13The Material and MechanicalProperties of the Healthyand DegeneratedIntervertebral Disc

Ron Alkalay

13.1. Introduction

The intervertebral disc is an essential part of the basic functional spinal unit(FSU), defined as the disc and two vertebrae with their associated ligamen-tous structures. The biomechanical performance of the spine throughout itslength is the sum of the properties of the individual FSUs. The mainfunctions of the intervertebral disc are to allow motion between adjacentvertebrae while being able to resist compressive and bending loads and, toa lesser degree, torsional and shear loads. The disc provides dynamic loadattenuation as a result of its viscoelastic mechanical properties and, throughits mechanical interaction with the vertebral body, it may have a role in theregulation of cancellous bone structure. In the succeeding sections theanatomy of the disc, the material and mechanical properties of its variousstructures, its overall mechanical behavior, and the effects of degenerationon these properties will be described. The chapter culminates with adiscussion on past and current approaches in the design of artificial discreplacement implants.

Ron Alkalay Orthopaedic Biomechanics Laboratory, Beth Israel Deaconess MedicalCentre, Boston, Massachusetts, 02215, United States.

Integrated Biomaterials Science, edited by R. Barbucci. Kluwer Academic/Plenum Publishers,New York, 2002.

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13.2. Anatomy

The disc is composed of three histologically different, yet functionallyand physically interdependent, elements: the nucleus pulposus, the annulusfibrosus, and the cartilaginous end plates (Figure 13.1).

13.2.1. Nucleus Pulposus

The nucleus is situated at the center of the disc occupying 30–50% of itscross-sectional area. It is composed of an irregular network of collagen fibers,0.10–0.15 mm in diameter, embedded in a granular gel with 70–90% water(Buckwalter, 1982). This extracellular tissue, the framework being a complexmatrix of proteoglycan macromolecule aggregates, is composed of centralhyaluronate filaments and multiple attached aggrecan molecules, with thecomplex stabilized by link proteins (Buckwalter, 1982; Buckwalter et al.,1989; Eyre et al., 1989). The cells of the nucleus are generally large, round,chondrocyte-type cells, which produce mainly collagen type II (Table 13.1).

The proteoglycan network of macromolecules is responsible for theconsiderable potential for swelling of the nucleus by up to more than 200%of its initial volume (Urban and Maroudas, 1981). Experimental andtheoretical work on related tissues, particularly knee (Mow et al., 1990;Holmes et al., 1990), has demonstrated that the proteoglycan network,through its interaction with both the collagen fibers and the interstitialwater, play a major role in the static and dynamic mechanical properties ofthe intervertebral disc (Best et al., 1994; Iatridis et al., 1997).

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13.2.2. Annulus Fibrosus

The annulus, which encloses the nucleus pulposus, is formed from aseries of concentric collagenous laminae embedded in a ground substance(Figure 13.1). In contrast to the nucleus pulposus, the annulus cell popula-tion consist of long, thin, fibroblast-type cells (Buckwalter, 1982; Butler,1989), with the gradual replacement of collagen type I at the periphery byweaker collagen type II in the inner layers (Eyre et al., 1989) (Table 13.1).The inner laminar fibers are connected to the vertebral end plates, while theouter laminar fibers, known as Sharpey’s fibers, are connected to the outerrim of the vertebra. Previously, the fibers of each lamina were described asbeing inclined at about ± 30° with respect to the vertebral transverse planewith fibers from successive laminae perpendicular to each (Galante, 1967;Brown et al., 1957) (Figure 13.1). However, recent studies have suggestedthat, depending on the topographical location and the locations of laminarirregularities, the angle of the fibers shows considerable departure from this30° orientation (Marchand and Ahmad, 1990; Tsuji et al., 1993). Theannulus structure was observed to be highly irregular circumferentially withmany of the layers, up to 40% at any sector of 20°, being noncontinuouswith the laminae exhibiting sharp discontinuities at their termination orinitiation. Similarly, the mean thickness of the lamellae, ranging from 0.14to 0.52 mm, was noted to vary with increasing thickness toward the innerlayers. It is noteworthy that both the observed irregularities and the highestfiber orientation, up to 70°, were most common posterolaterally. Thesestructural irregularities expose the posterolateral areas of the disc to early

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failure and may serve as the underlying cause for the increased pattern ofannular tears observed in these regions (Farfan et al., 1972). With age, thenumber of distinct laminae underwent a significant reduction, while both thethickness and the spacing of intercollagen bundles within an individual layerincreased (Marchand and Ahmad, 1990).

13.2.3. End Plate

The cartilaginous end plates, separating the disc from the adjacentvertebrae, play an important role in disc metabolism (Buckwalter, 1982;Ogata and Whiteside, 1981), in the growth of the vertebrae (Brenick et al.,1980) and in transferring the stresses to the vertebral cancellous bone (Kelleret al., 1987). In the young (under 10 years), the end plates are composed ofan inner growth plate zone, an articular region bordering the intervertebraldisc, and a thin calcified articular cartilage layer adjacent to the diaphysealbone trabeculae (Brenick and Caillet, 1982). With maturation (10 to 20years of age), there is a gradual reduction in the width of the growth plateand thickening of the cartilaginous layer. In the adult, the cartilaginous layeris composed of 0.6 to 1 mm thick hyaline- and fibre-cartilage with a verytight collagen framework arranged parallel to the vertebral bodies (Robertset al., 1989). They possess a relatively low water content, 58%, and a highcollagen content, up to 71% of dry tissue (Roberts et al., 1989; Setton et al.,1993). With increasing age, there is a resorption of the cartilage layer beingreplaced by bone, with the blood vessels at the endplate–vertebral boneborder becoming partially or completely blocked (Brenick and Caillet,1982).

13.3. Material Properties of the Structures of the Disc

13.3.1. Nucleus Pulposus

The material properties of the nucleus have been widely described asthose of incompressible fluid, i.e., the material is unable to sustain appliedstress when not confined. Thus the main thrust of research has been toelucidate the hydrostatic pressure developed in the material in vivo(Nachemson, 1960) and in vitro (McNally et al., 1992), or the swellingpressure in vitro (Urban and Maroudas, 1981; Urban and McMullin, 1988).However, in recent years, results from several experimental studies havesuggested the material to exhibit a more complex behavior. Panagiota-copulos et al. (1987), although employing a limited sample size due to thedifficulties encountered in gripping the specimens, found that the tensile

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relaxation modulus of a nondegenerated material ranges from 30 to 40 kPa(Figure 13.2). Similarly, both the complex dynamic shear modulus (|G*|),derived by dividing the peak shear stress, by the peak shear strain,(Figure 13.3), found to range from 10–50 kPa, and the material energydissipation, measured by the phase angle between the loading andresponse curves, were noted to increase with increasing loading frequency(Iatridis et al., 1997; Bodine et al., 1982). It is noteworthy that the tests

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revealed the phase angle to remain below 45°, indicating more solid-likebehavior. For a material exhibiting a fluid-like behavior, this angle wouldapproach 90°. However, when subjected to a fast ramp-relaxation test, thenucleus exhibited an extremely rapid relaxation to less than 50% of the peakstress in less than one second. This rapid relaxation is then followed by agradual relaxation with the material stress approaching zero at t = 600 s(Bodine et al., 1982). This behavior is indicative of that of a fluid. Despiteexperimental drawbacks which include the need to maintain appropriatehydration levels and significant alterations in the boundary conditionswhich these materials are likely to experience in vivo, these observationssuggest a rate-dependent behavior for the nucleus. At high loading rates, thenucleus behaves as a viscoelastic material, while at low loading rates, itpossesses a more “fluid-like” behavior.

13.3.2. Annulus Fibrosus

The mechanical properties of the annulus have been studied extensive-ly using experimental (Best et al., 1994; Galante, 1967; Brown et al., 1957;Ebara et al., 1996; Skaggs et al., 1993, 1994; Acaroglu et al., 1995; Kirsmeret al., 1996), theoretical (Broberg, 1983, 1993), and computational (Shirazi-Adl, 1989, 1991). The great interest in the mechanical properties of theannulus stems from its role in the overall mechanical behavior of the disc,its involvement in disc herniation, and its potential role in the aetiology oflow back (Eyre et al., 1989; Kelsly, 1980; White et al., 1981). The annulustissue could be regarded as a complex porous fiber-reinforced compositematerial consisting of a dense network of collagen fibers with its composi-tion dependent on the radial location of the lamellae (Table 13.1). Thematrix is made from a dense ground substance, having a heterogeneousstructure made from a wide distribution of constituents, with water itsmost abundant component comprising 65–75% by weight (Kraemer et al.,1985).

Galante (1967) showed the annulus fibrosus tissue to be anisotropicwith tissue strength directly dependent on fiber orientation and locationabout the disc. Tissue samples oriented parallel to the main axis of the fibersexhibited a tensile stiffness threefold greater than those along the transverseplane. The stress–strain tensile behavior of single (Skaggs et al., 1994) andwhole lamellae (Galante, 1967; Ebara et al., 1996) exhibit nonlinear behavior(Figure 13.4a). The stress–strain behavior is best described by a cubicrelationship given in equation (13.1) (Ebara et al., 1996; Skaggs et al., 1994;Araroglu et al., 1995),

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where the constant A corresponds to the tensile modulus of the material atzero strain and the constant B is a measure of the nonlinear dependence ofthe stress–strain curve. The tensile modulus for the tissue along the majoraxis of the collagen fibers, as estimated from the linear portion of thestress–strain curve, ranges from 60 to 140 MPa for single layer specimens(Skaggs et al., 1994) and 1 to 50 MPa for whole lamellae (Ebara et al., 1996;Acaroglu et al., 1995). The failure stress and the failure strain of wholelamellar specimens were reported to range from 1 to 3 MPa (Ebara et al.,1996; Acaroglu et al., 1995) and 10 to 18% (Ebara et al., 1996), respectively.In contrast, the tensile modulus for the tissue perpendicular to the majoraxis of the collagen fibers was observed to be as low as 0.2 to 0.5 MPa(Fujita et al., 1995; Marchand and Ahmed, 1989). The tissue Poisson’s ratiowas estimated to range from 0.46 to 1.63 (Acaroglu et al., 1995) clearlyshowing the tissue to be anisotropic (Figure 13.4b).

An important finding of the above-mentioned studies was the signifi-cant dependence of the estimated parameters, apart from Poisson’s ratio, onthe radial and, to a lesser extent, the anterior versus posterior location ofthe tissue. Tissue specimens taken from the outer annulus exhibited signifi-cantly higher modulus and failure strength than those from the innerlamellae (Skaggs et al., 1993; Acaroglu et al., 1995). Conversely, the tissuestrain to failure was lower in specimens from the outer and anterior annulusthan those from the inner and posterior annulus. In contrast, the tensileproperties of adjacent laminae with alternating fiber orientation fromsimilar locations did not differ distinctly, suggesting the lamina to be thebasic structural unit of the annulus (Skaggs et al., 1994). These marked

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differences may also reflect the higher content of collagen type I in the outerlamellae as compared to the higher content of type II collagen content inthe inner lamellae (Eyre et al., 1989).

The high tensile modulus of the outer lamellae limits the tensile (hoop)stresses created in the disc due to the swelling of the tissue and as a resultof the external loads applied to the disc. Indeed, experimental studies haveshown the outer fibers to undergo relatively small circumferential strains,less than 5% (Shah et al., 1978; Stokes and Greenapple, 1985). The hightensile strength of the outer lamellae, when combined with the inner lamellaslower modulus and higher strain to failure, allows the tissue to create a moreuniform stress distribution across the disc in situ.

A significant determinant of tissue properties was the level of tissuehydration with, at equilibrium, the inner layers being significantly morehydrated than the outer layers (Acaroglu et al., 1995). It is noteworthythat several studies reported the dependence of the tensile (Galante, 1967;Panagiotacopulos, 1987) and compressive (Best et al., 1994) properties onwater content (Figure 13.5). This dependency reflects the reduction in thetissue resistance to solute motion with increasing water content andhighlights the role of the proteoglycan interaction with the collagenfibers in maintaining the latter spatial orientation, thus affecting themechanical behavior of the tissue (Lai et al., 1991). Furthermore, theinteraction of the inner lamellae with the nucleus through the transport offluids resulting in flow-induced dissipation and changes in fluid volumeare thought to provide considerable energy dissipation within the tissue(Best et al., 1994). This mechanism, involving both flow dependent andindependent viscoelastic phenomena with frictional drag caused by the

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interactions between the matrix macromolecules and the interstitialwater, has been coined the “biphasic behavior” of the tissue (Mow et al.,1990).

13.3.3. End Plate

A recent study on the end plates of baboons (Setton et al., 1993)showed the end plate to exhibit a coefficient of permeability, k, of

nearly three orders of magnitude higher than that

tissue failing to reach equilibrium by seconds (Setton et al., 1993). Bycontrast, the annulus fibrosus reaches equilibrium at seconds (Best et al.,1994). It has been suggested that the increased permeability of the tissuewith consequent high creep rate allows for the rapid pressurization of thetissue (Setton et al. 1991). This rapid pressurization, when combined withthe relatively impermeable adjacent vertebral structures, may lead to auniform stress distribution across the end plate. This uniform stress distribu-tion, being resisted by the tight collagen framework (Roberts et al., 1989),could act to facilitate load transfer between the vertebral bodies and theintervertebral discs.

13.4. Mechanical Behavior of the Intervertebral Disc

In term of its in vitro characteristics, the disc exhibits a nonlinearload-displacement curve with two distinct phases similar to that of itscomponents (Figure 13.4a) (Brown et al., 1957). The first phase is charac-terized by large displacements for relatively low axial or bending loads,resulting in minimum energy expenditure required by the muscles to initiateand maintain physiological motion. The second phase is characterized bydecreasing levels of displacement in response to increasing loads until amaximum is achieved. This behavior is the result of the increase in the loadcarried by the solid constitutes of the matrix, i.e. the collagen fibers. Asimilar response is seen when the disc is loaded in torsion (Farfan et al.,1970). Estimated disc stiffness under several load conditions is presented inTable 13.2.

reported for human annulus fibrosus, (Best et al.,1989). This high level of permeability underlies their role in the transport ofwater and solutes from blood vessels at the border of the vertebral bone andend plates into the avascular disc. The end plate tissue demonstrated a highcompressive creep rate, characterized by the sudden and rapid deformationon the application of load, followed by slower but continuous creep with the

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Throughout the day, the disc in a normal subject is exposed toalternating periods of static and dynamic loading. However, due to therelative inaccessibility of the disc to external measurements, little is knownabout both the type and magnitude of such loads in vivo. Nachemson (1960),employing a pressure-sensitive needle inserted into the disc of live volun-teers, estimated the compressive loads occurring at the L3-4 disc level torange from 0.6 times body weight (BW) in standing and sitting to 3.0 BWin a subject holding a weight of 20 kg in the hands. However, as the pressuremeasurements were calibrated against the results from cadaver spinalsegments with their discs likely to have been superhydrated, these resultsmay not portray an accurate picture. There exists a large body of researchwhich has used noninvasive modeling techniques, including electromyogra-phy (McGill and Norman, 1986; Granata and Marras, 1993) and linksegment models (McGill and Norman, 1985; Trafimow et al., 1993) in anattempt to predict the loads occurring on the disc under flexion/extensionand torsional moments. Although these models produce widely varyingpredictions of possible loads, they do suggest that the disc tissue is exposedto complex stresses in response to both single and multiple external andinternal loads. At present, such models still await a more direct experimentalvalidation.

The effect of both the duration and the frequency of the applied loadson the mechanical properties of the disc is well documented, suggesting thatthe response of the disc as a whole is governed by its viscoelastic properties(Burns et al., 1984; Kalpes et al., 1984; Kazarian, 1975; Markolf and Morris,1974; Virgin, 1951). Viscoelastic models composed of ideal Hookian (spring)elements and Newtonian (dashpot) elements and their combinations havebeen successfully employed to model the gross creep and stress-relaxationbehavior of the disc. The former represents the instantaneous elasticresponse of the material with the latter representing the time-dependentresponse of the material. For example, the three-parameter model (Keller etal., 1987) represents the time-dependent strain behavior of the disc via the

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following characteristic equation:

where is the time-dependent strain, is the applied stress, is themean compressive viscous modulus, is the mean compressive modulusof the disc, and t is the relaxation time constant. The model fitted to theexperimental data yielded a mean(SD) value of 6.3 (2.7) MPa and 1.6 (0.9)MPa for and respectively (Keller et al., 1987; Kalpes et al., 1984). Thecoefficient of viscosity obtained from the relationship (13.3) between t,

and was found to have a mean (SD) value of 5.4 (3.9) GPa.s.

Although successful in modeling the experimental creep and stress-relax-ation behavior of the disc, these models cannot account either for thecontribution and time-dependent deformation of individual disc structuresor the effects of interstitial fluid flow. The effect of the latter could beobserved most clearly in the disc response to compressive loads. Initially,the hydrostatic pressure generated at the nucleus caused the disc to bulgesimultaneously radially via the deformation of the annulus and axiallythrough the deformation of the vertebral end plates into the vertebral bodies(Brinckmann et al., 1983; Holmes et al., 1993). It has been suggested thatthis mechanism allows the disc to resist compression while reducing theradial bulge of the disc, thereby minimizing the likelihood of neural cordcompression. However, with time, the nucleus experiences a loss in pressuredue to water transport through the end plates, which causes the disc to loseup to 20% of its volume (Broberg, 1983; Kraemer et al., 1985; Adams andHutton, 1983; Botsford et al., 1994). Of this 20%, up to 25% is due to creepof the collagen fibers in the annulus. The diurnal changes in the hydrationstate of the disc, manifested externally by the 15–25 mm change in stature(Krag et al., 1993), could lead an increasing proportion of the inner annulusto lose up to 36% of its stress-carrying ability (McNally et al., 1992; Adamset al., 1996).

Further insight into the mechanical function of the disc has beenobtained by the use of an instrumented pressure needle to map the pressuredistribution across the disc (McNally and Adams, 1992). In the non-degenerated disc (Figure 13.6), the nucleus and the inner regions of theannulus were observed to act as a “functional nucleus” exhibiting arelatively uniform stress distribution.

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The middle and outer layers of the annulus acted as a “functionalannulus” with the posterior region registering higher compressive stressvalues. Under flexion/extension moments and constant loading conditionsstress, gradients were seen in the pressure profile with high peaks noted forthe posterior annulus and the nucleus pressure decreasing markedly. These

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changes indicate stress shielding by the annulus and the osseous structuresof the vertebra. Under cyclical loads, likely to exist in daily activities suchas driving, the hysteresis demonstrated by the disc was dependent on age,spinal level, and the magnitude and frequency of loading (Kazarian, 1975;Markolf and Morris, 1974).

13.5. The Effect of Degenerationon the Mechanical Properties of the Disc

With age, the intervertebral disc undergoes remarkable changes in itsshape, volume, structure, and composition, causing the alteration of the discmaterial properties and mechanical behavior (Jhonstone and Bayliss, 1995).These alterations are directly related to the changes in the biology of thedisc. Subsequent to the maturation of the spine, the disc exhibits loss ofheight, a gradual replacement of the nucleus with fibrocartilaginous tissuethrough the expansion of the inner annulus, and the appearance of fissuresand cracks in the annulus (Vernon-Roberts, 1987). These gross structuralchanges are accompanied by a sharp decrease in the concentration of viablecells, with the highest reduction observed at the central regions (Buckwalter,1982; Trout et al., 1982a, b) and by a decrease in water and proteoglycanconcentration (Urban and McMullin, 1988). At the same time, the noncol-lagenous protein concentration increases (Dickson et al., 1967) while densegranular material, likely to contain degraded matrix molecules (Buckwalteret al., 1993), accumulates throughout the matrix (Buckwalter, 1982; Troutet al., 1982a, b). The outcome of these processes in conjunction with boththe reduced vascularization of the outer annulus and the ossification of theend plates (Repanti et al., 1998), results in the reduction of the diffusion rateof solutes across the disc (Maroudas et al., 1975; Urban, 1993), furthercompromising cell nutrition. In the elderly, almost the entire disc tissueturns into fibrocartilage with additional loss of height and evidence of deepfissures at the center of the disc (Vernon-Roberts, 1987).

Several studies investigating the effect of degeneration on the mechan-ical properties of the annulus fibrosus tissue reported conflicting results.Galante (1967) and Best et al (1994), reported the tissue tensile andcompressive modulus to exhibit weak correlation with the grade of degen-eration. However, recent studies suggested that the material properties ofthe tissue, including compressive and tensile modulus, failure strength,Poisson’s ratio, and strain energy, were strongly affected by the grade of discdegeneration (Ebara et al., 1996; Acaroglu et al., 1995; Iatridis et al., 1998).Age, although found to affect the tensile properties of the tissue (Galante,1967; Ebara et al.), has far less influence than that of disc degeneration (Best

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et al., 1994; Ebara et al., 1996; Skaggs et al., 1994; Acaroglu et al., 1995).These material changes affect the gross behavior of the disc, namely, thereduction in the energy dissipation of the disc (Kazarian, 1975), increasedcreep rate (Keller et al., 1987; Virgin, 1951), and the increase in tissuestresses under dynamic loads (Kasra et al., 1982). The changes in tissueproperties, as reflected by the changes in the overall mechanical behavior ofthe disc, indicate the transference of tissue loading from a uniform “fluidpressurized” mechanism to an alternative nonuniform mechanism whichpredominately loads the solid matrix. Further evidence of this transferencecould be observed by the reported decrease in radial swelling pressure withdegeneration and by the significant changes in the pressure distributionacross the disc (McNally and Adams, 1992) (Figure 13.7).

These changes could cause local buckling and, ultimately, separationof the laminae with the inner laminae buckling toward the nucleus (Adamset al., 1993), a phenomenon observed in degenerated and failed discs(Gunzburg et al., 1992; Tanaka et al., 1993). This structural disruption andthe accompanying changes in hydrostatic pressure, demonstrated to causecell-mediated degenerative changes (Osti et al, 1990; Pfeiffer et al., 1994), arelikely to have a marked effect on chondrocyte metabolism and proteoglycansynthesis in the annulus (Bayliss et al., 1988; Hall et al., 1991). The increasedloading of the solid matrix, when combined with the changes in tissueproperties, may explain the increased predisposition of the annulus tissue topremature failure.

13.6. Intervertebral Disc Prostheses

At present, treatment for a severally degenerated disc with neurologi-cal involvement ranges from a removal of nucleus material for the bulgingdisc to a complete removal of the disc and the creation of bone fusionbetween the two bounding vertebrae in the case of a severally ruptured disc.These procedures, although relatively effective in alleviating pain, predisposethe adjacent segments to early degeneration due to increased functionaldemands (Lee and Langrana, 1984; Leong et al., 1983). In order for anarticulating device to present a significant advantage over spinal fusion, asuitable design needs to comply with a stringent set of safety, functional,dynamical, and material based requirements for up to 40 years of operation.Kostuik (1997) and Lemaire et al., (1997) provide a detailed discussion ofsome of these requirements. Clearly, due to the intimate relationshipsbetween the disc space and the neural cord and major blood vessels, theoperational safety of these devices is of utmost importance.

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For the past 35 years, large numbers of alternative designs for totaldisc replacement have been proposed. Low friction designs, composed ofsliding congruent surfaces made from either metal (Patil, 1982) and/or acombination of metal and polymeric materials (Salib and Pettine, 1989),although providing a degree of flexibility, lack inherent kinematical con-

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straints. This results in reduced “stability” to the affected spine and,ultimately, may cause early failure due to wear. An all-metal spacer with aposterior “floating” hinge connecting two cobalt chromium platens with twointerposed coiled titanium springs was proposed by (Hedman et al., 1988,1991). This device possesses high fatigue strength, primary flexion/extension,and a degree of lateral bending motion, and was shown in a sheep model toperform successfully in the short term with no evidence of fibrous tissuegrowth (Kostuik, 1997). However, no clinical application of this device hasbeen reported.

The majority of these alternative designs have employed single ormultiple components made from a combination of polymeric elastomerswith or without metal components and have a geometry closely related tothat of the natural disc (Stubstad et al., 1975; Edeland, 1989). Examples ofcomposite disc designs included end plates made of silicone elastomer witha fluid filled core element bonded by Dacron fibers (Stubstad et al., 1975)and two polyurethane end plates with an interposing porous silicone(Edeland, 1989). In an attempt to replicate disc structure, Lee et al. (1990)and Parsons et al. (1992) proposed a design with a soft elastomer core whichwas encircled by reinforced fiber sheets. These fiber sheets have alternat-ing fiber orientation arranged in six to fifteen lamellae, are embeddedin a second polymer, and are capped by two metal end plates. A similarprinciple was employed by Steffee (1991), using a hexene-based, carbon-black-filled polyolefin, rubber core, vulcanized to two titanium end plates.Currently, the most successful design, known as the SB Charite’III,combines two concave end plates made of cobalt chromium, with a bi-convex contoured polyethylene oval spacer (Büttner-Janz et al., 1989;Zippel, 1991). In a clinical multicenter retrospective study, Griffith (1994)found that up to 70% of patients reported a resolution or a reduction inpreoperative lag pain.

A second approach, particularly in the case of an early degenerateddisc with structurally intact annulus, has been to develop a prosthesis toreplace the degenerated nucleus. This approach has the advantage of beingless invasive and, depending on the degenerative state of the disc, mayrestore the function of the annulus. Past endeavors attempted to fill thenucleus void ether with polymeric materials such as self-curing silicon(Scheider and Oyen, 1974; Roy-Camille et al., 1978; Fassio and Ginestie,1978) and contained systems made from either a polymer sac filled with anincompressible medium (Baumgartner, 1992) or a fiber wave based materialfilled with thixotropic gel (Ray and Corbin, 1990). Bao and Higham (1993),in an attempt to mimic the fluid transport in a disc, in addition to itsmechanical performance, employed a hydrogel-based material having 70%water content. Although this design resulted in a biomechanical behavior

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similar to that of a natural disc (Odrdway et al., 1994), no clinicalapplication has been reported. Recently, a novel bioelastic polymer manu-factured using genetic engineering methods, which is able to imbibe up to25% water thus causing it to swell, was employed to restore the mechanicalfunction and anatomical relationship of a denucleated lumbar disc (Alkalayet al., 2001) Preliminary tests, which suggested that the material restored theaxial compressive stiffness and hysteresis to that of the intact disc, indicatedthat the use of such a polymer could provide immediate restoration of discfunction and anatomy while exhibiting excellent long term biocompatibility.In recent years, the rapid development, in tissue engineering and tissueregenerative methods, is being increasingly explored for the treatment ofearly degenerated disc. These methods use sophisticated three-dimensionalscaffolds, seeded with cells and growth factors, and employ viral vectors totransfect disc cells to stimulate the production of necessary proteins to affectthe regeneration of disc tissue. Although these methods face many difficultchallenges with respect to their effective delivery, integration, and regener-ation of the tissue, and their short- and long-term viability, ultimately theyhold great promise.

13.7. Summary

In this chapter, the anatomy and material properties of the interver-tebral disc as a whole and its individual structural components werereviewed. The contribution of each of these structures to the mechanicalbehavior of the intervertebral disc was discussed, with the effects of degen-eration on these structures highlighted. The state of hydration and theinteraction of charged structural macromolecules with the interstitial fluidflow were demonstrated to critically affect the compressive viscoelastic creepstress-relaxation and dynamic load attenuation of the disc. These mechan-isms, in conjunction with the intrinsic nonlinear properties of the collagenproteoglycan matrix, which form the structures of the annulus and vertebralend plates, underlie the anisotropic properties of the disc. Thus, the variedfunctions of the intervertebral disc are intimately related to its biomechani-cal composition and architectural arrangements. The process of disc degen-eration was found to affect significantly the biology and structure of theintervertebral disc, thereby leading to marked changes in its mechanicalbehavior. Similarly, mechanical fatigue was demonstrated to cause a break-down of the tissue, thus altering the state of stress–strain in the disc. Bothof these mechanisms were suggested to disrupt the nutritional pathways andto accelerate tissue catabolism, culminating in the failure of the disc. The

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efforts to devise a whole intervertebral disc or nucleus prosthesis have, to alarge extent, met with limited success.

At present, our knowledge pertains predominantly to the in vitroquasi-static behavior of the intervertebral disc. Such a condition seldomoccurs in daily life. New developments in noninvasive techniques, particu-larly MRI imaging, may prove valuable in extending our knowledge of thefunction of the intervertebral disc in vivo and the effect of degeneration onthese functions. The use of such techniques may well point to possibletreatments of some of the underlying causes of disc degeneration and spinalback pain. When combined with new materials and biologically basedengineering methods, the developmental effort to design intervertebral discprosthesis and to reverse the effects of degeneration directly could showgreater success than that which has been achieved thus far.

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Bao, Q.B., Higham, P.A. 1993. Hydrogel intervertebral disc nucleus, US Patent 5,192,326.Baumgartner, W. 1992. Intervertebral prosthesis, US Patent 5,171,280.Bayliss, M.T., Jhonstone, B., O’Brien, J.P. 1988. Proteoglycan synthesis in the human interver-

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