Chapter 2 literature review design and development of tissue engineering scafflods using rapid prototyping technology

Therefore, the modern tissue engineering approach utilizes three-dimensional porous scaffolds made of natural or synthetic polymers which provide temporary substrate for cell attachment,

Chapter Two

Literature Review

2.1 Tissue Engineering Scaffolds

By far tissue engineering strategy that involves cell transplantation has shown high potential in treating damaged or malfunctioning organs Because of the fact that many cell types are anchorage-dependent, their direct transplantation

in the recipient’s body might result in death or loss of function and thus require the presence of a substrate It has also been observed that dissociated cells tend to organize themselves to form a tissue structure when they are provided with a proper guiding template (Vacanti et al, 1998) Therefore, the modern tissue engineering approach utilizes three-dimensional porous scaffolds made

of natural or synthetic polymers which provide temporary substrate for cell attachment, proliferation and function

The key parameters in designing a suitable scaffold for tissue engineering applications include the material properties and the macro/micro structure of the scaffolds (i.e porosity and pore morphology) An interconnected internal structure of the scaffold is important for adequate flow of nutrients into and transport of metabolic waste out of the structure (Langer and Vacanti, 1993) The scaffold materials must be biocompatible which means they must not trigger any adverse reactions with tissues Surface properties of the scaffold are also important for proper attachment of the cells onto the structure Often, additives such as hydroxyapatite (HA) are added to the basic scaffolding material to promote cell attachment In case of tissues that are subjected to stress and strain, e.g arteries, heart valves, bones etc the scaffold matrix must provide sufficient mechanical strength to withstand in vivo stresses and loading The mechanical properties of the bioresorbable 3D scaffold/tissue construct at

the time of implantation should match that of the host tissue as closely as possible Lastly, the scaffold material is expected to be reabsorbed by the tissue after the cells have established themselves Hence, the scaffold material should essentially be selected and/or designed with a controlled degradation and resorption rate such that the strength of the scaffold is retained until the tissue engineered transplant is fully accommodated by the host tissue and can assume its structural role (Stephen et al, 1998)

2.2 Scaffold Materials

One of the fundamental issues with regard to tissue engineering is the choice of suitable material Currently, polymeric materials have drawn great attention from the scientific and medical communities for tissue engineering applications (Maquet et al, 1997) Natural polymers such as collagens, glycosaminoglycan, starch, chitin and chitosan have been used to repair nerves, skin, cartilage, and bone (Mano et al, 1999) These naturally occurring biomaterials might most closely simulate the native cellular milieu However, large batch-to-batch variation upon isolation from biological tissues and availability are the main limitations for their wide applications Poor mechanical performance is also a drawback for transplanted scaffolds made from natural polymers On top of that natural polymers such as collagen and gylcosaminoglycan could also provoke adverse tissue reactions and immune responses

Synthetic polymers have been developed to overcome the aforementioned problems associated with natural polymers Synthetic polymers are well known for their enormous availability, high processability, and controllable mechanical

and biochemical properties Most synthetic polymers degrade via chemical hydrolysis and insensitive to enzymatic processes so that their degradation behaviours do not vary from patient to patient Many synthetic bioresorbable polymers such as poly (α-hydroxy ester)s, polyanhydrides, polyorthoesters, and polyphosphazens, have been studied for temporary surgical and pharmacological applications (Vert et al, 1992; Pitt et al, 1981a) These polymers have been found to be suitable to construct bioresorbable 3D scaffolds for tissue engineering applications Properties of different synthetic polymers are summarized in Table 2.1

Table 2.1: Properties of Biodegradable Polymers (Shalaby et al, 1994; Maquet

et al, 1997; Perrin and English, 1997; Middleton et al, 1998; Ali and Hamid, 1998; Huang Ming-Hsi et al, 2003; http://www.physics.iisc.ernet.in)

Polymer

Type

Melting Point (°C)

Glass Transi tion Temp (°C)

Degra dation Time (months) a

Density (g/cc)

Tensile Strength (MPA)

Elonga tion %

Modu lus (GPA)

phous

55-60 12-16 1.25 27.6-41.4 3-10 1.4-2.8 L-PLA 173-178 60-65 >24 1.24 55.2-82.7 5-10 2.8-4.2 PGA 185-225 25-65 6-12 1.53 >68.9 15-20 >6.9 PCL 58-68 - 70 >24 1.11 20.7-34.5 300-500 0.21-0.34

2.2.1 Poly( ε-Caprolactone)

Poly (ε-caprolactone) is one of the earliest polymers synthesised by the Carothers group in the early 1930s (van Natta et al, 1934) Commercially it became available following efforts to identify synthetic polymers that could be degraded by micro-organisms Poly (caprolactone) can be prepared by either ring-opening polymerisation of caprolactone using a variety of anionic, cationic and coordination catalysts or via free radical ring-opening polymerisation of 2-methylene-1-3-dioxepane Poly (caprolactone) is a semicrystalline polymer This semi-crystalline, linear aliphatic polyester has a repeating molecular unit of five non-polar methylene groups and a single relatively polar ester group (Figure 2.1) Its crystallinity tends to decrease with increasing molecular weight Degradation occurs largely due to the presence of the hydrolytically unstable aliphatic-ester linkages

Figure 2.1: Repeating molecular structure of PCL

The high solubility of poly (caprolactone), its low melting point (59 to 64oC) and exceptional ability to form blends has stimulated research on its application as a biomaterial In 1981 Pitt and co-workers (Pitt et al, 1981b) first reported an in vivo study of PCL drug-delivery capsules in a rabbit model It was observed that the polymer degraded in a two-phase process, with the majority of molecular weight loss occurring primarily in the first phase, and the subsequent

loss in mass and strength beginning at the onset of the second phase at an average molecular weight of 5000

PCL was susceptible to both enzymatic and non-enzymatic degradation Woodward and co-workers (Woodward et al, 1985) studied the intracellular degradation of low molecular weight (Mn 3000) PCL powders (106 to 500 µm)

in rats They reported that the PCL powders were hydrolytically degraded in phagosomes secreted by macrophages and giant cells Their studies suggested that in an in-vivo environment, enzyme-mediated intracellular degradation might be the principal pathway of degradation once the polymer was sufficiently pre-degraded by earlier non-enzymatic bulk hydrolysis Degradation of PCL preceded by random hydrolytic chain scission of the ester linkages, eventually producing the monomeric hydroxyacid Pitt (1992) also reported that in rat PCL was metabolized to ε-hydroxycaproic acid, the end product of ester hydrolysis in vivo The hydroxyacid was respired and broken down to CO2 and H2O when exposed to tissue fluids (Pitt et al, 1979)

Poly (caprolactone) has slow degradation and resorption kinetics and can therefore be used in drug delivery devices that remain active for over one year The toxicology of PCL had been extensively studied as part of the evaluation of CapronorTM, a one-year implantable subdermal contraceptive device as reported by Darney et al (1989) and Ory et al (1983) Based on these clinical products, ε-caprolactone and PCL were regarded as non-toxic and hard and soft tissue compatible materials Capronor had undergone FDA-approved phase I and II clinical trials (Pitt, 1990) Extensive in-vitro and in-vivo

biocompatibility and efficacy studies had also been performed in the research leading to the introduction of the MonocrylTM monofilament sutures as reported

Among the bioresorbable polymers used for biomedical applications, PCL has

an unusually low glass transition temperature (Tg) of – 65°C It also has a low melting temperature of 59-64°C and exists in a rubbery state at room temperature Another unusual property of PCL is its high thermal stability It has a much higher decomposition temperature (Td) of 350°C, in compared to other tested aliphatic polyesters that have decomposition temperatures (Td) between 235 and 255°C (Suggs and Mikos, 1996) Solid PCL also exhibits moderate mechanical properties as shown in Table 2.1 In comparison to other commercially available bioresorbable polymers, PCL is one of the most flexible and easy to process materials Even though it has one of the slowest degradation rates of all such polymers, the structural stability of PCL permits the study of fabrication and characterization as a tissue engineering scaffold

2.2.2 Copolymers

Presently, PCL is regarded as a soft- and hard-tissue compatible biodegradable material and often selected as a suitable material for thermoplastic processing of scaffolds for tissue engineering (Perrin and English, 1998a) The first generation of bioresorbable scaffolds for tissue-engineering applications has been fabricated from synthetic polymers of the aliphatic polyester family (Hutmacher et al, 2000a; Vats et al, 2003) However, the number of such bioresorbable polymers is limited when polymers with different properties are needed for the design and fabrication of devices and scaffolds adapted to specific applications (Hutmacher, 2001a; Saltzman, 1999) Polymers such as poly(ethylene glycol) (PEG), poly(ε-caprolactone) (PCL) and poly(DL-lactide) (P(DL)LA) have been used to make in vivo degradable medical and drug-delivery devices with Food and Drug Administration approval (Pitt, 1990; Li and Vert, 1999) Polyester–polyether block co-polymers composed of PCL or PLA and PEG have been considered

as suitable because they offer possibilities to vary the ratio of hydrophobic/hydrophilic constituents by copolymerization and to modulate degradability and hydrophilicity of corresponding matrices and surfaces (Rashkov et al, 1996; Li et al, 2002) Despite of favorable rheological properties and thermal stability in molten state, PCL degrades very slowly due to its high hydrophobicity and crystallinity (Pitt, 1990; Moore and Saunders, 1997) Introduction of hydrophilic blocks and/or fast degrading blocks into PCL main chains can be a means to prepare novel degradable and bioresorbable polymers Hydrophilic polyether blocks such as poly(ethylene glycol) (PEG) are introduced into PCL chains to enhance the hydrophilicity of the parent PCL

homo-polymer (Li et al, 2002; Lee et al, 2001) Likewise, block co-polymerization of PCL with faster degrading polyesters, such as poly(lactide) (PLA), allows to modify the degradability of the parent PCL homo-polymer (Feng et al, 1983; Deng et al, 1997) However, both types of co-polymers present specific disadvantages PLA is a hydrophobic polymer, whereas PEG is hydrophilic but not degradable in vivo Therefore, PCL-based co-polymer (PEG-PCL-PLA) was synthesized by combining both PEG and PLA blocks with PCL chains to produce novel hydrophilic and bioresorbable co-polymer with the aim of enhancing hydrophilicity and degradability

Similarly, if ε-caprolactone is copolymerised with ethylene oxide (EO) or poly(ethylene glycol) (PEG) to prepare PCL/PEG(PEO) block copolymers, their physical property, hydrophilicity and biodegradability can also be improved, and thus they may find much wider applications Recently, several research groups (Nagasaki et al, 1998; Li et al, 1996; Kricheldorf et al, 1993; Yuan et al, 2000; Dobrzynski et al, 1999; Longhai et al, 2003; Huang et al, 2004) prepared bioresorbable polyester–PEG diblock or triblock copolymers by using a monohydroxy or α,ω-dihydroxy PEG as initiator for the polymerization of lactone monomers employing various techniques Longhai et al (2003) synthesized and characterized the PCL-PEG-PCL triblock copolymers by ring-opening polymerization of ε-caprolactone (CL) in the presence of poly(ethylene glycol) using calcium catalyst The differential scanning calorimetry and wide-angle X-ray diffraction analyses revealed the micro-domain structure in the copolymer The melting temperature, Tm and crystallization temperature, Tc of the PEG domain were observed to be

influenced by the relative length of the PCL blocks They mentioned that it was because of the strong covalent interconnection between the two domains Huang et al (2004) performed degradation and cell culture studies on PCL homopolymer and PCL/PEG diblock and triblock copolymers prepared by ring-opening polymerization of ε-caprolactone in the presence of ethylene glycol or PEG using zinc metal as catalyst They performed the degradation of PCL and PCL/PEG diblock and triblock copolymers in a 0.13 M, pH 7.4 phosphate buffer at 370C The results indicated that the copolymers exhibited higher hydrophilicity and degradability compared to the PCL homopolymer They cultured primary human and rat bone marrow derived stromal cells (hMSC, rMSC) on the scaffolds manufactured with PCL homopolymer and PCL/PEG diblock and triblock copolymers via solid free form fabrication Light, scanning electron and confocal laser microscopy as well as immunocytochemistry showed cell attachment, proliferation and extracellular matrix production on the surfaces along with inside the scaffold architectures of all polymers However, the copolymers showed better performance in the cell culture studies than the PCL homopolymer

Some other researchers investigated the block-copolymers poly(ethylene glycol)-terephthalate/poly(butylene terephthalate) (PEGT/PBT) and polyethyleneoxide-terephtalate /polybutylene-terephtelate ((PEOT/PBT) to process into 3D scaffolds that can modulate their viscoelastic properties in order to mimic a large collection of natural tissues (Woodfield et al, 2004; Moroni et al, 2006) These polyether-ester multiblock copolymers belong to a class of materials known as thermoplastic elastomers that exhibit good physical

properties like elasticity, toughness and strength in combination with easy processability (Bezemer et al, 1999) This family of copolymers has drawn great attention for tissue engineering and drug delivery applications, because

by varying the molecular weight of the starting poly(ethylene glycol) (PEG) segments and the weight ratio of PEOT and PBT blocks it is possible to tailor-make properties, such as wettability (Olde et al, 2003), swelling (Bezemer et al, 1999; van Dijkhuizen-Radersma et al, 2002; Deschamps et al, 2002) [26,28,29], biodegradation rate (Deschamps et al, 2002), protein adsorption (Mahmood et al, 2004) and mechanical properties (Woodfield et al, 2004) Furthermore, PEOT/PBT block copolymers have shown to be extensively biocompatible both in vitro and in vivo (van Blitterswijk et al, 1993; Beumer et al, 1994) and reached clinical applications (PolyActiveTM, IsoTis Orthopaedics S.A.) as cement stoppers and bone fillers in orthopedic surgery (Mensik et al, 2002; Bulstra et al, 1996) Being polyether-esters, degradation occurs in aqueous media by hydrolysis and oxidation, the rate of which varying from very slow for high PBT contents to medium and fast for larger contents of PEOT and longer PEO segments (Bezemer et al, 1999; Deschamps et al, 2002)

2.3 Structures and Properties

The second important issue to be addressed in tissue engineering is the macro- and microstructures of the scaffolds From materials engineering point of view, tissues are considered to be cellular composites representing multiphase systems Cellular composites are then seen as consisting of three main structural components: (1) cells that are organized into functional units, (2)

extracellular matrix, and (3) scaffold architecture This architecture is increasingly believed to contribute significantly to the development of specific biological functions in tissues and thought to provide appropriate nutritional conditions and spatial organization for cell growth The regeneration of specific tissues aided by synthetic materials has been shown to be dependent on the porosity and pore size of the supporting three-dimensional structure (Cima et

al, 1991) A large surface area favours cell attachment and growth, whereas a large pore volume is required to accommodate and subsequently deliver a cell mass sufficient for tissue repair Clinically approved synthetic bioresorbable polymers were primarily considered as the material of choice to build such porous three-dimensional (3D) scaffolds It is important to recognize some of the salient features of porous solids to understand how a scaffold design could affect its physical and mechanical properties In tissue engineering applications, scaffolds require a balance of high inter-connectivity of pores (3D internal architecture) with overall structural stability (mechanical characteristics) Porous solids consisted of different types of natural and synthetic materials are generally classified as honeycombs and foams (Gibson and Ashby, 1997)

2.3.1 Honeycombs and Foams

A honeycomb consists of a regular two-dimensional array of polygonal pores each defined by a wall shared between adjacent pores The pores are packed

to fill a plane area like the hexagonal cells of the bee’s honeycomb Figures 2.2a–d show the structural plan-views of synthetic honeycombs with hexagonal, triangular and square pores, otherwise known as two-dimensional porous materials On the other hand, foam has a structure in which polyhedral

pores are surrounded by faces or edges, and are packed in 3D to fill space Figures 2.2e – f illustrate both polymeric foams, termed as three-dimensional porous materials, with open and closed pores

According to the ASTM terminology (American Standard, 1999) pores are classified into three groups: interconnecting (open pores), non-connecting (closed pores), or a combination of both When the pores are open, the foam material is usually drawn into struts forming the pore edges (Figure 2.2e) A network of struts produces a low-density solid with pores connecting to each other through open faces When the pores are closed, a network of interconnected plates produces a higher-density solid The virtually closed pores are sealed off from adjacent neighbours (Figure 2.2f)

Figure 2.2: Porous materials: (a) 2D aluminium honeycomb, (b) 2D paper-phenolic resin honeycomb, (c) 2D ceramic honeycomb with square pores, (d) 2D ceramic honeycomb with triangular pores, (e) 3D open-pore polyurethane, (f) 3D closed-pore polyethylene (Gibson and Ashby, 1997)

2.3.2 Mechanical Properties

In tissue engineering applications, porous scaffolds must have sufficient mechanical strength to retain their initial structures after implantation, particularly in the reconstruction of hard, load-bearing tissues, such as bones and cartilages The structure must not collapse or compress together under

pressure, causing damage to the cells within The biostability of many implants depends on factors such as strength, elasticity, absorption at the material interface and chemical degradation Therefore, the investigation of compressive properties is of primary importance in determining the suitability of the designed scaffold Other mechanical properties like, tensile or flexural properties are secondary to the compressive properties for some basic reasons Firstly, it should be considered that at the end the scaffolds would be utilized in the physiological environment where the primary loading is compressive e.g bone or cartilage Secondly, the presence of numerous tiny voids in the porous solid poses significant structural flaws to magnify the effect

of crack propagation in stretching or bending In addition, the specimen specification for tensile (dog bone shape) or flexural test is also a limiting factor

as it often remains tedious to prepare the scaffold samples accordingly Therefore, the majority of research findings on tissue engineering scaffolds had been focused on their compressive properties when reporting on scaffolds mechanical properties Most formulations, as reviewed by Gibson and Ashby (1997), found that the mechanical properties of a porous solid depended mainly

on its relative density, the properties of the material that made up the pore edges or walls and the anisotropic nature, if any, of the solid Figure 2.3 represents the model of a honeycomb and an open-pore foam

Figure 2.3: (a) Honeycomb with prismatic hexagonal pores (b) Cubic model for

an open-pore foam showing the edge length l and the edge thickness t (Gibson and Ashby, 1997)

The deformation behaviours of porous solids under compressive loads for honeycombs are shown in Figures 2.4 and 2.5 The mechanical properties of honeycombs are classified in two groups: in-plane properties and out-of-plane properties The in-plane properties are those relating to loads applied in the X1– X2 plane Responses to loads applied to the faces normal to X3 are referred

(b) (a)

Figure 2.5: A schematic diagram for a honeycomb loaded in compression, showing linear elastic, collapse and densification regimes, and the way the stress-strain curves changes with t/l; (a) compressed in X1 – X2 plane, (b) compressed in axial (X3) direction (Gibson and Ashby, 1997)

When a honeycomb is compressed in-plane, the pore walls at first bend (Figure 2.4a), giving linear elastic deformation (shown on stress-strain curves in Figure 2.5a) Beyond a critical strain the pores collapse by elastic buckling, plastic yielding (Figure 2.4b), creep or brittle fracture, depending on the nature of the pore wall material Pore collapse ends once the opposing pore walls begin to touch each other; as the pores close up the structure is densified and its stiffness increases rapidly On loading the honeycomb out-of-plane, the pore walls experience compression under both axial and bending stresses The moduli and collapse stresses are much larger Figure 2.5b shows the family of curves of honeycombs with different relative density (∝ t/l), compressed out-of-plane

The deformation behaviours of porous solids under compressive loads for open-pore foams are shown in Figures 2.6 Like honeycombs, foams (Figure

2.6 a & b) also show linear elasticity (Figure2.6c) at low stresses followed by a long collapse plateau during which the pore edges buckle (Figure 2.6b) This

is truncated by a final regime of densification in which the stress rises steeply when the pores are completely collapsed Experimental results, checked by using honeycombs (Patel and Finnie, 1970; Abd El-Sayed et al, 1979; Warren and Kraynik, 1987) and foams (Patel and Finnie, 1970; Menges and Knipschild, 1975; Warren and Kraynik, 1988) with a wide range of materials, structures and density, have been reported to produce good agreements in general

Figure 2.6: Deformation behaviour of an open-pore foam; (a) Pore edge bending, (b) Pore edge buckling, (c) Schematic compressive stress-strain curve, showing the three regimes of linear elasticity, collapse and densification (Gibson and Ashby, 1997)

2.4 Scaffold Fabrication Techniques

2.4.1 Basic Requirements

Besides the material issues (non-mutagenic, non-antigenic, non-carcinogenic, non-toxic, non-teratogenic), it is of major importance in scaffold production to maintain sufficient accuracy over the macro- (e.g., spatial form, mechanical strength, density, porosity) and microstructural (e.g., pore size, pore distribution, pore interconnectivity) properties A large variety of natural or

synthetic scaffolding materials have been developed depending on the targeted tissue Each scaffolding material or combination of materials possesses different processing requirements and varying degrees of processability to form scaffolds The key requirements necessary to assess a fabrication technique for scaffold production include the followings (Leong et al, 2003a):

Processing Conditions: The material properties of the scaffolds should not be

adversely affected by the material processing procedures and conditions This means the technique should not change the chemical properties and biocompatibility of the scaffold nor cause any deterioration in its mechanical properties

Process Accuracy: The technique should produce spatially and anatomically

accurate three-dimensional scaffolds that fit the intended spaces at the implant site The capability to vary and maintain accurate pore sizes and morphologies will enable a wide variety of scaffolds to be produced to suit different targeted

TE applications Accuracy in construction of scaffolds will enable the application of computer-aided engineering (CAE) methods to perform strength and degradation analyses to predict the scaffolds’ performance By this means, optimized scaffold designs can be realized with minimal experimentation

Consistency: The technique should be able to produce scaffold with highly

consistent pore sizes with a narrow range of size distribution over the entire volume Consistency in pore size, shape, distribution, density and interconnectivity in all three dimensions is required to produce highly regular three-dimensional structures

Repeatability: The same set of processing parameters and conditions should

exhibit minimal variations in physical forms and properties of different scaffold batches The technique should also allow achieving highly consistent and reproducible results without encountering any major difficulty

2.4.2 Limitations of Conventional Techniques

Conventional scaffold fabrication techniques have been developed mainly on the basis of textile and polymer processing technologies These techniques include non-weaving (Ma and Langer, 1995), fibre bonding (Wang et al, 1993; Brauker et al, 1995), phase separation (Lo et al, 1995; Ma and Zhang, 1999), solvent casting/particulate leaching (Mikos et al, 1993b; Holy et al, 2000), membrane lamination (Mikos et al, 1996), melt moulding (Thomson et al, 1995a), gas foaming/high pressure processing (Baldwin et al, 1995; Mooney et

al 1996), hydrocarbon templating (Shastri et al, 2000), freeze drying (Whang et

al, 1995; Healy et al, 1998) and combinations of these techniques (e.g., gas foaming/particulate leaching (Harris et al, 1998), etc.) The working principles, procedures, applications and potentials of these techniques can be found in several other works (Vacanti et al, 1988; Widmer and Mikos, 1998; Thomson et

al, 2000; Yang et al, 2001; Sachlos and Czernuszka, 2003) To date conventional techniques have shown great promise in the scaffold fabrication and a wide range of scaffold characteristics, such as porosity, pore size etc have been reported (Widmer and Mikos, 1998; Hutmacher, 2000b) However, these techniques remain impractical to manufacture scaffolds as required because of a number of limitations They are as follows (Leong et al, 2003a):

Manual Intervention: All conventional techniques involve multi-stage manual

processes that are labour-intensive and time-consuming For example, particulate leaching involves mixing of salt(s) with the scaffold material, casting the object and further dissolving the salt(s) to produce porous scaffold The system heavily relies on user skills and experiences and thus often results in non-uniformity and poor repeatability of the scaffold architectures and properties

Reproducibility of Processing Procedures: Conventional techniques are

unable to precisely control the pore size, pore geometry and spatial distribution

of pores which results in inconsistent macro- and micro-structure of the scaffolds Scaffolds produced by solvent-casting and/or particulate-leaching cannot guarantee interconnection of pores because this is dependent on whether the adjacent salt particles are in contact

Use of Toxic Solvents: Most conventional techniques require extensive use of

toxic organic solvents to dissolve the raw stock (granules, pellets or powders) and convert into the final scaffold Thereupon, it becomes difficult to remove the toxic solvents completely from the fabricated scaffolds especially, in thicker constructs The residual toxic solvents cause adverse effects on adherent cells, incorporated biological active agents or nearby tissues (Healy et al, 1998)

Use of Porogens: Some techniques (e.g., particulate leaching, hydrocarbon

templating, etc.) utilize salts or waxes as porogens to create porosities in the scaffolds The use of porogens limits the scaffold thickness to approximately

2mm (Lu and Mikos, 1996) because of the problems in complete removal of porogens In addition, it becomes difficult to prevent the agglomeration of porogen particles and thus to achieve uniform porogen dispersion This phenomenon results in uneven pore size and densities, and morphologies of the scaffolds which give rise to anisotropy in scaffold properties (Hutmacher et

al, 2001b)

Shape Limitation: Some of the techniques use moulds or containers to

manufacture scaffold as thin membranes or three-dimensional constructs These techniques are confined to create certain simple shapes and cannot produce scaffolds with complex and desired structural architectures

Limited Cell Growth: Conventional techniques produce scaffolds mostly in the

form of foams Cells are then seeded and expected to grow into the scaffold However, this approach has resulted in the in vitro growth of tissues with cross-sections of less than 500µm from the external surface (Ishaug-Riley et al, 1997; Freed and Vunjak-Novakovic, 1998) This is probably due to the diffusion constraint of the foam which causes scarcity in nutrients and oxygen supply, and insufficient removal of waste products

Table 2.2: Summary of the advantages and disadvantages of conventional scaffold fabrication techniques (Leong et al, 2003a)

Textile technique Larger pores and high

porosity

Structurally unstable and lacking in mechanical properties

Fibre bonding Easy process, high porosity

and high surface area to volume ratio

High processing temperature for non-amorphous polymer, limited range of polymers, lack of mechanical strength, poor control over micro-architecture, problems with residual solvent

Phase separation

Highly porous structures, allows incorporation of bioactive agents

Lack of control over micro-architecture, limited range of pore sizes, problems with residual solvent

Limited membrane thickness, lack of mechanical strength, problems with residual solvent and residual porogens

Membrane

lamination

Macro shape control, independent control of porosity and pore size,

Lack of mechanical strength, problems with residual solvent, tedious and time-consuming procedure, limited interconnected pores

Melt moulding

Independent control of porosity and pore size, Macro shape control

High processing temperature for non-amorphous polymer, Residual porogens

Problems with residual solvent and residual porogens

of porosity and pore size

Limited interconnected pores, Residual porogens

Freeze drying Highly porous structures

High pore interconnectivity

Limited to small pore sizes

Hydrocarbon

templating

No thickness limitation, independent control of porosity and pore size

Problems with residual solvent and Residual porogens