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

Chapter Two: Literature Review Chapter Two Literature Review 10 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 11 Chapter Two: Literature Review 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 12 Chapter Two: Literature Review and biochemical properties. Most synthetic polymers degrade via chemical hydrolysis and insensitive to enzymatic processes so that their degradation behaviours 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) 45-55 Degra dation Time (months) a Density (g/cc) Tensile Strength (MPA) Elonga tion % Modu lus (GPA) PLGA L-PLA Amor phous Amor phous 173-178 6-12 41.4-55.2 3-10 1.4-2.8 55-60 12-16 1.271.34 1.25 27.6-41.4 3-10 1.4-2.8 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 PEG 67-69 - 72 - 1.05 - - - PCL-PLA 61-68 - 55 - - - - - PCL-PEG 60-67 - 69 - - - - - PCL-PEG-P 59-69 - 69 - - - - - 58-67 - 67 - - - - - 58-65 - 68 - - - - - DL-PLA CL PLA-PCL-P LA PEG-PCL-P LA a Time to complete mass loss. Time also depends on part geometry. 13 Chapter Two: Literature Review 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 14 Chapter Two: Literature Review 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 15 Chapter Two: Literature Review biocompatibility and efficacy studies had also been performed in the research leading to the introduction of the MonocrylTM monofilament sutures as reported by Bezwada et al (1995). Several researchers (Hutmacher et al, 2000a; Marra et al, 1999a; Corden et al, 2000) could show the good biocompatibility of two-dimensional PCL specimens in human osteoblast-like cell cultures. Marra et al (1999a) concluded from their 2-D cell studies that PCL is superior to PLGA for bone cell growth. Cell migration requires a biomolecular healthy and dynamic interaction among the cell, scaffold surface and its cytoskeleton. Human osteoblast-like cell culture data showed the evidence of good biocompatibility of PCL for hard tissue formation. 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. 16 Chapter Two: Literature Review 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 17 Chapter Two: Literature Review 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 18 Chapter Two: Literature Review 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 19 Chapter Two: Literature Review of 30% and compressive strength of 13.8 MPa. After 12 weeks of implantation the results of rabbit model revealed great extents of bone tissue in-growth into the implants. Likewise, the results obtained from dog model after weeks of implantation exhibited excellent integration of the implant with the host bone. Vail et al (1999) also conducted the augmentation of alveolar ridge defects in canines to also assess the safety and efficacy of the SLS-fabricated calcium phosphate implants. Histological examination showed the basic biocompatibility of the implant material and mineralized bone formation in the macro pores of the implant. Using a simplified selective laser sintering apparatus, Rimell and Marquis (2000) fabricated clinical implants with ultra high molecular weight polyethylene (UHMWPE). It was reported that solid linear continuous bodies could be fabricated, but in case of sheet-like structure there happened material shrinkage. The material degradation in terms of chain scission, cross-linking and oxidation was observed when exposed to the laser beam. It was concluded that improved starting powder with increased density was required to apply the SLS technique to fabricate UHMWPE devices. Tan et al (2003) assessed the suitability of SLS technique for the processing of different compositions of non-degradable polyetheretherketone (PEEK)/HA powder blends. Leong et al (2003a) have also investigated the suitability of different biopolymers for SLS processing. Based on the sintering experiments thus far conducted using a commercialized SLS system, sinterstation 2500 have ascertained the feasibility of processing PLLA, PCL and polyetheretherketone (PEEK) powders by SLS technique with varying degrees of success. 38 Chapter Two: Literature Review More recently, Williams et al (2005) fabricated and evaluated the PCL scaffolds for bone tissue engineering via SLS. Fabricated PCL scaffolds with porous architecture had sufficient mechanical properties for bone tissue engineering applications. Compressive modulus and yield strength values were found to be in range of 52 to 67MPa and 2.0 to 3.2 MPa, respectively, lying within the lower range of properties reported for human trabecular bone. Scaffolds were seeded with bone morphogenetic protein-7 (BMP-7) transduced fibroblasts and implanted subcutaneously to evaluate biological properties and to demonstrate tissue in-growth. Histological evaluation and micro-computed tomography (µ-CT) analysis of implanted scaffolds showed in vivo bone generation. They designed and fabricated a prototype mandibular condyle scaffold based on an actual pig condyle to demonstrate the clinical application of this technology. SLS technique is capable of processing polymers, ceramics and metallic powders to manufacture scaffolds with irregular shapes to fit complex anatomic locations including channels and overhangings. Another important advantage of SLS is that there is minimal risk of material contamination as it is solvent free and does not require any secondary binder system. However, the major limitation of SLS is that this technique is confined to processing of only thermally stable polymers, because it involves high processing temperatures. In addition, in many cases the pores of the scaffolds are dependent on the particle size of the powder stock and the compaction pressure used. 39 Chapter Two: Literature Review Laminated Object Manufacturing (LOM): LOM is a process where individual layers (foils) are cut from a sheet by a computer-controlled laser, after which the individual layers are bonded together to form a 3D object. LOM has been used for fabrication of bioactive bone implants, using HA and calcium phosphate laminates. The undersurface of the foil has a binder that when pressed and heated by the roller causes it to glue to the previous foil. Once the parts have been built, the exterior of the slice is hatched to help the removal of the excess material, as opposed to fluid-based processes (e.g. the SLA process), where the interior is hatched. The disadvantage is the production of burnt edges due to the laser cut, not an issue with most applications, but may create unwanted and possibly harmful debris in biomedical applications. Material degradation in the heated zone may also occur. Newer technology (Kira’s Paper Lamination Technology) has substituted the laser with a blade, and thus this technology might have a greater potential to be applied in the fabrication of scaffolds. 2.4.3.2 Systems Based on Print Technology Three-dimensional Printing (3DP): 3D printing (3DP) technology was first developed by Sachs et al (1989) at Massachusetts Institute of Technology (MIT) and became one of the most investigated RP techniques in tissue engineering and drug-delivery applications (Curodeau et al, 2000; Zeltinger et al, 2001; Lam et al, 2002; Sherwood et al, 2002). This technology is based on the ink-jet printing of a binder, much like a conventional desktop printer. The 3D printer constructs the 3D model by first spreading a layer of fresh powder over a building platform. An ‘inkjet’ print head prints or deposits the binder solution 40 Chapter Two: Literature Review onto the powder bed. The binder dissolves and joins adjacent powder particles. After the 2D layer profile is printed, the piston chamber is lowered (z-axis control) and a fresh layer of powder is laid down. The printing cycle continues and the layers merge together when fresh binder is deposited until the whole scaffold structure is constructed. The unbound powder acts to support overhanging or unconnected features and needs to be removed after component completion. The completed scaffold is embedded inside a cake of unprocessed powders and is extracted by brushing away the loose powders. Thus far, a number of researchers (Kim et al, 1998; Park et al, 1998; Zeltinger et al, 2001, Lam et al, 2002) fabricated scaffolds employing 3DP and tested them for tissue engineering applications. Kim et al (1998) fabricated cylindrical (8mm diameter & mm height) porous scaffold by means of 3DP with particulate leaching technique using polylactide-coglycolide (PLGA) powder mixed with salt particles and a suitable organic solvent. Distilled water was used to leach out the salt particles after 3DP fabrication. The scaffold resulted in pore sizes of 45–150 µm and 60% porosity. In vitro cell culture study with hepatocytes (HCs) revealed the successful attachment of large numbers of HCs to the scaffolds’ outer surfaces and internal channels. Histological examination also showed the in-growth of HCs into the pore spaces. In 2001, Zeltinger and his group employed 3DP to fabricate disc shaped (10mm diameter & 2mm height) porous scaffolds of poly(l-lactic acid) (l-PLA) and evaluated the cell culture performance of these scaffolds using three different cell types, canine dermal fibroblasts, vascular smooth muscle cells and 41 Chapter Two: Literature Review microvascular epithelial cells. The scaffolds produced two different porosities (75% and 90%) and four different pore size distributions ([...]... diffusion constraint of the foam which causes scarcity in nutrients and oxygen supply, and insufficient removal of waste products 30 Chapter Two: Literature Review Table 2. 2: Summary of the advantages and disadvantages of conventional scaffold fabrication techniques (Leong et al, 20 03a) Process Advantages Textile technique Larger Disadvantages pores and high Structurally unstable and lacking in mechanical... strength and the final desired Ca/P ratio The pore sizes of the fabricated implants were measured to be up to 20 0 µm with an overall porosity 37 Chapter Two: Literature Review of 30% and compressive strength of 13.8 MPa After 12 weeks of implantation the results of rabbit model revealed great extents of bone tissue in-growth into the implants Likewise, the results obtained from dog model after 4 weeks of. .. the capabilities of 3DP for developing tissue engineering scaffolds include the work of Sherwood et al (20 02) , Lam et al (20 02) and Roy et al (20 03) In all their works, 3DP demonstrated its potential to fabricate scaffolds with a range of porosities and pore distributions which further showed success in cell culture studies Sherwood et al (20 02) developed an osteochondral scaffold using the TheriFormTM... such as wettability (Olde et al, 20 03), swelling (Bezemer et al, 1999; van Dijkhuizen-Radersma et al, 20 02; Deschamps et al, 20 02) [26 ,28 ,29 ], biodegradation rate (Deschamps et al, 20 02) , protein adsorption (Mahmood et al, 20 04) and mechanical properties (Woodfield et al, 20 04) Furthermore, PEOT/PBT block copolymers have shown to be extensively biocompatible both in vitro and in vivo (van Blitterswijk... smaller and more intricate objects are fabricated because of absorption and scattering of the laser beam Therefore, the overall use of SLA in the biomedical industry is currently limited to the development of anatomical models for surgical planning or teaching (Bibb and Sisias, 20 02; Vrielinck et al, 20 03; Cohen and Letelier, 20 03) Selective Laser Sintering (SLS): The SLS technique utilizes a CO2 laser... 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 23 Chapter Two: Literature Review (a) (b) Figure 2. 3: (a) Honeycomb... to the faces normal to X3 are referred to as out -of- plane properties (a) (b) (c) Figure 2. 4: In-plane compression of honeycomb pores; (a) Initial elastic bending of pore walls, (b) Buckling of pore edges at higher stress levels and (c) Out -of plane compression of honeycomb pores (Gibson and Ashby, 1997) 24 Chapter Two: Literature Review (a) (b) Figure 2. 5: A schematic diagram for a honeycomb loaded... joint replacement Lam et al (20 02) produced different scaffold designs with different pore sizes and interconnectivities from blend of commercially available biomaterials of natural origin (cornstarch, dextran and gelatin) utilizing distilled water as binder 42 Chapter Two: Literature Review via commercial 3D printer A series of infiltration and post-processing methods by using aliphatic polyesters were... al, 1998; Curodeau et al, 20 00) The component resolution and efficiency of removal of trapped materials are also affected by the surface roughness and aggregation properties of the powdered materials (Zeltinger et al, 20 01; Lam et al, 20 02; Sherwood et al, 20 02) On top of these, the scaffold fabricated by 3DP lacks in mechanical properties 2. 4.3.3 System Based on Assembly Technology Shape Deposition... Newer technology (Kira’s Paper Lamination Technology) has substituted the laser with a blade, and thus this technology might have a greater potential to be applied in the fabrication of scaffolds 2. 4.3 .2 Systems Based on Print Technology Three-dimensional Printing (3DP): 3D printing (3DP) technology was first developed by Sachs et al (1989) at Massachusetts Institute of Technology (MIT) and became one of . 15 -2 0 >6.9 PCL 5 8-6 8 - 70 > ;24 1.11 20 . 7-3 4.5 30 0-5 00 0 .21 -0 .34 PEG 6 7-6 9 - 72 - 1.05 - - - PCL-PLA 6 1-6 8 - 55 - - - - - PCL-PEG 6 0-6 7 - 69 - - - - - PCL-PEG-P CL 5 9-6 9 - 69 - - - -. 4 5-5 5 6- 12 1 .27 - 1.34 41. 4-5 5 .2 3-1 0 1.4 -2 . 8 DL-PLA Amor phous 5 5-6 0 1 2- 16 1 .25 27 . 6-4 1.4 3-1 0 1.4 -2 . 8 L-PLA 17 3-1 78 6 0-6 5 > ;24 1 .24 55 . 2- 82. 7 5-1 0 2. 8-4 .2 PGA 185 -2 2 5 25 -6 5 6- 12 1.53. - - - PLA-PCL-P LA 5 8-6 7 - 67 - - - - - PEG-PCL-P LA 5 8-6 5 - 68 - - - - - a Time to complete mass loss. Time also depends on part geometry. 13 Chapter Two: Literature Review 2. 2.1