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Spinal Facet Joint Biomechanics and Mechanotransduction in Normal, Injury and Degenerative Conditions


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The facet joint is a crucial anatomic region of the spine owing to its biomechanical role in facilitating articulation of the vertebrae of the spinal column. It is a diarthrodial joint with opposing articular cartilage surfaces that provide a low friction environment and a ligamentous capsule that encloses the joint space. Together with the disc, the bilateral facet joints transfer loads and guide and constrain motions in the spine due to their geometry and mechanical function. Although a great deal of research has focused on defining the biomechanics of the spine and the form and function of the disc, the facet joint has only recently become the focus of experimental, computational and clinical studies. This mechanical behavior ensures the normal health and function of the spine during physiologic loading but can also lead to its dysfunction when the tissues of the facet joint are altered either by injury, degeneration or as a result of surgical modification of the spine. The anatomical, biomechanical and physiological characteristics of the facet joints in the cervical and lumbar spines have become the focus of increased attention recently with the advent of surgical procedures of the spine, such as disc repair and replacement, which may impact facet responses. Accordingly, this review summarizes the relevant anatomy and biomechanics of the facet joint and the individual tissues that comprise it. In order to better understand the physiological implications of tissue loading in all conditions, a review of mechanotransduction pathways in the cartilage, ligament and bone is also presented ranging from the tissue-level scale to cellular modifications. With this context, experimental studies are summarized as they relate to the most common modifications that alter the biomechanics and health of the spine-injury and degeneration. In addition, many computational and finite element models have been developed that enable more-detailed and specific investigations of the facet joint and its tissues than are provided by experimental approaches and also that expand their utility for the field of biomechanics. These are also reviewed to provide a more complete summary of the current knowledge of facet joint mechanics. Overall, the goal of this review is to present a comprehensive review of the breadth and depth of knowledge regarding the mechanical and adaptive responses of the facet joint and its tissues across a variety of relevant size scales.
Content may be subject to copyright.
Nicolas V. Jaumard
William C. Welch
Dept. of Neurosurgery,
University of Pennsylvania,
HUP - 3 Silverstein,
3400 Spruce Street,
Philadelphia, PA 19104
Beth A. Winkelstein
Dept. of Neurosurgery,
University of Pennsylvania,
HUP - 3 Silverstein,
3400 Spruce Street,
Philadelphia, PA 19104;
Dept. of Bioengineering,
University of Pennsylvania,
210 S. 33rd Street,
Room 240 Skirkanich Hall,
Philadelphia, PA 19104
Spinal Facet Joint Biomechanics
and Mechanotransduction in
Normal, Injury and Degenerative
The facet joint is a crucial anatomic region of the spine owing to its biomechanical role
in facilitating articulation of the vertebrae of the spinal column. It is a diarthrodial joint
with opposing articular cartilage surfaces that provide a low friction environment and a
ligamentous capsule that encloses the joint space. Together with the disc, the bilateral
facet joints transfer loads and guide and constrain motions in the spine due to their ge-
ometry and mechanical function. Although a great deal of research has focused on defin-
ing the biomechanics of the spine and the form and function of the disc, the facet joint
has only recently become the focus of experimental, computational and clinical studies.
This mechanical behavior ensures the normal health and function of the spine during
physiologic loading but can also lead to its dysfunction when the tissues of the facet joint
are altered either by injury, degeneration or as a result of surgical modification of the
spine. The anatomical, biomechanical and physiological characteristics of the facet joints
in the cervical and lumbar spines have become the focus of increased attention recently
with the advent of surgical procedures of the spine, such as disc repair and replacement,
which may impact facet responses. Accordingly, this review summarizes the relevant
anatomy and biomechanics of the facet joint and the individual tissues that comprise it.
In order to better understand the physiological implications of tissue loading in all condi-
tions, a review of mechanotransduction pathways in the cartilage, ligament and bone is
also presented ranging from the tissue-level scale to cellular modifications. With this con-
text, experimental studies are summarized as they relate to the most common modifica-
tions that alter the biomechanics and health of the spine—injury and degeneration. In
addition, many computational and finite element models have been developed that enable
more-detailed and specific investigations of the facet joint and its tissues than are pro-
vided by experimental approaches and also that expand their utility for the field of biome-
chanics. These are also reviewed to provide a more complete summary of the current
knowledge of facet joint mechanics. Overall, the goal of this review is to present a com-
prehensive review of the breadth and depth of knowledge regarding the mechanical and
adaptive responses of the facet joint and its tissues across a variety of relevant size
scales. [DOI: 10.1115/1.4004493]
Keywords: spine, facet joint mechanics, mechanotransduction, articular cartilage,
capsule, biomechanics
1 Introduction
The zygapophyseal, or facet, joints are complicated biomechan-
ical structures in the spine, with complex anatomy, mechanical
performance and effects on overall spine behavior and health. At
each spinal level, there is a pair of facet joints located on the
postero-lateral aspects of each motion segment, spanning from the
cervical to the lumbar spine (Fig. 1). These facet joints are typical
diarthrodial joints with cartilage surfaces that provide a low-fric-
tion interface to facilitate motion during normal conditions in a
healthy spine. Owing to the anatomy of the spine, the mechanical
behavior of the facet joint is both dependent on the responses dic-
tated by the overall spine’s response and also can directly affect
the spine’s response, via its relationship to the intervertebral disc,
its anatomic orientation, and its own mechanical behavior. The ki-
nematics and mechanical properties of the facet joint and its tissue
components have been studied extensively for a variety of differ-
ent loading conditions [111]. Recently, there is growing interest
in the facet joint—its biomechanics and physiology—with the
advent of disc arthroplasty and there has been increased attention
to the relationship between spinal degeneration and its effects on
the mechanical environment of the different tissues in the facet
joint [1216]. Therefore, it is the primary goal of this review to
present an updated perspective of the anatomy and global
mechanics of the spinal facet joint and its individual tissue com-
ponents in conjunction with their loading during physiologic and
nonphysiologic motion. In addition, this review will summarize
the mechanotransduction processes by which mechanical loading
to the specific tissues of the joint translate into signals that drive
physiologic responses in health, injury and trauma, and spinal
degeneration. Computational models of the facet joint are also
reviewed since there has been quite a bit of work in this area to
complement and expand findings from biomechanical experi-
ments and to provide insight about facet joint mechanics other-
wise not measureable in typical cadaveric studies. Overall, this
review focuses on synthesizing this anatomical, biomechanical
and physiological information to give an overview of the facet
joint’s response to mechanical loading from the macroscopic to
the cellular scale, with implications and perspective for future
studies of this spinal joint.
Corresponding author.
Contributed by the Bioengineering Division of ASME for publication in the
JOURNAL OF BIOMECHANICAL ENGINEERING. Manuscript received February 12, 2011;
final manuscript received June 21, 2011; published online August 2, 2011. Editor:
Michael Sacks.
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C2011 by ASME
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2 Anatomy and Tissue Mechanics
The facet joints, together with the intervertebral discs and spi-
nal ligaments, connect the adjacent vertebrae of the spine at all
regions and provide support for the transfer and constraint of
loads applied to the spinal column. These articulations insure the
mechanical stability and also overall mobility of the spine, while
protecting the spinal cord running through it. At each spinal
level, the bilateral facet joints are positioned symmetrically rela-
tive to the mid-sagittal plane in the postero-lateral regions of the
spine (Fig. 1(a)). Because the facet is a diarthrodial synovial
joint, cartilage covers the sliding surfaces and ligamentous cap-
sules guide, couple, and limit the relative translations and rota-
tions of adjacent vertebrae. Broadly, the facet joint is made up of
a variety of hard and soft tissues: the bony articular pillars of the
lateral mass provide the opposing surfaces that are covered by
cartilage, the synovium which is a connective tissue lining that
maintains lubrication for the articular surfaces and enables their
frictionless motion, and a ligamentous capsule that envelops the
entire joint [1720]. The bony articular pillars support compres-
sive loads and the facet capsular ligament resists tensile forces
that are developed across the joint when it undergoes rotations
and translations [1,6,21]. Together, this collection of tissues
functions to transfer the different loads across the joint during a
variety of loading modes for the spine. Here, we provide a more
detailed presentation of the facet anatomy in order to describe
the response to mechanical loading for each of the soft and hard
tissues composing the facet joint.
2.1 Bony Articular Pillars. The articular pillars are the bony
protuberances that extend superiorly and inferiorly from the lam-
ina of each vertebra along the long-axis of the spine (Fig. 1(a)).
They are located at the junction between the lamina and the lateral
masses in the cervical region of the spine; whereas, in the thoracic
and lumbar regions, they are joined to the vertebral body via the
bony pedicles. At each intervertebral joint along the spine, the ad-
jacent articular pillars are aligned to establish two postero-lateral
columns that provide mechanical support for axial loading along
the spine, together with the anterior column comprised of the ver-
tebral bodies joined by their interconnected intervertebral discs
[22,23]. In general, the inclination angle of the articular surfaces
of the facet joint in the sagittal plane ranges from 20–78in the
cervical region, 55–80in the thoracic region, and 82–86in the
lumbar region (angle bin Fig. 1(a)). The angle between the articu-
lating surfaces in the axial plane range from 70–96,85
and 15–70off of the midline in the cervical, thoracic, and lum-
bar regions, respectively (angle ain Fig. 1(b)), with increasing
orientation angles moving towards the lower levels in the lumbar
spine [2427]. Lastly, the superior articular surfaces transition
from having a postero-medial orientation in the cervical region to
a more postero-lateral orientation in the thoracic region, although
asymmetrical orientations have also been reported [26].
The facet joint is formed by two adjacent vertebrae with the in-
ferior facet of the superior vertebra meeting the superior facet of
the inferior vertebra (Fig. 1(a)). As such, each articular pillar of a
vertebra has both a superior and an inferior articulating surface.
The surfaces of the pillars that form the articulation of the joint
have elliptically-shaped faces that are covered by cartilage (Fig.
2). The morphometry of these surfaces also differs between the
regions of the spine, as well as at each vertebral level [26,2832].
The superior facet of the inferior vertebra is rather flat in the cer-
vical and thoracic regions and more convex in the lumbar region
[26]. The opposing inferior facet of the superior vertebra is con-
cave and forms an arch with its apex pointing towards the verte-
bral body [20,3336] (Fig. 2). Articular surfaces are more
horizontally-oriented in the cervical and upper thoracic spinal
regions [26,36], which enables the great degree of coupling of
axial rotation and lateral bending that exists in the cervical spine
[3739]. In the lower thoracic and lumbar regions of the spine the
facets gradually become more vertically-oriented [25], which also
limits the flexibility of the spine in both lateral bending and rota-
tion in these regions. But, this decrease in flexibility protects the
intervertebral discs and spinal cord from nonphysiological kine-
matic and kinetic exposures that could cause injury and/or create
pathological conditions [6].
2.2 Cartilaginous Articular Surfaces. An avascular layer of
hyaline cartilage, with varying thickness across spinal regions and
the genders, covers the articulating surfaces of each facet [19,40].
The cartilage is thinner at the edges of the opposing surfaces and
gradually increases to its thickest (1 mm) towards the center of
the articulating joint, in both the antero-posterior and medio-lat-
eral regions of the joint [41]. Based on experimental studies, the
thickness of cervical facet cartilage has been described to have a
half sinusoidal shape with a maximum thickness (t
) at its cen-
ter and thinning out along its radius (r) towards the facet perimeter
), according to Eq. (1) [41]:
t¼tmax coskr
 (1)
where both the maximum thickness and the shape coefficient (k,
ranging 0.38–0.63) were both determined by minimizing the dif-
ference between the experimental and theoretical thickness distri-
butions [41].
Further, reports have found that the bony extremity of the pillars
is not always completely covered by a cartilage layer, leaving a
region of exposed subchondral bone at the outermost edges of the
bony pillar [19,41]. Yoganandan et al. [19] reported the gap of
Fig. 1 Lateral view of a cervical (a) and axial view of a lumbar (b) vertebra showing the overall
anatomy and the facet joints, articulations, and orientation relative to its angle with each of the
axial plane (b) and of the sagittal plane (a)
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exposed subchondral bone to be nearly three times wider in the
upper region of the cervical spine than in the lower cervical spine,
especially in the posterior and anterior regions of the facet articu-
lating surfaces. The cartilage layer may be thinner in these regions
because of the presence of synovial folds and meniscoids, which
also provide additional protection from compressive and shear
loads across the joint (see Sec. 2.1 for more details) [19]. However,
subchondral bone can be exposed to mechanical compression dur-
ing some loading scenarios. This not only presents the possibility
for direct trauma to the bone but has also been hypothesized to
lead to pain in some cases [42]. The gap in cartilage coverage is
greater in females [19], which may play a role in the greater sus-
ceptibility of women to suffer neck traumas [43,44].
2.2.1 Cartilage Composition. The cartilaginous layers cover-
ing the articulating facet surfaces enable frictionless motion
between the adjacent vertebrae, while also bearing compressive,
tensile, and shear loads. Such mechanical capabilities are due to
the specific structure of the cartilage tissue and the mechanical
properties of the matrix of the cartilaginous layer. The cartilage
matrix consists of collagen fibers, glycosaminoglycans (GAGs),
proteoglycans, and chondrocytes [4547]. This matrix is actually a
laminate composed of three main zones along the depth of the car-
tilage, with the outermost surface (lamina splendens) being the
region where contact with the articulating surface of the opposing
pillar occurs and the innermost region being at the subchondral
bone of the articular pillar (Fig. 2). Each cartilage zone has a
different structural organization, as well as variable and specific
mechanical properties [4749]. The superficial zone contains rela-
tively few flattened chondrocytes and collagen fibers that are ori-
ented tangentially to the surface of the cartilage; the horizontal
fiber alignment provides resistance to both the tensile and shear
stresses that develop during the relative sliding developed in the
joint between the opposing articular surfaces when the spine bends
or rotates [47,48]. In the middle or transitional zone, more chon-
drocytes are interspersed between larger collagen fibers with a
pseudorandom arrangement [4850]. Finally, in the deep zone,
chondrocytes align in columns perpendicular to the articular sur-
face and parallel to the collagen fibers. In the transitional and deep
zones, the concentration in proteoglycans embedded in the chon-
drocytes and collagen structure increases with the depth. These
proteoglycans trap water and increase the incompressibility of the
structure, thereby supporting compressive and hydrostatic stresses
that may be developed in the joint. At the bottom of the deep zone
a tidemark separates the deep zone from a zone of calcified carti-
lage that transition into subchondral bone, making the change in
elastic modulus from cartilage to bone more gradual [4649].
2.2.2 Cartilage Mechanics. As with articular cartilage of any
diarthrodial joint (knee, hip, facet joint), the specific mechanical
properties are heavily dependent on the cartilage composition
(water content, collagen fiber orientations), the specimen’s age
and relative health, and the specific loading conditions that the
joint undergoes [5153]. Although the facet joint cartilage has
been described macroscopically, there have been little-to-no spe-
cific investigations of its mechanical responses. However, there
have been extensive reports for cartilage of the knee; those are
provided briefly here, as they are relevant to the broader context
of joint cartilage. The equilibrium modulus (H) of human patella
samples subjected to confined compression has been reported to
decrease linearly with age, structural disorganization (I), and
water content (WAT), according to the set of linear equations
(Eqs. (2),(3), and (4)) that were optimized to fit experimental data
[51]. According to those relationships, the equilibrium modulus
was found to depend more on the water content than with the
other two parameters, which is expected since water is incompres-
sible and a greater volume of it retained in the cartilage structure
would stiffen it.
H¼1:04 0:0045 ðageÞ(2)
H¼0:95 0:065 ðIÞ(3)
H¼5:29 0:058 ðWATÞ(4)
For example, a 30–40% decrease in water content leads to a 161%
increase in the equilibrium elastic modulus of cartilage explants
Fig. 2 Schematic drawings of the facet joint and the primary tissues that compose it, as
well as the cartilage and menisci of the facet articulation. The blowup illustrates the differ-
ent zones of the articular cartilage layer with the collagen fibers and chondrocytes orien-
tations through its depth. A cut through of the facet joint (A-A) is also drawn to show the
elliptically-shaped inter-articular surfaces with the cartilage surface on the inferior facet,
the synovium, and meniscoids. Adapted collectively from Martin et al., 1998, Pierce et al.,
2009, and Bogduk and Engel, 1984 [48,49,73].
Journal of Biomechanical Engineering JULY 2011, Vol. 133 / 071010-3
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subjected to axial compression [52]. The stress-strain relationship
is anisotropic as it depends on the dimensions and organization of
the collagen fibers, and the cellular and proteoglycans content that
differ across the depth of the cartilaginous layer. In addition, artic-
ular cartilage deformation results from a reorganization of its colla-
gen structure and loss of fluid during loading. Fluid loss is a much
slower process than the polymer network re-arrangement and so an
initial deformation occurs first without any volume change, and a
second deformation then results from a change in volume due to
fluid loss which produces a nonlinear load-displacement response
that is exhibited during unconfined compression [54,55]. This type
of behavior highlights the biphasic and time-dependent mechanical
properties of articular cartilage in diarthrodial joints [56,57].
Accordingly, creep studies have also demonstrated that the time
constant (T) of cartilage to reach equilibrium under maintained
compressive loading is a function of the thickness (h), the equilib-
rium modulus (E), and specific properties of the cartilage, as well
as the applied load (p)[51,56]. This equilibrium time constant also
depends on porosity (U
), permeability (k), and the drag coefficient
(K) as described in Eq. (5):
In contrast, because of the interactions between the collagen fibers
and the proteoglycans during uniaxial confined compression, the
relationship between axial and radial stresses (r) and axial strain
(e) is defined by a linear isotropic constitutive relation (Eqs. (6)
and (7))[58]. In this relationship, the compression axial modulus
) and the chemical stress (b) imposed by the surrounding mi-
lieu, as well as the Lame´ constants (kand G), all depend on the
concentration (c) of the environment.
Articular cartilage is a composite material composed of fluid
(water) and solid (chondrocytes, collagen, proteoglycans) phases
that has anisotropic nonlinear mechanical properties and load-
bearing capacity [59]. The difference in response time of the two
phases contained in articular cartilage makes its mechanical
response dependent on the rate of loading. The dynamic stiffness
of the cartilage lining in diarthrodial joints increases with strain
rate [6062]. For example, in a study of cartilage impacts during
knee graft implantation, fissures in the cartilage matrix were pro-
duced for both single high energy impacts (over 25 MPa) and re-
petitive impacts (26–35 MPa) across a variety of human, bovine,
and porcine species [60,63]. Chondrocyte viability was also
reduced by up to 60% for impacts of 1 J compared to a 5% and
20% decrease in cell viability for impacts of 0.25 J and 0.5 J,
respectively [52,62,63]. Fissures in the articular surface can allow
the enzymes that are contained in the synovial fluid in the joint,
such as collagenase and hyaluronidase, to penetrate and break
down the cartilage matrix [48]. At the same time, an increase in
chondrocyte death can also impair the subsequent synthesis of car-
tilage proteins that are required for the proper maintenance of the
avascular cartilage matrix [63]. A damaged cartilage matrix can-
not effectively support compressive loads, distribute pressure, or
resist stresses because fissures can penetrate as far as the transition
zone and disrupt the matrix structure [62]. In addition, the repair
of the cartilaginous matrix and its functionality are compromised
by the death of the chondrocytes because the production of mole-
cules imperative for matrix regeneration is reduced, which is then
followed by the denaturation of the collagen fibers and the release
of proteoglycans, which are needed to retain water and to provide
compressibility for the damaged cartilage structure [48,63].
Impact(s) on the articular cartilage can; therefore, cause signifi-
cant loss of mechanical properties and cellular damage which also
may provide the stimulus for the onset of degeneration in that tis-
sue and/or can also accelerate it [52,64,65]. However, the energy
transferred to the facet joint cartilage during physiologic and/or
nonphysiologic loading of the spine remains to be measured.
Explicit experimental studies of facet joint cartilage are limited.
Currently, there is only one investigation of canine lumbar facet
cartilage, reporting its aggregate modulus to be 554 kPa at equilib-
rium after an indentation with a 1 mm flat-ended porous-tip [66].
That study also found that the aggregate modulus of cartilage from
the facet joint was similar to the modulus of cartilage from other
canine diarthrodial joints such as the knee lateral condyle, patellar
groove, and shoulder, suggesting that the similarities between
human and canine articular cartilage could also extend to facet joint
cartilage [66]. It was also reported that human cartilage from the
knee and the hip has a compressive stiffness comparable to that of
the distal femur in canine models and the proximal femur in baboon
models, respectively [67,68], which suggests that the mechanical
properties of articular cartilage may be similar among any diarthro-
dial joints in the body. However, further characterization of human
facet joint tissue is needed to verify if the mechanical properties of
facet joint cartilage are similar across species as well.
2.3 Synovium, Menisci and Capsular Ligament. Extending
from the superior to the inferior articular pillar, two superposed
membranes, the synovium and the ligamentous capsule, maintain
the articular surfaces in a low-friction environment and provide
mechanical resistance to their separation and relative motion. The
synovium of the facet joint is a thin and soft periarticular connec-
tive tissue [17] with two main layers that secrete synovial fluid
components involved in the maintenance of the synovial fluid
used to lubricate and nourish the cartilaginous articular surfaces
[69,70]. The synovium also regulates the exchanges between the
blood and synovial fluid, and contains macrophage cells that
phagocytose cell debris and waste contained in the joint cavity
[70]. Although the functional role of this structure has been inves-
tigated at the cellular level [70,71], it has not been investigated
mechanically, most likely because it is difficult to isolate since it
is very thin and its outermost layer is intimately merged to the
inner surface of the capsular ligament [71]. For the same reasons,
and also because the innermost synovial layer is loose, the syno-
vium also likely does not play a substantial role in the mechanical
behavior of the joint as a whole. Although, the synovial mem-
brane is very thin, its loose innermost layer bulges into the joint
cavity in some areas, forming folds that wedge between the
opposing articular surfaces of the facet joint [7274].
The synovial folds, or meniscoids or menisci, are intra-articular
structures that protect the articular cartilage when opposing articu-
lating surface glide on each other during joint motion [75]. This
protection is realized since the meniscoids compensate for the
incongruence of the joint’s articular surfaces, guiding and smooth-
ing their relative motion, and distributing the load over a greater
surface area [72,76,77]. Three main types of menisci have been
identified in the facet joints across all of the regions of the spine:
adipose tissue pads, fibro-adipose meniscoids, and connective tis-
sue rims (Fig. 2). The adipose pads and meniscoids are located
mainly at the periphery of the articular surface in the anterior and
posterior region of the joint, where they only partially extend cir-
cumferentially along the rim of the articular pillar. These tissues
are crescent-shaped and have a triangular cross-section in the sag-
ittal plane (Fig. 2), with the base being attached to the capsule and
the point extending as much as 9 mm inward toward the interior
of the joint [72,75,78]. The connective rims of the synovial tissue
are ring-shaped, wraparound the edge of the bony pillar, and are
tapered inward towards the center of the joint [72,74,75,77,78].
The meniscoids are composed of fat, fibrous connective tissue
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and/or a mix of fat and fibers covered by a cellular synovial lining
[7275,77,79]. Although the meniscoids are known to cover the
gap of exposed subchondral bone at the articular surface in order
to reduce friction during articular motion, their mechanical role is
still unclear [77]. They have been speculated to moderate the load
transferred to the cartilage when the articular surfaces engage in
compression during any joint motion by distributing the pressure
as they move freely in and out of the inter-articular space during
motion [72,7577,80]. This putative function is probably linked to
the meniscus entrapment theory that was developed to explain
how low back pain symptoms could be caused, and then treated
by simple manipulation [73,81]. Although this may be the case
under normal loading, these intra-articular folds can become torn
at their base under combined substantial compression and shear
loading which can also lead to subcapsular hemorrhage and
entrapment of the torn pieces in the joint, eventually inducing fur-
ther physiologic dysfunction and even pain [73,74,82]. Although
postmortem and in vivo MRI studies provide evidence of the pres-
ence of meniscoids in the spinal facet joints and help to character-
ize their dimensions and composition, the role of these structures
in the biomechanical behavior of the whole facet joint remains
As in the other joints in the body, such as the knee and the hip,
the facet capsular ligament covers the synovium to fully enclose
the facet joint, enveloping it in the superior-inferior direction and
with nonuniform thickness. For instance, the lumbar facet capsule
has been reported to be 2.0 mm thick in the posterior region, and
as much as 3.2 mm thick in the anterior region, whereas the supe-
rior and inferior regions are approximately 2.4 mm thick [83].The
capsular ligament is comprised of dense collagen fiber bundles
linked by proteoglycans, with elastin fibers and fibroblasts inter-
spersed [18,84]. The collagen and elastin fibers extend between
the laminae of adjacent vertebrae connecting to the ligamentum
flavum both in the antero- and postero-medial regions of the facet
joint, and completely surrounding the joint’s articular surfaces in
three dimensions. The collagen fibers are oriented differently
along the superior-inferior axis of the capsular ligament [84] and
they are crimped [18]. The crimped collagen fibers allow the cap-
sule to undergo substantial excursions without reaching its me-
chanical limit or inducing local injury. Under load, the fibers can
become uncrimped which allows the overall joint to translate and
rotate without offering any mechanical resistance.
The capsule, as well as the subchondral bone, synovium and
folds, are richly innervated with mechanoreceptive, proprioceptive
and nociceptive nerve endings [21,8588].Therefore, mechanical
loading of any of those innervated tissues in the facet joint could
activate nerve endings and modulate the signals in the nervous sys-
tem to initiate the development and maintenance of pain and/or
cellular dysfunction. The nervous system is also involved in modu-
lating the overall mechanical response of the facet joint and its tis-
sues since the intensity and frequency of the mechanical stimuli
experienced by these nerve endings also provide feedback to the
central nervous system which is used to adjust the activity of the
surrounding muscles and correct the loading of the joint in real-
time [89–91].
3 Facet Joint Macromechanics
3.1 Facet Joint and Spinal Stability. The facet joints guide
and constrain the motion of the vertebrae, while also facilitating
the transmission of the loads applied to the spine [2,21,66,92].
The facet joints also contribute to and help maintain the stability
of the spine. A structural column, like the spine, is considered
mechanically stable when the sum of the forces and moments
applied to it equals zero. Mechanical stability of the spine is
achieved when the paraspinal musculature effectively counteracts
the external loads via modification of the shape of the vertebral
column. Clinically, the term ‘spinal stability’ has taken on the
definition of the spine’s ability to maintain its alignment and to
provide protection to the neural structures it encloses during
physiologic loading [93]. The clinical assessment of spinal stabil-
ity/instability is required for a variety of different clinical scenar-
ios, including degeneration with altered kyphosis or lordosis,
surgical management or when motions become painful. The
assessment is performed using imaging to measure the relative
position of the vertebrae, and to detect any malalignment
White and Panjabi [24] defined clinical instability of the spine
as the spine’s loss of ability to maintain its normal motions under
physiologic loads which leads to initial or additional neurologic
deficit [24]. Although most clinicians agree on the clinical defini-
tion of spinal instability, there is still ambiguity in using the term
“spinal stability” because its quantitative assessment remains
challenging and subjective in the clinical setting [93,95]. Cur-
rently, clinicians consider the spine as a three-column system in
their assessment of spinal instability with the first column con-
taining the anterior longitudinal ligament and the anterior half of
the body and discs, the middle column contains the posterior half
of the vertebral body and disc, and the posterior column contains
the interspinous ligaments, spinous processes, pedicles, and the
facets [93] (Fig. 1). The spine is considered unstable when two
of the three columns are not intact. This rule is substantiated by
the more complex system implemented by White and Panjabi
[24] in which the spine is judged unstable when translations are
greater than 3.5 mm and rotations greater than 20 degrees in the
sagittal plane during flexion-extension bending [24]. Although
injured or damaged facet joints do not a priori dictate that the
spine is mechanically unstable, the proprioceptive and nocicep-
tive nerve endings in the facet joint can respond to overload,
damage, or injury to alter the musculature feedback and control
for providing support to the spinal column. Moreover, injured
nerves can also become nonresponsive to loading or motion or
exhibit dysfunctional performance, both of which can result in
abnormal sensory feedback for the central nervous system’s
coordination of the various spinal tissues and paraspinal muscles
to insure mechanical stability [94].
3.2 Mechanical Contributions of the Facet Joint. The role
of the facet joints in the mechanical stability of the spine has been
established from biomechanical and mathematical studies. The
facet joints prevent two adjacent vertebrae from engaging in rela-
tive motions that could overload and damage the surrounding spi-
nal structures, such as the intervertebral disc, the nerve roots that
exit the spinal column, and the spinal cord. Consequently, the facet
joint tissues are themselves mechanically loaded. For example,
Yang and King [2] reported that between 75–97% of the compres-
sive load applied to the lumbar spine is borne by the intervertebral
discs, and they estimated that 3–25% is carried by the posterior
elements of the vertebral column in what they referred to as “facet
force” [2]. In similar experiments using lumbar motion segments,
Adams and Hutton [96] measured that under 2 degrees of exten-
sion and 560–1030 N of compression, 16% of the load is borne by
the facet joints) [96]. Pal and Routal [4] assumed the spine to be
mechanically equivalent to three columns; an anterior column
composed of the vertebral bodies and discs, and two posterior col-
umns consisting of vertically-connected articular processes. Those
authors considered that any compressive load applied to the spine
was distributed over the whole vertebral body and areas of the
entire facet joints and that the ratio of the articular facet area to
vertebral body area could be used as a metric of the load-sharing
between the anterior and posterior columns [4]. Using an analysis
of detailed facet joint morphology (facet articular area, vertebral
body horizontal cross-section area, lordosis angle) Pal and Routal
[4] computed that 23% of any axial compressive load is transmit-
ted by the facet joints in the cervical and upper thoracic regions of
the spine [4]. They reached the same conclusion in a matched
study using the lower thoracic and lumbar regions of the spine, in
which their anatomical observations and cross-sectional measure-
ments of the vertebrae showed that the posterior vertebral elements
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were actually connected over a single larger area formed by the
stacking of laminae instead of the two smaller areas formed by the
articular pillars as in the cervical spine [5].
In addition to transmitting compressive loads along the spine,
facet joints also provide torsional stiffness, and resistance to shear,
lateral and antero-posterior vertebral translation, and joint distrac-
tion [24,97]. The specific contribution of the facet joint in resist-
ing these mechanical scenarios has been most widely studied and
demonstrated in facetectomy studies, in which the facet joints are
surgically removed either in total or partially [98104]. For exam-
ple, the shear strength of cadaveric cervical motion segments was
shown to diminish by 29% after 70% of the facet joints were
removed bilaterally [98]. In a later investigation, Raynor et al.
[99] found that a partial bilateral facetectomy in which only 50%
of the cervical facet joints were removed significantly reduced
coupled motions. The lateral translation, axial rotation, and supe-
rior-inferior translation all decreased when a lateral bending
moment of 3.4 Nm was applied to the head. Also, both lateral
translation and rotation were smaller when a lateral force of 89 N
was applied to the head, as compared to the intact condition [99].
These changes in coupled motion following facetectomy led to
the conclusion that more force must be applied to reach the same
degree of neck motion. However, the degree of vertical distraction
in response to a tensile load increased after facetectomy, in com-
parison with the intact condition, suggesting there is greater risk
for facet dislocation when a tensile load is combined with a model
of loading that also opens the facet joint, like flexion, since less
force is required to further separate the facets in that combined
loading scenario [99].
3.3 Capsular Ligament Mechanical Properties. Because
the facet capsular ligament is composed of elastin and collagen
fibers, that structure can only support tensile and/or shear load-
ing. Accordingly, it can only provide mechanical resistance
when the vertebrae that it encloses undergo relative translations
and rotations (Fig. 1). As such, the capsule can contribute to lim-
iting the motions of the facet joint. This role was demonstrated
in studies using C3-C7 cadaveric spine segments subjected to
100 N of compression coupled with either 2 Nm of sagittal bend-
ing or 5 Nm of axial torsion, before and after graded bilateral re-
moval of the capsular ligament [105]. After a removal of 50% of
the facet capsule, axial rotation increased by 19% when torsion
was applied while the vertical distance between the C4 and C6
spinous processes increased by 5% under flexion [105]. The
extent of axial torsion and posterior displacement increased by
25% and 32%, respectively, after 75% of the facet capsule was
removed [105]. In a recent investigation using cadaveric cervical
motion segments we also measured a significant increase in rota-
tion in flexion after even a unilateral transection of the facet cap-
sule [106].Theincreaseintherangeofmotionobservedafter
capsule transection or removal supports that this ligament pro-
vides substantial contribution to constraining vertebral motion,
particularly in flexion and lateral bending or torsion when the
capsular fibers are stretched.
Strains are measured because they inform on the strength, stiff-
ness, and deformability of the capsule and also help to quantify
both the failure mechanisms and thresholds of the facet capsular
ligament. In particular, the tensile stiffness, ultimate tensile
strength, and failure strain of the facet capsular ligament have
been measured from isolated cervical and lumbar facet joints in
tension to estimate the risk of capsular injury during physiologic
bending [8] and also to characterize the anisotropic viscoelastic
properties of the ligament for inputs for computational models
[9,10]. Both Winkelstein et al. [8] and Yoganandan et al. [9]
reported average failure strains ranging from 100 to 150% for the
cervical capsular ligament. Little and Khalsa [10] subjected iso-
lated lumbar facet joints to tensile stretch in directions parallel
and perpendicular to the principal orientation of the collagen
fibers up to a strain of 50% of their measured length in order to
characterize the static and dynamic mechanical properties of the
capsule. An exponential strain-stress relationship was determined
for the capsules stretched in the direction parallel to the collagen
fibers, while a linear relationship was obtained for those loaded
in the direction perpendicular to the collagen fibers [10]. In these
relationships, represents the strain and rthe stress for the vis-
cous (V) and elastic (E) cases:
Parallel ðviscousÞ:rV¼0:0034 e10:35eðÞ (8)
Perpendicular ðviscousÞ:rV¼2:02 e0:1732 (9)
Parallel ðelasticÞ:rE¼0:0030 e10:09eðÞ (10)
Perpendicular ðelasticÞ:rE¼1:04 e0:097 (11)
Both Yoganandan et al. [9] and Little and Khalsa [10] reported a
nonlinear relationship between the strain and the stress with mod-
uli of the same magnitude (>10 MPa) although the capsules were
from cervical and lumbar spines, respectively (Table 1). The me-
chanical properties of the capsular ligament do not seem to vary
across the spinal regions despite different demands for their load-
ing throughout the different regions of the spine and their varied
anatomy and orientation of the facets in those regions. The similar
mechanical properties could indicate that the mechanical role of
the facet capsule is not specific to each spinal region and is sec-
ondary to that of the intervertebral disc and supplemented by the
other more-robust spinal ligaments in restricting vertebral
3.4 Regional Capsular Strains During Spinal Loading.
Capsular strains measured in human cervical and lumbar spinal
segments subjected to flexion, extension, lateral bending, and
axial rotation moments have been shown to not be uniform over
the entire capsule surface and also that they can reach very large
values without being associated with any macroscopic evidence of
tissue damage or failure [8,107,108]. Panjabi et al. [107] subjected
lumbar motion segments to 15 Nm moments about each of the
three anatomical axes and measured the strains in the capsule as
an average of the change in distance between five pairs of points
materializing the ligament superior and inferior attachments. The
strains measured outside of the neutral zone, when vertebral rota-
tions reached approximately 5 degrees, were up to 13% in flexion,
6% in extension, 8.7% in axial rotation, and 8.8% in lateral bend-
ing in the left and right facet capsules [107]. Capsular strains have
also been measured in cervical motion segments subjected to a 2.5
Nm sagittal moment with and without a 1 Nm axial pretorque but
defining full-field strains [8]. Using an array of 30 miniature beads
affixed to the lateral region of the right capsule, maximum princi-
pal strains were 12% in flexion and extension. Further, in that
study, when an axial pretorque was applied directed away from
the facet joint being studied for strain, the maximum principal
strain significantly increased to 23% in flexion and 17% in exten-
sion. For an axial pretorque towards the facet joint, strains
increased to 16% in flexion and to 13% in extension but were not
significantly greater than in the loading condition without pretor-
que [8]. In a companion study, flexibility tests on human cadav-
eric cervical motion segments found that the maximum capsular
strain under 135 N of posterior shear was independent of any com-
bined axial compressive loading, and stayed around 17%, with a
primary direction oriented along the antero-posterior axis under
combined shear, bending, and compression [109]. Under shear
loading of isolated facet joints, the maximum principal strain in
the capsular ligament reached 35 621% and 94 685%, corre-
sponding to the conditions when the facet joints underwent suffi-
cient loading to induce minor “subcatastrophic” and frank
“catastrophic” failures, respectively [109]. Using an array of 6–9
infrared markers implemented on the left and right capsules of the
lumbar vertebrae, Ianuzzi et al. [108] implemented the same
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approach as Winkelstein (2000) and Siegmund (2000) to measure
principal strains in the capsular ligament of T12-S1 lumbar speci-
mens during flexion, extension, and lateral bending [8,109]. The
maximum principal strains were generally smaller than those
measured in the cervical spine, reaching 15% in flexion, 8% in
extension, and 9% in lateral bending [108]. The differences
between the studies in terms of spinal region, number of vertebral
levels tested, and magnitude and application method of the
moment, do not permit a direct comparison of the capsular strains
reported (Table 1). Yet, the large capsular strains measured at fail-
ure during shear loading indicate that the facet capsule can elon-
gate significantly when it is loaded. However, the small capsular
strains reported for simple loading conditions such as pure flexion
or axial torsion, further demonstrate that the capsule is very strong
in resisting deformation and opposing vertebral rotation and trans-
lation. This ability also explains the significant increase in verte-
bral range of motion observed in the experimental studies
employing capsulotomy [105,106].
Table 1 Summary of experimental studies of the facet capsule material and mechanical properties
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Straining of the fibers in the facet capsule not only results
from vertebral motion but also from the activation of the muscles
that can occur during mechanical loading of the spine as the
outer surface of the capsular ligament is covered by the sur-
rounding paraspinal muscles [110]. As such, the individual fibers
of the capsule can become stretched when the muscles that insert
on it contract [83,88,110]. In fact, muscle insertions have been
found to cover nearly 23% of the capsule area in the cervical
spine with a nonuniform spatial distribution [110], which can
give rise to unequal capsular strains and stresses in all the
regions (posterior, anterior, lateral) of the capsule when different
muscles are activated to stabilize the spine during loading. An in-
homogeneous mechanical loading environment (either in direc-
tion and/or magnitude) of the capsular fibers can also occur due
to vertebral motion alone in the absence of any muscle activa-
tion. For instance, during flexion, the fibers of the posterior cap-
sule region are stretched while the fibers in the anterior region of
the capsule remain lax [8]; in contrast, during extension this pat-
tern is reversed. Variation in strains in the capsule has also been
observed in the facet capsule of the rat, in association with pain-
ful and nonpainful mechanical loading conditions imposed to the
vertebral bones. Upon facet distractions that are considered to be
physiologic, the strains reached 21 64% in the posterior region,
17 64% along the postero-lateral ridge, and 18 64% in the lat-
eral region of the C6-C7 facet capsule. Similar differences in
strains of the capsule regions were also reported for painful facet
joint distractions in that study [111]. Quantitative polarized light
imaging was used to measure fibers kinematics during tensile
loading of the capsule [111,112] and it was found that distrac-
tions of the joint that correspond to those producing pain also
produced a significantly greater fiber dispersion in the posterior
(23.0 64.9) than in the lateral (16.8 62.6) regions of the cap-
sule [111]. Therefore, stretching of the capsular fibers in each
region of the capsular ligament depends on the type of loading
and on the extent of muscle insertion in that area.
3.5 Capsular Stretch and Neural Activity. Strains in the
capsular ligament and stiffness have also been defined in animal
models in which a distraction was imposed across the facet joint
in the cervical spine in order to investigate pathomechanisms of
facet-based pain [113123]. In these approaches, an array of 25
to 35 miniature beads was placed on the exposed capsule to cal-
culate strains during cervical distraction. Specifically, Quinn
et al. [111] reported a maximum principal strain of 50% for a
700 lm subcatastrophic distraction of the rat C6-C7 facet joint
that did not tear the capsule but did produce sustained behavioral
hypersensitivity [111].Kallakurietal.[121] reported a strain of
73% at failure of the C5-C6 ligament that occurred between 12
and30mmtensilestretchinthegoat[121]. The capsular strain
of 50% measured in the rat model [111]isslightlylargerthan
the 35% strain measured at the first sign of tissue rupture in the
human capsule under shear [109] but compares well with the
analogous measurement of 65% strain for the human capsule
under tension [8]. Similarly, the tensile failure strain of 73%
reported for the goat model [121] compares well with the failure
strains of 94 685% and 104 681% reported by both Siegmund
and Winkelstein for the human capsule [8,109] (Table 1). The
similarity between the capsular strain values in the human and
animal specimens may be a reflection of their similar mechanical
function and composition.
When the capsule is stretched, the nerve afferents that inner-
vate it are also stretched, which has been shown to trigger the
generation of neuronal signaling to the central nervous system
(CNS) in cases of noxious stretch [113,116,117,123]. Lu et al.
[116] stretched the C5-C6 facet joint in a goat model and quanti-
fied capsule strains as well as the associated activation of affer-
ents from the joint. They found that the capsule contained
afferents that responded with firing at both low- and high-thresh-
olds of strain (10% and 47%, respectively) and also that afferents
of both types exhibited persistent generation of afterdischarge for
up to 5 mins after the release of the applied strain (39–57%) that
did not produce tissue rupture [116]. That work strongly impli-
cated afferent injury in the capsule as a possible mechanism of
pain because the afterdischarges were hypothesized as potentially
having long-term effects in the CNS. Using a rodent model, Lee
et al. [113] distracted the C6-C7 facet joint along the long-axis
of the spine and measured a three-fold increase in behavioral
hypersensitivity, as well as a significant sustained increase in
astrocytic activation in the spinal cord in the absence of any liga-
ment failure. Activated astrocytes modulate immune activation,
neuronal synapses and play a role in pain signaling [113]. Using
the same rodent model, we have found that after a high-rate facet
joint distraction, expression of a glutamate receptor is also ele-
vated in the spinal cord and positively correlated with both the
degree of strain in the capsule and the amount of behavioral sen-
sitivity [123]. Collectively, the integration of biomechanics with
physiological and behavioral outcomes in these in vivo studies
indicate that the loading environment of the afferents in the cap-
sular tissue may be responsible for signaling injury and dysfunc-
tion (i.e., pain) in that tissue of the facet joint. In fact, from that
combined work it has been suggested that the strain threshold for
sustained painful capsular distraction may be between 20 and
47% [113,116,117,123,124].
3.6 Facetectomy Alters the Motion Segment Mechanical
Response. Cusick et al. [100] reported that both unilateral and
bilateral cervical facetectomies produced a loss of strength by as
much as 32% and 53%, respectively. In those same cadaveric
studies, rotations increased by 18% and joint distraction
increased by 19% for application of combined compression-flex-
ion [100]. Zdeblick et al. [102] showed that progressive bilateral
facetectomy in multisegment cervical spine specimens subjected
to 100 N of compression and 5 Nm of torsion significantly
decreased torsional stiffness from 0.37 Nm/degree in an intact
specimen to nearly half (0.18 Nm/degree) after a complete C5-
C6 facetectomy [102]. When the specimens were subjected to 2
Nm of flexion they measured a 25% increase in the vertical dis-
tance between the C4 and C6 spinous processes after a 75% fac-
etectomy, which was not significant but did show an increase in
C4-C6 rotation [102]. Nowinski et al. [103] proceeded with a
similar graded facetectomy procedure on C2-C7 segments after a
C3-C6 laminoplasty had already been performed. Applying
moments of up to 1.5 Nm about all three axes, they measured an
increase of 7 degrees in sagittal rotation, 9 degrees in axial rota-
tion and 3 degrees in lateral rotation [103]. They also measured
an increase in translation but no significant change in coupled
motion, after 25% or more facetectomy, which is in disagree-
ment with the results reported by Raynor et al. [99].
In the lumbar spine, partial stepwise and total facetectomies
also significantly increase rotation in flexion and axial rotation in
motion segments loaded in compression (200 N) and subjected to
8 Nm about the three axes [101]. Tender et al. [104] resected the
L5 pars interarticularis followed by a total unilateral facet removal
on L5-S1 cadaveric motion segments subjected to 280 N of com-
pression and 7.5 Nm of axial torsion. They found that the unilat-
eral facetectomy significantly increased ipsilateral axial rotation
by 1.4 degrees and overall axial ROM by 3 degrees. The increase
in rotation, the loss of strength, and the decrease in stiffness in the
spinal motion segment following facetectomy demonstrate that
the facet joint contributes to spinal mechanical stability in a vari-
ety of directions and loading scenarios by limiting the linear and
rotational motions during physiological loading [104]. The restric-
tion of motion and the assurance of spinal stability provided by
the facet joint stem from the biomechanical properties of the cap-
sular ligament, articular cartilage, and bony pillars that together
facilitate the functions of the joint as a whole.
3.7 Cartilage Mechanical Properties. Since the capsule
provides support to help keep the facet joint intact during
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physiologic motions of the spine, the articular surfaces also
remain in contact during those normal conditions. During such
joint motions, the superficial layer of the cartilage is exposed to
both tensile and compressive stresses as the cartilage of the oppos-
ing facet makes contact, glides over it and applies compression
[49]. With increasing tensile strain, the collagen fibers untangle
and straighten to exhibit nonlinear-to-linear r-ebehavior (Eq.
(12)), referred to as the fiber-recruitment model [56].
r¼Aexp BeðÞ1½ (12)
Although the tensile strength of cartilage is provided by collagen
fibers, its compressibility depends on the water content
[64,66,125,126]. Since joint cartilage contains both fluid and
solid elements it exhibits viscoelastic properties [125130]. This
response has been demonstrated in pure-shear tests during which
a cartilage specimen is subjected to a sinusoidal angular dis-
placement, while measuring the torque that is generated. From
such studies, the dynamic viscoelastic shear modulus (G*) of
cartilage has a complex value that has been described by a sinu-
soidal function of the phase shift angle (d) between the applied
angular displacement and the torque [56]. The viscoelastic na-
ture of cartilage is highlighted by the storage (G’) and loss mod-
uli (G”) that compose the complex shear modulus (G*). The
magnitude of the complex shear modulus depends on the ampli-
tude of the angular displacement input (h
), the torque (T
), the
thickness (h) and polar moment of inertia (I
of the specimen
(Eq. (13)):
Although many studies have reported the mechanical properties
of cartilage from other diarthrodial joints ([64] provides a sum-
mary from several investigations), the mechanical properties of
healthy facet cartilage tissue have not been well-studied. In fact,
only one investigation reports a Young’s modulus of
10.08 68.07 MPa and an ultimate strength of 4.44 62.40 MPa
for dog-bone-shaped specimens of canine lumbar facet cartilage
under tensile loading [66]. Even less has been defined regarding
the compressive properties of facet cartilage. The surface of the
human facet articular cartilage is curved and has a nonuniform
thickness (with a maximum thickness of only approximately 1
mm), making it challenging to harvest. Although the techniques
employed by Elder et al. [66] provide a potential method to col-
lect human facet cartilage for compressive and tensile testing,
further biomechanical investigations of human facet cartilage are
currently lacking. Such techniques may soon enable additional
testing to provide a more complete understanding of this tissue’s
properties in the human.
3.8 Facet Forces and Pressures. Because the compressive
force between articular facets in the joint is transferred to the
underlying bone, pressure measurements are important to identify
the loading experienced by the cartilaginous matrix. Further, cer-
tain loading conditions may place the cartilage matrix at risk for
damage and the underlying bone at risk for compressive fracture.
However, direct measurement of the contact pressure between the
articular cartilage surfaces in the intact facet joint is quite chal-
lenging without rupturing the capsule and altering the macroscale
mechanics of the joint. Since pressure and force are related by
contact area, facet contact pressure has been measured indirectly
using proxies such as force applied to the facet. Such force meas-
urements have been made during different modes of loading
(compression, extension, flexion, lateral bending, and rotation) in
spines from different species. Lumbar and cervical facet forces
have been estimated using strains measured on the articular pillar
and lamina during flexion, extension, lateral bending, and axial
rotation with applied moments varying from 1 to 7.5 Nm with a
100 N axial preload [11,131,132]. In that approach, uniaxial strain
gauges were aligned along the major axis of the articular pillar
(supero-inferior direction; Fig. 1) and the measured strains from
the gauges were used to indirectly interpolate the force transferred
through the joint [11,131]. After testing in the motion segment,
the facet joint was removed en bloc and tested using loads that
were applied at different locations on the exposed articular carti-
lage, perpendicular to the surface, to establish a strain-force rela-
tionship that correlated the strains measured during testing to the
actual compressive load develop in the joint at these locations
[11,131]. Using that approach, average facet forces of 74 N were
estimated for the canine lumbar spine under 2 Nm of extension
[131]. Chang et al. [132] reported 205 N under a 10 Nm extension
moment combined with 190 N of axial compression, and Sawa
et al. [11] reported 51 N under a 7.5 Nm extension moment of
human lumbar segments. Although the investigations by Butter-
mann et al. [131] and Sawa et al. [11] differed in the specifics of
the testing methods and specimen species, the facet force in both
studies was found to increase when an axial compressive load was
superposed on the primary loading vector. Also, both studies iden-
tified contraletral axial rotation away from the joint being investi-
gated as the loading condition generating some of the highest
facet forces (Table 2). However, further comparisons cannot be
made because of the differences in the testing methods of these
investigations. In addition, a recent study using the same tech-
nique as Buttermann et al. (1991) [131], with strain gauges
mounted on the lamina of an isolated cadaveric lumbar vertebra,
showed that considerable error can stem from determining facet
force from extra-articular strains in all loading configurations
except axial rotation [133]. Nevertheless, this strain gauge tech-
nique for the evaluation of the facet force preserves the facet cap-
sule and enables comparison of load transfer through the facet
joint before and after implantation of a medical device such as a
fusion cage or an artificial disc
Thin and flat pressure-sensitive paper or sensors can also be
inserted between the articular surfaces of the joint after capsule
transection to measure facet force [134136] and contact pressure
[137139]. Using pressure-sensitive paper Dunlop [137] was one
of the first to localize the regions and maximal magnitudes of con-
tact pressure that are established in the human cadaveric lumbar
facet joint during combined loading, with sagittal bending coupled
with a 1000 N compressive load and a 200–400 N shear load
applied to motion segments. Contact pressures of up to 3.7 MPa
and 6.1 MPa were noted in the central-medial and central-inferior
(dorsal) regions of the articular surface near its periphery for 4
degrees of flexion and 6 degrees of extension, respectively [137].
Later, Wiseman et al. [138] measured mean (0.93 MPa) and peak
(3.73 MPa) pressures with the same technique in lumbar joints
under more aggressive loading scenarios (a combined 700 N axial
compression and 15 Nm extension) [138]. More recently Niosi et
al. [136] implemented a flat electroresistive pressure sensor array
in the L3-L4 facet joints of lumbar motion segments subjected to
a 7.5 Nm moment (with and without a 600 N compressive pre-
load) in sagittal bending, lateral bending, and axial torsion. The
calibrated sensor measured facet forces of 4 N in flexion, 14 N in
extension, 16 N in lateral bending, and 56 N in axial torsion
[136]. Using pressure-sensitive paper and a tip-mounted pressure
probe fitted through the superior articular facet, our group has
measured facet contact pressure in cervical motion segments sub-
jected to 0.8–1.7 Nm extension moments [106] (Fig. 3). The pres-
sure paper localized the area of articular contact in the posterior
region of the facet near the periphery of the joint and measured an
average pressure of 92 kPa, while the pressure transducer meas-
ured an average pressure of 158 kPa [106] (Table 2). Together, all
of these investigations showed that contact is not uniform over the
articular surface and that the location of contact varies during dif-
ferent loading conditions likely owing to the shape and incongru-
ence of the facet surfaces.
Although the flat pressure sensors enable spatial mapping of the
location of contact between the articular surfaces of the facet joint
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Table 2 Summary of estimated facet forces and pressures during loading of intact specimens
Loading Details
Level Direction Magnitude Property Technique Reference
C4-C5 Compression 80 N Force (N) 30 38 FEM [323]
C5-C6 Pressure (MPa) 0.13 0.19
C4-C5 Flexion 1.3 (60.3) Nm Pressure (MPa) 0.086 (60.012) Pressure paper between [106]
C6-C7 Extension 1.7 (60.5) Nm 0.092 (60.014) facet surfaces
Flexion 1.3 (60.3) Nm 0.047 (60.057) Tip-mounted pressure sensor
Extension 1.7 (60.5) Nm 0.158 (60.040) in posterior region
of superior facet
C5-C6 Flexion 2.7 (60.3) Nm Pressure (MPa) 0.010 (60.010) Tip-mounted pressure sensor [146]
Extension 2.4 (60.3) Nm 0.068 (60.027) in posterior region
of superior facet
C5-C6 Compression 73.6 N Force (N) 4.2 N FEM [322]
Flexion None
Extension 1.8 Nm 37.6
Axial torsion 28.5
Cervical Compression 4.5 mm Compressive 0.29 0.55 0.39 0.05
Flexion stress (MPa) 0.04 0.23 3.98 0.28
Extension 18 deg 0.29 0.30 4.90 0.23
Lateral bending 0.02 0.14 3.81 0.23
T12-L2 Flexion Force (N) 46.1 (641.3) Uniaxial strain gages on the [11]
Extension 51.5 (639.0) outer lateral portion of L2
Axial torsion ipsi 31.3 (633.4) superior articular processes
Axial torsion contra Up to 7.5 Nm 70.3 (643.2)
Lateral bending ipsi (1.5 Nm increments) 32.0 (644.4)
Lateral bending contra 30.6 (629.1)
Axial compression 400 N 45.5 (640.4)
Flexion Unspecified 46.6 (641.9)
Extension Unspecified 75.4 (639.0)
L1-L2 Neutral position Pressure (MPa) 4.5 (61.6) Pressure paper between facet surfaces [137]
L3-L4 Flexion 4 deg 3.7 (61.3)
L5-S1 Extension 4 deg 5.8 (61.6)
6 deg 6.1 (61.9)
L1-L5 Compression 500 N Force (N) 43 (@ L2-L3) FEM [268]
Extension 7.5 Nm (@ L1) 86 (@ L2-L3)
Extension 20 deg (@ L1) 117 (@ L4-L5)
Axial compression 100 N 23 (616) Uniaxial strain gages on the [131]
þouter lateral portion of
Flexion 1 Nm Force (N) None right L3 superior articular process
Extension 2 Nm 74 (623)
Axial torsion 4 Nm 92 (627)
Lateral bending ipsi 1 Nm 40 (632)
Lateral bending contra 1 Nm 54 (619)
Axial torsion 2.3 Nm 32 Pressure paper between
6.0 Nm 210 facet surfaces
Left Right [136]
L2-L5 Flexion Force (N) 2 (65) 4 (64) Pressure film between
Extension 7.5 Nm 13 (614) 14 (610) facet surfaces
Lateral bending 11(611) 16 (614)
Axial torsion 56 (617) 55 (618)
L4-L5 Flexion None FEM [271]
Extension 7.5 Nm Force (N) 50
Lateral bending 36
Axial torsion 105
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during spinal motion, that approach does require that the capsule
be cut in order to insert the sensor in the joint. Capsule transection
has been shown to contribute to hypermobility of the facet joint
[105] and can be hypothesized as also potentially inducing non-
physiological joint loads and/or articular contact, and in locations
that are not usually loaded in an intact joint. Capsule transection
does not likely affect the joint’s behavior in extension since the
capsule is not stretched and does not bear load during that direc-
tion of loading. But, force measurements in flexion, lateral bend-
ing, and axial rotation can be biased since the joint’s overall
mechanical behavior is modified by the capsule transection itself
[105,139141]. This could explain why the facet force values ex-
trapolated from pressure sensor measurements in the study by
Niosi et al. [136] were much smaller than those obtained from
strain gauge measurements (Table 2). Furthermore, in any loading
condition, the pattern and magnitude of contact are modified by
the presence of the sensing device [142144]. A similar, but less-
invasive, method was developed by el-Bohy et al. [145] that
maintains the integrity of the facet capsule. In that approach, a 1.5
mm-diameter pressure gauge implemented at the tip of a 13-gauge
steel tube was positioned below the posterior bony tip of a lumbar
inferior facet just above the cartilage covering the lamina of the
vertebrae below [145]. Contact pressures of up to 0.3 MPa were
measured at the edge of the articular surface of the lowest vertebra
when a combined 600 N compression and 15 Nm flexion loading
was applied to three-vertebrae lumbar segments. A comparable
sparing-capsule technique was recently implemented in cadaveric
cervical motion segments to determine average facet pressures of
10.3 69.7 kPa and 67.6 626.9 kPa for 2.7 Nm flexion and 2.4
Nm extension moments, respectively [144] (Fig. 3).
4 Mechanotransduction
Since a portion of the spine’s mechanical loading is supported
by the facet joint, a variety of mechanical, physical, and chemical
cascades are initiated in response to loading of the individual tis-
sue components comprising the facet joint. These physiologic
responses occur across several scales, ranging from the macro-
scopic tissue-level, to cellular and molecular levels via many
mechanotransduction mechanisms. Although mechanotransduc-
tion can control and contribute to maintenance of the tissues in the
joint [57,147149], this process can also lead to and/or accelerate
tissue degeneration and dysfunction [150,151]. The mechanisms
of mechanotransduction in articular cartilage, ligaments, and bone
have been described in other synovial joints. Broadly, as the first
step the external primary spinal input (load or motion) is trans-
formed into a secondary tissue-specific loading profile (Fig. 4).
Then, the tissue-specific loads elicit a host of cascading mechani-
cal, electrical, and chemical responses from the various elements
that compose the tissue. These responses trigger further chemical
changes that affect the intracellular milieu (protein translation,
gene transcription, post-translational signaling) and the intercellu-
lar signaling (Fig. 4). Both the initial mechanical, electrical, and
chemical changes and the modification of the intracellular milieu
alter the intercellular signaling as well as the cellular activity (i.e.,
proliferation, differentiation, apoptosis). Modification of cellular
activity can result in the release of chemical agents and electrical
signals that influence the maintenance of the extracellular milieu,
but can also alter the secondary tissue-specific loading (Fig. 4).
Together, these physiological responses can modify the mechani-
cal behavior of the tissue and lead to further changes in its
response to mechanical loading and degeneration. Although this
cascade has been well defined through a large body of elegant
work, very few articles specifically address and detail these proc-
esses in the tissues of the facet joint. Therefore, this section
reviews the mechanotransduction mechanisms known for the facet
joint tissues and also provides a more global review of such mech-
anisms in similar tissues from other synovial joints.
4.1 Mechanotransduction in the Facet Joint and its
4.1.1 Capsular Ligament. Since spinal loading and motion
are both guided and constrained by the facet joints the primary me-
chanical loading of the facet joint induces primarily capsular
Fig. 3 Representative data quantifying the spinal rotations and
pressure responses in the facet of a multisegment (C2-T1)
cadaveric cervical spine during a range of bending moments
applied in continuous flexion-extension. The pressure response
in the C5-C6 facet joint increases with applied extension as con-
tact is developed in the articulating facets, but exhibits a differ-
ent pattern than the rotation angle. In contrast, during flexion,
when the joint opens up there is no pressure detected.
Tab le 2 Continued
Loading Details
Level Direction Magnitude Property Technique Reference
L5-S1 Axial compression 650 N Force (N) Pressure film between [134]
Shear 550 N facet surfaces
þGroup 1 Group 2
Flexion 40 (613) 45 (610)
Extension 6 deg 54 (618) 65 (618)
Lateral rotation ipsi 50 (613) 54 (619)
Lateral rotation contra 9 (64) 33 (610)
Note: C-Cervical, T-Thoracic, L-Lumbar; FEM Finite Element Model; SL-slideline model, CS-contact surface model, HE-hyperelastic model, FL-
incompressibe fluid model of articular cartilage; ispsi ipsilateral, contra – contralateral.
reported here from non-tabular data of two separate test groups.
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ligament stretch and compression of the cartilaginous articular
surfaces and the subchondral bone. In the capsular ligament under
stretch, the collagen fiber structure and the nerve endings embed-
ded in that network [152] and cells (fibroblasts, macrophages) are
all distorted and activated [153]. Accordingly, capsular deforma-
tions of certain magnitudes can trigger a wide range of neuronal
and inflammatory responses [124,154,155]. Neurophysiologic
studies with a goat model have shown that the nerve endings in the
capsule possess different stretch thresholds for activation [116].
Although most of the proprioceptive and nociceptive afferents
have a low-strain threshold (10%) for activation, a few receptors
have a high-strain threshold (42%) for signal generation via neural
discharge. In addition, capsular strains greater than 47% activate
nociceptors with pain signals transmitted directly to the central
nervous system [116]. Among both the low- and high-strain thresh-
old neural receptors in the capsular ligament a few sustain their fir-
ing even after the stretching of the capsular ligament is released
[116,154]. This persistent afterdischarge evident for strains above
45% constitutes a peripheral sensitization that may lead to central
sensitization with long-term effects in some cases [154]. Also, in
vivo stretch of goat cervical capsule until its rupture (up to 30
mm) showed that the strains in the capsule averaged 73% and were
sufficient to induce changes in axons taken as indicators of dys-
function (i.e., swelling, retraction beads, vacuolations). The effect
of the capsule distraction on axonal changes was significant, with
the ratio of abnormal to normal axons being greater in the stretched
(94/186) than in unstretched capsules (29/108) [121]. Such axonal
changes can be a source of hyperexcitability, spontaneous firing,
and persistent pain [156] since that axonal dysfunction subse-
quently disrupts gene transcription of substance P, a neuropeptide
protein involved in pain signaling [157]. In addition, inflammation
in the facet joint also increases the discharge rate of multiunit
nerves, sensitizes the nerves to mechanical stimulation, and acti-
vates previously inactive nerves [158].
The neural signals from the capsule travel via the primary affer-
ents to the dorsal root ganglion (DRG) and spinal cord, and can
induce several hallmarks of neuroinflammation, including glial
activation [155] and cytokine upregulation [159,160]. These
inflammatory responses have been reported after failure of the
facet capsular ligament and also after its subfailure distraction in a
rat model [161]. In response to the injurious stimuli, neuropepti-
des involved in pain signaling, such as substance P, are also modi-
fied. Substance P protein expression in the DRG after painful
capsule distractions was twice that of nonpainful distractions of
controls [115,124]. Although no gross damage of the capsule was
observed after a painful distraction, spinal astrocytic activation
was 61% greater and pain symptoms were also increased [113].
Capsular strains causing damage to the ligament structure can
also activate fibroblasts directly or indirectly for structural repairs.
While strains causing excessive failure of the collagenous liga-
ment structure trigger an inflammatory-driven cellular response,
subfailure strains elicit a fibroblast-mediated remodeling response
to restore integrity to the damaged collagen structure [153]. Com-
plete tissue tearing elicits an inflammatory response of the tissue
that results in macrophage infiltration in order to clear any debris
from the damaged collagen fibers and matrix. During the phago-
cytosis of the debris these cells release molecules that also trigger
the recruitment of additional fibroblasts with increased collagen
expression and this response can also lead to the formation of a
provisional collagenous scar [153]. In the case of a subfailure
loading scenario, no inflammatory response is observed and an
increase in proteoglycans (decorin, fibromodulin) might actually
help to modulate the fibrillogenesis of newly synthesized collagen
by the fibroblasts [153].
4.1.2 Cartilage. Compression of the articular cartilage in the
joint can occur during any mechanical loading of the facet joint
[136]. Although compressive load is transferred via the facets
between adjacent spinal levels and contact pressure develops in
the facets’ articular cartilage, contact is not uniform and the facet
surface presents both load-bearing and nonload-bearing regions
[162165]. Given the difference in material properties between
the various zones of the same tissue, the mechanisms by which
mechanical signals modulate physiologic responses likely also
lead to different spatial distributions of the responses in the
affected tissues. However, the particular relationship between the
mechanical, chemical, and cellular responses to compression in
the cartilaginous matrix of the different zones remains largely
unreported for the human spinal facet joint. Nevertheless, damage
to the cartilage structure elicits an inflammatory response
[43,166], which itself can also elicit not just osteoarthritis of the
joint but can modulate pain signals from other regions of the joint.
For example, one study showed that the inflammatory cytokines
IL-6 and IL-1bwere present in the facet cartilage retrieved from
patients undergoing surgery for lumbar spinal canal stenosis and
disc herniation [160]. This result led to the conclusion that pain
symptoms might be due not only to mechanical tissue insults but
also to chemical irritation of the tissue from the inflammatory
agents leaking from the facet joint into the spinal space.
4.2 Mechanotransduction Processes in Articular Cartilage
of other Synovial Joints. Mechanical stimuli elicit a cascade of
multistep responses including mechanocoupling, mechanotrans-
duction, intracellular conversion, and cellular response from artic-
ular cartilage (Fig. 4)[165]. These steps differ between the thick,
proteoglycan-rich load-bearing areas and the mechanically weaker
nonload-bearing areas of the articular cartilage layer because the
extracellular environment (collagen fibers, proteoglycan and water
content) varies along the depth of the cartilage layer (Fig. 2).
These structural and compositional variations imply that the cellu-
lar responses to mechanical loading vary within each zone of the
cartilage layer as well [46,167].
Tensile stresses that arise in the more superficial zone of the carti-
lage layer and hydrostatic pressure increases in the transitional and
deep zones are converted at the tissue and cellular levels into electri-
cal, chemical, and biomechanical stimuli [168]. Distortion of the
Fig. 4 Schematic representation of the generalized processes
of mechanotransduction in synovial joints, across the scales
ranging from tissue to molecule
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chondrocyte membrane and nucleus, changes in membrane poten-
tial, electric stimulation from streaming potentials and changes in
matrix water content, ion concentrations and pH are all likely
involved in the metabolic changes of compressed cartilage [168].
Chondrocytes are embedded in the collagen matrix of cartilage
and deform with it under compressive and shear strain [46,168
170]. Round chondrocytes cultured in a bioengineered cartilage
have been shown to become polygonal, doubling their area, and
spread after cyclic compressive load (1800 cycles at 1 Hz); they
resumed their initial round shape and size within six hours after
the cessation of stimulation [170]. Changes in chondrocytic shape
and spreading are linked to an increase in their secretion of matrix
metalloproteinases that augment the accumulation of newly syn-
thesized proteoglycans [170], and maintain the tissue function by
synthesizing matrix molecules such as aggrecans and type II colla-
gen [169,171]. When chondrocytes are deformed, ion channels
present on their membrane are activated [165,172]; the levels of
intracellular calcium Ca
have been shown to increase under
hydrostatic pressure [173]. An increase in calcium concentration
can inhibit the accumulation of cyclic adenosine monophosphate
(cAMP), a second messenger used for intracellular signal trans-
duction. This inhibition and reduction in cAMP may induce cell
proliferation [173]. Under mechanical loading, chondrocytic pro-
liferation and differentiation [46,165] result in a greater number
of cells for the synthesis of extracellular matrix (ECM) compo-
nents (collagen, proteoglycans). For example, intermittent tensile
stresses applied to chondrocytes from the rat rib growth plate
increased both DNA and proteoglycan synthesis by about 1.5-fold
[174]. Newly formed collagen is used for the maintenance and
repair of damaged extracellular matrix, while providing a support
medium for the proteoglycans to trap water in order to resist com-
pression. However, if the hydrostatic pressure is too high (5–50
MPa) cytoskeletal elements of the chondrocyte, such as the Golgi
apparatus and microtubules, can disorganize, and there may be a
reduction in protein synthesis and inhibition of membrane trans-
port [175]. Changes in chondrocytic growth and cellular division
are modulated by mechanical loading as the physicochemical
mechanisms within the cells in which the coding for protein syn-
thesis occur via gene transcription, protein translation, and post-
translational modifications [41,147,165,176].
The mechanical loading of cartilage also influences tissue me-
tabolism indirectly through electric stimulation from streaming
potentials. Streaming potentials develop when the cations con-
tained in the synovial fluid penetrate the matrix to interact with
the increasing concentration of proteoglycans that bear negative
charges, as fluid flows in and out of the cartilage matrix during
compression [46]. Streaming potentials are likely associated with
an electric potential jump across the chondrocytes’ membrane
[172], which can stimulate the biosynthesis of these cells during
dynamic compression [177]. Considering the electrokinetic trans-
duction taking place during cyclic compression of cartilage led to
the development of a solid-fluid interaction relationship (see Eq.
(14)) in which the total area-averaged fluid velocity (U) and the
current density (J) both depend on the electric potential generated
across the specimen (V) and the fluid pressure of the surrounding
bath (P) via a combination of the circuit hydraulic permeability
), the electrokinetic coupling coefficients (k
) and the
electrical conductivity (k
) of the specimen [177].
¼k11 k12
k21 k22
 (14)
Upon this electrical stimulation, as well as a shape alteration
caused by a volume change in the surrounding extra-cellular ma-
trix, chondrocytes synthesize proteins involved in the maintenance
of the cartilaginous matrix other than collagen and proteoglycans.
Neu et al. [178] reported that transforming growth factor (TGFb)
mediates the secretion of lubricin, a glycoprotein synthesized by
the chondrocytes of the superficial zone, in bovine condylar
explants subjected to shear loading. Lubricin is a lubricative gly-
coprotein that maintains the tribological properties of the synovial
joint and inhibits synovial cell overgrowths. Knowing its regula-
tory mechanisms can provide insight on the progression and
potentially treatment of degenerative processes of cartilage [178].
Generally, the tissue and cellular responses to mechanical stimu-
lation in the joint depend on the frequency, amplitude, and rate of
loading. The rate of matrix synthesis decreases as the static hydro-
static pressure increases up to 50 MPa [46,179]. In contrast, cyclic
loading can stimulate matrix synthesis [165,174]. However, physi-
ologic responses to cyclic loading have been shown to be tempo-
rally- and spatially-dependent since extracellular osmolality is not
homogeneous across the depth of each cartilage layer
[46,50,165,167,178,180]. However, it is not possible to fully define
the tissue and cellular responses of cartilage to cyclic compression
because there are also variations and inhomogeneities in the extrac-
ellular oxygen content, water content, pH[48,167] and in the mag-
nitude, frequency, duration, and location of the applied loads
[41]—all of which complicate these responses. The mechanotrans-
duction processes and their synergistic or antagonistic interactions
in articular cartilage are not yet fully understood. However, from a
global perspective, mechanical energy is converted into a type of
energy useful for the cells to proliferate, differentiate, communi-
cate, and synthesize proteins for the maintenance of the extracellu-
lar-matrix in response to loading. The mechano-electrochemical
processes that take place in cartilage are thus very similar to those
that take place in the surrounding bone and ligaments. However,
such mechanotransduction mechanisms could be limited in carti-
lage because, unlike those other tissues, it is both avascular and
aneural [48].
4.3 Mechanotransduction Processes in Bone and Ligament.
In bone, as in cartilage, a multistep process including mechano-
coupling, biochemical coupling, transmission of intercellular sig-
nal, and effector cell response (Fig. 4) was identified between
mechanical strains and the tissue response [181]. Briefly, mecha-
nocoupling defines how mechanical energy is detected by the
bone cells in the tissue (osteocytes, osteoblasts). The process con-
sists in a transformation of mechanical strains into fluid pressure
in the canaliculae, which generate fluid shear stresses on the
osteocytes’ membranes. Fluid flow also generates streaming
potential, an electrical energy, which stimulates bone cells for
remodeling and repair as proven by bone fracture healing from ex-
posure to electromagnetic fields [182185]. Both the magnitude
and the frequency of the mechanical stimulation, and also the
strain rate, influence the bone cells response. Turner et al. [186]
reported that mechanically induced bone formation was not
increased in the rat tibia subjected to bending until the loading fre-
quency increased over 0.5 Hz [186]. Biochemical coupling likely
takes place at the binding interface between the cells’ membrane
and the extracellular matrix. This attachment of the cell generates
tensile forces on the cytoskeleton that alter the shape of the cell,
its phenotypic expression, the binding of protein to the cell’s
membrane, and the activation of ionic channels on the cell’s mem-
brane. Release of prostaglandins and nitric oxide by the osteocytes
activates the proliferation of and matrix synthesis by osteoblasts
[187]. The complex biochemical interactions between the ECM
and the bone cells and within the cells are quite complicated and
can be the focus of a review themselves. Since this is not the cen-
tral focus of this review and they have been elegantly detailed in
recent specialized reviews [148,187], they are only broadly pre-
sented here. In essence, several peptides and proteins, such as
insulinlike growth factor (IGF), epidermal growth factor (EGF),
bone morphogenic protein (BMP), transforming growth factor
(TGFb), bind to the osteoblastic membrane and modulate kinase
activity within the cell which induces expression of activator pro-
tein 1 (AP-1) in the cell nucleus. Intracellular chemical signaling
is also modulated by the entrance of extracellular calcium ions
into the intracellular milieu. Ultimately, the mechanotransduction
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mechanisms permit the adaptation and maintenance of the bone
structure to mechanical loading by acting on bone-regulating
genes contained in the nucleus of bone cells leading to their prolif-
eration, differentiation, and survival [148,187].
Mechanical loading of the facet capsule imparts structural
changes to the collagen structure that can result in the degradation
of its mechanical properties, loss of function, and pain-activating
protein generation [88,112,155]. But ligaments also contain fibro-
genic cells that are directly and indirectly affected by mechanical
loading as chondrocytes in cartilage [188]. Fibroblasts deform
when the tissue in which they are embedded deforms. Matyas et al.
[189] observed that the nucleus of fibroblasts contained in the rab-
bit medial collateral ligament were 4 lm longer and 1 lm thinner
when the ligament was under 6% tensile strain than at rest.
Accordingly, the nuclear roundness decreased from 0.4 to 0.19
[189]. Upon deformation of the cell membrane, stretch-activated
ion channels might be activated which permits the penetration and
increase of cation concentration in the intracellular milieu and can
eventually alter cellular activities as was described above for bone
and cartilage cells. Studies of the periodontal ligament showed that
mechanical stimuli also indirectly affect fibroblastic activity via
trans-membrane and intracellular signaling as for osteoblasts. Dis-
turbance of the ECM homeostasis leads to an intracellular conver-
sion of the mechanical signal into a biochemical one via the
transduction of focal adhesion molecules such as FAK and MAP-
kinases in the fibroblast [190]. The cells then synthesize and
release matrix metalloproteinase in the ECM for the regulation,
modification, or degradation of ECM components [169], to modu-
late the mechanical loading state of the tissue. Similarly, loading
of periodontal ligament fibroblasts was shown to activate various
kinase proteins (ERK, JNK, p38) that communicate with the inner
cellular milieu and activate AP-1 in the cells’ nucleus. AP-1 can
up-regulate the COL I gene in the nucleus of the fibroblasts, stimu-
lating collagen expression by these cells [191]. Collagen expres-
sion is used to either maintain or repair the extracellular matrix.
Fibroblast activity and interaction with the extracellular environ-
ment is directly and indirectly affected by mechanical loading and
tissue deformation. Such mechanotransduction mechanisms are
similar to those described for cartilage and bone. Although the gen-
eral mechanisms of mechanotransduction in bone, ligaments, and
cartilage have been identified, as illustrated in Fig. 3,theyremain
to be elucidated more specifically for the facet joint tissues.
5 Injury and Trauma of the Facet and Other Synovial
Facet joint injuries result most-often from motor vehicle and
sports trauma, such as skiing, snowboarding, cycling, and diving
[82,192195], and include a wide range of bony and ligamentous
lesions depending on the extent and type of tissue trauma. Inter-
estingly, unilateral and bilateral facet injuries make up nearly 6%
of all cervical injuries, with undisplaced fractures, subluxations,
and dislocations being the most commonly reported facet injuries
[192,193,195,196]. Facet injuries often directly damage the hard
and soft tissues that compose that joint (Fig. 2)[8,197200]. In
addition, facet trauma is also associated with the occurrence of
damage to other soft tissues of the spine, such as disc tearing, spi-
nal cord trauma, and/or nerve root compression, all of which can
also lead to a transient or even permanent loss of mechanical and
neurological function of the facet joint, spinal column and/or
physiologic sequelae [193,201206].
5.1 Facet Joint Injuries. Because the facet joints comprise
the integrative biomechanical structure of the spine, any violation
of their mechanical integrity as can be caused by injury or trauma
directly affects the mechanical behavior of a motion segment or
even the overall spinal region [195,206]. For example, a unilateral
locked facet at C5-C6 produced by combined lateral bending and
flexion of the cervical spine has been reported to significantly
reduce the segmental range of motion (ROM) by 2.7 – 3.6 degrees
in all modes of loading except in ipsilateral axial rotation [206].
Once unlocked, the ROM of the C5-C6 motion segment was fur-
ther increased compared to its preinjury intact condition by an
additional 3.5 – 8.0 deg. This report, suggests that the capsular lig-
ament had been damaged due to the facet injury. Indeed, Crawford
et al. [206] also reported that the laxity of the capsular ligament
was significantly increased after that locked facet condition com-
pared to that in the intact condition. Also, the increase in laxity
was associated with some ligament tearing, supporting both the hy-
pothesis that the ligament sustained damage and exhibited altered
mechanical properties [206]. Taken with reports of instability fol-
lowing facet dislocation [193,201,207], these findings imply that
there may be a continuum between the degree of instability and
trauma to the facet, with greater instability for more severe facet
trauma, including dislocation and the more-extreme fracture.
In a clinical study of patients with cervical unilateral lateral mass
facet fractures, Lee and Sung (2009) found that the degree of axial
rotation and the segmental kyphosis were significantly greater in
those patients whose facet was both fractured and dislocated than in
those sustaining only a facet fracture [195]. Despite these differen-
ces, both types of injury were associated with instability for the cer-
vical spine in rotation; surgical treatment was required to
sufficiently restore stability, again demonstrating the role of the
facet joint in limiting spinal motions, in particular rotation. Also,
unilateral fracture of the facet joint has been shown to lead to spon-
dylolisthesis, an anterior translation of the superior vertebra, some-
times associated with an axial rotation of the superior vertebra
around the intact contralateral mass [195,208]. Such a fracture
injury can lead to a variety of neurological disorders since the
motion segment is unstable and can compress the spinal cord and/
or nerve roots during certain motions. In these cases there is also
the potential for capsule injury when the fractured vertebra exhibits
abnormal kinematics during physiologic motions that can also
impose nonphysiologic compressive stresses on either or both the
capsule and cartilage of the contralateral facet joint [208]. In the
same way, excision of the capsule and cartilage during a surgical
procedure has been shown to increase the sagittal and axial ranges
of motion by 38% and 57%, respectively [209]. Most simply, frac-
tures of the articular pillar or lateral mass impose an overt disrup-
tion of the facet joint’s mechanical properties since they eliminate
the joint’s ability to support any load and, in so doing, can cause
spinal instability and neurological impairment.
5.2 Surgical Treatments of Facet Injuries and Effects on
Facet Biomechanics. Fractures of the bones of the facet joint
leading to joint separation, comminution, split, and traumatic spon-
dylolysis, require surgical treatment to reduce the anterior transla-
tion and axial rotational deformity associated with these injuries
[195,210]. However, the type of treatment varies with the type and
severity of the fracture [210]. A separation fracture that isolates the
entire lateral mass can be treated with a pedicle screw that provides
stability and strength while also encouraging bone growth [210].
However, if the separation fracture is also associated with disc and/
or ligamentous damage, a one-level reduction and stabilization is
recommended to avoid any slippage of the vertebra; fractures which
can also result in the development of multiple bone fragmentations
and traumatic spondylolisthesis also require only a single-level sta-
bilization to treat the unstable anterior translation of the superior
vertebra. Split and severe fractures have been shown to be success-
fully treated with two-level posterior fixation that resolves both the
spinal instability and restores the spinal alignment [210]. In the
most severe cases, the articular surface of the facet joint can
become completely obliterated and the articulation so disrupted
that the constraining and guiding functions of the facet joint cannot
be restored; in that case, fusion is necessary.
Surgical fusion can relieve many of the physiologic symptoms
caused by facet fracture but the mechanical function of the joint is
not fully returned to normal. In fact, the normal range and pattern
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