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Chapter 6 Biomechanics of the Spinal Motion Segment 103
TABLE 6.3 Average Neutral Zone (Degrees of Motion) for Dierent
Spinal Motion Segments in Dierent Motion Planes
Lateral
Vertebral Segments Flexion-Extension
C0–C1 1.1 1.6 1.5
C1–C2 3.2 1.2 29.6
C3–C6 4.9 4 3.8
C7–T1 and T11–T1 1.5 2.2 1.2
L1–L2 and L3–L4 1.5 1.6 0.7
L5–S1 3 1.8 0.4
From White AA III, Panjabi MM. Clinical Biomechanics of the Spine, ed 2. Philadelphia: JB
Lippincott; 1990.
Bending
Axial
Rotation
zones for rotary motions as a function of the plane of motion
and the spine level.2 For the most part, the neutral zone is
limited in range except for certain vertebrae in certain axes of
rotation. From a clinical perspective, one must be sensitive to
the fact that normal and abnormal neutral zones can be very
dierent for dierent vertebrae.
A large neutral zone can be an indication of several biomechanical factors. First, the neutral zone has been observed to
increase with age.30 Second, a larger-than-expected neutral
zone can indicate injury to the tissue.31 ird, some clinicians
contend that low resistance to movement is an indication of
clinical instability.32 ere are several reasons to consider carefully the range of movement within the neutral zone.
Load Tolerance of the Spinal Motion Segments
Tendon stress has been estimated to be between 60 MPa and
100 MPa.
34,35
ere seems to be a safety margin between the
muscle failure point and the failure point of the tendon by a
factor of about twofold35 to threefold.
34
Ligament and Bone Tolerance
Ultimate ligament stress has been estimated at approximately
20 MPa. e ultimate stress of bone has been found to depend
on the direction of loading. Bone tolerance can range from
51 MPa in transverse tension to 190 MPa in longitudinal
compression.
A temporal component to ligament recovery has also been
reported. One study found that ligaments required extended
periods to regain structural integrity. During the recovery
period, compensatory muscle activities have been observed.
Recovery time has been observed to be several times the
loading duration.
Because the spinal ligaments oen are the structure that
protects the spinal system, it is important to appreciate the
failure limits of the various spinal ligaments; these are shown
in Table 6.4. Note that the load tolerance of these ligaments
and the deformation characteristics of the ligaments vary
markedly according to the region of the spine and the specic
ligament involved. Generally, the lower the level of the spinal
ligament, the greater is the tolerance of the ligament. ere are
notable exceptions to this trend, however. Spinal ligaments are
viscoelastic and can increase their length under load. ey can
be responsible for an increase in the neutral zone; excessive
movement can also initiate muscle activities intended to
regain stability.
36,44
36–43
SECTION
I
e precise tolerance characteristics of human spinal tissues—
such as muscles, ligaments, tendons, and bones—loaded under
various conditions have been dicult to establish. Structure
tolerances have been observed to vary greatly even under
similar loading conditions because of their dependence on
many factors, such as strain rate (rate of loading), age of the
structure, frequency of loading, physiologic inuences, heredity, conditioning, and other unknown factors. In addition, it
has been impossible to measure these tolerances under in vivo
conditions. Many of the estimates of tissue tolerance have
been derived from various animal or theoretical constructs.
Tolerance data limits have been derived primarily from
cadaveric tissue. e obvious compromise in this approach is
that in vitro tissue when tested does not have the ability to
adapt or recover (and potentially increase tolerance) as does
a live human. e material properties of cadaveric tissue vary
depending on the manner in which the specimen was prepared
for testing. At least one study suggests that living tissue failure
might occur at magnitudes below those observed in cadaveric
specimens.
33
Muscle and Tendon Strain
Muscle has the lowest tolerance among the tissues of the spine.
e ultimate strength of a muscle has been estimated at
32 MPa.34 Muscle oen ruptures before a (healthy) tendon.35
Contact Force Tolerance
Contemporary logic suggests that pain secondary to biomechanical loading of the spine may result from direct stimulation to the facet joints, pressure on the anulus, or pressure on
the longitudinal ligaments.9 At these sites, inammatory
responses and analgesic responses are thought to be involved
in the development of pressure and pain. It is much more
dicult to specify load tolerance thresholds for contact pressures because the body’s individual responses to the imposed
loads collectively dene the pressure imposed on the spinal
structure. e tolerance limits for these structures has not
been well dened at this time.
Tolerance of Specic Spine Structures
e general structure tolerance, or failure, limits in response to
loading of the lumbar spine have been well investigated. Table
6.5 provides a summary of these tolerances reported as a func-
tion of the nature of the loading for the spinal motion segment
structures and the disc and vertebral body structures.
18
Compression
e compression dimension of spine tolerance has been
widely examined. Of all the structures in the spinal motion
104 BASIC SCIENCE
TABLE 6.4 Failure Strength of Spinal Ligaments
LOAD (N) DEFORMATION (mm) STRESS (MPa) STRAIN (%)
Average Range Average Range Average Range Average Range
Upper Cervical
C0–C1
Anterior atlanto-occipital membrane 233 18.9
Posterior atlanto-occipital membrane 83 18.1
C1–C2
ALL 281 170–700 12.3
Atlanto-axial membrane 113 8.7
CL 157 11.4
Transverse ligament 354
C0–C2
Apical 214 11.5
Alar 286 215–357 14.1
Vertical cruciate 436 25.2
Tectorial membrane 76 11.9
Lower Cervical
ALL 111.5 47–176 8.95 4.2–13.7
PLL 74.5 47–102 6.4 3.4–9.4
LF 138.5 56–221 8.3 3.7–12.9
CL 204 144–264 8.4 6.8–10
ISL 35.5 26–45 7.35 5.5–9.2
SSL — — — —
Thoracic
ALL 295.5 123–468 10.25 6.3–14.2
PLL 106 74–138 5.25 3.2–7.3
LF 200 135–265 8.65 6.3–11
CL 168 63–273 6.75 3.9–9.6
ISL 75.5 31–120 5.25 3.8–6.7
SSL 319.5 101–538 14.1 7.2–21
Lumbar
ALL 450 390–510 15.2 7–20 11.6 2.4–21 36.5 16–57
PLL 324 264–384 5.1 4.2–7 11.5 2.9–20 26 8–44
LF 285 230–340 12.7 12–14.5 8.7 2.4–15 26 10–46
CL 222 160–284 11.3 9.8–12.8 7.6 7.6 12 12
ISL 125 120–130 13 7.4–17.8 3.2 1.8–4.6 13 13
SSL 150 100–200 25.9 22.1–28.1 5.4 2–8.7 32.5 26–39
From White AA III, Panjabi MM. Clinical Biomechanics of the Spine, ed 2. Philadelphia: JB Lippincott; 1990.
ALL, anterior longitudinal ligament; CL, capsular ligament; ISL, interspinous ligament; LF, ligamentum avum; PLL, posterior longitudinal ligament; SSL, supraspinous ligament.
segment, the endplate is considered to be the “weak point of
the system,” or the structure with the lowest tolerance to force.
Compression failure limits are a function of age, with older
endplates failing at lower levels of force, and a function of
gender, with female tolerances lower than male tolerances.
45,46
Fig. 6.13 shows a summary of the compression strength for
much of the spine. e magnitude of force required for endplate tissue failure follows a normal distribution that ranges
from 2000 to greater than 14,000 N. When compression forces
increase on a spinal motion segment, the rst signs of damage
usually occur at the endplate or the trabeculae that support
the endplate. e endplate must be a thin structure to serve
its nutrition transport function. Because it is thin, it is also a
very weak structure, however, and subject to early failure
when load is applied.
Failure is believed to be initiated by the NP of the adjacent
disc. is nucleus causes the endplate to bulge and compromise
the vertebral body. e superior endplate is damaged more
oen than the lower endplate. In some cases, it is possible for a
portion of the NP to make its way vertically through a herniation of the endplate into the bone.18 is herniation can calcify
and form a Schmorl node. Endplate fractures are dicult to
detect via routine radiographs; however, magnetic resonance
imaging (MRI) can indicate biologic (modic) changes that are
characteristic of vertical displacement of the NP.
18
When the endplate experiences excessive compressive load,
the endplate can bulge into the vertebral body, increasing the
volume available to the nucleus. is decompression of the
nucleus means that it cannot resist compression well, and
more of the load is borne by the AF. e anulus can become
unstable and the lamellae can become compressed, and cannot
be supported any longer by the nucleus. It is believed that this
form of disc loading can result in internal derangement of the
disc and potentially reverse bulging of the inner lamellae.
Chapter 6 Biomechanics of the Spinal Motion Segment 105
(1800 lbf)
Compression strength in newtons (pound-force)
Vertebral level
Compressive strength (kN)
08
Age (y)
TABLE 6.5 Tolerance of Lumbar Motion Segment and Disc Structures
as a Function Load and Motion Characteristics
Failure Site Average Tolerance
Motion Segments
Compression Endplate
Shear Neural arch 2 kN
Flexion Posterior
ligaments
Extension Neural arch 26–45 N-m
Torsion Neural arch 25–88 N-m
Flexion and compression Disc or vertebra 5.4 kN
Disc Plus Vertebral Bodies
Shear Anulus 0.5 kN
Flexion Posterior anulus
Torsion Anulus 10–31 N-m
From Adams MA, Bogduk N, Burton AK, et al. The Biomechanics of Back Pain, ed 2.
Edinburgh: Elsevier; 2013.
C3
C4
C5
C6
C7
T1
T2
T3
T4
T5
T6
T7
T8
T9
T10
T11
T12
L1
L2
L3
L4
L5
0 2000
(450 lbf)
4000
(900 lbf)
5.2 (±1.8) kN all specimens
6.1 (±1.8) kN men
(20–50 y)
73 (±18) N-m with
compressive load of
0.5–1 kN
33(±
13 N-m)
6000
109
107
1957
106
1880
1967
8000
Messerer,
Perry,
Bell et al.,
(1350 lbf)
Strength/kN = a + b • age/decade
15
a = 10.53
b = –0.974
2
r
= 0.39
10
5
0
a = 7.03
15
b = –0.591
2
r
= 0.35
10
5
0
a = 8.60
15
b = –0.728
r2 = 0.27
10
5
0
020406
FIG. 6.14 Strength tolerance to static lumbar compression derived from
the literature as a function of age and gender. (From Jager M, Luttmann A,
Laurig W. Lumbar load during one-hand bricklaying. Int J Indust Ergo.
1991;8:261–277.)
Male
n = 174
Female
n = 132
Total
n = 342
0
endplate tolerance dierently between men and women,
however. e decrease in tolerance with age is nearly two
times greater for men compared with women.
45,46
In addition,
the strength of the vertebrae is nearly 0.8 kN lower than that
of the disc.46 Finally, strength increases as one moves down
the lumbar spine by approximately 0.3 kN per lumbar level.
47
Repetitive loading also seems to inuence the tolerance to
load of the motion segment. Fig. 6.15 shows how the number
of load repetitions and the relative magnitude of the load collectively have a dramatic impact on probability of failure of
the segment. As can be seen in this gure, when the relative
load becomes greater, the chances of failure increase the risk
signicantly when the number of loading cycles increases.48
Studies have also shown that, as the exion angle increases, the
number of cycles required for failure is dramatically reduced.
47,49
SECTION
I
FIG. 6.13 Estimates of vertebral compression tolerance (strength) under
slow load rates for the various vertebrae from C3 to L5.
III, Panjabi MM. Clinical Biomechanics of the Spine, ed 2. Philadelphia: JB
Lippincott; 1990.)
As noted earlier, endplate tolerance seems to be a function
of gender and age.
the literature are shown in Fig. 6.14. Although great variability
is evident, women generally have lower compression tolerance
by an average of almost 2 kN compared with men. In addition, tolerance reduces signicantly with age. Age inuences
106–108
(From White AA
45,46
Tolerance estimates based on a review of
Shear
e disc bers and intervertebral ligaments are inadequately
oriented to resist shear forces. Shear causes the disc to creep
during repetitive loading.50 Under many situations, the neural
arch resists shear force, however. e articular process resists
on average 2 kN of load before failure; however, this can range
from 0.6 kN to 2.8 kN.51 e specic point of load application can also greatly aect tolerance of the neural arch to
shear. Fig. 6.16 shows how diering methods of shear force
application can result in dramatically dierent neural arch
load tolerances.
52,53
106 BASIC SCIENCE
Probability of failure
60–70%
Load cycles
1000 N (224 lbf)
3000 N (670 lbf)
3000 N (670 lbf)
Lamy and colleagues,
1975
100
80
50–60%
Load range
40–50%
30–40%
20–30%
FIG. 6.15 Probability of vertebrae failure as a function of load magnitude and number of cycles of loading.48
(Modied from Marras WS. The Working Back: A Systems View. Hoboken, NJ: John Wiley & Sons; 2008.)
Weiss,53 1975
52
FIG. 6.16 Force tolerance of neural arch varies greatly as a function of
shear force application method.
Biomechanics of the Spine, ed 2. Philadelphia: JB Lippincott; 1990.)
22,53
(From White AA III, Panjabi MM. Clinical
60
40
20
0
10 100 500 1000 5000
Torsion
e motion segments oer little resistance to small angles of
axial rotation. Torsion is rst resisted by collagen bers in the
anulus that simply stretch slightly.
the articular surfaces make contact at one of the zygapophyseal
joints, and motion is limited to 1 or 2 degrees.21 is ROM
increases, however, with greater disc degeneration.
typical loading conditions (involving torsion and compression),
the loads imposed on the spine are shared by several structures.
At the limit of the natural range of movement, 30% to 70% of the
applied torque is resisted by the zygapophyseal joint as a compressive load, 20% to 50% is resisted by the disc, and less than 15%
is resisted by all of the intervertebral ligaments, collectively.
e lower limit for initiation of damage owing to torque
application seems to begin at about 10 to 30 N-m.21 Many clinicians believe that damage owing to torsional movements occurs
at the zygapophyseal joint before damage occurs to the discs.
18,60
With further axial motion,
61–63
Under
18,21
18
Repetitive shear loading can also reduce the tolerance to
380 N.51 Some authors have concluded that the limit at which
shear begins to increase risk is 750 to 1000 N,
this is also known to vary according to load rate.
tion, studies have reported failure occurring at the pars under
these conditions. Fig. 6.17 shows a summary of results of
ultimate shear strength of human cadaveric lumbar spines
obtained from in vitro studies.59 Gallagher and Marras conducted a Weibull analysis on shear failure data of human
cadaveric lumbar spines and recommended a maximum
permissible shear limit of 1000 N for occasional exposure to
shear loading (≤100 loadings/day) during occupational tasks.
However, for activities resulting in more frequent shear loadings (100–1000 loadings/day), they recommended a shear
limit of 700 N.
59
54–56
although
57,58
In addi-
Flexion and Extension
Signicant repositioning of the spine results when exion
and extension of the spine occurs. Dierent structures are
responsible for resisting force, and the tolerance of the spine
can change. During extension of the spine, 60% to 70% of the
applied load is resisted by the neural arch. Studies have reported
damage resulting from 3 to 8 degrees of extension under
bending moments of 28 to 45 N-m.
is oered by the disc and the anterior longitudinal ligament.18
Of particular concern is the risk of the anulus bulging into the
vertebral canal and compromising canal space.
It is hypothesized that the zygapophyseal joint would be the
structure damaged rst owing to extension. However, it is also
believed that the interspinous ligament may be at risk because
it would be compressed by opposing spinous processes. Rapid
load rates, possibly resulting from athletic endeavors, are also
thought potentially to increase risk.
64,65
Resistance to extension
(5 mm/s)
(5 mm/s)
(2 mm/min)
(0.5 mm/s)
(50 mm/s)
(50 mm/s)
3000
Ultimate shear strength (N)
Chapter 6 Biomechanics of the Spinal Motion Segment 107
SECTION
Flexion can lead to injury when imposed moments reach
50 to 80 N-m.
segment reaches 5 to 9 degrees per motion segment in the
upper lumbar spine and 10 to 16 degrees per segment in the
lower lumbar spine. e rst structures to sustain damage
are the interspinous and supraspinous ligaments.68 During
complex motions involving exion and lateral bending, the
capsular ligaments can also be compromised. e nal tissue
to fail is the outer posterior AF. In isolation (without the
ligaments), the disc can fail when exed at 18 degrees with an
application of 15 to 50 N-m of load.69 As with most structures,
load rate also plays a role in tolerance. Resistance to exion
can increase by more than 10% when rapid motions (10
seconds) are compared with slow motions (1 second).70 Static
postures seem to reduce resistance to bending by very large
amounts, probably owing to the interrelationship between the
ligamentous system and muscular control.
Lateral Motion
Less has been reported about the tolerance associated with
lateral bending moment exposure. Some studies have reported
that a lateral bending moment of 10 N-m results in 4 to 6
degrees of lateral bending in the lumbar spine, with most of
the resistance occurring at the disc.
degeneration, the ROM is greatly reduced to 3 to 4 degrees,
however, practically eliminating the neutral zone.
In Vivo Spine Biomechanics
Overview
Our fundamental knowledge of spine biomechanics has been
primarily gained through in vitro studies on animal and
2000
1000
1994
110
Begemann
1994
110
Frei
1998
111
51
Cyron
1976
FIG. 6.17 Summary of in vitro studies measuring the ultimate shear stress of human lumbar segments (error
bars represent the range of shear tolerance values). (From Gallagher S, Marras WS. Tolerance of the lumbar
spine to shear: a review and recommended exposure limits. Clin Biomech [Bristol, Avon]. 2012;27[10]:973-978.)
Cyron
1976
51
Begemann
Bisschop
2012
112
human cadaveric models in laboratories. However, the value
66–68
Damage occurs when the spinal motion
of these biomechanical results increases signicantly only
when it can be directly correlated to clinical outcomes.
Intuitively and experientially, we know that there are several
limitations associated with in vitro studies—such as specimen
integrity, lack of complex neuromuscular control, and proprioceptive and nociceptive inputs—that all play a signicant
role in producing the natural, graceful, and ecient motion of
the spine. Of particular clinical signicance is the understanding of pain-modulated motion, which is an in vivo phenomenon. Several studies have explored this complex phenomenon
in symptomatic individuals with low back disorders and found
signicant modications to their kinematics due to underlying pathology and pain when compared to asymptomatic
individuals.
74–76
Unfortunately, current in vitro testing methods
are unable to replicate pain-modulated kinematics, estimate
38
pain, or determine the impact of altered kinematics due to
pain avoidance on the overall mechanical response of the
spine and potential injury risk or damage. is creates a
substantial impetus for improving our understanding of the
mechanical behavior of the spine under in vivo conditions and
developing strategies to better translate biomechanical parameters from benchtop to bedside.
To address these limitations, over the past 15 years, several
62,71
If the disc experiences
researchers have developed tools to enhance our understanding
of in vivo spine biomechanics using advanced medical imaging,
62
motion-capture systems, and ecient numerical techniques
to help provide clinically measurable biomechanical metrics.
ese tools have the potential to help determine accurate in
vivo spinal motions in three-dimensional (3D) load exposures,
provide insights into mechanisms of spinal injury and pathology,
and facilitate overall assessment of treatment outcomes, design
of novel spinal implants, and improve current prevention and
rehabilitation strategies. e following sections highlight key
areas of research on in vivo spine biomechanics.
72,73
I
108 BASIC SCIENCE
30
flex
ext
flex
ext
flex
ext
C
Quantitative Assessment of in Vivo Spinal Motion
Overall Spine Kinematics (Extrinsic Measurements)
To appreciate the dierences involved in spine impairment, it
is important to understand the normal motion or kinematics
of the spine. It has been observed that people with low back
pain move more slowly.
be a result of the “guarding” that occurs in an attempt to
minimize the stimulation of pain-producing nociceptors.
Abnormal coupling of movement has also been shown to be
associated with low back pain.
Spine kinematic proles associated with asymptomatic
individuals and people with low back pain have been reported
in the literature at least for the lumbar spine. Fig. 6.18 summarizes how trunk ROM, velocity, and acceleration change
as a function of low back pain in the sagittal, lateral, and
transverse planes of the body. ere seem to be no dierences in ROM between the low back pain group and the
asymptomatic group. Signicant dierences are apparent,
however, when trunk velocity and acceleration are considered.
is seems to be the case in all motion planes of the body.
More recent studies have shown that kinematic ability can be
used to document the extent of a low back disorder.
dierences in velocity and acceleration are believed to be a
result of protective “guarding” employed by patients with low
74,75,77
Motion reduction is assumed to
76
74,77
ese
back pain through the excessive coactive recruitment of the
trunk muscles. is coactivity is believed to slow the motions
of the torso. e use of objective quantitative biomechanical
metrics to augment traditional subjective measures such as
pain questionnaires or Oswestry Disability Index (ODI) may
provide new insights during clinical evaluation and potentially
improve overall treatment outcomes.
Spine Kinematics (Intrinsic Measurements)
Several studies have investigated noninvasive techniques to
quantify normal in vivo spinal kinematics to aid in the clinical
diagnosis of spinal impairments and instability.
ity of these studies relied on static planar radiographs to assess
in vivo spinal ROM.
27,82
Fig. 6.19 illustrates the estimated
normal movement characteristics of the lumbar spine measured in living subjects. is gure indicates signicantly
dierent normal movements, particularly in exion-extension,
between in vivo and in vitro observations.
lights this dierence between the in vitro and in vivo observations in the sagittal plane.
27
ere is a general overestimation of extension movement
range in vitro and a general underestimation of exion range
in vitro. In addition, signicant dierences can be seen
between levels between the two states. It should be noted
that the measurements from these studies were from static
27,78–81
A major-
18,27
Fig. 6.20 high-
25
20
15
Degrees
10
5
0
A
300
250
2
200
150
100
Degrees/sec
50
0
Normal
Patient
Sagittal Lateral Transverse
Normal
Patient
Sag.
Sag.
FIG. 6.18 (A) Spine range of motion characteristics (mean and standard deviation [SD]) associated with
asymptomatic patients versus patients with low back pain in sagittal, lateral, and transverse planes of the body.
(B) Spine velocity characteristics (mean and SD) associated with asymptomatic patients versus patients with
low back pain in sagittal, lateral, and transverse planes of the body. (C) Spine acceleration characteristics (mean
and SD) associated with asymptomatic patients versus patients with low back pain in sagittal, lateral, and
transverse planes of the body.
Lat.
Lat.
Trans.
Trans.
60
40
20
Degrees/second
0
B
Sag. flex Sag. ext Lat. flex Lat. ext Trans. flex Trans. ext
Normal
Patient
Chapter 6 Biomechanics of the Spinal Motion Segment 109
Range of movement (degrees)
NORMAL MOVEMENTS IN THE LUMBAR SPINE
Degrees
RANGE OF MOTION OF LUMBAR
Degrees
B
14
12
10
8
6
4
2
0
L1–2
FIG. 6.19 Ranges of motion in lumbar spine during exion, extension,
lateral bending, and rotation.
et al. The Biomechanics of Back Pain, ed 2. Edinburgh: Elsevier; 2013.)
L2–3 L3–4 L4–5 L5–S1
Lumbar level
Flexion
Extension
Lateral bend
Axial rotation
25,27
(From Adams MA, Bogduk N, Burton AK,
MOTION SEGMENTS IN VIVO
20
Flexion
Extension
15
SECTION
I
10
5
0
L1–2 L2–3 L3–4 L4–5 L5–S1
A
RANGE OF MOTION OF LUMBAR
MOTION SEGMENTS IN VITRO
20
Flexion
Extension
15
Lumbar level
two-dimensional (2D) positions; thus, there are several inherent limitations with these measurements, such as kinematic
dierences between static and dynamic motions, inaccuracies
in measurements using radiographs, and inability to measure
multiplanar motion.
Recent advances in imaging technologies have helped to
address the limitations of static 2D measurements and have
facilitated 3D dynamic motions of the spine to be measured
in vivo with high accuracy and precision using a system of
synchronized biplanar radiographs.
spine models developed from computed tomography or MRI
can be directly matched to the biplane radiographs, and segmental kinematics can be determined in real time. Using this
technique, one study reported that, for healthy subjects, the
upper vertebrae in the lumbar spine had larger ROMs than
the lower vertebrae during functional exion-extension.
However, during lateral bending, the lower vertebrae showed
higher motion than the upper vertebrae. ey also found no
signicant dierence between levels during axial rotation
motion.80 A subsequent study on patients with degenerative
disc disease (DDD) found signicant dierences in spinal
kinematics between patients and healthy controls, especially
at L3–L4. ey found that L3–L4 showed the largest ROM in
patients in all planes of motion.81 Fig. 6.21 shows the results
of in vivo spinal rotations during dynamic functional motion
of patients with DDD and healthy controls.
Adjacent-level degeneration is a common occurrence clinically following a fusion surgery; however, its etiology is unclear
and controversial. Several in vitro studies have shown that
there is a signicant increase in adjacent-level kinematics
following fusion; however, the ndings of these studies are
83–85
78–81
3D person-specic
10
5
0
L1–2 L2–3 L3–4 L4–5 L5–S1
Lumbar level
FIG. 6.20 Range of exion and extension motion in lumbar spine
measured (A) in vivo and (B) in vitro.
AK, et al. The Biomechanics of Back Pain, ed 2. Edinburgh: Elsevier; 2013.)
18,27
(From Adams MA, Bogduk N, Burton
based on assumptions and testing protocols that are not necessarily true under in vivo conditions.86 A group investigating
in vivo segmental kinematics using dynamic biplanar radiography recently showed that cervical spine patients followed 1
year postoperatively aer a fusion surgery at C5–C6 showed
no increase in adjacent-level motions, contrary to the ndings
of in vitro studies.
78,79
Rather, they observed that, during
exion-extension, a redistribution of adjacent-level motion
occurred with more extension motion, less exion occurring
at segments rostral to the fusion, and more posterior translations rostral and caudal to the fusion. ese dierences in
motion may be attributed to iatrogenic factors such as alteration in neutral sagittal alignment and disruption of the anterior
longitudinal ligament during fusion.79 It appears also that,
under in vivo conditions, overall ROM of the entire spine
actually decreases aer fusion.
87
110 BASIC SCIENCE
ROM (deg)
Level LevelLevel
Twist
BendFlexion
160
B
Intradiscal pressure (kPa)
2500
A
24
Spinal load (N)
10.0
8.0
6.0
4.0
2.0
0.0
#
23 34 45 51
FIG. 6.21 Range of motion (ROM) of vertebral levels of patients with degenerative disc disease (DDD) and
normal healthy controls. *Signicant dierence within group; #Signicant dierence between normal healthy
controls and patients. (From Passias PG, Wang S. Kozanek M, et al. Segmental lumbar rotation in patients with
discogenic low back pain during functional weight-bearing activities. J Bone Joint Surg Am. 2011;93[1]:29-37.)
2000
1500
1000
500
DDD
Normal
10.0
8.0
6.0
4.0
ROM (deg)
2.0
0.0
23 34 45 51
Standing position
Sitting position
DDD
Normal
#
140
120
100
80
60
40
20
0
10.0
8.0
6.0
4.0
ROM (deg)
2.0
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MildNormal
#
23 34 45 51
Vertical
Horizontal
Moderate Severe
DDD
Normal
Quantitative Assessment of in Vivo Spinal Loading
Accurate in vivo measurement of internal spinal loads requires
placement of a measuring device or sensor invasively into the
region of interest. is would not only pose ethical concerns,
but there is potential risk associated with the implantation
of load sensors in living subjects. Due to these factors, there
are only a few documented studies that have investigated in
vivo spinal loads. e earliest attempt at quantifying in vivo
compressive loads was conducted by Nachemson in 1964,
who measured intradiscal pressures using a needle-mounted
pressure gauge.
other groups in 1999.
summarized as follows: the lowest compressive loads were seen
when lying down (144–250 N), standing upright showed loads
of 500 to 800 N, and sitting erect was 700 to 996 N.
studies also showed that both forward and backward bending
caused an increase in spinal loads. Fig. 6.22A shows intradiscal pressures measured in vivo from dierent postures. One
of these studies also compared intradiscal pressure (horizontal
and vertical pressures based on orientation of pressure gauge)
with respect to progression of disc degeneration, and found a
signicant reduction in pressure with grade of degeneration.
0
0
246810 12 14 16 18 20 22
Angle of motion segment
FIG. 6.22 (A) In vivo intradiscal pressure at dierent postures. (B) Intradiscal pressure variation with progression
of disc degeneration. (From Sato K, Kikuchi S, Yonezawa T. In vivo intradiscal pressure measurement in healthy
individuals and in patients with ongoing back problems. Spine. 1999;24[23]:2468-2474.)
88,89
Similar attempts were made again by two
90,91
e results from these studies can be
88–91
ese
P < .0001
Grade of disc degeneration
Fig. 6.22B shows reduction in intradiscal pressure with disc
degeneration.
91
Rohlmann and colleagues implanted telemeterized verte-
bral body replacements (VBRs) on ve patients with L1 or L3
compression fractures and measured the spinal loads in their
anterior spinal column. Using this setup, they investigated the
eect of locomotion on spinal loads and found that walking
caused signicantly higher loads than standing.92 ey also
found that ascending stairs caused higher loads than descending (Fig. 6.23).
ey also conducted a longitudinal study and observed 10
everyday activities that caused signicant increases in spinal
loads.93 Fig. 6.24 shows 10 activities that caused the highest
increase in spinal loads (compression and shear forces) for ve
patients. ey observed large individual variations in loads for
the various activities.
In Silico Modeling in the Spine
e structural architecture of the human spine exhibits a
hierarchical organization spanning from the whole system
level (macroscale), to the organ, tissue, and cellular levels
Chapter 6 Biomechanics of the Spinal Motion Segment 111
500
Patient
Resultant force (% STG)
(microscale). Within this complex organization, there is a
network of biologic and mechanical interactions between the
dierent levels that dictates overall biomechanical responses
of the spine. Unfortunately, it is extremely dicult to obtain
biomechanical parameters, such as internal stress and strain
distributions, especially at lower spatial scales (cellular).
is knowledge would improve our understanding of the
complex micromechanical environments in relation to normal
structure–function relationships as well as the underlying
mechanisms behind structural and functional breakdown due
to disease.
In silico models, more commonly known as computational
or biomechanical models, are seeing an increased utilization
in spine-related research for investigating complex mechanobiologic phenomena. ese models provide a viable and
practical alternative to relate the physical and material characteristics of the spine to its mechanical function. Using
advanced numerical and imaging techniques, detailed anatomic and material representations of each hierarchical level
(macroscale to microscale) can be developed and used for
biomechanical evaluations.
enables whole body level simulations of spinal kinematics to
be used to predict spinal loads within each segment, and then
quantify tissue- and cellular-level stresses and strains. ese
models have the exibility of precisely controlling a variety of
parameters, then observing the eects of these changes on the
biomechanical response of the modeled structures. ey
provide a unique platform to complement in vitro and in vivo
experimental techniques.
Within the spine, there are several areas of application for
in silico models; for instance, mechanical loads are believed to
play a major role in the initiation of degenerative changes in
the disc. However, the underlying mechanisms by which onset
of damage occurs is not well understood.
how whole body level mechanical loads translate to deformations at the cellular levels that lead to localized damage and the
onset of a degenerative cascade. Using in silico models, we can
450
400
350
300
250
200
150
100
50
0
WP1 WP2 WP3 WP4 WP5
FIG. 6.23 Comparison of peak resultant forces while ascending stairs, descending stairs, and level walking
normalized to standing.92 STG, percentage relative to standing. (From Rohlmann A, Pohl D, Bender A, et al.
Activities of everyday life with high spinal loads. PloS One. 2014;9[5].)
Ascending stairs
Descending stairs
Level walking
now begin to explore these complex mechanical relationships
across spatial and temporal scales. It also paves the way for
the exploration of various other mechanobiologic scenarios,
such as age-related degeneration, eect of endplate microfractures, nutrient transport, and tissue remodeling and repair.
is whole systems approach to predict the impact of spinal
loads on the mechanical behavior of the spine would provide
invaluable information for the development of appropriate
preventive and therapeutic strategies against back injury.
Clinically, in silico models demonstrate a great potential to
aid clinicians in the management of complex spinal ailments.
Patient-specic computational models can be developed for
use in presurgical planning and evaluation, and an optimized
therapy can be implemented for the patient. ese models can
also be used for conducting comparative analysis of spinal
implants.
99,100,103,104
In essence, an in silico model can serve as
a valuable, cost-eective tool for the modication of existing
implants or the design of new spinal implants aimed at stabilizing and/or preserving motion. Spinal stability following a
surgical intervention on an individual can be simulated and
94,95
is multiscale approach
evaluated, providing the clinician with valuable insight and
quantiable metrics to help guide the decision-making process
prior to actual implementation of desired course of treatment.
Knapik and colleagues investigated the biomechanical consequences of a total articial disc replacement (TDR) at L5–S1
under various simulated dynamic loading conditions obtained
from real-life task performance, such as forward bending and
liing of dierent weights (9.5 and 19 kg).
96–100
mechanical stress distribution following a TDR.
found a signicant increase in spinal loads between intact and
TDR at insertion level (Fig. 6.26). ey also found that motion
increased in all three planes (sagittal, lateral, and twisting) at
insertion level. Fig. 6.27 shows sagittal motion across lumbar
18,101,102
It is not clear
levels as a function of intact, TDR, and external loading. eir
model was able to eectively show in detail the biomechanical
trade-os with TDR specic to that subject’s spine under
realistic loading conditions.
105
Fig. 6.25 shows
105
eir study
SECTION
I
112 BASIC SCIENCE
B
150
Arm elevation with weight in hands
A
1750
Lifting weight from ground
Moving weight in front of body
Standing up/sitting down
Staircase walking
Tying shoes
Upper body flexion
Lifting a carried weight
Washing face
Moving from lying to sitting
Walking
Arm elevation with weight in hands
Lifting weight from ground
Moving from lying to sitting
Tying shoes
Upper body flexion
Carrying weight in hands
Staircase walking
Upper body extension
Moving arms laterally in circles
Standing up/sitting down
Axial rotation
Lateral bending
Washing face
0 250 1250500 1500750
Maximum resultant force (N)
WP5
WP4
WP3
WP2
WP1
1000
The System
As can be seen through this review, the spine performs several
important functions: it transmits force, allows motion, and
protects the spinal cord. Although these functions have been
considered independently here, it is important to develop an
appreciation for the systematic nature of these spine functions.
Although these functions have been described independently,
they interact in such a way that the inability to perform one
of these functions can also aect the ability to perform other
functions.
–250 –200 –150 –100 –50 050 100
Range of shear force (N)
FIG. 6.24 (A) Ten activities with highest compressive forces. (B) Anterior (positive) and posterior (negative)
shear forces for the 10 activities.93 WP, patient label. (From Rohlmann A, Dreischarf M, Zander T, et al. Loads on a
vertebral body replacement during locomotion measured in vivo. Gait Posture. 2014;39[2]:750-755.)
If the disc becomes compromised in its mechanical integrity,
and disc space is reduced, it can alter the load transmission
between vertebrae. With less disc space, more of the load may
be transmitted through the posterior elements; this repeated
loading may change the biochemical behavior of the system.
is change may result in an upregulation of proinammatory biochemical activity and increased pain transmission.
Similarly, reduced disc space height may alter the motion
characteristics of the spinal motion segments. With less disc
space, the stability of the joint can be compromised, and
the contact points of the posterior elements can be altered.