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Composites: Part B 38 (2007) 291–300

A critical review on polymer-based bio-engineered materialsfor scaffold development

Hoi-Yan Cheung a, Kin-Tak Lau a,*, Tung-Po Lu b, David Hui c

a Department of Mechanical Engineering, The Hong Kong Polytechnic University, Hung Hom, Kowloon, Hong Kong SAR, Chinab EBM Biodegradable Materials Limited, Hong Kong SAR, China

c Department of Mechanical Engineering, University of New Orleans, New Orleans, LA 70148, USA

Received 2 February 2006; accepted 6 June 2006Available online 20 November 2006

Abstract

Since the last decade, tissue engineering has shown a sensational promise in providing more viable alternatives to surgical proceduresfor harvested tissues, implants and prostheses. Due to the fast development on biomaterial technologies, it is now possible for doctors touse patients’ cells to repair orthopedic defects such as focal articular cartilage lesions. In order to support the three-dimensional tissueformation, scaffolds made by biocompatible and bioresorbable polymers and composite materials, for providing temporary support ofdamaged body and cell structures have been developed recently. Although ceramic and metallic materials have been widely accepted forthe development of implants, its non-resorbability and necessity of second surgical operation, which induces extra for the patients,limit their wide applications. This review article aims at introducing (i) concept of cartilage tissue engineering, (ii) common types ofbio-engineered materials and (iii) future development of biomaterial scaffolds.� 2006 Elsevier Ltd. All rights reserved.

Keywords: A. Recycling; B. Environmental degradation; Biocomposites

1. Introduction

In the musculoskeletal system of a human body, itinvolves the integration of tissues and other structuralmembers such as bones, cartilages, tendons, ligaments,peripheral nerves and spinal nerve roots, skeletal musclesand etc. In connecting those bones, there are three typesof joints exist in the human body namely fibrous, cartilag-inous and synovial. Only the synovial or diarthrodial jointsallow a large degree of motion. The coverage of the diar-throdial joints is a layer of hyaline articular cartilage,which is a thin, dense, white but translucent connective tis-sue, which forms articulating surfaces of the diarthrodialjoints. The articular cartilage is a soft tissue composed pri-marily of a large extracellular matrix (ECM) with a sparse

1359-8368/$ - see front matter � 2006 Elsevier Ltd. All rights reserved.

doi:10.1016/j.compositesb.2006.06.014

* Corresponding author. Tel.: +852 2766 7730; fax: +852 2365 4703.E-mail address: mmktlau@polyu.edu.hk (K.-T. Lau).

population of chondrocytes distributed throughout the tis-sue. ECM is mainly composed of collagen, which formsinsoluble tightly woven fibers, water and proteoglycanare dispersed through the collagen framework as a solublegel making the matrix biphasic. Collagen fibrils can with-stand tension but not in compression in general, and thusthe matrix possesses high tensile strength, whereas for pro-teoglycan macromolecules which are the main protein inhyaline articular cartilage to withstand compression byattracting and entrapping large amount of water. The artic-ular cartilage, as a load-bearing material, supports jointsmovement with relatively low coefficient of friction andhigh wear resistance [24] (Fig. 1).

The articular cartilage is predominantly loaded in com-pression and is viscoelastic in nature, its defects are usuallycaused by congenital diseases, paediatric growth plate dis-orders, trauma-induced injuries and others. Once injured,the capacity of tissue regeneration and self-recovery forfull-thickness defects is very low, because of its nature of

Fig. 1. Articular cartilage coated on the surfaces of two bones [24].

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no contact with blood and exhibits low cell to matrix ratio[6]. Statistically, It has been reported that more than900,000 Americans suffer from articular cartilage injurieseach year and they have been waiting for organ trans-plants. This leads to the situation that the transplantationof organs from compatible donors is severely insufficientto go around. Driven by the shortage of the organ trans-plants, tissue engineering becomes a tremendous promisefor providing more viable alternatives to surgical proce-dures for harvested tissues, implants and prostheses [39].In the past decades, most of the surgical inventions torepair damaged cartilage have been directed to the treat-ment of clinical symptoms rather than the regeneration ofhyaline cartilage, such as pain relief and functional restora-tion of joint structures and articulating surface [9].

The term of Tissue Engineering was initially defined byattendees of the first NSF (National Science Foundation,USA) sponsored meeting in 1988 as ‘‘application of theprinciples and methods of engineering and life sciencestoward fundamental understanding of structure–functionrelationship in normal and pathological mammalian tissuesand the development of biological substitutes for the repairor regeneration of tissue or organ function’’ [3]. A compre-hensive review was given by Langer and Vacanti [18] andhas addressed that tissue engineering is ‘‘an interdisciplin-ary field that applies the principles of engineering and lifesciences toward the development of biological substitutesthat restore, maintain, or improve tissue or organ function.It means that this field involves multi-disciplinary knowl-edge between life science, biological cells, engineering bio-materials, biomedical factors and biotechnology. Untilrecently, the adoption of micro-mechanics for the predic-tion of the mechanical properties of tissue has appeared[20].

Currently, there are three approaches in tissue engineer-ing: (1) usage of segregated cells or cell substitutes toreplace those functional cells; (2) delivery of tissue inducingsubstances to targeted locations including growth and dif-ferentiation factors; and (3) seeded and growing cells inthree-dimensional scaffolds [29]. On the other hand, thereare three general strategies for joint surface restoration;

they are (i) enhancement of intrinsic healing capacity ofcartilage, (ii) replacement of damaged articular cartilagewith osteochondral transplants and (iii) regeneration ofarticular cartilage surface with the growth of new cartilage[10]. Autologous chondrocyte implantation, donor tissue isharvested and separated into isolated cells which are thenattached and cultured onto some suitable substrates ulti-mately implanted to the wound site, has yielded promisingresults in treating articular cartilage defects [6], however,the usage of suspended cells and a periosteal flap for hold-ing chondrocytes in positions have certain difficulties ontheoretical and experimental results [7]. Moreover, tissuesin vivo are generally three-dimensional construction; cellscultured in three-dimensional substrates are more likelyto reflect in vivo scenarios. In order to address these prob-lems, three-dimensional cell seeded matrices are studied incartilage tissue engineering to facilitate in vivo implanta-tion and promote cartilage repair [8].

Three-dimensional scaffolds are built for the accommo-dation of mammalian cells, the guidance of the cell growthand the regeneration of three-dimensional tissues withproper structures and functions. The scaffolds are fabri-cated in proper size and shape, cells are seeded on scaffolds,which are harvested from the patient and implanted to thewound sites. Cells migrate and proliferate in all directionsusing the three-dimensional scaffold architecture to popu-late all regions of the constructs. The growth rate of neo-tissues is generally affected by the properties of scaffoldsincluding their structures, composition, architecture andbiocompatibility of the biomaterial; the adhesion, migra-tion and proliferation of cells and the external stimuliincluding growth factors, nutrients and other bioactiveagents that modulate the functions of cells [2]. In Fig. 2,the effect of seed density versus tissue growth rate is shown.

As the growth and regeneration of neo-tissues can begreatly influenced by the biological environment that essen-tially provides to the cells and also the scaffold characteris-tics including biocompatibility, biodegradability, micro- ormacro-porosity, and interconnectivity etc. Different sortsof considerations on the design and build of scaffolds hasto be clearly understood by all researchers and scientists

Fig. 2. Effect of seed density on growth rate [2].

Table 1Commercially available resorbable orthopedic implants

Polymer Abbreviation Purpose

Poly(glycolode) PGA 3-D polymer scaffolds for celltransplantation

Poly(lactic acid) PLA 3-D polymer scaffolds for celltransplantation

Poly(L-lactide) L-PLA Fracture fixation, suture anchor,suture anchor, ACL reconstruction,Rotator cuff repair, meniscus repair

Poly(D,L-lactide) D,L-PLA Fracture fixation, ACL repair,suture anchor

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working in or intending to fit into this area. This paperintends to provide a comprehensive review on the effectsand relationships of the choice of different types of bioma-terials for fabricating scaffolds. The advantages and limita-tions of these biomaterials are also discussion in detail.

2. Scaffolds for tissue engineering

Scaffolds used in tissue engineering applications, partic-ularly for articular cartilage repair should demonstratecompatible biological and physical properties which matchthe physiological conditions in vitro and in vivo. The majorfunction of the scaffolds is to provide a temporary supportto body structures to allow the stress transfer over-time toinjured sites, and facilitate tissue regeneration on the scaf-folds. Recognition of this phenomenon raises the possibil-ity that the scaffolds should have sufficient mechanicalintegrity to retain their shape and at implantation site untiltheir function completed. The selection of appropriatematerials for implant development requires satisfying somecriteria, in which the effect to the human has to keep asminimal as possible. Besides, the biodegradability and theability of cell growing from the scaffolds are highly depend-ing on their physical and structural properties, such asinterconnectivity, porosity and surface morphology. Thedesign of the scaffold systems for articular cartilages, inhuman musculoskeletal system, is rather complicated dueto the combination of repeated shear and compressivestresses co-exist. In the sack of simplicity, most of the pre-vious studies were simply focused on a pure compressivestress acted on the articular cartilage system. In general,the design of scaffolds should allow the temporary supportof body’s structures or injured sites through the stresstransfer from one part to another, and thereafter bedegraded over time. To select an appropriate type of mate-rials for designing and manufacturing scaffolds, several keyfactors, listed as the follows, have to be studied in detail.

2.1. Biocompatibility

The biocompatibility of materials inside the humanbody plays a key role in tissue engineering, which ensures

the materials are safe for use in the human body and inthe endogenous fluids. Being an ideal implant for repairingtissue defects and regenerating neo-tissue for wound sites,the materials used must be biocompatible, i.e., the materi-als must not induce any inflammatory response, extremeimmunogenicity or cytotoxicity to native cells, tissues ororgans in vivo. Since the implants are normally imbeddedinto the human body and last for a period of time, bi-prod-ucts result from the degrading process of the implantsshould not produce any harmful material and/or elementto the body also. For recent polymer-based scaffold designand development, the basic requirement is that the scaf-folds must be naturally degraded with time, and graduallyabsorbed by the human body itself without generating anyside effect. This process allows the functionality of thehuman body gradually recovers without too relying onthe support of the scaffolds. Due to the demand on scaf-folds and implants has been increasing, many comprehen-sive research on the biocompatibility of scaffolds and thetoxicity of their decomposed products have been pro-gressed rapidly in the past ten years.

2.2. Biodegradability and bioresorbability

The field of biodegradable polymers is a fast growing areaof polymer science because of the interest of suchcompounds for temporary surgical and pharmacologicalapplications. As referred to the definitions of these twophenomena, biodegradable materials, like polymers can bedecomposed naturally but their degraded products willremain inside the human body. For bioresorbable materials,they will degrade after a certain period of time of implanta-tion, and non-toxic products will be produced in the ways ofelimination with time and/or metabolism. For the chemicaldegradation, two different modes are defined, they are (1)hydrolytic degradation or hydrolysis which is mediated sim-ply by water and (2) enzymatic degradation which is mainlymediated by biological agents such as enzymes. Scaffoldsshould be biodegradable allowing ECM to occupy the voidspace when the biomaterial is degraded. Today, there aremany types of materials as listed in Table 1 that are widelyused as bioabsorbable implants and their market is expand-ing rapidly worldwide. In fact, the advantages of using bio-degradable polymers over the traditional metallic materials

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for implant development include the reduction of stressbearing capability over time since the polymer will bedegraded naturally, and the alleviation of pains, both phys-ically and physiologically due to the need of second surgicaloperation for removing metallic implants.

Fig. 4. Decomposition rate of porous (m) and original (d) PCL films, andporous (j) and original (r) PLA films [23].

Fig. 5. Percentage of cells attached to the collagen-glycosaminoglycan(CG) scaffolds with different pore sizes [28].

2.3. Degradation rate, pore size and surface morphology

The major function of a biodegradable polymer-basedscaffold must be able to retain at the implantation site withmaintaining its physical characteristics and ordinarymechanical properties, as well as supporting the attach-ment, proliferation and differentiation of cells, until theregeneration of tissues on an injured site. Ideally, the deg-radation rate of the scaffold should be matched with therate of neo-tissue formation so as to provide a smoothtransition of the load transfer from the scaffold to the tis-sue. However, based on recent research investigations, itwas found that the degradation rate of different types ofbiodegradation polymer has difference, depending on thecomposition of the polymer, conditions of loading andambient environment. The enzymatic degradation rate pro-ceeds from the surface of the polymer, therefore surfacearea and condition (porosity) of the polymer is one ofthe important factors to control the degradation rate [23].

Up-to-date, there are many ways to manufacture porousscaffolds, they are porogen leaching, emulsion freeze-dry-ing, expansion in high temperature gas, 3-D printing, phaseseparation techniques, thermal phase separation [26]. InFig. 3, two porous films of PLA and PCL were fabricatedby using the freeze-dry method is shown. It was found thatthe degradation rate of the films varied with the porous sizeand density. The two porous films decomposed faster thanthat of the films without pores (as shown in Fig. 4) [23].Large number of pores may be able to enhance the masstransport and neo-vascularization within the implants,whereas smaller diameter of pores is more preferable toprovide larger surface per volume ratio. Moreover, insteadof adjusting the pore size, the shape of the pores is also akey to affect the efficiency of tissue regeneration. It isimportant to note that some specific pore sizes canenhances the cellular activity, but optimal size and geome-

Fig. 3. SEM photographs of porous PLA

try are highly dependent on specific cell types grown oninjured sites. For bone in-growth, the optimal pore size isin the range of 75–250 lm. On the other hand, for ingrowth of fibro-cartilaginous tissue, the recommended poresize ranges from 200 to 300 lm [1]. In Fig. 5, it shows thepercentage of cells attached onto scaffolds with differentpore sizes [28] and it further proves that the control of poresize for different scaffolds is necessary.

(left) and PCL (right) surface [23].

Table 2Mechanical properties of the biodegradable polymer SPLA over time [27]

Material Immersiontime (days)(0.154 M ofNaCl)

Compressivemodulus(MPa)

Compressivestrength(MPa)

Strain atbreak (%)

SPLA 70 0 200 ± 45 14 ± 5 30 ± 93 22 ± 3 5 ± 3 35 ± 8

14 39 ± 9 4 ± 2 32 ± 2

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Most organ cells require suitable substrate to retain theirability to adhere, proliferate and perform the later differen-tiated functions including spreading and differentiatingseeded cells. Therefore, the design of scaffolds for tissueregeneration, the control of surface topography and rough-ness is important to allow the migration of the cells on thescaffolds’ surface. The rougher surface also enhances thediffusion rates to and from the scaffolds as well as facilitatesvascularization and improves the oxygen and nutrientssupply and waste removal. In Fig. 6, the osteoblast-likecells were grown on poly(L-lactic acid) (PLLA) substrateswith different surface morphologies, one was a smooth flat-ted surface while another with PLLA islands [37]. Theheight of the cells on smooth surface was higher than thatof the cells on the island-patterned PLLA, which impliedthat the cells preferred to stretch on a rough surface. Theuniformity of the cells grown on the smooth PLLA sub-strate was also relatively poor as compared with thatgrown on the island-patterned PLLA substrate. Besidesin the figure, it also clearly indicates that the total contactarea of the cells to the PLLA was much larger as comparedwith the one without the island-patterned structure built onthe PLLA surface [37].

2.4. Mechanical integrity

In order to fulfil the physiological loading requirement,tissue engineered scaffolds should act as temporary physicalsupports (or called ‘‘structural members’’) to withstand theexternal and internal stresses until neo-tissues are gener-ated. As for a native cartilage, its metabolism, synthesis,and the organization of ECM are affected by the mechanicalenvironment experienced by chondrocytes. These forces areusually generated during both the implantation procedureand the mechanical forces experienced at the joint surfaces.In vitro, the scaffolds have to be able to resist the culturingmechanical environment such as direct compressive, hydro-static pressure and static loads. Especially, for the dynamicfunctional environment of the scaffolds, it can have a signif-icant effect on the scaffolds in vitro degradation as well asrelease of bioactive agents. Consequently, the propertiesof the scaffolds should be as similar as the properties of

Fig. 6. SEM images of OCT-1 osteoblast-like cells attached on the (left) smoo[37].

neo-tissues generated, in order to provide a proper struc-tural support in the stage of healing. At the later stage, allthe loads have to be totally transferred to the neo-tissuessince the scaffolds should be degraded gradually. Neveset al. [27] have studied the change of mechanical propertiesof starch with poly (lactic acid) (SPLA) with time and theresults are shown in Table 2. It is obvious that the strengthof the SPLA dropped substantially after immersed into iso-tonic saline solution (0.154 M of NaCl) for 3 days in spite ofthe change of its deformation was small.

3. Types of scaffold biomaterials

As mentioned in the previous sections, the major con-cern in developing scaffolds for different surgical andorthopedic operations is the selection of suitable biomate-rials, which must be biocompatible and biodegradable.Potential materials that have been proven with experimen-tal data on their validity for biomedical applications aremetal, ceramics, polymers and the combinations of thesematerials. For metallic materials and ceramics, they havecontributed to lists of medical applications, particularlyin orthopedic tissue replacements. However, there aretwo major limitations, they are (i) not biodegradable exceptbiodegradable bio-ceramics, and (ii) poor processability.For articular cartilage and scaffold design, normally thestiffness and coefficient of friction should be relativelylow and the shape is highly irregular, in which the metallicmaterials cannot fulfil this requirement.

To successfully adopt biodegradable polymers for carti-lage tissue engineering applications, understanding thecompressive and shear stress transfer mechanisms at

th surface of PLLA substrate and (right) island-patterned PLLA substrate

Fig. 7. Alginate gels of various moulded shapes [17].

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cartilage joints is essential to ensure the physical andmechanical properties of articular joints are strong enoughto sustain different loadings. Basically, there are four typesof polymer-based materials used for cartilage joints, theyare (i) natural polymers, (ii) synthetic polymers, (iii) hydro-gels and (iv) composites [6]. Many natural polymers foundin living organisms of known biocompatibility, actually arehydrogels and some are being popularly used in combina-tion with other materials as composite scaffolds. Such poly-mers can be used to replace or regenerate native tissuestructures and allows positive cell interactions withsurrounding tissues. Conversely, synthetic polymers areformed through controllable chemical processes to achievedesirable material and chemical properties for a wide rangeof biomedical applications. The synthetic polymers have apromising advantage over the natural polymers for scaffolddevelopments because their mechanical and proliferationproperties are comparatively more predictive and repro-ducible. Hydrogels are primarily composed of fluid thatswells the polymer network to form a biphasic construct.Although the hydrogels can be synthesized, they are alwaysformed naturally. Composite scaffolds are built by mixingtwo or more materials to achieve desirable properties andcharacteristics by taking advantages from each of thematerials.

3.1. Natural polymers

Natural polymers used in cartilage engineering applica-tions include collagen [10,30,32], alginate [7,11,25,31], aga-rose [22,38], chitosan [5,21], fibrin [12,33] and hyaluronicacid–based materials [11–13,16]. Natural polymers oftenpossess highly organized structures and may contain anextracellular substance, called ligand, which can be boundto cell receptors. Although they are of known biocompati-bility, lack of large quantities and difficult in processinginto scaffolds limit them for clinical applications. More-over, as natural polymers can guide cells to grow at variousstages of development, they may stimulate an immuneresponse at the same time. This leads the concerns overantigenic and deliver of diseases for allograft. Since thedegradation of these polymers depends on the enzymaticprocesses, degradation rate may vary from patient topatient.

Collagen is a fibrous protein that is the major compo-nent in connective tissues, which has been used widely fortissue regeneration applications, particularly for soft tissuerepair. It provides cellular recognition for regulating cellattachment and function, but may lead to the concern ofadverse immune response. Collagen favours cell adhesionthat is normally found in joint tissues and those exogenouscells embedded in a collagen delivery device. Although ithas to be purified before seeded with chondrocytes to makeit less antigenic, it may transmit pathogen and induceimmune reactions. On the other hand, it has poor mechan-ical properties, difficult to handle and fabricate, and is alsoless control over biodegradability [42].

Alginate is a polysaccharide extracted from algae. It canbe formed as gels or beads (as shown in Fig. 7), which arealso used for cartilage repair, as a support to encapsulatecells within the scaffold [4]. Through encapsulation, chon-drocyte’s phenotype during culture can be maintainedand this allows cell that has been cultured in monolayerto be re-differentiated. Alginate, as a liquid, is injectedand cross-linked with calcium to prevent the migration ofthe defect and allow the formation of a tissue with similarmorphological characteristics as the native hyaline carti-lage. Conversely, there are concerns on its slow degrada-tion rate as it may cause problems when the neo-tissuestarts to grow; insufficient mechanical integrity, and long-term implants are impossible as alginate scaffolds will loseits functionality within a year.

Agarose is also a polysaccharide but it is extracted fromseaweeds. It can be used to form hydrogels, which makesthe seeded cells with more uniform distribution throughoutscaffolds. It can be prepared as an injectable scaffold andprovide a three-dimensional environment helps to maintainchondrocyte in a rounded shape during culture. However,it degrades slowly like alginate. Agarose exhibits a temper-ature-sensitive solubility in water for encapsulating cells. Ithas been used widely for chondrocyte behaviour studiesin vitro [40] and showed that agarose transmits the appliedmechanical forces to the chondrocytes during compression,which excites the cells to produce more extracellular matrixproteins than during static controls.

Chitosan is a biosynthetic polysaccharide deacetylatedderivative of chitin which is a naturally occurring polysac-charide extracted from crab shells, shrimps etc. or from afungal fermentation process. Chitin and chitosan are semi-crystalline polymers having a high degree of biocompatibil-ity in vivo. Chitosan can be prepared as temperature-sensitive carrier materials and are injectable as fluids. Itforms gels at body temperature and has the ability to deli-ver and interact with growth factors and adhesion proteins.Degradation of chitosan is controlled by the residualamount of acetyl content and it can degrade rapidlyin vivo according to the deacetylation with the polymer.

Fig. 8. Chitosan molecular structure.

Fig. 9. A synthetic polymer scaffold synthesized from D,D-L,L polylacticacid [15].

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In addition, porosity of chitosan scaffolds can be controlledwhich can affect the strength and elasticity of the implants.In Fig. 8, the Chitosan molecular structures is shown.

Fibrin glue is formed from the mixture of fibrinogen andthrombin and allows them to solidify. As the patient’s ownfibrinogen and thrombin can be used, it is not a concern onsterilization, biocompatibility and temperature change dur-ing setting. It is always being used as a carrier for cells andin conjunction with other scaffold materials. This materialcan be completely degraded and injected but with a draw-back on lack of mechanical strength for articular cartilageengineering applications.

Hyaluronan (or hyaluronic acid, HA) is an anionic poly-saccharide, which can be obtained from natural sources orthrough microbial fermentation and has often been used ascarriers for cells to regenerate various tissues. It can becross-linked to form scaffolds and seeded with chondro-cytes. As it is injectable, it can be used in irregularly shapeddefects and implanted with minimal invasion. Hyaluronicacid can be easily chemically modified, like collagen. Andit is a desirable biomaterial since it is not antigenic and elic-its no inflammatory or foreign body reaction. However, ithas limited range of mechanical properties for applications.

3.2. Synthetic polymers

Synthetic polymers are man-made polymers, which havethe advantages over the use of natural origin polymers asthey are more flexible, more predictable and processableinto different size and shapes. The physical and chemicalproperties of a polymer can be easily modified and themechanical and degradation characteristics can be alteredby their chemical composition of the macromolecule. Thefunctional groups and side chains of these polymers canbe incorporated, i.e., the synthetic polymers can be self-cross-linked or cross-linked with peptides or other bioac-tive molecules, which may be a desirable biomaterial forcartilage tissue engineering. Additionally, synthetic poly-mers are generally degraded by simple hydrolysis that isdesirable as the degradation rate does not have variationsfrom host to host, unless there are inflammations andimplant degradation etc. to affect the local pH variations.The most extensively used synthetic polymers are poly-(glycolic acid) (PGA), poly(lactic acid) (PLA) and theirco-polymers; polycaprolactone (PCL) and polyethyleneglycol (PEG).

Poly(lactic acid) (PLA) is an alpha polyester used widelyin medical applications and it has been approved by the

FDA for implantation in human body. PLA degradesslower than PGA due to its hydrophobic characteristic,which limits the water absorption of thin films and slowsdown the backbone hydrolysis rate. Based on availabledata to date, the duration of degradation can be rangedfrom 12 months to over 2 years. PLA can have a relativelylow tensile strength and modulus of elasticity alone,depending on which isomer is used in the polymer.Although many researchers have tried to investigate por-ous PLA scaffolds for the usage in orthopaedic applica-tions, PLA is primarily used as a non-woven mesh fortissue engineering applications. Ishaug-Riley et al. [43] haveshown that fewer chondrocytes were attached to PLLAthan to PGA at the initial stage, but as both surfaces allowextensive cells proliferation, giving the similar total numberof cells at confluence. Lactic acid exists in two stereoisom-erism forms, which can be separated into four morpholog-ically distinct polymers namely D-PLA, L-PLA, D,L-PLAand meso-PLA (Fig. 9). Degradation products of thesematerials reduce local pH, accelerate the polyester degrada-tion rate and induce inflammatory reaction.

Poly(glycolic acid) (PGA) is the simplest linear aliphaticpolyester. It has a highly crystalline structure with a highmelting point and low solubility in organic solvents. PGAdegrades too rapidly that makes it difficult to process andhandle. The degradation period can be from 4 to 12months, which is much shorter when comparing withPLA. The degradation products of PGA are naturallyresorbed by the body makes it desirable to be a biomate-rial. In comparison to PLA, PGA has a relatively high ten-sile strength and modulus of elasticity. However, thesecould not directly translate to the scaffold stiffness as thematerial is generally formed into a mesh or a felt for carti-lage tissue engineering purposes. PGA is often extruded asthin polymer strands, which is around 13 lm in diameter, itmust be moulded into non-woven mesh discs as a scaffold.This provides a high porosity environment for cells to growand proliferate. To the contrary, this limits its immediateapplications on hard tissue engineering as the tissue needto fill much of the void space in order to maintain the

Fig. 10. Compressive modulus of PLG scaffolds with and without PGAfiber reinforcements [34].

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construct with sufficient mechanical integrity. Conse-quently, the mechanical properties of PLA and PGA arerelatively weak for high porosity scaffolds, which limit theirusage on the hard tissue regeneration applications.

PLA–PGA co-polymers have no linear relationship onthe ratio of PLA to PGA but it is critical as PLAs hydro-phobic characteristic reduces the rate of backbone hydroly-sis, whereas PGA is high crystalline and crystallinitystructure, which leads to rapidly lost in PLA–PGA co-poly-mers. That is, the rate of hydration and hydrolysis will beincreased by these morphological changes. Therefore, thetypically 50 (PLA):50 (PGA) co-polymers degrade morerapidly than either PLA or PGA. PLA or PGA based poly-mers can be processed by injection and/or compressionmouldings but the mechanical properties of the scaffoldslimit them from any load-bearing applications. In order touse them for clinical applications, there are two self-rein-forcing techniques namely sintering (highly oriented PLAfibers are joined together by a same material matrix, sinter-ing at 5–25 K below the respective melting temperature atpressures around 50–80 MPa) and fibrillation (polymer isoriented in solid state through a drawing process, a meltmoulded polymer billet is passed through a die, transform-ing the partially crystalline structure into a highly orientedfibril structure by mechanical deformation). These methodsinduce highly oriented polymer morphology and enhancethe mechanical properties. Like PLA and PGA, these co-polymers are fabricated as non-woven mesh and felt forthe usage of cartilage engineering scaffolds and the intercon-nectivity, porosity, pore size and void space can be adjustedduring the fabrication process to give proper construction.

Another type of synthetic polymer is polycaprolactone(PCL), which degrades through hydrolytic scission withresistance to rapid hydrolysis. PCL has the presence ofhydrolytic aliphatic–ester linkage, which is unstable to leadit biodegradability. Its degradation duration can be as longas 24 months for being completely degraded. Thus, it isalways used to co-polymerize with other materials to havedesired degradation properties. Its biocompatibility, degra-dation and mechanical strength characteristics are suitableand compatible for orthopaedic applications. In essence, itis more compatible for PCL to be long-term implants andrelease control applications. Honda et al. used poly(L-lac-tide-epsilon-caprolactone) as a biodegradable sponge andimplanted it into a nude mice. After 4 week experimentallyprocess, the result showed that there was the formation ofcartilage-like structures in the construct [14].

The last but not least type of synthetic polymer is poly-ethylene glycol (PEG), also known as polyethylene oxide(PEO). This is a newer material that is used to restrictand control the attachment of cells and proteins on scaf-folds since it is very hydrophilic. With the increase ofhydrophilic properties, antibodies and other proteins aredifficult to attach to the scaffolds, which lessen any adverseimmune response. By using PEG in a co-polymer, research-ers can control the cell attachment characteristics of thescaffolds and enhance the biocompatibility of the co-poly-

mer. PEG alone has a relatively high compressive moduluscorresponding to a higher molecular weight. Investigationshave been done on PEG co-polymers including PEG–PLAconstructs, PEG–PPF, PEG dimethacrylate and lactide-based PEG networks which have better degradation char-acteristics than PEG alone and still retaining the desiredbiocompatibility.

3.3. Hydrogels

Hydrogels are sub-class of natural and synthetic poly-mers. They are cross-linked hydrophilic polymers that withthe characteristic of swelling significantly when placed in apolar, liquid solution. That is, they contain large amountsof water without dissolution. They can retain their shapeeven they have up to 99% water by volume. This highlyhydrated composition can be linked to that of native artic-ular cartilages. Hydrogels are ideally used as injectablescaffolds due to their mass is composed of water primarily,they can be used to fill irregularly shaped defects, allowminimally invasive surgical procedures and act as facilita-tor to incorporate with cells and bioactive agents. In carti-lage engineering, hydrogels are being used to encapsulatecells and growth factors in a polymer network, whichimmobilizes the cells and allows differentiation in chondro-cytes more effectively by forcing them to retain in arounded shape. Also, the growth factors encapsulatedcan remove any complexity might be present due to theconcentration gradients in the media. Agarose and simplehydrogels are often used for direct compression tests onchondrocytes. It is because the dynamic compression isusually conducted at the frequency of around 1 Hz, hydro-gel scaffolds can be easily fabricated to have sufficient elas-ticity to recover from the compressive deformation andstay in contact with the loading platen rather than theusage of fibrous mesh. Besides, as hydrogels can exert con-trolled compressive forces on encapsulated cells, it is moresimilar to the physiological conditions.

Most of the natural polymers can be formed into hydro-gels including collagen, alginate, agarose, chitosan, fibringlue etc. Conversely, none of the natural hydrogels have

Fig. 11. 3-D scaffolds made by PLA/HA ceramic (a) [36] and PCL laminate (b) [41].

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sufficient stiffness to function immediately in vivo. For syn-thetic hydrogel scaffolds, many works have been studied onPEG-based polymers. PEG-diacrylamide can be performedin contact with cells through a photopolymerization step,although there is still concern on the toxicity of cross-link-ing agents. Since PEG hydrogel is lack of degradability,researchers try to impart it by formulating co-polymersof PEG with other synthetic biodegradable polymers(PLA, PGA and PPF), or by having enzymatic degradationlinkages to the PEG backbone. Other than that, poly(N-isopropylacrylamide) and linear poly(acrylic acid) formsemi-interpenetrating polymer networks [35]. However, inorder to produce sufficiently stiff synthetic scaffold, theproblem is almost all cross-linking agents are cytotoxic atthe required levels.

3.4. Composites

Scaffolds that are made by composites consist of two ormore materials are discussed previously, these materialsused together to produce a better scaffold taking the advan-tages from each of the composed materials independently.Very often, some reinforcements, like fibers are made bythe same or different bulk materials. They will be addedto the composition, such as reinforced PLGA with PGAfibers developed by Slivka et al. [34], which enhanced thecompressive modulus and also the yield strength of the ori-ginal materials. In Fig. 10, the mechanical properties ofPLG scaffolds with and without PGA fiber reinforcementsare shown. Notwithstanding the reinforcements can be usedto enhance the materials’ properties, we have to consider thecriteria of the required application. For example, there areonly proteoglycan, type II collagen and water inside thearticular cartilage, any material that stimulates other thanthese compositions in the wound sites should be avoided.

In increasing the mechanical properties of biodegrad-able polymers, silk-based fibers reinforced biodegradablepolymer composites have emerged recently. Spider andsilkworm silks have been recognized as high strength ani-mal silks that can be resorbed by human body. Combiningthese silks with biodegradable polymers can produce amoderate strength and durable biocompatible and biore-sorbable polymer-based composite for biomedical and

orthopedic applications [19]. Taboas et al. [36] and Zeinet al. [41] developed 3-D scaffolds with controlled localporous and internal architectures, using polymer–ceramiccomposites and laminated PCL, respectively (Fig. 11) forcell proliferation. It was found that the 3-D scaffolds pro-vided better mechanical performance depending on theirporosity and lay-up orientations, for woven type 2-D scaf-fold materials.

4. Conclusion

Recently, research of new types of biodegradable materi-als for scaffold development has been the hottest topic inboth advanced composites and biomaterials communities.The integration of composite techniques to fabricate high-strength and durable biodegradable polymer-based com-posites is of a great interest to all researchers and engineers.Properly applying right polymer with the considerations ofporous size, degradation rate and surface morphology, as aparent material with mixing with high-strength fibers canindeed provide a good alterative for existing polymer andmetallic based biocompatible materials for scaffold applica-tions. This review article provides fundamental informationfor researchers and engineers working in advanced compos-ite industry, to open a new step to the development of bio-engineered composites. Cross-disciplinary research effortsare definitely needed in bridging expertise from bio-, nano-and advanced composite areas, to work closely along thisnew scientific and engineering direction.

Acknowledgement

This project was funded by The Hong Kong PolytechnicUniversity Research Grant.

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