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Nano hydroxyapatite bone tissue engineering ceramics

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https://www.eduzhai.net American Journal of Biomedical Engineer ing 2013, 3(6): 148-168 DOI: 10.5923/j.ajbe.20130306.04 Nanometre Scale Hydroxyapatite Ceramics for Bone Tissue Engineering Gé rrard Eddy Jai Poine rn*, Ravi Krishna Brundavanam, De rek Fawce tt 1M urdoch Applied Nanotechnology Research Group, Department of Physics, Energy Studies and Nanotechnology School of Engineering and Energy, M urdoch University, M urdoch, Western Australia 6150, Australia Abstract The consequences of bone traumatisation, loss or damage, resulting fro m injury or disease can dramatically reduce the quality of life for a patient at a significant socioeconomic cost. The aim o f bone tissue engineering is not only to repair, but also to init iate natural bone regeneration. The u ltimate goal is to develop a synthetic tissue scaffold that uses biocompatible materials to produce an effective functional replacement for damaged bone tissue. Thus, avoiding all the problems associated with current bone transplantation procedures. However, repairing and regenerating damaged bone tissue is still a challenging task. Since the skeletal tissues are co mp lex and the presence of foreign materials used to construct a tissue scaffold within the body’s environment will initiate an inflammatory response, which ultimately leads to failure of the repair procedure. This review discusses a number of materials currently being used or has the potential to be used in bone tissue engineering applications. In particu lar, the advantages and limitations of hydroxyapatite are discussed at length, since its desirable properties such as biocompatibility, b ioactivity, osteoconductivity and osteoinductivity make it an ideal starting material for bone tissue engineering applications. Keywords Bone Tissue Engineering, Bio materials, Tissue Scaffolds, Hydro xyapatite 1. Introduction Today, there is a high demand for advanced biosynthetic bone-like materials for the development o f b io medical devices and imp lants for use in tissue engineering applications. For examp le, t issue loss as a result of injury and diseases in an increasingly ag ing population reduces the quality of life for many at a significant socioeconomic cost. This is co mpounded by the fact that many conventional implants fail due to non-integration with the surrounding normal tissue. This leads to complications that require revision surgery to remove or repair the imp lant device. Furthermore, problems also exist for bio medical devices implanted within the body for the controlled release of drugs and similar c linica l applications. Hence there is an a ll round interest to develop novel tissue engineering applications that incorporate “living” constructional materials that are similar to natural bone tissue and posses the potential to integrate with the surrounding native tissue for a faster healing outcome. Con vent ion al med ical p rocedu res such as autog raft (patient’s own bone) and allograft, (sourced fro m another donor) treatments for the rep lacement of bone loss resulting * Corresponding author: g.poinern@murdoch.edu.au (Gérrard Eddy Jai Poinern) Published online at https://www.eduzhai.net Copyright © 2013 Scientific & Academic Publishing. All Rights Reserved fro m damage or disease have met with varying degrees of success to date. Both procedures have significant risks associated with them. In the case of an autograft, donor site morb idity is a frequent outcome[1, 2], wh ile disease transmission is potentially a serious side effect of an allograft procedure[3]. Furthermore, both procedures suffer fro m the limited supply of viable bone tissue. An attractive alternative to nature bone grafts is the use of tissue engineering techniques to create synthetic scaffolds that can effectively replicate the various physical, chemical and mechanica l properties found in natural bone tissue. There has been considerable development over the past few decades in mu ltidiscip linary field of tissue engineering, which has being able to produced engineered imp lantable human tissues such as bone, cartilage and skin[4, 5]. The research to date has clearly demonstrated that a major function of tissue engineering is to create an environment that can promote productive and efficient cellu lar act ivity for regenerative purposes. Therefore, it is very important that any potential tissue scaffold be capable of replicating the extracellular matrix (ECM) and solicits favourable cell responses. Fro m the cellular perspective, the interaction between the cell and nanometre scale structures is crucial for controlling a variety of cell functions such as adhesion and proliferation[6]. The operational demands placed on an engineered biomed ical device or t issue scaffold are numerous and presents many challenges that must be overco me to achieve American Journal of Biomedical Engineer ing 2013, 3(6): 148-168 149 a successful clinical outcome. For examp le, the biological compatibility of the device or scaffold material is crucial in preventing any cytotoxicity, immunological reactions, and inflammat ion responses from the body[7–9]. The presence of any foreign material within the body environment will initiate an immediate inflammatory response at the site. As a consequence, a complex biochemical cascade of events takes place in which cells arrive and start producing chemokines, cytokines and growth factors to init iate the repair o f damaged tissues surrounding the foreign material. When a scaffold is imp lanted into the body environment, the presence of these cells on the surface of the scaffold can initiate a foreign body reaction to the biomaterial used to manufacture the scaffold. These cells produce oxygen radicals and enzy mes that have the potential to degrade the scaffold, which can u ltimately lead to the failure o f the s caffo ld [1 0] . Historically, a variety of different metals, ceramics and polymers have been used to repair or replace damaged bone tissue with varying degrees of success. Despite having a wide selection of biomaterials and surgical implants, no single material to date can exactly match the composition, structure, che mica l and physical properties of any part icular body part. This is especially true of bone tissue, with its very complex h ierarchical structure and its associated biological functions such as providing a natural reservoir of healing cells and mineral ions that play an important part in maintaining the biochemical balance within the body[11, 12]. Since the first total hip replacement in 1970 by Hamadouche et al, based on ceramic alu mina, the use of hip and knee replacements has steadily increased worldwide [13]. Recently, the Australian Orthopaedic Association’s National Jo int Replacement Registry noted that there are more than 75,000 hip and knee replacements being carried out annually. The register also indicated that there was an increasing trend of around 8% per annu m in surgical procedures involving implant replacements. This trend was also reflected in the aging population statistics and also in the increasing number of implants and biomedical devices being used in younger patients[14]. This health issue is not unique to Australia, but in fact is a major g lobal health problem facing the world today. And due to its global importance, the United Nat ions, the World Health Organization and 37 countries declared the period from 2000 to 2010 as the Bone and Joint Decade[15]. For the last 40 years, several approaches have been emp loyed to develop biosynthetic bone grafting substitutes for the reconstruction of osseous defects[16, 17]. These procedures offer patients restored mobility and long-term pain relief. However, it should be mentioned that the biomaterials used to manufacture the bio medical device or implant will have to operate in an environ ment that is both chemically hostile and at the same time very sensitive to the presence of foreign materials. Conventional metallic materials (316 stainless steel, cobalt chro miu m and titaniu m alloys) used for hip and knee replacements, unfortunately can fail due to an incompatibility between the modulus of elasticity of the orthopaedic imp lant and the surrounding bone tissue (stress shielding)[18]. They also fail due to low wear, corrosion resistance and lack of bioco mpatibility. There are currently considerable c linica l concerns regarding the known toxicities associated with elements in metal implants and known pathologies such as particle induced inflammat ion (part icles co ming fro m wear debris) and hypersensitivity associated with metal implants degradation. The body tissues are extremely sensitive to the presence foreign part icles and materials, which solicit an immune response that ultimately result’s in the rejection of the foreign particles[19]. Despite a new generation of Ti implants and a closer e lastic modulus to that of natural bone, the resulting wear resistance under normal loading conditions is still poor[20]. Alternative materials such as polymers and ceramics have also been investigated for possible use in a wide range of tissue engineering applications. Poly mers are large mo lecules composed of smaller repeating structural units called monomers. These high purity mono mers are usually attached by covalent chemical bonds, with cross-linking taking place along the length of the macro molecule. It is the amount of cross-lin king that gives the polymer its physiochemical properties. There are many methods of producing polymer materials, but electro-spinning is one of the most effective processes used for producing nanometre sized fibres which can then be woven into a variety of components[21]. Poly mers have been used in a variety of applications ranging fro m bone and dental cements, bone grafting materials, plates and fixation devices to load bearing applications in orthopaedics[22]. In addit ion to polymers, several other b iosynthetic ceramic based bone substitutes have been developed[23-27]; however, to be acceptable, the material should have a high porosity and exhibit good mechanical properties, otherwise osteoconductivity and resorption at the implant site will be affected by the body’s immunological responses[28]. There has also been considerable research work into combin ing materials with desirable propert ies, wh ile at the same time trying to avoid some of their less attractive properties. The comb ination of t wo or more materials for their favourable propert ies creates a new co mposite material with a set of unique properties that each individual components or phases do not possess. For example, bone is a natural two phase organic-inorganic ceramic co mposite composed of collagen fibrils forming one phase, with a well embedded array of inorganic nano-crystalline co mponents forming the second phase[29]. Recently, a b iosynthetic bone grafting composite material based on nanometre sized hydroxyapatite (nano-HAP) and silica were evaluated by Gerike et al and Got z et al, and shown to exhib it good osteoinductivity[30, 31]. However, the biosynthetic bone graft granules had agglomerated HAP crystals attached to their surface. To fu lly benefit fro m the intrinsic physiochemical properties of nano-HAP that originates fro m its large surface area, it is necessary that the 150 Gérrard Eddy Jai Poinern et al.: Nanometre Scale Hydroxyapatite Ceramics for Bone Tissue Engineering nano-HAP crystals are co mpletely d ispersed and are individually anchored to the underlying matrix of the composite. This problem h ighlights the need for careful preparation and efficient manufacture of the co mposite to take full advantage of the respective phases when forming the new co mposite bio material potential t issue engineering ap p licatio n s . 2. Bone The skeletal system of the human body is important for two main reasons; the first is the structural support of the organs and other related tissues, the second is related to the attachment of muscle groups that permits the motion of the body. The skeleton is constructed of a hard natural t issue called bone and cartilaginous materials. The adult hu man skeleton consists of some 206 bones[32]; some of these provide protection to the internal organs of the body from external forces, wh ile others do specialized functions such as the inner ear bones that transmit sound vibrations for the sense of hearing. The bone matrix also provides a natural reservoir of healing cells and mineral ions that play an important part in maintaining the biochemical balance within the body. For examp le, the calciu m level in the body is closely monitored and regulated (in a process called homeostasis) since calciu m is an important element involved in muscular action and nerve conduction[11]. 2.1. Bone Formation Bone format ion (osteogenesis) commences as early as the foetal stage and continues throughout adulthood. During this period the softer bones of the infant slowly increase in hardness during childhood, eventually becoming quite hard by the adult stage. This hardening process or ossification is the result of the initial hyaline cartilage being transformed into bone. There are two methods in which ossificat ion can take place, the first is by intra me mbranous and the second is endochondral. Both methods produce bone tissue by replacing existing cartilage; however the mechanism that is used in each case is quite d ifferent[33]. Intramembranous ossification occurs during the formation of the craniofacial bones in which mesenchymal cells instead of cartilage are transformed direct ly into osteoblasts. In the remainder of the body endochondral ossification takes place and begins in the embryonic stage with mesenchymal cells forming a cartilage mat rix. In the following stage, osteoblasts and osetoclasts attach themselves to the cartilage matrix. During this stage the osteoclasts generate an acidic microenvironment that degrades the cartilage, while the osteoblasts build a bone matrix onto the cartilage, which now forms a scaffold structure[12, 34]. 2.2. Bone: Its Hierarchical Structure and Mechanical Properties The structural architecture of bone is h ierarch ical and complex[35]. It is an irregular co mposite material with various combinations, arrangements and orientations of its component materials at various scales that gives bone its heterogeneous and anisotropic nature[36]. The overall structure of bone makes it a remarkable b io material with unique physical and mechanical properties. For example, the toughness of bone is achieved by the interplay of the organic collagen scaffold that supports the inorganic minera l phase[37, 38]. The hierarch ical structural of bone consists of five size scales as shown in Figure 1. The first scale is the largest and is called the macrostructure, which consists of cortical and cancellous bone; the second is the microstructure, which ranges in size fro m 10 to 500 µm and consists of single trabeculae or osteons[35]. The oseteon or Haversian system consists of cylindrical structures composed of concentric layers or lamellae surrounding a central duct called the Haversian canal. This canal delivers the blood supply and provides connection for the nerves in surrounding bone tissue[39]. The next step down in size brings us to the sub-microstructure scale, which ranges from 1 to10 µm and contains the lamellae which are thin plate like structures that form the basic structural unit of the bone in the sub-microstructure. The fourth scale is the nanostructure, which ranges in size fro m around 500 n m up to 1µm and contains collagen fibre assemblies co mposed predominantly of collagen fibrils with imp lanted minerals[40]. The final and smallest scale is the sub-nanostructure, which covers the range from 500 n m and below. This size range covers the mo lecular structure of the indiv idual elements, such as collagen mo lecules, bone mineral crystals and the non-collagen organic proteins[41]. Bone is a natural two phase organic-inorganic ceramic composite consisting of collagen fibrils with embedded well-o rdered inorganic nano-crystalline co mponent. The primary organic phase of the bone matrix is Type I co llagen and is secreted by osteoblast cells, wh ich in turn forms self-assembled collagen fibrils. The fibrils are bundled together and predominately self orientate themselves parallel to the load-bearing axis of the bone. Du ring the self assembly p rocess the collagen fib rils, which are typically 300 n m long, develop a 67 n m periodic pattern in which a 40 n m gap or hole is formed between the ends of the fibrils while a 27 n m overlap results fro m the bundle behind[43]. This pattern creates discrete and discontinuous sites for the deposition of plate-like hydro xyapatite (HA P) crystals of bone, which form the second phase of the matrix. HAP is a mineral co mposed of calciu m phosphate which has the general chemical formula of[Ca10(OH)2 (PO4)6]. It is the main inorganic co mponent of natural bone, accounting for up to about 65% by weight of cortical bone and in teeth it accounts for 97 % by weight of dental enamel in mammalian hard t issue[44]. The d iscontinuous discrete sites limit the growth of the crystals and force the crystals to grow with a specific crystalline orientation, wh ich is parallel to the load-bearing axis of the bone and collagen fibrils. The crystal plates typically have a length of 50 n m, a width of around 25 n m and on average a thickness of 3 American Journal of Biomedical Engineer ing 2013, 3(6): 148-168 151 nm[45-48]. The hydroxyapatite has also trace amounts of potassium, manganese, sodium, ch lorine, hydrogen phosphate, citrate and carbonate[49]. The final co mponent of the bone composite consists of the non-collagen organic proteins such as the phosphoprotein group, wh ich are believed to regulate the formation of the inorganic crystal phase by influencing the size, orientation and the depositional environ ment within the spaces between the collagen fibrils. Th is organic phase is also believed to be a source of calciu m and phosphate ions, which are later used in the format ion of the minera l phase[50]. The comb ination of the organic and inorganic phases of the composite bone structure perform together to give bone its unique mechanical properties. Properties such as elastic modulus, toughness, strength, stiffness, and fracture properties of this complex co mposite structure gives bone and the skeletal system in general its remarkab le ab ility to withstand the various mechanical and structural loads encountered during normal and extreme physical activity[51]. Typical mechanical and physical properties of bone are presented in Table 1 along with other biocompatible materials that are currently being used or under investigation. When an implant is used within the skeletal structure for load bearing applications, its mechanical properties should be as close as possible to those of natural bone, since both will be exposed to the same external loads. If there are significant differences in the mechanical properties, with the imp lant having the greater loading bearing capacity, the resulting mechanical mis match produces a much lower stress distribution in the surrounding bone. The lower stress or stress shielding, results in necrosis of the surrounding bone tissue, which ultimately leads to imp lant failure[52, 53]. Th is is why it is very important that the bio material used for the imp lant should have sufficient strength to withstand the stress created by the load, have sufficient stiffness to resist the effects of deformat ion without failure and have its mechanical properties match the surrounding bone tissue[54]. The type of stress the implant will experience is dependent on the application of the load. The loads that can be encountered singularly or in co mbinations are tensile, compressive, shearing, and tensional. Fi gure 1. Hierarchical st ruct ure of bone represent ed by different levels of organizat ion from Bonzani et al.,[42] 152 Gérrard Eddy Jai Poinern et al.: Nanometre Scale Hydroxyapatite Ceramics for Bone Tissue Engineering Table 1. Some typicalmechanical properties of bone tissue and a selection of several Tissue/Mat erial Density (g cm-3) Compressive St ren gth (MP a) Natural Materials Arterial wall Collagen Collagen(Rat tail tendon) Cancellous bone Cortical bone Magnesium Alloys Pure magnesium AZ31 (Extruded) AZ91D (Die cast) Other metal alloys Cobalt-Chrome Alloys Stainless Steel Titanium Alloys Ceramics Synthetic- Hydroxyapatite Alumina Ceramics (Al2O3 80% - 99%) Polymers Po lymethy lmet hacry late (P MMA) Polyethylene- terephthalate (P ET) 1.0 – 1.4 1.8 – 2.0 1.74 1.78 1.81 7.8 7.9 4.4 3.05 – 3.15 3.30 – 3.99 1.12 – 1.20 1.31 – 1.38 1.5 – 9.3 160 Trans. 240 Long. 20 - 115 83 - 97 160 - 100 - 900 2000 – 4000 45 – 107 65 – 90 biocompatible materials[58-64] Tensile St ren gth (MP a) Elast ic Modulus (GP a) 0.50 -1.72 60 - 1.5 – 38 35 Trans. 283 Long. 0.001 1.0 3.75 – 11.5 0.01 – 1.57 5 - 23 90 - 190 45 241 - 260 45 230 45 450 - 960 480 – 620 550 – 985 195 - 230 193 – 200 100 – 125 40 – 200 - 70 – 120 260 – 410 38 – 80 42 – 80 1.8 – 3.3 2.2 – 3.5 2.3. Bone Repl acement Therapies Since the dawn of humanity, the organs and bones of the body have been subjected to a variety of medical complications such as damage, wear, d isease or just the normal ag ing process. Unfortunately, there are times when the natural healing processes of the human body are unable to effectively repair and heal the damage or disease to the organs and bones. It is under these circumstances that med ical intervention is required to correct or alleviate the problem with a suitable procedure that returns functionality and removes the pain associated with damage. Historically, the demand fo r b io materials started with the Egyptians over 4000 years ago with the use of elephant’s tusks to manufacture art ificial legs and teeth. Since then many civilizat ions such as the ancient Chinese, Indians, Greeks, Ro mans and Aztec’s have used gold for dentistry, gold wires for repairing fractures, glass for artificial eyes and wood for artificial legs and teeth[55-57]. Today, medicine has been able to extend the average life expectancy in many countries and in turn has increased the average lifespan of the population. One significant problem that has arisen from the aging population is the need fo r suitable bone replacement therapies. These bone replacement procedures can range from bone grafts needed to fill bone cavities resulting fro m infections and tumours resulting from disease, reconstructive surgery to repair bone damage resulting fro m accidents or total joint replacements to restore a patient’s mobility after their own joint has failed due to disease, injury or wear. 2.3.1. Bone Transplantation Apart fro m b lood, bone is the most frequently transplanted tissue. Historically the preferred bone replacement procedure involved the use of autologous bone tissue (autograft), since it displays both excellent biocompatibility and osteogenic properties. The use of autograft’s is well established with good clinical results, which makes this procedure the gold standard for bone transplantation. There are basically two types of bone transplant procedure carried out; the first is small frag ment, in which the frag ments are used to induce osteogenesis. The second uses larger bone fragments, which are used to compensate for the large bone loss produced by the removal of tumours, bone cysts and traumas[65]. Every year, hundreds of thousands of people worldwide require a bone transplant[66, 67]. But problems such as donor site morb idity and limited supply sources have led researchers to investigate alternative sources of bone material. Bone tissue can be sourced from a suitable donor of the same species, which is called an allograft or it can be sourced fro m another species, xenograft. Medical procedures that uses xenogenic and allogenic bone grafts generally result in a significant response from the body’s immune system. In addition, there are also problems associated with the possible transmission of pathogens to the donor fro m the graft that needs to be addressed. In addition, obtaining bone tissue for an allograft and/or a xenograft is co mplicated, (medically, ethically and legally) and the grafts are expensive to process for human use. Because of these concerns there has been a great deal of interest in using man-made materials commonly called bio materials for use in bone tissue engineering applications. American Journal of Biomedical Engineer ing 2013, 3(6): 148-168 153 3. Biomaterials as Bone Substitutes An alternative to using natural bone grafts is to use a biocompatible synthetic material that possesses unique properties that can affectively replace bone tissue. To do this the material needs be biologically co mpatible, i.e. it should be nontoxic to the body tissues, non-immunogenic and should be chemically stable at body temperature and pH[68]. Other important properties include a high surface area for maximu m cell coverage; the surface chemistry should maximize cell and tissue adhesion, support the transfer of body fluids and promote vascular activity, the material should be cable of being sterilized to remove any contamination. And finally the material should also be mechanical strong enough to withstand normal physical loading encountered in vivo[69, 70]. Bio materials can be classified into four main groups; metals, poly mers, ceramics and composites. A short discussion of each is presented below for comp leteness, before a detailed discussion of nanometre sized hydro xyapatite (nano-HAP) in section 5, which is the focus of this research. 3.1. Metals and Metal Alloys Metallic materials such as stainless steel, cobalt-chromiu m based alloys (Co -Cr), t itaniu m (Ti) and its alloys have been widely used for decades to replace failed hard tissues. The high elastic modulus and yield strength, coupled with good ductility and fracture toughness makes metallic materials the ideal choice fo r high load bearing applications[71]. Because of these superior mechanical properties metal alloys have been used to manufacture total hip joints, bone plates, dental co mponents, pins, screws and a variety of load bearing imp lants. Unfortunately, the elastic modulus and strength of metallic imp lants is much higher than that of natural bone tissue. Depending on the metal alloy used the elastic modulus maybe as much as 10 times greater than that of bone in the case of Co-Cr alloys and as much as 5 times for Ti alloys[72]. The mis match between the mechanical propert ies of the implant and the surrounding bone tissue results in stress shielding, wh ich in turn produces bone resorption and imp lant failure. In addition, because metallic materials are bio-inert, they do not biologically or chemically interact with the surrounding tissues and as a result there is very little interfacial bonding or osteointegration taking place[73]. One technique currently being used to imp rove the osteointegration of metallic implants is to coat it with a bioactive material such as HAP. The main d isadvantages of metallic imp lants are; they are difficult to manufacture, are heavier than bone and they tend to suffer the effects of corrosion induced by the chemical environ ment of the body flu ids. Furthermore, the corrosion bi-products of the metallic implants are to xic and induce an immune response. This is why the current generation of metallic implants are being manufactured fro m Ti alloys; they have a low elastic modulus and have a higher resistance to corrosion than conventional stainless steel and Co-Cr alloys[74]. In addition, Ti can be anodized to produce a porous oxide layer that can be used directly to improve integration with the surrounding tissues or be used as an anchoring structure for a coating of bioactive material that can integrate with the surrounding tissues[75, 76]. 3.2. Pol ymers The application of b io materials in tissue engineering for tissue repair and regeneration has generally favoured bio-inert materials for permanent bio-imp lants such as hip-replacements. And in the case of scaffolding materials, both natural and synthetic polyme rs have been investigated. Poly mers are long chain mo lecules co mposed of repeated small simple chemical mono mer units. Poly mers have low densities, are flexib le, resilient and are surface modifiab le which makes them suitable for a variety of applications[77]. Typical applications of polymers are tissue scaffolds, skin augmentation, breast imp lants, denture bases, hip joints, blood vessels, cartilage replacement, sutures, bone screws, pins, plates and drug delivery devices[78]. The operational demands placed on polymeric materials used in tissue scaffolds or imp lants are numerous and there are many challenges to overcome to achieve successful clin ical outcome. For examp le, bioco mpatibility is crucial in preventing any cytotoxicity, immunological reactions and inflammat ion responses fro m the body[79-81]. In addit ion, the scaffolding material must be able to cope with the mechanica l stresses resulting from the growth stages as well as prevent any rapid bulk degradation crating voids. The material must also be easily sterilised prior to application without any significant changes to its surface chemistry [82-84]. At the molecu lar level, this is extremely important as the surface’s topography and chemistry are key factors in gaining the correct response fro m the cells, which will in return promote cell adhesion and proliferat ion [85-88]. The findings of Andersson et al suggests that in the case of epithelial cell attach ment to surfaces of similar chemistry, the morphology and cytokine production are strongly dependent on the underlying nanometre sized topography [89]. Poly meric materials can be manufactured fro m many natural sources or fro m synthetic materials. Natural biodegradable materials have been extensively investigated for use in t issue engineering since the body’s natural pathways can easily handle the breakdown of their metabolic by-products. Natural poly mers such as polysaccharides[90-94], chitosan[95-99], hyaluronic based derivatives[100-103] and protein-based materials such as fibrin gel[104, 105] and collagen[106, 107], have shown favourable outcomes. On the other hand, synthetic biodegradable polymers have been fabricated under controlled conditions to produce scaffolds with tuneable, predictable mechanical and physical properties. And being based on simple h igh purity constituent monomers, these biopolymers have lower to xicity reactions with the body and their degradation rate can easily be controlled. Examples of bulk b iodegradable poly mers include Po ly 154 Gérrard Eddy Jai Poinern et al.: Nanometre Scale Hydroxyapatite Ceramics for Bone Tissue Engineering (lactic acid), PLA[108-113], Po ly (L-lact ic acid ), PLLA [114-117], Poly (lact ic-co-glycolic acid), PLGA[118-121], Poly-caprolactone PCL[122-125] and Po ly (g lycolic acid), PGA[126-129]. These are generally poly-α-hydro x esters that de-esterifies in the body as the polymer degrades to simp le metabolites[130]. Currently available biodegradable sutures in clinical use are made fro m PLA and PGA. These synthetic polymers can also be made into different shapes and structures, such as pellets, disks, films and fibres as required for the specific application. Due to the fact that the extracellular matrix (ECM) has a fibrous nature and has features at the nanometre scale or sub micron level, engineering materials mimicking the ECM is the ultimate goal of many research teams wo rld wide. Two recent developments in nanotechnology, has been the refinement of electro-spinning and phase separation techniques to produce polymeric nanoscaffolds[131-134]. Initially developed by the text ile industry, electro-spinning has been used for the past 100 years. Refinements in the past decade have seen this technique being used increasingly in the manufacture of nanometre sized fibrous polymer scaffo lds. Membranes of PLA and PLGA, fibro in and collagen have been made using this technique. As shown by Agarwal et al[135], this technique is favourable and conducive for cells attaching and pro liferat ing onto this material. Th is nano-fibrous electro-spun material can be further bioengineered to resemble the ECM at the nanometre scale level by coating the material with collagen macro molecules. This technique is still evolving and there are still challenges ahead. Other nano-polymeric substrates have been manufactured using phase separation techniques that produce a film o f the desired polymer. Recent developments to this technique have produced nanometresized features over its surface to enhance the value of this material[136-138]. The studies by Ma have shown that nano-polymers by this technique have an advantage in terms of surface area and enhanced 3D connectivity for tissue engineering[139]. Furthermore, poly meric materials are effective bioco mpatible materials that have also been extensively investigated for the controlled delivery of drugs to specific organs within the body[140-142]. Although polymers have been effectively used in a number of tissue engineering applications, polymers have a lower mechanica l strength compared to metals and cera mics, so they are generally used predominantly in low load bearing applicat ions such as soft tissue engineering. In addition there are still issues to be resolved like local inflammat ion resulting fro m the degradation products and the uneven bio scaffold degradation, which can disrupt the gradual integration of the surrounding tissues. Because of this, alternate scaffold and imp lant materials are currently under investigation. For examp le, recent studies using bioactive glass as a scaffold; revealed that when the glass was seeded with osteoblasts, a positive effect could be seen in cell pro liferat ion[143]. 3.3. Biological Ceramics Ceramics are non-metallic and inorganic materials that are used in hard tissue engineering applications where they are collect ively termed as biological ceramics or bioceramics. The advantages of using bioceramics are; they are physically strong, both chemically and thermally stable, provide good wear resistance and are durable[144]. In addition, these materials are readily available, can be shaped to suit the application, they are also bioco mpatible, hemoco mpatible, nontoxic, non-immunogenic and can be easily sterilised[145]. The only d isadvantages of using bioceramics are, they tend to be brittle, have lo w fracture toughness and are not resilient. But they have found application in hip jo ints, coatings on imp lants, maxillofacial reconstruction, bone tissue engineering and drug delivery d ev ices . Bioceramics can be categorized into four types, which are dependent upon the tissue response they solicit. The first type is bio-inert ceramics, (i.e. alu mina and zirconia) which are generally used in dental imp lants and orthopaedics [146,147]. These materials do not biologically or chemically interact with the surrounding tissues; hence they don’t solicit a noticeable response fro m the t issues. The second category is bioactive (i.e. bioactive g lasses, hydroxyapatite and glass ceramics), these materials interact with the surrounding tissues and induce a strong osteointegrative response[148]. A typical applicat ion of these materials is the coating of titaniu m alloy imp lants to improve their osteointegration with the surrounding tissues [149]. The third type is resorbable ceramics, (i.e. t ricalciu m phosphate (TCP)) in this case the TCP forms a scaffold structure for new bone cells to adhere to and proliferate [150,151]. During this process the new bone tissues continue to grow and rep lace the scaffo ld structure as it is slowly resorbed[152]. And the final type is porous; this type of bioceramic permits the surrounding host tissues to penetrate into the pores of the implant (HAP coated metal and alumina imp lants). Because of the attractive features and properties of bioceramics there has been a significant effort into researching and developing new synthetic biomaterials such as porous coralline, calciu m sulphates, and calciu m phosphates, i.e. HAP[153-155]. The importance of calciu m phosphates as bioceramics is discussed in detail in section 4 and particular attention is paid to HAP in section 5, a member of the calciu m phosphate family and the focus of this review. 3.4. Composites A composite material consists of two or more distinct parts or phases[156]. The major advantage of using composite biomaterials stems from the fact a single phase material may not have all the required properties for a particular applicat ion[157]. Ho wever, by co mbining this phase with one or mo re other phases with differing physical and chemical properties it is possible to create a co mposite material with superior propert ies to those of the indiv idual components. A good example is bone, which is a co mposite of Type 1 co llagen and HAP, and an examp le of a man made American Journal of Biomedical Engineer ing 2013, 3(6): 148-168 155 composite is coating a titanium imp lant with a bioactive material such as HAP or b ioactive glass to promote bone attachment[158]. Similar co mposites have been developed to produce biomaterials for hard t issue engineering applications that have used a 2 phase HAP-poly mer mixtu re to produce a composite with a modulus of elasticity close to natural bone[159]. 4. Calcium Phosphates Ceramics The inorganic phase of bone is composed of a calciu m phosphate (CaP) co mpound, namely, HAP which forms crystal platelets, a description of which was presented in section 2.2. (Bone: Its hierarchical structure). CaP compounds exist in several phases which can be discriminated by the stoichiometric ratio of calciu m to phosphate (Ca:P) which is important for crystallin ity, solubility and strength[160]. There are nu merous Ca:P compounds, each with their own unique crystalline structure and Ca:P ratio, wh ich means that they have different properties and form under different conditions[161]. For example, the stability of CaP co mpounds is influenced by controlling the Ca:P ratio, water p resent, formation temperature and pH[162]. The Ca:P ratio varies fro m 0.5 to 2.0, see Table 2, however, for rat ios below 1, the solubility rates and the acidity are both high making these CaP compounds unsuitable for bio med ical applications. Several members of CaP family have been used in a variety of tissue engineering applications to date. In addition, CaP compounds have been combined with collagen, to form composite biomaterials that display some similarity to natural bone tissue[163]. Amorphous calcium phosphate (ACP) is the least stable solid phase of the CaP family formed in solution[164]. It is usually the first phase formed by precipitation in solutions containing high concentrations of calciu m and phosphate. ACP has a Ca:P ratio of around 1.5, generally has spherical morphology and lacks any internal crystalline structure normally associated with CaP co mpounds[165]. It is usually considered the precursor crystalline CaP co mpounds, since heating can convert ACP to poorly crystalline apatite at 600℃, tri-calciu m phosphate (TCP) at 800℃ and HAP around 800℃ respectively[166]. ACP has been used in mineral releasing composites for dental fillings, bone cements and for non-load bearing imp lants in bone tissue engineering[167]. However, the two most pro minent calciu m phosphate bioceramics currently being used in a variety of medical procedures are TCP and HA P. Both materials are biocompatible, osteconductive and directly bond with natural bone[168, 169]. TCP is a polymorph material with two distinct phases (α-TCP and β-TCP) each phase is formed by varying the humidity and the sintering temperature. The α phase is formed in dry heat with temperatures above 1300℃ then quenching in water, while the β phase is produced in a humid at mosphere during the sintering process[166]. There are also significant differences in the propert ies, for examp le β-TCP is less soluble and less reactive than α-TCP which is unstable in water and reacts to produce HAP. A major advantage of using β-TCP in vivo is that its solubility, dissolution and re-precip itation results in a gradual phase change into carbonated apatite. The resorption of the carbonated apatite by the macrophages, giant cells and osteclasts allows the β-TCP to be gradually replaced by natural mineralized bone tissue. Both α-TCP and β-TCP have been used in bone tissue engineering applications; α-TCP is mainly used in some CaP cements while β-TCP has been used in bone cements and fillers, low load bearing implants, used to provide a degradable coating on metallic imp lants to induce a favourable biological response and increase oseointegration of the implant[170, 171]. Table 2. A selection of common calcium phosphate compounds,[172, 173] Ca: P rat io 0.5 Chemical formula Ca(H2P O4 )2 .H2 O Compound Mono calcium phosphate mono hydrate (MCP M) pH stability range in aqueous solutions at 25ºC So lubilit y at 25ºC (g/L) 0.0 – 2.0 18 So lubilit y at 25ºC -log(Ks) 1.14 1.0 CaHP O4.2H2O Di-calcium phosphate Di-hydrate (DCPD) 2.0 – 6.0 0.09 6.59 1.5 α-Ca3 (P O4)2 1.5 β-Ca3(PO4)2 α-Tricalcium phosphate (α-TCP ) β-Tricalcium phosphate (β-TCP ) * ~0.0025 25.5 * ~0.0005 28.9 1.67 Ca1 0(P O4 )6(OH)2 Hydroxyapatite (HAP) 9.5 – 12 ~0.0003 116.8 * Compound cannot be precipitated from an aqueous solution 156 Gérrard Eddy Jai Poinern et al.: Nanometre Scale Hydroxyapatite Ceramics for Bone Tissue Engineering 5. Hydroxyapatite and Nanometre Scale Hydroxyapatite 5.1. Hydroxyapatite Due to the close chemical similarity of synthetic HAP to the natural inorganic bone matrix co mponent, there has been an extensive research effort to emp loy synthetic HAP as a bone substitute or replacement in several clinical procedures[174, 175]. There are four main of advantages of using synthetic HAP; firstly, it has good biocompatib ility to surrounding body tissues, secondly its biodegradability in situ is slow, thirdly it provides good osteoconductivity and finally it has good osteoinductivity capabilit ies[176-179]. An investigation by Taniguchi et al[180] revealed that sintered HAP p rovided a good bioco mpatible response to soft tissues such as skin, muscle and gums. It is this favourable tissue response that makes synthetic HAP an ideal candidate for orthopaedic and dental imp lants. This is the reason why synthetic HAP has been widely used for a variety of hard tissues applications such as; bone repair, bone augmentation, the coating of metal imp lants and used as a filling material in both bone and teeth[181,182]. Despite its many advantages, pure HAP has unfortunately very poor mechanical strength. As a result, pure HAP ceramics are restricted to low load-bearing applicat ions. However in some cases, comb ining HAP with other materials such as polymers and/or glass to form a co mposite material can alleviate these inherently low mechanical properties. For example materials such as high-density polyethylene (PE) and polypropylene have been successfully used to improve the load bearing capabilit ies of HAP[183, 184]. Another significant advantage of using HAP as a component of a co mposite material is that its co mplex structure provides a good absorption matrix for other mo lecules. The absorption properties of HAP have been exploited fo r a nu mber o f in situ applications, such as HAP-antib iotic and HAP-drug co mposites for slow release[186-188]. These composites have also proved to be effective in the treatment of diseases such as osteomyelitis, where the slow drug release has resulted in a successful recovery fro m the disease[189]. Figure 2. The unit cell of hydroxyapatite in P63/m space group[185] 5.2. Nanometre Scale Hydroxyapati te Research to date in both nanoscience and nanotechnology has highlighted the need to investigate the formation of HAP in the nano metre size range and to clearly define the properties of HAP at this scale. This is very important in nanotechnology because different forms of a material at the nanometre scale can result in significant differences in its physicochemical properties[190-193]. Parameters such as the structure of the HAP and how it is processed can influence the properties. For examp le when the parameters are adjusted it is possible to change the particle shape of the HAP being produced. This means that it is possible to produce rods, tubes, plates and spheres[46, 194-196]. Major improvements in the properties of HAP can be achieved when the material is synthesised in the nanometre scale [197-200]. The improved properties of the HAP can also have a significant effect in interacting with surrounding body tissues[201, 202]. For examp le, in a HAP particle study conducted by Sun et al[197], the influence of particle size on in vivo osteoblast cells found that the inflammatory response was inhibited when smaller part icles sizes (0.5 – 3.0 µm) were used. Furthermore, Hu et al was able to demonstrate that nano-HAP particles used in their study displayed a strong anti-tumour effect, and in part icular the nano-HAP particles were ab le to inhibit the growth of hepatic tumour cells[203]. In a similar study, Li et al were able to demonstrate the beneficial effect of a nano-composite material co mposed of HAP and chitosan on human gastric cancer cells[204]. Similar studies have also looked using nano-HAP as a delivery platform for the slow release of drugs and gene therapies for the treat ment of tumours[205, 206]. In a co mparable study using nano-HAP impregnated with an antibacterial or antifungal material, it was possible to effectively inhibit the growth and proliferation of infect ious microorganisms[207]. Bioceramics composed of nano-HAP are currently being researched and in some cases used in a variety of medical procedures other than drug delivery. For example, nano-HAP has been extensively used in a range of bone cements and fillers which are used to provide a scaffold structure that encourages the growth of natural bone forming tissue and form an effective alternative to bone grafts in a nu mber o f procedures[208]. Currently many orthopaedic imp lants are based on metallic materials such as stainless steel and titanium alloys[209]; unfortunately these materials are bio-inert and do not interact or form direct bonds with the surrounding bone tissue. The use of nano-HAP has opened new opportunities in designing biocompatible coatings for metallic imp lants. For example, the use of plasma spraying techniques to create thin coatings that can deliver controlled thicknesses of nano-HAP that can create an effective environ ment for osteoconduction, protein adhesion and bone in growth[210]. Another interesting feature of using nano-HAP as a scaffold is that it can be pre-treated to produce micro meter sized American Journal of Biomedical Engineer ing 2013, 3(6): 148-168 157 cracks to simulate the cracks naturally found in bone tissue that results from repetit ive stress injuries such as running or repeated lifting of heavy loads. As in the case of natural bone, the cracks in the nano-HAP can pro mote bone re-growth[211]. The advantage of using nano-HAP lies in its ability to integrate with bone tissue, promote and support bone in growth. In fact, one interesting property of nano-HAP is that when it bonds with bone, it fo rms an indistinguishable union with the surrounding bone tissue. Unfortunately, there are some disadvantages of using nano-HAP; they range fro m the in itial synthesis process, where the composition of the final product can be greatly affected by even small changes in the init ial reaction conditions such as pH and formation temperature. Furthermore, nano-HAP cannot be used in bulk form to a make a load bearing orthopaedic implant because the HAP has low mechanical properties such as strength and fracture toughness. Also the solubility and bioactivity o f nano-HAP is influenced by its phase purity, ionic purity and its crystalline structure. To overcome the d isadvantages in the properties of pure nano-HAP, in particular its mechanical strength, composite materials composed primarily of nano-HAP co mbined with a variety of other materials with desirable properties have been investigated. For examp le, organic materials such as collagen, gelat ine, chitosan and poly (lactic acid) (PLA) have been used to reinforce the nano-HAP matrix, in a similar way collagen does in the natural bone matrix, and significantly improve the mechanical p roperties[212-216]. Also adding nano-HAP into a polymer matrix has been found to turn a non-bioactive polymer into a bone bonding composite with improved mechanical properties such as elastic modulus and hardness[215]. The possibility of making a nano-HAP co mposite as strong as cortical bone with good biocompatibility, slo w b iodegradability in situ, and osteoconductivity and osteoinductivity capabilities by adding an extra reinforcing phase or phases makes the composite a very attractive alternative[213]. 5.3. Synthesis of Nanometre Size Hydroxyapatite Traditionally, several procedures have been developed and used to produce HAP and calciu m phosphate ceramics. These diverse procedu res include ho mogeneous precipitati on[217, 218], sol-gel[219, 220], spray dry method[221], plasma spray[222], hydrothermal[223, 224], reverse micelle [225] and ultrasonic spray freeze drying processes[226]. Of these the sol-gel is the most appealing method since it is based on a wet chemical technique that has a simp le and straight forward p rocedure that can economically produce HAP without rely ing on expensive specialised equipment. Furthermore, this technique can be scaled up to meet high demands. However; the main problem in using this technique is being able to effectively fine-tune, with in a small parameter range, the size and morphology of the nano-particles produced. This fine-tuning of the size and morphology is crucial in determining the properties of the resulting ultrafine nano-HAP powders. Crystalline materials can be produced from solutions, using wet chemical techniques, but a subsequent thermal treatment at elevated temperatures is required to produce the specific crystalline phases. Both the particle size and morphology of HAP produced using this technique can be efficiently controlled by vary ing the experimental parameters that regulate the nucleation, the aging p rocess and the growth kinetics of the forming particles. The parameters that control the production of mono phase HAP are the initial reactants, the preparation temperature and the pH value. Both Khopade et.al[227] and An et.al[228]; using a wet chemical method were able to produced HAP particles that with a plate like structure or morphology. This structure is in complete contrast to the spherical-shaped HAP part icles that are normally synthesized when ultrasonic irradiation is used during the processing stage. The most commonly used wet chemical technique is the precipitation method and it can be used to produce homogeneous or inhomogeneous calcium phosphate ceramics[229]. It is important to note that variables such as Ca:P ratios, structural defects, crystal size, temperature and the preparation procedures can have a significant impact on the electronic properties of the manufactured HAP[230]. A number of chemical routes have been used to manufacture HAP; each p roducing a significant variation fro m the normal HAP phase, and this clearly demonstrates that the HAP produced is procedure dependent[231]. Furthermore, different HAP phases can be produced by the influence of other molecules present in the manufacturing process. Wang et al successfully studied the in fluence of other mo lecules as molecular temp lates, for producing nanometre sized HAP rods[232]. Guo and Xiao studied the properties of nano-crystalline HAP produced during a hydrothermal process and their investigation revealed that the particle size decreased when the thermal treat ment temperature increased[233]. In a similar investigation by Meissner et al. found that both size and morphology of the produced HAP could be controlled, both particle size and morphology were dependent on the precipitation temperature and the ultrasonic power used[234]. The part icle size and temperature relationship has also been investigated by Laquerriere et al who were able to demonstrate that increasing thermal treat ment temperatures produced a decreasing particle size[235]. In addition, Li-Yun et al confirmed the dependence of the particle’s size and morphology under ultrasonic irradiat ion, as they prepared a mono-phase nano-HAP material using a 300W ultrasonic transducer[236]. An important procedure carried out during sample preparation is wet milling. Ultrasonic irradiation is used during the wet milling procedure as an efficient means of dispersing and de-agglomerat ing the sample part icles during the grinding process. The sonochemical effect, which produces acoustic cavitations, also promotes both chemical reactions and physical effects that directly influence the particle mo rphology during the growth phase. There are a number of advantages in using ultrasonic irradiation during

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