sirolimus that was re-leased from the stent retained its original activity in inhibiting RASMC prolifera-tion during the G1 phase, which finding was consistent with the proven effects of
the agent on cell cycle signaling and proliferation.
To investigate the efficacy of the prepared DES in inhibiting restenosis, test stents were individually deployed into the infrarenal abdominal aorta of rabbits. Six weeks after stenting, stent-released sirolimus had greatly altered the intimal response to stent implantation. In the group treated with an empty stent, a marked neointima was observed (inset in Fig. 16a) [171] and many RAM11-positive mac-rophages (green) were present around the stent struts (Fig. 16d) [171]. In contrast, the neointimal area was markedly reduced (insets in Figs. 16b and 16c) [171], and less macrophage infiltration was observed (Fig. 16e and Fig. 16f) [171] following implantation of the control or experimental stent, revealing the anti-proliferative and anti-immune effects of the released sirolimus.
Reendothelialization occurs after vascular injury and similarly following stent placement [147]. Endothelial cells are conventionally believed to proliferate and migrate from intact neighboring artery segments, eventually leading to the reendothelialization of the injured segment [147]. The degree of CD31-positive cellular coverage (indication of reendothelialization, blue arrows in Fig. 16a−16c) [171] was notably higher in arteries that were implanted with an empty stent than in those that received sirolimus-loaded stents (control or experimental stents).
Recent studies have reported that sirolimus reduces neointimal formation by inhibiting the proliferation and migration of VSMC and delays the normal healing processes of the injured arterial wall [150]. In the control group, poor endothelial cell junction formation (Fig. 16b) [171] and thrombus-like substance deposition around the stent struts (indicated by the black arrow in the inset) were observed at six weeks after implantation, possibly owing to the adverse side effects of the ini-tial burst release of sirolimus. In contrast, the experimental stent with nanoscale si-rolimus-entrapment within L121 micelles provides sustained release without any apparent burst (Fig. 14) [171], preventing undesirable side effects such as delayed endothelial healing that would be caused by the overdose of sirolimus (Fig. 16c) [171]. The patient must receive a safe and effective dosage of the drug.
3.4 Polymeric Stent Crosslinked by Genipin
Most polymeric stents are flexible along their longitudinal axes, facilitating the delivery of stents through a tortuous vessel. However, one of the limitations of the use of polymers as stent matrices is their inherent mechanical weakness. Cova-lent crosslinking has been extensively utilized for improving structural stability and the mechanical properties of many engineering materials [178,179]. The fixa-tion of materials using various crosslinking agents reportedly may form distinct crosslinking structures, which affect their mechanical properties and crosslinking characteristics [180]. A naturally-occurring crosslinking agent, genipin, which has a cyclic molecular structure has been utilized to crosslink chitosan-based stents (genipin stents) as polymeric stents with the enhanced mechanical strength [181].
The results of mechanical tests indicate that the genipin stent has a signifi-cantly higher ultimate compression load (1123 ± 77 mN) and collapse pressure (2.5 ± 0.1 bar) than the epoxy stent (856 ± 148 mN, 1.9 ± 0.1 bar). These results are attributable to the various crosslinking structures that are formed in stent ma-trices (Fig. 17) [181].The two epoxide functional groups in epoxy compound that was used in the study crosslinked the amine groups of chitosan in the stent matrix [182]. In this way, a linearly crosslinked structure between the adjacent chitosan molecules may be formed intermolecularly.
In contrast, genipin reacts spontaneously with the amine groups of chitosan to form a nitrogen-iridoid, which undergoes dehydration to an aromatic monomer. Dimerization occurs in the second stage, perhaps via a radical reaction [183]. Hence, genipin can form a heterocyclic intermolecular crosslinking structure in the stent matrix. The heterocyclic crosslinking structure that is formed in the gen-ipin stent matrix is denser than the linear crosslinking structure in the epoxy stent matrix, because of it is bulky and cyclic(Fig. 17) [181].Therefore, the mechanical strength of the genipin stent greatly exceeds that of its epoxy counterpart. These results suggest that the cyclic crosslinked structures formed within the genipin stent matrix improved its mechanical properties.
The in vivo vascular responses of the genipin stent were investigated in rabbit infrarenal abdominal aortas. At three months postoperatively, the retrieved arteries remained patent, and no thrombosis was observed. An almost intact layer of endo-thelial cells was observed on the stent-implanted vessel wall. Endoendo-thelialization is one of the most promising mechanisms for reducing thrombogenicity of any car-diovascular prostheses [184]. To evaluate its potential use in developing a drug de-livery vehicle, the genipin stent was loaded with sirolimus. The genipin stent that was coated with hydrophobic heparin (DurafloTM, Edwards Life Science) had a linear sustained-release profile (Fig. 18) [181] and the released sirolimus main-tained its original activity in inhibiting smooth muscle cell proliferation. These findings suggest that the genipin stent with the improved mechanical strength can be used for effective local drug delivery.
4 Drug Eluting Medical Devices
Chitosan-based materials have a wide range of applications, including in wound dressings, cartilage and bone grafts, and nerve guidance conduits. In recent years, much attention has been paid to chitosan-based devices owing to their min-imal foreign body reactions, intrinsic antibacterial nature, biocompatibility, bio-degradability, and ability to be molded into various geometries and forms, such as porous or tubular structures, enabling cell ingrowth and nerve conduction [185]. Degradable polymeric implants eliminate the need for a second surgical operation and can prevent some of the problems that are associated with stress shielding dur-ing post-healdur-ing, and they can be used simultaneously to deliver therapeutic drugs for treating infections or to deliver growth factors to accelerate new tissue growth [186].
Many investigations have demonstrated that tissue regeneration is optimized by the association of proteins or growth factors with a sustained release carrier. Chitosan-based devices have been shown to deliver such agents in a controlled fashion. The following sections discuss chitosan-based drug-eluting medical de-vices for use in tissue engineering and their promotion of the ingrowth and biosyn-thetic ability of tissues.
4.1 Wound Dressing/Artificial Skin
Wound healing is a complex process that involves inflammation, granulation tissue formation, extracellular matrix deposition, angiogenesis, and remodeling [187]. Chitosan oligosaccharides are known to have a stimulatory effect on mac-rophages and polymorphonuclear cells, and chitosan is a chemo attractant for neu-trophils. Accordingly, chitosan promotes the formation and re-epithelization of granulation tissue. Therefore, it is useful for healing open wounds [188]. Many studies have shown that chitosan-based materials in the forms of non-wovens, nanofibers, composites, films and sponges can accelerate wound healing and der-mal regeneration [189]. Various wound dressings that are made from chitosan are commercially available. These include Chitosan Skin, HemCon, TraumaStat and others.
To increase healing efficacy, chitosan can be used in combination with many pharmacological agents, such as growth factors, antibiotics, and antibacterial agents. The application of growth factors that can induce fibroblast and/or endo-thelial cell proliferation for healing-impaired wounds may increase the rate and extent of formation of granulation tissue, and thereby stimulate wound repair. Many growth factors have been incorporated into chitosan-based devices to
accel-erate healing; these include basic fibroblast growth factor (bFGF, FGF-2) [190
–
193], platelet-derived growth factor (PDGF) [194], and recombinant human epi-dermal growth factor (rhEGF) [195].
However, owing to the high diffusibility and the very short half-life time of growth factors, their use in treating healing-impaired wounds is not always suc-cessful. A photocrosslinkable chitosan hydrogel can control the release of fibro-blast growth factors, serving as a drug carrier and inducing neovascularization in vivo [190,196]. Chitosan hydrogel gradually releases incorporated FGF-2 mole-cules upon its biodegradation, and chitosan hydrogels that incorporate FGF-2 greatly improve wound healing in healing-impaired db/db mice.
DNA-incorporating chitosan matrices have been developed as platforms for gene delivery and as reservoirs for the localized and sustained expression of growth factors [197]. Plasmid DNA, encoding vascular endothelial growth factor-165 (VEGF-factor-165)/N,N,N-trimethyl chitosan chloride (TMC) complexes were load-ed into a bilayer porous collagen-chitosan/silicone membrane dermal equivalents, which were used in the treatment of full-thickness burn wounds [198]. TMC/pDNA complexes can be released in a sustained manner from collagen-chitosan scaffold for up to 28 days. The supercoiled structure of the released DNA remains, but its content decays with incubation time. Relevant results indicate that the TMC/pDNA-VEGF group had more newly-formed and mature blood vessels than other groups, and the faster regeneration of the dermis.
Wound infection is one of the most important factors that delay healing. As well as facilitating normal physiological wound repair, a dressing can importantly prevent infection. Adding antimicrobial agents to wound dressings is an effective method. The direct delivery of such agents to the wound site is favorable, espe-cially when systemic delivery can cause organ damage because of toxicological issues that are associated with the preferred agents. Many groups have successful-ly used chitosan-based wound dressings for the controlled release of antimicrobial agents to prevent infection.
As noted above, polyelectrolyte multilayers that are assembled from cationic chitosan and anionic polymer are useful platform for loading drugs inside multi-layers. Since LbL is assembled by electrostatic interaction or hydrogen bonding, charged or polar substances can be sandwiched within the multilayers. This ap-proach has also been used to fabricate drug-loaded wound dressing. Tetracycline (TC), the most effective antibiotics, was sandwiched between a poly(vinylacetate) (PVAc) layer and a chitosan/alginate nanosheet (named PVAc-TC-nanosheet) [199]. Under physiological conditions, TC was released from the nanosheet for 6 h. The potential efficacy of the PVAc-TC-nanosheet as an anti-microbial material was studied using a murine cecal puncture model. An in vivo study showed that
treatment with a PVAc-TC-nanosheet increases survival rates and reduces the in-cidence of bacterial peritonitis.
Silver has been used as an antimicrobial agent for a long time in the forms of metallic silver, silver sulfadiazine ointments, or silver nanoparticles. A chitosan-based wound dressing with improved hemostatic and antimicrobial properties was prepared by incorporating a procoagulant (polyphosphate) and an antimicrobial agent (silver) [200]. The silver-loaded wound dressing has been found to exhibit a significantly greater bactericidal activity than the unloaded control, completely killing Pseudomonas aeruginosa and consistently killing > 99.99% of Staphylo-coccus aureus. A bilayer chitosan wound dressing, consisting of a dense upper lay-er (skin laylay-er) and a sponge-like lowlay-er laylay-er (sublaylay-er), has been designed for the controlled delivery of silver sulfadiazine to treat infected wounds [201]. An in vitro drug release study reveals that the silver that is released from a bilayer chi-tosan dressing has a slow release profile. In vivo antibacterial tests have confirmed that such a wound dressing effectively inhibits the growth of Pseudomonas aeru-ginosa and Staphylococcus aureus in an infected wound over the long term.
4.2 Cartilage Graft
Chitosan is regarded as a potential material for use in cartilage tissue engi-neering to modulate chondrocyte morphology, in differentiation, and in the stimu-lation of chondrogenesis [202]. Chitosan-based scaffolds not only efficiently sup-port chondrogenic activity [203,204], but also allow the ability of chondrocytes to synthesize cartilage extracellular matrices (ECM) proteins [204]. Chondrocytes that are cultured in chitosan scaffolds may maintain their round shape, a normal phenotypic characteristic [205].
Although much attention has been paid to chitosan scaffolds, their use is re-stricted because of limited cell adhesion [206,207]. Efforts have been made to solve this problem by incorporating ECM molecules into chitosan matrices. Carti-lage ECM is composed of mainly type II colCarti-lagen and glycosaminoglycans (GAGs). To mimic the natural environment of cartilage ECM, chitosan is com-monly combined with collagen or/and hyaluronan as a fundamental material for regenerating cartilage. A recent investigation demonstrated that chitosan fibrous scaffolds that are coated with Type II collagen promote the attachment and distri-bution of mesenchymal stem cells as well as chondrogenic differentiation. Addi-tionally, the number of cells, the matrix production, and the expression of genes that are specific to chondrogenesis were improved [206].
Many studies [208,209] have addressed the regeneration of osteochondral de-fects by promoting the repair of articular cartilage, which can be further enhanced using controlled release approaches. An ideal scaffold for tissue engineering not only provides a temporary three-dimensional support of the formation of tissues,
but also acts as a carrier for important signal molecules [210]. The cationic nature of chitosan allows it to electrostatically interact with anionic GAGs and proteo-glycans that are distributed widely throughout the body and with other negatively charged species [211]. This property is critical because many cytokines/growth factors are known to be modulated by GAGs [212].
Insulin-loaded chitosan particle-aggregated scaffolds were developed as a controlled release system [210]. Insulin was used as a potent bioactive substance to induce chondrogenic differentiation. In vitro release studies have shown that the concentrations of insulin that are released to the medium were correlated closely with initial loading. Prechondrogenic cells (ATDC5) that are seeded in 5% insu-lin-loaded scaffolds typically exhibit a chondrocytic round morphology with visi-ble cell condensation. Insulin-loaded scaffolds up-regulated the Sox-9 and aggre-can expression of ATDC5 cells after 4 weeks of culture above those of unloaded scaffolds.
To promote the ingrowth and biosynthetic ability of chondrocytes, chitosan-based scaffolds were adopted to deliver growth factors in a controlled manner. Transforming growth factor beta1 (TGF-β1) is an important regulator of the pro-liferation and differentiation of chondrocytes and can promote the synthesis of specific ECMs [213]. Porous collagen/chitosan/GAG scaffolds that are loaded with TGF-β1 reportedly promote cartilage regeneration for cartilage defects [211,214]. A porous freeze-dried chitosan scaffold that incorporates TGF-β1-loaded microspheres has also been used to treat cartilage defects. TGF-β1 is re-leased in a sustained fashion, promoting the proliferation of chondrocytes and ma-trix synthesis [214].
Like the direct delivery of recombinant proteins, gene therapy can be used to introduce growth factors. Gene therapy favors local expression more than does the continuous injection of recombinant growth factors, because DNA is more stable and flexible than proteins, and is therefore likely to be compatible with established sustained delivery systems. The feasibility of the use of gene-activated chitosan-gelatin matrices for primary chondrocytes culture and the expression of the TGF-β1 gene in vitro has been established [215]. The incorporation of plasmid DNA in-to the scaffolds is associated with burst release in the first week and sustained re-lease for the following two weeks. A gene that is the transfected into chondrocytes expresses TGF-β1 protein stably for three weeks. The histological and immunehis-tochemical results confirm that the primary chondrocytes that are cultured into the chitosan-gelatin scaffold maintained their round shape and their owned characters of high secretion of specific ECM.
4.3 Bone Graft
Ideal bone substitutes should exhibit osteogenic, osteoinductive, and oste-oconductive properties. They should be resorbed and gradually replaced by newly formed bone [216]. Chitosan has been suggested as a potential material for pro-moting bone regeneration because of its apparent osteoconductive [217] and bio-degradable properties. It has been shown to promote the growth of osteoblasts and the mineral rich matrix deposition in culture [217]. Additionally, chi-tosan promotes the recruitment and attachment of osteogenic progenitor cells, fa-cilitating bone formation [218].
However, after grafting chitosan matrices, bone forms over a long period (of many months or years) [219]. The chitosan matrix serves physically in bone de-fects as a bone substitute, and seems not to suffice to induce rapid bone regenera-tion in the initial stage of bone regeneraregenera-tion [219]. Chitosan matrices must have the additional function to shorten the bone-forming period and improve the bone forming efficacy. Chitosan has been demonstrated to be effective in regulating the release of bioactive agents. Incorporating bioactive materials such as growth fac-tors may be very advantageous for improving bone-forming efficacy [219
,
220].Platelet-derived growth factor-BB (PDGF-BB) is an osteoblastmitogen and chemotaxin that has been shown in many preclinical studies to accelerate bone-healing [221]. A PDGF-BB-loaded chondroitin-4-sulfate-chitosan sponge is re-portedly useful in controlling PDGF-BB release and physically serves as a scaf-fold to facilitate osteoblast proliferation. Owing to the interaction between posi-tively charged PDGF-BB and negaposi-tively charged chondroitin sulfate, incorporated chondroitin sulfate effectively controls the release of PDGF-BB from the sponge and increases the porosity of the sponge. Relevant results indicate that PDGF-BB that is released from the sponge retains its biological activity and enhances the migration and proliferation of osteoblasts [219].
The delivery of multiple growth factors involved in tissue regeneration could mimic the conditions of natural tissue formation. In the bone-forming process, the optimal release kinetics of growth factors must be established, not only with re-spect to local concentrations but also to the duration of action of the growth fac-tors in the damaged tissue. A brushite-chitosan system has been developed to real-ize physiologically relevant VEGF/PDGF profiles for bone repair [216]. PDGF acts in the first stage of bone repair, and so this growth factor is dispersed in the brushite for fast release [222]. VEGF acts after PDGF. Consequently, VEGF is pre-encapsulated in alginate microspheres that are present in the chitosan sponge, reducing the release rate and thereby prolonging the presence of the GF at the damage site.
In vivo studies have shown that 80% of the PDGF was released from the brushite-chitosan composite scaffold in two weeks, while only 70% of VEGF was delivered in three weeks. Both growth factors that were released from the con-structs remained near the implantation site (5 cm) with negligible systemic expo-sure. The results indicate that the brushite-chitosan system can control the release rate and localization of both growth factors in a bone defect. PDGF/VEGF-loaded brushite-chitosan scaffolds greatly promoted bone formation [216].
Another investigation also demonstrated the synergistic enhancement of bone formation by both growth factors in combination [223]. Bone morphogenetic pro-teins (BMPs) can improve bone formation by inducing the chondroblastic and os-teoblastic differentiation of mesenchymal stem cells (MSCs) [224]. BMP-2 and BMP-7 have been shown to be the most effective BMPs for stimulating complete bone morphogenesis [225], and have been approved by FDA for clinical use [226,227]. A chitosan-based scaffold that contains two populations of nanocap-sules for delivering BMPs sequentially has been developed [223]. Poly(lactic acid-co-glycolic acid) nanocapsules that are loaded with BMP-2 and poly(3-hydroxybutyrate-co-3-hydroxyvalerate) nanocapsules that are loaded with BMP-7 enable the early release of BMP-2 and the longer-term release of BMP-7. The re-sults indicate that sequential growth factor delivery is more effective than the use of an individual growth factor in tissue engineering because it mimics the natural process of healing.
Chitosan-based composite scaffolds are also used as drug delivery systems in the antibiotic treatment of osteomyelitis, which is a common bone disease that is caused by a bacterial infection of the bone modullar cavity, cortex, and/or perios-teum upon the implantation [228,229]. Macroporous chitosan scaffolds that are re-inforced by β-tricalcium phosphate (β-TCP) and calcium phosphate invert glasses were originally designed as both drug carriers for controlled drug release and scaf-folds for bone regeneration [228]. Related results indicate that incorporating β-TCP and glass into the chitosan matrix effectively reduces the initial burst release of antibiotic gentamicin-sulfate (GS) from the composite chitosan scaffolds, per-haps because calcium phosphates are more stable than chitosan in physiological media and have a much lower dissolution rate therein. Therefore, controlled bio-degradation and drug release can be achieved by the hybridization of chitosan and calcium phosphates. MG63 osteoblastlike cells that are seeded on the composite scaffolds grow and migrate into the scaffolds, suggesting good cell biocompatibil-ity of the composite scaffolds. Calcium phosphate/chitosan composite scaffolds were originally fabricated for potential use as a bone substitutes and a synergistic drug carrier, and they can be utilized to deliver biological active agents to promote bone regeneration.
4.4 Nerve Guidance Conduit
Large-gap nerve damage that cannot be directly repaired using sutures has usually been treated using nerve autografts, but this approach suffers from donor site morbidity, inadequate return of function, aberrant regeneration, and shortage of donor tissue [230,231]. An alternative method is to use a nerve guide conduit. Nerve conduits (NC) are tubular structures that are used to bridge the gap of a severed nerve, and thereby acting as guides for the regenerating axons, and as bar-riers to the in-growth of scar-forming tissue [232].
Chitosan-based composites have also been regarded as promising materials for nerve repair [230,233
–
236]. In particular, nerve guide conduits that are made from chitosan and poly(caprolactone) (PCL) nanofibrous matrix have excellent mechanical and biological properties in vitro [230]. Additionally, a pilot in vivo study demonstrated the regeneration of nerve fibers after implantation of the chi-tosan-PCL fibrous matrices in a sciatic nerve defect in rats for one month. Some groups have also attempted to use a nerve graft that comprises an outer mi-croporous conduit of chitosan and interior oriented filaments of polyglycolic acid (PGA) for bridging a 30 mm defect in sciatic nerves in six Beagle dogs [236]. Six months post-operatively, in the chitosan/PGA graft group, the dog sciatic nerve trunk had been reconstructed with restored nerve continuity and functional recov-ery, and its target skeletal muscle had been re-innervated, improving the locomo-tion of the the limb that had been operated upon. These investigalocomo-tions establish the feasibility of a chitosan-based nerve graft for regenerating peripheral nerves.However, physical nerve guidance by an NC may not suffice to optimize re-covery [232]. The transience of the increase in growth factor expression following nerve injury, mainly by Schwann cells [237], is thought to compromise the ability of cut axons to maintain their regenerative state [238]. The prolonged delivery of exogenous growth factors to the injured nerve may thus sustain the drive to regen-erate injured axons when endogenous growth factor expression has been down-regulated [239].
Attempts to improve the regeneration process by including growth factors in the nerve conduit have had partial success [240,241
]
. One method of loading con-duits with growth factors involves blending of the factor with the scaffold polymer, making the factor an intrinsic part of the conduit, which is released as the conduit is degraded. This method has been used to integrate glial cell line-derived neu-rotrophic factor (GDNF) into chitosan conduits before experimental sciatic nerve injuries are inflicted. The GDNF-loaded conduit can promote axon regeneration in the early stages of recovery around six weeks post-implantation, but after 9 and 12 weeks, no differences between the GDNF that contained the conduit and the emp-ty chitosan controls can be observed [240].Like growth factors, laminin peptides facilitate nerve regeneration. One study utilized GDNF-laminin-blended chitosan (GLC) nerve guides in rat peripheral nerve injury models to promote functional nerve recovery, as determined by gait analysis and measurements of muscle mass [241]. The results indicated a greater functional restoration in the group that was treated with GLC than was achieved using the unblended chitosan nerve guides. Muscle weights of the GLC group re-vealed less atrophy and greater restoration of functional strength than in the un-blended control groups. Additionally, according to behavioral testing, the GLC group regained sensation while the control groups exhibited no restoration. The study that involved those tests verified that adding GDNF and laminin to chitosan nerve guides accelerated both functional and sensory recovery.
Although modification of conduits by the inclusion of growth factors has been shown to promote nerve regeneration under some conditions, little is known about the effect of their delivery kinetics on peripheral nerve regeneration [232]. The timing of growth factor delivery probably is important in determining the de-gree of axon regeneration [239]. Clearly, further development and refinement of the drug delivery technique are required.
One group has begun to develop an NC with an adjustable rate of release of nerve growth factor (NGF) [232]. The NC, which comprises a polyelectrolyte al-ginate/chitosan complex, is coated with layers of poly(lactide-co-glycolide) to control the release of embedded NGF. The related study revealed that the release kinetics could be efficiently adjusted by incorporating NGF at various radial loca-tions within the NC. The release of bioactive NGF in the low nanogram per day range was sustained for at least 15 days. The designed NC are thus promising can-didates for exploring the effect of release kinetics on nerve regeneration in the fu-ture.
