Biodegradable Ppf 3d Scaffold By Photopolymerization Biology Essay


The first known strategy for the preparation of PDPs was through the polycondensation of depsipetide monomeric units using methods similar to those used in peptide synthesis. This method is particularly troublesome turning its implementation at large scale impracticable [124-125]. In 1985, Helder proposed the synthesis of PDPs by the ROP of morpholine-2,5-dione derivatives [124,126] (Error: Reference source not found) which still is the most used method.

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The morpholine-2,5-dione derivatives can be copolymerized with other cyclic monomers or macrodiols (Error: Reference source not found), yielding materials with unique and tailored properties [124].

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Recently, Katsarava [127] proposed the synthesis of PDPs by the active polycondensation method, where a nitrophenyl ester derived from an α-hydroxy acid is made to react with a diester salt derived from an α-amino acid (Error: Reference source not found).

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As already mentioned, PDPs show interesting features that make these materials suitable for applications in the biomedical field, namely in DDS and tissue engineering [122b].

Ouchi [128] prepared microspheres based on PLGA for the controlled release of proteins using PDP-b-PLLA block copolymers as biodegradable surfactants. The purpose of such study was to evaluate the effect of the PDP block copolymer in the entrapment efficiency of the protein and in its release profile. It was found that in the presence of the ionic PDP-b-PLLA, the proteins had a sustained release for more than two months without the initial burst release. This was attributed to the electrostatic interactions between the protein and the ionic block copolymer.

Xie [129] prepared the amphipilic polymer poly((lactic acid)-co-[(glycolic acid)-alt-(l-glutamic acid)])-block-poly(ethylene glycol)-block-poly((lactic acid)-co-[(glycolic acid)-alt-(l-glutamic acid)]) (P(LGG)-PEG-P(LGG)) to be used in the controlled release of paclitaxel. The amphiphilic polymer had carboxyl pendant groups that were used to covalently attach paclitaxel. The authors found that the polymer drug-conjugate self-assembled into micelles in aqueous medium with a unimodal distribution and a mean size of 119 nm. In vitro release tests showed that the paclitaxel release is pH dependent, being faster for lower pH. The in vitro antitumour activity of the conjugate against rat brain glioma C6 (RBG-6) cells was evaluated and the results showed almost the same cytotoxicity as paclitaxel alone.

A poly(lactic acid-co-L-lysine) (PLA-PLL) copolymer bearing pendant amine groups was prepared by Yu [130] for application in cancer treatment. The tripeptide arginine-glycine-aspartic acid (RGD) was covalently attached to the pendant amine groups, yielding a polymer-tripeptide conjugate (PLA-PLL-RGD). More recently, Liu [131] prepared nanoparticles from the PLA-PLL-RGD conjugate and their applicability as a targeted delivery vehicle for cancer treatment was evaluated in vivo. The results of target imaging and biodistribution showed that the nanoparticles can significantly target to the tumour of Bacp-37 breast cancer bearing mice, which means that such nanoparticles have potential to be used as targeted delivery carrier.

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Ohya [132] prepared biodegradable ABA-type triblock copolymers, PDP-PEG-PDP, using glycine (PGG-PEG-PGG), L-leucine (PGL-PEG-PGL) and L-phenylalanine (PGF-PEG-PGF) in the PDP block. The effect of the -amino acid pendant group in the thermoresponsive behaviour of the resulting triblock copolymer was studied. It was found that the glycine-based copolymer did not exhibit any thermoresponsive behaviour, contrarily to the L-leucine and L-phenylalanine containing triblock copolymers. PGF-PEG-PGF exhibits a cloud point 10 °C lower than that of PGL-PEG-PGL. Moreover the BAB-triblock copolymer containing L-leucine (PEG-PGL-PEG) showed a thermoresponsive sol-gel transition. The authors believe that such block copolymer systems can be used as injectable formulations for both drug delivery or scaffolds for tissue engineering.

3.3. Poly(ortho esters)

Poly(ortho esters) (POE) are biocompatible and bioerodible hydrophobic synthetic polymers. These polymers have high stability in alkaline medium due to their resistant ortho ester linkages [133].

The first POE to be applied as a biomaterial was commercialized by Alza Corporation in the 1970s under the trade name Chronomer and later Alzamer®, and was designed for applications in DDS [134].

Currently there are four different classes of POEs (I to IV) [133,135].

The synthesis of POE I (Error: Reference source not found) occurs by transesterification between a diol and a dietoxytetrahydrofuran. In this reaction, γ-hydroxybutyric acid is produced, due to an autocatalytic effect while the degradation of the polymer occurs.

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In an aqueous environment, the polymer decomposes producing a diol and γ-butyrolactone (Error: Reference source not found), that is rapidly converted into γ-hydroxybutyric acid. An additional erosion of the polymer occurs due to the catalysis of the hydrolysis reaction by the γ-hydroxybutyric acid [6a,101b,133].

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In order to prevent the production of this acid, the synthesis of POE II was proposed. POE II is obtained by the condensation of diols or triols with the diketene acetal 3,9-bis(ethylidene-2,4,8,10-tetraoxaspiro[5,5]undecane) (Error: Reference source not found). From the degradation of POE II only neutral molecules are obtained. The rate of degradation of this polymer can be accelerated by adding an acid [101b].

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The condensation of a flexible triol with an alkyl orthoacetate allows the preparation of POE III (Error: Reference source not found). POE III is a semi-solid material at room temperature, due to its very flexible backbone. This brings advantages such as the incorporation of therapeutic agents at room temperature without the use of solvents, and also the fact that these materials can be easily injectable.

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POE IV was developed due to the high viscosity of POE III and consequent difficulties in the scaling up of its synthesis. POE IV (Error: Reference source not found) is obtained by the incorporation of short segments based on LA and GA in the polymer backbone of POE II. Compared to POE II, this allowed a considerable degradation rate without the addition of an acid excipient [133,136].

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Within the four different classes of POEs, POE IV is a promising biomaterial, because of its good biocompatibility, accessible synthetic procedure, and also its ability to offer well-controlled release profiles for various pharmaceutical agents, including proteins [1b,101b,135-136].

The incorporation of different diols in the POE structure influences the rate of degradation of these polymers, their pH sensitivity and the transition temperature. The use of basic or acidic excipients and the pH sensitivity of POEs turn these polymers very suitable materials for DDS [135].

The delivery of DNA vaccines is a possible application of POE carriers. Wang [137] showed that DNA shielded with POE microspheres was protected from degradation, enabling uptake by antigen-presenting cells. The release of DNA was also facilitated in response to phagosomal pH [137]. Furthermore, the POE microspheres did not produce any cytotoxic effect. Recently, Nguyen [138] proposed a new approach for the release of DNA vaccines with a 'modified' microparticulate system based on POE's combined with PEI. PEI increases POE microsphere gene transfection efficiency and slows the DNA release rate.

Sustained release of analgesic [139] and antiproliferative [140] drugs, insulin [141] and contraceptive esteroids [142] were also achieved using POEs-based carriers.

3.4. Polyanhydrides

Polyanhydrides are a class of polymers with attractive characteristics for DDS, namely a controlled rate of degradation [30] (Error: Reference source not found). This class of materials was approved by the Food and Drug Administration (FDA) in 1986 as a vehicle for drug delivery [101b,143]. Among the different types of bonds in polymer structures, the anhydride bond is one of most labile in the presence of water due to a facile hydrolysis process. The exception is some aromatic polyanhydrides which degrade after a few years while generally aliphatic polyanhydrides degrade in days. Therefore, the development of aliphatic-aromatic copolymers allows the control of degradation rates according to the monomer composition [1b,6a,143].

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The ability to control the hydrolysis reaction within the polymer device and the improvement of their mechanical properties are two critical parameters that need to be carefully considered in the use of polyanhydrides [143].

The main disadvantage of polyanhydrides as DDS is related with the fact that these polymers can react with drugs containing free amine groups or other nucleophilic functional groups, especially during encapsulation at high temperatures, limiting the type of drug that can be successfully incorporated within this polymer [1b]. Polyanhydrides have, in general, excellent biocompatibility in vivo [144]. In recent years, polyanhydrides have been mainly investigated to deliver proteins [145], for use in tissue engineering [146] and to induce immune protection [147]. Gliadel® wafer, an implant from Eisai Inc., is nowadays the only polyanhydride-based biomedical product in the market being used to treat certain malignant brain tumours.

3.4.1. Polyanhydride copolymers

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The most extensively studied polyanhydride copolymer is poly[(1,3,bis-p-carboxyphenoxypropane)-co-(sebacic anhydride)] (CPP-SA). This copolymer is approved by the FDA for use in biomedical devices, and used in the treatment of brain cancer, for the delivery of carmustine, an alkylating agent in chemotherapy [148].

For orthopedic applications, polyanhydrides are not the ideal candidates due to their poor mechanical strength. For example, the Young's modulus of the polyanhydride poly(1,6-bis-(carboxyphenoxy) hexane) is only 1.3 MPa. In order to improve the mechanical properties of polyanhydrides, copolymers such as poly(anhydride-co-imide)s were developed. These copolymers present better surface erosion properties than polyanhydrides. For application in bone tissue engineering, as scaffolds, Laurencin and Nair [101b] investigated the mechanical properties and the biocompatibility of poly[pyromellitylimidoalanine-co-1,6-bis(p-carboxyphenoxy)hexane] (PMAala:CPH) (Error: Reference source not found). This polymer exhibited good mechanical performance and good osteocompatibility. Other poly(anhydride-co-amide)s based on succinic acid, trimellitylimidoglycine (e.g., poly[trimellitylimidoglycine-co-1,6-bis(p-carboxyphenoxy)hexane]) and trimellitylimidoalanine, had high compressive strength and were suitable for orthopedic applications [101b,149]. The degradation of poly(anhydride-co-imide)s occurs firstly in the anhydride bonds and then, by hydrolysis, in the imide bonds.

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For intramuscular or dermal applications, poly(terephthalic acid-sebacic acid) (P(TA-SA)) was used to study the local tissue response by the implantation of polymer samples into the cornea of rabbits. No inflammatory response was observed during a six weeks implantation period [148b].

Poly(anhydride-esters) are considered to be useful for biomedical applications because of their favourable surface erosion degradation, and hydrolytic bond cleavage that leads to the release of active therapeutics and water-soluble biocompatible by-products. In the case of salicylic-comprised poly(anhydride-esters), their biodegradation releases salicylic acid that is therapeutically effective as an antiseptic, an analgesic and an anti-inflammatory drug [146b].

For the delivery of antigens, adjuvants and for the stimulation of dendritic cells, copolymers based upon sebacic acid (SA), 1,3-bis(p-carboxyphenoxy)propane (CPP), 1,8-bis-(p-carboxyphenoxy)-3,6-dioxaocatane (CPTEG) and also 1,6-bis(p-carboxyphenoxy) hexane (CPH) have been studied [147c].

Carrillo-Conde [147b] used a CPTEG:CPH copolymer to prepare nanoparticles whose surface was subsequently 'decorated' with sugar residues (dimannose and lactose). The authors figured out that the nanoparticles containing the dimannose residue at the surface had similar composition to pathogen surfaces and were efficiently internalized by dendritic cells.

In order to improve the mechanical properties of polyanhydrides, the incorporation of acrylic functional groups in the monomer units allows the preparation of crosslinkable polyanhydrides. These injectable anhydrides can be used for filling bone defects or to repair soft tissues. Error: Reference source not found shows the structure of end capped polymers of methacrylated anhydride monomers of sebacic acid (MSA) (A), 1-3 bis(p-carboxyphenoxy)propane) (MCPP) (B) and 1,6-bis(carboxyphenoxy) hexane (MCPH) (C) used by Muggli and Kumar [143,150]. The degradation time of these polymers changed with the network composition, from 2 days for poly(MSA) to 1 year for poly(MCPH).

The products of hydrolytic degradation of these polymers, diacid and linear molecules of methacrylic acid, are non-toxic, non-mutagenic and biocompatible both in vitro and in vivo. This class of polymers has compressive strength comparable to the lower range of cancellous bone (30-40 MPa) [101b,147c].

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Xing [151] showed that mPEG-b-poly(octadecanoic anhydride)-b-mPEG (mPEG-b-POA-b-mPEG) is a potential candidate delivery system for hydrophobic drugs such as paclitaxel, commonly used for cancer chemotherapy.

Poly(fatty acid dimer-sebacic anhydride) (P(FAD-SA)) copolymers have found a wide-range application in controlled drug delivery due to their low melting point, hydrophobicity and flexibility of the polymer [148b,149c].

3.5. Poly(iminocarbonates)

Poly(iminocarbonates) are structurally similar to polycarbonates, with the substitution of the carbonyl group for an imine group. Synthetically they derive from the reaction of a diol and a dicyanate compound. Polycarbonates have excellent mechanical strength and are easily processed, but do not degrade under physiological conditions. The replacing of the oxygen of the carbonyl group in polycarbonate by an imine group decreases the stability of the polymer due to the sensitivity of the imine group to hydrolysis (Error: Reference source not found - (A)). The products of the biodegradation of poly(iminocarbonates) in water are ammonia, carbon dioxide and bisphenol (Error: Reference source not found - (B)) [6a,152].

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Many other structurally similar polymers have been prepared starting from less toxic bisphenols. For example, the derivatives of tyrosine amino acid have been used to produce poly(iminocarbonates) with good mechanical properties, such as poly(desaminotyrosyl-tyrosine hexyl ester iminocarbonate) (Error: Reference source not found) [6a,153].

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A similar approach can be used to produce a polycarbonate of derivatives of tyrosine dipeptides (Error: Reference source not found) where the physico-chemical properties of the final polymer can be modified by changing the nature of the R groups [6a,154].

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Regarding the biocompatibility of the polymers of Error: Reference source not found and Error: Reference source not found, when implanted subcutaneously in rats, the tissue response was similar to the one produced by implantation of PLA or polyethylene discs [6a]. Poly(iminocarbonate)s are good candidates for implantable devices useful as prosthesis or DDS [155].

3.6. Polymers containing phosphorus

Phosphorus-containing polymers are interesting materials for biomedical applications due to their good biocompatibility and hemocompatibility, and protein adsorption resistance. They have been widely used in dentistry, regenerative medicine and drug delivery areas as well described by Monge in his recent review [156].

3.6.1. Polyphosphazenes

Polyphosphazenes (Error: Reference source not found) are polymers with a main chain alternating among phosphorus and nitrogen atoms, containing two organic or organometallic lateral groups attached to the phosphorus atom. These polymers have a structure that can be considered in the interface between organic and inorganic polymers, presenting unusual properties [1b,6a,101b,157]. They were developed during the 60s by Kugel and Allcok [158]. Numerous biological compounds, proteins or peptides, have been attached to the pendant chains of their chemical structure, being then released via hydrolysis [159].

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The chemical structure of polyphosphazenes provides considerable flexibility in the design of biomaterials. By selecting the side groups of the polymeric chain it is possible to control the hydrophobicity/hydrophilicity ratio, crystallinity, solubility, thermal transitions and degradation rate of the polymer [101b]. The most investigated method for polyphosphazene synthesis is based on the use of a reactive macromolecular intermediate, poly(dichlorophosphazene) (Error: Reference source not found - (B)), produced by the ROP of the cyclic trimer, hexachlorocyclotriphosphazene (Error: Reference source not found - (A)). Then, the chlorine side units in this polymer are replaced by an organic side group. Another method for the synthesis of polyphosphazenes with well-controlled structure is the living cationic polymerization process, involving a catalysed condensation of the monomer, (CH3)3Si-N = PCl3, with loss of (CH3)3SiCl. This polymerization process allows the synthesis of phosphazene-based block copolymers at room temperature with controllable chain lengths and narrow molecular weight distribution [6a,101b,157].

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A wide range of bioactive compounds can be linked to polyphosphazene molecules allowing the formation of bioactive water-soluble macromolecules or polymer surfaces with biological activity [6a,160].

The nitrogen-phosphorus linkage is hydrolytically stable. The incorporation of groups such as imidazole, glucosyl or aminoacid ester makes the hydrolysis process easier to occur and also allows the preparation of an extensive variety of structures derived from polyphosphazenes [157].

Weikel [161] developed phosphazene tissue engineering scaffolds with bioactive side groups using choline chloride, a biological buffer. The authors synthesized cyclic phosphazene trimer mixed substituents (as model systems) and polymers with choline chloride and glycine ethyl ester, alanine ethyl ester, valine ethyl ester, or phenylalanine ethyl ester. Two different synthetic processes were used to prepare these polymeric species: a sodium hydride mediated route which resulted in polyphosphazenes with a low amount of choline and a cesium carbonate mediated route giving polyphosphazenes with high choline content. The polymers were mixed with PLGA (50:50) or PLGA (85:15). The results demonstrated that the polymers obtained through the first process were miscible with both ratios of PLGA, whereas the polymers obtained through the second route had reduced miscibility. In vitro results demonstrated that the choline-based blends, compared to PLGA, supported primary rat osteoblast cell growth with an elevated osteoblast phenotype expression.

Recently, Potta [162] developed photo-crosslinkable and thermoresponsive polyphosphazene hydrogels with properties of mechanically suitable strength and controllable biodegradation for injectable biomedical applications. Polyphosphazenes were modified with methacrylate groups and then photo-crosslinked upon UV light under mild conditions, resulting in the formation of compact three-dimensional networks. The thermoresponsive hydrophobic interaction forces in the polymer network at body temperature facilitated the rapid dual crosslinking accomplishment of the photo-crosslinking even under mild conditions, having a great potential as injectable, biodegradable, and controllable carriers for various biomedical applications.

Polymethyl methacrylate polyphosphazene microspheres were CE marked in 2006 and FDA approved in 2008 for the embolization of hypervascularized tumours and arteriovenous malformations [163].

Polyphosphazenes have been investigated for orthopedic applications [164] and as immunologic adjuvants in vaccines [165]. One of the most studied polyphosphazenes for biomedical applications are poly(amino acid ester phosphazenes) (Error: Reference source not found), which are highly osteocompatible [101b].

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Recently, an amphiphilic graft copolymer comprising a polyphosphazene backbone, PEG branches and a pH-sensitive moiety was prepared by Zheng [166]. The copolymer self-assembled in aqueous medium and presented the ability to encapsulate both hydrophobic and hydrophilic drugs, as demonstrated by the results obtained with Nile red and dextran-fluorescein isothiocyanate (FITC) as model drugs. The vesicles were also tested as carriers for DOX. A sharp pH-responsive drug release behavior was observed when the pH decreased from 7.4 to 5.5 and an accentuated enhancement of DOX cytotoxicity to DOX resistant MCF-7 human breast cancer cells.

3.6.2. Polyphosphoesters

Polyphosphoesters are an additional class of bioabsorbable polymers containing phosphorus (Error: Reference source not found), developed in the 1970s by Lapienis and Penczek [167]. This class of polymers, due to phosphoester bonds in their chain, attain good biocompatibility. Depending on the functional group linked to the phosphorous atom, polyphosphoesters can be divided in polyphosphites (R1=H), polyphosphonates (R1=alkyl group/aryl) and polyphosphates (R1=aryloxy/alkyloxy group) [168].

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Various synthetic routes including ROP, polycondensation, transesterification, and polyaddition are used to synthesize polyphosphoesters. The choice of the R and R1 groups alters the physical and chemical properties of the polymers [168b].

The degradation of polyphosphoesters occurs by hydrolytic and enzymatic cleavage of the phosphate bonds in their chain, giving phosphates, alcohols and diols as final products. Due to the pentavalency of the phosphorous atom, the polyphospazenes are able to chemically bind drugs or proteins. This is an interesting feature of this family of polymers that enables the preparation of new polymer pro-drugs, with good in vivo tissue compatibility [101b].

This class of polymers is widely used in DDS, usually as copolymers, to release chemotherapeutics [169] and in tissue engineering [170]. Polyphosphoester-based microspheres have been used as a delivery system of paclitaxel. This system is commercially known as PACLIMER®.

3.7. Natural based polymers

3.7.1. Proteins

A common strategy in the design of bioabsorbable polymers for medical applications has been to use naturally occurring monomers, since the degradation of the polymers originates non-toxic molecules. Thus, amino acids appear as an obvious choice as monomers for the synthesis of new polymeric bioabsorbable materials; moreover, the combination of the amino acids available provides a tool for the synthesis of a vast number of materials. In this section, only natural proteins such as collagen, fibrin and elastin will be discussed since they are the most commonly used in biomedical applications. Also 'pseudo'-poly(amino acids) will be discussed since they result from amino acid condensation but they do not have the same type of bonds.

Amino acids are the basic structural units of proteins. An α-amino acid comprises an amine group and a carboxyl group linked to the same carbon atom, a hydrogen atom and a group with several possible structures (R) (Error: Reference source not found) [171].

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Naturally, there are 20 types of R groups, varying in structure, size, shape, charge and ability to form hydrogen bonds. In proteins, amino acid polymerization is due to the reaction of the α-carboxyl group of one amino acid with the α-amine group of another amino acid to form an amide bond, also known as peptide bond. Error: Reference source not found shows the formation of a peptide bond from two amino acids with the loss of one water molecule [171]. The reaction is reversible which means that the hydrolytic cleavage of an amide bond originates the two starting amino acids.

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Collagen is one of the most plentiful proteins present in mammals, being the major component of skin and connective tissues, and is one of the primary initiators of the coagulation cascade [101b]. The basic collagen molecule presents three polypeptide α-chains, each consisting of more than 1000 amino acids, arranged in a unique triple-helix forming sequence [172]. Repeating triplets of (glycine-X-Y)n, where X and Y are often proline and hydroxyproline, are part of the primary structure of these proteins (Error: Reference source not found).

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The increase of flexibility of the backbone of collagen is associated to the increase of the presence of the smallest glycine amino acid [101b,172-173]. Due its small side group (R=H) and its repetition at every third position, the sequence enables close package of the chains into a helix, reducing the space for residues in the core [172] contributing to the helix assembly and mechanical strength. In order to provide suitable mechanical forces to collagen, fibrils formed by the secretion of pro-collagen molecules into the extracellular space, gather spontaneously to form in intricate structures. These fibrils can be oriented in a different way according to the tissue nature defining their mechanical properties [101b,173].

In the human body, there are over 29 identified types of collagen [174] that confer distinct biological features to the various types of tissues in the body [172]. Over 90% of the collagen is of type I, being the prominent form of collagen. Collagen type I has a tertiary structure where, wrapped in a triple helix, are three polypeptide subunits comprising approximately 1050 amino acids with approximately 33% glycine, 25% proline and 25% hydroxyproline, being also abundant in lysine [101b]. It is predominant in the skin, tendon and bone where extreme forces are transmitted [172].

In the body, enzymatic degradation of collagen occurs by collagenases such as matrix metalloproteinase (MMP), being broken down into amino acids [172-174]. The degradation rate can be changed by an enzymatic pre-treatment or by crosslinking. In this latter case, crosslinking can be achieved by heat radiation or high energy or with different types of crosslinking agents, such as difunctional or multifunctional aldehydes, carbodiimides, hexamethylene diisocyanate, and compounds of polyepoxy and esters of succinimidyl PEG. Since collagen is soluble in acidic solutions, it is possible to use this material in different forms, such as: sheets, tubes, sponges, foams, nanofiber matrices, powders, viscous injectable solutions and dispersions [172-173]. Collagen is known as an ideal material for tissue engineering [175], coatings [176] and release of pharmaceutical products [177].

Regarding biomedical applications, collagen is mostly derived from skin of porcines and bovines and tendons of equines and bovines. The slight immunogenicity transmitted by the terminal area of this material and its antigenic sites in the central helix are the key issues of collagen use. Other disadvantages are the high commercial cost of pure collagen, its allogeneic or xenogeneic derivation that increases the risk of contagious illnesses, and its variable physico-chemical degradation properties [101b].

Collagen has been used, using a variety of shapes, as a drug carrier of various active compounds, ranging from high [178] to low molecular weight [179] drugs, growth factors [180], oligonucleotides [181], and proteins [30,182].

Tissue engineering is a field where collagen-based devices demonstrated to be very useful. Collagen fibers are important constituents of tissues, and consequently in tissue, with significant prominence in tendon tissue engineering [183]. Regarding extensive skin loss from burns and other injuries, skin substitutes are the ideal treatment option. Collagen-based matrixes are the most preferred for wound repairs [184].

Kawashima [178b] developed a device that consisted of collagen microparticles embedded in a PEG membrane for transcleral drug delivery using dextran-FITC (40kDA) and recombinant human brain-derived neurotrophic factor (rhBDNF) as active compounds. This research group showed that the release of the drug followed a zero-order kinetic that can be controlled by the concentration of collagen microparticles within the membrane.

The combination of the advantages of synthetic PLGA knitted meshes and porous collagen sponges composed of natural type I collagen was studied by Dai [185], in order to obtain 3D biodegradable scaffolds. In the formation of these scaffolds the PLGA mesh served as support and the collagen microsponges enabled cell seeding and tissue formation. In this work the authors prepared three types of scaffolds with different structural designs and they demonstrated that these scaffolds could be used for tissue regeneration of articular cartilage with an adjustable thickness. The use of these devices was tested in vivo: bovine chondrocytes were cultured in these scaffolds and transplanted subcutaneously into nude mice for 2, 4, and 8 weeks. The results indicated that all three groups of transplants showed a spatially even cell distribution, natural chondrocyte morphology, abundant cartilaginous ECM deposition, and excellent biodegradation in vivo.

For use in the field of cartilage regenerative medicine, Tanaka [186] developed an implant type tissue-engineered cartilage with firmness and 3D structure: the authors combined a porous biodegradable polymer scaffold with a atelocollagen (a pepsin-solubilized type I collagen solution) hydrogel, with the objective of optimizing the structure and the composition of porous scaffolds. These devices had different polymers in their composition [PLLA, PDLA, poly(L-lactide-co-ε-caprolactone) (PLLA-CL) and PLGA], various kinds of porosities (80-95%) and pore sizes (0.3-2.0 mm). The authors optimized the structure and composition of porous scaffolds by administrating chondrocytes/atelocollagen mixtures into the scaffolds and transplanting the constructs into the subcutaneous areas of nude mice. The scaffold effectively retained the cells/hydrogel mixture, enabling its use in fair cartilage regeneration. The optimal structure for the porous scaffold in combination with the atelocollagen was regarded to be that with the porosity of 95% and pore size of 0.3 mm made by the sugar leaching method. PLGA and PLLA were shown to be the most recommended polymers to use in scaffolds of cartilage, improving the quality of the tissue-engineered cartilage.

A "smart" drug delivery device based on collagen microcapsules was prepared by Pastorino [187]. This system is sensitive to disease-associated overexpressed MMP, an enzyme responsible for cartilage damage. MMP triggers the drug release of the active compounds. Many diseases as arthritis, cancer and cardiovascular problems are related to the overexpression of MMP. The up-regulated expression of MMPs is the cause of collagen's degradation. The results showed that the shape and morphology of the microcapsules was only significantly modified in the presence of MMP. Therefore, these findings suggested that the presence of MMP was related to the release of some active compounds, present in the capsule.

Collagen has shown to be promising in cardiac tissue engineering where there is a need of regeneration of the damaged myocardium by combining cells to a biocompatible matrix [188].

Injectable collagen microspheres for the release of recombinant VEGF (rhVEGF) were prepared by Nagai [180b]. The factor was released in a sustained manner over four weeks, remaining bioactive after this period. Thus, the proposed system had very favourable features for tissue engineering applications.

Collagen-based biomaterials are also presently used as cell culture scaffolds in tissue engineering approaches [189]. As collagen is as a native component of the ECM, this polymer has been used to obtain scaffolds as these are able to mimic the native ECM [190]. A collagen scaffold was prepared with covalently immobilized VEGF. This scaffold was used as a patch in cardiac repair, which improved tissue formation as there was cell proliferation within the graft both in vitro and in vivo. This led to an increase in the blood vessel density, supporting vascularization within the graft [191].

Collagen has also been used as part of a composite tissue system for use in musculoskeletal tissue reconstruction since it improves cell attachment. Both PLLA and PCL were mixed with collagen where the PLLA/collagen side of the scaffold was the stiffest with the lowest strain, while the PCL/collagen side had the highest strain. The middle region of the scaffold possessed an intermediate stiffness and strain, analogous to the tendon, muscle, and junction respectively.

Teixeira [180c] prepared a hydroxyapatite scaffold coated with type I collagen film for the release of bone morphogenetic protein 2 (BMP-2). The presence of heparin in the scaffold reduced the initial undesirable burst release phase of BMP-2, being continuously released for seven days.

The delivery of recombinant human bone morphogenetic protein-2 (rhBMP-2) for bone formation or repair was studied using collagen and collagen-chondroitin sulfate microparticles as carriers [180d]. The authors showed that an increase of the initial burst of rhBMP-2 occurred when chondroitin sulfate was incorporated into the collagen scaffold, an essential effect in the acceleration of the bone regeneration process.

Duncan [192] studied a mechanically stable and optically transparent collagen gel from an acidified atelocollagen solution, appropriate for use as corneal stromal equivalent. It was demonstrated that even with minor changes the environmental conditions of the gels (e.g., pH) affect the optical and mechanical properties of the constructs. For this reason, a balance between the solution pH and the crosslinker concentration was established. The hydrogel was implanted in vivo into mid-depth corneal stromal pockets of male New Zealand white rabbits and the results demonstrated favourable biocompatibility, and showed the effect of higher levels of fibrillogenesis and crosslinking on increased long term survival of the gels. These gels can be used in the construction of full thickness artificial corneas, as potential functional stromal equivalents for use in stromal grafting, as drug carriers and as stromal implantation for tissue replacement and regeneration.

Collagen has also contributed to the understanding of tumour progression. Collagen I hydrogels cultured with MDA-MB-231 human breast cancer cells were bioengineered as a platform for in vitro solid tumour development [193].

Regarding tumour development, collagen I hydrogels have been used to reproduce the microenvironmental conditions of a solid tumour due to their biocompatibility and 3D architecture [193-194].

More recently, collagen-based scaffolds have been recognized as adequate for human bladder

regeneration [195] and collagen membranes have been used to repair the uterine horn defects in the rat uterine with potential application in human uterine regeneration [196]. High density collagen has also been used for the construction of tubular, smooth muscle containing structures, such as the ureter or the urethra [197].

There are some collagen-based products already on the market, as for example, InductOs® from Wyeth Pharmaceuticals for treatment of acute fractures, DuraGen® Dural Graft Matrix of Integra Lifesciences Corporation for spine surgery, NeuraGen® of Origin BioMed Inc., a topical solution or gel for nerve pain relief, or Zyderm®, Zyplast®, EvolenceTM, CosmoDerm® and Cymetra® Micronized AlloDerm® Tissue, where these last materials are collagen injectable fillers used in aesthetic dermatology.


Elastin is an ECM protein that is known for providing elasticity to tissues/organs (such as pulmonary and vascular tissues) [198]. Tropoelastin, the soluble precursor of elastin, formed by a variety of cells including smooth muscle cells, endothelial cells, fibroblasts and chondrocytes, is the main constituent of elastin. The molecules of tropoelastin are then covalently linked, leading to elastin, as a highly crosslinked and insoluble polymer [198-199]. Tropoelastin is composed of many repetitions of sequences of tetrapeptide, pentapeptide, hexapeptide and nonapetide - VPGG, VPGVG, APGVGV and VPGFGVGAG, respectively, where V is valine, P is proline, G is glycine, F is phenylalanine and A is alanine. For instance, the sequence VPGVG is repeated 50 times in one molecule [101b]. An important property of tropoelastin, also common in elastin-like peptides, is their potential to self-assemble under physiological conditions [200]. The resulting insoluble elastin has a half-life of 70 years, being one of the most stable proteins [198].

The insolubility of elastin in water and a mild immune response limits the application of this material in medical applications. Nevertheless, it has been used as an organic coating for synthetic vascular grafts as it showed minimal interaction with platelets. Elastin matrices are considered as favourable flexible biomaterials, prepared by transferring the tropoelastin solution into suitable molds, where a formation of lysine crosslinks occurs on tropoelastin initiated by the enzyme lysyl oxidase [101b,199,201]. Recently, genetically engineered elastin-like polypeptide-based hydroxyapatite bionanocomposites have been prepared to use as multifunctional bone cements [202]. Elastin-like polypeptides exhibited sequence-specific capacities to interact with ions, and bind and disperse hydroxyapatite nanoparticles.

Annabi [203] prepared elastin-based hybrid hydrogels and studied the effect of high pressure CO2 on their synthesis and characteristics. The hydrogels were prepared by chemically crosslinking tropoelastin/α-elastin solutions with glutaraldehyde at high pressure CO2. Dense gas CO2 had an important role in characteristics such as porosity, swelling ratio, compressive properties and modulus of elasticity tropoelastin/α-elastin hybrid hydrogels. The results of this work demonstrated that the generation of larger pores with an average pore size of 78-17mm was fabricated by the use of high pressure CO2. The use of high pressure CO2 also eliminated the skin formation on the top surfaces of hydrogels as well as increased their mechanical properties. In vitro results showed that the large pores throughout the hydrogel enabled the migration of human skin fibroblast cells 300 µm into the construct.

Fabrication of vascular grafts from a recombinant elastin-like protein reinforced with collagen microfiber was developed by Caves [204]. The authors obtained an acellular arterial substitute consisting of a multilamellar structure formulated from integrated synthetic collagen microfibers and a recombinant elastin-like protein. The process facilitated control over collagen density and collagen microfiber orientation. The study showed that the suture retention strength, burst strength and compliance were modulated by the fiber architecture and processing method of the elastin-like protein.

Sallach [205] developed a self-assembling, recombinant elastin-mimetic triblock copolymer (ABA triblock; A - block: VPAVG[(IPAVG)4(VPAVG)]16IPAVG; B - block: VPGVG[(VPGVG)2VPGEG(VPGVG)2]48VPGVG) that showed a robust in vivo stability for more than 1 year, causing a minimal inflammatory response, in the absence of either chemical or ionic crosslinking. The authors were able to predict the use of these elastin-mimetic triblock copolymers as structural components of artificial organs, engineering living tissues, devices for controlled drug release and biocompatible surface coatings.

Elastin has also been used to enhance the properties of PCL-based scaffolds in cartilage repairs. The integration of elastin in the scaffolds' matrix exhibited lower energy loss compared to PCL scaffolds, showing that elastin increased the elastic properties of the construct [206].

More recently, protein-based composite patches consisting of collagen microfibers in a recombinant elastin protein polymer matrix prevented hernia recurrence in Wistar rats over an 8-week period with the formation of new tissue and a sustained structural integrity [207].

For dermal tissue engineering, elastin was used to produce scaffolds by electrospinning that served as efficient dermal substitutes for the treatment of large and deep burns [208].

Matriderm® is an elastin- and collagen-based three dimensional matrix that guides autologous cells for dermal 'reconstruction'.


Fibrin, a biopolymer comparable to collagen, was one of the first biopolymers used as biomaterials. The monomeric form of fibrin is fibrinogen which undergoes fibrillogenesis as the enzyme thrombin cleaves specific fibrinopeptides, allowing the lateral association of the molecules and consequent formation of linear fibrils. The process of fibrinolysis, an important process that prevents blood clotting, leads to the breakage of the fibrin clot, where the fibrils are degraded by a series of enzymes in the body [101b,209]. Fibrin is extensively used in tissue engineering [210]. Aprotinin or other protease inhibitors have been added to fibrin to diminish its relatively rapid resorption in vivo, conferring extended longevity to fibrin matrices [211].

Fibrin contains protein nature materials ("cryptic" matrikines) that are released upon their activation/degradation, providing regenerative properties to this protein. Besides this, fibrin is biodegradable, biocompatible and has injectability properties. These characteristics led fibrin to be used as a material for regenerative medical applications in tissue and organ sealing, in plastic and reconstructive surgery, including skin grafting. Fibrin is used as fibrin matrices on endothelial cell branching [212], cell matrix material on its own or with a polymeric scaffold, and as delivery systems for bioactive and therapeutic agents [213]. Recently, Chen described fibrin scaffolds that were applied in tissue engineering on treatment of eriodontitis, an inflammatory disease that causes the destruction of the tooth-supporting apparatus, such as the alveolar bone, the periodontal ligament, and the root cementum, potentially leading to tooth loss [214].

As a commercial application, fibrin was first used as a sealant. Nowadays, fibrin sealants are used for clinical haemostasis, among other medical procedures [101b,211,215].

The proliferation and differentiation of stem cells can be attained in a fibrin matrix. Hydrogel microbeads capable of encapsulating stem cells, with the purpose of quickly degrading to release the cells, have been studied [216]. The microbeads were composed of fibrin-alginate which had encapsulated human umbilical cord mesenchymal stem cells (hUCMSCs). The results demonstrated that the cells showed good proliferation and osteogenic differentiation.

In combination with PLLA and PLGA sponges, fibrin matrices provided added mechanical strength to support in vitro construct vascularization, improving neovascularization upon implantation [217].

A hemostatic fibrin pad that is able to control surgical-related bleeding was recently developed by Harmon [218]. This pad consists of an absorbable composite matrix scaffold that is covered with human-derived active biologics. When in contact with targeted bleeding surfaces, the human biologics instantly form a fibrin clot. The results of the research group showed that the pad can increase wound healing in in vitro experiments, supporting also wound healing at the site of in vivo application.

A product already present on the market based on this bioabsorbable material is Bioseed® from BioTissue AG, which consists of a mixture of keratinocytes and fibrin, and used for treating chronic wounds, and also Tisseel® (Baxter Healthcare Corporation) and Evicell® (OMRIX Biopharmaceuticals Ltd.), both fibrin sealants.

'Pseudo'-poly(amino acids)

Poly(amino acids) are a group of poly(ionic) molecules with various biological functions which play an important role both in nature and in human life [219]. The reactive side chain groups of poly(amino acids) allow the binding of drugs or molecules and the products of degradation are amino acids.

Poly(amino acids) are highly insoluble in most common organic solvents, difficult to be processed, have low mechanical properties and are extremely antigenic [1b,6a,101b,220]. These disadvantages of poly(amino acids) limit their applicability. In order to improve their mechanical and biological properties, the amino acids were polymerized in a way that leads to the formation of bonds other than peptide bonds, resulting in 'pseudo'-poly(amino acids). The first modified chain of 'pseudo'-poly(amino acids) was introduced in 1984 [152a]. They have been used as suture materials, artificial skin replacements and also as systems for controlled drug delivery [1b]. The chemical bonds present in pseudo-poly(amino acids) are carbonate, ester or iminocarbonate bonds. The number of carbon or oxygen atoms in the polymer backbone modifies its properties such as cristallinity, Tg, hydrophilicity and hydrophobicity. The mechanical properties of the polymer can be easily designed by changing the structure of the pendant chain.

One of the most used 'pseudo'-poly(amino acids) as bioabsorbable materials are the tyrosine-derived polycarbonates (Error: Reference source not found). These types of polymers demonstrated a high degree of bone conductivity, good biocompatibility and processability [1b,221].

Recently, an interesting review about protein-based block copolymers encompassed the synthesis, structure, assembly, properties, and their main applications [222].

<Insert Figure 32>

Tyrosine-derived polycarbonates are ideal candidates for use in long-term orthopedic implants due to their mechanical properties and good in vitro and in vivo osteocompatibility [101b]. Their high strength and hydrophobic properties preserve their shape during an extended period of time, even in the advanced stages of degradation [101b]. Polymers derived from tyrosine carbonates are applicable in other medical areas, as DDS [223] and scaffolds [224]. Other 'pseudo'-poly (amino acids) widely used as bioabsorbable materials are tyrosine-derived polyarylates. These polymers are a particular type of tyrosine derived polycarbonates, differing only in the length of alkyl ester pendant chains (R) (Error: Reference source not found) [101b,225]. These tyrosine-derived polyarylates have a faster degradation rate than the tyrosine-derived polycarbonates and do not show any cytotoxicity [225].

<Insert Figure 33>

3.7.2. Polysaccharides

Polysaccharides are a class of biopolymers composed of simple sugar monomers, monosaccharides, obtained from different sources: microbial, animal or plant. These are linked by O-glycosidic bonds that can occur on any hydroxyl group of the monosaccharide, originating linear and branched structures. The differences in monosaccharide composition, the shape of the chain and the molecular weight, alter the polysaccharides' properties such as their solubility, gelation time and surface properties. The main advantages of using polysaccharides are [173,226][173,226][174,227][174,227]174,227[174,227]174,227their general low cost and availability and their non-toxicity (with a few exceptions) 174,227. Heparin is an example of a polysaccharide with excellent properties of hemocompatibility.

Polysaccharides have been designed as materials for scaffolds for use in tissue engineering, as well as carriers of drugs in controlled release systems [173].

Hyaluronic acid

Hyaluronic acid (HA) (Error: Reference source not found) is a polymer composed of linear polysaccharides of alternating disaccharide units of β-1,3-N-acetyl-D-glucosamine and α-1,4-D-glucuronic acid connected by β(13) links, belonging to the glycosaminoglycan family [101b,173]. Its molecular weight can reach a few million Da [30]. HA is the biggest macromolecule present in the intercellular matrix of conjunctive tissues such as the vitreous humor, cartilage, umbilical cord blood and synovial fluid [173,227].

HA has good viscoelastic properties and is soluble in water [101b]. This polymer has been used in the composition of implantable devices and for DDS with different administration routes, such as: subcutaneous, nasal, pulmonary, parental via and in modified liposomes [101b,173,227]. For DDS purposes, nanoparticles and hydrogels with HA were developed to deliver antibiotics [228], genetic material [229], analgesics [230] and chemotherapeutics [231]

Upadhyay [232] used a poly(-benzyl-L-glutamate)-b-hyaluronan copolymer to encapsulate DOC, a hydrophobic anticancer drug, by a nanoprecipitation method. The loaded nanoparticles showed high stability both in solution and in its lyophilized form and allowed a sustained release of DOC over 10 days. The in vitro cytotoxicity tests using MCF-7 and U87 cells showed a higher cytotoxicity of the encapsulated DOC than the free DOC. The blood circulation time (t1/2) and the mean residence time was significantly higher for the encapsulated DOC than for the DOC in solution. Additionally, the uptake ratio of the nanoparticles containing DOC at the tumour site was higher than for the DOC solution. The nanoparticles based on the above copolymer demonstrated to be effective in improving DOC chemotherapy.

<Insert Figure 34>

This polymer has been investigated in tissue engineering, including cartilage tissue engineering [233], due to some peculiar characteristics: free radicals scavenging, bacteriostasis, and aiding function in tissue repair. However, the poor mechanical properties are a main problem for this type of application [30,173]. To overcome this drawback, HA has been crosslinked with several compounds, such as benzylic esters [234] or methacrylic anhydride [235] to create injectable hydrogels.

In the ECM, HA can be degraded by free radicals, such as nitric oxide or MMP. An important feature achieved with the crosslinking of HA is the possibility to control its degradation rate, depending on either physical or chemical methods employed [101b].

Sundararaghavan [236] prepared a scaffold based on methacrylated HA combining the electrospinning and photopatterning techniques. The obtained scaffolds had macrochannels that enabled cellular infiltration and vascularization and exhibited a fibrous structure that mimics the ECM. Other HA-based scaffolds were prepared by Ekaputra [237] which contributed beneficially to vascular tissue engineering.

Even though HA has been extensively used for tissue engineering, it does not upkeep cell attachment and spreading, needing chemical modification to support cellular adhesion [238]. The attempt to improve the properties of HA-based gels and the spreading, activation and proliferation rate of cells on their surface has been done recently [239]. Also, the functionalization of HA hydrogels has been performed with deposition of PLL with the HA multilayer films being made on the surface of the hydrogel by the layer-by-layer technique. This modification enhanced the physical and chemical properties of the hydrogel, yielding an improvement on the spreading of the cells on these surfaces and on the cell adhesion [238].

HA hydrogels were synthesized by Lei [240] in order to culture mouse MSC that could be degraded through a combination of cell-released enzymes. These can be used as culture systems that mimic the natural stem cell niche. The authors found that mMSC proliferation occurred in the absence of cell spreading and that mMSCs only spread when RGD and MMP degradation sites were present in the hydrogel scaffold. Cells in stiffer hydrogels exhibited less spreading, migration and slower proliferation rates.

Toh [241] demonstrated a safe and highly-efficient strategy of utilization of non-tumourigenic lineage restricted chondrogenic cell lines derived from human embryonic stem cells (hESCs) to produce ECM-enriched cartilaginous construct when cultured in HA-based hydrogels (Glycosilâ„¢) for cartilage tissue engineering and regeneration.

Photo-cured HA hydrogels loaded with simvastatin, an efficient drug for the induction of osteoblastic differentiation of human adipose-derived stromal cells (hADSCs), were studied. It was demonstrated that they have an important role in bone tissue regeneration as they have proper biocompatibility and show effectiveness towards osteogenesis [242].

HA has also been very investigated as a targeting moiety of the drug conjugates or nanoparticles for cancer therapy as it can specifically bind to various cancer cells that over-express CD44, an HA receptor [243]. Choi [244] studied the PEGylation of HA nanoparticles which resulted in a prolonged blood circulation, a reduced uptake by the liver and improved tumour targetability.

Considerable importance has been given to HA regarding brain tumours, since it is a major component of the brain ECM. A study showed that HA is an important component of implantable scaffolds used to understand the mechanisms of matrix regulation of the tumour and to treat specific central nervous system pathologies [245].

The interesting properties of HA are the reason for the vast portfolio of products available in the market (Table 4).

<Insert Table 4>


Chitosan is a linear natural polysaccharide composed of units of β(14)-2-amido-2-deoxy-D-glucan (glucosamine) and β(14)-acetoamido-2-deoxy-D-glucan (acetyl glucosamine). It is obtained by partial deacetylation of chitin (Error: Reference source not found), a polysaccharide widely found in nature (e.g., cell walls of fungi, exoskeletons of crustaceous and insects) [173,246].

<Insert Figure 35>

The term chitosan is used to describe a series of polymers of different degrees of deacetylation, defined in terms of percentage of primary amino groups on the polymer chain and the average molecular weights. The typical degree of deacetylation of commercial chitosan is usually between 70% and 95% and the molecular weight varies between 10 and 1000 kDa. The biodegradability and biological role of chitosan are dependent on the relative proportions of residues of N-acetyl-glucosamine and D-glucosamine [173,247].

The presence of amino and hydroxyl groups in chitosan that can be modified provide a high chemical versatility. This material can be hydrolysed by certain human enzymes, especially lysozyme, and therefore is considered biodegradable [173].

The crosslinking of chitosan is the common strategy to enhance its mechanical properties. The most commonly used crosslinking agents are dialdehydes such as glyoxal (Error: Reference source not found - (A)) and glutaraldehyde (Error: Reference source not found - (B)). The aldehyde groups form imino bonds with the amine groups of chitosan. The dialdehydes allow crosslinking by direct reaction in aqueous medium under mild conditions. In addition, the crosslinking resulting from the dialdehydes reaction stabilizes the structure of chitosan, preventing digestion by enzymes or bacteria and reducing the antigenicity of the material. Natural derived crosslinking agents, for example, genipin (Error: Reference source not found - (C)) are also used in the crosslinking of chitosan [173,246b].

<Insert Figure 36>

Chitosan is not soluble at physiological pH, being only soluble in acidic medium. In order to increase its solubility at physiological pH, it is necessary to substitute the amine groups of chitosan with hydrophilic groups, usually negatively charged, as methylcarboxy and ethylcarboxy [246b]. However, these components can compromise the biocompatibility of the material [246a].

The largest areas of application of chitosan as a bioabsorbable material are as DDS [248] (e.g., drugs, growth factors), wound healing and tissue engineering. Its success is due to its excellent biocompatibility, low toxicity and immunostimulatory activity, sensitivity to pH, bioadhesive properties, anti-fungal and anti-bacterial action and anti-clotting (hemostatic) properties.

Moreover, the degradation products of chitosan are non-toxic, non-immunogenic and non-carcinogenic [246-247,249].

Chitosan was used in the encapsulation of different active compounds through nanoparticles [12,250], in routes of administration such as ocular, nasal, oral [251], intravenous and mucosal [250b,c].

Gu [252] used a fly-larva shell-derived chitosan sponge to study the possibility of its application as an absorbable, implantable agent for promoting perioperative hemostasis. The great ability to enhance platelet activation, erythrocyte aggregation, morphological alteration and thrombin generation at sites makes chitosan sponges more effective than a gelatin sponge or oxidized cellulose as local hemostatic agents.

Particles based on chitosan can be prepared by different methods using mild conditions, such as ionotropic gelation, polyelectrolyte complex and complex coacervation [12,250a], where the integrity of the drug to be encapsulated is maintained. The encapsulation efficiency and release of protein can be influenced by parameters such as the chitosan molecular weight and deacetylation degree, chitosan concentration, chitosan/polyanion mass ratio, protein concentration and the presence of stabilizers [253].

Results obtained in ophthalmologic applications of chitosan particles as drug carriers show that the ocular surface tissue tolerates very well these particles [254]. This subject was recently reviewed by de la Fuente [255]. Other authors have shown the potential of the chitosan-based carriers on cancer treatment [250c,256]. On this matter, Ta [256b] reviewed different methods of preparing injectable chitosan hydrogels used for the entrapment of anti-cancer drugs.

Hollow chitosan-silica nanospheres were synthesized for breast cancer therapy. In order to produce pH sensitive nanoparticles, chitosan is added to the surface of nanoparticles using the cross-linking reaction with (3-glycidyloxypropyl) trimethoxysilane. This forms nanoparticles with a pH-sensitive polyelectrolyte layer that is conjugated to the antibody molecule, producing the desired nanocarriers for targeted tumour necrosis factor drug delivery in tumour cells [257].

Chitosan/gelatin/glycerol phosphate hydrogels have also been applied as controlled release systems for ferulic acid (FA) (a phenolic antioxidant) in inner nucleus pulposus of the vertebral disc. The injectable thermosensitive hydrogels showed to be an appropriate slow release system for FA delivery in the nucleus pulposus [258]. Brunel [259] prepared nanohydrogels based on chitosan in the absence of any toxic solvent or crosslinker and the adsorption of different biomolecules onto the nanohydrogel surface was studied. Taking into account the results obtained, the authors believe that the gels are versatile carriers that can find broad application in the biomedical field.

Highly porous 3D chitosan hydrogels were developed for use in tissue regeneration. The porosity has a very important role, as its presence improves the swelling ratio of chitosan hydrogels and enhances the cell proliferation rate [260].

Also, chitosan has been used to synthesize nerve conduits for peripheral nerve regeneration. The PDLLA/chondroitin sulfate/chitosan based conduits prepared by Xu [261] are effectively used for peripheral nerve damage repair due to their proper mechanical strength, good nerve regeneration ability and biodegradability.

Chitosan has also been used as a biomaterial substrate for the maintenance of self-renewal for adult hMSCs, since stem cells can lose their primitive properties during the in vitro culture [262].

The cationic nature of chitosan makes it a prospective material to be used in gene therapy as it easily binds to nucleic acids, forming polymer-nucleic acid complexes [30,250b,263]. Chitosan's molecular weight, its degree of deacetylation, the chitosan/nucleic acid ratio and the type of cell transfected influences the transfection efficiency [263-264]. These complexes are able to protect the genetic material from nuclease attack, maintaining their bioactivity.

To effectively deliver siRNA to cancer cells, Veiseh [265] developed a cancer-cell specific magnetic nanovector. The authors used a superparamagnetic iron oxide nanoparticle core coated with PEG-grafted chitosan and PEI. To induce tumour specificity and potency, the aggregate was functionalized with siRNA and a tumour-targeting peptide, chlorotoxin. The results showed that this nanovector system has a high potential to be used for treatment of malignant tumours because of it specificity and potency as well as it could enhance MRI contrast in vitro, potentially enabling monitoring of treatment in vivo through MRI.

The influence of the chitosan chain architecture in the transfection efficiency of DNA in HeLa cells was studied by Malmo [266]. Chitosan complexes, namely self-branched chitosan, self-branched trisaccharide-substituted chitosan and linear chitosan were studied showing that self-branched chitosan gave very low toxicity, being the most efficient in transfecting HeLa cells. In turn, the self-branched chitosan containing trisaccharides showed low transfection efficiency but a high intracellular uptake. Interesting to note that with linear chitosan no transfection in the HeLa cellular line was observed.

Opanosopit [267] demonstrated that mixing chitosan with poly(L-arginine)-pDNA complex enhances the transfection efficiency of such complex in HeLa cells. The controlled modification of chitosan using stimuli-responsive polymers prepared by controlled/ "living" radical polymerization methods have been focus of particular attention due to the enormous potential of this approach to fine tune the properties of chitosan. Recently, Carreira [268] highlighted the most relevant achievements on the modification of chitosan using temperature and pH responsive polymers.

Some commercial products based on chitosan materials are HemCom® bandages from HemoCon® Medical Technologies Inc., consisting of an adhesive film to stop the bleeding and the geniaBeads® CN of Genialab® GmbH, which is a hydrogel in the form of granules with various biomedical applications.


Dextran (Error: Reference source not found) is a biocompatible branched polymer consisting of links of D-glucose residues by α(16) bonds with some degree of branching using links of α(13) bonds. It is produced by different bacteria from sucrose through the action of the enzyme dextransucrase. Dextran is available in a wide range of molecular weights, along with several derivatives, being biodegraded predominantly by enzymatic pathways. Dextran is very stable, since the glycosidic bonds are hydrolysed only under strongly acidic or alkaline medium [173,269].