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`first century: biology, chemistry or
`
`
`
`−
`Bone repair in the twenty
`engineering?
`
`Karin A. Hing
`
` 2004 Phil. Trans. R. Soc. Lond. A
`2004
`
`
`
`362
`
`, doi: 10.1098/rsta.2004.1466
`
`, published 15 December
`
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`MILLENIUM EXHIBIT 2031
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`10.1098/rsta.2004.1466
`
`Bone repair in the twenty-first century:
`biology, chemistry or engineering?
`By Karin A. Hing
`Interdisciplinary Research Centre in Biomedical Materials,
`Queen Mary, University of London, London E1 4NS, UK
`(k.a.hing@qmul.ac.uk)
`
`Published online 24 September 2004
`
`Increases in reconstructive orthopaedic surgery, such as total hip replacement and
`spinal fusion, resulting from advances in surgical practice and the ageing population,
`have lead to a demand for bone graft that far exceeds supply. Consequently, a number
`of synthetic bone-graft substitutes (BGSs) have been developed with mixed success
`and surgical acceptance. Skeletal tissue regeneration requires the interaction of three
`basic elements: cells, growth factors (GFs) and a permissive scaffold. This can be
`achieved by pre-loading a synthetic scaffold with GFs or pre-expanded cells; however,
`a ‘simpler’ approach is to design intrinsic ‘osteoinductivity’ into your BGS, i.e. the
`capability to recruit and stimulate the patient’s own GFs and stem cells. Through
`investigation of the mechanisms controlling bone repair in BGSs, linking interactions
`between the local chemical and physical environment, scientists are currently devel-
`oping osteoinductive materials that can stimulate bone regeneration through control
`of the scaffold chemistry and structure. Moreover, this body of research is providing
`the foundations for future generations of BGSs and bone-repair therapies and may
`ultimately contribute towards improving the quality of life through maintenance of
`the skeleton and reversal of disease states, as opposed to the mending of broken
`bones that we currently practice. Will we be able to grow our own bones in a bio-
`reactor for use as autologous graft materials in the future? Could surgery be limited
`to accidental trauma cases, with greater restoration of function through biochemical
`or gene therapies? The technology and research probes necessary to this task are cur-
`rently being developed with the advent of nanotechnology, genomics and proteomics:
`are we about to embark on a chemical revolution in medicine? This paper aims to
`discuss some of the current thinking on the mechanisms behind bioactivity and bio-
`compatibility in bone and how a fuller understanding of the interactions between
`cells and the materials used today could bring about completely new approaches for
`the treatment of bone fracture and disease tomorrow.
`Keywords: tissue engineering; cell–material interactions;
`bone formation; bone-graft substitute
`
`1. Introduction
`From a biological perspective, bone is a remarkable living tissue that performs sev-
`eral key functions within the body. Bone not only provides structural support and
`
`One contribution of 17 to a Triennial Issue ‘Chemistry and life science’.
`
`Phil. Trans. R. Soc. Lond. A (2004) 362, 2821–2850
`2821
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`K. A. Hing
`
`protection to bodily organs, but it is involved in the metabolism of minerals such as
`calcium, and is the primary site for the synthesis of blood cells. Furthermore, it is
`capable of maintaining an optimal shape and structure throughout life via a continual
`process of renewal through which it is able to respond to changes in its mechanical
`environment by ‘remodelling’ to meet different loading demands, so maintaining an
`optimal balance between form and function (Wolff 1870). However, as a living tissue,
`bone requires a constant supply of oxygen and nutrients, and is limited in the size
`of fracture or defect it is able to restore to healthy tissue. Furthermore, bone can
`suffer from pathological conditions, (e.g. cancer) and is subject to degeneration as
`a result of age and disease (i.e. osteoporosis). In these cases, patient comfort and
`bone function can often only be restored by surgical reconstruction. 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 first established in the
`1800s (Czitrom & Gross 1992; Meeder & Eggers 1994; Sanan & Haines 1997). Evi-
`dence 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 (Sanan & Haines 1997). Furthermore, archaeolog-
`ical 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 (Wangensteen & Wangensteen 1978). In modern
`medicine, autografting is regarded as the ‘gold standard’; however, the amount of
`bone that can be safely harvested is limited, while 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 disease
`(Barriga et al. 2004; McCann et al. 2004) and healing can be inconsistent (Togawa et
`al. 2004). Consequently, there is an increasing demand for synthetic bone-graft prod-
`ucts that would avoid these complications, in addition to overcoming the problem of
`an inadequate supply of material.
`In 2001 worldwide sales for orthopaedic products approached $15 billion and con-
`tinue to expand at an annual growth rate of 13% (Clinica Reports 2002). Further-
`more, the bone-grafting segment, valued at over $1 billion globally, has been esti-
`mated to represent 408 000 procedures in Europe and 605 000 procedures in the USA
`alone (Clinica Reports 2002). In the period 2002–2003, over 77 000 primary hip oper-
`ations were performed within a total of 617 000 bone and joint procedures recorded
`by the NHS in England alone (Government Statistical Service 2003). Moreover, an
`estimated further 19 000 (20%) hip operations are performed annually within the
`private sector; thus the true number of bone-repair procedures performed annually
`within England is likely to be well in excess of this figure. This ‘boom’ in reconstruc-
`tive surgery is due in part to favourable demographics. Currently, there is a projected
`annual increase of 2–3% in the global 65+ population (i.e. a population increase of
`100 000 000 people aged 65 and over from 2000 to 2010). However, it also results
`from lifestyle changes, increased expectations regarding the quality of life, develop-
`ments in surgical technique (such as minimally invasive surgery) and advancements
`in technology (leading to new innovations in implant materials) resulting in a greater
`demand for and a wider application of orthopaedic devices. Ironically, the desire to
`lead an active healthy lifestyle results in an increase in sports-related injuries and
`joint damage, while those that lead an overly sedentary lifestyle can suffer from poor
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`Bone repair in the twenty-first century
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`posture leading to spinal problems. Moreover, joint replacements performed in the
`1980s are now reaching the end of their lifetimes and will require revision. Fortu-
`nately, advancements in biomaterials technology have led to the development of new
`more durable or more ‘biocompatible’ materials that make devices such as hip pros-
`theses more suitable for younger patients who wish to continue to lead an active
`lifestyle. For instance, the use of tough alumina (Al2O3) or yttria-stabilized zirco-
`nia ceramic heads coupled with durable high molecular weight polyethylene cups in
`total hip replacement has been shown to significantly reduce the generation of wear
`particle debris as compared with metal/polymer combinations (Hernigou & Bahrami
`2003; Santavirta et al. 2003; Tanaka et al. 2003). These submicrometre or nanosized
`wear particles, when produced in a high enough concentrations, have in turn been
`found to upregulate or ‘switch on’ the body’s natural immune response, which sets in
`motion a cascade of events ultimately leading to loosening and failure of the implant
`(Orishimo et al. 2003; Warashina et al. 2003). Moreover, the inclusion of a ‘bioactive’
`coating on the stem of the femoral component can also increase the rate of healing
`and extend the lifetime of the implant by promoting direct bonding between the
`bone and the implant (Kroon & Freeman 1992; Reikeras & Gunderson 2003; Skinner
`et al. 2003).
`However, as a result of bones’ load-bearing function in the body, much of the
`orthopaedic surgery performed today uses engineering solutions; for instance, in a
`simple break the fracture site will be externally or internally fixated using splints and
`casts or metal plates and pins, respectively, which provides the local stability neces-
`sary to facilitate bone regeneration. Similarly, spinal injuries are treated with the use
`of metal cages and hips and knees are reconstructed using components selected for
`their mechanical properties rather than their biological functionality. Where internal
`fixation is required to facilitate healing the patient often has to undergo a second
`operation to remove the metalwork, particularly in the young and active where it is
`vital to retrieve devices once they become surplus to requirements lest they interfere
`with normal patterns of bone growth, lead to wasting of bone tissue due to stress
`shielding (a side effect of bone’s ability to respond to functional demand, which
`is why astronauts loose bone mass in space and why highly trained athletes have
`denser bones), or have the potential to severely complicate a second fracture should
`the patient be unlucky enough to injure themselves again. In order to spare patients
`the ordeal of retrieval (and to save healthcare resources) there is a considerable body
`of research focused on the development of bioresorbable fixation devices (Steinmann
`et al. 1990) that use bones’ natural remodelling characteristics to degrade the device
`and dispose of the degradation products, similar to the resorbable sutures commonly
`used in wound closure (Chu et al. 1996). Examples of the materials used in these
`devices include resorbable polymers such as polyurethane, and poly(L-lactic acid).
`However the application of bioresorbable plates, screws and cements is generally
`limited to the treatment of small defects, such as in maxillofacial surgery, and to
`the fixation of tendon grafts into bone (McGuire et al. 1999), as a result of their
`relatively low strength and concerns regarding the biocompatibility of their degra-
`dation products (B¨ostman 1998; Ignatius & Claes 1996). However, they do have the
`potential to act as drug-delivery devices, releasing growth factors (proteins that can
`stimulate bone repair) to promote rapid bone healing as they degrade, the main
`technical challenge being the control of the rate at which they degrade and thus also
`control of the targeted area of delivery and the drug dosage (Di Silvio et al. 1994).
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`K. A. Hing
`
`Table 1. A selection of bone-graft substitutes available for clinical use
`
`name
`
`ApaPore
`
`Bonesave
`
`chemistry
`
`morphology
`
`features
`
`HA
`
`macroporous
`
`HA/TCP
`(60/40)
`
`macroporous
`
`three porosity grades
`60, 70, 80%;
`interconnected microporosity
`50% porosity
`
`Endobon
`
`HA
`
`macroporous
`
`bovine cancellous
`structure
`
`Grafton
`
`demineralized
`bone matrix
`
`putty, gel, granules,
`sheet, matrix
`
`natural cocktail
`of proteins/GFs
`
`Hedrocel
`trabecular metal
`
`tantalum
`
`macroporous
`
`InFuse
`
`Norian
`
`OsSatura
`
`BMP 2
`
`CO3 apatite
`
`HA/TCP
`(80/20)
`
`GF loaded onto
`collagen sponge
`
`dense dahllite
`cement on setting
`
`macroporous
`
`OssiGraft (OP-1) BMP 7
`
`powder
`
`Osteoset
`
`Pro Osteon
`
`CaSO4
`
`HA
`
`dense pellets
`
`macroporous
`
`high
`toughness
`
`contains GF
`
`injectable cement
`
`scaffold with
`biomimetic coating
`
`contains GF
`
`resorbable
`
`coral exoskeletons;
`two 200 and 500 mm
`pore-size grades
`
`macroporous
`
`resorbable
`
`Pro Osteon-R
`
`Skelite
`
`Vitoss
`
`CaCO3
`HA coated
`
`Ca–PO4
`
`macroporous
`
`TCP
`
`macroporous
`
`resorbable;
`interconnected microporosity
`
`resorbable;
`interconnected microporosity
`
`In orthopaedics the closest we come to harnessing bone’s natural regenerative
`powers is in bone grafting. In severe trauma cases, some hip revisions and in the
`correction of large ‘bony defects’, where a significant part of the bone is missing or
`damaged, bone grafts are used to replace or augment the missing or fractured bone
`(generally in combination with fixation devices as mechanical stabilization demands
`must still be met). As discussed earlier, this can be done using either the patient’s own
`bone (autograft), removed from another site, or bone allograft (human bone obtained
`from a bone bank). The graft should not only replace missing bone but encourage
`osseointegration, i.e. act as a scaffold for guided bone growth into the graft, so helping
`the body to repair its own lost bone. This bone in-growth strengthens the grafted
`area by forming a bridge between the existing bone and the graft material. Ideally
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`Bone repair in the twenty-first century
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`(a)
`
`(b)
`
`1 mm
`
`10 µm
`
`Figure 1. Scanning electron micrographs of some typical ceramic BGSs. Note
`the variation in (a) macrostructures and (b) microstructures and surface topographies.
`
`with remodelling over time, the newly formed bone should replace much of the graft.
`However, with the rising usage of bone grafts there are now not enough to meet
`demand, surgeons have reported that in autografting there is generally insufficient
`bone 20% of the time, while a single hip-revision procedure can require 4–6 femoral
`heads’ worth of banked allograft.
`In response to this a number of synthetic bone-graft substitutes (BGSs) have been
`developed and are in use clinically, with mixed success and surgical acceptance (see
`figure 1 and table 1). As can be seen from table 1 the term ‘bone-graft substitute’
`includes materials with a wide range of chemistries and structural morphologies;
`all these materials can be said to be biocompatible and most are osteoconductive
`(i.e. they support the formation of bone on their surfaces by mature bone-forming
`cells, osteoblasts). Some of these materials are intended for use with cells or growth
`factors to produce a biologically active graft—a practice known as tissue engineer-
`ing. However, some also claim to be ‘osteoinductive’ (i.e. they are able to induce
`bone formation by influencing the differentiation or maturation of stem cells into
`bone-forming cells) by interaction of the material’s surface with local cells and pro-
`teins, through either their chemistry or their micro-topography. The availability of
`a synthetic BGS that would reliably replicate the best results observed with the use
`of fresh healthy autograft, which contains the patient’s own bone cells and growth
`factors, would be of great benefit to both surgeon and patient alike.
`This paper aims to discuss and elaborate on some of the current thinking on
`the mechanisms behind bioactivity and biocompatibility in bone and how a fuller
`understanding of the interactions between cells and the materials used today could
`bring about completely new approaches for the treatment of bone fracture and disease
`tomorrow. However, we will first start by considering bone, its make-up and its
`normal healing and remodelling cycles.
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`2826
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`K. A. Hing
`
`cancellous bone
`
`cortical bone
`
`
`Figure 2. Sectioned human femoral head demonstrating the variation in bone structure.
`
`(a) Basics of bone
`Bone is astonishing. It is a living, highly vascular, dynamic, mineralized, connective
`tissue, which is characterized by its hardness, resilience and growth mechanisms, and
`its ability to remodel and repair itself. Simply, bone is a dense multi-phase material or
`‘composite’ made up of cells embedded in a matrix composed of both organic (colla-
`gen fibres, lipids, peptides, proteins, glycoproteins, polysaccharides and citrates) and
`inorganic (calcium-phosphates, carbonates, sodium, magnesium and fluoride salts)
`elements (Cameron 1972). However, its structure and 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. Moreover,
`as previously mentioned, bone is not purely a structural tissue, but is also respon-
`sible for maintaining mineral homeostasis and providing a source of haematopoietic
`stem cells, i.e. it acts as a mineral (notably calcium) and blood cell reservoir for
`the rest of the body. Consequently, bone in its natural environment is engaged in
`a constant cycle of resorption and renewal, undergoing continual chemical exchange
`and structural remodelling, due to both internal hormonal regulation and external
`mechanical demands. Moreover, the mineral function can supersede the structural
`function, resulting in loss of integrity in the bone structure. Indeed, many patholog-
`ical bone diseases are instances where an imbalance in the body’s normal hormonal
`regulatory system results in depletion of bone (osteoporosis) or overproduction of
`bone (Paget’s disease).
`
`(i) Bone structure
`Mature bone is composed of two types of tissue, one of which is relatively dense,
`known as cortical bone, while the other consists of a network of struts or trabeculae
`surrounding interconnected spaces or cancelli and is known as trabecular or cancel-
`lous 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,
`while cancellous bone is found in the interior of bones, such as within the femoral
`head, and vertebra (figure 2).
`There are two kinds of cancellous bone: coarse and fine. Coarse cancellous bone
`is characteristic of healthy adult mammalian skeleton, while fine cancellous bone is
`characteristic of the foetal skeleton or early fracture callus and comes in two forms,
`fine cancellous membranous bone and fine cancellous endochondral bone. There are
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`Bone repair in the twenty-first century
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`2827
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`Table 2. Data for the mechanical properties of bone (Bonfield 1989; Cowin et al. 1986; Goldstein
`et al. 1983; Kuhn et al. 1989; Martens et al. 1983), HA (Akao et al. 1981; Best 1990; de With
`et al. 1981) and collagen (Bennet et al. 1986)
`
`testing
`direction
`
`compressive
`strength (MPa)
`
`Young’s
`tensile
`strength (MPa) modulus (GPa)
`
`cortical bone
`
`cancellous bone
`
`HA
`collagen
`
`longitudinal
`transverse
`longitudinal
`transverse
`n/a
`longitudinal
`
`193
`133
`3.6–9.3
`0.6–4.9
`—
`—
`
`133–150
`50
`—
`—
`9–120
`100
`
`17–25
`12
`0.26–0.90
`0.01–0.40
`80–117
`1.5
`
`ropocollagen molecule
`
`characteristic banding
`
`collagen fibre
`
`64 nm
`
`peptide chain
`
`collagen fibril
`
`Figure 3. Hierarchical organization of collagen fibres.
`
`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.
`This variation in structure leads to considerable variation in its 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 site (Goldstein 1987) and
`age (Burstein et al. 1976; Wong et al. 1985). In addition, cancellous bone is often
`anisotropic due to the orientation of major trabeculae along lines of principle stress
`(figure 2). Table 2 summarizes some of the data available from the literature.
`On 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 and acts as the body’s mineral store, the various
`bone cells are responsible for maintaining the structure of the matrix, regulating its
`oxygen and nutrient supply, and storing or releasing minerals as required, while the
`marrow and vasculature provides the source of stem cells and the main means of
`communication and interaction with the rest of the body.
`
`(ii) The matrix
`The extra-cellular matrix has two main components: the organic collagen fibres and
`the inorganic bone mineral crystals. Together they make up ca. 95% of the dry weight
`of bone, the remainder being composed of other organic molecules (known collectively
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`2828
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`K. A. Hing
`
`Table 3. Variation in the composition of bone mineral
`(References: 1, McConnel (1973); 2, Driessens (1980); 3, Aoki (1991); 4, Le Geros & Le Geros
`(1993).)
`
`reference
`
`Ca
`
`P
`
`Mg
`
`Na
`
`K
`
`CO3
`
`F
`
`Cl
`
`Sr, Zn, Cu
`
`1
`2
`3
`4
`
`26.7
`36.7
`34.0
`24.5
`
`12.5
`16.0
`15.0
`11.5
`
`0.44
`0.46
`0.50
`0.55
`
`3.48
`0.06
`0.73
`0.77 — 8.00
`0.80
`0.20
`1.60
`0.70
`0.03
`5.80
`
`0.08
`0.07
`0.04 —
`0.08
`0.2
`0.02
`0.10
`
`Sr = 0.04
`—
`—
`traces
`
`Ca:P
`ratio
`
`1.66
`1.77
`1.75
`1.65
`
`as the non-collagenous proteins) 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 a wide range of mechanical
`properties and to retain elasticity, toughness and hardness for a minimal weight.
`
`Collagen. Collagen is the most abundant protein found in the body, and occurs
`in a number of different connective tissues both calcified and non-calcified. Collagen
`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. Although all α-chains contain the glycine–X–Y sequence, differ-
`ent types of collagen may be produced via the combination of different amounts and
`sequences of other amino acids within the tropocollogen molecule. To date, 13 dif-
`ferent 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 3
`4 stagger, parallel array (figure 3). As a result of this assembly, the
`fibrils exhibit characteristic cross-striations or banding, which occurs in a repeating
`pattern every 55–75 nm, average 64 nm (Robinson & Watson 1952). 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 the fibrils are wound into bundles
`to form collagen fibres that range in diameter from 0.2 to 12 µm (Kielty et al. 1993).
`
`Bone mineral. 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 (de Jong 1926),
`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 defi-
`ciency (reported Ca:P ratios of 1.37–1.87) (McConnel 1973; Posner 1969), internal
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`Bone repair in the twenty-first century
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`crystal disorder and ionic substitution within the apatite lattice resulting in the pres-
`ence of significant levels of additional trace elements within bone mineral (table 3).
`It is not a direct analogue of HA as is commonly believed, but more closely related to
`an A–B type carbonate-substituted apatite (Elliot 1994; Le Geros & Le Geros 1993).
`These factors all contribute to an apatite that is insoluble enough for stability, yet
`sufficiently reactive to allow the in vivo sub-microscopic (5–100 nm) crystallites to
`be constantly resorbed and reformed as required by the body.
`While many investigators of calcification agree that mineralization originates
`within matrix vesicles (Anderson 1980), there is some disagreement about the exact
`mechanisms of the process. It seems likely that both cellular and physiochemical
`factors are involved. Generally amorphous calcium phosphates with a Ca:P ratio
`varying between 1.44 and 1.55 are believed to be deposited under the control of
`osteoblasts. Once deposited, this amorphous tricalcium phosphate is present as a
`reservoir for HA crystallite nucleation and growth independently of the cells (Posner
`1969). In vitro studies have been used to try to provide an insight into the forma-
`tion of bio-apatite in vivo, where variation in the Ca and PO4 saturation above and
`below physiological levels influenced nucleation and crystallite shape (Blumenthal &
`Posner 1973; Boskey & Posner 1976). Other organic substances present in the matrix
`have long been thought to be responsible for promoting the initial nucleation and
`deposition of bio-apatite, and regulating the orientation, size and growth rate of the
`crystals. More recent studies have focused investigation on the role of some of the
`non-collagenous proteins found in bone with interesting results.
`
`Non-collagenous proteins. The non-collagenous organic constituents of bone matrix
`include a number of sulphated and acid mucopolysaccharides. There are four proteins
`of particular interest, osteocalcin (OC), bone sialoprotein (BSP) osteopontin (OP)
`and osteonectin (ON). These are members of the group generally referred to as the
`non-collagenous proteins (NCPs), are produced by bone cells and are believed to
`regulate bone mineralization and remodelling. Only OC and BSP are bone specific,
`all appear to be multi-functional and their relative composition within the bone
`matrix appears to be self-regulating through a feedback effect on the expression of
`NCPs by osteoblasts (Butler 2000; Gerstenfeld et al. 2000).
`Osteocalcin, also known as bone Gla protein, is one of the most abundant NCPs
`in bone, comprising up to 20% of the total NCPs in bone. OC has a strong affinity
`for HA but not amorphous Ca–PO4, binding to HA through orientation of the Gla
`residues with the Ca ions in the mineral lattice (Butler 2000). Furthermore, OC
`has been shown to be a potent inhibitor of HA formation by delaying nucleation
`(Hunter et al. 1996); its expression is mediated by Ca2+ regulated hormones, and
`recent experiments have demonstrated that OC acts to regulate remodelling through
`suppression of bone formation by osteoblasts (Butler 2000). Several metabolic bone
`diseases such as Paget’s and osteomalacia are associated with elevated levels of OC.
`Bone sialoprotein (BSP) is an acidic sialoglycoprotein and comprises 15% of the
`total non-collagenous proteins in bone. BSP is thought to be involved in both bone
`mineralization and remodelling (Wuttke et al. 2001). It is a multi-functional protein
`promoting both osteoblast differentiation and bone resorption in a dose-dependent
`manner (Wuttke et al. 2001). BSP has an arginine–glycine–aspartate (RGD) tripep-
`tide sequence, the minimal structure required for cell binding and has been shown to
`promote osteoblasts adhesion in vitro (Gerstenfeld et al. 2000). Furthermore, BSP
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`has a high affinity for HA and has been shown to facilitate HA nucleation (Hunter
`et al. 1996).
`Osteopontin is an acidic phosphorylated glycoprotein present in many tissues. OP
`is a highly potent inhibitor of HA formation, acting by inhibition of crystallite growth
`(Hunter et al. 1996). Furthermore, protein confirmation has been shown to affect its
`ability to inhibit nucleation. It appears that the high level of phosphorylization and
`associated negative charge density, which may be regulated, is key to its inhibition
`potency. This supports the theory that crystallite growth is suppressed by OP bind-
`ing to the HA and sterically preventing further ionic growth through electrostatic
`repulsion of ions (Pampena et al. 2004). OP also has an RGD sequence that promotes
`cell attachment and is involved in the regulation of osteoclast motility during bone
`resorption (Butler 2000) and it has been implicated in the regulation of both osseous
`and ectopic calcification (Pampena et al. 2004).
`Osteonectin is expressed by osteoblastic cells. However, its abundance in the bone
`matrix is highly variable and it is also present in many other tissues. As a result of
`this variability, its role in bone formation is unclear, ON has been reported to inhibit
`HA crystallite growth but only when it is present in sufficiently high concentrations
`(Hunter et al. 1996). Furthermore, studies failed to detect ON in newly formed
`bone, questioning previously proposed roles for this protein in tissue mineralization
`(Kasugai et al. 1991; Nagata et al. 1991). ON is a single-chain polypeptide containing
`two glutamate-rich segments that can bind eight Ca2+ ions. Unsurprisingly, ON binds
`strongly to HA (Romberg et al. 1986). However, it also binds to other extracellular
`matrix proteins including collagens I and V. ON will also bind to, and inhibit the
`spreading of, endothelial and smooth muscle cells (Sage et al. 1989). ON is therefore
`believed to be involved in cell–matrix interactions rather than being directly involved
`in mineralization, facilitating changes in cellular shape and cell disengagement from
`the matrix.
`
`Growth factors. Growth factors (GFs) are peptides that regulate cell growth, func-
`tion and motility, resulting in the formation of new tissue. Bone GFs influence the
`synthesis of new bone by acting on the local cell population present in bone mar-
`row and on bone surfaces. They either act directly on specific osteoblasts as local
`regulators of cell growth and function or by inducing angiogenesis (vascularization)
`(basic fibroblast growth factors 1 & 2, bFGF-1/2, vascular endothelial growth factor,
`VEGF), or osteogenesis by promoting endothelial or osteoprogenitor cell migration
`and differentiation (Urist 1965). Bone matrix contains a great number of growth fac-
`tors (Solheim 1998; Yoon & Boden 2002) including fibroblast growth factors (FGFs),
`insulin-like growth factor I and II (IGF-I, IGF-II), platelet-derived growth factors
`(PDGF), and the transforming growth factor beta (TGF-β) supergene family, which
`currently has 43 members and includes, among others, TGF-β 1–5 and the bone
`morphogenic proteins, BMP 2–16 (Burt & Law 1994). The proteins of the TGF-β
`superfamily regulate many different biological processes, including cell growth, dif-
`ferentiation and embryonic pattern