67

REVIEW Cartilage and diarthrodial joints as paradigms for hierarchical materials and structures Van C. Mow and Anthony Ratcliffe Orthopaedic

Research

Laboratory,

Columbia

University,

New York, NY 10032, USA

A. Robin Poole Department of Surgery, McGill University Children, Montreal, Canada H3G IA6

and Joint Disease Laboratory,

Shriners

Hospital

for Crippled

The anatomic forms of diarthrodial joints are important structural features which provide and limit the motions required for the joint. Typically, the length scale of topographic variation of anatomic forms ranges from 0.5 to 15 cm. Articular cartilage is the thin layer of hydrated soft tissue (0.550 mm thick) covering the articulating bony ends in diarthrodial joints. This tissue has a set of unique mechanical and physicochemical properties which are responsible for its load-carrying capabilities and near-frictionless qualities. The mechanical properties of articular cartilage are determined at the tissue-scale level and these properties depend on the composition of the tissue, mainly collagen and proteoglycan, and their molecular and ultrastructural organization (ultra-scale: 1O-s-1O-6 m). Because proteoglycans possess a high density of fixed negative charges, articular cartilage exhibits a significant Donnan osmotic pressure effect. This physicochemically derived osmotic pressure is an important component of the total swelling pressure; the other component of the total swelling pressure stems from the charge-to-charge repulsive force exerted by the closely spaced (l-l.5 nm) negative charge groups along the proteoglycan molecules. Thus these interactions take place at a nano-scale level: 10-io-lO-g m. Finally, cartilage biochemistry and organization are maintained by the chondrocytes which exist at a micro-scale level (10-7-10-6 m). Significant mechanoelectrochemical transduction occurs within the extracellular matrix at the micro-scale level which affects and modulates cellular anabolic and catabolic activities. At present, the exact details of these transduction mechanisms are unknown. In this review, we present a summary of the hierarchical features for articular cartilage and diarthrodial joints and tables of known material properties for cartilage. Also we summarize how the multi-scale interactions in articular cartilage provide for its unique material properties and tribological characteristics. Keywords:

Review,

Accepted 25 July

cartilage,

bone,

Evolution in mammals, reptiles, birds and fish has developed extraordinarily efficient musculoskeletal systems for generating and controlling motion. However, the musculoskeletal system is not only an efficient system for delivering useful mechanical energy and load support, but is also capable of synthesizing, processing and organizing complex macromolecules to fashion tissues and organs for specific mechanical functions. An important subset of organs of the musculoskeletal system are the joints. Many types of joints exist in the body. Freely moving joints [ankle, elbow, hip, knee, shoulder, and those of the fingers and wrists) are known as diarthrodial or synovial joints. The intervertebral Correspondence to Dr V.C. Mow, Orthopaedic Research Laboratory, Columbia Presbyterian Medical Center, BB1412, Street, New York, NY 10032, USA.

0 1992

Butterworth-Heinemann

0142-9612/92/020067-31

Ltd

joints,

tribology

1991

630 West 166th

joints of the spine are not diarthrodial joints as they are fibrous and do not move freely. They do, however, provide the flexibility required by the spine. Diarthrodial joints enable locomotion and other activities of daily life to take place. They perform their functions so well that we are often not even aware of their existence nor the functions they provide until injury strikes or arthritis develops. From an engineering point of view, these natural bearings are very uncommon structures. Under healthy conditions, they function in a nearly frictionless and almost entirely wear-resistant manner throughout our lives. Failure of the bearing surfaces of the body [i.e. articular cartilage), as with engineering bearings, means a failure of these bearings to provide their essential functions. In biomedical terms, failure of diarthrodial joints leads to arthritis. Arthritis which develops primarily from mechanical causes is .___.~__ Biomaterials

1992, Vol. 13 No. 2

Cartilage

and diarthrodial

joints:

TRIBOLOGY

known as osteoarthrosis or osteoarthritis, or simply OA. While this is an oversimplified depiction of the arthritic process, the analogy between an engineering bearing and a diarthrodial joint is apt, and it conveys the essential biomechanical processes involved in joint failure. Repair of arthritic joints means orthopaedic surgery to replace the worn-out joint by a prosthesis or by a biological graft. This is an enormous medical and economic problem 32-38 million Americans suffer from this crippling disease. Major opportunities towards developing synthetic material replacements for cartilage may be derived from knowledge of diarthrodial joint function. A brief description is presented below, outlining the compositional and hierarchical features of diarthrodial joints and articular cartilage; these are responsible for the efficient functioning of animal joints. Figure 1 shows some of the hierarchical features found in the organization of a typical joint, and Table 2 summarizes the dimensional scales found in this hierarchy. At each hierarchical level, the structure has a significant influence on articular cartilage properties and joint function. Over the past two decades, significant progress has been made toward understanding the physicochemical and deformation behaviour of cartilage at each of these levels. Indeed, much can be learned from knowing how the connective tissues of our bodies are made and organized, how they function and how they behave under loading. Engineers would benefit from an appreciation of nature’s design of the extraordinarily efficient and long-lasting animal bearings: conversely, biomedical researchers on arthritis can learn much from engineers with detailed knowledge of the deformational behaviour of nature’s bearing materials. Also, knowledge of friction, lubrication and wear of bearings (tribology: tribo - to rub; science of rubbing) will assist in understanding joint function.

Bearings, linear or rotational, are essential for the function of practically every kind of man-made mechanical machinery. Since the classic work of Osborne Reynolds (1888) on viscous fluid-film lubrication, the science of tribology has been applied to every conceivable bearing system made by man - from the heavily loaded, slow-moving caterpillar space shuttle transports to the bearings of the ordinary automobile, to the lightly loaded, long-life gas bearings of gyroscopes, commonly used in inertial guidance systems. Various theories have been proposed for the study of lubrication of these bearings. The most fundamental ones are the classical hydrodynamic lubrication theory of Reynolds’, the elastohydrodynamic lubrication theory of Dowson and Higginson (1977)' and the boundary lubrication theories [see Bowden and Tabor, 19843;Peterson and Winer, 19804). In this survey, we present a brief accounting of tribology to facilitate the understanding of biotribology phenomena. In so doing the functional roles of the hierarchical organization of articular cartilage and diarthrodial joints will be developed.

Modes of lubrication In hydrodynamic and elastohydrodynamic thin fluid-film separates the two bearing Table 1 Multiscaled structural cartilage and diarthrodial joints. Nano-scale Ultra-scale Micro-scale Tissue-scale Macro-scale

Tissue (10%10’*m)

67 “In Co0aga-n

Pmteogtycans

I

Triple Helix

Chondrqcyte

Atticular Cartilage

[+-1200 “In-+( Proteoglycans Nano (10”~-10

Figure 1 Biomaterials

*rn)

Some of the important 1992,

Vol. 13 No. 2

structural

of

articular

(1O-1o-1O-g m): Na+, Ca2+, SO,-, COO(10-*-10-6 m): collagen and proteoglycan (10-7-10-4 m): cells and extracellular matrix (1O-4-1O-2 m): articular cartilage and bone (0.5-15 cm): diarthrodial joints

Cancellous and Marrow

Macro (OScm-15cm)

lubrication, a surfaces which

organization

Joint Cavi

bllagen

V.C. Mow et al.

Ultra (lOdm-10’m)

features found in a typical joint and the relevant length scales.

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and diarthrodial

joints:

69

V.C. Mow et al.

are sliding relative to each other. This fluid-film is usually less than 15 pm thick, and it is created by the motion of one of the solid surfaces relative to the other as the relative motion drags the viscous lubricant into the converging load-bearing contact region. In hydrodynamic lubrication, the major elements of the theory include: (1) the viscosity of the linear newtonian lubricant, (2) the geometry of the rigid and smooth solid bearing surfaces, (3) the functional load support requirements, and (4) the speed of operation of the bearing. In elastohydrodynamic lubrication, this theory is extended to include the elastic deformation of the bearing surfaces (with every other parameter remaining the same). In the elastohydrodynamic theory, the bearing surfaces are usually considered to be isotropic and linearly elastic solids (Hookean). A generalization of this theory is the elastorheodynamic theory where the lubricant is considered to be viscoelastic or non-newtonian, which may be linear or non-linear. These modes of lubrication require relatively high sliding speeds to create a lubricant film with significant load-carrying capacity (see, for example, Ref. 5). Bearings operating under fluid-film lubrication conditions are ideal, since no surface-to-surface contact occurs, and thus, theoretically, adhesive surface wear is not possible. However, because of the high pressures within the load-bearing region, fatigue failure often occurs deep within the bearing materia14. Historically, boundary lubrication is a term used to describe the mode of lubrication when a fluid film does not exist, e.g. at very high loads or low operating speeds for the lubricant or the bearing material used3’ 4. More recently, boundary lubrication has been used to refer to situations in which molecular rather than bulk properties of the lubricant play significant roles in lubrication, For example, adsorption of a monolayer of polymeric species in the lubricant, charged or uncharged, may play a significant role in protecting the bearing surfaces from adhesive surface wear. This is also a relevant concept in biotribology. The smooth functioning of bearings is an enormously important and costly aspect of our industrialized civilization. All these considerations have their biological and biomedical counterparts.

Modes of wear Wear, in general, is defined as an unwanted removal of material from an object. It is an extremely complex phenomenon and, to date, no satisfactory theory exists to explain even the simplest wear processes4. It is, at best, an empirical science, based upon years of practical engineering experience. For bearings, mathematical solutions derived from the two theories of fluid-film lubrication*% ‘*‘, and those for contact mechanics problems between two solid bodies (elastic or plastic)6 are used as guides for understanding the wear processes. For any real engineering bearing problem, one must first appreciate that there are many forms of wear in bearings3*4. There is two-body adhesion (metal-metal), corrosion (oxide film formation), cavitation (liquid drops impinging on a solid surface - rain erosion), three-body wear (debris particles trapped between bearing surfaces), ploughing (a hard surface moving over a softer surface) and delamination (debonding of laminar structures). Some of these processes are also relevant to wear of

biological materials. Wear of bearing surfaces designed to function under fluid-film lubrication conditions has a severe detrimental effect on lubrication function, since the hydrodynamic flow process in the very thin fluid-film between the two bearing surfaces is significantly altered by even the most minute scratch on the surfaces. Thus the smooth surface topography must be preserved for any bearing which is designed to function under fluidfilm conditions. From the viewpoint of mechanics of materials, adhesive wear that occurs between sliding metal-metal surfaces results from welding of the contact points between micro-asperities on the two moving surfaces3,4. Wear occurs when the micro-asperities of the softer metal fracture and are transferred to the harder metal. Eventually, these torn particles are removed from the adherent surface as wear debris. This type of wear can exist even for bearings designed to operate under fluidfilm conditions. This occurs when the fluid-film thickness becomes less than the mean height of the microasperities, and thus the lubricant film is insufficient to prevent contact of the micro-asperities. Fatigue wear occurs when micro-cracks on the surface or interior, formed by the high cyclic stresses of bearing function, coalesce to form wear particles. This type of wear is generally responsible for delamination of laminated structures, due to bonding agents between the sheets being, in general, weaker than the laminates. Plastic deformation occurs during ploughing when the heavily loaded, harder, smooth bearing surface moves over the softer surface (Figure 2). Here, micro-cracks can be created in the plastically deformed softer surface, and again, with repeated cyclic loading, these micro-cracks coalesce to cause delamination and eventually wear particles. Wear due to ploughing is affected by the surface topography since contact stresses are determined by the shape of the contacting bodies”.

Friction Classically, dry friction, i.e. surface friction without a lubricant has been described by the three laws enunciated by Amonton (1699) and Coulomb (1785):

LOAD v MOTION

Figure2 Ploughing caused by a hard smooth surface as it moves over the softer surface. Biomaterials

bearing

1992, Vol. 13 No. 2

Cartilage

70 1. Frictional force (F) is directly proportional to the applied load (IV). 2. F is independent of the apparent area of contact. 3. The kinetic F is independent of the sliding speed (V), and it is less than static friction. These laws define a coefficient of friction p by the simple, well-known equation F = p W and are operative for hard elastic bodies in contact. In general, kinetic friction is less than static friction. Another type of friction, called bulk friction, occurs because of the internal energy dissipation mechanisms within the bulk material or within the lubricant. Bulk friction is important for relatively soft polymeric viscoelastic materials where internal dissipation or damping is pronounced. In contrast to surface dry friction, bulk friction is dependent on the deformation rate3* 4. The coefficient of friction for fluid-film lubrication also depends on the relative sliding speed, though the dissipation depends on the viscosity of the thin layer of lubricant.

Biotribology In studying the friction, lubrication and wear processes in animal bearings (ankle, elbow, finger, hip, knee, shoulder, wrist), all the elements considered above play significant roles in providing the efficiently lubricated bearings for our bodies. Under normal (healthy] circumstances the coefficient of friction of animal bearings is extraordinarily low (better than most engineering bearings), and they will last seven or eight decades. In this survey we shall describe how the hierarchical nature of the animal bearings provide a unique bearing system with superior tribological functions. Figure 1 and Table 1 summarize the length scales found in a typical joint and in articular cartilage. The lubricant of these joints, i.e. synovial fluid (from Syn - like, ovial - egg white] has significant non-newtonian flow properties which appear to be required for fluid-film lubrication7. The nonnewtonian behaviour is provided by a linear charged (COO- 1.5 nm apart) polymer known as hyaluronic acid (glucuronic N-acetylglucosamine). In synovial fluid from healthy joints, the molecular weight of the hyaluronate is approximately 0.5 X lo6 and its concentration range is 0.1-5mg/ml. The fluid also contains a specific molecule for boundary lubrication of joints” ‘. The hyaluronate of synovial fluid from diseased joints, notably rheumatoid arthritic patients, is depolymerized and the fluid does not exhibit a non-newtonian flow behaviour7. This is believed to be detrimental to fluid-film lubrication in the joint.

Diarthrodial joints Diarthrodial joints have some common structural features. First, all diarthrodial joint are enclosed in a strong fibrous capsule (top left Figure 1). Second, the inner surfaces of the joint capsules are lined with a metabolically active tissue, the synovium, which secretes the synovial fluid (long thought to be the lubricant for diarthrodial joints’) and provides the nutrients required by the avascular cartilage within the joint. Third, the articulating bone Biomaterials

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V.C. Mow et a/.

ends in the joint are lined with a thin layer of hydrated soft tissue known as articular cartilage. These linings, i.e. the synovium and the two articular cartilage layers, form the joint cavity which contains the synovial fluid. Thus, in animal joints, the synovial fluid, articular cartilage, and the supporting bone form the bearing system which provides the smooth, nearly frictionless bearing system of the body. While diarthrodial joints are subjected to an enormous range of loading conditions under cyclical conditionsl’, the cartilage surfaces undergo little wear and tear under normal circumstances.

Joint loading The human hip or knee joint may sustain loads up to ten times body weight during daily activities such as walking or stair climbing. In the hip these forces may yield compressive stresses as high as 18 MPa acting between a metallic endoprosthesis and the acetabulum in vivo while the individual is rising from a chair’l. During prolonged standing, or when a joint is held in a static loaded position, loads several times body weight are commonly found in the knee and hip. Even for the glenohumeral joint of the shoulder, a 50 N weight with the arm held fixed in an extended position will develop a 1500 N joint reaction force because of moment equilibrium. Thus it is known that the mechanical disadvantage of the human joints (recognized by Aristotle: lever law) can cause stresses as high as 6 MPa even in the joints of the fingers, wrist and shoulder. Indeed, most human joints must be capable of functioning effectively under very high loads and stresses and at very low operating speeds. These performance characteristics demand efficient lubrication processes to minimize friction and wear of cartilage in the joint. Severe breakdown of the joint cartilage by either biochemical or biomechanical means leads to arthritis, and thus arthritis is defined as a failure of the animal bearing system*‘.

Friction in animal joints Precise and ingenious measurements have been made on the frictional coefficients of joints and wear coefficients theories have been for cartilage *3-18. Novel lubrication proposed to describe these extraordinary efficient frictional and wear properties1s-22. At present, no comprehensive or consistent theory exists for diarthrodial joint lubrication under all operating conditions. However, ploughing friction (internal energy dissipation] and interface friction appear to be the two dominant mechanisms of energy dissipations, ‘, 16,23. Currently, with the knowledge of the many components of the diarthrodial joint (i.e. the deformational properties of articular cartilage, the rheological properties of synovial fluid, the microanatomy of the articulating surfaces and the kinematics and load-bearing characteristics of these joints), the field of biotribology is poised to make significant advances in understanding joint lubrication, and the breakdown of these lubrication mechanisms which leads to failure of these animal bearings. For an updated description of the biomechanics of diarthrodial joints, where all these topics [including some specialized

Cartilage

Table2 diarthrodial

and diarthrodial Coefficient joints.

joints:

of friction

V.C. Mow et al. for

articular

cartilage

investigator

Coefficient of friction

Joint tested

J. Charnley (1960) C.W. McCutchen (1962) F.C. Linn (1968) A. Unsworth (1975) L.L. Malcolm (1976)

0.005-0.02 0.02-0.35 0.005-0.01 0.01-0.04 0.002-0.03

Human Porcine Canine Human Bovine

Table 3

Coefficients

71 10IOol

in

knee shoulder ankle hip shoulder

of friction for typical materials.

Materials

Coefficient

Gold on gold Aluminum on aluminum Silver on silver Steel on steel Brass on steel Glass on glass Wood on wood Nylon on nylon Graphite on steel Ice on ice at 0°C

2.8 1.9 1.5 0.6-0.8 0.35 0.9 0.25-0.5 0.2 0.1 0.01-0.1

of friction

theories on the lubrication of diarthrodial joints) are treated in detail see Ref. 24. The first reported quantitative measure of the coefficient of friction of animal joints dates back to 1936 by Jones13. Since then many studies on friction and wear have been performed and reported in the literature13-17* 25. Table 2 provides a summary of the range of coefficients of friction available in the literature for some joints. Table 3 provides the coefficients of friction without lubricant for some common engineering materials for comparisonz6. Clearly, animal joints enjoy a very low coefficient of friction when compared with common engineering materials. In addition, for animal joints, the static coefficient of friction seems to be lower than the kinetic coefficient of frictionI”. This may be a significant tribological fact for the function of diarthrodial joints which are always loaded in a cyclical manner. Historically, the fact that diarthrodial joints enjoy such favourable frictional characteristics has been the major motivational force behind a world-wide effort to study cartilage and animal joint biomechanicsz4. Mechanism for ploughing friction Joint cartilage is a soft tissue with a compressive modulus of less than 1.5 MPa and a shear modulus of less than 0.5 MPa; the Poisson’s ratio of the tissue ranges from 0 to 0.42. It is a highly hydrated tissue with its water content ranging from 60 to 85% (see section on cartilage composition for more detail). This water occupies molecular-size ‘pores’ with a diameter estimatedz41 27,” to range from 2 to 6 nm. Most of these pores are ‘open’ and ‘connected’, allowing passage of fluid. A small portion of the water occupies space within the collagen fibres, and this water is not free to move”. Thus, when cartilage is deformed, the water in the open compartment of the tissue will flow and distribute throughout the tissue. This interstitial fluid flow is resisted by a high drag force [experimental evidence suggests that the

Hysteresis

loop

Figure3 Predicted hysteresis loop and experimental data from cartilage. Analytical solution was derived from the nonlinear strain-dependent biphasic theory.

chondroitin sulphate chains on the proteoglycan molecules are responsible for the high resistance to fluid flow3’). The permeability coefficient [k) is inversely related to the drag coefficient (1(1 by the simple relationship (k = @/K; @ = porosity). The permeability of normal cartilage is very low with a coefficient ranging from lo-l4 to lo-l5 m4/N s (Ref. 27). When a joint moves, ploughing of cartilage occurs (see Figure 2)“j. This ploughing deformation will cause interstitial fluid flow, and thus internal dissipation. This internal dissipation is the mechanism that gives rise to ploughing friction. From uniaxial, saw-tooth compression tests, the hysteresis of cartilage has been theoretically predicted and experimentally measured. Figure 3 is a comparison of the predicted energy dissipation density [area enclosed in the stress-strain diagram) from the interstitial fluid flow mechanism to the experimental data. The predictions were mathematically derived from biphasic mixture theory for hydrated soft tissues where the permeability depends exponentially on the volume change during compression (seeEquation 7 later, and Ch. 8 of Ref. 24). Lubrication

theories

for animal

joints

Many experiments have been performed to determine the role of synovial fluid or its components in joint lubrication’, ‘I 15-17.This has been difficult to do since the deformation and flow properties of cartilage and synovial fluid differ significantly with varying physiological circumstances, anatomical form of the joint and different joint loadings. Further, the mechanics of flow of such non-linear fluids as the synovial fluid in a thin-film gap between two porous-permeable hydrated soft tissues has many formidable theoretical difficulties. To date no problem addressing fluid-film lubrication in a conBiomaterials

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figuration modelling diarthrodial joints has ever been solved. In a series of experiments examining the molecular aspects of boundary lubrication, synovial fluid specimens were passed through fine-pore filters with various pore sizes, and/or treated with various degradative enzymes to digest the proteins and hyaluronate chains in the fluid” ‘. The components from the separation procedure were subsequently tested for their lubrication efficacy. From these studies a glycoprotein in synovial fluid was found which acts as a lubricant under boundary lubrication conditions. Also it was found that the lubrication properties of synovial fluid treated with hyaluronidase differed significantly from untreated synovial fluid. These two results suggest that a mixed lubrication process might occur within diarthrodial joints where the non-newtonian fluid plays a role in fluid-film lubrication and the lubricating glycoprotein (LGP) plays a role in boundary lubrication, Since animal joints operate at very high loads and slow speeds, it is unlikely that a sufficiently thick fluid film can be generated by any of the classical fluid-film lubrication mechanisms (e.g. elastorheodynamic theory) which would allow appreciable load-carrying capacity. Nevertheless, the literature does argue that the extremely low coefficient of friction prohibits ruling out fluid-film lubrication. Thus, after more than half a century of research on diarthrodial joint lubrication, the enigma still remains, namely, how can fluid-film lubrication be achieved in a diarthrodial joint to provide the extraordinarily low coefficients of friction? The answer may lie in the biphasic nature of cartilage and the mechanics of interstitial fluid flow.

Wear of articular

cartilage

It is estimated that a human knee or hip joint may experience one million cycles of loading per year. These large cyclical stresses and strains may cause fatigue micro-cracks on the articular surface or within the bulk material, and these may grow and accumulate into microscopically observable damage to the articular surface. Figure 4a is a scanning electron micrograph (SEM) at 3000X original magnification showing a microcrack on the surface of human hip joint cartilage removed during surgery. Collagen fibres spanning the crack tips are clearly visible in Figure 4a. It is likely that the collagen fibres at the articular surface serve as crack arresters. Micro-cracks can be formed within cartilage subjacent to the surface. These micro-cracks can coalesce, eventually causing delamination of the surface membrane from cartilage. Figure 4b is another SEM at 1000X original magnification showing a surface layer being removed from a different human hip joint cartilage specimen. In time, if the rate of this kind of damage exceeds the rate at which the cartilage cells can repair the tissue, an accumulation of damage will occurs which will lead to bulk tissue failure”, 31. Thus the fundamental difference between wear of articular cartilagein vivo and wear of an engineering bearing is that in the biological system, there is a balance of mechanical attrition and biological synthesis and repair. However, the rates of collagen synthesis and turnover in cartilage are very low, usually considered to be less than 1% per year (see later). There is a hypothesis in the literature which states that Biomaterials

1992, Vol. 13 No. 2

Figure 4 a, A typical scanning electron micrograph (original magnification 3000X) showing a micro-crack on the surface of human hip joint cartilage removed during surgery. b, SEM (original magnification 1000X) showing a surface layer being peeled off of a human hip joint cartilage specimen.

fatigue of the collagen network is the primary factor leading to cartilage damage and arthritis3’. As attractive as this hypothesis is in its simplicity, no definitive progress has been made to verify its validity in the nearly 20 years since it was proposed. We have found that, because of confounding viscoelastic effects, it is difficult to measure the tensile fatigue characteristics of collagen fibres. Wear from cartilage rubbing against cartilage is not the result of adhesive wear since cartilage surfaces do not weld together as do metal surfaces when their microasperities come into contact. However, cartilage must articulate against a metal endoprosthesis when one side

73

Cartilage and diarthrodial joints: V.C. Mow et al.

of a damaged joint is replaced by a metallic device, Studies of wear of cartilage rubbing against a polished smooth metal surface have been reported”. In these in vitro experiments excised cartilage specimens were bathed in a normal physiological solution and loaded against a polished stainless steel surface at moderately high pressures (4.62 MPa) under steady harmonic sliding motion, Loss of collagen (hydroxyproline] from the tissue was used as the indicator of wear rate. It was found that even at moderately high pressures, the wear rates were generally low ( aggregates

Shear-dependent

viscosity

q @ 10 s-l = O(10 Pas = 100 poise) @ 50 mg/ml q @ 100 s-l = 0 (1 Pas = 10 poise) @ 50 mg/ml proteoglycan monomers < aggregates larger proteoglycans + higher viscosity link-stabilized aggregates > link-free aggregates

Figure8 The collagen network interacting glycan network in the extracellular matrix reinforced composite.

with the proteoforming a fibre-

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Cartilage

Cartilage

Interactions between collagen and proteoglycan are also required in the biology of tissue maintenance. First, biochemical evidence exists indicating that collagen and GAG synthesis are closely interdependent4’. Also, it is known that proteoglycan monomers can accelerate fibril formation while proteoglycan aggregates have little influence on fibril formation, Thus, it is possible that different proteoglycans in the matrix affect fibril formation differently, giving rise to the possibility of biologically controlled collagen fibre architecture throughout the tissue.

Interstitial

water

The intermolecular space within the tissue is filled with water and dissolved electrolytes. The water content of cartilage ranges from 60 to 85% by wet weight27-2g* 4gV 54-5g. From hydraulic permeability experiments, the pore size in articular cartilage has been estimated to range from 2 to 6 nmZ7* ‘s. Such small pores provide an effective barrier against transport of large molecules through the tissue. Thus frictional interaction between interstitial water and the walls of the ‘nano-size’ pores of the solid matrix also occurs within the nano-scale range. A small percentage of this water is contained in the intracellular compartment, about 30% exists in the intrafibrillar compartment of the collagen fibres and the remainder exists in the solution domain of proteoglycan molecules”’ ‘j”,The amount of water present in the interstitium depends largely on several factors: (1) the FCD associated with the proteoglycans [i.e. the charge-charge repulsive force exerted by the closely spaced negative charge groups on the proteoglycans) and the concentration of ions dissolved in the interstitial fluid - Donnan osmotic pressure effect; the combined effect is known as the of the collagen swelling pressure5’* 52; (2) the organization network strengths and stiffnesses (i.e. anisotropic and inhomogeneous distribution of the collagen): and (3) the material properties of the collagen-proteoglycan solid matrix38-40,51' 52,54,61

and diarthrodial

structure

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(micro-scale:

V.C. Mow et al.

10-7-10-4 m)

The hierarchical nature of cartilage becomes even more apparent at the micro-scale level. Light and low-power electron microscopy studies have shown that articular cartilage can be regarded as having three separate structural zones. At times this layering structure is even visible to the naked eye. Figure 9 shows the arrangement of the cells throughout the surface, middle and deep zones. There are some data in the literature which suggests that cells from different zones have different metabolic and synthetic activities33. Also, the arrangement of these cells is reflected in the organization of the dense collagen network of the extracellular matrix of cartilage [Figure 10). It has been suggested that these two aspects of tissue organization are related by the synthetic activities of the cells in each zone. The collagen and proteoglycan content in the tissue also varies with depth from the articulating surface. The collagen content is highest in the surface zone, and there is a decrease of approximately 15% in the middle and deep zones. The proteoglycan content is inversely related to the collagen content, being lowest at the surface and rising 15% in the middle and deep zones. In human articular cartilage this rise may be as high as 50%“~. Water content variations reflect those seen for collagen, being the highest at the surface (-80%) and decreasing to -65% in the deep zone. On the articular surface, a specific collagen fibre orientation can be demonstrated by using the old technique of generating a ‘split-line pattern’ on the articular surface by puncturing it with a blunt-round aw1”3V”4. Indeed, many tensile testing procedures have used this split-line pattern to provide the appropriate orientation for harvesting specimens4” 54V 65*‘j6. In a section below, we provide the data from tensile testing of specimens obtained parallel and perpendicular to the predominant split-line directions. From sections of cartilage taken perpendicular to the articular surface, the layer appearance of collagen ultrastructure is apparent. In the superficial tangential zone (lo-20% of the total

,Articular

surface STZ (lo-20%)

Mlddle zone (40.60%)

Deco zone (30%) Calclfled zone

Subchondral

a

b

Figureg

The arrangement of chondrocytes during ageing thinning cartilage.

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1992, Vol. 13 NO. 2

throughout

the surface,

\

bone

‘Tide mark Chondrocyte

middle and deep zones of cartilage.

The tidemark

advances

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and diarthrodial

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77

V.C. Mow et al.

Superficial tangential (10-200/o)

-

Calcified cartilage ”

Subchondral bone Cancellous bone

a Deep zone

Figure 10

The organization

of the dense collagen network throughout

tissue thickness), fine collagen fibrils are organized parallel to the articular surface (Figure 20, top insert)67* 68. In the middle zone (40-6070 of the total thickness), the collagen fibres have a larger diameter, and appear to be randomly arranged, though a slight preference of 45’ orientation has been noted6g. The preference for this orientation is depicted in Figure 10. Below this, in the deep zone (about 30% of the total thickness) the fibres appear to be woven together thus forming large fibre bundles organized perpendicular to the surface7*. These bundles cross the ‘tidemark’ and insert into the calcified cartilage providing a strong anchoring system for the tissue on to the subchondral bone7’. 71.

Surface

roughness

To assess whether or not a predicted fluid-film lubrication mode is realistic, it is necessary to know the surface roughness of articular cartilage. As seen in Figure 4, damaged articular cartilage is roughened and characteristically shows micro-cracks and surface delaminations. This micro-structural damage alters the fluid transport properties of cartilage across the surface and also the ability of the collagen-rich surface zone to carry tensile stress. However, the surface of normal living articular cartilage is not smooth either7z*73. There are four distinct orders of surface ‘roughness’, The primary surface features are the anatomical contours of the joint (i.e., macro-scale: see later). These contours form the bearing geometries of the joint. The secondary contours are irregularities of less than 0.5 mm diameter, often seen on low-power SEM photographs. The tertiary surface features are hollows or protrusions of cells subjacent to the surface, and they are characteristically 20-45 pm in diameter and 0.52,Oprn in depth, Figures 128 and b. --

the three major zones of cartilage.

b

Finally, there are regularly appearing shallow quarternary ridges l-10pm in wavelength on the surface. The amplitude of these regularly appearing waves is estimated to be less than l.O,~m. All these surface features have implications for the fluid-film lubrication mechanism existing within the diarthrodial jointZ0-26. For example, the declivities on the articular surface may serve to trap pockets of fluid or hyaluronate gel for ‘boosted lubrication’ 20*25. Also, micro-surface protrusions are likely to be flattened during Ioading73, thus yielding a reserved pool of fluid lubricant trapped in these surface regions. Recently, a new theory on joint lubrication has been proposed which includes the micro-roughness of the articular surface”, 24. From theoretical analysis, the micro-surface roughness seems to enhance the prospect of fluid-film lubrication.

Anatomy of diarthrodial 0.5-15 cm)

joints (macro-scale:

The diarthrodial joints of mammalian bodies come in many forms. The hip (5-10 cm) is a close-fitting deep ball-and-socket joint that in the young is nearly spherical. The acetabulum (socket] is a deep cup forming a very stable joint. However, at times, congenital childhood abnormalities of acetabulum shape (shallow cup) lead to severe hip diseases. Also deviations from sphericity of the articular and subchondral bone surfaces on the acetabulum and femoral head (ball] have been noted7’. In arthritis, both sides of the hip are anatomically grossly distorted by the formation of bony spurs (osteophytes) which in their advanced stages severely restrict the motion of the hip12375. The shoulder (5-10 cm) is also a close-fitting ball-and-socket joint, though the glenoid surface is a very shallow cup covering less than one-third ---..---__~__ ~iomaterials

1992, Vol. 13 No. 2

Cartilage

and

d~a~hrodial

..-

__.....

joints:

V.C. how __--...

et a/.

__

provide quantitative measures of anatomic surfaces’** 77S 7a. The knee @O-15 cm) and the ~arpometac~al (l-2 cm: basal joint of the thumb) joint have been extensively characterized7g9 “. Figures lL?a, b and c show the surface contours and cartilage thickness contours of the three surfaces of a knee joint (patella, tibia1 plateau, and distal femur). The patella articulates and glides in the trochlear groove of the femur, and this forms the patellofemoral joint within the knee. This articulation is congruent in the coronal plane, but very incongruent in the sagittal plane, This joint is frequently subject to compressive loads as high as 10X body weight (during deep knee bends) and, particularly in the female population, it is a frequent site of orthopaedic problems and arthritis. Clearly, along with the applied load and joint motion, the anatomical form of the joint determines the magnitude and distribution of the contact stress, as well as its mode of lubrication. These factors, and the thickness contour, determine the state of stress acting within the cartilage. The carpometacarpal (CMC) joint is widely regarded as a saddle-shaped joint suitable for the bidirectional motion required by the thumb. Figures ~fa and b show the location of this joint in the thumb and the orientation P

Thickness

L

mm

T.00 I.00 D.00

a

Thickness 3.00 6.00 5.00 4.00

i‘: L

‘* 3.00 2.00 1.00

mm

Rgurelt Normalhealthy smooth

but commonly

subjacent

to the surface

living

shows, and,

b,

articulaf

a, protruding wave-like

surface

is

not

chondrocytes

surface

‘ripples’.

of the spherical bumeral head7”. Of all the joints in the body the shoulder provides the greatest range of motion. This is facilitated by the motion of the scapula and a loose capsule around the glenohumeral joint. Because of this, however, the shoulder is very susceptible to instability and dislocation, i,e, shifting of contact region within the joint, and eventually the development of OA. For both the hip and the shoulder, under normal conditions, the articulation is characterized by a sliding motion of one articular surface over the other. Thus, the close-fining sliding contact would promote the formation of ~uid-film lubrication in these joints. a close-range analytical stereophotoRecently, grammetry (SFG) technique has been developed to Biomateriais

1992, Vol. 13 No. 2

0.00

Thickness 4.00 3.20 2.40

1.60 ““’ 0.80 mm

0.00

c FIgurei maps for: human

The

a, the

surface contours and cartiiage thickness patella; b, tibiai plateau and, c, distal femur of a

knee joint.

79

Cartilage and diarthrodial joints: V.C. Mow et al.

Metacarpal Thumb

CMC Joint 1

b

Trapezium

dorsal

radial

ulnar

volar

C

volar Trapezium

Metacarpal

Figure 13 Location of the thumb carpometacarpal (CMC) joint and the orientation of the saddles relative to the bones, a, b elevation contour maps for each side of the CMC joint (metacarpal and trapezium) demonstrating its saddle-like nature, c.

of the saddles relative to the bones, and Figure 23~ shows an e~e~atton contour map for each side of the CMC joint (metacarpal and trapezium) which clearly demonstrates its saddle-like nature. From recent studies, it was found that the human CMC joint may be precisely described by

a biquintic function So. The numbers on the contours of Figure 13~ show elevation in millimetres for the articular surface away from the saddle point. F&ures 14a and b show the principal curvature directions on the metacarpal dictate the and trapezium surfaces. These directions

dorsal

dorsal

ulnar

ulnar

radial

volar

a Figure 14

b The principal curvature directions on, a, the metacarpal and, b, trapezium (b) surfaces. Biomateri~ls

1992, Vol. 13 No. 2

80

Cartilage and diarthrodial joints: V.C. MOW et al.

normal directions of motion for the CMC joint. Again joint instability (for example, due to mechanical insufficiencies in the ligament constraints around the joint) and shifting of contact areas may cause arthritis to develop in this joint. Arthritis in this joint is more severe in the female population.

BIOLOGICAL MAINTENANCE OF THE EXTRACELLULAR MATRIX Articular cartilage is a metabolically active tissue that is synthesized and maintained by the chondrocytes. Cartilage is an unusual tissue in that the extracellular matrix contributes most of the tissue volume, and the cells occupy only a small proportion of the total volume (>cF, the difference between the total number of ions in the tissue (2c + cF) and the total number of ions outside (2c*) is given by: (zc

+

_ zc*

,$I

=

k?

(13)

(4c*) 0.6

0.2

r

: 0

0.1 Fixed charge

0.3

0.2

density

0.4

cF (mEq/ml

0.5

water)

Figure26 The total ion concentration in cartilage as a function of the fixed charge density cF at an external concentration c’ = 0.15 M NaCI. -, ion concentration in cartilage 2c + cF; ---, ion concentration in solution 2~‘. Biomaterials

1992, Vol.

13 No. 2

90

___-..._

Cartilage and diarthrodiaf joints: V.C. A9ow et a/.

~~_.

Thus, the difference approaches zero as the external concentration c* + co. The right side of Equation 13 provides a simple expression for calculation of the Donnan osmotic pressure. From the classical theory for osmotic pressure, the Donnan osmotic pressure n due to the excess of ion particles inside the tissue is given by: Jr = RT[@(Zc + 2) - z@*c*] + P,

(141

where R is the universal gas constant, T is the absolute temperature, d, and # * are osmotic coefficients and P, is the osmotic pressure due to the concentration of proteoglycan particles in the tissuez8* 49,‘**“* I”. Usually, because of the size of proteoglycan molecules, P, is negligible. Thus, Equations 13 and 24 show that the Donnan osmotic pressure TT= RT(cF)‘/(4c*) would approach zero as c* + co. Also, with increasing ion concentration within the tissue, the equivalent Debye length between the ion cloud and the fixed charges is decreased. This results in charge shielding which decreases the net charge-to-charge repulsive force and thus decreases the proteoglycan solution volume51* 52. Decrease of the size of the PC molecule in solution with increased electrolyte concentration, i.e. charge shielding, has been measuredl”. Changes in the dimensions of the tissue specimen corresponding to these proteoglycan volumetric changes have also been measured. Articular cartilage swelling has been measured dimensionally when c* is changed”‘. These measurements show that swelling of cartilage is non-homogeneous (Figure 27a) and anisotropic (Figure 27b). The length and the directions correspond to directions parallel and perpendicular to the split-line directions. The dimensional change is the largest along the thickness direction and in the deep zone of the tissue where the proteoglycanlcollagen ratio is highest. As can be seen in Figure ,??‘a, the dimensional change along the split-line direction of cartilage under free swelling conditions varies linearly with c* over the range 0 to 0.2 M NaCl. This result is described by a linear constitutive relationship :

E, =

-a,c*

(151

where a, is known as the coefficient of chemical contraction, and E, is the chemically induced strain’Z8; the negative sign indicates that an increase in c* produces a contraction of the tissue. For articular cartilage, the value a, ranges from 0.05 to 0.3 M-’ depending on the depth within the tissue. If the tissue is isometrically constrained, the equilibrium chemicalexpansion stress T, would also vary linearly with c*, given by the relationship I&E, = -Esocc*. Thus the chemical-expansion stress T, in Equaijon 11 must depend on cr, the fixed charge density, and the external bathing solution concentration c*, i.e. T, = Tc(cF, c*), When the c varies over a greater range, experimental data on dimensional swelling suggests that T, may decrease exponentially with increasing c* and that chemicalexpansion effect must vanish when cF is zero’*. This suggests that a good constitutive model for the chemicalexpansion stress T, should be: T, = aocFexp(-Kc*)

@Sal

where a, and Kare two material coefficients. By virtue of the Donnan equilibrium ion distribution law (Equafiun 12) this expression may be written entirely in terms of cF and the internal counter-ion concentration c: T,(c,

cF]= aocFexp

For commonly used values, i.e. cF = 0,l mEq/ml, yt/ rf = 1, @/@*= 1, and with @*(c*)given by”‘, the values fora, and Khave been calculated to bea, = 2.5 MPa/(mEq/ ml) and K = 7.5 M-‘+ A summary of the mechanoelectrochemical properties of articular cartilage is provided in Table 9. A triphasic mechanoelectrochemical theory (comprising a miscible ion phase along with the two immiscible fluid and solid phases of the biphasic theory) has been developed to account for the swelling effects from the Donnan osmotic pressure and the chemical-expansion stress5’q “, along with the biphasic deformational

6

3

0 0.00

a

0.05 NaCI

0.10 concentration

0.15 fM)

0.20

I

I

Length

Width

Thickness

b

Figure 27 Swelling of cartilage, as measured by dimensional changes, is, a, non-homogeneous and, b, anisotropic. The length and width directions correspond to directions parallel and perpendicular to the split-line directions. a, El, surface zone; I), middle zone; A, deep zone. b, Surface zone only. Biomaterials 1992, Vol. 13 No. 2

Cartilage

and diarthrodial

Table 9

V.C. Mow et al.

joints:

Summary of mechanoelectrochemical

91

properties

of properties

of articular cartiiage5*.

cF (mEq/ml)

rr (MPa)

a, W’)

a, (MPa/(mEq/ml))

K (M-l)

0.01-0.30

0.1-0.2

0.05-0.3

2.5 (o-lo)*

7.5(0-lo)*

*The range of values shown are estimates.

effectsz7% lzl. In this theory, analogous to Equation 8 for the biphasic theory, the total stress oT may now be decomposed into partial stresses, an elastic stress (0’)s and the swelling pressure P, which includes an osmotic pressure: cT = (cTS)s- P,Z

(17)

matrix. (From this relationship, clearly, for an anisotropic elastic solid matrix, the swelling will be anisotropic.) From this expression, the change of volume per unit volume due to swelling may be obtained: AV

~7+ T,

n + T, B

V

where p, = cp + T,) At equilibrium, the fluid pressure p = TI is the Donnan osmotic pressure given by Equation 14. For cF = 0.1 mEq/ml, y2 lyf- = 1, Q/Q* = 1, with Q*(c*) given by Ref. 129, a, = 2.5 MPa/(mEq/ml) and K = 7.5 M-l, Figure 28 shows the relative contribution of Donnan osmotic pressure and chemical-expansion stress to the total swelling pressure P,. At normal saline concentration, i.e. 0.15 M NaCl, the Donnan osmotic pressure and the chemical-expansion stress contribute in nearly equal proportions. For an unloaded tissue at swelling equilibrium, the total stress crT acting on the tissue is zero. Thus, the stress acting on the elastic collagen-proteoglycan solid matrix (~3)s due to the internal swelling pressure is given by: (d)E = (n + T,)Z

(181

Since the assumed constitutive law for the solid matrix stress (0’)s is linearly elastic and isotropic, Equation 18 may be written as: (1%

A,(tr&)Z + 2,ug = (n + TJZ where 0.5

As, ,us are the Lame’s

coefficients

of the solid

r

0.15 Bathing

1.o

0.5 solution

concentration

1.5

where B is the bulk modulus. Thus at every point within the tissue there is a volumetric change due to the swelling pressure. We see that as long as there is a Donnan osmotic pressure R and a chemical-expansion stress T, associated with cF, the porous-permeable, collagenproteoglycan solid matrix will be in an inflated state, i.e. AVIV > 0, Figure 23. Clearly, the balance between the elastic stress developed within the porous-permeable solid matrix and the swelling pressure associated with the FCD determines the degree of tissue swelling or hydration. An important consequence of these two charge swelling mechanisms is that they can be used to modulate the tension, compression and shear stiffnesses of articular cartilage or any charged-hydrated soft tissue5’, 52X128.By simply controlling the nature of the electrolyte (concentration and valance, e.g. Na+ or Ca’+) environment around the tissue, the Donnan osmotic pressure and chemical-expansion stress will be altered. The FCD, and therefore the internal ion concentration, provide the fundamental mechanoelectrochemical mechanisms for tissues (and likely, any other aqueousbased organis,ms with similar characteristics) to control their material properties. Also, by regulating their internal FCD, the tissues may control their electrochemical responses to external stimuli. This phenomenon is vital in all biological processes. To conclude our discussion, we shall briefly describe some simple mechanoelectromechanical phenomena. The electrochemical potentials of the counter-ions and co-ions, e.g. Na+ and Cl-, within the interstitium of cartilage are given by:

where & (a = + for cation and a = - for anion) are the reference potentials of the anions and cations, M” are the atomic weight of these ions, F, is the Faraday constant and y is the electric potential13’. From these electrochemical potentials, it can be shown that the streaming potential of electrolyte solutions flowing through a charged hydrated soft tissue such as articular cartilage is given by:

c*(M)

Figure28 The relative contribution of Donnon osmotic pressure (---) and chemical-expansion stress to the total swelling pressure P, (-).

Aw =

-FCcFk Ap .

_

KO

and it can be shown that from electro-osmosis, Biomaterials

the fixed

1992. Vol. 13 No. 2

92 --

Cartilage

charge density formula:

cF may

be determined

by the simple

In Equation 22, K,, is the electric conductivity of the tissue. From the basic theoretical construction of the mechanoelectrochemical theory5’l 52 one can obtain, in the limiting case of infinitesimally thin membranes or small motive forces applied across a tissue specimen, a set of governing equations entirely equivalent to classical electrochemical theory for charged semipermeable membranes13’. For example, from the triphasic mechanoelectrochemical theory, the Kedem-Katchalsky force-flux phenomenological flow equations may be derived: LpDAn

In = LD, AP t LD An

in cartilage

properties

4

18 -

1992. Vol. 13 No. 2

Q) = 2 g

664-

0123456789

-

CollagerVProteoglycan (Ratio) Figure29 Decrease of the equilibrium tensile modulus of human knee joint cartilage with collagen/proteoglycan ratio (ages 24 to 27 y.0.); the tissue from each joint was assessed histologically as normal (0), mildly fibrillated (A), or osteoarthritic ( n ). Table 10 Summary of changes on the tensile modulus (E)*, and water, collagen and cross-link contents.

Controls Experimental

E (MPa)

H,O%

OH-Pro/wet

13.2 6.17

78.4 84.3

12.37 9.27

w-t PYD/Coll 1.74 1.55

during arthritis

The descriptions provided above show that the composition and ultrastructure of nature’s bearing material has been designed to function specifically in the high loading environment of diarthrodial joints. There is an exquisite balance of force due to the swelling pressure P, associated with the charge groups on the proteoglycans and the restraining force developed within the surrounding collagen network62. This balance of force keeps the collagen network in an inflated state for load carriage. Within the framework of the triphasic mechanoelectrochemical theory, this balance is expressed in terms of volumetric change AVW and the swelling pressure P,; the proportionality constant is the bulk modulus B of the solid matrix (Equation 20). If the collagen network is damaged in some manner as typically occurs during OA (e.g. see Figures 4 and 381, the tensile modulus would dramatically decrease. Figure 29 shows a decrease of the equilibrium tensile modulus of human knee joint cartilage (ages 24 to 74 years); the tissue from each joint was assessed histologically as normal, mildly fibrillated, or 0A3’. Clearly there is a significant drop of cartilage tensile stiffness with increasing severity of lesion, reflective of the damaged collagen network. Table 10 provides the tensile stiffness, water content, OH-proline (collagen) and PYD/ collagen [cross-link density) data determined from a well-controlled set of canine cartilage specimens obtained from anterior cruciate ligament transected knees 4 months post-op [i.e. the Pond-Nuki Biomaterials

W.C. Mow et al.

(241

where L,, LD and LpD ( =LDp) are phenomenological flow coefficients. In these equations, Jv is the total volume flux of water through the specimen and In is the anion transport relative to the water. These flow coefficients depend entirely on the fundamental set of coefficients defined in the triphasic mechanoelectrochemical theory, e.g. Equations 21-2351* 52. Further, the mechanoelectrochemical theory provides a method to actually calculate the stress-strain, ion concentration, electric and flow fields within a layer of charged hydrated soft tissue of finite thickness. Al1 these effects (mechanical, chemical, electrical] may be important in controlling chondrocyte metabolism and maintaining tissue homoeostasis. All these effects are manifestations of interactions occurring at the nano-scale level.

Changes

joints:

20 1231

JV = L,AP+

and diarthrodial

*The equilibrium stress-strain relationship Pro, proline; PYDKoll, PYDkollagen.

for canine cartilage

is linear.

procedure)131. This surgical procedure is known to produce early OA changes in the canine model similar to human OA changes13’. These results show that as early as 4 months post-op, there is a dramatic drop in the tensile modulus of the canine knee joint tissue similar to human tissue, and there is also a significant increase of water content, and a loss of collagen and cross-link content within the tissue from the impaired knee. Typically, OA tissues also show an increase in water content and a decrease in proteoglycan content 56, 57,133,134. Associated with the loss of proteoglycan content is an increase in the collagen/proteoglycan ratio. Figure 29 shows the correlation between this ratio and the tensile modulus of normal, mildly fibrillated and OA human knee joint cartilage. This correlation suggests not only that the cbllagen network has failed during OA but that the collagen-proteoglycan solid matrix has also failed. Clearly, from the preceding discussion on the biphasic nature of cartilage, one can anticipate that water content of the tissue is a major factor determining its compression properties. Changes in the water content can occur when an imbalance exists in proteoglycan metabolism. For example, the net loss of proteoglycan from within the tissue leads to an increase in space available for water within the intrafibrillar space. It would also produce an associated decrease of cF, A decrease of cF results in a

Cartilage

and diarthrodial

joints:

V.C. Mow et al.

93

decrease of the Donnan osmotic pressure, see Equation 14. However, a weakening of the collagen-proteoglycan solid matrix, see Equation 20, would allow the tissue to swell more. The net effect of these three competing processes is an increase in tissue hydration5”l 57,1331134,135. Figures 30a and b show the effects of increased water content and decreased GAG content (chondroitin sulphate, keratan sulphate and hyaluronate). The loss of the equilibrium compressive modulus is due to an increase in porosity and a decrease in FCD (Donnan osmotic pressure). Also associated with these changes is an increased permeability56. These material property changes act to defeat the fluid pressurization mechanism for load support in cartilage, thus the weakened collagen-proteoglycan solid matrix must support

3

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0.6

s’ 5

0.5

35 9 w Z pE $ 8

2 IL!

0.4 0.3 0.3

1 0.0

)

0.74

0.76

a

0.78

0.80

0.82

0.84

0.86

WATER CONTENT

l

.

. .

ii

2

s

.I

0.0 0.01

0.02

0.03

GLYCOSAMINOGLYCAN

0.04

0.05

(PER WET WEIGHT)

b Figure 30 Changes in the equilibrium compressive modulus cartilage of articular with water content (a) and glycosaminoglycan content (b).

additional stresses and strains. Eventually, failure of this animal bearing material occurs and arthritis develops.

SUMMARY Articular cartilage and diarthrodial joints are hierarchical materials and structures which seem to have been specifically designed by nature to provide the excellent lubrication and load-carrying capacity required by joint function. At the macro-scale level [O.5-15 cm], each joint of the mammalian body has a regular but complex anatomical form. The form of the joint is quantifiable using such techniques as stereophotogrammetry, and each anatomical form is related to the function that the joint must provide. Changes in the anatomical form such as those occurring during arthritis hinder joint function. The synovial fluid has non-newtonian behaviours and contains a lubricating glycoprotein which is required for joint lubrication. The deformational behaviour of cartilage is best described by a fluid-solid mixture theory known as the biphasic theory. The pore size of the collagen-proteoglycan solid matrix is exceedingly small (

Cartilage and diarthrodial joints as paradigms for hierarchical materials and structures.

The anatomic forms of diarthrodial joints are important structural features which provide and limit the motions required for the joint. Typically, the...
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