Biomimetic bone regeneration

8 Biomimetic bone regeneration K. A. HING, Queen Mary University of London, UK DOI: 10.1533/9780857098887.2.207 Abstract: This chapter reviews basic bone biology and bone graft technology to give an insight into the increasingly biomimetic approach being pursued in this field. Many current synthetic bone graft substitute materials have calcium phosphate chemistries reflecting the composition of bone mineral, and hierarchical pore structures similar to cancellous bone. Significant research effort has been devoted to understanding how slight variations in chemistry and pore structure impact a graft material’s ability to support or stimulate bone healing. It appears that the best results are obtained when a graft chemistry or structure is designed to mimic the natural tissue. Key words: bone graft substitute, bioactive chemistry, hierarchical pore structure, osteoconductive, osteoinductive.

8.1

Introduction

The first scientific documentation of the basic principles behind modern biomimetic bone regeneration dates back to the late 1700s with the pioneering work of Dr John Hunter. During his groundbreaking studies into the immune system, bone repair, regeneration and remodelling,1 he performed many grafting studies using a variety of tissues, performing possibly the first systematic study of bone autografting (the use of the patient’s own bone taken from a donor site to augment or repair a bone defect or fracture) on a group of cockerels.2 However, he was well ahead of his time and the world had to wait 100 years for Lister to develop antiseptic surgery for the potential of his work to be fully appreciated. Bone grafting, the procedure of replacing missing bone with material from either the patient’s own body (autografting) or that of a donor (allografting), was not established until the 1800s,3 becoming clinical mainstream in the early 1910s.2 Evidence for the use of artificial, synthetic or natural substitutes, however, predates this in the form of gold and silver plates and pieces of coconut shell found in cranial defects within prehistoric skulls.4 Furthermore, archaeological studies of the skeletons of ancient Egyptian mummies have demonstrated the successful practice of external fracture fixation using splints made of bamboo, reeds, wood or bark, padded with linen. Thus, the use of natural templates in the repair of bone has a long history in human medicine.

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8.2

Demand and supply: clinicians, engineers and biologists

When presented with a patient with broken or damaged bone tissue, clinicians have the task of selecting the best strategy to efficiently return pain-free function to a patient. However, with the exception of Imhotep who was both a physician and engineer (and later a God), development of strategies to facilitate bone repair and regeneration has traditionally been studied from a number of distinct perspectives – as a biomedical engineering challenge to restore efficient loadbearing function (by engineers) or as the restoration of structurally or metabolically ‘normal’ tissue that is physiologically and biomechanically responsive (by clinicians and biologists). Although both these approaches can exploit bones’ natural capacity for adaptation and repair, when taken to extremes they can result in very different recommendations for treatment, ranging from amputation and the fitting of an osteointegrated state-of-the-art limb prosthesis, to the introduction of powerful growth factors that stimulate bones’ natural repair cascade. To date, history would suggest that when engineers, biologists and clinicians work together, or borrow from each other’s fields to take a multi-disciplinary biomimetic view to bone repair, then the best results are achieved. A classic example of this approach was that of Sir John Charnley, whose pioneering work in the development of ‘low friction’ total hip replacement (THR) underpins clinical practice in modern articulating joint replacement. Charnley applied the interest that he had developed in mechanics of the hip while at medical school, together with his evidence that ‘the coefficient of friction of normal articulate cartilage is phenomenally low and in fact lower than anything encountered between solid substances in engineering practice’, to advance the concept that if a damaged or fractured femoral head was to be replaced with a material selected for strength then the opposing damaged articulating surface should also be replaced so as to mimic cartilage and possess a low coefficient of friction. This resulted in the introduction of total hip arthroplasty in which the joint was completely replaced with a low friction polymeric acetabula cup to mimic nature, coupled with a metallic femoral component that was designed to possess the smallest diameter ball still able to cope with the expected loads so as to ‘fit in with best engineering practice’. He then further refined his surgical technique with the addition of polymethyl methacrylate (PMMA) cement to act as a grout to mechanically fixate the prosthesis in place and ‘distribute the load of the body weight over a wider area of bone’. The fundamental concepts of combining low friction articulation, load transfer in the physiological range and prosthesis stability still direct modern advancements in THR; however, even with additional understanding of the importance of sub-micron wear particle generation, average prosthesis life spans are still considered to be around 10–15 years. Replacement of living bone with monoliths of inert metal or polymer is clearly not ideal.

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8.2.1 Bone structure, biology and function Bone is a complex living tissue with an intricate hierarchical structure (Fig. 8.1) that performs several key functions within the body. Not only does bone provide structural support and protection to bodily organs, but it is also involved in maintaining mineral homeostasis (i.e. acting as a mineral reservoir for the rest of the body) and providing a source of mesenchymal and haematopoietic stem cells. From a materials engineering perspective, bone is a (often anisotropic) porous composite containing bone cells and blood vessels embedded in a bi-phasic matrix of organic (collagen fibres, lipids, peptides, proteins, glycoproteins, polysaccharides and citrates) and inorganic (calcium phosphates, carbonates,

8.1 Complex, hierarchical structure of cortical and cancellous bone and the systematic distribution of living cells and blood vessels that maintain and sustain it.

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sodium, magnesium, and fluoride salts) elements.5 The main component of the organic phase is a highly elastic protein, collagen, and the main inorganic component, bone mineral, is a highly substituted form of carbonated hydroxyapatite. It is the combination of mineral and collagen that gives bone its hardness and toughness. Moreover, its ‘macro-structure’ and the proportion of its components differ widely with age, site and history, resulting in many different classifications of bone that exhibit very different mechanical and functional characteristics; the principal ones being cancellous (or spongy) and cortical (or dense) bone. Bone surfaces consist of cortical bone and the thickness of this protective skin increases in mechanically demanding regions such as the shafts of long bones, whereas cancellous bone is found in the interior of bones, such as within the femoral head, and vertebra. There are two kinds of cancellous bone, coarse and fine. Coarse cancellous is characteristic of healthy adult mammalian skeleton, whereas fine cancellous bone is characteristic of the fetal skeleton or early fracture callus and comes in two forms, dependant on the route of osteogenesis, fine cancellous membranous bone (bone formed de novo) and fine cancellous endochondral bone (bone formed from a cartilaginous template). There are also several types of cortical bone: surface, primary and secondary osteonal cortical bone, and, as with the cancellous bone, the distinctions are dependent on the age and origins of the bone. Furthermore, the microstructure of bone also varies with its rate of formation, rapidly formed bone, such as that formed in fracture callus, often has a disordered, less dense structure and is known as woven bone. Whereas bone formed more sedately during normal growth or remodelling has a more ordered dense structure and is known as lamellar bone as a result of its striated structure. This variation in structure leads to considerable variation in stiffness, strength and toughness in both cortical and cancellous bone. The osteonal microstructure of cortical bone makes it highly anisotropic, although its density is relatively consistent (1.85–2.05 g.cm−3 in human bone). The mechanical properties of cancellous bone (which can be considered to be a foam) are highly dependent upon porosity and architecture, both of which vary widely with anatomic site6 and age.7 In addition, cancellous bone is often anisotropic because of the orientation of major trabeculae along lines of principal stress. Despite its complex structure, bone also possesses an innate regenerative capacity and is capable of maintaining an optimal shape and structure throughout life via a continual process of renewal. This characteristic enables bone to respond to changes in local mechanical environment by ‘remodelling’ to meet different loading demands. Remodelling is the term used to describe the phenomenon in which old bone is sequentially removed by phagocytic cells, including osteoclasts – specialist boneresorbing cells, and replaced with new bone by osteoblasts – committed boneforming cells. This process enables bone to maintain an optimal balance between form and function throughout life via a continual cycle of renewal, and is the key

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to the skeleton’s ability to grow, mature and meet different loading demands while maintaining an optimal balance between form and function.8 Remodelling is believed to be regulated by both internal hormonal and external mechanical demands. Consequently, bone is engaged in a constant cycle of resorption and renewal, undergoing continual chemical exchange and structural remodelling as directed by local cells sensitive to changes in their physiochemical and/or mechanical environment. These cells, known as osteocytes, are differentiated osteoblasts that reside in osteocyte lacunae within mineralized bone and are responsible for the maintenance of the surrounding tissue. However, like any engineering material, bone will fracture spontaneously when overloaded, and, as a living tissue, it also requires a constant supply of oxygen and nutrients, is subject to infection, degenerative (age-related), metabolic and metastatic disease. Although it possesses the ability to repair small fractures or defects without external intervention, it is limited in the size of fracture or defect it is able to restore to healthy tissue. Alternately, where there is an imbalance in the body’s normal hormonal regulatory system which results in metabolic bone disease and either depletion (osteoporosis) or overproduction (Paget’s disease) of bone, spontaneous restoration to healthy tissue is unlikely. For these reasons clinical intervention is sometimes necessary, which can take the form of external or internal stabilization to facilitate bone regeneration or the introduction of medical devices to facilitate bone replacement.

8.2.2 Bone regeneration and bone replacement Bone regeneration is generally accepted as being preferable to bone replacement in the long term. Despite the best efforts of many leading researchers and clinicians, the materials and devices available for bone replacement still fall short of ideal. One of the biggest hurdles in bone replacement is the mechanical mismatch between traditional orthopaedic biomedical materials such as stainless steel, cobalt-chrome and titanium alloys and natural bone tissue, leading to a phenomenon known as stress shielding. Stress shielding occurs when a high modulus (metal) prosthesis is attached to or implanted in bone, initially the metal component stabilizes motion or channels load providing either a protective environment for bone to heal around the prosthesis or an alternate mechanism for the bone or joint to perform its normal load-bearing or locomotive function. However, with time, the fact that the stiffness of the prosthesis is greater than that of natural bone results in the local tissue receiving a reduced biomechanical stimulus leading ultimately to resorption of bone, loosening of the prosthesis and the requirement for further surgical intervention in an environment now depleted of bone stock. The need for a more biomimetic solution with matched mechanics is clear; however, providing a biocompatible, low modulus material with sufficient fatigue resistance still presents a technical challenge to medical engineers and biomedical

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materials scientists alike, hence the continued use of metal-based bone replacement strategies in mechanically demanding indications such as total hip replacement, to date. Bone regeneration on the other hand makes a virtue of the fact that bone is an adaptive living tissue and harnesses bone’s natural ability to heal itself. The hurdles to this approach are the limited volume of tissue that bone can spontaneously regenerate and the rate at which the regeneration of fully functional load-bearing tissue occurs. These hurdles place an interim requirement for a support or scaffold that, at worst, does not hinder the regeneration process, at best stimulates it. Again, a biomimetic approach, copying or borrowing the structures and materials utilized by nature seems an obvious solution and has proved highly successful.

8.3

Bone grafting: the ultimate biomimetic regeneration procedure

As previously mentioned, the concept of bone (or any form of tissue) grafting was around long before anyone coined the term ‘biomimetic’, and for this reason it is often not considered biomimetic but rather more as a form of tissue donation. However, although this may be partially true in the case of fresh autologous bone grafting which could be considered to be the transfer of live tissue, bone-grafting procedures using some form of allograft, xenograft or completely synthetic material are employing the use of a material that has been deliberately chosen and manipulated to have properties that mimic natural cortical or, more usually, cancellous bone.

8.3.1 Allografts, autografts, biologics and synthetics Bone grafting is the procedure of replacing missing bone with, traditionally, material from either the patient’s own body (autograft) or that of a donor (allograft – human bone obtained from a bone bank),3,9 and is utilized in severe trauma cases, oncology, total hip revisions and in the correction of large ‘bony defects’, where a significant piece of bone is missing or damaged. Bone grafts are generally used in combination with fixation devices (both temporary and permanent) to ensure adequate mechanical stabilization. Ideally, the graft should not only replace the missing tissue but encourage new bone ingrowth into the grafted area, thereby reinforcing the repaired area and forming a living bridge between the existing bone and the graft material. Moreover, with time, the graft should be replaced with healthy bone tissue via the normal bone remodelling process. Thus, bone grafting exploits bone’s unique properties, namely its ability to repair small fractures or defects without external intervention via the formation of fully functional new bone and its participation in a continuous cycle of regeneration, features that enable the skeleton to grow, mature and meet different loading demands while maintaining an optimal shape and structure.

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In modern medicine, autografting is regarded as the ‘gold standard’; however, the amount of bone that can be safely harvested is limited, and the additional surgical procedure may be complicated by donor site pain and morbidity. Modern allografting using material stored within regulated bone banks overcomes these difficulties; however, the demand far outstrips the supply, there is no assurance of freedom from disease10,11 and healing can be inconsistent.12 Consequently, there is an increasing demand for alternative therapies that would avoid these complications, in addition to overcoming the problem of an inadequate supply of material. This has led to the development of two approaches, the development of osteogenic bone grafts (collectively known as synthetics) and treatment with osteoinductive growth factors (collectively known as biologics). In tissue engineering these two approaches come together with the introduction of live cells, the aim being to generate live functional tissue ex vivo for implantation, while the delivery of biologics is usually achieved via a carrier scaffold so the delineation is somewhat artificial. Indeed, you could even argue that implantation of an osteogenic bone graft scaffold is, in fact, a less externally controlled form of tissue engineering, you are merely using the patient as the source of cells and growth factors and as an incubator. However, the biomimetic principle behind all these approaches is clear. The development of bone grafts has been driven by the desire to replicate the chemistry and/or structure of bone matrix so as to encourage bone formation within the graft material, whereas the development of biologics was inspired by the ability of certain natural proteins to stimulate bone formation when implanted in muscle tissue or to stimulate bone cell differentiation in vitro. Interestingly, the use of biologics has recently been associated with increasing levels of controversy as to whether or not their efficacy outweighs their risks. It could be argued that this may be related to the fact that the very high supraphysiological doses that have been used clinically depart somewhat from the biomimetic principles underling the initial concept behind using these proteins, and that their future in bone regeneration may depend on solving that dichotomy.

8.3.2 Biomimetic materials selection At an elementary level the fabric of bone may be split into three main components: bone matrix, bone cells, bone marrow and its associated vascular network. The bone matrix provides mechanical strength, acts as the body’s mineral store, and is made up from two main components: organic collagen fibres and inorganic bone mineral crystals. Together they make up approximately 95% of the dry weight of bone, the remainder being composed of other organic molecules (known collectively as the non-collagenous proteins, NCPs) and ‘amorphous’ or poorly crystalline inorganic salts. It is this combination of highly ordered elastic collagen fibres reinforced by sub-microscopic inorganic crystallites together with some latitude in composition and density at any one point, that enables bone to display

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a wide range of mechanical properties and to retain elasticity, toughness and hardness for a minimal weight. There is currently no artificial engineering material that can match the performance of bone gram per gram. Researchers have long recognized that it is the composite nature of bone matrix that is key to its success; however, we are only just beginning to understand how to manipulate materials at the sub-micron ‘nano’ level required for the level of control found in bone structure. Collagen is the most abundant protein found in the body, accounts for 70–90% of the non-mineralized component of the bone matrix and varies from an almost random network of coarse bundles to a highly organized system of parallel-fibred sheets or helical bundles. Collagen consists of carefully arranged arrays of tropocollagen molecules, which are long rigid molecules (300 nm long, 1.5 nm wide) composed of three left-handed helices of peptides (‘monomers’ of proteins composed of amino acid sequences) known as α-chains that are bound together in a right-handed triple helix (Fig. 8.2). Although all α-chains contain the glycineX-Y sequence, different types of collagen may be produced via the combination of different amounts and sequences of other amino acids within the tropocollogen molecule. To date, 16 different types of collagen have been identified. Bone contains mostly type I collagen with some type V collagen. Type I collagen is the most abundant form, accounting for 90% of the body’s total collagen; it contains two identical and one dissimilar α-chains (α1(I)2α2) within its tropocollogen molecule. Molecules of both types I and V are organized into collagen fibrils, which are formed by the assembly of tropocollagen molecules in a ¾ stagger, parallel array. As a result of this assembly, the fibrils exhibit characteristic crossstriations or banding which occurs in a repeating pattern every 55–75 nm, average 64 nm.13 The fibrils are stabilized by inter- and intra-molecular cross-links (the number and distribution of which determine whether the tissue will mineralize), and have individual diameters of 40–120 nm, average 100 nm. In type I collagen,

8.2 Hierarchical structure of collagen.

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the fibrils are wound into bundles to form collagen fibres that range in diameter from 0.2 to 12 μm. Significant research has been done on the use of collagen and gelatine-based scaffolds or composites; however, strength, sterility and disease transmission are always limiting factors in commercialization of these materials. Other biomimetic organic materials, such as chitosan and wood, are also receiving increasing interest as disease transmission is less of an issue with these compounds, and have delivered promising results. The main inorganic phase within bone is usually incorrectly referred to as hydroxyapatite (HA) a hydrated calcium phosphate ceramic, with a similar (but not identical) crystallographic structure to natural bone mineral,14 which has a chemical formula of Ca10(PO4)6(OH)2 and a Ca:P ratio of 5:3 (1.66). However, bone-apatite is characterized by calcium, phosphate and hydroxyl deficiency (reported Ca:P ratios of 1.37–1.87),15,16 internal crystal disorder and ionic substitution within the apatite lattice resulting in the presence of significant levels of additional trace elements within bone mineral (Table 8.1). It is not a direct analogue of HA as is commonly believed, but more closely related to an A-B type carbonate substituted apatite.17,18 These factors all contribute to an apatite that is insoluble enough for stability, yet sufficiently reactive to allow the in vivo submicron (5–100 nm) crystallites to be constantly resorbed and reformed as required by the body. The use of synthetic bone graft substitutes (BGS) have been considered for over 30 years,19 initially with the use of inert, relatively strong materials such as alumina or polyethylene sponges. However, recognition that a scaffold material needed to be more than biocompatible for optimal performance followed work that recognized that the performance of an orthopaedic implant was greatly enhanced by the use of a ceramic which supports direct bonding of bone to its surface20,21 and elicits a response similar to the normal healing cascade of bone within its porous structure.22,23 Unsurprisingly, the first ceramic found to exhibit these properties was HA, for which the fabrication of high purity, high porosity, foam-like structures presents a particular challenge.24,25 Despite this, there are a considerable number of apatite or calcium phosphate-based synthetic BGS on the market produced via a number of routes such as ceramic slip foaming,26 positive

Table 8.1 Typical levels of trace elements found in bone mineral Reference

Ca

P

Mg

Na

K

CO3

F

Cl

Sr, Zn, Cu

McConnel16 Driessens144 Aoki145 Le Geros18

26.7 36.7 34.0 24.5

12.5 16.0 15.0 11.5

0.44 0.46 0.50 0.55

0.73 0.77 0.80 0.70

0.06 – 0.20 0.03

3.48 8.00 1.60 5.80

0.07 0.04 0.08 0.02

0.08 – 0.2 0.10

Sr = 0.04 – – Traces

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replication of reticulated foam scaffolds,27 sol-gel processing, burn-out of sacrificial porogens such as polymer beads24 and techniques which exploit naturally occurring porous calcium-based structures, as in the hydrothermal conversion of either coral28 or bone.29 Unsurprisingly, both the chemical and physical properties, such as the mechanical behaviour and the pore architecture, of the various BGS are subject to considerable variation,24,28,30–32 making comparison between materials and identification of common trends a difficult task, exacerbated by the broad range of methods used to quantify the physical and chemical characteristics of a BGS let alone the biological response to it. However, most of these materials can be said to be osteoconductive, that is they support the formation of bone on their surfaces by committed osteoblasts (bone-forming cells). This tolerance of osteoconductive behaviour to a relatively wide range of highly specified materials and structures, coupled with the variability in the level of bioactivity attained, reflects the fact that bone apposition within a porous bioceramic implant is mediated by the combination of physiochemical factors, from the bulk level of porosity and the macropore pore geometry and connectivity down to the micro-topography of the bioceramic struts and their surface chemistry; what is an optimum pore structure for one material is not necessarily so for another. Therefore, as correctly identified in the field of tissue engineering, successful bone regeneration is dependent on having the right combination of structure, chemistry and biology. Traditionally, the field of BGS has concentrated on optimization and study of the first two elements, with the view that if these are optimized then once in situ the native tissue will provide the right biology.

8.4

Bone graft substitute pore structures: balancing space, permeability and mechanics

The structure of ceramic implants has been considered since the use of porous material was first described in 1963.33 Hulbert et al.34 demonstrated that porous disks of a near inert ceramic exhibited thinner fibrous encapsulation with faster healing in surrounding muscle and connective tissue than dense disks, as a result of a mechanical interlock which reduced motion between host tissue and implant. Subsequently, many studies have demonstrated a greater degree and faster rate of bone ingrowth or apposition with percentage porosity; however, there still seems to be some dispute regarding the optimum ‘type’ of porosity. The rate and quality of bone integration has, in turn, been related to a dependence on pore size, porosity volume fraction, interconnection size and interconnection density, both as a function of structural permeability and mechanics (Fig. 8.3). More recently, the role of the strut microstructure and pore geometry has been considered with respect to the influence of these parameters on entrapment and recruitment of growth factors and matrix proteins35 and the phenomenon of osteoinductivity, which will be discussed in more detail in sections 8.5 and 8.6.36–41

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8.3 The effect of increasing porosity (at either the macro- or microstructural level) on cellular environment within a bone graft scaffold.

8.4.1 Scaffold permeability: recreating osteons and canaliculae It is generally accepted that a greater volume and faster rate of bone ingrowth may be obtained by increasing BGS macroporosity (i.e. pores >50 μm in size);24,28,29,42,43 however, there is some confusion as to whether this is a reflection of a dependence between volume or rate of integration and pore size23,30,42,44–47 or other structural parameters, such as pore morphology, porosity volume and pore connectivity.22,48–50 A pore size of 100 μm is often cited as a minimum requirement for healthy ingrowth following the work of Klawitter et al., who actually observed mineralized

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bone ingrowth in pores as small as 40 μm, but reported a greater penetration of bone ingrowth into polyethylene implants containing increasing pore interconnection size (often misquoted as pore size) up to a 100–135 μm limit.42 Earlier work conducted by members of the same group had previously shown that interconnection sizes were critical in determining the quality of tissue ingrowth within porous calcium aluminate ceramics where interconnection sizes of >100 μm, >40 μm and >5 μm were required for ingrowth of mineralized tissue, osteoid and fibrous tissue respectively.43 However, the authors also reported that they believed that leaching of Al from the ceramic may have inhibited mineralization within their study. More recently Lu et al. demonstrated that when using either HA or βTCP the critical pore interconnection size for bone ingrowth was only 50 μm,49 corroborating the earlier work of Holmes23 who found that, when implanted in cortical bone, coral structures with interconnections of osteonic diameter were required for sustainable bone ingrowth.23 This would suggest that pore size is not the controlling factor, but that in fact it is the pore interconnection size, which is often related to both pore size and the extent of porosity,51,52 which is key to rapid and sustained bone ingrowth. This was elegantly demonstrated by improved integration in structures with well interconnected 50–100 μm pores compared with less connected but larger pores of 200–400 μm with similar levels of porosity.48 Moreover, this dependence is unsurprising when you consider that bone is a mineralized tissue that relies heavily on the presence of an internal blood supply for supply of nutrients and oxygen which do not readily diffuse through it. Any new bone formation or repair must always be preceded by the formation of a vascular network, the rapidity and extent of which is strongly influenced by the degree of structural interconnectivity between pores.53 Therefore, is it surprising that a minimum interconnection size exists which is in line with that of osteonal diameter? There is a general consensus that a porosity threshold exists around 60%, below which sustainable bony integration into central pore chambers cannot be expected within non-resorbable scaffolds.31,52 This can be explained by considering the geometrical constraints of porous structures comprising predominantly monomodal spherical porosity above and below 60% with respect to the size and density of pore interconnections.54 Thus, both the degree of scaffold porosity and the interconnection size are likely be responsible for altering the perceived bioactivity of a BGS as a function of increasing the structural permeability. Interestingly, pore connectivity and porosity volume cease to be such a critical factor for resorbable bioceramics such as βTCP and bioglasses, as the resorption exhibited by these materials acts to open up the structure. Thus the optimal connectivity and porosity of resorbable scaffolds may be lower than those established for non-resorbable materials, which may explain the relative insensitivity and conflicting data of scaffolds containing these materials to pore interconnection size and porosity reported in the literature.47,49

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Although it is recognized that both the rate of integration and the volume of regenerated bone may be dependent on features of the macroporosity, recent in vitro and in vivo studies have demonstrated biological sensitivity to the level of microporosity within the ceramic struts.38–41,55–59 As can be seen in Fig. 8.4, this so-called strut porosity or microporosity is of the order of only 1–20 μm in size. There is some evidence that this enhancement in bioactivity may be through mediation of cell attachment55,56,60,61 and/or selective sequestering and binding of adhesion proteins and growth factors35 as a function of either a larger surface area, thereby increasing the quantity of adsorbed growth factors above a critical level for cell recruitment and activation. Alternatively, it may be that the surface texture is a geometrically more suitable substrate for specific adhesion molecule or growth factor adsorption by its influence on surface roughness and surface energy61 resulting in selective and/or ‘functionally advantaged’ adsorption leading to enhanced cell anchorage, regulation or differentiation. The most striking evidence for the importance of the geometrical configuration of a BGS is in the work of Ripamonti and coworkers, who demonstrated that osteoinductivity in HA was linked to the precise shape of surface concavities in implants.62–64 Using immunolocalization, they demonstrated that this osteoinductivity occurred as a result of a concentration of BMP-3 and BMP-7 within the surface concavities. A similar mechanism was proposed to explain the bioactivity seen in bioactive glass microspheres, in which de novo bone formation is believed to initiate from the centre of the hollowed out beads following dissolution of the Si-rich glass and reprecipitation of a CaP-rich shell loaded with adsorbed proteins.65 In a recent in vivo study, where apatite-based grafts with varied levels of strut microporosity were implanted intramuscularly, significant ectopic bone formation was only observed within those grafts with microporosity levels ≥20%, suggesting that the surface concentration or potency of adsorbed native osteogenic growth factors was sensitive to strut porosity levels in these scaffolds, assuming adsorption of these factors to be behind stimulation of osteogenesis within these synthetic grafts.59 Furthermore, studies on the influence of BGS microporosity on the rate and quality of bone healing in vivo demonstrated that faster apposition in microporous scaffolds with strut microporosity levels of ≥20% was linked to the rate of development of the vascular network.58 Immunohistochemical localization had previously demonstrated an association of VEGF with HA surfaces in addition to a close relationship between the HA surface and newly formed capillaries.66

8.4.2 Scaffold mechanics: the goldilocks principle Unlike many simple physical interactions or laws, which often follow predictable relationships where an increase or decrease in a specific input leads to a correlated, predictable change in a specific output, complex biological systems often have so-called goldilocks points where a specific output (the porridge being edible) is

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only achieved when the input is just right (not too hot nor too cold). One example is bone’s sensitivity to just the right level of mechanical stimulation to avoid resorption (i.e. loading is too low, as in astronauts in space) or necrosis and stress fracture (i.e. loading is too high, as in overtraining athletes). Therefore, when optimizing BGS porosity, the effect of structural morphology on mechanical properties must also be considered. Traditionally, most BGS are placed with either internal or external fixation devices so the over-riding mechanical requirement was that the porous structure survived the surgical procedure intact, a stipulation that will vary with the type of procedure employed (e.g. impaction grafting versus void filling). Thus, the mechanical behaviour of various BGS varied widely and was not considered critical.26,30,31,67 For example, the strength of one type of BSG was found to vary between 2 and 10 MPa as a function of porosity (50–80%) and to a lesser extent with fabric, a strength-porosity relationship that is typical of all ceramic BGS. However, any BGS must permit even load distribution and should not be overly stiff, so as not to produce load concentrations or stress shielding.23 Stress analysis investigations have demonstrated that anisotropic structures promote even loading and reduce stress concentrations in comparison with isotropic ones when placed in systems likely to be preferentially loaded in one direction.68 Thus, appropriate modification of the macrostructure to distribute loads and match the intended host tissue will improve the biocompatibility, such as has been demonstrated with the use of hydrothermally converted corals with highly porous structures similar to anisotropic cancellous bone in the filling of cancellous bone defects.30 The converted corals were reported to possess a similar stiffness to cancellous bone, but to have significantly lower strengths and no plastic behaviour, but after six months in vivo the resultant bone/implant composite exhibited similar behaviour to the host bone. It is well known that bone is functionally adaptive, that is it responds to external mechanical stimuli to either reduce or increase its mass as required8 as a result of the mechano-sensitivity of many cell types, including osteoblasts and osteocytes.69 Moreover mechanical forces have recently been shown to mediate osteoblastic differentiation of osteoprogenitor cells.70 Therefore, it is unsurprising that a number of studies have demonstrated that in structures where the level of pore interconnection is sufficient to support adequate vascularization for full bony integration of internal porosity,48 there is a degree of adaptation of bone ingrowth within the porous BGS with time51,57 sometimes leading to the loss of bone volume,44 suggesting that the variation of local strain in scaffold struts with macroporosity may induce or inhibit bone formation within BGS. Additionally, it has been demonstrated that in the longer term both micro and macro-porosity influence bone adaptation,56–58 where it was proposed that a reduction in strut modulus associated with increasing microporosity levels was sufficient to shift the strut modulus below a threshold value resulting in a swing in the equilibrium local bone cell activity towards a greater degree of stable bone apposition. Presumed to result from the sensitivity of cells associated with remodelling within

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normal bone to micro-fracturing and consequent changes in micro strain within the bone.71–73 Moreover, in an experiment comparing integration within a series of four BGS having total porosities of 70% and 80% with low and high levels of strut porosity (and thus bulk compressive strengths and moduli that varied significantly by up to a factor of 10), different levels of equilibrium bone regeneration were observed, where the highest volumes of bone retained in specimens retrieved after a period of six months were found in scaffolds with the highest strut porosities. However, the compressive strengths and moduli of the explants tested after this period of time were all statistically similar to each other and control bone from the same site retrieved and tested under identical conditions, irrespective of the starting properties of the original scaffold.51,58 These results suggest that the equilibrium level of bone ingrowth attainable by a BGS may thus be highly sensitive to a scaffold’s capacity to stress shield integrated bone, again pointing to an optimal position where the scaffold mechanics mimic that of natural bone tissue. Similar findings have been reported by researchers investigating biomechanical modulation of metaphyseal fracture healing, where they have demonstrated strain dependence in a controlled metaphyseal fracture model. In areas with interfragmentary strains below 5%, significantly less bone formation occurred compared with areas with higher strains (6–20%). For strains larger than 20%, fibrocartilage layers were observed. Moreover, low interfragmentary strain (<5%) led to intramembraneous bone formation, whereas higher strains additionally provoked endochondral ossification or fibrocartilage formation.74 Alternatively, introduction of interconnected strut porosity will enable nutrient transport through the struts as well as via the macropore structure, mirroring the canaliculae network found in the microstructure of natural bone tissue (Fig. 8.4).

8.4 Typical biomimetic pore structure of a synthetic bone graft substitute with both (a) interconnected macroporosity and (b) interconnected microporosity within the ceramic struts.

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Thus, the structural dependence of bone ingrowth within a porous scaffold argues for careful selection of BGS material where both rapid assimilation and long-term integrity are favoured by more porous structures, whereas mechanical resilience during surgery and the belief that stability during the initial period of integration is superior for more dense structures. At first, these appear to be conflicting requirements; however, consideration of the native mechanical adaptability of bone suggests that the ideal structure of a scaffold is dictated by its ability to physically permit nutrient, cell and vascular penetration while also having mechanical properties that are matched to the demands of the local environment – reflecting the fact that the properties of native cancellous bone vary significantly throughout the skeleton. This would suggest that precise scaffold requirements may well vary with site of application.

8.5

Bone graft substitute chemistry: creating interactive interfaces

Chemistry is one of the first factors considered when screening new materials for medical or dental application. This screening is often described as assessing ‘biocompatibility’. However, in the first instance it is establishing whether or not a material is toxic. This has resulted in the classification of materials by the type of biological response that their chemistry elicits, as described by Hench et al.,75 where materials with responses ranging from near-inert to resorbable (Table 8.2) can be considered to be biocompatible depending on the requirements of the clinical indication for which they are under evaluation. Surgery always results in an initial inflammatory response and no material is ever totally inert when introduced to the body. At worst, materials can be toxic causing the death of the surrounding cells and tissue usually by the release of soluble products that may, in extreme cases, migrate with tissue fluids causing systemic damage to the patient (Fig. 8.5).76 However, the most common response, displayed by most metals and polymers, is fibrous encapsulation where initially a thin 1–3-μm thick, loosely organized fibrous layer is absorbed to the surface of the implant75 resulting in no direct chemical bond between implant and host tissue (Fig. 8.5). This fibrous layer effectively isolates the implant from the tissue and the ultimate thickness of the fibrous layer is dependent on a combination of the chemical reactivity of the material and the relative motion between tissue and implant. More reactive materials, such as metals that undergo corrosion or polymers containing residual monomers that may be leached under physiological conditions, induce a thicker layer as the body attempts to isolate the source of irritation. Similarly, increased relative motion between an implant and the host tissue, or sub-micron particle release also induces a thicker fibrous inter-layer. Bioactive materials are distinct in their ability to participate in dynamic surface exchange with the physiological environment attaining a degree of dynamic chemical equilibrium with the host tissue at the implant interface

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Table 8.2 Classification of tissue response as proposed by Hench et al.75 Classification

Tissue response

Implant/ tissue bond

Examples

Toxic

Tissue dies

None

Lead oxide Arsenic oxide Most organic monomers

Near inert

Formation of a non-adherent fibrous membrane around implant

None

Alumina Zirconia Carbon

Bioactive

Formation of an interfacial bond with the implant

Chemical

Hydroxyapatite Bioactive glasses and glass- ceramics

Resorbable (may also be bioactive)

Tissue replaces None/ implant as it degrades chemical

Tricalcium phosphate Calcium sulphate Bioactive glasses Calcium phosphate bone cements

(Fig. 8.5). No fibrous layer is formed and the material is said to bond directly with the tissue.77,78 A further extension of bioactive materials are the resorbable materials which undergo biodegradation via a combination of chemical degradation and/or cell-mediated phagocytosis. However, for these materials to retain their bioactivity it is important that the degradation products are non-toxic, may be easily disposed of by the cells and that the rate of degradation is matched to the rate of tissue formation. Since the early 1970s, researchers have been investigating the use of HA (Ca10(PO4)6(OH)2) for use in the treatment of bone fractures or defects as a result of its crystallographic similarity to bone mineral. However, as discussed earlier, stoichiometric HA and bone mineral are fundamentally different at the nanoscale. Stoichiometric HA comprises a highly ordered repeating pattern or lattice of Ca, PO3 and OH ions, whereas bone mineral is substituted with CO3, and a number of trace elements, the exact proportions of which vary according to local physiological conditions (Table 8.1). Moreover, synthetic HA is generally employed as micronsized polycrystalline plasma-sprayed coatings, sintered ceramic monoliths or particles, whereas bone mineral is present as individual nanosized crystals. This morphology gives the natural material greater surface area, whereas, on balance, the presence of the chemical impurities act to distort the crystal lattice, both of these factors contribute to reduce the solubility of bone mineral to facilitate its chemical dissolution at body temperature under the enzymatic and acidic conditions produced by activated osteoclasts. This is critical to bone’s ability to

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8.5 Characteristic events behind varied biological response to toxic, near-inert and bioactive materials.

readily undergo cell-mediated (i.e. controlled) remodelling rather than being resistant to osteoclastic attack (as is dense phase pure stoichiometric HA) or susceptible to cellularly independent chemical degradation under the steady state physiological conditions of the osseous environment (as is phase pure stoichiometric β-TCP). Over the years this disparity in the stability of HA and bone mineral under physiological conditions has led to much disagreement in the classification of the biological response of HA, initially it was considered by some to be near-inert as it did not readily undergo chemical exchange under acellular physiological conditions such as those provided by ‘simulated body fluid’ screening tests. In these tests a protein- and amino acid-free fluid is prepared with an ion content and pH similar to human blood plasma, the test subject is then judged to be ‘bioactive’ if it supports the epitaxial growth of a ‘bone-like apatite’ (BLA) on its surface within a reasonable time frame. This sort of test is critical as a screening method for the assessment of bioglasses, where the presence of a BLA precipitate is

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critical to protein, cell and ultimately tissue attachment. However, in the assessment of calcium phosphates whose native surface chemistry may readily support osteogenic protein and cell attachment, this form of test is less relevant, and false-positive or false-negative results may be obtained where ion exchange pushes super saturation of Ca and PO4 ion concentrations resulting in autoprecipitation of BLA from the fluid. For example, calcium sulfate hemihydrate (CSH) forms apatite in SBF but resorbs rapidly and displays poor in vivo bonebonding ability.79,80 HA is now generally recognized to be bioactive along with other selected calcium phosphates such as β-tricalcium phosphate (β-TCP, Ca3(PO4)2) in addition to specific formulations of quaternary (SiO2-P2O5-CaO-Na2O,81,82), ternary (SiO2-Na2O-CaO, SiO2-P2O5-CaO,75 P2O5-CaO-Na2O,83) and binary (SiO2CaO)84 glass systems, collectively known as bioactive glasses, and a number of glass-ceramics crystallized from SiO2, CaO, P2O5-based melts, such as pseudowollastonite (psW, predominate crystalline phase ~ CaSiO3)85 and apatitewollastonite (AW-GC, predominate crystalline phases ~ Ca10(PO4)6(OH)2-CaSiO3).86 However, the bioactivity of these bioceramics and bioactive glasses is extremely sensitive to phase and glass composition, respectively, for instance the osteoconductivity of HA can be masked by the presence of 5wt% CaO as a second phase. This reflects the dual principle mechanisms through which chemistry induces bioactivity:

•

•

directly through dissolution and release of ionic products in vivo, elevating local concentrations of species which interact directly with local cells via ion channels (Ca2+), influence cell behaviour by their effect on local pH (OH−) or interact directly with key proteins to modulate protein conformation and activity; indirectly through the effect that surface chemistry has on protein adsorption and conformation, hence the species and their activity resident at the material’s surface, which impacts on subsequent cell attachment and function.

For example, the presence of second phases such as CaO in HA can result in both inappropriate ion exchange and ‘poisoning’ of the material’s surface to block the adhesion of osteoconductive proteins, resulting in the phase impure material evoking a near-inert response.87,88 This behaviour may account for some of the early disagreement as to the bioactivity of HA, as a result of underspecification of phase purity in early products where the impact of relatively low levels of second phases was poorly understood. As a result of the recognition that bone repair is highly sensitive to chemistry, there has been considerable activity in the development of substituted apatites, the rationale being to more closely mimic the chemistry of bone mineral.89–93 In addition to calcium, phosphate and hydroxyl ions, bone mineral contains significant concentrations of other ions such as carbonate, sodium, magnesium

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and trace levels of silicon and zinc and aluminium (Table 8.1). The inclusion of some of these ions within the apatite lattice has thus been postulated to improve the bioactivity of the material. Carbonate substitution was historically one of the first substituted apatites to attract interest and was reported to increase the solubility of HA and to facilitate resorption by osteoclasts;94,95 however, there are conflicting reports as to the efficacy with which osteoclasts are able to remodel carbonated apatites reflecting in turn the sensitivity of osteoclastic activity to local pH and the surface charge of carbonated apatites which is related to the extent and type of carbonate substitution.96,97 Magnesium has long been recognized to play an important role in the mineralization of bone98 and in the promotion of bone formation.99 Studies of Mg ion incorporation by ion-beam implantation in a range of biomaterials, including HA,100,101 have established its stimulatory effect on bone formation, although given the role of Mg2+ as a signalling molecule, whether this is through its influence on surface chemistry or by direct interaction of Mg ions with cells is unclear. One of the more widely studied and clinically applied forms of substituted apatites to date has been silicate substitution, which has been demonstrated to have a significant influence on the rate and pattern of bone formation in vivo.102,103 However, it has also been reported that the level of Silicate (SiO4) substitution is critical to determining the best response, where an optimum level of around 0.8wt% Silicon (Si) has been reported,102,104 and the incorporation of Si into the HA structure appears to affects its bioactivity by both direct and indirect mechanisms.105–107 The direct influence of bioavailable Si on bone formation and in bone repair has been of interest for a number of years. In 1970, Carlisle first reported that silicon deficiency resulted in abnormal bone formation,108 although there is some evidence to suggest that this may have resulted from the blocking action that Si has on the uptake of Al.109 The role of silicon has clearly been demonstrated in promoting bone induction in bioactive glasses and glass ceramics with demonstration of upregulation of proliferation110 and gene expression by osteoblasts including BMP-2 when exposed to the ionic products of these bioactive glasses.111,112 Interestingly, elevated concentrations of these products have been reported to result in cell death or apoptosis,113,114 while Reffitt et al. demonstrated that orthosilicic acid present in physiological concentrations (5–20 μM) stimulated collagen type I synthesis and enhanced osteoblast differentiation while treatment at a higher concentration resulted in a smaller increase in collagen synthesis.115 Furthermore, work on the mechanisms behind the bioactivity of psW has demonstrated the additive effect that co-treatment with SiO4 and Ca ions has on stimulation of bone cell activity compared with treatment with either element in isolation,116 and is reflected in the differential release of these ions from silicate substituted apatites.107 There is also evidence that bioavailable Si affects multiple anabolic and catabolic processes in bone. In silicate substitute apatites, the presence of Si has been shown to accelerate capillary penetration, and an optimal

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level of 0.8wt% Si also increases the rate of early bone formation and elevates the equilibrium volume of bone within the scaffold. In contrast, non-optimal levels of silicate substitution (0.2, 0.4 and 1.5wt%) resulted in a significant reduction in equilibrium bone volume despite having elevated mineral apposition rates at later time points suggesting a direct impact on the catabolic response of bone.102 This suggests that the presence of optimal Si levels promotes the retention of dense bone morphology within the osseous environment, despite local fluctuations in biomechanical demand. This hypothesis is supported by a decrease in the humeri strength of broiler chickens with one wing immobilized on a diet supplemented with sodium fluoride compared with no variation in the humeri strength of sodium silicate supplemented birds.117 Moreover, a similar observation with regard to retention of bone density was reported in a study of bone apposition around bioglass and HA/TCP-coated implants118 where the HA/TCP and control groups suffered a significant decline in bone volume and trabecular thickness with time in comparison to the bioglass-coated implants. This disparity in remodelling behaviour suggests that optimal levels of bioavailable silicon may stimulate bone development under both healing and steady state conditions. Therefore, it appears that in the field of bone regeneration, to stimulate an appropriate response, then both a biomimetic structure and a biomimetic chemistry are required. However, empirical and clinical evidence demonstrates that this does not mean that the chemistry is limited to only a specific crystallographic form of Ca-PO4 which is similar to BLA. The success of bioactive glasses and surface-treated titanium implants suggests that complete mimicry is not required, moreover, the chemical differences between these materials appears at first highly contradictory. Study of the dynamic exchange processes occurring at the material interfaces, however, demonstrates similarities in the way that hydroxylation of titinate and silica gel-like layers support epitaxial growth of BLA, which then interacts with the dynamic protein population to present an osteogenic surface to the local cell population.119 The selectivity or sensitivity of osteogenic cells to underlying chemistry was elegantly demonstrated in the early 2000s by work with either HA-polyethylene (PE) or HA-PMMA composites,60,120 where osteoblast proliferation and differentiation increased with HA addition and were directly related to increased focal contact (with HA acting as ‘stepping stones’ for cell adhesion plaques) and rates of cytoskeletal organization. This work translated in vivo to a critical HA loading level (of 40wt%) that related to the surface availability of HA particles for osteointegration. In order to understand these phenomena, a number of studies quickly demonstrated the importance of surface charge and wettability in controlling cell attachment and activity,96,121–123 and this has been linked to their influence on protein attachment124,125 in terms of the selectivity126 quantity127 and the ‘quality’65 of protein attachment. Finally, there is increasing evidence to suggest that either form of chemical interaction may not only act directly on the bone or stem cells but also may act

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further downstream, influencing events such as angiogenesis prior to ossification or the production of specific cytokines or signal molecules by ‘controller’ T cells that then direct the differentiation of stem cells and the function of bone cells. For example, in a study of angiogenesis following HA implantation a close relationship was demonstrated between the HA surface and newly formed capillaries where immunohistochemical localization demonstrated association of VEGF with the HA surface.66 Similarly, vascularization was found to progress more rapidly within a macroporous HA orbital implant compared with a macroporous PE implant,53 whereas investigation of T lymphocyte activity (controller cells that express inflammatory cytokines) on exposure to HA particles combined with histological evaluation of distant organs suggests that implantation of HA may have a short-term systemic effect.128,129

8.6

Future regeneration therapies: biologics and tissue engineering

The fact that chemistry can induce bioactivity, through the influence that surface charge and ionic speciation has on the population and conformation (and thus potency) of proteins (such as matrix proteins or growth factors) absorbed on its surface and their subsequent influence on cell attachment behaviour, has led to considerable interest in co-delivery130 or direct surface modification of implants or scaffolds with specific peptide sequences or whole protein segments131–133 to facilitate the rapid incorporation and appropriate functionality of devices. Alternatively (and sometimes concurrently), pre-colonization of scaffolds or surfaces with native cells (both committed osteoblasts and stem cells) has been employed and demonstrated to accelerate repair.134,135 These solutions are particularly attractive in situations where we are dealing with patients with compromised biology or where we are trying to induce bone formation outside of normal physiological conditions (such as following chemotherapy or in spinal fusion, respectively).136 This represents a blurring of the lines between provision of traditional medical devices, tissue donation and biologic or pharmaceutical therapies. From a biomimetic perspective then, moving from the crude use of natural templates to the more intelligent harnessing of a tissue’s natural capacity for regeneration via the use of native cells and stimulants seems an obvious advancement for medical science. Regulatory headaches aside, this is the most likely direction for future bone grafting therapies; however, recent clinical experience with bone morphogenic proteins (BMPs) suggests that considerable care must be exercised when manipulating the body’s repair processes at this level. BMPs were first identified in 1965 by Marshal Urist137 as being ‘osteoinductive’, that is having the capacity to induce de novo bone formation in an ectopic site (such as a muscle pouch) through both morphogenic and mitogenic action on local stem cells.138 The development of biologic therapies using use supra-physiological doses of

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powerful growth factors such as rhBMP-2 or rhBMP-7 delivered with the use of putty or collagen sponge carriers initially generated significant interest and investment as a viable alternative to autografting.139 However, complications associated with osteolysis, graft subsidence, graft migration, cyst formation and neuritis and spontaneous ectopic bone formation, particularly when used off label in the spine where reported side effects include paralysis and death, has resulted in restricted regulatory approval and some caution in the use of these products.140,141 There is considerable dispute as to whether or not some ceramic bone grafts can be considered to be osteoinductive. De novo bone formation has been reported with the implantation of some graft materials ectopically; however, it appears to be strongly associated with the presence of strut porosity40,59 or specific structural features,64 and may reflect the fact that the material itself is not truly inductive in the sense of a morphogenic growth factor. It is more likely that a combination of a biomimetic chemistry and biomimetic pore structure works to facilitate sequestering and enrichment of native growth factors, which then act to induce bone formation locally. The native ‘remodelling’ process of bone which enables it to maintain an optimal balance between form and function throughout life via a continual process of renewal8 and to repair itself following minor trauma is a highly complex process modulated by a combination of mechanical, biochemical and hormonal factors, an imbalance in any of which leads to metabolic bone diseases such as osteoporosis. As a consequence of the mapping of the human genome and the relative ease and accuracy with which gene analysis can now be performed using micro-arrays and RT-PCR of both cell and tissue samples, significant advances have been made into the understanding of these processes. Mechanistic studies into the pathways behind cell and tissue response to specific implant surfaces,142 surfaces grafted with proteins133 and mechano-biologic control of osteogenesis74 are informing new bone regeneration therapies in a manner once considered more appropriate to pharmacological studies of bone metabolism143 and demonstrating genetic basis behind nanoscale sensitivity at the chemical, structural and mechanical levels.

8.7

Conclusion

Bone is a living, breathing adaptive tissue, able to respond to changes in both metabolic and biomechanical demands on a daily basis. The development of an increasingly biomimetic approach to the development of synthetic BGS coupled with greater understanding of the mechano-biological pathways behind regulation and regeneration of bone structure may well result in the next step change in orthopaedic medicine, perhaps even beyond the choice between regeneration or replacement surgery to minimally invasive therapeutic alternatives superseding surgical intervention.

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8.8

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