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DOI: 10.1177/154411130301400207 2003 14: 138CROBM

Eiji Tanaka and Theo van EijdenBiomechanical Behavior of the Temporomandibular Joint Disc

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138 Crit Rev Oral Biol Med 14(2):138-150 (2003

Introduction

Since 1980, numerous studies have been conducted to assistinvestigators in comprehending the biomechanics of thetemporomandibular joint (TMJ). Experimental and modelstudies have verified that the TMJ is loaded during function(macaque—Hylander and Bays, 1978, 1979; Brehnan et al.,1981; Boyd et al., 1990; human—Hatcher et al., 1986; Smith et al.,1986; Throckmorton and Dechow, 1994). However, detaileddata about the distribution of the loads are still lacking.

Mathematical models of the human masticatory system,including the TMJ, have served as powerful tools to predict theloads acting on this joint. These models have yielded valuableinformation on, for example, the role of the disc and that of other tissues of the TMJ in its functioning. Many studies, how-ever, have oversimplified the geometry of the articular surfacesand assumed them to be rigid (e.g., Koolstra et al., 1988; Ferrarioand Sforza, 1994). Therefore, the deformations and the distri-

bution of loads inside the joint could not be analyzed.The articular surfaces of the TMJ are highly incongruent.

Due to this incongruence, the contact areas of the opposingarticular surfaces are very small. When joint loading occurs,this may lead to large peak loads, which may cause damage tothe cartilage layers on the articular surfaces. The presence of a

fibro-cartilaginous disc in the joint is believed to prevent thesepeak loads (Tanne et al., 1991; Scapino et al., 1996), since it iscapable of deforming and adapting its shape to that of the artic-ular surfaces. These deformations ensure that loads areabsorbed and spread over larger contact areas. In addition, theshape of the disc and the area and location of its contact areaswith the articular surfaces change continuously during jawmovement to adapt to the changing geometry of the articularsurfaces of the mandible and temporal bone. As a result, therewill be a continuous change in the magnitude and location of the deformations that occur. For example, according to the

work of Beek et al. (2001b), when loading occurs in the jawclosed position, the deformations in the disc are spreadthroughout the entire intermediate zone, while translation othe condyle in the forward direction to obtain a protrusive oropen jaw position leads to a concentration of the deformationin the lateral part of the disc. This suggests that certain areas ofthe disc are more heavily loaded than other areas.

The magnitude of the deformation and resulting stress othe disc is primarily determined by the nature of the appliedloads and by the biomechanical properties of the disc, such as

stiffness and strength. An understanding of these properties isimportant for several reasons. First, they determine the role othe disc as a stress-distributing and load-absorbing structure(Nickel and McLachlan, 1994; Beek et al., 2001a). Therefore, theproperties of the disc will also influence the stresses and strainsthat occur in the cartilage layers on the bone surfaces. Thesestresses and strains are of critical importance for adaptation andwear. For example, mechanical stress affects the proteoglycansynthesis in the disc (Carvalho et al., 1995), resulting in an adaptation of stiffness. Second, precise information on the biomechanical properties of the disc is required to develop suitable

joint simulation models, with which the distribution of stresand strain in the structures of the joint can be estimated. In thelast decade, several three-dimensional finite element models o

the joint have been developed (Korioth et al., 1992; Tanaka et al.1994, 2001c; Nagahara et al., 1999; Beek et al., 2000). Howeverthus far, the available models do not include all relevant properties, such as the shock-absorbing capabilities of the disc. Finallyinformation on the biomechanical properties of the disc is indispensable for the development of replacement materials for TMJprostheses.

In this paper, the fundamental concepts of the biomechani-cal behavior of the TMJ disc are reviewed. The review is divid-ed into four parts. The first part introduces some basic defini-tions and the general physical properties of soft tissues, to facil-

BIOMECHANICAL BEHAVIOR OF THETEMPOROMANDIBULAR JOINT DISC

Eiji Tanaka1*Theo van Eijden2

1Department of Orthodontics and Craniofacial Developmental Biology, Hiroshima University Graduate School of Biomedical Sciences, 1-2-3 Kasumi, Minami-ku, Hiroshima 734-8553, Japan; an2Department of Functional Anatomy, Academic Center for Dentistry Amsterdam, Meibergdreef 15, 1105 AZ Amsterdam, The Netherlands; *corresponding author, [email protected]

ABSTRACT: The temporomandibular joint (TMJ) disc consists mainly of collagen fibers and proteoglycans constrained in theinterstices of the collagen fiber mesh. This construction results in a viscoelastic response of the disc to loading and enables thedisc to play an important role as a stress absorber during function. The viscoelastic properties depend on the direction (tension,compression, and shear) and the type of the applied loading (static and dynamic). The compressive elastic modulus of the discis smaller than its tensile one because the elasticity of the disc is more dependent on the collagen fibers than on the proteoglycans. When dynamic loading occurs, the disc is likely to behave less stiffly than under static loading because of the differenceof fluid flow through and out of the disc during loading. In addition, the mechanical properties change as a result of variousintrinsic and extrinsic factors in life such as aging, trauma, and pathology. Information about the viscoelastic behavior of thedisc is required for its function to be understood and, for instance, for a suitable TMJ replacement device to be constructed. Inthis review, the biomechanical behavior of the disc in response to different loading conditions is discussed.

Key words. Stress and strain, viscoelasticity, elastic modulus.



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itate comprehension of the later parts, in which the biomechan-ical behavior of the TMJ disc is discussed. Part 2 relates this

behavior to the composition and organization of the disc, suchas its collagen fiber and proteoglycan contents. In part 3, the bio-mechanical properties of the disc, including elastic modulusand viscoelasticity, are summarized. Finally, in part 4, adaptivechanges of the disc in relation to its biomechanics are discussed.

(I) Nature of Loading Applied to the TMJ Disc

(A) ELASTICITY UNDER VARIOUS LOADINGSThe disc is subject to a multitude of different loading regimensduring mandibular movements. Basically, three types of load-ing can be distinguished: compression, tension, and shear (Fig.1). Obviously, during natural loading of the joint, combinationsof these basic types of the disc loading do occur. During com-pressive loading, the disc becomes shorter in the loading direc-tion; during tensile loading, it is stretched in the loading direc-tion; and during shear loading, one boundary surface of thedisc moves parallel to an adjacent surface. During every type of loading, the disc undergoes a deformation, while internalforces are produced within the tissue. The amount of deforma-tion is quantified by the amount of strain, which is defined as

the change in length per unit of original length. The internalforces are quantified by the amount of stress, which is definedas force per unit area in Pa (1 Pa = 1 N/m2).

The relationship between stress and strain of an elasticmaterial can be described by a stress-strain curve (Fig. 2). Thecurve has both elastic and plastic deformation regions. If thestructure is not loaded beyond the elastic region, it will returnto its original shape once the load is released. If the structure isloaded up to its plastic region, it will not return to its originalshape when the load is released. After plastic deformation, thestress will cause permanent damage of the tissue. In the elasticregion, a toe region and a transition zone can be distinguished(Fung, 1981; Li et al., 1983), which are more or less arbitrarilydivided by the so-called critical point (Tanne et al., 1991; Tanaka

et al., 2000). The toe region can be considered as the physiolog-ic range of stress and strain in which the tissue normally func-tions (Fung, 1981). Beyond the toe region, in the transitionzone, the tissue usually will have a large reserve of strength

before it ruptures and fails. The physiologic range can be verydifferent for various connective tissues (mesentery, 100-200%;skin, 40%; tendon, 2-5%; Fung, 1981). The physiologic range of the disc is reported to be approximately 4% (canine, tensilestrain; Teng et al., 1991).

Another important feature is the strength of a tissue. Theyield strength is defined as the stress at the yield point, beyondwhich deformation causes tissue damage. This point separates theelastic strain region and the plastic region of the stress-straincurve. The ultimate strength is the stress that the tissue can maxi-

mally sustain; the breaking strength is the stress at which the tis-sue will break. The value of the ultimate strength of the discdepends on the direction of the applied stress and the regionwhere it is applied. For example, the ultimate strength of the inter-mediate zone of the disc is 37.4 MPa (1 MPa = 106 N/m2) when atensile stress is applied antero-posteriorly, while it is 1.6 MPawhen the stress application is medio-lateral (porcine disc; Beatty etal., 2001); the ultimate strength is significantly larger in the anteri-or and posterior regions of the disc (46.7 MPa and 69.7 MPa) thanin its central region (14.7 MPa) (canine disc; Teng et al., 1991).

For evaluation of the basic biomechanical characteristic of

a tissue, the elastic modulus or Young's modulus E is common

ly calculated. This modulus is defined as the slope of the elas-tic region (almost linear part) of the stress-strain curve. The tensile and compressive moduli are a measure of the ability of thetissue to resist deformation in the direction of the applied loadThese are defined as E = /, where is the stress and is thestrain. The shear modulus G is a measure of the ability of thetissue to resist shear stress in a particular plane. It is defined asG = /, where is the shear stress and is the shear strainHere, the shear strain is the displacement in the direction of theapplied force per original thickness (Fig. 1). In general, the shearmodulus tends to be 1/3 to 1/2 of the value of the tensile or

14(2):138-150 (2003) Crit Rev Oral Biol Med 139

Figure 1. Diagram showing the different types of strain for threedirections of loading. During compressive loading, the disc becomesshorter in the loading direction; during tensile loading, it is stretchedin the loading direction; and during shear loading, one boundarysurface of the disc moves parallel to an adjacent surface.

Figure 2. Typical stress-strain curve for connective tissue. The elasticand plastic regions of the curve are divided by the yield pointbeyond which deformation causes tissue failure; the elastic region isfurther divided into a toe region and a transition zone.





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compressive modulus.When the disc is compressed or stretched in one direction,

not only will it deform in that direction (primary strain), but itwill also become thicker or thinner, respectively, in a directionperpendicular to it (secondary strain). Poisson's ratio is a mea-sure of the ability of a structure to resist deformation in a direc-tion perpendicular to that of the applied load. Poisson's ratio(v) is defined as v = y/x, where y is the secondary strain andx is the primary strain. Because the volume does not increaseupon loading, the Poisson's ratio is less than 0.5.

The elastic constants of a tissue (elastic modulus, shearmodulus, and Poisson's ratio) describe the relationship

between a load placed on the tissue and the resulting deforma-tion within the elastic range. If the elastic constants have thesame values in different loading directions, the tissue is calledisotropic. For an isotropic material E, G and v are related as:

EG = _________

2(1 + v)

However, the elastic constants are generally not equal in alldirections. This directional dependency is called anisotropy.

(B) VISCOELASTICITY

The disc exhibits not only elastic but also viscous characteristics.The combination of these effects is called viscoelasticity, i.e., if thedisc is subjected to a constant force or deformation, its responsevaries over time. This results in a characteristic stress-strain

behavior when the disc is loaded, i.e., a rapid deformation uponload application, followed by a time-dependent creep or stress-relaxation phase of increasing deformation or reducing stress(Fig. 3). Creep is defined as the strain over time with constantstress; stress-relaxation is the decrease of stress in time with con-stant strain. Here, the viscoelastic properties are mainly the

result of fluid flow through and out of thedisc. Immediately after loading occurs, thesmall permeability of the collagen networkimpedes instantaneous fluid flow throughthe collagen network. With time, the loadcauses the fluid to be driven away from theloaded site, through pores in the collagennetwork (Scapino et al., 1996). This fluidflow also explains another feature of vis

coelasticity, i.e., that the biomechanica behavior of the tissue is dependent on thestrain rate and on the time after stress application. For instance, when the strain andstress are applied rapidly, the slope of thestress-strain curve will be steeper thanwhen they are applied slowly.

(a) Stepwise loading

How the viscoelastic properties of the discchange over time during constant loadingcan be characterized by stress-relaxationtests, creep tests, and restoration tests (Fig3). The parameters obtained from thesetests provide valuable information on thetissue behavior as a function of time andare of great importance for our betterunderstanding of mechanical properties o

the disc, such as energy dissipation and stress absorption. In astress-relaxation test, a stepwise deformation with a specificstrain level is applied to a specimen, and this strain level is keptconstant until the stress reaches an almost steady level. Fromthis test, the relaxation time and the relaxed modulus areobtained. The relaxation time is defined as the time in which thesteady level of stress is reached. One calculates the relaxed modulus by dividing the stress value after relaxation, i.e., the relaxedstress, by the specified strain. In a creep test, a certain amount ofstress is applied instantaneously and kept constant, and thetime-dependent changes of strain are observed. In a restorationtest, the stress and strain are observed after removal of the loading. The remaining deformation is called residual strain.

In a viscoelastic material, the stress-strain curve duringloading is essentially different from that during unloading, andthis feature is called hysteresis (Fig. 3). The area enclosed by thestress-strain curves during loading and unloading is a measureof the amount of energy dissipation. The lower the energy, themore elastic the response of the tissue. The hysteresis energymay dissipate as heat or in the drag of the fluid that is exudedand absorbed during loading and unloading, respectively.

There are many physical models to represent viscoelastic behavior during creep and stress relaxation. In these modelsthe elastic behavior is usually represented by a spring with an

elastic constant. The spring produces a resistance which is proportional to the applied force and determines the stress leveafter stress-relaxation. The viscous behavior is represented by adashpot with a viscous constant. Although the dashpot exhibitsresistance to stress at the onset of stress application, the resilienforce disappears gradually. Of the many available models, thefollowing three models are commonly used (Fig. 4). Maxwell'smodel is a serial combination of a linear spring and a dashpoand is used to characterize restoration and hysteresis after stressremoval (Fung, 1969). Voigt's model is a parallel combination oa linear spring and a dashpot and is used to characterize creep


Figure 3. Stress-relaxation (A) and creep and restoration curves (B). In the hysteresis curve(C), the area enclosed by the stress-strain curve is a measure for the energy dissipation.





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features (Fung, 1969). Kelvin's model is a combination of Maxwell's and Voigt's model andhas been applied to characterize both creepand stress-relaxation behavior (Fung, 1969).The relationship between stress and strain at a particular time is formulated as +

· =ER( +

· ) where and are time constants,and ER is the relaxed modulus after stress-relaxation; the time constant is a measure of

the relaxation time. These parameters are com-monly used in studies describing the vis-coelastic properties of the disc.

(b) Cyclic loading

The loading on the disc can be classified intotwo types: static and dynamic loading. Staticloading occurs, for example, during clenching,grinding, and bruxism; dynamic loadingoccurs during, for example, talking and chew-ing. The methods mentioned above are notsuitable for evaluation of the characteristics of the disc during dynamic loading. This requiresassessment of the responses to cyclic loading

at a wide range of frequencies. In dynamictests, a cyclic or sinusoidal stress is commonlyproduced for determination of the behaviorduring dynamic loading (e.g., Beek et al.,2001a; Tanaka et al., 2002b).

Under cyclic loading, the disc will quick-ly settle into a steady-state response. Usually,this steady state is achieved in fewer than 10cycles (Beek et al., 2001a). Sometimes, there isalso some slow secondary creep of the steady-state cycle.However, this is generally negligible compared with the differ-ences in response over the first few cycles of loading. The sizesof the hysteresis loops of the stress-strain curves may changedramatically during the first few cycles of loading and unload-

ing. This is called pre-conditioning (Fung, 1972). Cycle tests of connective tissue specimens at a constant strain rate and atmoderate strain levels will show repeatable stress-strain rela-tionships within 7 to 10 cycles of pre-conditioning. The hys-teresis loop during loading and unloading, at a constant rateafter pre-conditioning, will remain essentially unchanged.

Due to the viscoelastic behavior of the disc, the stressresponse to a cyclic strain is, in general, out of phase, and thephase difference between the stress and strain is somewhere

between 0 and 90º. The schematic representation of the rela-tionship between stress and strain, with a sinusoidally varyingstress, for a perfectly elastic material, a viscoelastic material,and a perfectly viscous liquid is shown in Fig. 5. If the materialis perfectly elastic (Hookean body), the strain is exactly in

phase with the stress, which implies that the phase difference is0º. If the material is perfectly viscous (Newtonian fluid), thestrain is 90º out of phase with stress. For viscoelastic materials,the phase angle is somewhere in between 0 and 90º. The com-plex dynamic modulus E* can be determined experimentally

by the application of a sinusoidal strain (Murata et al., 2000).The modulus E* consists of a real part, the storage modulus E,and an imaginary part, the loss modulus E, shown vectoriallyin Fig. 5. The magnitude of the complex modulus E* is deter-mined by E* =/. With the phase angle , the storage andloss moduli, E and E, are determined by


Figure 4. Relaxation and creep function of (A) Maxwell, (B) Voigt, and (C) Kelvin mode(standard linear solid). Maxwell's model and Voigt's model consist of only two elements which are a linear spring with a spring constant and a dashpot with a coefficient o viscosity , whereas Kelvin's model is composed of a combination of two springs and adashpot. A linear spring produces an instantaneous deformation proportional to thestress. A dashpot produces a velocity proportional to the stress at any instant.

Figure 5. Schematic representation of the relationship betweenstress and strain during a sinusoidal oscillating strain (, angular velocity) for a perfectly elastic solid ( A , Hookean body), a vis-coelastic material (B), and a perfectly viscous liquid (C, Newtonianbody). In a viscoelastic material, the phase difference betweenstress and strain is somewhere between (/2 > > 0), and thecomplex modulus E* is resolved into two components, i.e., the stor-age modulus E and the loss modulus E, shown vectoriallyFurthermore, the tangent of the phase angle () between stress andstrain is a measure of the ratio of energy loss to energy stored dur-ing a cyclic deformation.





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E* = E + iE

where E = E* cos, E = E* sin, i = - 1, and tan = E/E isthe loss tangent. E describes the elastic deformation under stressand is directly proportional to the energy storage in a cycle of deformation. E denotes the viscous deformation and is propor-tional to the average dissipation or loss of energy as heat in acycle of deformation. In addition, the tangent of the phase angle

() between stress and strain, i.e., the loss tangent (tan), is a mea-sure of the ratio of energy loss to energy stored during cyclicdeformation. For a tissue with a high value of the loss tangent,the viscous behavior is stronger than the elastic behavior. In sucha tissue, the energy used for its deformation is dissipated as heatand causes changes in the inner structure by movement of fluid.

(II) Biomechanical Properties of the DiscRelated to its Composition and Organization

The disc is composed of variable amounts of cells and an extra-cellular matrix. The matrix consists of macromolecules and fluid.The macromolecules constitute about 15-35% of the wet weight of the disc, while the tissue fluid constitutes about 65-85%. Thesemacromolecules consist mainly of collagen (85-90%) and proteo-

glycans (10-15%) (Nakano and Scott, 1989a; Sindelar et al., 2000).The mechanical properties of the disc are largely dependent on itscollagen fiber and proteoglycan composition and organizationand on their interaction with the tissue fluid.

(A) COLLAGENThe collagen fibers maintain the shape of the disc, while elastinrestores shape during unloading (Scapino et al., 1996). Collagenfibers commonly exhibit waviness ("crimping"). The functionalsignificance of this crimping can be seen in the stress-strain curve(Fig. 2). When a tension is applied to the disc, the first effect is to

straighten the crimp, and that accounts for theinitial toe region of the curve (Stegenga et al.1991; Berkovitz, 2000). Thus, the initial toeregion reflects a high compliance and corre-sponds to the straightening of the crimpingwithout any lengthening of the collagen fibers(Gathercole and Keller, 1991). Beyond this initial phase, the collagen fibers begin to extendand become load-bearing. Furthermore, the

small permeability of the collagen networkimpedes interstitial fluid flow through this network (Mow et al., 1984, 1993). Therefore, theloads acting on a cartilaginous structure as thedisc are initially transmitted by a pressurizationof the incompressible fluid without muchdeformation of the collagen network (Soltz andAteshian, 1998). Nonetheless, fluid flowthrough the collagen network leads to a gradual transfer of the load from the fluid to the collagen fibers. When further loaded, the collagennetwork deforms, and water is squeezed out ofthe disc while the orientation of the collagenfibers is re-arranged (Mow et al., 1986; Woo

1986). The movement of fluid out of the discand the re-arrangement of the collagen fibersare reversible when the disc is not deformed

beyond the physiologic strain range. Evenapplication of significant long-term stresses

beyond the physiologic strain range introduce but minor changes in fiber waviness and alignment within thedisc (Scapino et al., 1996). This enables the disc to adapt its shapecontinuously to fit in the space between the opposing articularsurfaces and to distribute loads suitably in the TMJ. Collagengives the disc much of its tensile stiffness and strength.

The thin surface layers of the disc have an architecture differentfrom that of the thick inner layer (Fig. 6). In the superior and inferi-or surface layers, the collagen fibers are more or less perpendicular-

ly arranged in an antero-posterior and medio-lateral direction(Minarelli et al., 1997). In the inner layer, the orientation of collagenfibers varies markedly in different regions of the disc. The fibers runprimarily antero-posteriorly in the intermediate zone and mediolaterally in the anterior and posterior bands. The antero-posteriorfibers from the intermediate zone are interlaced with the medio-lat-eral fibers in both bands (Teng and Xu, 1991). In the central regionof the bands, the fibers from the intermediate zone flare superiorlyand inferiorly and turn medially and laterally, merging structurallywith those of the bands (Mills et al., 1994; Scapino et al., 1996). In themedial and lateral regions of the disc, near the condylar poles, theantero-posterior fibers of the intermediate zone are attached tightlyto the poles of the condyle (Teng and Xu, 1991). As a result of thesedifferences in collagen fiber orientation, regional differences and

anisotropy in the mechanical properties of the disc can be expected(B) PROTEOGLYCANS

Proteoglycans consist of a core protein to which glycosaminoglycan (GAG) sulfate side-chains are attached (Iozzo, 1998). In thedisc, there are several proteoglycans. Biglycan and decorin belongto the group of small proteoglycans, consisting of a core protein oapproximately 38 kDa to which either one (decorin) or two (biglycan) chondroitin/dermatan sulfate side-chains are attached(Chopra et al., 1985; Fisher et al., 1989). Aggrecan is a large proteo-glycan containing both chondroitin sulfate and keratan sulfate


Figure 6. Schematic illustration of the disc (antero-lateral view) with various orientationsof collagen fibers.





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(Nakano and Scott, 1989b). Sulfated GAGs such as chondroitinsulfate, dermatan sulfate, and keratan sulfate are synthesized asproteoglycan in the Golgi apparatus. Hyaluronic acid, anotherGAG, is not attached to a core protein and is synthesized as a freeGAG molecule in the cell membrane (Iozzo, 1998).

Proteoglycans are enmeshed in the network of collagen fibersand are virtually immobile. Proteoglycan molecules possess ahigh viscosity and a large molecular size that reduce their capaci-ty to diffuse through the collagen network, thus resulting in the

retention of large amounts of water (Muir, 1973). The result is astiff viscoelastic material surrounding the collagen fibers. Becauseof their molecular structure, proteoglycans are ideally suited toresist compressive loadings. Proteoglycans can indirectly modu-late the stiffness of the collagen network, since an increase in pro-teoglycan concentration leads to an increase in the osmotic pres-sure, which in turn affects the synthesis of collagen (Muir, 1981).

In the disc, the large proteoglycans and the related chondroitinsulfate are preferentially localized in the central area of the intermedi-ate zone and in the anterior and posterior bands (Mills et al., 1988;Nakano et al., 1993; Nakano and Scott, 1996; Mizoguchi et al., 1998). Ithas therefore been suggested that these parts of the disc encounterheavy compressive loading during function and may be responsiblefor maintaining the resilience of the disc (Nakano and Scott, 1996). The

small proteoglycans, decorin and biglycan, are mainly found in the lat-eral and medial parts of the intermediate zone and are present in less-er amounts in the central part of the intermediate zone and in the ante-rior and posterior bands (Scott et al., 1995; Mizoguchi et al., 1998). Theexpression of decorin mRNA in tendon is promoted by tensile stress(Robbins and Vogel, 1994). Consequently, the distribution of decorinmay reflect the distribution of tensile stress in the disc. Decorin can alsointeract with type I collagen and cause small increases in fibrillar diam-eter (Scott et al., 1995). This is in line with the observation that the col-lagen fibrils in the decorin-rich peripheral region of the disc are thethickest (Kuc and Scott, 1994). These biochemical findings for the pro-teoglycan concentration in different areas of the disc are consistentwith results obtained from mechanical tests (see Part III).

(C) TISSUE FLUID

The tissue fluid is a viscous gel and contains mostly water. Thisfluid can move both inside and through the surface layer of thedisc. The collagen and proteoglycans are dispersed in the fluid,making the cartilage a microporous material with a certain per-meability. The amount of permeability is particularly significantfor compression, since the mechanical response of the disc willdepend on it (Beek et al., 2001b). A low permeability means thatany substantial exchange of fluid between the inside and outsideof the disc must take place over a substantial period of time (e.g.,minutes) compared with the physiological loading cycle (1 sec).As a consequence, the disc will maintain its stiffness under com-pression. In case of a high permeability, a rapid exchange of fluidis possible, which results in a substantial decrease of stiffness.

(III) Biomechanical Behavior of the Disc

(A) ELASTIC CONSTANTS

Many studies have been conducted on the elastic properties of the disc since Fontenot's initial investigation (1985). In general,the elastic modulus has been measured on small specimens bymeans of static tests (Tanne et al., 1991; Teng et al., 1991; Chin etal., 1996; Lai et al., 1998; Tanaka et al., 1999, 2000, 2001a; Beattyet al., 2001; del Pozo et al., 2002). However, due to the largeinterspecies variations and different experimental protocols,

the results of the various studies cannot be easily comparedDifferences in experimental techniques include the size, hydra-tion fluid and holding of the specimens, and different testingmachines and protocols. Due to this variation, the reportedmoduli show a large range (from 1 to 100 MPa). Most information has been obtained from human and from porcine and

bovine discs, because of their structural and functional similarity to the human disc. Structurally, the bovine and porcine TMhave a more or less similar shape, and their mandibular halves

are rigidly fixed, as in the human (Bermejo et al., 1987, 1993González et al., 1991). Functionally, the bovine and porcine TMexhibit lateral translatory movements (Berg, 1978; Hatton andSwann, 1986). The masticatory pattern of pigs is different fromthat in humans and cattle (Langenbach and Van Eijden, 2001)Pigs chew faster than humans and cattle, but the pattern of pigchewing is probably more similar to that of the human than tothat of cattle; cattle have much greater excursions than humansIn addition to these species, information has been obtainedfrom canine discs. However, functionally, the dog exhibitsmainly chopping strokes during mastication.

Another problem when the results of various studies arecompared is that different tests (for example, compression andtension) have not been performed on the same specimen. Also, as

mentioned before, the relationship between stress and strain forthe disc is non-linear and time-dependent. For example, the tensile modulus of the porcine disc is about 27 MPa at a strain rate o0.5 mm/sec, whereas it is about 83 MPa at a rate of 500 mm/sec(Beatty et al., 2001). Thus, when data on elastic moduli are evaluated, consideration should also be given to the strain rate, themagnitude of the applied strain, and the measuring time. In addi-tion, because of structural differences within the disc, the locationof the loading and its direction and type (tension, compressionshear) are major factors for determining its elastic properties.

(a) Tensile modulus

The tensile modulus is mainly dependent on the orientation ocollagen fibers, because they can resist tension only in the direc-

tion parallel to their orientation. As mentioned in Part II, the intermediate zone of the disc consists mainly of antero-posteriorly oriented fibers. Therefore, the tensile modulus and tensile strengthof the intermediate zone are larger in the antero-posterior direction than in the medio-lateral direction (Beatty et al., 2001; Tanneet al., 1991; Teng et al., 1991). For example, the tensile modulus ofthe porcine disc was 76.4 MPa in the antero-posterior directionwhereas it was 3.2 MPa in the medio-lateral direction (strain rate500 mm/min; Beatty et al., 2001). Because in the anterior and pos-terior bands, the collagen fibers run mainly medio-laterally, theyhave a relatively large tensile modulus and strength in the mediolateral direction, although no data for the antero-posterior direction are available because of their short antero-posterior lengthInterestingly, the antero-posterior tensile moduli of the centra

and lateral regions of the intermediate zone differ (Tanne et al.1991; Tanaka et al., 2001a). This suggests differences in collagenfiber distribution between the two regions. In addition, in the

bovine disc, the tensile moduli in the lateral and medial regionare higher than those in the central region (Tanaka et al., 2001a)whereas in the canine disc (Tanne et al., 1991), the reverse is true(Table). This difference may be dependent on the difference inmasticatory patterns between cattle and dogs. As mentionedabove, the dog exhibits mainly chopping strokes during mastication, which may be primarily associated with loadings in the central region. Cattle exhibit lateral translatory movements, as in






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humans, which may be associated with a shift of the loading tothe lateral side of the intermediate zone (Beek et al., 2000).

(b) Compressive modulus

The resistance to compression is mainly dependent on the densi-

ty of proteoglycans, especially of the large chondroitin-sulfateproteoglycan molecules. Since the distribution of the proteogly-cans is different in various regions of the disc, regional differencesin its compressive modulus can be expected. In the anterior andposterior bands, and in the central region of the intermediatezone, the compressive modulus is slightly higher than in themedial and lateral regions of the intermediate zone (del Pozo etal., 2002; see Table). The possible explanation for this regional dif-ference is that the large chondroitin-sulfate proteoglycans and therelated chondroitin sulfate are preferentially localized in the cen-tral region of the intermediate zone and in the anterior and pos-terior bands (Mizoguchi et al., 1998). During compression, theseproteoglycans with large molecules may interfere with the flowof fluid out of the disc, resulting in an increase of the compressive

modulus. In the medial and lateral regions of the intermediatezone, the major GAG is dermatan sulfate, which is related to thesmall proteoglycans, decorin and biglycan (Nakano and Scott,1996). In these regions, the interference of fluid flow is assumedto be lower than in the chondroitin-sulfate-rich regions, resultingin smaller values of compressive modulus and strength.

The compressive modulus of the disc is generally con-sidered to be smaller than its tensile modulus. There are onlytwo studies available in which this has been demonstrated bymeans of the same experimental protocol and material (bovinedisc: Tanaka et al., 2001a; del Pozo et al., 2002). In these studies,

the tensile modulusranged betweenabout 22 and 26MPa, and the compressive modulus

between 14 and 17MPa. The possibleexplanation for thelarger tensile modu

lus is that the elasticity of the disc ismore dependent onthe collagen fibersthan on the proteoglycans. During tension, the stiffness isprimarily due to theresistance produced

by the collagenfibers. During compression, the collagen fibers are proba

bly slack, which

makes them lesseffective in resistingcompressive stress.

(c) Shear modulus

Investigation oshear properties insynovial joints is oparticular interest

because shear stress can result in fatigue, damage, and deformation of cartilage (Spirt et al., 1989; Zhu et al., 1993, 1994)Therefore, data on the shear modulus might contribute to a better understanding of secondary tissue damage.

It is very likely that shear stresses occur during loading o

the disc, because the articular surfaces that compress the discare not parallel. As a result, not all areas of the disc aredeformed in the same direction, leading to local shear. Anotherreason why shear stress occurs in the disc is its non-homogeneous structure. Its inner layer consists mainly of antero-poste-rior running collagen fibers and the "leaflet-like" proteoglycan(Kuc and Scott, 1994; Nakano and Scott, 1996), whereas thesuperior and inferior surface layers consist mainly of anteroposteriorly and medio-laterally running collagen fibers andsmall proteoglycans (Nakano and Scott, 1996; Minarelli et al.1997). Therefore, these layers are considered to have differen

biomechanical properties (Nakano and Scott, 1996; Mizoguchet al., 1998), which might lead to shear stress. This is supported

by the results of a finite element study, in which a relatively

large shear stress was predicted in a disc consisting of three lay-ers (Tanaka et al., 1994).

With respect to the shear modulus of the disc, thus far, onlyone study has been published, in which the shear modulus of theintermediate zone of the human disc (strain rate, 0.02 mm/secwas evaluated (Lai et al., 1998). It appeared that in the centraregion the shear modulus (about 1.0 MPa) was lower than in thelateral and medial regions (about 1.75 MPa). It has been reported that the shear stress in cartilage is very sensitive to the fre-quency and direction of the loading and to the amount of compressive strain (Mow et al., 1992). Therefore, the shear behavior


TABLESummary of Instantaneous Elastic Moduli (in MPa) of TMJ Disc

RegionSpecies Lat a Cent a Meda Ant a Post a Nsa

TensionTanne et al . (1991) Dog 39.5b 50.2b 43.3b

83.7 c 101.1c 91.9c

Tenget al

. (1991) Dog 18.4e 30.0e

30.1e

Tanaka et al . (2000) Human 44.0d (29.9)de

95.7 f (61.2)fe

Beatty et al . (2001) Pig 1.2-3.2g

27.3-76.4h

Tanaka et al . (2001a) Cow 21.7-24.0 20.2-22.9 24.0-25.9

CompressionFontenot (1985) Human 300 (1.8)e

Pig 347 (1.8)e

Chin et al . (1996) Human 0.2-0.5Tanaka et al . (1999) Dog 30.9del Pozo et al . (2002) Cow 14.6 (2.0)e 15.5 (1.5)e 14.7 (2.3)e 17.3 (1.1)e 15.5 (1.4)e

Shear

Lai et al . (1998) Human 1.8 1.0 1.7

a Lateral (Lat), central (Cent), and medial (Med) regions of intermediate zone, anterior (Ant) and posterior (Post)bands, region not specified (Ns).

b Stress range, 0-1.5 MPa.c Stress range, 1.5-4.0 MPa.

d Strain range, 0-2%.e Relaxed modulus.f Strain range, 2-4%.g Mediolateral loading.h Anteroposterior loading.





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of the disc should also be considered as a non-linear, anisotropic, and time-dependent behav-ior. Future studies need to examine the effectsof dynamic shear properties on the disc.

(B) VISCOELASTIC PROPERTIES

(a) Quasi-static behavior

(i) Stress relaxation

When the disc is compressed with a constantstrain, it shows a time-dependent stress relax-ation (Fig. 7). The produced stress decreasesmarkedly during the initial 5 to 30 sec, afterwhich it levels off to a steady low, but not zero,level (Scapino et al., 1996; Tanaka et al., 1999;del Pozo et al., 2002). The relaxation time of thedisc ranges from 3 to 45 sec, and the relaxedstress ranges between ca. 5 and 50% of the ini-tial stress (Fontenot, 1985; Tanaka et al., 1999;del Pozo et al., 2002). These large variations inrelaxation time and relaxed stress are due to adependency of the applied stress level; whenlarge amounts of stress are applied, the curve

shows a relatively long relaxation time and alarge relaxed stress. These findings indicatethat the movement of fluid out of the discunder loading is relatively slow and not pro-portional to the fluid pressure. Because of thisrelatively long relaxation time, the loaded disc

becomes relatively stiff when it is cyclicallyloaded—during, for example, chewing andspeaking (Beek et al., 2001a).

In tensile stress relaxation (Teng et al.,1991), the relaxation time (about 30 sec) isalmost the same as during compression. Therelaxed stress is approximately 70-80% of theinitial stress, which is considerably largerthan that in compression. The difference inthe relaxed stress between the compressiveand the tensile relaxation is due to the role of proteoglycans and collagen fibers. Duringcompression, the proteoglycans enmeshed inthe collagen matrix inflate the collagen net-work and lead to an increase in the osmoticpressure, resulting in the interference of fluidflow out of the disc. This resistance andresilience to compression both reduce withtime. In contrast, during tension, the collagen fibers arestretched, and the resistance to tension increases when thestrain is applied. Because of the flow of fluid out of the discduring compression, reduced relaxation occurs with time,whereas the resistance to tension due to the stretched collagenfibers may remain within the physiologic strain region. Thesefindings indicate that the proteoglycans are more important tocounteract compression and the collagen fibers are moreimportant to counteract tension.

After stress relaxation, a biomechanical equilibrium willeventually occur, which implies a balance between the appliedstress and the resistance to this stress in the disc. Before relax-ation, a high initial stress acting on the disc is distributedthrough the whole disc. More than 50% of the initial stress is

dissipated. This behavior implies that the disc functions as astress absorber and a stress distribution material. Without thedissipation of strain energy, storage of excessive strain energycan lead to breakage of the disc and other components of theTMJ (Fontenot, 1985; Teng and Xu, 1991; Nickel andMcLachlan, 1994; Tanaka et al., 1999; del Pozo et al., 2002).

(ii) Creep

The creep curves of the disc (Fig. 8) show a marked increase instrain during the initial few seconds, followed by a slowincrease, reaching an almost steady level after 3-10 min (compression, Kuboki et al., 1997; tension, Tanaka et al., 2001a). Creeptime is longer in compression (more than 10 min; Kuboki et al.1997) than in tension (3 min; Tanaka et al., 2001a). These findings


Figure 7. Experimental stress-relaxation curve ( = strain level) under compression (A)Experimental stress-relaxation plots obtained from discs (mean + 1 SD) with a theoreti-cal curve calculated from the linear regression model with the time constants (B). Datafrom Tanaka et al . (1999).

Figure 8. Experimental creep and restoration curve under tension (A). Experimental creep(B) and restoration (C) plots with theoretical curves. The graphs (B) and (C) are enlargedat the onset of the stress application and stress removal. Data from Tanaka et al . (2002a)





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indicate that fluid flow within and out of the disc is slower dur-ing compression than during tension. Creep time appears not to

be dependent on loading region (central, lateral, and medialregions of the intermediate zone; Tanaka et al., 2001a) or stresslevel (10 N, 20 N, and 30 N; Kuboki et al., 1997; 1.0 MPa and 1.5MPa; Tanaka et al., 2001a). During tension, the initial response to

the load is due to stretching of the collagen crimping within thetoe region, and the secondary response is due to the elongationof the collagen network and the squeezing of fluid against inter-nal stress. Fluid flow through the network is nonetheless possi-

ble, which leads to a gradual transfer of the load from the fluidto the collagen network. This finding also indicates that stressdistribution occurs in the disc. Changes in the shape of the discreduce the amount of stress concentration and probablydecrease the progression of injury (Tanaka et al., 2001a).

The restoration curves (Fig. 8) exhibit a marked decrease instrain during the first 5 sec, and the decrease in strain ceasesafter a few minutes. This feature indicates that some of theenergy used to deform the disc is not released immediatelyafter unloading occurs, and that the return of fluid from out-

side the disc and the recovery to its original shape are relative-ly slow (Beek et al., 2001a). The residual strain after creep could be an important factor for the capability for recovery and theprediction of permanent deformation in the disc.

The residual strain after 20 min of creep is less than 1%,which is almost similar among the central, lateral, and medialregions of the intermediate zone (Tanaka et al., 2002a). Thisimplies that the disc has a great capability of recovery after sus-tained stress. The creep time of 20 min is long when comparedwith in vivo loading conditions, such as sustained clenching.However, sustained stress with high magnitude is considered to

generate catabolic effects (Burger et al., 1992). Therefore, information on creep time may help to assess the possible effects opermanent changes of the disc as a result of prolonged stress.

(b) Dynamic behavior

The above-mentioned quasi-static experiments have providedvaluable information on how the behavior of the disc changesover time. With quasi-static experimental set-ups, howeveronly the linear viscoelastic behavior of the disc can be studiedThe disc should essentially be approached as a structure withnon-linear behavior, and thus its dynamic viscoelastic properties need to be examined, although the mechanisms responsi

ble for stress distribution, energy dissipation, and stressabsorption are the same as those for quasi-static loadingTherefore, dynamic experiments have recently been performed(compression, Beek et al., 2001a; tension, Tanaka et al., 2002b). Ingeneral, the dynamic properties are determined in cyclic testsat a physiologic strain range (Fig. 9).

When a sinusoidal oscillating strain is applied to the discwith a constant amplitude, a small phase shift is presen

between the strain and stress signals; the stress reaches its maximum earlier than the strain (Fig. 9; Beek et al., 2001a). Thephase shift between the strain and stress depends on the dif-

ference between the viscous and elastic properties of the discIn subsequent cycles, the value of the maximum stress decreases asymptotically, and the stress shows an almost steady leveafter 7 to 10 cycles. Therefore, the hysteresis loop during load-ing and unloading at a constant rate after 7 to 10 cycles willremain essentially unchanged. The hysteresis loops show thaenergy is dissipated inside the disc. As a result, dynamic cycletests enable us to comprehend the viscous and elastic properties of the disc, and its energy dissipation function.

The dynamic properties of the disc are dependent on the frequency and strain of loading. For example, in dynamic compression tests on the human disc, the indentation amplitude (strain0.25, 0.30, and 0.35) and frequency (0.02, 0.05, and 0.1 Hz) had aproportional effect on the value of the maximal stress and the

amount of energy dissipation (Beek et al., 2001a). Furthermore, themaximal stress and the energy dissipation are significantly largerin the intermediate zone than in the anterior and posterior bands(Beek et al., 2001a). In contrast, the anterior band of the bovine discshows a greater compressive modulus than the intermediate zonein a quasi-static test (del Pozo et al., 2002). Cyclic testing of articular cartilage specimens at a constant strain amplitude and at mod-erate strain levels showed repeatable stress-strain relationshipswithin 7 to 10 cycles of pre-conditioning (Fung, 1972). Furthermore, it is known that the curves showing the experimental peakand valley stresses in cyclic testing match well with those obtained

by the theoretical load-relaxation curves according to thequasi-linear viscoelastic theory (Woo et al., 1988). Therefore, the dynamicproperties are likely to exhibit smaller values than the static ones

and present some resemblance to the static modulus after load-relaxation rather than the elastic modulus.Under dynamic tension, a storage modulus (E) of about 0.7

to 1.4 MPa and a loss modulus (E) of about 0.1 to 0.2 MPa werefound (Tanaka et al., 2002b). These values are very small buclose to the relaxed moduli described in previous studies(Fontenot, 1985; Chin et al., 1996). This finding implies thatunder cyclic conditions, the disc exhibits a dynamic equilibriumwhich is similar to that after stress-relaxation. The relaxed modulus of the intermediate zone is larger than that of the anterior

band (del Pozo et al., 2002). This suggests that the intermediate


Figure 9. Measurement signals of a cyclic test with a constant fre-quency. Data from Beek et al . (2001a). (A) Stress vs. time. (B) Strainvs. time. (C) Stress vs. strain.





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zone shows a great capacity of energy dissipation, especiallyduring dynamic loading such as mastication and chewing.

When one compares the results of tests with a relativelylarge strain (30%; Beek et al., 2001a) with those of tests with asmall strain (0.5%; Tanaka et al., 2002b), all dynamic moduli

become larger when a larger strain is applied. Presumably, thisdifference is due to differences in fluid flow. The small perme-ability of the collagen network (pore size of 10-60; Mow et al.,1984, 1993) impedes the fluid flow through the collagen network.

Under small strains, the hydrostatic pressure in the interstitialfluid due to the hydrophilic character of the proteoglycans is in balance with the applied force (Tanaka et al., 2002b). Therefore,the load acting on the disc can be assumed to be carried by pres-surization of fluid without much deformation of the collagennetwork (Soltz and Ateshian, 1998). This mechanism protects thecollagen network against extreme local deformations duringloading. Under relatively high strain, fluid flow through the col-lagen network is nonetheless possible, which may lead to a grad-ual transfer of the load from the fluid to the collagen network(Beek et al., 2001a). When loaded, the collagen network deforms.This enables the disc to adapt its shape continuously to fit thespace between the opposing articular surfaces. However, thereturn of the fluid squeezed out of the disc under loading is rel-

atively slow, as mentioned above. As a result, the disc cannotkeep sufficient fluid in itself during cyclic loading, resulting in ahigher stiffness than that under small strain.

The values of dynamic viscoelastic E-moduli (E*, E, andE) also increase as the frequency increased from 0.1 to 100 Hz(Tanaka et al., 2002b). In dynamic tensile tests, the dynamic vis-coelastic E-moduli are about 2 times larger at 100 Hz than thoseat 1 Hz (Tanaka et al., 2002b). This non-linear dependency on thefrequency is due to fluid flow and squeezing within the matrixof the disc. At higher frequencies, the proteoglycans occupyingthe interfibrilar spaces interfere with smooth fluid flow, whichleads to strain energy dissipation, resulting in a higher stiffness.

(IV) Adaptive Changes of the Disc

The biomechanical behavior of the disc may change due to vari-ous factors during life. Although the detailed mechanism is notyet clear, mechanical stress affects the GAG synthesis in the disc,especially that of chondroitin sulfate, dermatan sulfate, andhyaluronic acid (Carvalho et al., 1995; Sindelar et al., 2000). Staticloading decreases the proteoglycan synthesis in cartilaginousstructures, whereas dynamic loading is positively related to thissynthesis and is considered as an important factor for mainte-nance of the homeostasis of the joint cartilage (Burger et al., 1992;Kuboki et al., 1997; Quinn et al., 1998). In contrast, there are signsthat the water content of cartilage from joints that are maximallyloaded is less than that of cartilage from regions that are notheavily loaded (Thonar et al., 1978). Reduction of the water con-tent due to high load is probably caused by a reduction in pro-

teoglycan and GAG synthesis that tends to normalize when theexternal pressure is eliminated (Schneiderman et al., 1986). Thedisc has a capacity to modify its GAG composition continuouslyin response to specified mechanical stresses (Nakano and Scott,1989a,b). These changes will probably affect its mechanical prop-erties. In addition, the collagen fibers are strengthened due to theinfluence of decorin, which may increase type I collagen fibrildiameters, resulting in discs that are better able to withstand ten-sile forces (Kuc and Scott, 1997). Hence, investigation of themechanical properties of the disc is also necessary for an under-standing of the consequences of functional remodeling.

(A) AGE

Age-related changes in viscoelasticity have been extensivelystudied in connective tissues, such as skin and tendon (Walker eal., 1976; Vogel, 1980; Woo, 1986). The tensile modulus of the ratskin increases with maturation and decreases with senescence(Vogel, 1980). In canine tendon, an increased stiffness andreduced viscoelasticity were found during aging (Walker et al.1976). Age-related changes have also been demonstrated in thedisc. The calcium content of the human disc increases progres

sively with aging (Takano et al., 1999). The increase in calcificationmay be caused by an intrinsic aging process or by an alteredmechanical stress (Jibiki et al., 1999). With respect to the GAGcomposition in the disc, the amounts of total and sulfated GAGsmarkedly increase from mature fetuses to mature adults (Nakanoand Scott, 1996). An increase in the content of sulfated GAGs relative to the tissue fluid will elevate the osmotic swelling pressureand, hence, the compressive stiffness of the disc. Accordingly, thematerial properties of the disc can also be expected to be relatedto age. Lai et al. (1998) were the first who demonstrated that theshear modulus of the human disc increases with age and suggested that the increase of this modulus may be the result of adecrease of collagen-remodeling capacity. The elastic modulus onormal human discs has a constant value of about 45 MPa up to

50 years of age, but increases thereafter to about 65 MPa (Tanakaet al., 2001b). The elastic moduli of bovine discs increase slightly

but significantly from those of the young adult (about 22 MPa) tothose of the mature adult (about 25 MPa; Tanaka et al., 2001b). Theratio of collagen to water in the disc increases with age, while thewater content remains constant (Nakano and Scott, 1996). Youngdiscs contain relatively more fluid and are capable of releasingmuch more fluid out of the disc, resulting in a smaller elasticityIn contrast, mature discs consist of relatively abundant collagenfibers, and may exhibit less water loss after relaxation.

(B) TRAUMA AND PATHOLOGY

The biomechanical properties of the disc also change due to trau

ma and pathology of the TMJ. In the case of pathology or trau-ma, the quantity of damage exceeds the normal repair capacityof the articular tissues. It has been reported that the articular cartilage of the knee joint shows a reduction in tensile strengthprior to evidence of surface damage (McCormack and Mansour1998). This suggests that, under pathologic conditions, the biomechanical properties of the articular cartilage change with molecular composition. Similar observations have been reported forthe disc. For example, human pathological discs with extensivedegenerative alterations, such as hyalinization and cell-freeareas, had significantly greater relaxed tensile moduli than nor-mal discs (Tanaka et al., 2000). Proteoglycans and GAGs in thedisc are also redistributed and newly synthesized after an orasplint is worn for two months (Sindelar et al., 2000).

Continuous sustained loading in the joint can also induceincrease of joint friction, because only solid contact may exist

between the articular surfaces after prolonged loading(Forster and Fisher, 1996, 1999). Continuous loading may leadto an increase of surface roughness of the articular cartilageand subsequently to surface wear (Forster and Fisher, 1999)The resulting increase of friction may induce increased shearstress in the disc. As mentioned above (Part 3), shear stresscan be associated with fatigue and damage of the disc and canlead to changes in GAG composition and thus to changes inmechanical properties of the disc.






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(V) Concluding RemarksThe TMJ disc behaves as a viscoelastic structure. Through itsviscoelastic properties, the disc can function as a stressabsorber and stress distributor. Thereby, the disc contributes toprevent stress concentration and excessive stress in the carti-lage and bone components of the joint. It is likely that thesefunctions protect the joint from degeneration of the disc (perfo-ration and thinning) and osteoarthrosis (condylar deformationwith erosive and/or proliferative changes).

The finite element method has been proven to be a suitabletool for approximating the distribution of loads in the structuresof the TMJ. Since 1990, several three-dimensional finite elementmodels of the TMJ, including the disc, have been developed(Korioth et al., 1992; Chen and Xu, 1994; Tanaka et al., 1994; Chinet al., 1996; DeVocht et al., 1996; Nagahara et al., 1999; Beek et al.,2000). These models have shown, for example, that the predict-ed stress in the joint components and the sizes of the contactareas depend on the elastic modulus of the disc (DeVocht et al.,1996; Beek et al., 2000; Tanaka et al., 2001b). However, thus far, thematerial properties of the disc have been assumed to be linearlyelastic in these models. It will be a challenge to develop futuremodels into which the viscoelastic, anisotropic, and heteroge-neous properties of the disc are incorporated. A first approxima-

tion could be the application of an adequate material model of the disc that includes both fluid and solid constituents. To assessthe interaction of the solid matrix and the interstitial fluid, inves-tigators have developed a biphasic theory (Mow et al., 1980),which shows that a large part of the load acting on cartilaginousstructures is carried by interstitial fluid pressurization (Soltz andAteshian, 1998). In contrast to the presently available models of the TMJ, such a biphasic or poro-elastic model would alsoaccount for the shock-absorbing properties of the disc.

Wear of the disc has been found primarily in its anterior band (Kopp, 1976) or intermediate region (Werner et al., 1991).The former is supported by the biomechanical and biochemicalfindings that the anterior band receives the largest loads(Tanaka et al., 1994), resulting in the concentration of chon-

droitin sulfate in this region (Nakano and Scott, 1989a;Mizoguchi et al., 1998). The latter is also in agreement with var-ious studies predicting that the disc is predominantly loaded inits intermediate zone (DeVocht et al., 1996; Nagahara et al., 1999;Beek et al., 2000). Although the stress distribution in the disc isstrongly affected by the direction of loading, the two regionsprobably play a role in stress absorption and stress distributionduring clenching and mastication.

Translation of the condyle in the forward direction, to obtaina protrusive or open-jaw configuration, leads to a concentrationof loading in the lateral part of the disc (Beek et al., 2001a; Tanakaet al., 2001c). This would suggest that, during open-close move-ments, the lateral part of the intermediate zone is primarily sub-

jected to wear and friction. This is supported by Werner et al.(1991), who reported that wear leading to perforations of the discwas mainly located in the lateral part of its intermediate zone.Furthermore, Gallo et al. (2000) suggest that, during mastication,fatigue failure of the TMJ disc could result from shear stress pro-duced by medio-lateral translation of stress location.

More information about the biomechanical properties of thedisc is also required for the design of TMJ implants. Some of theavailable implant materials failed under functional conditions(Silastic—Dolwick and Aufdemorte, 1985; Eriksson andWestesson, 1986; Westesson et al., 1987; Sanders et al., 1990; Tucker

and Watzke, 1991; Proplast—Heffez et al., 1987; Florine et al., 1988Valentine et al., 1989; Wagner and Mosby, 1990). The major reasonfor these failures is that these biomaterials were not strongenough to withstand the functional loading applied to them.

REFERENCESBeatty MW, Bruno MJ, Iwasaki LI, Nickel JC (2001). Strain rate depen

dent orthotropic properties of pristine and impulsively loadedporcine temporomandibular joint disk. J Biomed Mater Res57:25-34

Beek M, Koolstra JH, van Ruijven LJ, van Eijden TMGJ (2000)Three-dimensional finite element analysis of the human temporomandibular joint disc. J Biomech 33:307-316.

Beek M, Aarnts MP, Koolstra JH, Feilzer AJ, van Eijden TMGJ(2001a). Dynamic properties of the human temporomandibular

joint disc. J Dent Res 80:876-880.Beek M, Koolstra JH, van Ruijven LJ, van Eijden TMGJ (2001b). Three

dimensional finite element analysis of the cartilaginous structuresin the human temporomandibular joint. J Dent Res 80:1913-1918.

Berg R (1978). Anatomía topográfica y aplicada de los animalsdomésticos. Madrid: AC Libros Científicos y Técnicos.

Berkovitz BKB (2000). Collagen crimping in the intra-articular discand articular surfaces of the human temporomandibular joint

Arch Oral Biol 45:749-756.Bermejo FA, Puchades-Orts A, Sánchez del Campo F, Panchón

Ruiz A, Herrera-Lara M (1987). Morphology of the meniscotemporal part of the temporomandibular joint and its biomechanical implications. Acta Anat 129:220-226.

Bermejo A, González O, González JM (1993). The pig as an animalmodel for experimentation on the temporomandibular articulacomplex. Oral Surg Oral Med Oral Pathol 75:18-23.

Boyd RL, Gibbs CH, Mahan PE, Richmond AF, Laskin JL (1990)Temporomandibular joint forces measured at the condyle o

Macaca arctoides. Am J Orthod Dentofac Orthop 97:472-479.Brehnan K, Boyd RL, Laskin J, Gibbs CH, Mahan P (1981). Direc

measurement of loads at the temporomandibular joint in Macaca arctoides. J Dent Res 60:1820-1824.

Burger EH, Klein-Nulend J, Veldhuijzen JP (1992). Mechanicastress and osteogenesis in vitro. J Bone Miner Res 7:S397-S401.

Carvalho RS, Yen EH, Suga DM (1995). Glycosaminoglycan synthesis in the rat articular disk in response to mechanical stress

Am J Orthod Dentofac Orthop 107:401-410.Chen J, Xu L (1994). A finite element analysis of the human tem-

poromandibular joint. J Biomech Eng 116:401-407.Chin LPY, Aker FD, Zarrinnia K (1996). The viscoelastic properties

of the human temporomandibular joint disc. J Oral MaxillofaSurg 54:315-318.

Chopra RK, Pearson CH, Pringle GA, Fackre DS, Scott PG (1985)Dermatan sulphate is located on serine-4 of bovine skin proteo-dermatan sulphate. Biochem J 231:277-279.

del Pozo R, Tanaka E, Tanaka M, Okazaki M, Tanne K (2002). Theregional difference of viscoelastic property of bovine temporomandibular joint disc in compressive stress-relaxation. Med EngPhys 24:165-171.

DeVocht JW, Goel VK, Zeitler DL, Lew D (1996). A study of the con-trol of disc movement within the temporomandibular joint usingthe finite element technique. J Oral Maxillofac Surg 54:1431-1437

Dolwick MF, Aufdemorte TB (1985). Silicone-induced foreign bodyreaction and lymphadenopathy after temporomandibular joinarthroplasty. Oral Surg Oral Med Oral Pathol 59:449-452.

Eriksson L, Westesson PL (1986). Deterioration of temporary siliconeimplant in the temporomandibular joint: a clinical and arthro-scopic follow-up study. Oral Surg Oral Med Oral Pathol 62:2-6.

Ferrario VF, Sforza C (1994). Biomechanical model of the humanmandible in unilateral clench: distribution of temporo






13/14

mandibular joint reaction forces between working and balan-cing sides. J Prosthet Dent 72:169-172.

Fisher LW, Termine JD, Young MF (1989). Deduced proteinsequence of bone small proteoglycan I (biglycan) shows homol-ogy with proteoglycan II (decorin) and several nonconnectivetissue proteins in a variety of species. J Biol Chem 264:4571-4576.

Florine BL, Gatto DJ, Wade ML, Waite DE (1988). Tomographicevaluation of temporomandibular joints following discoplastyor placement of polytetrafluoroethylene implants. J Oral

Maxillofac Surg 46:183-188.

Fontenot MG (1985). Viscoelastic properties of human TMJ discsand disc replacement materials (abstract). J Dent Res 64:163.

Forster H, Fisher J (1996). The influence of loading time and lubri-cant on the friction of the articular cartilage. Proc Inst Mech Eng[H] 210:109-119.

Forster H, Fisher J (1999). The influence of continuous sliding andsubsequent surface wear on the friction of the articular carti-lage. Proc Inst Mech Eng [H] 213:329-345.

Fung YC (1969). Viscoelasticity. In: A first course in continuummechanics. Englewood Cliffs, NJ: Prentice-Hall, pp. 174-179.

Fung YCB (1972). Stress-strain-history relations of soft tissues insimple elongation. In: Biomechanics: its foundations and objec-tives. Fung YC, Perrone N, Anliker M, editors. EnglewoodCliffs, NJ: Prentice-Hall, pp. 181-208.

Fung YC (1981). Biomechanical properties of living tissue. NY:Springer-Verlag, pp. 196-260.Gallo LM, Nickel JC, Iwasaki LR, Palla S (2000). Stress-field trans-

lation in the healthy human temporomandibular joint. J DentRes 79:1740-1746.

Gathercole LJ, Keller A (1991). Crimp morphology in fibre-formingcollagens. Matrix 11:214-234.

González SO, Bermejo FA, González GMJ (1991). Estudio de laarticulación temporo-madibular en diferentes animalsmamíferos. Un nuevo concepto. Av Odontoestomatol 7:445-451.

Hatcher DC, Faulkner MG, Hay A (1986). Development of mechan-ical and mathematic models to study temporomandibular jointloading. J Prosthet Dent 55:377-384.

Hatton MN, Swann DA (1986). Studies on bovine temporo-mandibular joint synovial fluid. J Prosthet Dent 56:635-638.

Heffez L, Mafee MF, Rosenberg H, Langer B (1987). CT evaluationof TMJ disc replacement with a Proplast-Teflon laminate. J Oral

Maxillofac Surg 45:657-665.Hylander WL, Bays R (1978). Bone strain in the subcondylar region

of the mandible in Macaca fascicularis and Macaca mulatta. Am J Phys Anthropol 48:408.

Hylander WL, Bays R (1979). An in vitro strain-gauge analysis of squamosal-dentary joint reaction force during mastication andincision in Macaca mulatta and Macaca fascicularis. Arch Oral Biol24:689-697.

Iozzo RV (1998). Matrix proteoglycans: from molecular design tocellular function. Annu Rev Biochem 67:609-652.

Jibiki M, Shimoda S, Nakagawa Y, Kawasaki K, Asada K, IshibashiK (1999). Calcifications of the disc of the temporomandibular

joint. J Oral Pathol Med 28:413-419.Koolstra JH, van Eijden TMGJ, Weijs WA, Naeije M (1988). A three-

dimensional mathematical model of the human masticatorysystem predicting maximum possible bite forces. J Biomech21:563-576.

Kopp S (1976). Topographical distribution of sulphated glycosamino-glycans in human temporomandibular joint disks. A histochemi-cal study of an autopsy material. J Oral Pathol 5:265-276.

Korioth T, Romilly D, Hannam A (1992). Three-dimensional finiteelement stress analysis of the dentate human mandible. Am J Phys Anthropol 88:69-96.

Kuboki T, Shinoda M, Orsini MG, Yamashita A (1997). Viscoelastic

properties of the pig temporomandibular joint articular soft tis-sues of the condyle and disc. J Dent Res 76:1760-1769.

Kuc IM, Scott PG (1994). Ultrastructure of the bovine temporomandibular joint disc. Arch Oral Biol 39:57-61.

Kuc IM, Scott PG (1997). Increased diameters of collagen fibrils precipitated in vitro in the presence of decorin from various connective tissues. Connect Tissue Res 36:287-296.

Lai W-FT, Bowley J, Burch JG (1998). Evaluation of shear stress of thehuman temporomandibular joint disc. J Orofac Pain 12:153-159.

Langenbach G, Van Eijden TMGJ (2001). Mammalian feeding

motor patterns. Am Zool 41:1338-1351.Li JT, Armstrong CG, Mow VC (1983). The effect of strain rate on

mechanical properties of articular cartilage in tension. InBiomechanics symposium. Woo S-Y, Mates R, editors. NewYork: American Society of Mechanical Engineers, pp. 117-120.

McCormack T, Mansour JM (1998). Reduction in tensile strength ocartilage precedes surface damage under repeated compressiveloading in vitro. J Biomech 31:55-61.

Mills DK, Daniel JC, Scapino RP (1988). Histological features and invitro proteoglycan synthesis in the rabbit craniomandibular

joint disc. Arch Oral Biol 33:195-202.Mills DK, Fiandaca DJ, Scapino RP (1994). Morphologic, micro

scopic and immunohistochemical investigations into the function of the primate TMJ disc. J Orofac Pain 8:136-154.

Minarelli AM, Santo MD Jr, Liberti EA (1997). The structure of thehuman temporomandibular joint disc: a scanning electronmicroscopy study. J Orofac Pain 11:95-100.

Mizoguchi I, Scott PG, Dodd CM, Rahemtulla F, Sasano YKuwabara M, et al. (1998). An immunohistochemical study othe localization of biglycan, decorin and large chondroitin-sulphate proteoglycan in adult rat temporomandibular joint disc

Arch Oral Biol 43:889-898.Mow VC, Kuei SC, Lai WM, Armstrong CG (1980). Biphasic creep

and stress relaxation of articular cartilage in compression: theory and experiments. J Biomech Eng 102:73-84.

Mow VC, Holmes MH, Lai WM (1984). Fluid transport andmechanical properties of articular cartilage: a review. J Biomech17:377-394.

Mow VC, Kwan MK, Lai WM, Holmes MH (1986). A finite defor

mation theory for nonlinearly permeable soft hydrated biological tissues. In: Frontiers in biomechanics. Schmid-SchonbeinGW, Woo SL-Y, Zweifach BW, editors. New York: Springer-Verlag, Chapter 13, pp. 153-179.

Mow VC, Ratcliffe A, Chern KY, Kelly MA (1992). Structure andfunction relationships of the menisci of the knee. In: Kneemeniscus: basic and clinical foundations. Mow VC, ArnoczkySP, Jackson DW, editors. New York: Raven Press, pp. 37-57.

Mow VC, Ateshian GA, Spilker RL (1993). Biomechanics odiarthrodial joints: a review of twenty years of progress. Biomech Eng 115:460-467.

Muir H (1973). The chemistry of the ground substance of joint car-tilage. In: Joints and synovial fluid. Vol. 18. Sokoloff L, editorNew York: Academic Press, pp. 1011-1020.

Muir H (1981). Proteoglycans: state of the art. Arthrit Rheum 24:7-10Murata H, Taguchi N, Hamada T, McCabe JF (2000). Dynamic vis-

coelastic properties and the age changes of long-term soft den-ture liners. Biomaterials 21:1421-1427.

Nagahara K, Murata S, Nakamura S, Tsuchiya T (1999)Displacement and stress distribution in the temporomandibular joint during clenching. Angle Orthod 69:372-379.

Nakano T, Scott PG (1989a). A quantitative chemical study of glycosaminoglycans in the articular disc of the bovine temporomandibular joint. Arch Oral Biol 34:749-757.

Nakano T, Scott PG (1989b). Proteoglycans of the articular disc othe bovine temporomandibular joint. I. High molecular weigh






14/14

chondroitin sulphate proteoglycan. Matrix 9:277-283.Nakano T, Scott PG (1996). Changes in the chemical composition of

the bovine temporomandibular joint disc with age. Arch OralBiol 41:845-853.

Nakano T, Imai S, Koga T, Dodd CM, Scott PG (1993). Monoclonalantibodies to the large chondroitin sulphate proteoglycan from

bovine temporomandibular joint disc. Matrix 13:243-254.Nickel JC, McLachlan KR (1994). In vitro measurement of the stress-

distribution properties of the pig temporomandibular joint disc. Arch Oral Biol 39:439-448.

Quinn TM, Grodzinsky AJ, Buschmann MD, Kim Y-J, Hunziker EB(1998). Mechanical compression alters proteoglycan depositionand matrix deformation around individual cells in cartilageexplants. J Cell Sci 111:573-583.

Robbins JR, Vogel KG (1994). Regional expression of mRNA forproteoglycans and collagen in tendon. Eur J Cell Biol 64:264-270.

Sanders B, Buoncristiani RD, Johnson L (1990). Silicone rubberfossa implant removal via partial arthrotomy followed byarthroscopic examination of the internal surface of the fibrouscapsule. Oral Surg Oral Med Oral Pathol 70:369-371.

Scapino RP, Canham PB, Finlay HM, Mills DK (1996). The behav-iour of collagen fibres in stress relaxation and stress distributionin the jaw-joint of rabbits. Arch Oral Biol 41:1039-1052.

Schneiderman R, Keret D, Maroudas A (1986). Effects of mechani-

cal and osmotic pressure on the rate of glycosaminoglycan syn-thesis in the human adult femoral head cartilage: an in vitrostudy. J Orthop Res 4:393-408.

Scott PG, Nakano T, Dodd CM (1995). Small proteoglycans fromdifferent regions of the fibrocartilaginous temporomandibular

joint disc. Biochim Biophys Acta 1244:121-128.Sindelar BJ, Evanko SP, Alonzo T, Herring SW, Wight T (2000).

Effects of intraoral splint wear on proteoglycans in the tem-poromandibular joint disc. Arch Biochem Biophys 379:64-70.

Smith DM, McLachlan KR, McCall WD (1986). A numerical modelof temporomandibular joint loading. J Dent Res 65:1046-1052.

Soltz MA, Ateshian GA (1998). Experimental verification and theo-retical prediction of cartilage interstitial fluid pressurization atan impermeable contact interface in confined compression. J Biomech 31:927-934.

Spirt AA, Mak AFD, Wassell RP (1989). Nonlinear viscoelasticproperties of articular cartilage in shear. J Orthop Res 7:43-49.

Stegenga B, DeBont LGM, Boering G, Van Willigen JD (1991).Tissue responses to degenerative changes in the temporo-mandibular joint: a review. J Oral Maxillofac Surg 49:1079-1088.

Takano Y, Moriwake Y, Tohno Y, Minami T, Tohno S, Utsumi M, etal. (1999). Age-related changes of elements in the human articu-lar disk of the temporomandibular joint. Biol Trace Elements Res67:269-276.

Tanaka E, Tanne K, Sakuda M (1994). A three-dimensional finiteelement model of the mandible including the TMJ and its appli-cation to stress analysis in the TMJ during clenching. Med EngPhys 16:316-322.

Tanaka E, Tanaka M, Miyawaki Y, Tanne K (1999). Viscoelasticproperties of canine temporomandibular joint disc in compres-sive load-relaxation. Arch Oral Biol 44:1021-1026.

Tanaka E, Shibaguchi T, Tanaka M, Tanne K (2000). Viscoelastic prop-erties of the human temporomandibular joint disc in patientswith internal derangement. J Oral Maxillofac Surg 58:997-1002.

Tanaka E, Tanaka M, Hattori Y, Aoyama J, Watanabe M, Sasaki A,et al. (2001a). Biomechaical behaviour of bovine temporo-mandibular articular discs with age. Arch Oral Biol 46:997-1003.

Tanaka E, Sasaki A, Tahmina K, Yamaguchi K, Mori Y, Tanne K(2001b). Mechanical properties of human articular disk and its

influence on TMJ loading studied with the finite elemenmethod. J Oral Rehabil 28:273-279.

Tanaka E, Rodrigo DP, Tanaka M, Kawaguchi A, Shibazaki T, TanneK (2001c). Stress analysis in the TMJ during jaw opening by useof a three-dimensional finite element model based on magneticresonance images. Int J Oral Maxillofac Surg 30:421-430.

Tanaka E, Tanaka M, Aoyama J, Watanabe M, Hattori Y, Asai D, eal. (2002a). Viscoelastic properties and residual strain in a tensile creep test on bovine temporomandibular articular discs

Arch Oral Biol 47:139-146.

Tanaka E, Aoyama J, Tanaka M, Murata H, Hamada T, Tanne K(2002b). Dynamic properties of bovine temporomandibular

joint disks with age. J Dent Res 81:618-622.Tanne K, Tanaka E, Sakuda M (1991). The elastic modulus of the

temporomandibular joint disc from adult dogs. J Dent Re70:1545-1548.

Teng SY, Xu YH (1991). Biomechanical properties and collagen fibeorientation of TMJ discs in dogs: part 1. Gross anatomy and collagen fiber orientation of the discs. J Craniomandib Disord FaciaOral Pain 5:28-34.

Teng SY, Xu YH, Cheng MH, Li Y (1991). Biomechanical propertieand collagen fiber orientation of TMJ discs in dogs: part 2Tensile mechanical properties of the discs. J Craniomandib DisordFacial Oral Pain 5:107-114.

Thonar EJMA, Sweet MBE, Immelman AR, Lyons G (1978)Hyaluronate in articular cartilage: age related changes. CalcifTissue Res 26:19-21.

Throckmorton GS, Dechow PC (1994) In vitro strain measurementin the condylar process of the human mandible. Arch Oral Bio39:853-867.

Tucker MR, Watzke IM (1991). Autogenous auricular cartilage graffor temporomandibular joint repair. A comparison of techniquewith and without temporary silastic implantation. Craniomaxillofac Surg 19:108-112.

Valentine JD Jr, Reiman BE, Beuttenmuller EA, Donovan MG(1989). Light and electron microscopic evaluation of Proplast ITMJ disc implants. J Oral Maxillofac Surg 47:689-696.

Vogel HG (1980). Influence of maturation and aging on mechanicaand biochemical properties of connective tissue in rats. Mech

Ageing Dev 14:283-292.Wagner JD, Mosby EL (1990). Assessment of Proplast-Teflon disc

replacements. J Oral Maxillofac Surg 48:1140-1144.Walker P, Amstutz HC, Rubinfeld M (1976). Canine tendon studies

II. Biomechanical evaluation of normal and regrown canine tendons. J Biomed Mater Res 10:61-76.

Werner JA, Tillmann B, Schleicher A (1991). Functional anatomy othe temporomandibular joint. Anat Embryol 183:89-95.

Westesson PL, Eriksson L, Lindstrom C (1987). Destructive lesionsof the mandibular condyle following diskectomy with temporary silicone implant. Oral Surg Oral Med Oral Pathol 63:143-150

Woo SL-Y (1986). Biomechanics of tendons and ligaments. InFrontiers in biomechanics. Schmid-Schonbein GW, Woo SL-YZweifach BW, editors. New York: Springer-Verlag, Chapter 14pp. 180-195.

Woo SL-Y, Mow VC, Lai WM (1988). Biomechanical properties oarticular cartilage. In: Handbook of bioengineering, Skalak RChen S, editors. New York: McGraw-Hill Book Co., Chapter 4pp. 1-44.

Zhu W, Mow VC, Koob TJ, Eyre DR (1993). Viscoelastic shear properties of articular cartilage and the effects of glycosidase treatments. J Orthop Res 11:771-781.

Zhu W, Chern KY, Mow VC (1994). Anisotropic viscoelastic shearproperties of bovine meniscus. Clin Orthop 306:34-45.

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