Baixe Collagen-hydroxyapatite composites for hard tissue repair e outras Notas de estudo em PDF para Engenharia de Produção, somente na Docsity! 43 DA Wahl et al. Collagen-Hydroxyapatite Composites for Hard Tissue RepairEuropean Cells and Materials Vol. 11. 2006 (pages 43-56) ISSN 1473-2262 Abstract Bone is the most implanted tissue after blood. The major solid components of human bone are collagen (a natural polymer, also found in skin and tendons) and a substituted hydroxyapatite (a natural ceramic, also found in teeth). Although these two components when used separately provide a relatively successful mean of augmenting bone growth, the composite of the two natural materials exceeds this success. This paper provides a review of the most common routes to the fabrication of collagen (Col) and hydroxyapatite (HA) composites for bone analogues. The regeneration of diseased or fractured bones is the challenge faced by current technologies in tissue engineering. Hydroxyapatite and collagen composites (Col-HA) have the potential in mimicking and replacing skeletal bones. Both in vivo and in vitro studies show the importance of collagen type, mineralisation conditions, porosity, manufacturing conditions and crosslinking. The results outlined on mechanical properties, cell culturing and de- novo bone growth of these devices relate to the efficiency of these to be used as future bone implants. Solid free form fabrication where a mould can be built up layer by layer, providing shape and internal vascularisation may provide an improved method of creating composite structures. Keywords: Collagen Type I, hydroxyapatite, composite scaffolds, biocompatible devices, bone substitute, tissue engineering *Address for correspondence: Denys A Wahl, Department of Materials, University of Oxford, Parks Road, Oxford, OX1 3PH, UK E-mail: denys.wahl@materials.ox.ac.uk Introduction Bone tissue repair accounts for approximately 500,000 surgical procedures per year in the United States alone (Geiger et al., 2003). Angiogenesis, osteogenesis and chronic wound healing are all natural repair mechanisms that occur in the human body. However, there are some critical sized defects above which these tissues will not regenerate themselves and need clinical repair. The size of the critical defect in bones is believed to increase with animal size and is dependent on the concentration of growth factors (Arnold, 2001). In vivo studies on pig sinus (Rimondini et al., 2005) and rabbit femoral condyles (Rupprecht et al., 2003) critical size defects of 6x10mm and 15x25mm respectively were measured. These defects can arise from congenital deformities, trauma or tumour resection, or degenerative diseases such as osteomyelitis (Geiger et al., 2003). Bone substitutes allow repair mechanisms to take place, by providing a permanent or ideally temporary porous device (scaffold) that reduces the size of the defect which needs to be mended (Kohn, 1996). The interest in temporary substitutes is that they permit a mechanical support until the tissue has regenerated and remodelled itself naturally. Furthermore, they can be seeded with specific cells and signalling molecules in order to maximise tissue growth and the rate of degradation and absorption of these implants by the body can be controlled. Bioresorbable materials have the potential to get round the issues that occur with metallic implants, such as strain shielding and corrosion. Titanium particles produced from wear of hip implants, were shown to suppress osteogenic differentiation of human bone marrow and stroma-derived mesenchymal cells, and to inhibit extra cellular matrix mineralisation (Wang et al., 2003). Furthermore, these materials should help to reduce the problems of graft rejection and drug therapy costs, associated with for example the use of immunosuppressants (e.g. FK506) after implantation of bone grafts (Kaihara et al., 2002). When using a biodegradable material for tissue repair the biocompatibility and/or toxicity of both the material itself and the by-product of its degradation and subsequent metabolites all need to be considered. Further, at the site of injury, the implant will be subjected to local stresses and strains. Thus, the mechanical properties of the implant, such as tensile, shear and compressive strength, Young’s modulus and fracture toughness need to be taken into consideration when selecting an appropriate material. However, given a bone analogue is ideally resorbable, these properties are not as important as for an inert implant which does not (intentionally) degrade. It is important for the bioresorbable material to be osteoconductive and osteoinductive, to guide and to encourage de novo tissue formation. The current aim of the biological implant is to be indistinguishable from the surrounding host bone COLLAGEN-HYDROXYAPATITE COMPOSITES FOR HARD TISSUE REPAIR DA Wahl and JT Czernuszka Department of Materials, University of Oxford, Parks Road, Oxford, OX1 3PH, UK 44 DA Wahl et al. Collagen-Hydroxyapatite Composites for Hard Tissue Repair (Geiger et al., 2003). It is self evident that creating new tissue will lead to the best outcome for the patient in terms of quality of life and function of the surrounding tissue. Synthetic polymers are widely used in biomaterial applications. Examples in tissue engineering include aliphatic polyesters [polyglycolic acid (PGA) and poly- L-lactic acid (PLLA)], their copolymers [polylactic-co- glycolic acid (PLGA)] and polycaprolactone (PCL). However, the chemicals (additives, traces of catalysts, inhibitors) or monomers (glycolic acid, lactic acid) released from polymer degradation may induce local and systemic host reactions that cause clinical complications. As an example, lactic acid (the by-product of PLA degradation) was found to create an adverse cellular response at the implant site by reducing the local pH, in which human synovial fibroblasts and murine macrophages released prostaglandin (PGE2), a bone resorbing and inflammatory mediator (Dawes and Rushton, 1994). Nevertheless, a potential way to stabilise the pH is by the addition of carbonate to the implant (Wiesmann et al., 2004). Some polymeric porous devices also have the disadvantages of not withstanding crosslinking treatments such as dehydrothermal treatment (DHT) and ultraviolet (UV) irradiation (Chen et al., 2001). The drawback of requiring chemical crosslinking (glutaraldehyde) is the formation and retention of potential toxic residues making these techniques less desirable for implantable devices (Hennink and van Nostrum, 2002). The reader is referred to Athanasiou et al. (1996) for a review in the biocompatibility of such polymeric materials. Ceramics [eg HA, tricalcium phosphate (TCP) and/or coral] have been suggested for bone regeneration. Bone substitutes from these materials are both biocompatible and osteoconductive, as they are made from a similar material to the inorganic substituted hydroxyapatite of bone. However, the ceramic is brittle (Kc<1MPam -1) (Dewith et al., 1981) and does not match alone the mechanical properties of cortical bone (Kc~2-12MPam -1) (Bonfield, 1984). Therefore, calcium phosphates have been used in areas of relatively low tensile stress such as bone/ dental fillings or as coatings on implanted devices (Vallet- Regi and Gonzalez-Calbet, 2004). Collagen, as a natural polymer, is increasingly being used as a device material in tissue engineering and repair. It is, for example, found in bone (Type I), cartilage (Type II) and in blood vessel walls (Type III) and has excellent biocompatible properties. Collagen is easily degraded and resorbed by the body and allows good attachment to cells. However, its mechanical properties are relatively low (E ~100 MPa) in comparison to bone [E ~2-50GPa (Clarke et al., 1993)] and it is therefore highly crosslinked or found in composites, such as collagen-glycoaminoglycans for skin regeneration (O’Brien et al., 2004), or collagen- hydroxyapatite for bone remodelling (Kikuchi et al., 2004b). Both collagen and hydroxyapatite devices significantly inhibited the growth of bacterial pathogens, the most frequent cause of prosthesis-related infection, compared to PLGA devices (Carlson et al., 2004). Other approaches of bone repair have been to use autografts, allografts and xenografts. Although very successful in many operations, autografts have the disadvantages of insufficient supply and morbidity, as well as increasing surgery times and donor site pain (Uemura et al., 2003). Allografts and xenografts are associated with infection and inflammation and have perceived ethical disadvantages. Xenografts also carry the risk of species- to-species transmissible diseases (Meyer et al., 2004). Careful consideration of the bone type and mechanical properties are needed for bone substitutes. Indeed, in high load-bearing bones such as the femur, the stiffness of the implant needs to be adequate, not too stiff to result in strain- shielding, but stiff enough to present stability. However, in relatively low load-bearing applications such as cranial bone repairs, it is more important to have stability and the correct three dimensional shapes for aesthetic reasons. One of many approaches of tissue engineering is to create a device of similar mechanical and biological properties to the one of the substituted tissue. Therefore, this review will focus on the engineering of a bone substitute, from the understanding of the individual and main components of bone to the creation of a collagen-HA composite. It will bring forward the idea that the manufacturing process of such biocompatible device defines its final microstructure, which in turn will determine its mechanical and biological response, and therefore its efficiency in repairing a hard tissue defect. A Composite of Collagen and Hydroxyapatite Skeletal bones comprise mainly of collagen (predominantly type I) and carbonate substituted hydroxyapatite, both osteoconductive components. Thus, an implant manufactured from such components is likely to behave similarly, and to be of more use than a monolithic device. Indeed, both collagen type I and hydroxyapatite were found to enhance osteoblast differentiation (Xie et al., 2004), but combined together, they were shown to accelerate osteogenesis. A composite matrix when embedded with human-like osteoblast cells, showed better osteoconductive properties compared to monolithic HA and produced calcification of identical bone matrix (Serre et al., 1993; Wang et al., 1995). In addition, Col-HA composites proved to be biocompatible both in humans and in animals (Serre et al., 1993; Scabbia and Trombelli, 2004). These composites also behaved mechanically in a superior way to the individual components. The ductile properties of collagen help to increase the poor fracture toughness of hydroxyapatites. The addition of a calcium/ phosphate compound to collagen sheets gave higher stability, increased the resistance to three-dimensional swelling compared to the collagen reference (Yamauchi et al. 2004) and enhanced their mechanical ‘wet’ properties (Lawson and Czernuszka, 1998). This happened even when the collagen was highly crosslinked. Collagenase digestion can represent an in vitro measure of the rate of degradation and resorption of a biological implant. Uncrosslinked collagen and hydroxyapatite- collagen gel beads were analysed by collagenase digestion. HA-containing gel beads were less prone to collagenase and degraded more slowly than collagen gel beads. The 47 DA Wahl et al. Collagen-Hydroxyapatite Composites for Hard Tissue Repair have been mixed with biodegradable matrices, such as collagen and PGA (Vallet-Regi and Gonzalez-Calbet, 2004). The cellular response to HA particles, incorporated into a matrix or as coating, has been shown to depend on properties such as particle size and morphology (needle like, spherical or irregular plates), chemical composition, crystallinity and sintering temperatures. Due to such variability in HA properties, contradictions arise in the literature on the influence of one property over another and a need for a more systematic research was proposed by Laquerriere et al. (2005). However, it is generally accepted that needle shaped particles produce deleterious cellular response compared to spherical shaped particles. Indeed, macrophages have been found to release higher levels of inflammatory mediators and cytokines such as metalloproteinases (MMPs) and Interleukine-6 respectively (Laquerriere et al., 2005). In the case of collagen-HA implants, the size and crystallinity of the HA particles will have an importance to its stability and inflammatory response. In skeletal bones, carbonate substituted HA crystals are mineralised within small gaps of the collagen fibrils and have been quoted as 50nm×25nm×2-5nm in length, width and thickness respectively (Vallet-Regi and Gonzalez-Calbet, 2004). They provide a local source of calcium to the surrounding cells as well as interacting with collagen fibrils in order to achieve the relatively high mechanical properties of bones. However, small sintered particles of less than 1µm have been cautioned against in bone implants due to their increased inflammation response (Laquerriere et al., 2005) and cell toxicity in vitro (Sun et al., 1997). In contrast, Lawson and Czernuszka (1998) have shown that smaller plate-like particles of the order of 200nm×20nm×5nm produced an enhanced osteoblastic adhesion and proliferation compared to HA particles of an order of magnitude larger. These were carbonate substituted HA particles and produced at physiological temperatures (unsintered). Collagen-Hydroxyapatite Composite Fabrication This section will summarise the different methods for collagen-hydroxyapatite composite formation. It will include the production methods of composite gels, films, collagen-coated ceramics, ceramic-coated collagen matrices and composite scaffolds for bone substitutes and hard tissue repair. In vitro collagen mineralisation Direct mineralisation of a collagen substrate involves the use of calcium and phosphate solutions. Collagen can either be a fixed solid film through which calcium and phosphate ions diffuse into the fibrils (Lawson and Czernuszka, 1998), or as a phosphate-containing collagen solution (Kikuchi et al., 2004a), or an acidic calcium-containing collagen solution (Bradt et al., 1999). The advantage of using the first method is that the orientation of the collagen fibres can be controlled (Iijima et al., 1996). Indeed, it has been shown that the c-axis of HA crystals can be made to grow along the direction of collagen fibrils if the right conditions of mineralisation are met. These conditions (pH 8-9 and T =40°C) promote calcium ion accumulation on the carboxyl group of collagen molecules, leading to HA nucleation (Kikuchi et al., 2004a). The formation of HA is temperature and pH dependent as well as molar dependent (Ca/P ratio). Figure 1 shows an experimental set-up for calcium and phosphate diffusion and apatite crystallisation onto collagen substrates. Undenatured collagen films can be obtained from an acidic collagen suspension by air dehydration at different temperatures (4-37°C) or by applying cold isostatic pressure (200MPa) for 15h (Kikuchi et al., 2001). Figure 1. An experimental set-up for the direct mineralisation of a collagen sheet (Modified from Lawson and Czernuszka, 1998) 48 DA Wahl et al. Collagen-Hydroxyapatite Composites for Hard Tissue Repair Thermally-triggered assembly of HA/collagen gels Liposomes have been used as drug delivery system due to their ability to contain water-soluble materials in a phospholipids layer (Ebrahim et al., 2005). Pederson et al. (2003) have combined the direct mineralisation method with the ability of liposomes to exist in a gel for the potential use in injectable composite precursors. They reported a method whereby calcium and phosphate ions where encapsulated within the liposomes and the latter inserted into a collagen acidic suspension. After injection into a skeletal defect, the increase in temperature due to the body heat would start a gelation process, forming a collagen fibrous network where mineralisation occurs after reaching the liposome’s transition temperature (37ºC) (Pederson et al., 2003). Vacuum infiltration of collagen into a ceramic matrix Multiple tape casting is a method of ceramic scaffold production: an aqueous hydroxyapatite slurry containing polybutylmethalcrylate (PBMA) spheres is heated to high temperatures to burn out the BMPA particles, forming a porous HA green body (Werner et al., 2002). Collagen infiltration was performed under vacuum, and the collagen suspension filled in the gaps of the porous matrix. The final composite was then freeze dried to create microsponges within. Variation of the final product was dependent on process time and flow resistance during filtration (Pompe et al., 2003). Enzymatic mineralisation of collagen sheets Figure 2 shows a method of enzymatic loading of collagen sheets and the following cycle of mineralisation. The collagen containing an alkaline phosphatase was allowed to be in contact with an aqueous solution of calcium ions and phosphate ester. The enzyme provided a reservoir for PO4 3- ions for calcium phosphate to crystallise and mineralisation was found to occur only on these coated areas (Yamauchi et al., 2004). The sample was then coated again with a collagen suspension, air dried and cross-linked with UV irradiation. Repeating this cycle resulted in multilayered composite sheets of calcium/phosphate and collagen, with a sheet thickness of 7µm (Yamauchi et al., 2004). Water-in-Oil emulsion system Col-HA microspheres or gel beads have been formed in the intention of making injectable bone fillers. A purified collagen suspension mixed with HA powders at 4°C was inserted in olive oil and stirred at 37°C. The collagen aggregated and reconstituted in the aqueous droplets. The phosphate buffered saline (PBS) was added to form gel beads (Hsu et al., 1999). However, this method has the disadvantage of not being able totally to remove the oil content from the composite. Additionally, composite gels for injectable bone filler have the disadvantage that the viscosity of the mixture becomes too low on contact with body fluids, resulting in the “flowing out” from the defect (Kocialkowski et al., 1990). Freeze Drying and Critical Point Drying Scaffolds In order to form a sponge-like porous matrix, either freeze- drying or critical point drying (CPD) is required. A collagen/HA/water suspension can be frozen at a controlled rate to produce ice crystals with collagen fibres at the interstices. In the case of freeze-drying, ice crystals are transformed to water vapour at a specific temperature and pressure by sublimation. In the case of CPD, liquid and vapour become indistinguishable above a certain pressure and temperature, where both densities of the two phases converge and become identical (supercritical fluid shown in Figure 3). Above this critical point, surface tension is negligible resulting in little matrix collapse. The lower critical point of carbon dioxide (31.1°C, 7.3MPa) compared to water (364°C, 22.1MPa) makes the use of CO2 very popular when critical point drying. Freeze drying and critical point drying have the fewest residual solvent problems compared to other scaffold manufacturing techniques. Furthermore, the easy removal of ice crystals compared to porogens, used in conventional Figure 2. The cycle of enzymatic mineralisation of collagen sheets Figure 3. Phase Diagram of Carbon Dioxide. (a) Triple Point (-56.4ºC and 0.5MPa) (b) Critical Point (31.1ºC and 7.3MPa). 49 DA Wahl et al. Collagen-Hydroxyapatite Composites for Hard Tissue Repair polymeric-porogen leaching, eliminates any dimensional restrictions (Leong et al., 2003). In freeze drying and critical point drying processes, pore sizes are determined by the ice crystal formation. Changing the freezing rate and solubility of the suspension as well as the collagen concentration can alter the pore size. Additional solutes (ethanoic acid, ethanol) can create unidirectional solidification of collagen solutions (Schoof et al., 2000), and lower freezing rates generate larger pore sizes (O’Brien et al., 2004). Pore size is important in scaffolds as they will determine cell adhesion and migration, the mechanical properties of the scaffold and as a result the success of new tissue formation. Karageorgiou and Kaplan (2005) recommended biomaterial scaffolds to possess pore sizes of over 300µm in order to favour direct osteogenesis and to allow potential vascularisation. Solid Freeform Fabrication with Composite Scaffolds Figure 4 shows a schematic of a computer model of a bone, to the creation of a mould (3-D printing) and to the scaffold production. The model is first drawn with the help of computer aided design, the mould is then printed with a “support” and “build” material, the sacrificial mould dissolved to obtain the casting mould, Col-HA cast into the mould and frozen, ice crystals replaced with ethanol, ethanol-liquid CO2 exchanged and critical point dried to finally arrive with an exact porous replica of the original design. This method has been used extensively by Sachlos (Sachlos and Czernuszka, 2003; Sachlos et al., 2003) and is part of many solid freeform fabrication or rapid prototyping methods used to form scaffolds for tissue engineering and reviewed by Leong et al. (2003). Solid freeform fabrication techniques have recently been developed with artificial polymers and ceramic materials (Taboas et al., 2003; Hutmacher et al., 2004). These have the ability to change pore interconnectivity, pore size and pore shape, but have the disadvantage of not having the affinity of collagen to cell attachment. Another major advantage of Collagen-HA scaffolds produced through the SFF method is the ability to control variables at several length scales: Control of the external structure: computerised tomography (CT) or magnetic resonance imaging (MRI) scans can be used to engineer biocompatible scaffolds. Although CT scans are 2-dimensional and MRI scans less sensitive to skeletal tissues (bones), it is possible to obtain the relative dimensions of a defect. These scans can be directly converted to a computer design and then the mould or scaffold itself printed out with solid freeform fabrication techniques. Control of the internal structure: the Harvesian system of bone is a very complex system of vascularisation, in which cells are not found beyond 200µm of a blood supply (and therefore oxygen). 3-D printing can incorporate such architecture in its scaffold manufacture, with the hope that Figure 4. The use of Solid Freeform Fabrication in the design of composite scaffolds Figure 5. Scanning Electron Microscopy Image of a collagen scaffold with graded porosity. On the left of the scaffold, the mean pore diameter is ~70µm, and on the right the mean pore diameter is ~15µm. This type of scaffold could be used to engineer hybrid tissues. 52 DA Wahl et al. Collagen-Hydroxyapatite Composites for Hard Tissue Repair content combined with low porosity exhibited higher UTS and E, expressing the greater importance of the ceramic component. The strengthening effect of HA can be explained by the fact that the collagen matrix is a load transfer medium and thus transfers the load to the intrinsically rigid apatite crystals. Furthermore, the apatite deposits between tangled fibrils ‘cross-link’ the fibres through mechanical interlocking or by forming calcium ion bridges, thus increasing the resistance to deformation of the collagenous fibre network (Hellmich and Ulm, 2002). The mechanical properties of all the tabulated composites are lower than the natural properties of bones. The highly organised structure of cortical and cancellous bones is very hard to reproduce in vitro, and more research is needed on improving the Col-HA composites if they properly want to imitate skeletal bone structure. Cell culturing and in vivo implantation of composites Results of using collagen-calcium phosphate composites in vitro (using osteogenic cells) and in vivo (in bone defects) are presented in Table 4. The table illustrates the potential for fast surface tissue formation (in under 3 weeks) In addition, at least one study showed that angiogenesis had occurred. Inert materials will never give such behaviour. The aim of Col-HA composite scaffolds is to enhance the ease of application of tissue engineering, thus such demonstrably efficient bone forming capacity is of great value. Tissue engineered devices will have a direct impact on post-operative recovery times and overall costs of treatment. Conclusion This paper has examined the processing routes for fabricating collagen-hydroxyapatite composites and their effect on mechanical and biological properties. It is possible to vary the type of collagen, the crosslinking method and density, the porosity levels and to control the stoichiometry and defect chemistry of the hydroxyapatite, as well as the particle size. The volume fractions of each component are important. 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