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Biomechanics of Cartilage
J O S E P H M. M A N S O U R , P H. D.
COMPOSITION AND STRUCTURE OF ARTICULAR CARTILAGE.......................
MECHANICAL BEHAVIOR AND MODELING.....................................
MATERIAL PROPERTIES....................................................
RELATIONSHIP BETWEEN MECHANICAL PROPERTIES AND COMPOSITION............
MECHANICAL FAILURE OF CARTILAGE........................................
JOINT LUBRICATION......................................................
MODELS OF OSTEOARTHROSIS..............................................
SUMMARY...............................................................
The materials classed as cartilage exist in various forms and perform a range of functions in the body. Depending on its composition, cartilage is classified as articular cartilage (also known as hyaline), fibrocartilage, or elastic cartilage. Elastic cartilage helps to maintain the shape of structures such as the ear and the trachea. In joints, cartilage functions as either a binder or a bearing surface between bones. The annulus fibrosus of the intervertebral disc is an example of a fibrocartilaginous joint with limited movement (an amphiarthrosis). In the freely moveable synovial joints (diarthroses) articular cartilage is the bearing surface that permits smooth motion between adjoining bony segments. Hip, knee, and elbow are examples of synovial joints. This chapter is concerned with the mechanical behavior and function of the articular cartilage found in freely movable synovial (diarthroidal) joints.
In a typical synovial joint, the ends of opposing bones are covered with a thin layer of ar- ticular cartilage (Fig. 5.1). On the medial femoral condyle of the knee, for example, the cartilage averages 0.41 mm in rabbit and 2.21 mm in humans [2]. Normal articular carti- lage is white, and its surface is smooth and glistening. Cartilage is aneural, and in normal mature animals, it does not have a blood supply. The entire joint is enclosed in a fibrous tissue capsule, the inner surface of which is lined with the synovial membrane that se- cretes a fluid known as synovial fluid****. A relatively small amount of fluid is present in a normal joint: less than 1 mL, which is less than one fifth of a teaspoon. Synovial fluid is clear to yellowish and is stringy. Overall, synovial fluid resembles egg white, and it is this resemblance that gives these joints their name, synovia, meaning “with egg.”
Cartilage clearly performs a mechanical function. It provides a bearing surface with low friction and wear, and because of its compliance, it helps to distribute the loads between opposing bones in a synovial joint. If cartilage were a stiff material like bone, the contact stresses at a joint would be much higher, since the area of contact would be much smaller. These mechanical functions alone would probably not be sufficient to justify an in-depth study of cartilage biomechanics. However, the apparent link between osteoarthrosis and
C H A P T E R
Chapter 5 | BIOMECHANICS OF CARTILAGE 67
mechanical factors in a joint adds a strong impetus for studying the mechanical behavior of articular cartilage.
The specific goals of this chapter are to
■ Describe the structure and composition of cartilage in relation to its mechanical behavior ■ Examine the material properties of cartilage, what they mean physically, and how they can be determined ■ Describe modes of mechanical failure of cartilage ■ Describe the current state of understanding of joint lubrication ■ Describe the etiology of osteoarthrosis in terms of mechanical factors
A comment on terminology seems appropriate. Osteoarthritis is the term commonly used to describe the apparent degeneration of articular cartilage. Radin has argued that this is a mis- nomer since osteoarthritis does not directly involve inflammation. He suggests the term os- teoarthrosis, which is defined as “loss of articular cartilage with eburnation of the underlying bone associated with a proliferative response [68,69].” In this chapter, the term osteoarthro- sis is used in place of osteoarthritis****. Before proceeding through this chapter, the reader should be familiar with the basic concepts and terminology introduced in Chapters 1 and 2.
Bone
Bone
Articular Joint cartilage capsule Synovial membrane
Figure 5.1: Schematic representation of a synovial joint. Articular cartilage forms the bearing surface on the ends of opposing bones. The space between the capsule and bones is exaggerated in the figure for clarity.
Chapter 5 | BIOMECHANICS OF CARTILAGE 69
framework is an essential element in the mechanical response
of cartilage. When cartilage is compressed, the negatively
charged sites on aggrecan are pushed closer together, which
increases their mutual repulsive force and adds to the com-
pressive stiffness of the cartilage. Nonaggregated proteogly-
cans would not be as effective in resisting compressive loads,
since they are not as easily trapped in the collagen matrix.
Damage to the collagen framework also reduces the com-
pressive stiffness of the tissue, since the aggregated proteo-
glycans are contained less efficiently.
The mechanical response of cartilage is also strongly tied
to the flow of fluid through the tissue. When deformed, fluid
flows through the cartilage and across the articular surface
[42]. If a pressure difference is applied across a section of car-
tilage, fluid also flows through the tissue [51]. These obser-
vations suggest that cartilage behaves like a sponge, albeit one
that does not allow fluid to flow through it easily.
Recognizing that fluid flow and deformation are interde-
pendent has led to the modeling of cartilage as a mixture of
fluid and solid components [59–61]. This is referred to as the
biphasic model of cartilage. In this modeling, all of the solid-
like components of the cartilage, proteoglycans, collagen,
cells, and lipids are lumped together to constitute the solid
phase of the mixture. The interstitial fluid that is free to move
through the matrix constitutes the fluid phase. Typically, the
solid phase is modeled as an incompressible elastic material,
and the fluid phase is modeled as incompressible and invis-
cid, that is, it has no viscosity [60]. Under impact loads, car-
tilage behaves as a single-phase, incompressible, elastic solid;
there simply isn’t time for the fluid to flow relative to the solid
matrix under rapidly applied loads. For some applications, a
viscoelastic model is used to describe the behavior of carti-
lage in creep, stress relaxation, or oscillating shear. Although
the mathematics of modeling cartilage is outside the scope of
this chapter, some examples illustrate the fundamental fluid–
solid interaction in cartilage.
MATERIAL PROPERTIES
A confined compression test is one of the commonly used
methods for determining material properties of cartilage
(Fig. 5.4). A disc of tissue is cut from the joint and placed in
an impervious well. Confined compression is used in either
a “creep” mode or a “relaxation” mode. In the creep mode,
a constant load is applied to the cartilage through a porous
plate, and the displacement of the tissue is measured as a
function of time. In relaxation mode, a constant displacement
is applied to the tissue, and the force needed to maintain the
displacement is measured.
Axis of split line
Bone
Superficial
Intermediate
Radiate
Calcified
Subchondal
Calcified cartilage
Collagen leaves
Figure 5.3: Cross sections cut through the thickness of articular cartilage on two mutually orthogonal planes. These planes are oriented parallel and perpendicular to split lines on the cartilage surface. The background shows the four zones of the cartilage: superficial, intermediate, radiate, and calcified. The foreground shows the organization of collagen fibers into “leaves” with varying structure and organization through the thickness of the cartilage. The leaves of collagen are connected by small fibers not shown in the figure.
70 Part I | BIOMECHANICAL PRINCIPLES
In creep mode, the cartilage deforms under a constant
load, but the deformation is not instantaneous, as it would be
in a single-phase elastic material such as a spring. The dis-
placement of the cartilage is a function of time, since the fluid
cannot escape from the matrix instantaneously (Fig. 5.5). Ini-
tially, the displacement is rapid. This corresponds to a rela-
tively large flow of fluid out of the cartilage. As the rate of
displacement slows and the displacement approaches a con-
stant value, the flow of fluid likewise slows. At equilibrium,
the displacement is constant and fluid flow has stopped. In
general, it takes several thousand seconds to reach the equi-
librium displacement.
By fitting the mathematical biphasic model to the meas-
ured displacement, two material properties of the cartilage
are determined: the aggregate modulus and permeability.
The aggregate modulus is a measure of the stiffness of the
tissue at equilibrium when all fluid flow has ceased. The
higher the aggregate modulus, the less the tissue deforms
under a given load. The aggregate modulus of cartilage is
typically in the range of 0.5 to 0.9 MPa [2]. There is no anal-
ogous material constant for solid materials, but using the
aggregate modulus and representative values of Poisson’s
ratio (described below), the Young’s modulus of cartilage is
in the range of 0.45 to 0.80 MPa. For comparison, the
Young’s modulus of steel is 200 GPa and for many woods is
about 10 GPa parallel to the grain. These numbers show
that cartilage has a much lower stiffness (modulus) than most
engineering materials.
In addition to the aggregate modulus, the permeability of
the cartilage is also determined from a confined compression
test. The permeability indicates the resistance to fluid flow
through the cartilage matrix. Permeability was first introduced
in the study of flow through soils. The average fluid velocity
through a soil sample (vave) is proportional to the pressure
gradient (p) (Fig. 5.6). The constant of proportionality (k)
is called the permeability. This relationship is expressed by
Darcy’s law,
vave kp (Equation 5.1)
Constant load
Articular cartilage Porous plate
Impervious container
Figure 5.4: Schematic drawing of an apparatus used to perform a confined compression test of cartilage. A slice of cartilage is placed in an impervious, fluid-filled well. The tissue is loaded through a porous plate. In the configuration shown, the load is constant throughout the test, which can last for several thousand seconds. Since the well is impervious, flow through the cartilage will only be in the vertical direction and out of the cartilage.
Displacement
Time Figure 5.5: Typical displacement of cartilage tested in a confined compression test. A constant load is applied to the cartilage, and the displacement is measured over time. Initially, the deformation is rapid, as relatively large amounts of fluid are exuded from the cartilage. As the displacement reaches a constant value, the flow slows to zero. Two material properties are determined from this test.
Low pressure (P 1 )
High pressure (P 2 )
Porous plate
Fluid filled chamber
Fluid filled chamber
Direction of fluid flow
Articular cartilage
h
Figure 5.6: Schematic representation of a device used to measure the permeability of cartilage. A slice of cartilage is supported on a porous plate in a fluid-filled chamber. High pressure applied to one side of the cartilage drives fluid flow. The average fluid velocity through the cartilage is proportional to the pressure gradient, and the constant of proportionality is called the permeability.
72 Part I | BIOMECHANICAL PRINCIPLES
changes in mechanical behavior, and additional collagen–
proteoglycan interaction. Tensile tests of cartilage are per-
formed by first removing the cartilage from its underlying
bone. This sheet of cartilage is sometimes cut into thin slices
(200–500 m thick) parallel to the articular surface, using a
microtome. Dumbbell-shaped specimens are cut from each
slice with a custom-made cookie cutter.
A particularly thorough study of the tensile properties of
cartilage shows that samples oriented parallel to split lines
have a higher tensile strength and stiffness than those per-
pendicular to the split lines. In skeletally mature animals
(closed physis), tensile strength and stiffness decrease from
the surface to the deep zone. In contrast, tensile strength and
stiffness increase with depth from the articular surface in
skeletally immature (open physis) animals [76].
The relative influence of the collagen network and pro-
teoglycans on the tensile behavior of cartilage depends on the
rate of loading [77]. When pulled at a slow rate, the collagen
network alone is responsible for the tensile strength and stiff-
ness of cartilage. At high rates of loading, interaction of the
collagen and proteoglycans is responsible for the tensile be-
havior; proteoglycans restrain the rotation of the collagen
fibers when the tissue is loaded rapidly.
RELATIONSHIP BETWEEN MECHANICAL
PROPERTIES AND COMPOSITION
In addition to the qualitative descriptions given above, quan-
titative correlations between the mechanical properties of car-
tilage and glycosaminoglycan content, collagen content, and
water content have been established. The compressive stiff-
ness of cartilage increases as a function of the total gly-
cosaminoglycan content [35] (Fig. 5.8). In contrast, there is
no correlation of compressive stiffness with collagen content.
In these cases, compressive stiffness is measured in creep,
2 seconds after a load is applied to the tissue. Permeability
and compressive stiffness, as measured by the aggregate mod-
ulus, are both highly correlated with water content. As the
water content increases, cartilage becomes less stiff and more
permeable [1] (Fig. 5.9). Note that the inverse of permeabil-
ity is plotted in Figure 5.9 B. This is done for convenience,
Two-second creep stiffness x 10
-^6
(MPa)
60
20
0
40
60
80
120
140
160
100
80 100 120 140 160 Total glycosaminoglycan content (μg/mg dry weight)
Figure 5.8: Correlation of compressive stiffness with the total glycosaminoglycan concentration. As the total glycosaminoglycan concentration decreases, the compressive stiffness also decreases.
Aggregate modulus (MPa)
70
0
75 80 85 90 Water content (%)
A
1/Permeability x 10
-^14
(Ns/m
4 )
70
1
0
2
3
5
6
7
4
75 80 85 90 Water content (%)
B
Figure 5.9: A. Correlation of the aggregate modulus with water content of articular cartilage. A regression line obtained from tests of a large number of samples is plotted. As the water content increases, the aggregate modulus decreases. B. Correlation of the inverse of permeability with water content. A regression line obtained from tests of a large number of samples is plotted. As the water content increases, the permeability increases.
Chapter 5 | BIOMECHANICS OF CARTILAGE 73
since the permeability becomes very large as the water con-
tent increases.
CLINICAL RELEVANCE: MATERIAL PROPERTIES
OF CARTILAGE
The relationships between material properties and water content help to explain early cartilage changes in animal models of osteoarthrosis. Proteoglycan content and equi- librium stiffness decrease and the rate of deformation and water content increases in these models [38,56]. De- creasing proteoglycan content allows more space in the tissue for fluid. An increase in water content correlates with an increase in permeability. Increasing permeability allows fluid to flow out of the tissue more easily, resulting in a more rapid rate of deformation. Using confined compression, indentation, tension, and shear tests, the mechanical properties of cartilage can be determined. These properties are necessary for any analy- sis of stress in the tissue. However, material properties do not give any indication of the failure of cartilage. For example, simply knowing the value of aggregate modulus or Poisson’s ratio is not sufficient to predict if cartilage will develop the cracks, fissures, and general wear that is char- acteristic of osteoarthrosis. Various loading conditions have been used to gain better insight into the failure properties of cartilage.
MECHANICAL FAILURE OF CARTILAGE
A characteristic feature of osteoarthrosis is cracking, fibrilla-
tion, and wear of cartilage. This appears to be a mechanically
driven process, and it motivates numerous investigations
aimed at identifying the stresses and deformations responsi-
ble for the failure of articular cartilage. Since cartilage is an
anisotropic material, we expect that it has greater resistance
to some components of stress than to others. For example,
it could be relatively strong in tension parallel to collagen
fibers, but weaker in shear along planes between leaves of
collagen.
Tensile failure of cartilage has been of particular interest,
since it was generally believed that vertical cracks in cartilage
were initiated by relatively high tensile stresses on the artic-
ular surface. More-recent computational models of joint con-
tact show that the tensile stress on the surface is lower than
originally thought, although tensile stress still exists within the
cartilage [13–15]. It now appears that failure by shear stress
may dominate. Studies of the tensile failure of cartilage are
primarily concerned with variations in properties among
joints, the effects of repeated load, and age.
Kempson and coworkers report a decrease in failure stress
with age for cartilage from hip and knee [30–32, 34]. How-
ever, they find no appreciable age-dependent decrease in ten-
sile failure stress for cartilage from the talus (Fig. 5.10).
CLINICAL RELEVANCE: INCIDENCE OF OSTEOARTHROSIS
AT THE ANKLE
There is a low incidence of osteoarthrosis in the ankle com- pared with the hip or knee. The maintenance of tensile strength of cartilage from the ankle may play a role in the reduced likelihood of degeneration in this joint.
Repeated tensile loading (fatigue) lowers the tensile
strength of cartilage as it does in many other materials. As the
peak tensile stress increases, the number of cycles to failure
decreases (Fig. 5.11) [93–95]. For any value of peak stress,
the number of cycles to failure is lower for cartilage from
older than younger individuals.
Repeated compressive loads applied to the cartilage sur-
face in situ also cause a decrease in tensile strength, if a suf-
ficient number of load cycles are applied [53]. Following
64,800 cycles of compressive loading there is no change in the
tensile strength of cartilage, but after 97,200 cycles, tensile
strength is reduced significantly. Surface damage is not found
in any sample. This shows that damage may be induced within
the tissue before any signs of surface fibrillation are apparent.
Some caution must be exercised when interpreting the re-
sults of tests in which a large strain is applied to cause failure
of samples removed from the joint. The strain to failure may
be greater than that experienced in vivo. In addition, the prop-
erties of most biological materials change with the applied
strain; the collagen network becomes aligned with the direc-
tion of the tensile strain, and the material becomes strongly
anisotropic.
Tensile failure stress (MPa)
1
5
0
10
15
20
30
35
25
20 40 60 80 100 Age in years
Femoral head 40 Talus
Figure 5.10: Comparison of the tensile failure stress of cartilage from the hip and talus. There is a statistically significant drop in the failure stress, as a function of age, for cartilage from the hip, but not for cartilage from the talus. Interestingly, there is a relatively high occurrence of osteoarthrosis in the hip compared with that in the ankle (talus).
Chapter 5 | BIOMECHANICS OF CARTILAGE 75
cartilage experiences large lateral displacement (due to its
high Poisson’s ratio) when loaded in compression, but this ex-
pansion is constrained by the stiff underlying bone (Fig. 5.13).
Under these conditions, high shear stress develops at the
cartilage–bone boundary.
Most studies of cartilage failure are based directly on the
values of ultimate stress or strain. An alternative is to use pa-
rameters that more directly represent the propagation of a
crack in a loaded material sample. The feasibility of using two
methods to determine fracture parameters of cartilage is eval-
uated extensively by Chin-Purcell and Lewis (Fig. 5.14) [9].
The so-called J integral is a measure of the fracture energy
dissipated per unit of crack extension. As used, the J integral
also assumes that a crack propagates in the material, as op-
posed to deformation or flow of the material, which results
in a more ductile failure. Since cracks may not propagate in
soft biological materials, a tear test is also evaluated. The tear
test yields a fracture parameter similar to the J integral. As
with tensile-stress-based ideas of failure, it is necessary to
apply large strains to cause failure of the samples: these strains
may be far greater than those found in any in vivo loading
conditions. To date, the application of these fracture param-
eters is limited to the normal canine patella.
JOINT LUBRICATION
Normal synovial joints operate with a relatively low coeffi-
cient of friction, about 0.001 [40,54,86]. For comparison,
Teflon sliding on Teflon has a coefficient of friction of about
0.04, an order of magnitude higher than that for synovial
joints. Identifying the mechanisms responsible for the low
friction in synovial joints has been an area of ongoing research
for decades. Both fluid film and boundary lubrication mech-
anisms have been investigated.
For a fluid film to lubricate moving surfaces effectively, it
must be thicker than the roughness of the opposing surfaces.
The thickness of the film depends on the viscosity of the fluid,
the shape of the gap between the parts, and their relative ve-
locity, as well as the stiffness of the surfaces. A low coefficient
of friction can also be achieved without a fluid film through
a mechanism known as boundary lubrication. In this case,
molecules adhered to the surfaces are sheared rather than a
fluid film.
It now appears that a combination of boundary lubrication
(at low loads) and fluid film lubrication (at high loads) is
responsible for the low friction in synovial joints [41,74,75].
This conclusion is based on several important observations.
First, at low loads, synovial fluid is a better lubricant than
buffer solution, but synovial fluid’s lubricating ability does
not depend on its viscosity. Digesting synovial fluid with
hyaluronidase, which greatly reduces its viscosity, has no ef-
fect on friction. This shows that a fluid film is not the pre-
dominant lubrication mechanism, since viscosity is needed to
generate a fluid film. In contrast, digesting the protein com-
ponents in synovial fluid (which does not change its viscosity)
causes the coefficient of friction to increase. This result
suggests that boundary lubrication contributes to the overall
lubrication of synovial joints. A glycoprotein that is effective
as a boundary lubricant has been isolated from synovial fluid
[84]. Newer evidence suggests that phospholipids may be im-
portant boundary lubricant molecules for articular cartilage
[17,65,78]. At high loads, the coefficient of friction with syn-
ovial fluid increases, but there is no difference in friction
between buffer and synovial fluid. This suggests that the
boundary mechanism is less effective at high loads and that
a fluid film is augmenting the lubrication process. Numerous
mechanisms for developing this film have been postulated
[12,28,48,54,89,91,92]. If cartilage is treated as a rigid mate-
rial, it is not possible to generate a fluid film of sufficient thick-
ness to separate the cartilage surface roughness. Treating the
cartilage as a deformable material leads to a greater film thick-
ness. This is known as elastohydrodynamic lubrication: the
pressure in the fluid film causes the surfaces to deform. How-
ever, as the surfaces deform, the roughness on the surface
also deforms and becomes smaller. Models, which include de-
formation of the cartilage and its surface roughness, have
shown that a sufficiently thick film can be developed [28].
This is known as microelastohydrodynamic lubrication. De-
formation also causes fluid flow across the cartilage surface,
which modifies the film thickness, although there is some
question as to the practical importance of flow across the sur-
face [22,23,28].
MODELS OF OSTEOARTHROSIS
Animal models are used to provide a controlled environment
for studying the progression of osteoarthrosis. Although
osteoarthrosis may be induced by numerous means, models
based on disruption of the mechanical environment of
the joint, either by surgical alteration of periarticular struc-
tures or by abnormal joint load, are commonly used
[24,25,57,66,72,73,81].
Cartilage Modified single edge notch test
Trouser tear test
Cartilage
Force
Force
Bone
Figure 5.14: Sample shape and load application for the modified single-edge notch and trouser tear tests. Each test yields a specific measure of fracture, the energy required to propagate a crack in the material.
76 Part I | BIOMECHANICAL PRINCIPLES
Surgical resection of one or combinations of the anterior
cruciate ligament, the medial collateral ligament, and a par-
tial medial meniscectomy produce osteoarthrosis of the knee.
These models are thought to produce an unstable joint, but
kinematic studies show varying degrees of deviation from
normal joint kinematics.
Small differences in kinematics between control and op-
erated knees (anterior cruciate ligament release and partial
medial meniscectomy) are reported in rabbit [49]. At 4 weeks
after surgery, there is a statistically significant change in the
maximum anterior displacement of the knee, but anterior dis-
placement is not significantly different from normal at 8 or
12 weeks after surgery. The most notable kinematic changes
are in external rotation at 8 weeks and adduction at 4, 8, and
12 weeks after surgery. In dog, which has a more extended
knee, greater anterior-posterior drawer is found after anterior
(cranial) cruciate ligament release [36,88]. The relatively small
changes in kinematics in unstable joints (particularly in rab-
bit) suggests that altered forces and possibly sensory input
may be more important than joint displacements in the de-
velopment of osteoarthrosis [29].
Repetitive impulse loading also produces osteoarthrosis in
animal joints [70,72,73,81]. An advantage of this model is that
it is more controlled than surgical models; the force applied
to the limb is known and can be altered. This model has
demonstrated the effect of loading rate on the development
of osteoarthrosis. Impulsively applied loads were found to pro-
duce osteoarthrosis, while higher loads applied at a lower rate
do not. The importance of impulsive loading to the develop-
ment of osteoarthrosis also appears in humans; persons with
knee pain, but no history to suggest its origin, load their legs
more rapidly at heel strike than persons without knee pain.
Although biochemical, metabolic, and mechanical assays
have been used to evaluate the properties of cartilage from
animal models of osteoarthrosis, this chapter concentrates
on the mechanical properties of cartilage. Following resec-
tion of the anterior cruciate ligament in dog, tensile stiff-
ness, aggregate modulus, and shear modulus are lower than
those in cartilage from unoperated control joints [79]. Per-
meability increases significantly 12 weeks after surgery.
There is a significant increase in water content of samples
from the medial tibial plateau and the lateral condyle and
femoral groove.
In summary, various mechanical alterations of a joint lead
to the development of osteoarthrosis. The kinematic instabil-
ity induced by surgical alterations may be small, suggesting
that altered forces are primarily responsible for the develop-
ing osteoarthrosis. Models based solely on abnormal joint
loading support the view that alterations in force can lead to
osteoarthrosis. Following resection of the anterior cruciate lig-
ament, cartilage is less stiff in both compression and shear,
and fluid flows more easily through the tissue in joints with
osteoarthrosis. This implies greater displacement of os-
teoarthrotic cartilage than normal (decreased stiffness) and a
greater rate of deformation (increased permeability).
CLINICAL RELEVANCE: OSTEOARTHROSIS
Osteoarthrosis is a leading cause of disability in devel- oped countries [10]. In the United States, it is second to cardiovascular disease as the most common cause of dis- ability [63]. Despite the widespread occurrence of osteo- arthrosis, it is difficult to study in human populations. Early physical symptoms such as fibrillation and cracking of the articular surface cannot be detected by an individ- ual, since cartilage is aneural. Insults to the cartilage may take years to progress to the point where symptoms are detected by the surrounding joint structures and underly- ing bone. Although numerous epidemiological studies of osteoarthrosis have been performed, they have been described as “disappointing,” since they have not lead to an explanation of the mechanisms underlying the devel- opment of osteoarthrosis [63]. However, what seems to be clear is that the development of osteoarthrosis depends on a combination of factors including age, sex, heredity, joint mechanics, and cartilage biology and bio- chemistry [11,46,55]. Although it is not an inescapable consequence of aging, osteoarthrosis is more prevalent in the elderly [62,64]. In the United States, approximately 80% of people over the age of 65 and essentially everyone over the age of 80 has osteoarthrosis, although it is uncommon before the age of
40. After 55 years of age, osteoarthrosis is more common in women than in men. Typically the interphalangeal, first carpometacarpal and knees are the first joints that are affected [62]. However, specific links between aging and osteoarthrosis are not known. Excessive mechanical load- ing may also predispose joints to osteoarthrosis. Some studies have shown workers in physically strenuous occupa- tions (coal miners) have a higher incidence of osteoarthrosis than those in less strenuous lines of work (office workers) [63]. Interestingly, osteoarthrosis of the shoulder and elbow have been found in relatively young individuals in ancient populations who depended on hunting [63]. However, strenuous work may not be the only risk factor for osteoarthrosis, since persons who use pneumatic drills or physical education teachers do not have an increased risk of osteoarthrosis [63]. Obesity has also been found to increase the risk of osteoarthrosis, particularly in the tibiofemoral, patellofem- oral, and carpometacarpal joints [10]. Although increased weight would be expected to increase the load on joints of the lower extremity and possibly predispose an individual to osteoarthrosis, obesity would have no direct mechanical effect on the carpometacarpal joint. Injuries to the anterior cruciate ligament, collateral ligament, or meniscus have been implicated in the devel- opment of osteoarthrosis in the knee [39]. Loss of the anterior cruciate ligament may impair sensory function and protective mechanisms at the knee. Disruption of internal joint structures may alter joint alignment and the areas of
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