review polymers

Polymer International Polym Int 56:145–157 (2007)

ReviewBiodegradable polymers applied in tissueengineering research: a reviewMonique Martina1 and Dietmar W Hutmacher2∗1Department of Orthopedic Surgery, Yong Loo Lin School of Medicine, National University of Singapore, 10 Kent Ridge Crescent,Singapore 1192602Division of Bioengineering, Faculty of Engineering, Department of Orthopaedic Surgery, Yong Loo Lin School of Medicine,National University of Singapore, 10 Kent Ridge Crescent, Singapore 119260

Abstract: Typical applications and research areas of polymeric biomaterials include tissue replacement, tissueaugmentation, tissue support, and drug delivery. In many cases the body needs only the temporary presence ofa device/biomaterial, in which instance biodegradable and certain partially biodegradable polymeric materialsare better alternatives than biostable ones. Recent treatment concepts based on scaffold-based tissue engineeringprinciples differ from standard tissue replacement and drug therapies as the engineered tissue aims not only torepair but also regenerate the target tissue. Cells have been cultured outside the body for many years; however, ithas only recently become possible for scientists and engineers to grow complex three-dimensional tissue grafts tomeet clinical needs. New generations of scaffolds based on synthetic and natural polymers are being developed andevaluated at a rapid pace, aimed at mimicking the structural characteristics of natural extracellular matrix. Thisreview focuses on scaffolds made of more recently developed synthetic polymers for tissue engineering applications.Currently, the design and fabrication of biodegradable synthetic scaffolds is driven by four material categories:(i) common clinically established polymers, including polyglycolide, polylactides, polycaprolactone; (ii) noveldi- and tri-block polymers; (iii) newly synthesized or studied polymeric biomaterials, such as polyorthoester,polyanhydrides, polyhydroxyalkanoate, polypyrroles, poly(ether ester amide)s, elastic shape-memory polymers;and (iv) biomimetic materials, supramolecular polymers formed by self-assembly, and matrices presentingdistinctive or a variety of biochemical cues. This paper aims to review the latest developments from a scaffoldmaterial perspective, mainly pertaining to categories (ii) and (iii) listed above. 2006 Society of Chemical Industry

Keywords: scaffolds; biodegradable polymers; tissue engineering; matrices

INTRODUCTIONA great number of current tissue engineering strate-gies are based on the development of a cell–scaffoldconstruct whose role is to repair and regenerate tissuedefects (Fig. 1). During the first phase of tissue engi-neering in the 1990s research utilized either US Foodand Drug Administration (FDA)/CE mark approveddevices or used so-called conventional scaffold fabrica-tion technologies in combination with FDA/CE markapproved biomaterials of synthetic and natural origin.This work has been reviewed in detail elsewhere.1–3

Currently, the design and fabrication of syn-thetic scaffolds is driven by four material categories:(i) biodegradable and bioresorbable polymers, whichhave been effectively used for clinically establishedproducts, including polyglycolide (PGA), polylactides(PLA), poly-L-lactic acid (PLLA), poly-D,L-lacticacid (PDLA), polycaprolactone (PCL); (ii) novel di-and tri-block polymers which predominantly incor-porate PGA, PLA, and other resorbable polymersin different chain arrangements which confer both

degradation and mechanical property customization;(iii) polymers that are regulatory approved for spe-cific applications and/or are in clinical trials, e.g.polyorthoester (POE), polyanhydrides, polyhydrox-yalkanoate (PHA), and newly synthesized polymericbiomaterials, such as polypyrroles (PPy), poly(etherester amide)s (PEEA), and elastic shape-memory poly-mers; and (iv) biomimetic materials, supramolecularpolymers formed by self-assembly, and matrices pre-senting distinctive or a variety of biochemical cues.The aim of this review is to capture the latest develop-ments from a scaffold material point of view coveringthe aforementioned categories (ii) and (iii).

DI- AND TRI-BLOCK POLYMERS BASED ONALIPHATIC POLYESTERSIn attempts to tune aliphatic polymer propertiesto wider applications for scaffold-based tissue engi-neering, di- and tri-block polymers are seen aspromising alternatives. By modifying the backbone

∗ Correspondence to: Dietmar W Hutmacher, Division of Bioengineering, Faculty of Engineering, Department of Orthopaedic Surgery, Yong Loo Lin School ofMedicine, National University of Singapore, 10 Kent Ridge Crescent, Singapore 119260E-mail: [email protected](Received 3 October 2005; revised version received 28 February 2006; accepted 21 March 2006)Published online 13 October 2006; DOI: 10.1002/pi.2108

2006 Society of Chemical Industry. Polym Int 0959–8103/2006/$30.00

M Martina, DW Hutmacher

Figure 1. Schematic of scaffold-based tissue engineering.

of the polymers, some of the characteristics, suchas degradation properties, mechanical properties, andeven biocompatibility, of the polymers can be changedto a significant extent.4–6

Li et al.4 developed a series of di- and tri-blockpolymers based on PCL, PLA, and poly(ethylene gly-col) (PEG)/poly(ethylene oxide) (PEO). They used atri-block made of PLA/PEO/PLA, with PEO as thehydrophilic and swollen component; and PLA chainsas the nanometric nodes in the gel network. Thesephysically crosslinked hydrogels possessed interestingdegradation properties. Incorporation of PLA blocksat the end of the PEO segments decreased the degra-dation rate when compared to pure PEO. Weightlosses of 49% after 10 days and 77% after 30 dayswere observed. A fast initial loss was observed due tothe release of PEO-rich segments. Enzymatic degra-dation promoted a much faster degradation rate whichreached 59% weight loss after 80 h of immersion, com-pared to 5% under hydrolytic degradation conditions.The author of this review and co-workers prepared aseries of scaffolds from these materials using a rapidprototyping system (Fig. 2) and the scaffolds werestudied in vitro and in vivo.6

PCL/PGA di-block systems5 have certain elasticproperties. Scaffolds with pore sizes of 250 ± 50 µmand a porosity of 93% were produced using a sol-vent casting and particulate leaching method. Due

to its elasticity (elongation up to 250% and recoveryup to 98% after applied strain of 120%) Lee et al.5

predicted the potential of this di-block in muscle tis-sue engineering. Preliminary in vitro studies using ratsmooth muscle cells (SMC) displayed growth and tis-sue formation on the PCL/PGA scaffold. A summaryof the properties of this system is given in Table 1.

PCL-based copolymers for soft tissue engineeringhave been investigated by Cohn’s group,7 whosynthesized tri-blocks based on PCL/PEO/PCL (softsegment with hexamethylene diisocyanate chainextension as the hard segment). Incorporation of PEOenhanced water permeability. These tri-blocks havea water uptake of up to 94%, which also influencesthe mechanical properties and biodegradability. Thelength of the blocks affect polymer properties: forexample, longer PEO segments lead to a decreasein the degree of crystallinity, an increase in wateruptake, and an increase in degradation rate. Possibleapplications include soft tissue replacement and drugdelivery systems. The cell biocompatibility of thissystem was studied by the authors of this review(Fig. 3(a)–(c)).

One of the current challenges facing polymerusage in cell therapy is the ability to add bioac-tive molecules to enhance cytocompatibility. Con-ventional biodegradable polymers lack this charac-teristic. A strategy to overcome this limitation is

146 Polym Int 56:145–157 (2007)DOI: 10.1002/pi

Biodegradable polymers in tissue engineering

Tab

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Polym Int 56:145–157 (2007) 147DOI: 10.1002/pi


Tab

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148 Polym Int 56:145–157 (2007)DOI: 10.1002/pi


Figure 2. Polyester–polyether block copolymers composed of PCL or PLA and poly(ethylene glycol) (PEG) have attracted much attention as theyoffer the possibility of varying the ratio of hydrophobic/hydrophilic constituents to modulate the degradability and hydrophilicity of the polymermatrix and surface. Scaffolds were fabricated using a rapid prototyping (RP) machine built in-house at the National University of Singapore. Systemdetails are reported elsewhere.7 Briefly, in contrast to traditional RP systems such as fused deposition modelling, three-dimensional printing,stereolithography, and selective laser sintering, which mainly focus on a single mode of material processing, this system can accommodate a muchlarger variety of synthetic and/or natural biomaterials. A lay-down pattern of 0/90 was used to form the honeycomb patterns of square with a singlefill gap (FG) of 1 mm. Porous sheets measuring 40 × 40 × 3 mm were fabricated using the system, with a 0.5 mm diameter nozzle. PCL–PEGdiblock copolymer was synthesised by Li et al.4 The extrusion temperature was set at 110 ◦C for PCL–PEG. PCL–PEG was dispensed with an airpressure of 3.5 bar and a speed of 30 mm min−1 in the x/y axis. Scanning electron micrographs (right) of melt extruded PCL–PEG scaffolds displaytypical honeycomb morphology (A). Alternate layers of filaments were positioned at right angles to one another creating pore sizes of 600 µm.Confocal laser microscopy of scaffold/cell constructs was used to study cell attachment and proliferation.

to use polymers containing functional side groups.Guan and co-workers8 attempted to synthesize anABA tri-block polymer that had this particularproperty. Fabrication from ABA-type tri-block poly-mer PLGBG–PEG–PLGBG consisting of PEG andpoly[(lactic acid)-co-(glycolic acid)-alt-(γ -benzyl-L-glutamic acid)] that had undergone catalytic hydro-genation resulted in PLGG–PEG–PLGG, whichcontained carboxyl pendant groups. These carboxylpendant groups may be useful to facilitate the attach-ment of oligosaccharides, drug molecules, or shortpeptides. Due to its amphiphilic nature, the polymerformed micelles and may find useful applications indrug delivery systems.

Another new tri-block system are thermogellingcopolymers9 made of PEG/PCL/PEG. At a copolymerconcentration of less than 1% and at a temperatureof 20 ◦C, micelle aggregates of 10 and 23 nm coex-isted, but at 25–45 ◦C micelle aggregates of 23 nmwere dominant. Increasing the copolymer concentra-tion also increased the micelle aggregation size. Uponvarying the PCL block’s length, sol–gel temperaturedecreased due to the hydrophobic interaction thatdrove gel formation. This system is appropriate forin situ gel forming whereby entrapment and depotformation can be envisaged with minimally invasivetherapy.

NEWLY SYNTHESIZED OR STUDIEDPOLYMERIC BIOMATERIALSPolycarbonatePolycarbonate in its pure form is an amorphouspolymer that possesses low moisture absorption,and is not susceptible to microbial attack, whichimplies non-biodegradability of the polymer. How-ever, the mechanical properties of the polymer attractresearchers keen to develop polycarbonate-basedmaterial that can withstand mechanical loading whiletissue is being regenerated. Since resorbability is themain problem, Bourke et al.10 investigated a member

of the tyrosine-derived polycarbonates that was notonly resorbable, but also possessed high strength. Thestructure of this material is shown in Fig. 4(a). How-ever, in that study, the degradation rate was shown tobe very slow, with no mass loss observed after 30 weeksof incubation in PBS, at 37 ◦C.

In the first phase, Bourke et al. fabricated this mate-rial by melt extrusion at 60–90 ◦C, which was abovethe glass transition temperature. In the second phase,they tried the melt-spinning technique at 181–183 ◦C.The fabricated fibres were aligned to mimic thestructure of ACL (anterior cruciate ligament). Theultimate tensile strength of this material was foundto be comparable with natural ACL (57 MPa). Sub-cutaneous implantation using a rat model revealedtissue ingrowth 4 weeks post-operative with mainlyfibroblast-like structures observed surrounding thefibres. Interestingly, although the structure was intact8 weeks post-implantation, strength retention droppedto 40%; this was in contrast to in vitro studies whichshowed almost 90% strength retention, which impliedthat the in vivo environment induced molecular relax-ation, causing reduction in strength.

Further studies based on polycarbonate’s goodbiocompatibility and ease of biochemical modificationtowards cell adhesivity11 have been undertaken.Using the polycondensation technique, copolymers ofpoly ethylglycol/poly(desamino tyrosyl-tyrosine ethylester carbonate) (PEG/poly(DTE carbonate)) wereobtained to study keratinocyte migration. It wasconcluded that PEG provided a better attachment andmigration substrate for fibroblasts and keratinocytes,whereas poly(DTE carbonate) acted to maintain thestructural integrity of the graft.

PolyphosphazeneThe search for polymeric materials that are versatilefor various hard and soft tissue engineering appli-cations has generated interest in another group ofpolyphosphazene-based materials.12–14 The uniqueproperties arise from the unusual flexible backboneallowing torsional and angular freedom within the

Polym Int 56:145–157 (2007) 149DOI: 10.1002/pi


(a)

(b)

(c)

Figure 3. (a) Soft-PCL scaffold fabricated via a rapid prototyping technique (unpublished data). (i) High flexibility of scaffold is demonstrated; (ii, iii)swelling ability is demonstrated (soft segments in polymer chain allow scaffold to swell up to 1.5 times). (b) Light micrographs of soft-PCL(P) alginate/thrombin (M) construct seeded with human adipose derived precursor cells (hADAS). Formulations of cell aggregates and extracellularmaterial (white arrows) illustrate cell proliferation upon culturing: (i) day 1, (ii) day 7, (iii) day 22, (iv) day 42. Bar represents 100 µm. (c) Environmentalscanning electron micrograph (left) of scaffold/cell construct after 3 weeks of culturing exhibited intact cell/alginate/thrombin structure (M) inside thesoft-PCL pore morphology. Confocal laser micrograph (right) of soft-PCL (P) alginate/thrombin construct after 3 weeks of culturing. Cell viability isshown by life/death assay (FDA; green fluorescence in confocal laser micrograph) inside the alginate/thrombin matrix. (Cells are indicated by whitearrows).

P–N skeletal system (Fig. 5(a)). The most com-mon fabrication technique is thermal ring-openingpolymerization of dichlorophosphazene. The current

technology enables fabrication of this polymer derivedfrom polydichlorophosphazene using two differentsubstituents. These substituents induce highly tuned

150 Polym Int 56:145–157 (2007)DOI: 10.1002/pi


(a) (b)

(c)

(d)

Figure 4. Structures of (a) poly(desamino tyrosyl-tyrosine ethyl ester carbonate) (poly(DTE carbonate)), modified from of Bourke et al.;10

(b) PPF/PPF-DA crosslinked, modified from of Horch et al.;25 (c) poly DTH-adipate for R = hexyl and y = 4, modified from of Schachter and Kohn;42

(d) soft segment and hard segment of poly(ether ester amide), modified from Deschamps et al.44

(a)

(b)

(c)

Figure 5. Preparation schemes of (a) polydichlorophosphazene, modified from Ambrosio et al.;12 (b) PEUU, modified from Stankus et al.;24

(c) poly(glycerol sebacate), modified from Wang et al.33

Polym Int 56:145–157 (2007) 151DOI: 10.1002/pi


properties such as crystallinity, degradability, andhydrophilicity/hydrophobicity which provide the ver-satility of the polymers.

Ambrosio et al.12 explored amino acid esterpolyphosphazene in bone tissue engineering. Theychose poly[(ethylglycinate phosphazene)-co-(p-me-thylphenoxy phosphazene)] (PPHOS) in combinationwith hydroxyapatite (HA) to develop a composite.Degradation was achieved by incorporating side chainsthat sensitized the polymer backbone to hydrolysis;in this case, the amino acid side chains. Phenoxyside chains inhibited degradation, and in this study,PPHOS/HA remained stable mechanically and main-tained a compressive modulus of around 200 MPa,after 12 weeks’ degradation in vitro. PLAGA/HA con-structs that were used as a control showed almostinsignificant compressive modulus. Degradation prod-ucts could easily be detoxified by the body as thedegradation by-products are amino acids, phosphates,and ammonia. PPHOS possesses a combination ofsurface- and bulk-eroding properties.

Other types of polyphosphazene, those withpolyethyloxybenzoate and polypropyloxybenzoate,were combined with HA resulting in compositessuitable for bone grafts.15 Syntheses of these polyphos-phazene/HA composites were achieved throughacid–base reactions to form HA precursors, whichincreased the pH of the solution and in the presenceof polyphosphazenes resulted in carboxyl formation inthe surface layer. These carboxyl groups then reactedwith calcium ions to form calcium crosslinks on thesurface, which played a role in nucleation and deposi-tion of HA.

In another study, a PPHOS/PLAGA blend exhibiteda higher pH in the degradation solution compared toPLAGA.13 The buffering phenomenon was attributedto phosphates which were shown to be present in thesolution using 31P NMR. Cell adhesion experimentsusing MC3T3-E1 cells revealed PPHOS/HA topossess a similar cell attachment and proliferationwhen compared to tissue culture plastic (TCPS).

Nair and colleagues14 investigated the effect of sol-vent, needle diameter, solution concentration, andapplied voltage on the fabrication of polyphosphazenenanofibre scaffolds. Using chloroform as the solventproduced the most uniform fibres. Decreasing the nee-dle diameter resulted in decreased fibre diameter, buttoo fine a needle produced fibres with beads. Increas-ing the solution concentration generally resulted inlarger fibre diameters. The applied voltage affected theshape of the nanofibres generated: 27–30 kV resultedin a curly or distorted shape, 33 kV resulted in cylindri-cal and rod-like structures, while 36 kV gave even finerand more distinct rod-like structures. Investigationof the polyphosphazene nanofibres extended to cellcompatibility studies. In vitro cell adhesion of bovinecoronary artery endothelial cells showed adhesion 24 hpost-seeding. The materila also exhibited cytocom-patibility to osteoblast cell lines, MC3T3-E1, whichadhered and proliferated for 7 days. This presents an

opportunity for this material to be utilized in appli-cations such as coatings to enhance tissue integrationand wound dressings.

Trimethylene carbonate-based materialsThe elastomeric properties of poly(trimethylene car-bonate) (poly(TMC)) make it a potential can-didate for scaffold-based soft tissue engineeringapplications.16–19 Interestingly, this polymer exhibitsslow degradation in vitro, but rapid degradationin vivo.16 Poly(TMC) showed negligible mass lossafter 2 years in physiological solution, in which degra-dation mainly occurred due to hydrolysis of esterbonds. In contrast, in vivo implantation of polymerdiscs into Wistar rats resulted in a high degradationrate based on a cell-triggered surface erosion mecha-nism. After 52 weeks, implanted poly(TMC) was 1%of its initial mass.

Incorporation of other monomers such as D,L-lactide acid (DLLA) or ε-caprolactone (CL) mod-ulated the degradation properties in vitro as well asin vivo. In vivo degradation characteristics of thesecopolymers coincided well with in vitro degradationbehaviour, in which hydrolysis was mainly responsi-ble for both conditions. TMC/DLLA decreased to1% of the initial mass after 52 weeks, with rapid lossobserved after week 12, while TMC/CL displayed alinear and continuous reduction throughout a year ofinvestigation, with mass loss less than 7% of the initialmass. The difference in degradation behaviour of thetwo copolymers was caused by autocatalytic activity ofDLLA, which did not occur in TMC/CL.

Furthermore, the in vivo tissue response of thesecopolymers is similar to sterile inflammatory reactionsfollowed by normal foreign body reactions, which arecommonly seen after implantation of biodegradablepolymers.

Due to its flexibility, poly(TMC/DLLA) was tai-lored towards heart tissue engineering applications.17

Poly(TMC/DLLA), although glassy at room temper-ature, was rubbery at body temperature which made itdifficult to achieve a porous structure. A combinationof coprecipitation, compression moulding, and saltleaching techniques were used to produce a porousstable scaffold with a pore size of around 100 µm, aporosity of 85–95%, and a compressive modulus upto 430 kPa. In vitro degradation studies in physiolog-ical solution showed that porous scaffolds shrunk byup to 65% after the third week. Preliminary investi-gation of cell compatibility using rat cardiomyocytesshowed average cell attachment and proliferation onthe copolymer surface compared to normal TCPSplates.

Poly(TMC/CL) was explored for nerve grafting.16

Human dermal fibroblasts (HDFs) were subjectedto extracts of the polymer. Results showed that HDFssubjected to the extraction vehicle maintained high cellviability and cell metabolism activity. Schwann cellswere plated onto copolymer discs and demonstratedgood proliferation, with a higher proliferation rate after

152 Polym Int 56:145–157 (2007)DOI: 10.1002/pi


6 days compared with those cells plated onto PLGA.Implanted tubular poly(TMC/CL) scaffolds were stillsmooth and flexible at 60 days post-implantation, andadhered well to the surrounding tissue. In anotherin vivo study, a lesion in the spinal cord of femaleWistar rats exhibited extensive growth of Schwanncells and extracellular matrix after implantation ofpoly(TMC/CL) to the lesion.20

Grijpma et al.19 also explored the possibility of aphotocrosslinkable polymer via UV functionalizationof poly(TMC/CL) or poly(TMC/DLLA) with fumaricacid monoethyl ester (FAME).

PolyurethanesPolyurethanes are another area of investigation,especially for soft tissue engineering applications,in contrast to aliphatic linear polyesters that arebetter suited to hard tissue engineering due to theirhigh glass transition temperature and high modulus.Polyurethanes exhibit a wide range of propertiesthrough variability of the hard segment (diisocyanate),the soft segment (polyethers or polyesters), the chainextenders, and the ratios in which they are reacted.Previously, polyurethanes had a limited usage dueto the toxicity of their degradation product (2,4-diaminotoluene). Hence, the challenge presented wasto develop polyurethanes with non-toxic degradationproducts.

Non-toxic polyurethanes with diisocyanate replace-ments which could give rise to non-toxic degrada-tion products have been developed by at least twogroups.21–24 Guan et al.23 synthesized the polymersmade from PCL and 1,4-diisocyanatobutane (BDI)with putrescine as chain extender (poly(etherurethaneurea), PEUU) (Fig. 5(b)). BDI was used because,upon degradation, it would release putrescine, apolyamine that is essential for cell growth and pro-liferation. Zhang et al.22 synthesized polyurethanefrom highly pure lysine diisocyanate (hard segment)and polymerized it with glucose, which resulted inmajor degradation products lysine and glucose (LDI-glucose).

Scaffold fabrication routes for these polyurethanesinclude thermally induced phase separation, electro-spinning, and water foaming. These create differ-ent porosities, surface-to-volume ratios, and three-dimensional structures, with concomitant changes inmechanical properties which are wide-ranging and canbe varied to suit potential applications in the biomed-ical field, such as engineering blood vessels and bone(Table 1).

The degradation mechanisms of the polymersare important and need to be investigated further.Non-toxic degradation products are necessary and,moreover, mechanical properties are also influencedby degradation mechanisms. LDI-glucose polymer,for example, is degraded by hydrolysis of urethanebonds to liberate lysine, glucose, ethanol, and CO2.Ethanol could inhibit cell–cell adhesion, but a studyreported that concentrations less than 30 mM (0.5%

v/v) are harmless to the cell.21 Moreover, in contrastto PLA and PLGA degradation mechanisms, thestudy showed no significant increase in pH of thesolution. PEUU degradation products were alsoshown to be non-toxic to endothelial cells. Thepolymer showed a linear degradation with no signsof autocatalytic effects when compared to PLAor PLGA degradation behaviour.23 Scaffolds wereprepared by thermally induced phase separation and asubsequent solvent extraction technique. The resultswere porous structures (porosity >80%) with porediameters ranging from 12 to 232 µm. The pore shapewas dependent upon the quenching temperature andpolymer concentration. The tensile strength of thesepolymers during degradation decreased by about 12%after 1 week and 31% after 2 weeks from the initialvalue of 0.57–1.69 MPa.

Stankus et al.24 were able to machine thesepolyurethanes via electrospinning to produce scaffolds.The tensile strength of the nanofibres produced rangedfrom 2 to 13 MPa and breaking strains from 160 to280%. Incorporation of type I bovine collagen reducedthe tensile strength, but also reduced water contactangles. This represents an increase in hydrophilicityand, in turn, enhanced attachment of rat smoothmuscle cells.

In vitro and in vivo observations demonstrated thatthese polymers had no harmful effects on cell viability,growth, and proliferation. Subcutaneous implantationusing rat models revealed that LDI-glucose polymerdid not enhance capsule formation, accumulation offoreign body giant cells, or tissue necrosis.22

The versatility of these polyurethanes was demon-strated through attempts to enhance cell and tissuecompatibility via: addition of soft segments with PEGto the backbone in order to enhance hydrophilicity;mixing with collagen type I for better cell attachment;and addition of ascorbic acid to the polymer mixtureto enhance osteoblast lineage progression. All of theseproved to be feasible. Polyurethane-based materialsare consequently considered promising candidates forbiomaterial-based applications.

PolyfumaratePolyfumarate-based materials have been developedmainly for bone tissue engineering.25–28 The mainadvantages of these materials are their injectableand in situ crosslinkable properties.29,30 WithN-vinylpyrrolidone (N-VP) as crosslinker, Payneet al.28 showed during an in vitro study that fullycrosslinked poly(propylene fumarate) (PPF) hadpotential as a substrate for supporting rat osteoblasts.Cell proliferation, ALP activity, and osteocalcin andcalcium production of cells on fully crosslinked PPFdid not exhibit significant differences with thosecells grown on TCPS. Poly(caprolactone fumarate)(PCLF) and poly(ethylene glycol fumarate) (PEGF)were also investigated as injectable, self-crosslinkablepolymers which circumvented the requirement forcrosslinking agents that may be toxic. The polymers

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were shown to harden and self-crosslink when underphysiological conditions, and tissue compatibilitystudies using rat models demonstrated no inflamma-tory reactions.29,30

Bone and soft tissue responses have been inves-tigated in New Zealand white rabbit models.27 APEG block was added in order to make a hydrogel(oligo(PEG fumarate), OPF). Tissue response andin vitro degradation behaviour of four groups wereinvestigated. The groups included longer PEG blocks,higher crosslinking density, and the addition of cell-adhesive peptides to the hydrogels. Hydrogels withhigher crosslinking were observed to have the leastmass loss after 12 weeks in vitro. As the degradationproduct was fumaric acid, a drop in pH was expected.A pH drop of not more than 0.5 was reported (whichwas considered to be harmless to the cells). In contrast,in vivo results suggested that hydrogels with longerPEG block resulted in more tissue infiltration, andhence more active degradation. Fibrous capsule for-mation with 5–15 cell layers including inflammatorycells was observed for all groups.

Degradation rate and degree of swelling were alsodependent on the degree of crosslinking per macromerchain. Less crosslinking gave rise to an increase inwater uptake and increased swelling. As degradationoccurred mainly due to the cleavage of ester bondswithin the crosslinked network by hydrolysis, a greaternumber of chains were accessible to water and sothe polymer degraded faster. The higher the wateruptake, the faster the degradation. One type of OPFinvestigated even appeared soft and jelly-like after12 weeks under physiological conditions.

As the applications of polyfumarate-based materialswere first investigated for bone substitutes, mechanicalproperties are an ultimate concern. Cortical bone hasa compressive modulus in the range 17–20 GPa, com-pressive strength of 106–144 MPa, flexural modulusof 15.5 GPa, and flexural strength of about 180 MPa.This presents a major challenge for the tissue engineerusing polyfumarate-based materials. This unique char-acteristic of bone is a result of its composite make-upcomprising the interaction of inorganic material, i.e.HA, with organic material such as collagen fibres.Researchers have considered ways to optimize themechanical properties of poyfumarate-based polymersusing the principle of bone architecture.25,26

Horch et al.25 developed PPF/poly(propylenefumarate diacrylate) (PPF/PPF-DA) (Fig. 4(b)) witha surface modification using carboxylate alumoxanenanoparticles. The polymer composites were gener-ated by mixing in a chloroform solvent, and the solventwas then removed by rotary evaporation and high-vacuum drying. Crosslinking was achieved by UVradiation. Interactions of the inorganic and organicmatrix were achieved by covalent bonding, and thepresence of organophilic chains in the organic matrixenhanced dispersion. Covalent bonding alone was notenough; dispersion played a key role in the mechanical

properties of the composite. Without proper disper-sion, the inorganic particles tended to aggregate andcreate crack propagation sites, which, in turn, madecompressive fracture and flexural fracture strengthworse than that of PPF/PPF-DA.

Flexural testing was done on samples fabricatedby injecting a nanocomposite mixture into a mould,and crosslinking the mixture using UV radiation.The samples that showed enhanced dispersion andcovalent bonding with PPF/PPF-DA matrix exhibitedthe best flexural modulus of more than threefoldhigher (5.4 GPa) compared with blank PPF/PPF-DA,which has a flexural strength of 1.5 GPa. Compressivestrength, however, was not significantly affected bythese modifications.

Incorporation of β-tricalcium phosphate (β-TCP)was also attempted to enhance the mechanicalproperties of polyfumarate-based polymers.26 Thepolymer was produced by radical polymerization usingbenzoyl peroxide and dimethyltoluidine as initiatorand accelerator. N-VP was used as crosslinkingreagent. Scaffolds were then fabricated by mixingPPF with β-TCP and leachable porogen NaCl.Groups with β-TCP concentrations of 0.5 g g−1 ofPPF exhibited bending strengths of up to 16 MPa,compressive strengths of up to 79 MPa, moduliin bending of up to 1270 MPa, and moduli incompression of up to 1020 MPa. An approximatelytwofold increase in bending and compressive strengthand bending and compression modulus for a twofoldincrease in β-TCP concentration was observed.Mechanical testing showed compression and bendingmodulus of elasticity to be of the same order ofmagnitude as that of trabecular bone.

PolyorthoesterPolyorthoester is a hydrophobic polymer fabricated bypolycondensation of diketene acetals and diols. Thisfabrication creates ortho-ester bonds that are stable atneutral pH, but hydrolyse rapidly at phagosomal pH,i.e. pH 5.5. Today, it is mainly used in drug deliverysystems.31,32

Poly(glycerol sebacate)The search for soft and mechanically stable elastomericmaterials to be implanted in dynamic environmentsled to the investigation of polymers that are anal-ogous with vulcanized rubber, having a crosslinkedthree-dimensional network in combination with ran-dom coil characteristics. Polycondensation of glyceroland sebacic acid renders a polymer having hydro-gen bonding interactions through hydroxyl prolinehydroxyl groups.33,34 With building blocks made ofglycerol (a basic building block of lipids) and sebacicacid (a natural metabolic intermediate in ω-oxidationof medium- to long-chain fatty acids), such materi-als are expected to be biocompatible and non-toxic(Fig. 5(c)).

The hydrophilic characteristics of the material area result of the hydroxyl groups attached to its

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backbone. The material is insoluble but swells inwater by approximately 2%. It is totally amorphousat 37 ◦C like vulcanized rubber, a thermoset polymer.However, the uncrosslinked polymer can be melted toa liquid form which is soluble in common organicsolvents. It is tougher than hydrogels with tensilestrength of less than 0.5 MPa and tensile strainmore than 300%.33 In vivo and in vitro investigationsshowed an acceptable biocompatibility. In vitro studiesusing NIH 3T3 fibroblast cell lines grown in PGS-coated Petri dishes exhibited higher cell growthcompared with PLGA-coated Petri dishes. An in vivostudy with Sprague-Dawley rats showed less fibrouscapsule formation compared to PLGA control aftersubcutaneous implanation.

While simple foams were fabricated using a saltleaching technique, lithography was used to fabricatecapillary networks.34 The surface of the wafer-likescaffold was coated by a pentapeptide derived fromfibronectin to improve cell attachment. Adherenceof the HUVEC to the PGS was observed andproliferation occurred for at least 10 days. After14 days, the surface of the network was nearlyconfluent and cultures were kept for up to 4 weeks.

PGS has another potential application as nerveguides.35 In vitro studies demonstrated PGS asnon-toxic to Schwann cells and supported betterproliferation compared to PLLA. In vivo studiesrevealed less inflammatory response possibly due toits degradation profile via surface erosion.

Elastic shape-memory polymersLendlein and Langer36,37 investigated polymers thatcould change shape by increases in temperature. Thisthermally induced shape-memory effect required twocomponents: a switching segment, having transitiontemperature to fix temporary shape, and a segment forcrosslinking to determine the permanent shape. Theswitching segment used was oligo(ε-caprolactone)diolor dimethacrylates, while crosslinking segments wereeither n-butylacrylate or oligo(p-dioxanone)diol.

The mechanism of changing the shape frompermanent to temporary relies on the transitiontemperature of the material. A polymer in a permanentshape is heated above Ttrans while applying an externalstress. The temporary shape is obtained by reducingthe temperature to below Ttrans. Releasing the externalstress and heating up the material will reform thetemporary shape back to the permanent shape.

Using this mechanism, prior to surgery, implantscan be compressed to a smaller and more compacttemporary shape, inserted by minimally invasivesurgery, and then using body heat, it will expand backto the permanent shape. Another possible applicationis as sutures in endoscopic surgery, whereby the sutureknot can be applied loosely in its temporary shape,followed by an increase in the temperature, whichwould tighten the knot as it goes back to its permanentshape.

Another development by Lendlein and co-workers38

as regards elastic shape-memory polymers was light-induced shape-memory polymers. With a similarmechanism to that of a heat-sensitive shape-memorypolymer, this too consisted of two segments: molecularswitches that were photoresponsive, and netpoints forcovalent crosslinking that determined the permanentshape, using cinnamic acid as a molecular switch.When the polymer was stretched, the coiled segmentsof amorphous polymer chain were elongated. Uponexposure to UV radiation of wavelength >260 nm,the elongated segments were partially fixed due to theformation of new photoresponsive crosslinks. If theloading was removed and the polymer exposed to UVradiation of wavelength <260 nm, the crosslinks werecleaved and the polymer returned to its original shape.The discovery of this material eliminated temperatureconstraints associated with thermally induced shape-memory polymers for medical applications.

PolypyrroleElectronic interaction is one of the factors that mayinfluence neuronal tissue regeneration and growth.A class of polymers that have also been explored inthe field of nerve tissue engineering are conductingpolymers. Polypyrrole (PPy) is an electrodepositedpolymer that can be doped to modify its physical,chemical, and electrical properties, and has conse-quently emerged as a potential candidate for scaffoldsin neuronal tissue regeneration.39 Using PPy dopedwith polystyrene-sulfonate (PSS) and sodium dode-cylbenzene sulfonate (NaDBS) at different depositiontemperatures and solvents, George et al.39 investi-gated the influence of these parameters on neuraltissue growth in vivo. Immunofluorescence analysis ofthe neural tissue displayed a more complete bridgingwhen compared to a Teflon implant, which is cur-rently a gold standard for neural implants. The datapresented implied that PPy implants generally exhib-ited greater tissue integration and less inflammation.The feasibility of incorporation of neuronal growthfactors (NGFs) was also shown, with even more neu-ral tissue formation observed for NGF-incorporatedimplants.

Moreover, the conductivity of PPy can be exploitedfor the fabrication of bioelectrical circuits that integrateelectrical and neural signals. Cui et al. investigatedthe use of PPy as a neural probe.40,41 These neuralprobes facilitated the functional stimulation andrecording from the peripheral or central nervoussystem. The conductivity of PPy might induce selectiveneurons to attach to the electrode and achieveneuronal tissue regeneration. However, problems wereencountered, including loss of the ability to recordneural activity with time. Modifications by patterningpeptide or peptide/protein polymer blends on thesurface were explored to enhance interaction andanchorage between electrode and neurons.

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PolyarylatesSchachter and Kohn introduced an interestingconcept of using a group of polyarylates for drugdelivery vehicles as an alternative to conventionalPLA, PGA, or PLGA systems.42 The polyarylates(Fig. 4(c)) used were tyrosine-derived polymers whichpossessed sites for interaction with peptides. Somepossible interactions included hydrogen bonding andhydrophobic interactions. Addition of PLGA systemas a ‘delayed excipient’ induced a decrease ofinternal pH during its degradation, and weakened thesensitive interaction between peptide and polymer.Using Integrillin as the drug model, the glasstransition temperature of the system varied in therange 36–40 ◦C as the amount of loaded drug varied.The delayed time of the Integrillin release increasedas the initial molecular weight of the PLGA increased.While degradation of PLGA affected the release of thepeptide, degradation of the polyarylates did not havesuch an effect. Preliminary in vitro studies showed thepolymer to be versatile towards drug release kineticsby tuning in the pendant chain and the backbone unit.

Poly(ether ester amide)Another candidate for drug delivery systems arisesfrom the family of poly(ether ester amide)s (PEEAs).They are mainly fabricated by polycondensation ofPEG and diester-diamide to create an amphiphilicsystem.43,44 Diester-diamide acts as a hydrophobicblock for creating reversible physical crosslinks thataccount for stronger mechanical properties in theswollen state. In this case, PEG is the ‘soft’ segmentand diester-diamide acts as the ‘hard’ segment(Fig. 4(d)).

The structure gives versatility allowing the tailoringof hydrogels to suit different drug release profiles.Bezemer et al.43 fabricated the polymer by a two-stepmechanism. The first step involved transesterificationof diamide-dimethyl ester monomers with PEG, andthe second step was polycondensation at 220 ◦C. Theresulting polymers had an intrinsic viscosity rangingfrom 0.58 to 0.78 dL g−1, and they were not completelyamorphous; both depended on the length of thepolymer blocks. It was shown that the soft segmentlength in a microsphere influenced the degree ofswelling, in vitro degradation, and release rate.43

Cytocompatibility was investigated usingHUVEC.44 These surfaces did not perform as pos-itively as TCPS, possibly due to the PEO con-tent. Deschamps et al.44 implanted a low-content-PEG PEEA subcutaneously in rats. Mass loss was7–12 wt% at 14 weeks post-implantation and a slowdecrease in intrinsic viscosity led to bulk hydrolysisdegradation. Tissue responses were found to be sim-ilar to tissue reactions observed upon implantation ofother biodegradable polymers. Fibrous capsules wereobserved, with macrophages and blood vessel infiltra-tion accompanied by cracks on the polymer surface.

A porous scaffold was also fabricated from this mate-rial using compression moulding followed by a saltleaching technique.

Another amphiphilic drug delivery system uses aPCL-based polymer construct containing hydrophilicPEEA.45 The possibility of using PCL macromersof different molecular weights changed the PEEAproperties including crystallinity and the hydrophilic-ity/hydrophobicity, and this would affect the drugrelease rate. In vitro drug release studies using threedifferent drugs of different natures revealed that PEEAsuccessfully enhanced the completion of drug release.An acidic drug was released rapidly after 2 h incuba-tion, while a drug of a basic nature exhibited longercontrolled release.

Poly(amido amine)The use of hydrogels as matrices for tissue engineeringapplications and the search for versatile, easily fab-ricated, biodegradable, and biocompatible matricescontinue. One novel hydrogel is made of poly(amidoamine), a polycationic polymer in nature. It con-tains ter-amino and amido groups regularly arrangedalong the polymer chain, obtained by Michael-typepolyaddition of primary or secondary amines to bis-acrylamides. The method of fabrication made itpossible to have side substituents with biomimick-ing properties, such as carboxyl, ter-amino, hydroxyl,allyl, or bioactive molecules like proteins and peptides.In vitro assays performed by Ferruti et al.46 inves-tigated cytotoxicity, compatibility, and proliferationof the materials and the cells in contact with them.Biological evaluation results exhibited non-toxicity,comparable cell proliferation (more than 70%) to nor-mal TCPS plates, and normal cell morphology.

Furthermore, degradation products were non-toxicand the rate of degradation could be fine tuneddepending on the structure and degree of crosslinking,which, in turn, affected the degradation via a hydrolysismechanism. For example, hybrids of PAA and BSA(bovine serum albumin) degraded fully up to 8 monthsin simulated body fluids.46 It was suspected that thiswas because of the protective characteristics of BSA.Apart from providing a protective layer, BSA made thehybrid substrate more supportive to cell adhesion andproliferation. An attempt to make it more bioactive wasachieved by incorporating agmatine, a decarboxylatedproduct of arginine, derived from RGD sequences.47

Cell adhesion and proliferation assays exhibited up to80% capacity compared to normal polystyrene cultureplates.

CONCLUSIONSBiodegradable synthetic polymers, from a materialstandpoint, are a key area of interest for the devel-opment of new scaffold-based tissue engineeringstrategies. A critical issue in scaffold-based tissueengineering is the assembly of cells and extracel-lular material into a three-dimensional architecture

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that allows for both structure and functionality thatmimics the native tissue that is being replacedand/or repaired. As the scaffolds for tissue engi-neering will be implanted in the human body, thescaffold materials should be non-antigenic, non-carcinogenic, non-toxic, non-teratogenic, and possesshigh cell/tissue biocompatibility so that they willnot trigger pathological reactions after implantation.In addition to materials issues, the macro- andmicrostructural properties of the scaffold are also veryimportant. In general, the scaffolds require individualexternal shape and well-defined internal structure withinterconnected porosity to host most cell types. Froma biological point of view the designed matrix shouldserve various functions, including (1) as an immobi-lization site for transplanted cells, (2) formation of aprotective space to prevent unwanted tissue growthinto the wound bed and allow healing with dif-ferentiated tissue, (3) directing migration or growthof cells via surface properties of the scaffold, and(4) directing migration or growth of cells via release ofsoluble molecules such as growth factors, hormones,and/or cytokines. Some technology platforms of FDA-approved polymeric degradable scaffold systems havealready found promising clinical applications. Newpolymers are under development and their applica-tions need to be a part of the systemic approachto tissue engineering; namely the need to coordi-nate interaction between the parameters of scaffolds,cells, bioreactors, and biomolecular factors as well asthe controlling features of the host response to re-implanted constructs, including phenomena of angio-genesis, inflammation, and immune response. Basedon these results, novel therapeutic options in the areaof polymeric scaffold-based tissue replacement can beexpected in the 21st century.

ACKNOWLEDGEMENTSThe work reported in this review was supported in partby the Biomedical Research Council (BMRC) grantsR-397-000-005-305 (to DWH). The authors thankMaria Woodruff for editing the manuscript.

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