Engineering functionally graded tissue engineering scaffolds

J O U R N A L O F T H E M E C H A N I C A L B E H AV I O R O F B I O M E D I C A L M A T E R I A L S 1 ( 2 0 0 8 ) 1 4 0 – 1 5 2

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Review article

Engineering functionally graded tissue engineering scaffolds

K.F. Leonga,∗, C.K. Chuaa, N. Sudarmadjia, W.Y. Yeonga,b

a Rapid Prototyping Laboratory, School of Mechanical and Aerospace Engineering, Nanyang Technological University, North Spine,50 Nanyang Avenue, Singapore 639798, Singaporeb Forming Technology Group, Singapore Institute of Manufacturing Technology, 71 Nanyang Drive, Singapore 638075, Singapore

A R T I C L E I N F O

Article history:

Received 20 April 2007

Received in revised form

25 October 2007

Accepted 3 November 2007

Published online 17 November 2007

A B S T R A C T

Tissue Engineering (TE) aims to create biological substitutes to repair or replace failing

organs or tissues due to trauma or ageing. One of the more promising approaches in TE

is to grow cells on biodegradable scaffolds, which act as temporary supports for the cells to

attach, proliferate and differentiate; after which the scaffold will degrade, leaving behind

a healthy regenerated tissue. Tissues in nature, including human tissues, exhibit gradients

across a spatial volume, in which each identifiable layer has specific functions to perform

so that the whole tissue/organ can behave normally. Such a gradient is termed a functional

gradient. A good TE scaffold should mimic such a gradient, which fulfils the biological and

mechanical requirements of the target tissue. Thus, the design and fabrication process

of such scaffolds become more complex and the introduction of computer-aided tools

will lend themselves well to ease these challenges. This paper reviews the needs and

characterization of these functional gradients and the computer-aided systems used to

ease the complexity of the scaffold design stage. These include the fabrication techniques

capable of building functionally graded scaffolds (FGS) using both conventional and rapid

prototyping (RP) techniques. They are able to fabricate both continuous and discrete types

of FGS. The challenge in fabricating continuous FGS using RP techniques lies in the

development of suitable computer aided systems to facilitate continuous FGS design. What

have been missing are the appropriate models that relate the scaffold gradient, e.g. pore

size, porosity or material gradient, to the biological and mechanical requirements for the

regeneration of the target tissue. The establishment of these relationships will provide the

foundation to develop better computer-aided systems to help design a suitable customized

FGS.c© 2007 Elsevier Ltd. All rights reserved.

Contents

1. Introduction .................................................................................................................................................................................141

2. Functional gradients in natural structures and functionally graded scaffolds (FGS) .....................................................................142

2.1. Biological requirements .....................................................................................................................................................142

∗ Corresponding author. Fax: +65 6791 2859.E-mail address: [email protected] (K.F. Leong).

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J O U R N A L O F T H E M E C H A N I C A L B E H AV I O R O F B I O M E D I C A L M A T E R I A L S 1 ( 2 0 0 8 ) 1 4 0 – 1 5 2 141

2.2. Mechanical requirements ..................................................................................................................................................143

2.3. Anatomical requirements ..................................................................................................................................................144

2.4. The effects of FGS ..............................................................................................................................................................144

2.5. Spatial gradients of FGS .....................................................................................................................................................145

3. Methods for building FGS and their mechanical properties ..........................................................................................................145

3.1. Conventional methods.......................................................................................................................................................145

3.2. Rapid prototyping (RP) methods ........................................................................................................................................146

4. Computer-aided tissue engineering and FGS................................................................................................................................148

5. Conclusion....................................................................................................................................................................................149

References ....................................................................................................................................................................................150

1. Introduction

The loss or failure of organs or tissues due to trauma orageing are a major concern in healthcare as they are a costlyand devastating problems (Langer and Vacanti, 1993; Risbud,2001). This has led to the development of tissue engineering(TE), which aims to create biological substitutes to repair orreplace the failing organs and tissues (Risbud, 2001; Tan et al.,2005a; Vacanti, 2006). One of the more promising approachesin TE is to grow cells on biodegradable scaffolds — highlyengineered structures that act as a temporary support forcells to facilitate the regeneration of the target tissues (Langerand Vacanti, 1993). The main challenge in TE scaffold is todesign and fabricate customizable biodegradable constructswith desirable properties that promote cell adhesion, supportcell growth, proliferation and differentiation, and facilitatethe formation of extra cellular matrix (ECM) (Gross andRodríguez-Lorenzo, 2004). Several important criteria for goodtissue scaffolds include having the appropriate materialswith suitable internal architecture and surface properties,sufficient mechanical properties that match the host tissue,including elastic modulus, compressive strength and fatigueproperties, biocompatibility with predictable degradation rateand sterilizability (Ma, 2004; Mikos and Temenoff, 2000;Hutmacher, 2000).

Identifying suitable biomaterials for building TE scaffoldsthat can degrade appropriately without eliciting too muchforeign body reactions is the first important step towardsbuilding good TE scaffolds. Table 1 shows the commonbiomaterials that are used for producing TE scaffolds, alongwith their degradation times and mechanisms, where naturalpolymers generally degrade much faster than syntheticpolymers and bioceramics. Natural polymers, which areextracted from natural tissues, include collagen and chitosan,while synthetic polymers are manufactured chemicallyand include poly-L-lactic acid (PLLA), poly-glycolic acid(PGA) and poly-ε-caprolactone (PCL). Bioceramics includehydroxyapatite and tricalcium phosphate and are usuallyused in bone regeneration. These biomaterials can be blendedor mixed with one another in order to tailor the degradationtime.

Currently, there are two broad categories of scaffoldfabrication methods, conveniently named conventional andadvanced processing methods. Conventional methods areusually manual processes and include fibre bonding, solventcasting/particulate leaching, gas foaming, phase separationand freeze-drying (Yang et al., 2001; Leong et al., 2003; Ma,2004). These techniques have inherent limitations, e.g. the

heavy reliance on user skills that result in inconsistencyof the scaffold architecture, and the use of toxic solventsthat may be toxic to the cells if they are not properly andcompletely removed (Yeong et al., 2004).

Advanced methods include the use of Computer-AidedDesign (CAD) and Rapid Prototyping (RP) techniques tofabricate TE scaffolds. RP processes have been shown to beviable to produce TE scaffolds, both via direct and indirecttechniques (Yeong et al., 2004). They include Fused DepositionModelling (FDM) (Zein et al., 2002; Too et al., 2002), 3D Printing(3DP) (Seitz et al., 2005; Lam et al., 2002), 3D Bioplotting(Landers et al., 2002a,b), ModelMaker II (MMII) (Manjubalaet al., 2005; Yeong et al., 2006), Stereolithography Apparatus(SLA) (Chu et al., 2001), and Selective Laser Sintering (SLS)(Tan et al., 2005a; Williams et al., 2005). Using RP techniquesto fabricate TE scaffolds are not without limitations andchallenges. These include the use of non-biocompatiblecommercial RPmaterials, limited number of biomaterials thatcan be process on RP machines, the small features attainablethat are dependent on the machine-specific parametersand also the issue of trapped materials within the porousscaffolds (Yang et al., 2002). However, RP techniques holdmuch promise over conventional techniques in terms of partconsistency, design repeatability and the control of scaffoldarchitecture, both at the micro-level (e.g. pore size, porosityand surface-to-volume ratio) and the macro-level (e.g. theexternal shape of the scaffold) architecture (Leong et al., 2003;Yeong et al., 2004; Hutmacher et al., 2004). The control overscaffold micro architecture such as pore size and porosityis particularly important, as TE scaffold mimics the originalenvironment of the healthy organ to successfully regeneratedamaged tissues (Langer and Vacanti, 1993). RP processes thatare used to fabricate TE scaffolds are able to achieve thedesired porosities. The attainable porosities by the various RPmethods are summarized in Table 2.

In mimicking the native environment of the healthyorgan, attention has to be paid to the typical functionsof this environment (Sherwood et al., 2002; Taguchi et al.,2004; Chong and Chang, 2006). These functions in biologicalsystems are not homogeneous and thus the environmentwill vary according to the functional requirements. Thiscan be observed at three levels — biological requirements,mechanical requirements and anatomical requirements. Thispaper will discuss the motivation and need for functionallygraded scaffolds (FGS), as well as the need for a sound strategyto design continuous FGS.

142 J O U R N A L O F T H E M E C H A N I C A L B E H AV I O R O F B I O M E D I C A L M A T E R I A L S 1 ( 2 0 0 8 ) 1 4 0 – 1 5 2

Table 1 – List of biomaterials, their degradation times and degradation mechanisms

Biomaterials Degradation time Degradation mechanism

Natural polymers:� Collagen 2–24 weeks (depending on the cross-linking

degree)Enzymatic degradation (Rothamel et al.,2005)

� Chitosan Half-life weight: 10–56 days for 52%–62%deacetylation degree (DD), more than 84 daysfor DD > 72%

Enzymatic degradation (Ren et al., 2005)

Synthetic polymers:� Poly(L-lactic acid) (PLLA) 2–12 months Hydrolytic mechanism (Yang et al., 2001)� Poly(glycolic acid) (PGA) 4–6 months Hydrolytic mechanism (Yang et al., 2001)� Poly(caprolactone) (PCL) 1–2 years Hydrolytic mechanism (Engelberg and

Kohn, 1991; Holmbom et al., 2005)Bioceramics:� Hydroxyapatite (HA) In the order of years, poor degradation Dissolution, resorbed by osteoclast

(Dellinger et al., 2006)� Tricalcium phosphate (TCP) 8–24 weeks Dissolution, resorbed by osteoclast (Dee

et al., 2002; Fujita et al., 2003; Yamada et al.,2007)

Table 2 – Scaffold porosity values that can be achievedusing RP processes

RP process Porosity achieved

Laminated objectmanufacturing (LOM)

Almost completely dense(Yang et al., 2002)

3D pinting (3DP) 40%–60% (Lam et al., 2002),75%–90% (Zeltinger et al., 2001)

Selective laser sintering(SLS)

37.5%–55% (Williams et al.,2005), 70%–74% (Tan et al.,2005b)

Stereolithographyapparatus (SLA)

70%–95% (not scaffold)(Langton et al., 1997), 50%(indirect RP) (Padilla et al.,2007)

Fused deposition modeling(FDM)

21%–68% (Too et al., 2002),48%–77% (Zein et al., 2002)

3D fiber deposition (3DF) 75%–80% (Malda et al., 2005),70%–87% (Woodfield et al.,2004)

2. Functional gradients in natural structuresand functionally graded scaffolds (FGS)

The structural organizations found in nature are largelydictated by their functions, e.g. load-bearing function,biomechanical function, etc. (Ford et al., 1999a). Simpleobservations of natural tissues and organs show that thesestructures are not homogeneous and there exist naturalfunctional gradients in their structure, as seen in theexamples of bamboo, mollusc shells and human tissuessuch as bone and skin (Ford et al., 1999a; Parenteau et al.,2000). When each layer of the tissue or organ has oneor more specific functions to perform and the tissue ororgan has more than one layer, the tissue or organ is saidto be functionally graded across the layers. Simply, thetissue or organ is described as functionally graded. As such,to regenerate the natural tissue, a successful TE scaffoldshould also be functionally graded. This is to facilitate theappropriately seeded cells to proliferate at the desired layer

and to perform their functions properly and normally toachieve the regeneration of healthy tissues. This can beobserved and described from three perspectives: biological,mechanical and anatomical. Matching these perspectives willprovide the favourable environment required for cell growthand proliferation for successful organ or tissue regeneration.

2.1. Biological requirements

From a biological standpoint, functional gradients are veryoften observed in human organs and tissues. Different layersof the tissue perform different roles in maintaining the organfunctions. One layer may possess cell types or phenotypesthat may be different from other layers. The different celltypes or phenotypes have their own specific functions and areimportant for the tissue or organ to function normally.

To illustrate this, one can take the example of cartilage.Native articular cartilage has anisotropic cell distribution,from the articular surface to the part mating with bone.The cells are arranged differently and can be classifiedinto three zonal organizations and they affect the ECMproduction that is specific to their zonal organizations (Kimet al., 2003). They are, namely, superficial (the one nearestto the articular surface), middle and lower (the one nearestto the bone) zones. Chondrocytes, the cells that maintainand secrete the ECM of cartilage, appear more flattenedin the superficial zone with collagen fibrils orientatedtangentially to the articular surface, while chondrocytes inthe middle zone are more rounded with randomly orientedcollagen fibrils. In the lower zone, chondrocytes are sphericalwith collagen fibrils oriented perpendicular to the jointsurface (see Fig. 1). These conditions relate directly to thebiomechanical functions of the articular cartilage (Woodfieldet al., 2005). The chondrocytes in the superficial surfaceare densely populated and secrete proteoglycan 4 (PRG4),molecules that may aid the surface lubrication. The densityof chondrocytes decreases and the matrix secretion suchas collagen II and aggrecan increases with increasing depthfrom articular surface. The increase in matrix componentsand the varying orientation of collagen fibres makes the


Fig. 1 – Chondrocytes distribution and collagen fiberorientation in articular cartilage.

compressive properties increase as it goes deeper from thesurface, while the superficial zone, which is the softest,increases the contact area and distributes the load (Kleinet al., 2007). Woodfield et al. illustrated that scaffold designwith anisotropic pore architecture could be used to instructzonal cells and ECM distribution to a certain extent inconstructing tissue engineered cartilage (Woodfield et al.,2005).

In the case of regenerating multiple tissues, e.g. boneand cartilage tissues (Sherwood et al., 2002; Taguchi et al.,2004; Chong and Chang, 2006), more than one cell typeis necessary. These different cell types have obviouslydifferent environments in vivo; hence they have differentscaffold requirements, such as different pore sizes andporosities. Thus, a well-engineered TE scaffold should betailored with the appropriate pore sizes and porositiesaccording to the needs of the specific cells to betteraccommodate the proliferation and growth of the cells. Forthe regeneration of multiple tissues and tissue interfaces,functionally guided gradients in pore size and porosity willbe critical (Karageorgiou and Kaplan, 2005).

It has been found that the pore size and substrate surfacemorphology influence cell morphology and phenotypicexpression (Nehrer et al., 1997; Miot et al., 2005). Cells respondto substrate topography differently due to the differences incell size and cell-matrix adhesion mechanism (Salem et al.,2002; Sun et al., 2007). For example, fibroblasts (cell size20–50 µm), being smaller than endothelial cells (cell size60–200 µm), can bridge greater pores than endothelial cells(Salem et al., 2002). Fibroblasts can span void spaces up to200 µm (Sun et al., 2007), while endothelial cells can onlybridge pores with size 30–80 µm (Salem et al., 2002). Thus, thedimensions of scaffold pores can be exclusive and depend onthe cell type (Zeltinger et al., 2001; Beckstead et al., 2005). Thepreferred scaffold pore sizes for different cell types as foundby several of these research are summarized in Table 3.

Besides pore size requirements, generally scaffolds withhigh porosity (around 90%) facilitate better cell growth,infiltration and ECM deposition (Zeltinger et al., 2001).Moreover, by controlling the scaffold porosity and poreinterconnection, one can design suitable flow channels forthe scaffold so as to result in the appropriate transport ofnutrients and wastes and cellular signals for proper tissueregeneration (Sun et al., 2004). Such flow channels will be

Table 3 – The preferred scaffold pore sizes for differentcell types

Tissueregeneration

Cell size (µm) Preferred porediameter (µm)

Vascular 60–200 (Salemet al., 2002)

5 [for neovascularisation](Tanaka et al., 2007)

Hepatocytes 20–40 (Galarneauet al., 2007)

20 (Yang et al., 2001)

Fibroblast 20–50 (Salem et al.,2002)

90–360 (Wang et al., 2005)

Bone 20–30 (Oota et al.,2006)

100–350 (Yang et al.,2001)

Table 4 – Cells proliferation on polished and modifiedsurfaces for human gingival fibroblasts on high-purityaluminium oxide substrate

Surfacemodification

type

Averagesurface

roughness(nm)

Result after 72 h

Pressed, milled,densely sintered

90.162 ± 38.848 Significantly more cellsattached than polishedsurface (around 66% ofcell attachment)

Pressed, denselysintered

42.256 ± 9.172 Significantly more cellsattached than polishedsurface (around 66% ofcell attachment)

Pressed, denselysintered, polished

2.736 ± 0.471 Around 61% cellattachment

very useful if the cell-seeded scaffold is cultured in a suitablebioreactor to mimic the fluid flow inside the body (Martinet al., 2004).

Another important biological requirement is the surfaceproperties of the fabricated scaffold. It has been knownthat surface properties such as surface composition, surfaceroughness and hydrophilicity play an important role inregulating cell attachment, spreading and ECM deposition.It is also known that different cell types reacted differentlyto the same surface property (Singhvi et al., 1994). Ingeneral, hydrophilic surfaces are preferred and coating ascaffold using suitable protein to improve its hydrophilicitycontributed positively to cell attachment and proliferation(Zhao et al., 2003; Cheng and Teoh, 2004). Regarding scaffoldsurface roughness, it is believed that there exists a range ofpreferred surface roughness value for each cell type. Seedinghuman gingival fibroblasts on high-purity aluminium oxidesubstrate with different surface roughness, Mustafa et al.found that cells proliferate better on rough surfaces thatpolished ones (Mustafa et al., 2005). Table 4 summarizes theresults of these experiments.

2.2. Mechanical requirements

Generally, a scaffold should have sufficient mechanicalstrength to maintain integrity until the new tissue regen-erates, maintain the space for cell ingrowths and nutrienttransport in vitro (Leong et al., 2003) and support physiologi-cal loadings in vivo (Temenoff et al., 2000). The scaffold should


Table 5 – Mechanical properties of human bone (Yanget al., 2001)

Tissues Tensilestrength(MPa)

Compressivestrength(MPa)

Young’smodulus(GPa)

Cancellous bone N/a 4–12 0.02–0.5Cortical bone 60–100 130–180 3–30

match its mechanical properties to that of the native tissue toboth prevent stress shielding (Prendergast, 2001) and give thecells proper mechanical cues as the ones they normally re-ceive in their native environment (Kreke et al., 2005; Tan andTeoh, 2007).

Human body architectures have complex geometry withvarying mechanical properties distributed spatially withrespect to the anatomical site (Ford et al., 1999a). The tissuesconsist of layers having different microstructures, which canvary in terms of porosity, density and pore size. To guide thenew tissue to behave like the original tissue, some kinds ofmechanical cues are needed (Kreeger and Shea, 2002). Forexample, as a natural load-bearing structure, human bonesadapt their structure to their function to sustain a variety ofexternal loads, i.e. the collagen fibres in bone are orientatedaccording to the direction of the stresses. Such orientationsare not found in foetuses (McGowan, 1999). When observedacross its transverse section, human bone has a gradedstructure varying its pore size and porosity distribution. Theirouter layer, cortical bone, is solid and dense, while the innerlayer, cancellous bone, is a spongy honeycombed structurefilled with blood vessels and bone marrow maximizing thestrength to weight ratio for bending and compression loads(Ford et al., 1999a). As a result, bone structure has functionallygraded mechanical properties, which are shown in Table 5.

The scaffold’s mechanical properties should match thoseof the host tissue, this is particularly important for bonetissues as they are the main load-bearing tissues in thebody. For a load-bearing tissue, if the scaffold has superiormechanical properties as compared to the bone itself, thescaffold will take the load previously taken by the bone andthus shield the bone from load, causing bone resorptionat the surrounding tissue. The opposite scenario of havingan inferior scaffold will result in scaffold failure, as thescaffold does not have sufficient strength to withstand thephysiological loads (Prendergast, 2001). This is also true forsoft tissues, indicated by fibroblasts, which proliferated betterin a lower stiffness PCL membrane (0.05 ± 0.01 N/mm),which was close to their native environment (Tan and Teoh,2007). Griffith and Swartz found that ECM stiffness influencedthe degree of which the cells can contract the matrix,which was related to the cells migration and functions.The stiffness of ECM affected the response of the ECMto externally applied loads, which in turn affecting theinterstitial fluid flow, cell-cell distance and signalling and alsoECM components gradients (Griffith and Swartz, 2006). Thus,the important mechanical properties to consider include thescaffold stiffness and strength, as well as the scaffold fatigueproperty under cyclic loading experienced by the nativetissue.

Besides bone, articular cartilage regeneration is alsoknown to be influenced by these mechanical cues. Generally,articular cartilage is loaded by compression loadings invivo. Such mechanical loading affects changes at cellularlevel such as tissue deformation, fluid flow and nutrientand ion concentration gradients. Thus, compressive loadingspromote chondrocytes viability, gene expression and thesecretion of various ECM components such as collagen,proteoglycan and fibronectin in articular cartilage constructs.It was also found that fluid flow in turn affected themechanical properties of the chondrocytes cultured inscaffolds, where generally laminar flow increased themechanical properties of the construct, in contrast toturbulent flow. It appeared that the equilibrium modulusof the cartilage construct cultured in laminar flow wasabout four times the modulus of the constructs cultured inturbulent flow and static culture (Vunjak-Novakovic et al.,1999). Besides, having fluid flow inside the scaffold wouldincrease the effectiveness of nutrients and transport ofwastes (Lee et al., 2005).

To better suit its functions, the scaffold needs to bemodelled as a multilayer structure with varying mechanicalproperties across the scaffold. This can be done by changingthe porosity across the scaffold, as porosity is more dominantin determining the scaffold mechanical properties than poresize (Bignon et al., 2003). There should be no incompatibilityor separation at the interfaces of the layers. Thus a soundapproach is to mimic such bonding at the scaffold layerinterface (Ford et al., 1999a).

2.3. Anatomical requirements

Anatomically, the external geometry and size of the scaffoldshould be the same as those of the tissue defect inorder for the scaffold to fit and anchor onto the defectsite (Sun et al., 2005). With appropriate external geometryand size of the scaffold, better fixation can be achieved,while facilitating better stress distribution at the interfacebetween the surrounding tissue and the scaffold. Theanatomical requirement can be fulfilled by making useof CAD systems developed to aid scaffold design process,for example: Computer Aided System for Tissue Scaffolds(CASTS) (Naing et al., 2005) and Computer-Aided TissueEngineering (CATE) (Sun et al., 2004). The general steps inobtaining the appropriate information for designing andfabrication patient-specific scaffold are discussed in thesection on Computer-Aided Tissue Engineering.

2.4. The effects of FGS

The gradient of the structure in the scaffold to facilitatethese mechanical cues can be modeled using FGS withvarying pore size and porosity values, as these parametersdirectly affect the mechanical properties of the scaffolds.These have been confirmed in the following studies. Moutoset al. generated two different 3D woven structures using104 µm-diameter continuous multifilament PGA yarn forarticular cartilage repair. The first one, called ‘small pore’structure, had interconnected internal pores of 390 µm ×

320 µm × 104 µm with porosity of ∼70%, while the second


Fig. 2 – Fibre orientation of a woven scaffold (adapted from Moutos et al. (2007).)

one, ‘large pore’ structure, had pore size of 450 µm×320 µm×

104 µm with porosity of ∼74%. Both scaffolds possessedanisotropic compressive and shear properties similar to thenative articular cartilage. Pore size was shown to greatlyaffect the Young’s modulus of the scaffolds both in confinedand unconfined compression, where the mean values of‘small pore’ structure (199 kPa and 77 kPa, for confinedand unconfined compression respectively) were significantlyhigher that those of ‘large pore’ structure (138 kPa and68 kPa). The schematic diagram of the woven scaffold isshown in Fig. 2, in which the fibres are classified accordingto their direction in the woven scaffold. The lengthwise fibresinterlaced the fibres laid in x and y directions. Moutos calledthe fibres parallel to the lengthwise fibres warp fibres, whilethose that are perpendicular to the lengthwise fibres weftfibres (see Fig. 2) (Moutos et al., 2007).

Another work by Lin et al. used solution coating andparticulate leaching technique to fabricate porous poly(L-lactide-co-DL-lactide) (PLDL) scaffolds for cancellous boneapplications with varying porogen volume fractions. Theresulted scaffolds possessed porosity values ranged between58% and 80%. The scaffolds were tested for averagecompressive modulus and ultimate strength to correlate thescaffolds microstructure with their mechanical properties.The test results showed that the strength ranged between1 and 10 MPa while the modulus raged between 50and 250 MPa. The mechanical properties decreased withincreasing porosity linearly (Lin et al., 2003).

To tailor and closely control the culture conditions, devicessuch as bioreactors can be utilized. Cells seeded ontoscaffolds can be subjected to medium flow at an appropriateflow rate that mimicked the natural conditions to facilitatenutrients and wastes transport. Furthermore, mechanicalforces resulting from the fluid flow are included to directcell activity and phenotype. The types of bioreactors availableinclude spinner flasks, rotating vessels and direct perfusionbioreactors (Martin et al., 2004).

2.5. Spatial gradients of FGS

Generally, spatial gradients can be categorized as continuousand discrete gradients. The first type has a continuous changein microstructure with spatial position, while the latterhas multilayer structures with distinctive interfaces betweenthe layers (Ford et al., 1999b). From a mechanical point ofview, discrete gradients have mismatches, incompatibilityand stress concentration at the interfaces between the layers,

which will likely cause delamination between the layers andweaken the structure.

This can be seen in sandwich structures, where eachpart had a foam core and two high-density face sheetssandwiching the foam core. The foam core had low densitywith stiffness much lower than the face sheets. Whenthe structure was loaded, there were incompatibilitiesbetween the core and face deformations, leading to stressconcentration at the interfaces between the core and facesheets. When the concentrated strain energy at the interfacewas released, the interfaces were delaminated (Hohe andBecker, 2001).

Mahfuz et al. performed a buckling test for three types ofsandwich structures with the same face sheets but differentfoam core densities with embedded delamination at theinterface. The results showed that the structure with thehighest core density (the closest to the face sheet density)had minimal delamination propagation. This showed thatless abrupt density transition will lessen the likelihood ofdelamination (Mahfuz et al., 2004). Thus, it is much preferredthat the scaffold functional gradients should avoid discretegradient as much as possible.

3. Methods for building FGS and their me-chanical properties

There are several techniques that are able to fabricatestructures with gradients and they can again be categorizedas conventional and advanced processing methods. Someexamples of the mechanical property variations with respectto the spatial position are also illustrated.

3.1. Conventional methods

Even though conventional techniques do not have full controlof the macro- and micro-architecture of the fabricatedscaffold, some of these techniques were still able to fabricatescaffolds with porosity and pore size gradients. Some ofthese are presented in the following sections. The first twotechniques result in scaffolds that have discrete gradients,while the next two techniques are capable of fabricatingscaffolds with continuous gradients.

1. Using porogensThis method involves stacking powder mixture layers

containing different volume fractions and/or particle sizes


Fig. 3 – The variation of porosity and pore size with respect to the position in the scaffold, adapted from: (a) Schwarz andEpple (1998) (discrete gradient) and (b) Oh et al. (2007) (continuous gradient).

of volatile sintering additives such as ionic salts. Werneret al. used a technique called multiple tape casting toproduce hydroxyapatite (HA) scaffolds with porosity andpore size gradients. After the multilayer scaffold wasprepared, the poly(butylmethacrylate) (PBMA) particlesthat acted as the porogens were burnt out and sintered,resulting in pore sizes ranging from 70 µm to 200 µm(Werner et al., 2002). Schwarz and Epple fabricatedpolyglycolide scaffolds resembling natural bone with acortical (compact) and a cancellous (porous) part usingsodium chloride crystals as porogens. The salt crystalswere then washed out with water, leaving interconnectedpores with size in the micrometer range. The sizes ofthe pores were controlled by the salt particle size whilethe porosity was controlled by the proportion of the saltcrystals (Schwarz and Epple, 1998). Vaz et al. incorporatedpolyvinyl polyacrylate (PVPA) to create pores in multipleslip casting technique and fabricated HA bone implantwith porosity gradient (Vaz et al., 1999).

2. Impregnating cellulosic sponges into HA slurriesTampieri et al. made HA slurries with dissimilar

characteristics using HA powders with different degrees ofcrystallinity. Two types of porous body were obtained fromsoaking cellulose sponges into the two different slurries(slurry L from 20% crystalline powder, slurry H from80% crystalline powder). First, the sponge was completelyimpregnated in slurry H. Subsequently the same spongewas partially impregnated in slurry L. Afterwards, thesponge was left to decant and air-dried, and finally it wassintered. The resulting scaffold possessed a low-porosityzone, resulting from the additional partial impregnation inslurry L, and a more porous zone, resulting from the firstcomplete impregnation in slurry H. The overall porosityof the sample was 67% while the individual porosities oflow and high porosity regions were not reported (Tampieriet al., 2001).

3. Self-foaming followed by pyrolysisZeschky et al. obtained silicone oxycarbide (Si–O–C)

ceramic foam with porosity gradient by setting a lowerfoaming temperature to control the bubble nucleation,growth, and bubble rise. The resulting structure possessedsmaller bubbles at the top of the foam while the biggerbubbles were found at the bottom of the foam (Zeschkyet al., 2005).

4. Centrifugation methodHarley et al. fabricated collagen-based, porous tubular

scaffolds to facilitate the study of myofibroblast migrationduring peripheral nerve regeneration (Harley et al.,2006). The collagen-glycosaminoglycan (CG) suspensionin acetic acid was spun in a cylindrical mould about itslongitudinal axis to separate the components (liquid andsolid phases) and rapidly frozen to lock the componentseparation. Finally the frozen solvent phase was removedvia sublimation. The resulting scaffolds possessed radiallyaligned pore structure with pore size (<5–20 µm) andporosity gradient along the radius, where the outer portionof the tube had smaller pores compared to the innerportion (Harley et al., 2006). Oh et al. obtained PCL scaffoldwith pore size (∼88–405 µm) and porosity (∼80%–94%)gradients using centrifugation method followed by fibberbonding by heat treatment. The scaffold was fabricated toinvestigate the interaction between cells and scaffold withvarying pore size (Oh et al., 2007).

The variations of porosity and pore size with their spatialposition across the scaffold made by conventional techniquesmentioned are shown in Fig. 3. The variations of the scaffoldflexural strength and compression strength with respect totheir porosity are shown in Fig. 4, where these figures showdecreasing values in mechanical properties with increasingporosity. Fig. 4(a) shows that flexural strength of thescaffold built using multiple tape casting technique decreasesexponentially with the scaffold porosity, while Fig. 4(b) showsthat the scaffold fabricated using centrifugation method hascompression strength values that decreases linearly with itsporosity.

3.2. Rapid prototyping (RP) methods

Beside these techniques, RP fabrication techniques have alsobeen explored to produce structures with discrete (usingFDM and 3D fibre deposition) and semi-continuous (usingTheriformTM 3D printing process) pore size gradient.

1. TheriformTM

Sherwood et al. fabricated an osteochondral scaf-fold with porosity and material gradient by usingTheriformTM three-dimensional printing process (Sher-wood et al., 2002). The porosity was varied by changing the


Fig. 4 – The variation of mechanical properties with respect to scaffold porosity, adapted from: (a) Werner et al. (2002) and(b) Oh et al. (2007).

Fig. 5 – The variation of porosity and pore size with respect to the position in the scaffold, adapted from: (a) Sherwood et al.(2002) (front view) and (b) Kalita et al. (2003) (top view, concentric rings diameters: 1, 2.125 and 3.125 in., porosity values notavailable).

amount of the porogen (NaCl) spatially across the scaffold.The porogens were then leached with water. The upper re-gion was the cartilage region and was made of PLGA/ PLA,while the lower region was bone region and was made ofPLGA/tri-calcium phosphate (PLGA/TCP) to promote boneingrowth. The difference in the porosity of the cartilage(90%) and bone region (55%) was meant to prevent thechondrocytes seeded in the cartilage region from migrat-ing towards the bone region of the scaffold and also togive the bone section more mechanical strength. A largerpore size was used in the bone region (>125 µm) to facili-tate the mineralized bone ingrowth. Between the cartilageand bone region, there existed a gradual transition of thematerial and porosity values (85%, 75%, and 65%) to over-come delamination at the interface of the two regions dueto shrinkage (refer to Fig. 5(a)) (Sherwood et al., 2002).

2. FDMKalita et al. used FDM to build several concentric

cylinders with different pore sizes radially from segmentto segment (Kalita et al., 2003). The material used was poly(propylene)/tricalcium phosphate (PP/TCP). The resultingscaffold had different discrete rings of constant porosityin the radial direction, and was directed for bone implantsespecially for cancellous bone grafts. Gradient porosity,pore size and pore shape were fabricated by varying theroad gaps and lay-down patterns for each concentric ring(Kalita et al., 2003). The scaffold shown in Fig. 5(b) wasfabricated using constant lay-down pattern of 0◦–60◦–120◦

with varying the road gap (2.03 mm at the centre, 1.27 mmin the middle section and 0.51 mm at the outer ring)and had decreasing pore size and porosity towards the

outer section. When the structure was closely observed,there are discontinuities of the structure at the interfacebetween two concentric rings, which can likely affect theinterconnections of the pores from one region to the otheras these are not well defined. They can also obstruct cellsinfiltration towards the centre of the scaffold (refer toFig. 5(b)).

3. 3D Fibre deposition (3DF)Woodfield et al. fabricated a graded structure for

articular cartilage application using 3DF by varying thefibre spacing. Unlike FDM, the input material for 3DF isin a pellet form rather than a filament. The fabricatedscaffolds have 100% interconnected pores with controlledmicrostructure and mechanical properties similar to thatof native articular cartilage (Woodfield et al., 2004). Thestudy was followed with in-vitro cell seeding. It wasfound that pore size gradient promoted anisotropic celldistribution imitating those of natural bovine articularcartilage. By controlling the cells distribution it was foundthat the zonal ECM production was affected to a certainextent to conform to those of native bovine articularcartilage, which in turn will influence the biomechanicalfunction of the cartilage (Woodfield et al., 2005).

Schematic models of RP-fabricated FGS are shown inFig. 5 to illustrate the variation of porosities and pore sizeswith respect to their spatial position across the scaffold.The variations of the compression modulus with respect toscaffold porosity are shown in Fig. 6. As seen from Fig. 6(a)and (b), the compressive modulus of the scaffold fabricatedusing Theriform and FDM, as expected, decreased linearlywith increasing porosity.


Fig. 6 – The variation of mechanical properties with respect to scaffold porosity adapted from: (a) Sherwood et al. (2002)and (b) Kalita et al. (2003).

Table 6 – List of biomaterials used in RP fabrication of TE scaffolds and the achieved mechanical properties

RP process Biomaterials Range of mechanical properties

3D printing (3DP) Cornstarch, dextran and gelatin Compressive modulus: 0.059–0.102 MPa (Lamet al., 2002)

HA (3DP followed by sintering) Compressive strength: up to 22 MPa (Seitz et al.,2005)

Poly(urethane) (PU) Young’s modulus: 580 MPa (Pfister et al., 2004)

Selective laser sintering (SLS) Poly(caprolactone) (PCL) Compressive modulus: 52–67 MPa (Williams et al.,2005)

PCL-HA Compressive modulus: 33–102 MPa (Wiria et al.,2007)

Fused deposition modelling (FDM) PCL Compressive modulus: 4–77 MPa (Zein et al., 2002)PCL Compressive strength: 2.4–20.2 MPa (Hutmacher

et al., 2001)Alumina (indirect fabrication) Compressive strength: 50 MPa (Leong et al., 2003)

3D fibre deposition (3DF) Poly(ethyleneglycol)-terephthalate–poly(butylenesterephthalate) (PEGT–PBT)

Dynamic compressive modulus: 4.33 MPa (Maldaet al., 2005)

As mentioned in the previous section, discrete gradientscan severely affect the interconnection between pores ofone region to the pores of the neighbouring region atthe interface of the two regions. This can severely hindercellular infiltration and migration into the deeper region ofthe scaffold. The discontinuities at the scaffold interfacecould also negatively influence the fluid flow between thetwo regions, which will in turn affect the nutrients andwastes transport in and out of the scaffold. Moreover, abruptstress transfer can occur at the interface and can result instress concentrations at specific locations, both of which willweaken the scaffold and cause delamination at the interfacebetween two regions. As such, there is a need for a strategyto design and build scaffolds with appropriate continuousfunctional gradients to support tissue regeneration.

Owing to the repeatability and control over the scaffoldmacro- and micro-architecture, RP methods can be mean-ingfully exploited to fabricate such complex continuous gra-dient FGS. To better assess the viability of RP processes, alist of biomaterials that have been used in the related RPprocesses are presented, together with the range of me-chanical properties that can be achieved and their respec-tive degradation time. Table 6 lists the biomaterials usuallyused in RP techniques that include both direct and indirectmethods. They ranged from natural polymers such as corn-starch, dextran and gelatin that can reach a compressive

modulus range of 0.059–0.102 MPa, to synthetic polymers likepoly(caprolactone) and poly(urethane) with much higher val-ues in mechanical properties, also the bioceramics and bio-composites.

4. Computer-aided tissue engineering andFGS

To build better TE scaffolds for the human body, two ofthe major challenges are to customize scaffold pore sizeand porosity to suit different types of cells and develop thescaffold’s biological, mechanical and anatomical propertiesmimicking the native tissue (Sun et al., 2005; Fang et al., 2005).Furthermore, the external geometry of the scaffold has toanchor and fit into the defect site as it will determine thestress distribution at the organ-implant interface and also thesurrounding tissue (Chua et al., 2003a,b). These challengesmake the design stage ever so complex. Several researchershave proposed the use of computer-aided tissue engineeringto ease the design and control of the scaffold macro- andmicro-structures. Several notable works on computer-aidedtissue engineering including the works by Hollister et al.(2000), Chua et al. (2003a,b); Cheah et al. (2004), Wettergreenet al. (2005a,b) and Naing et al. (2005) are presented in thissection.


The general approach of computer-aided tissue engineer-ing can be described as follows (Sun and Lal, 2002; Naing et al.,2005):

1. Acquiring the defect image from medical imagingtechnologies like CT scan or MRI to obtain the appropriategeometry and form of the defect to be repaired orregenerated.

2. From a library of basic shapes of unit cells for the scaffoldinternal architecture the appropriate unit cell shape ischosen for regeneration according to the needs of thedefect site.

3. Stacking up or stitching together the cellular units toconstruct a block larger than the defect size for Booleanoperation.

4. Performing Boolean operation between the defect imageand the stacked or stitched cellular units after the defectimage is placed over the arranged stack of cellular units togenerate the form and architecture of the scaffold.

Hollister et al. presented an image-based approach todesign TE scaffolds using voxels, which are the smallestunit volume of a computer image. The approach worked bysetting the voxels within a unit cube to be either 0 for nomaterial (hollow) or 1 for material (solid). Depending on thevoxel setting, both regular and random pores within the unitcube can be created. The unit cube was repeated afterwards(Hollister et al., 2000).

Wettergreen et al. created a library of polyhedral units,which exhibited symmetry along the three orthogonal axesand thus were considered as orthotropic. Two types ofpolyhedral units were created. The first one was spacefilling solid structures such as spheres and cubes with voidssuperimposed on them by Boolean operation to be createdas pores. The second type was a wireframe approximationof basic polyhedral shapes. To assist the mating betweentwo different polyhedral units, a common mating featureof a torus shape was accommodated into each polyhedralunit (Wettergreen et al., 2005b). The system was used toconstruct scaffold replacement for the entire vertebral body(Wettergreen et al., 2005a).

Chua et al. investigated suitable polyhedral shapes foruse in scaffold modelling in order to create a parametriclibrary of scaffold structures (Chua et al., 2003a,b). Cheahet al. continued the work by developing a novel algorithmto create internal architecture of TE scaffolds by assemblingopen polyhedral cells into a scaffold structure (Cheah et al.,2004). This work was further refined by Naing et al. and thesystem is called Computer-Aided System for Tissue Scaffolds(CASTS). Thirteen configurations of polyhedral unit cellswere available to be assembled into customized scaffoldswith the incorporation of calculation of important scaffoldparameters, such as pore size, porosity and surface area-to-volume ratio (Naing et al., 2005).

Computer-aided TE systems have greatly eased thedesign of uniformly structured TE scaffolds with complexrequirements, i.e. biological, mechanical and structuralneeds. Various types of scaffold structure library have beendeveloped, each having unique cellular unit structures,such as voxel-based, solid-based and wireframe-polyhedral-based unit cells. Scaffolds with controllable and repeatable

microstructures can be fabricated based on user’s input andrequirements. However, in terms of regenerating multipletissues with interfaces such as bone-cartilage tissue, the CADsystems mentioned have not been able to generate a suitableFGS design as the relationships between the complex tissuerequirements, including those of biological and mechanicalneeds, and the scaffold structural gradients are not yet wellestablished. Nevertheless, early efforts in the direction of FGSfabrication have been made. One example is the scaffoldlibrary constructed by Wettergreen et al. (2005b), in which acommon connector feature of a torus shape was added atevery sides/facets of each cellular unit. This feature actedas a bridge to allow a smooth change between two differentcellular units.

Preferably, after the relationships between scaffoldsrequirements, i.e. biological, mechanical and anatomicalrequirements and the scaffold physical characteristics havebeen well established, such relationships can be incorporatedinto the CAD systems in order to help tissue engineersdesign a customized FGS that fulfil the requirementsmentioned. This would be the likely direction of CAD aidedTE systems development in near future. The application ofmedical imagining systems and techniques together withCAD systems have pretty much dealt with the problemof matching anatomical requirements of the TE scaffolds.The next task to be carried out would be to establish therelationships of both mechanical and biological requirementsand the scaffold structural parameters such as pore size andporosity. The likely major controlling factor would be thefunctions of the tissues and cells at the local and cellularlevel. Thereafter, conflicting requirements, i.e. high porosity isdesired for cell proliferation (biological need) but detrimentalfor scaffold strength and stiffness (mechanical need), canbe resolved by finding the optimized scaffold structuresuch as pore size, porosity and the gradient to better suitboth biological and mechanical requirements. The difficultchallenge in matching these relationships to the physicalcharacteristics of the TE scaffolds can be met with developingthe understanding the different functional requirements ofthe tissues and organs themselves. Mimicking the physicalenvironment may be a first step and further experimentswill be needed to study the cellular behaviour and responseto these physical parameters. Much of this information isavailable in one form or another in literature, but there is aneed to assemble these bits of information in a meaningfulway such that a good model can be established upon whicha good computer-aided tissue engineering system can bedeveloped.

5. Conclusion

Close observations of natural tissues and organs shows thatthey are not homogeneous and there exist natural functionalgradients in their structure. Many TE research literatureshave shown that there is a necessity in developing scaffoldswith continuous functional gradients in regenerating tissues,particularly multiple cell types with tissue interfaces tomimic the complete organ. The generation of such scaffoldscan substantially be aided by CAD system coupled with


an automated fabrication system to ease the tedious andcomplex scaffold design stage, as well as addressing thechallenges of repeatability and reproducibility. However, thesuitable and meaningful model to represent continuous FGSthat can fulfil the biological and mechanical requirementsof the regenerated tissue is needed. This model can thenbe incorporated into the CAD systems in order to automatethe design of such FGS. Therefore, the next directionfor CAD aided TE will be first to establish a relationamong the scaffolds biological, mechanical and anatomicalrequirements to be implemented into the available CADsystems.

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Engineering functionally graded tissue engineering scaffolds