Добавил:
Sekretar
kiopkiopkiop18@yandex.ru
t.me/Prokururor I Вовсе не секретарь, но почту проверяю
Опубликованный материал нарушает ваши авторские права? Сообщите нам.
Вуз:
Предмет:
Файл:Ординатура / Хирургия / Библиотека им академика М.И. Перельмана / Книга_6034_Библиотеки_им_академика_М_И_Перельмана.pdf
CHAPTER 5/CLINICAL SPINAL INSTABILITY / 53
different animals, three graded spinal injuries (interspinous and supraspinous ligament transections, laminectomy, and facetectomy) at the C4-5 level were studied by
functional flexion-extension stereoradiographs for up to
24 weeks (17,18). In these in vivo animal experiments,
contrary to expectations, the spine at the injury site
became more stable (even compared with the intact
spine) as measured by standardized functional X-ray
studies during the healing period (Fig. 5-3). Although the
facetectomy resulted in the largest increase in motion
acutely, it also produced the largest decrease in motion in
vivo. At 6 weeks after the injury, the range of motion
(ROM) decreased from 23 degrees preoperatively to 5
degrees postoperatively. These findings are suppor ted by
studies using a canine (19) and a porcine model (20). In
the later study, explanation was provided for the
decreased motion. At 3 months post facetectomy, we
found hypertrophy of the facet joints, which limited the
range of motion.
Degeneration
For an in-depth description of degeneration, please see
Chapter 3.
The spine degenerates with age; this is a normal
process that results in altered mechanical characteristics.
It also may lead to low back problems. Kirkaldy-Willis
(6) provided a classification of degeneration of the spine
based on three stages.
Stage 1. Dysfunction. This includes low back pain with
nonspecific syndrome. The facet capsule may be lax
and disc degeneration is of grade 1 to 2 on a scale of 1
to 4.
FIG. 5-3. Average intervertebral range of motion at the injury
site as a function of healing time. The injury was bilateral
facetectomy at C4-5 in a canine model. The injury was
unprotected during the entire healing phase, and the motion
measurements were made using functional flexion-e xtension
stereoradiographs.(From Panjabi M, P elk er K, Crisco J, et al.
Biomechanics of healing of posterior cervical spinal injuries
in canine model. Spine 1988;13:803–807.)
Stage 2. Instability. This is marked by increased facet
joint laxity and moderate disc degeneration (grades 2
to 3). Clinical syndrome can be identified, and the
instability can be measured by functional X-ray studies.
Stage 3. Restabilization. This is characterized by fibrosis
in posterior joints and osteophytic formations leading
to decreased overall motion. Disc generation has
reached the final stage (g rades 3 to 4).
A recent study has confirmed the biomechanical
aspects of the preceding hypothesis using an intraoperative instrumented lamina spreader (21). It consisted of an
electric motor, which spread the adjacent laminae, and
the strain gauges attached to the spreader legs, which
measured the force applied. Based on a study of nearly
300 patients and 650 FSUs intraoperatively, we conclude
the following. The average stiffness reaches its peak of
120 N/mm at about 25 years of age, decreases thereafter
to less than 20 N/mm at about 55 years, and then
increases once more to about 80 N/mm above the age of
60 (Fig. 5-4). The stiffness seems to have an inverse relationship to the disc degeneration and range of motion.
The degeneration effect on the mechanical properties
of the spine is specific and direction dependent. Several
parameters may be obtained from load-displacement
curves of a lumbar spine specimen to quantify the
mechanical properties. These are: the neutral zone (NZ),
representing “looseness” of the specimen; the elastic
zone (EZ), which may equate with elastic deformation;
and the range of motion (Fig. 5-5). Another parameter is
the neutral zone ratio (NZR), equal to NZ divided by
ROM. In a study using fresh cadaveric lumbosacral spine
specimens, intervertebral flexion-extension, lateral bending, and axial rotation were measured and plotted against
disc degeneration grade (22). In flexion-extension, there
was some tendency for ROM to decrease and NZ to
increase. The lateral bending showed significant decrease
in ROM and significant increase in NZR. In axial rotation
there were significant changes in ROM, which decreased,
and in NZ and NZR, both of which increased.
The preceding knowledge has been obtained mostly
from in vitro experiments. The general degeneration of
the spine, seen on X-ray films as decreased disc height,
deformed end plates, and osteophyte formation, has not
been found to be a reliable predictor of subsequent low
back pain. On the other hand, evidence suggests that
increased disc degeneration carries a signif icantly higher
risk of low back problems (23). During discography , 23%
of patients with nondegenerated discs reported pain, and
77% felt either pressure only or no pain at all. On the
other hand, among patients with a severely degenerated
annulus, 90% reported pain during discography, whereas
only 10% felt no pain or simply some pressure. Thus, a
significant relationship seems to exist between disc
degeneration and low back pain, even though it may not
54 /SECTION I/BASIC SCIENCE
FIG. 5-4. Functional spinal unit (FSU) stiffness distribution with age. (From Brown MD, Holmes DC,
Heiner AD, et al. Intraoperative measurement of lumbar spine motion segment stiffness. Spine
2002;27(9):954–958.)
be a one-to-one correspondence. This and similar in vivo
studies provide a link to the in vitro biomechanical stud-
ies, by which the mechanical characteristics of the spine
and the clinical symptoms of low back pain may be
related.
Role of Spinal Muscles
The understanding of the primary role of musculature
in providing spinal stability and the extent to which the
FIG. 5-5. Load-displacement curve of a spine specimen.The
measurements for neutral zone (NZ), elastic zone (EZ), and
range of motion (ROM) are obtained from the curve.
musculature contributes to pain production, modulation,
and prevention is not well understood. Muscle dysfunction may result from muscle weakness, in the form of
decreased strength or endurance, and possibly from a disturbance in the neuromuscular control system, in the
form of altered recruitment patterns. Muscle spasm and
pain may be indicators of muscular overload owing to the
reduced efficiency in weakened passive structures of the
spinal system. As described, muscles form an important
subsystem of the overall spinal stabilizing system.
A lumbo-sacral (L1-sacrum) spinal column that is
devoid of musculature is a mechanically unstable structure, with a load-carrying capacity of less than 90 N (or
20 pounds) (24). However, with properly coordinated
muscle action, the spine can sustain large loads, which is
exemplified by the action of weight lifters. In the past, the
complexity of the muscular anatomy and physiology hindered the development of biomechanical models for
studying the stabilizing role of muscle, as well as various
passive components of the spine (e.g., ligaments, discs,
vertebrae, and facet joints). Detailed morphologic and
biomechanical analyses of the lumbar musculature are
now available (25–27). The spinal muscles may be conceptualized as local (intersegmental) and global (multisegmental), which helps us to understand their functions
of stabilizing the spine and producing motion (28,29).
Advanced mathematical models are helping us to better
understand the instability (30).
A modeling study based on radiographs from normal
subjects was performed to determine the effects of flexion on the forces exerted by the lumbar muscles (27). The
CHAPTER 5/CLINICAL SPINAL INSTABILITY / 55
act of flexing caused substantial elongation of many muscle fascicles, which consequently reduced the maximum
active tension they could exert. Consequently, it was
found that the compressive forces and moments exerted
by the back muscles in full flexion are not signif icantly
different from those in the upright posture. However,
major changes in shear forces were found, particularly at
L5-S1, where there was a re versal from a net anterior to a
net posterior force. These shear forces must be considered when prescribing therapeutic exercise, particularly
in patients with translatory instability in the lower lumbar
and lumbosacral region.
Using an anatomically detailed biomechanical model,
the role of the lumbar erector spinae musculature in offsetting the anterior shear forces on L4-5 (58 to 324 N)
and upper body mass during different dynamic lifts
(squat and stoop) were studied (31,32). They found that,
during the squat lift, the maximum peak forces supported
by the f acet joints and possibly the disc remained relatively constant at approximately 200 N, regardless of the
load mass. When comparing the two different lifting
styles, the stoop lift, which produced a more flexed lumbar spine than did the squat lift, and had greater contributions from the passive lumbar structures (e.g., ligamentous strain), although the peak moments provided by
these tissues were less than 60 nm.
The effects of simulated intersegmental muscle forces
on spinal instability in an in vitro experiment have been
investigated (33). In flexion loading, range of motion
increased and neutral zone decreased with the application
of muscle forces, whereas both variables decreased in
extension loading. Similar observations have been made
in an in vivo investigation using a porcine model to study
alterations in segmental kinematics as a result of injury to
the passive stabilizing components and stimulation of the
lumbar musculature (34). When compared with the unstimulated situation, stimulation of the paraspinal muscles produced significantly greater range of motion in
sagittal rotation and shear translation in the L3-4 motion
segment after injuries to the disc or facet joints. Although
it increased the range of motion, the increased muscular
activity also stabilized the injured motion segment. This
stabilization was indicated by a reduction in the abrupt
changes in the pattern of motion for sagittal rotation during the transitional phase between dynamic flexion and
extension (neutral region).
Electromyographic signals of the paraspinal and abdominal muscles have been studied both in normal subjects and in patients with low back pain. Some studies
have sho wn that the electromyographic patterns displayed
some abnormalities in patients with low back pain compared with the normal group (35–37). Also, the flexionrelaxation phenomenon of the erector spinae muscle
group is absent in some patients with acute low back pain
but returns after the pain has gone. The flexion-relaxation
phenomenon is the myoelectric silence at approximately
two thirds of maximum flexion angle, at which the load
moment is carried by the soft tissues (e.g., ligaments, fascia, and passive elongated muscle) (38,39). It is now
believed that intra-abdominal pressure stabilizes the
spine (40).
However, studies show diverging results as to whether
increased intra-abdominal pressure loads or unloads the
spine (30,41,42). The muscles not only apply loads and
provide stability, but also help control the posture and
movement (5). In a study of low back patients and healthy
controls, patients demonstrated poorer balance control
while sitting on an unstable hemisphere and had longer
reaction times to sudden horizontal loadings (43).
DIAGNOSTIC METHODS
Roentgenographic Motion Studies
Besides the grades of disc degeneration, which are
related to a greater risk of low back pain, other motion
and posture measures can be obtained from radiographs
or computed tomography or MRI images. Functional
radiographs (e.g., a pair of radiographs taken, generally,
at the extremes of a motion in a certain plane) form the
basis of most clinical studies of motion. Knutsson (44)
was probably the f irst to indicate a relationship between
excessive anteroposterior translation seen on flexionextension radiographs and low back problems. In another
study, patients with low back pain were examined in lateral bending, and centers of rotation were calculated for
various positions of the lumbar spine. An increased area
occupied by the locus of the centers of rotation at a particular level w as found to be directly related to the pain at
that level (45). In another study, motions were measured
from lateral radiographs taken in three specified postures
(46). Normal patients were found to be different from the
patients with spondylosis in translation and rotation and
in coupling between these motions.
The spinal movements of patients with low back pain
who are suspected of having instability ma y not al w a ys be
greater in magnitude. It is known clinically, quantif ied
using stereoradiographic analysis, that patients with low
back pain have restricted flexion-extension intervertebral
motion. The total flexion (L1-S1) of about 50 degrees in
normal individuals decreases to less than 20 degrees in
patients with low back pain and nerve root tension signs
(47). Associated with the restricted flexion-extension
spinal motion are increased coupled motions (i.e., lateral
bends and axial rotations). The coupled motion is defined
as the associated motion produced during the main
motion (e.g., lateral bending or axial rotation produced
during flexion). Theoretically, there are up to f ive coupled motions for every main motion. Both observations
may be explained by the fact that spinal instability
resulted in activation of the muscular system. Increased
muscle forces restricted the overall motion of the spine
56 /SECTION I/BASIC SCIENCE
and at the same time, owing to muscle imbalance,
resulted in asymmetric spinal movements (e.g., out-ofsagittal plane coupled motion during flexion-extension).
Functional flexion-extension X-ray studies were performed passively on a patient population that was subdivided into different groups having similar pathologic
conditions (48). When compared with a normal population, all patients exhibited less motion, except for highperformance athletes who showed more motion compared to the controls. Therefore, it was concluded that a
kinematic analysis of the lumbar spine using passiv e flexion-extension was not a clinically useful method.
Inferior-superior loading using functional X-ray e xaminations also has been investigated as a measure of spinal
instability (49,50). The motion was measured at two
extremes of motion obtained by (a) spinal traction (suspending the individual from his or her hands); and (2)
compression (using a weighted backpack during standing). Anteroposterior translation measurements were
taken from lateral X-ray films of patients who had
spondylolisthetic or retrospondylolisthetic displacement.
In accordance with the severity of symptoms, the patients
were divided into the following groups: (a) asymptomatic
patients; (b) those with moderate symptoms and (c)
patients with severe symptoms. The degree of primary
anterior slip was almost equal in the three groups, but the
translator movement differed signif icantly among them,
as follows: 0.7, 5.2, and 7.5 mm, respectively.
Stereoradiographic techniques have been used to analyze three degrees-of-freedom sagittal plane motion
(sagittal rotation, antero-posterior translation, and inferior-superior translation) in patients with low back pain
and suspected segmental spinal instability (51). The average angular ROM in patients at the unaf fected level (9.67
degrees) was not different from that at the affected level
(8.45 degrees). The same was true for antero-posterior
shear translation values, which were 1.54 and 0.92 mm,
respectively. However, the ratio (i.e., coupled shear translation divided by the flexion angle) was significantly different (+0.18 versus −0.13 mm/degree) at the unaffected
and affected levels, respectively. The retrodisplacement
(anterior-to-posterior translation during flexion from
extended position) was associated with the restricted
motion, especially for sagittal plane rotations of less than
5 degrees, but was not correlated with the specif ic clinically unstable levels.
In a recent study, three-dimensional coupled motions
were measured in low back pain patients (52). The
patients were asked to move in three planes (sagittal,
transverse, and frontal) while the intervertebral motions
of pedicle screws inserted into the vertebrae above and
below the suspected painful level were measured. During
flexion-extension, there w ere small out of plane rotations.
During axial rotation there was considerable v ariability in
the coupled motions. The same was true for the lateral
bending. The authors concluded that in contrast to well-
defined in vivo and in vitro coupling patterns observed in
the controls, the low back pain patients showed significantly greater variability. The inherent coupling pattern
of the osseoligamentous spine was modified by the altered muscle pattern or pain.
Other Measures of Instability
Measurement of ROM, especiall y flexion-extension, is
easy in vivo. For this reason, the ROM has been used
often as an indicator of instability (4). Unfortunately, the
ROM is not related to clinical instability, as exemplified
by a young gymnast who may have extensive ROM but
no clinical symptoms of instability (3). Further, the measurement of the ROM is affected by voluntary effort that
the subject applies at the time of examination and motion
limitation because of pain. Thus, investigating other measures of motion as possible indicators of instability has
merit.
One such variable is the neutral zone, w hich represents
looseness of the spinal column around the neutral position. Support for the coupled motions concept is provided
by an in vivo study, which documented the presence of
these motions in patients with suspected clinical instability (47). The neutral zone has been studied only in vitro.
The increase in the neutral zone was found to be associated with disc degeneration and its decrease was related
to simulated muscle force application (22,33). No direct
clinical evidence is yet available. Because both measures
are generally smaller in magnitude compared with the
ROM, new and more accurate diagnostic methods are
needed. In a recent study using ultrasound Doppler effect,
the neutral zones of the sacroiliac joint have been measured in subjects without pain (53). Future studies with
low back pain patients using this technique will be interesting to see if the neutral zone concept is clinically
useful.
Using an intervertebral motion device for continuously
measuring sagittal plane motion in the human lumbar
spine, the intervertebral motion, along with the overall
trunk angle, was measured dynamically during standing
flexion-extension, both in normal subjects in patients suspected of having clinical instability in a lumbar motion
segment (39). There exists a characteristic pattern of
motion during flexion-extension for normal lumbar
motion segments and patients (Fig. 5-6). The main f indings were the follo wing. Motion w as significantly less, by
at least 50%, in patients compared to the controls. A 78%
reduction in muscle activity at full flexion (flexion relaxation) occurred in controls, whereas only a 13% reduction was found in patients. These observations were
explained by hypothesizing that the neuromuscular control system provides acti v e stabilization needed to protect
the injured or diseased passive structures from movements that may cause pain, similar to the stabilization
concepts proposed by Panjabi (5).
FIG. 5-6. Segmental kinematics (sagittal rotation) and myoelectric experimental data during a flexion-extension (F-E)
cycle from the L4-5 motion segment plotted as a function of
trunk F-E angle for a control and a patient. (Neutral standing
position = trunk F-E angle = 0 degrees. Root mean square of
the right-side erector spinae myoelectric activity (RMS
EMG).) (From Kaigle AM, Wessberg P, Hansson T. Muscular
and kinematic behavior of the lumbar spine during flexionextension. J Spinal Disord 1998;11(2):163–174.)
TREATMENTS
Spinal instability is treated clinically by diverse conservative methods, some of w hich seem to be parado xic.
Both the flexion exercises, which strengthen abdominal
muscles, and the extension exercises, which strengthen
back muscles, have been effective (54). To increase
spinal stability co-contraction of both the front and back
muscles is needed (55). This may be the explanation for
the effectiveness of both the flexion and extension exercises. Rotational exercises have been found to be effective in patients who did not respond to other treatments
(56). In addition to strengthening the spinal muscles,
improving muscle coordination is important in enhancing spinal stability (5). Muscle stabilization has been
advocated and shown to be effective in treating back
pain patients (57). Various fusion techniques are reported to have clinical success (3).
CHAPTER 5/CLINICAL SPINAL INSTABILITY / 57
FUTURE RESEARCH
Several aspects of spinal instability need to be investigated from the biomechanical viewpoint. A short list is
provided.
1. By means of in vitro simulations (using human
cadaveric material), in vivo animal models, and
mathematical models, investigate the role of intersegmental (deep) as well as multisegmental (superficial) muscles in providing spinal stability.
2. Develop techniques that measure the dynamic intervertebral motion continuously.
3. Using in vivo animal models, study the role of heal-
ing and adaptation after injury in altering the spinal
stability.
4. Develop new and more accurate diagnostic methods
for determining abnormalities of coupled motion,
neutral zones, and other motion variables, w hich ma y
help to provide more sensitive and specific measurements of spinal instability than are presently available.
5. Conduct clinical studies (prospective, double b lind , and
controlled) that correlate carefully obtained measures
of instability of intervertebral motions (representing
spinal column) and muscle function (representing neuromuscular control) with the clinical symptoms. These
studies may help to bridge the gap between instability
indicators and clinical symptoms.
REFERENCES
1. White A, Gordon S. Synopsis: workshop on idiopathic low-back pain.
Spine 1982;7:141–149.
2. Moran FP, King T. Primary instability of lumbar vertebrae as a common cause of low back pain. J Bone Joint Surg 1957;39B:6–22.
3. White AA, Panjabi MM. The clinical biomechanics of the spine.
Philadelphia: JB Lippincott, 1990.
4. Nachemson A. A critical update and symposium summary. Spine 1985;
10:290–291.
5. Panjabi MM. The stabilizing system of the spine. Part I. Function, dysfunction, adaptation, and enhancement. J Spinal Disord 1992;5(4):
383–389.
6. Kirkaldy-Willis WH. Managing low back pain. New York: Churchill
Livingstone, 1983.
7. Nakana T, Nakano K, Nakano N. Does the disc narrowing of the lumbar spine cause symptoms? Proceedings of the International Society
for the Study of the Lumbar Spine. Boston, 1990.
8. Macnab I. The traction spur: an indicator of segmental instability. J
Bone Joint Am 1971;53(4):663–670.
9. Farfan HF, Cosette IW, Robertson GH, et al. The effects of torsion on
the lumbar intervertebral joints: The role of torsion in the production of
disc degeneration. J Bone Joint Surg 1970;52A:468–497.
10. Bongers P, Boshnizan H. Back disorders and whole body vibration at
work. Amsterdam: Akademisch Proefschrift, University of Amsterdam.
1990.
11. Abumi K, Panjabi MM, Kramer KM, et al. Biomechanical evaluation
of lumbar spine stability after graded facetectomies. Spine 1990;15:
1142–1147.
12. Posner I, White AA, Edwards WT, et al. A biomechanical analysis of
the clinical stability of the lumbar and lumbosacral spine. Spine 1982;
374–388.
13. Hirsch C. The mechanical response in normal and degenerated lumbar
discs. J Bone Joint Surg 1956;38A:242.
14. Markolf KI, Morris JM. The structural components of the intervertebral disc: J Bone Joint Surg 1974;56A:675.
58 /SECTION I/BASIC SCIENCE
15. Goel CK, Nishiyama K, Weinstein JN, et al. Mechanical properties of
lumbar spinal motion segments as affected by partial disc removal.
Spine 1986;11:1008–1012.
16. Panjabi MM, Krag MH, Chung TQ. Effects of disc injury on mechanical behavior of the human spine. Spine 1984;9(7):707–713.
17. Panjabi MM, P elk er K, Crisco J, et al. Biomechanics of healing of posterior cervical spinal injuries in a canine model. Spine 1988;13:
803–807.
18. Wetzel FT, Panjabi MM, Pelker RR. Biomechanics of the rabbit cervical spine as a function of component transection. J Orthop Res 1989;7:
723–727.
19. Whitehill R, Stowers SF, Fechner RE, et al. Posterior cervical fusions
using cerclage wires, methylmethacrylate cement and autogenous bone
graft. An experimental study of a canine model. Spine 1987;12:12–22.
20. Kaigle AM, Holm SH, Hansson TH. Kinematic behavior of the porcine
lumbar spine: a chronic lesion model. Spine 1997;22(24):2796–2806.
21. Brown MD, Holmes DC, Heiner AD, et al. Intraoperative measurement
of lumbar spine motion segment stiffness. Spine 2002;27(9):954–958.
22. Mimura M, Panjabi MM, Oxland TR, et al. Disc degeneration affects
the multidirectional flexibility of the lumbar spine. Spine 1994;19(12):
1371–1380.
23. Vanharanta H, Sachs BL, Spivey MA, et al. The relationship of pain
provocation to lumbar disc deterioration as seen by CT/discography.
Spine 1987;12:295–298.
24. Crisco JJ, Panjabi MM, Yamamoto I, et al. Euler stability of the human
ligamentous lumbar spine: Part II. Experiment. Clin Biomech 1992;7:
27–32.
25. Bogduk N, Macintosh JE, Pearcy MJ. A universal model of the lumbar
back muscles in the upright position. Spine 1992;17:897–913.
26. Macintosh JE, Bogduk N. The biomechanics of the lumbar multifidus.
Clin Biomech 1986;1:205–213.
27. Macintosh JE, Bogduk N, Pearcy M. The effects of fle xion on the geometry and actions of the lumbar erector spinae. Spine 1993;18:884–893.
28. Bergmark A. Stability of the lumbar spine. A study in mechanical engineering. Acta Orthop Scand Suppl 1989;230:1–54.
29. Crisco JJ, Panjabi MM. The intersegmental and multisegmental muscles of the lumbar spine: a biomechanical model comparing lateral stabilizing potential. Spine 1991;16(7):793–799.
30. Cholewicki J, VanVleit JJ. Relati ve contribution of trunk muscles to the
stability of the lumbar spine during isometric exertions. Clin Biomech
2002;17(2):99–105.
31. Potvin JR, McGill SM, Norman RW. Trunk muscle and lumbar ligament contributions to dynamic lifts with varying degrees of trunk flexion. Spine 1991;16:1099–1107.
32. Potvin JR, Norman RW, McGill SM. Reduction in anterior shear forces
on the L4-L5 disc by the lumbar musculature. Clin Biomech 1991;6:
88–96.
33. Panjabi MM, Abumi K, Duranceau J, et al. Spinal stability and intersegmental forces—a biomechanical model. Spine 1989;14:194–199.
34. Kaigle AM, Holm SH, Hansson TH. Experimental instability in the
lumbar spine. Spine 1995;(4):421–430.
35. Floyd WF, Silver PHS. The function of erector spinae muscles in certain movements and postures in man. J Physiol 1955;129:814–203.
36. Sihvonen T, Partanen J, Hänninen O, et al. Electric behavior of low
back muscles during lumbar pelvic rhythm in low back patients and
healthy controls. Arch Phys Med Rehabil 1991;72:1080–1087.
37. Triano JJ, Schulz AB. Correlation of objectiv e measure of trunk motion
and muscle function with low-back disability ratings. Spine 1987;12:
561–565.
38. Kippers V, Parker AW . Posture related to myoelectric silence of erectors
spinae during trunk flexion. Spine 1984;9:740–745.
39. Kaigle AM, Wessberg P, Hansson T. Muscular and kinematic behavior
of the lumbar spine during flexion-extension. J Spinal Disord 1998;11
(2):163–174.
40. Hodges PW, Cresswell AG, Daggfeldt K, et al. In vivo measurement of
the effect of intra-abdominal pressure on the human spine. J Biomech
2001;34:347–353.
41. Daggfeldt K, Thorstensson A. The role of intra-abdominal pressure in
spinal unloading. J Biomech 1997;30:1149–1155.
42. Gracovetsky S, Farfan H, Helleur C. The abdominal mechanism. Spine
1985;10:317–324.
43. Radebold A, Cholewicki J, Polzhofer GK, et al. Impaired postural control of the lumbar spine is associated with delayed muscle response
times in patients with chronic idiopathic low back pain. Spine 2001;26
(7):724–730.
44. Knutsson F. The instability associated with disc degeneration in the
lumbar spine. Acta Radiol 1944;25:593–609.
45. Dimnet J, Fischer LP, Gonon G, et al. Radiographic studies of lateral
flexion in the lumbar spine. J Biomech 1978;11:143–150.
46. Keeson W, During J, Beeker TW, et al. Recordings of the movement at
the intervertebral segment L5-S1: a technique for the determination of
the movement in the L5-S1 spinal segment by using three specif ied
postural postures. Spine 1984;9:83–90.
47. Pearcy M, Portek I, Shepherd J. The effect of low-back pain on lumbar
spinal movements measured b y three-dimensional x-ra y analysis. Spine
1985;10:150–153.
48. Dvorák J, Panjabi MM, Novotny JE, et al. Clinical validation of functional flexion-extension roentgenograms of the lumbar spine. Spine
1991;16:943–950.
49. Friberg O. Lumber instability: a dynamic approach by traction-compression radiograph. Spine 1987;12:119–129.
50. Käleho P, Kadziolka R, Sward L, et al. Stress views in the comparative
assessment of spondylolytic spondylolisthesis. Skel Radiol 1989;17:
570–575.
51. Stokes IAF, Wilder DG, Frymoyer JW, et al. Assessment of patients
with low back pain by biplanar radiographic: measurement of intervertebral motion. Spine 1981;6:233–240.
52. Lund T, Nydegger T, Schlenzka D, et al. Three-dimensional motion patterns during active bending in patients with chronic low back pain.
Spine 2002;27(17):1865–1874.
53. Buyruk HM, Snijders CJ, Vleeming A, et al. The measurements of
sacroiliac joint stiffness with colour Doppler imaging: a study of
healthy subjects. Eur J Radiol 1995;87(2):127–133.
54. Williams PC. Examination and conservation treatment for disc lesions
of lower spine. Clin Orthop 1955;5:28–35.
55. Cholewicki J, Ivancic PC, Radebold A. Can increased intra-abdominal
pressure in humans be decoupled from trunk muscle co-contraction
during steady state isometric exertions? Eur J Appl Physiol 2002;21:
117–121.
56. Polermo K, Panjabi MM. Role of trunk rotation endurance exercise in
failed back treatment. Arch Phys Med Rehabil 1986;67:620.
57. Hides JA, Jull GA, Richardson CA. Long-term effects of specif ic stabilizing exercises from first-episode low back pain. Spine 2001;26(11):
E243–248.
CHAPTER 6
Spinal Instrumentation
Vijay K. Goel, Manohar M. Panjabi, Huroshi Kuroki, Setti S. Rengachary,
D. McGowan, and N. Ebraheim
In recent years, surgeons have well-accepted surgical stabilization and fusion of the spine using instrumentation.
Accordingly, the number of available devices for use by a
surgeon has increased (1–4). The types and complexity of
procedures (e.g., posterior, anterior, interbody) (3) have
produced novel design challenges, requiring sophisticated testing protocols (3). In addition, most contemporary implant issues of stabilization and fusion of the spine
are mostly mechanical in nature (4). [Biologic factors
related to the adaptive nature of living tissue further
complicate mechanical characterization (3,5,6).] Accordingly, researchers have designed various methods of testing to assess the mechanical nature of the spine and
implants, both as separate and united entities. These evaluation regimens have produced valuable information and
have led to the design and de v elopment of state-of-the-art
systems. The most eff icient way to describe the biomechanical issues relating to stabilization and fusion in the
thoracolumbar region is to group the literature that concerns the major testing modalities. Results of specific
studies are presented to show the type of information provided by the various testing methods.
CLINICAL SCOPE AND OBJECTIVE OF SPINAL
FUSION
Low back pain is responsible for approximately 14%
of visits to physicians that do not in v olve pree xisting conditions (2). Others have estimated 70% of the population
in the United States has experienced back pain in their
lives, leading to surgical intervention of the lumbar spine
in 4% of the population (2). Surgical treatments most
often promote fusion of the painful segments, with an
estimated 25% of the 280,000 operations involving the
lumbar spine (4,7).
The objective of spinal fusion is to eliminate pain and
allow the patient to resume normal activities. Elimination
of relative motion between the affected joints often
reduces this type of segmental pain. Spinal fusion is also
performed to prevent or correct deformity (3) and stabilize the spine after trauma. Pathologic degeneration of the
bony elements, intervertebral disc, and soft tissues are
also indicators for fusion (3). Although intervertebral
disc conditions seem to account for a significant proportion of the lesions leading to fusion, other indications
include segmental instability, both degenerative and
iatrogenic, and failed previous surgery. Although the
aforementioned indications are commonly cited in the literature as grossly appropriate, there is considerable
debate as to the degree of the lesion that indicates fusion
(8).
Properly applied, spinal instrumentation maintains
alignment and shares spinal loads until a solid, consolidated fusion is achieved. As instrumentation procedures
have become increasingly popular, the number of available fixation systems has grown. With few exceptions,
these hardware systems are used in combination with
bone grafting procedures, and may be augmented by
external bracing systems.
BIOMECHANICAL EVALUATION OF
INSTRUMENTATION PERFORMANCE
Spinal implants typically follow loosely standardized
testing sequelae during the design and development stage
and in preparation for clinical use. The design and development phase goal, from a biomechanical standpoint,
seeks to characterize and define the geometric considerations and load-bearing environment to which the
implant will be subjected. Various testing modalities exist
that elucidate which components may need to be
redesigned. Not including the testing protocols for individual components of a device, plastic vertebrae (corpectomy) models are one of the first-stage tests that involve
placing the assembled device on plastic vertebral components in an attempt to pinpoint which component of the
59
60 /SECTION I/BASIC SCIENCE
assembled device may be the weakest mechanical link in
the worst case scenario, vertebrectomy. The in vivo effectiveness of the device may be limited by its attachment to
the vertebrae (fixation). Thus, testing of the implant-bone
interface is critical in determining the fixation of the
device to biologic tissue. Construct testing on cadaveric
specimens provides information about the effectiveness
of the device in reducing intervertebral motion across the
affected and adjacent segments during quasi-physiologic
loading. Animal studies provide insight with respect to
the long-term biologic effects of implantation. Analytic
modeling, such as the finite element method, is an
extremely valuable tool for determining how implants
and osseous loading patterns change with varying parameters of the device design. This type of modeling may
also provide information about temporal changes in the
bone quality due to the changing loading patterns as bone
adapts to the implant (e.g., stress shielding-induced bone
remodeling). After a certain level of confidence in the
implant’s safety and effectiveness is established through
all or some of the aforementioned tests, controlled clinical trials allow for the determination of an implant’s suitability for widespread clinical use. The following sections discuss each of these testing modalities, with
specific examples used to illustrate the type of infor mation that different tests can provide.
Implant-Bone Interface
Device-Ver tebra Interface
Depending upon the spinal instrumentation, the
device-vertebra interface may deal with laminae, pedicles, the vertebral body itself, or the end plates.
Interlaminar Hooks
Interlaminar hooks are used as a means for fixing the
device to the spine. Hook dislodgment, slippage, and
incorrect placement have led to loss of fixation, however,
resulting in nonfusion and pseudoarthrosis. Purcell et al.
(9) investigated construct stiffness as a function of hook
placement with respect to affected level in a thoracolumbar cadaver model. They created posterior ligamentous
defects through sectioning and imposed bony fracture at
T-12 and L-1 by flexion testing to failure. The unstable
spines were instrumented with Harrington distraction
instrumentation and interlaminar hooks placed initially
on T-11 and L-2. The hooks w ere relocated to v arious levels about the affected area and the construct retested. The
failure moment was found to be a function of the hook
placement. The authors recommended hook placements
three levels above and two levels below the affected area.
This placement reduced vertebral tilting (analogous to
intervertebral motion) across the stabilized area, where
fusion is to be promoted.
Transpedicular Screws
Proper application of screw-based anterior or posterior
spinal devices requires an understanding of screw biomechanics, including screw characteristics and insertion
techniques, as well as an understanding of bone quality,
pedicle and vertebral body morphometries, and salvage
options (10–12). This is best illustrated by the fact that
the pedicle, rather than the vertebral body, contributes
approximately 80% of the stiffness and about 60% of the
pullout strength across the screw-bone interface (10).
Carlson et al. (13) evaluated the effects of screw orientation, instrumentation, and bone mineral density (BMD)
on screw translation, rotation at maximal load, and compliance of the screw-bone interface in human cadaveric
bones. An inferiorly directed load was applied to each
screw, inserted either anteromedially or anterolaterally,
until failure of the f ixation was perceived. Anteromedial
screw placement with fully constrained loading linkages
provided the stiffest f ixation at low loads and sustained
the highest maximal load. Larger rotation of the screws,
an indication of screw-out failure, was found with the
semi-constrained screws at maximal load. BMD directly
correlated with maximal load, indicating that bone quality is a major predictor of bone-screw interfacial strength.
Peif fer et al. and Ryken et al. also found a significant correlation between BMD and torque (p < .0001, r < 0.42),
BMD and pullout force (p < .0001, r < 0.54), and torque
and pullout force (14–16).
Since the specimens used for pullout strength studies primarily come from older adult subjects, Choi et al. used
foams of varying densities to study the effect of BMD on
the pullout strength of several screws (17). Pedicle screws
(6.0 mm × 40 mm, 2 mm pitch, titanium alloy) of several
geometric variations were used for the study. They included
the buttress (B), square (S), and V-shape (V) screw tooth
profiles. For each type of tooth prof ile, its core shape (i.e.,
minor diameter) also varied, either straight (i.e., cylindrical,
core diameter < 4.0 mm) or tapered (i.e., conical, core
diameter < 4.0 mm/2.0 mm). In addition, for the cylindrical
screws the major diameter was kept straight or tapered. The
conical screws had their major diameters tapered only.
Therefore, screws with a total of nine different geometries
were prepared and tested (Fig. 6-1A). Nomenclature used
for identifying each screw type followed this sequence:
tooth profile, the shape of the major diameter, and core
type. For example, BST represents the screw with the buttress tooth profile and straight major diameter on a tapered
core. The screws were implanted in the rigid polyurethane
foams (77 cm × 127 cm × 77 cm) (Sawbones, Pacific
Research Laboratory , Vashon Island , WA) of three different
grades (grades 10, 12, and 15). These grades “simulated”
3
the variations in BMD (10 lbm/ft
3
, respectively) of the cancellous bone of a vertebra.
lbm/ft
, 12 lbm/ft3, and 15
Screws were implanted according to the American Society
for Testing and Materials (ASTM: F1839-97) protocol. The
CHAPTER 6/SPINAL INSTRUMENTATION / 61
FIG. 6-1. A: Different types of screws used in the foam model
to determine the pullout strengths of various designs. The
nomenclature used is as follows: thread shape—square (S),
buttress (B), V-shape (V); screw diameters—straight major
diameter on straight core (SS), straight major diameter on
tapered core (ST), tapered major diameter on tapered core
(TT). B: Regression analysis. The maximum and minimum val-
ues from pullout test for each foam gr ade were used regardless
of tooth or core profiles. (From Choi W, Lee S, Woo KJ, et al.
Assessment of pullout strengths of various pedicle screw
designs in relation to the changes in the bone mineral density.
Paper presented at: 48th Annual Meeting of the Orthopedic
A
Research Society; February 10–13, 2002; Dallas, TX.)
B
screws were pulled out at a loading rate of 5 mm per minute
(ASTM: F1691-98) using MTS858 Bionix Machine (MTS
Corp., Eden Prairie, MN). A one-way analysis of variance
(ANOVA) test was done for the statistical analysis with
SPSS 7.0 (SPSS, Inc., Chicago, IL). Comparison of the
TABLE 6-1. Axial strength (N) data for different types of screws pulled out in foam of different densities
Foam grade Body profile Square Buttress V-shape
SS 591 ± 22 497 ± 80 615 ± 36
10 ST 622 ± 43 598 ± 25 634 ± 19
TT 525 ± 36 547 ± 30 568 ± 74
SS 864 ± 50 769 ± 56 987 ± 55
12 ST 956 ± 30 825 ± 108 1,005 ± 92
TT 811 ± 41 808 ± 25 944 ± 32
SS 1,397 ± 93 1,303 ± 126 1,516 ± 78
15 ST 1,582 ± 82 1,438 ± 36 1,569 ± 79
TT 1,197 ± 43 1,352 ± 88 1,396 ± 68
SD, standard deviation; SS, straight major diameter on straight core; ST, straight major diameter on
tapered core; TT, tapered major diameter on tapered core.
Source: Choi W, Lee S, Woo KJ, et al. Assessment of pullout strengths of various pedicle screw
designs in relation to the changes in the bone mineral density. Paper presented at: 48th Annual Meeting or the Orthopedic Research Society; February 10–13, 2002; Dallas, Texas.
pullout strength between the screw types was assessed with
the Tukey test and Scheffe test. P values less than 0.05 were
regarded as statistically significant.
The maximum pullout strengths for various screw
designs are shown in Table 6-1. The highest purchasing
Tooth profile (mean ±SD)
62 /SECTION I/BASIC SCIENCE
pow er in an y screw design w as observed in foams with the
highest density (grade 15). Exponential increase in pullout
strength was seen when the foam density increased from
grade 10 through 15 (Fig. 6-1B). The VST screws exhibited the highest strength while the BSS the lowest with
grades 10 and 12. The SST type screws were strongest
against pullout with grade 15 foam while the STT the
weakest. Statistical analysis showed that regardless of the
foam grades or tooth profiles, the conical screws with
straight major diameter (i.e., ST types) were stronger than
the other two designs (i.e., SS or TT, p < .05). Within the
ST types, the buttress (B) tooth screws showed the lowest
pullout strength among the three tooth profiles (p < .05),
while there was no statistical difference between the
square and V-shape tooth with grades 12 and 15. Howe ver ,
with grade 10 foam, no significant difference was
observed statistically among the three. In a case for the SS
type screws, the buttress (B) tooth was the w eakest re gardless of the foam grades. Between the square and V-shape
tooth screws, no difference was found. As for the TT
types, V-shape screws had higher pullout strength than the
square with grades 12 and 15. No statistical differences
were found between the V-shape and the buttress (B)
screws with grades 12 and 15, nor were any found among
the three tooth types with grade 10.
The use of foam for pullout tests afforded a control on
the variability in the quality of bone that is prevalent in
other studies. Thus, the foam allowed for characterization
of the effects of screw variables on the pullout strength.
Overall, results demonstrate that the conical screws are
consistently more effective against the pullout than the
cylindrical designs. This is especially evident when the
major diameter of the screw is kept straight. In this case,
the contact area between the screw thread and surrounding foam is large. Although no consistent statistical superiority was found with the tooth profiles, results did suggest that the V-shape tooth screws ranked highest in many
statistical comparisons and the buttress types showed
comparativel y lo wer pullout strength than the other types.
This finding may be somewhat different from the literature. This can be due to the absence of the cortical purchase in foam model used in this study. On the other
hand, the square tooth screws faired well in terms of pullout strength when the major diameter was kept straight
but did not do so when tapered. Results also suggest that
as the density of the host site is decreased no clear choice
of tooth profile could be found.
Likewise, McKinley et al. de v eloped a synthetic model
to study the role of variations in pedicle morphology on
the loads in pedicle screws (18). Synthetic vertebral
analogs were fabricated, varying in pedicle height, length,
or width independently. Pedicle screws internally instrumented with strain gauges were used as load transducers
to determine screw-bending moments within the pedicle
and body of the analog. Analogs were loaded in compression to simulate loading of an unstable burst fracture.
Screw bending moments within the pedicle increased
incrementally with increasing pedicle length, rising 30%
as length increased from 8 mm to 12 mm. Screw moment
increased 20% when pedicle height dropped below 15
mm, consistent with a threshold effect. Changes in pedicle width did not affect screw loads within the pedicle.
Thus, in situ pedicle screw loads increased significantly
as pedicle length increased and as pedicle height
decreased.
Lim et al. investigated the relationship between the
BMD of the vertebral body and the number of loading
cycles to induce loosening of an anterior vertebral screw
(19). (Screw loosening was defined as 1 mm displacement of the screw relative to bone.) There was a positive
correlation between the number of loading cycles to
induce screw loosening and BMD (r < 0.8, p < .01). The
average number of loading cycles to induce screw loosening was significantly less for specimens with BMD
2
less than 0.45 g/cm
greater than or equal to 0.45g/cm
compared to those with BMD
2
. These findings suggest that BMD may be a good predictor of anterior vertebral screw loosening as well, just like the pedicle screws.
These findings of increase in pullout strength, number
of cycles to failure, and tightening torque with BMD,
however, are not fully corroborated with the cor responding in vivo work. For example, moments and forces during pedicle screw insertion were measured in vivo and in
vitro and correlated to BMD, pedicle size, and other
screw parameters (material, diameter) (20). The mean in
vivo insertion torque (1.29 Nm) was significantly greater
than the in vitro value (0.67 Nm). The linear correlation
between insertion torque and BMD was significant for
the in vitro data but not for the in vivo data. No correla-
tion was observed between insertion torque and pedicle
diameter. However, another investigation that clinically
evaluated 52 patients who underwent pedicle screw f ixation augmenting posterior lumbar interbody fusion
(PLIF) supports the in vitro findings. BMD was measured using dual energy X-ray absorptiometry (DEXA)
and radiographs were assessed for detecting loosening
and at the pedicle screw bone interface. BMD was found
to have a close relationship with the stability of pedicle
screw in vivo, and BMD values below 0.674 ± 0.104
2
suggested a potential increased risk of “nonunion”.
g/cm
Similar studies pertaining to screw vertebral body interface for the anterior instrumentation have yet to be undertaken.
The current literature is based on studies of cylindrical
pedicle screw designs. Conical screws have been introduced that may provide better “f it and f ill” of the dorsal
pedicle as well as improved resistance to screw bending
failure. However, there is concern about loss of fixation if
conical screws must be backed out after insertion (21).
Abshire et al. evaluated these issues by pulling out cylindrical and conical screws inserted in pedicles of porcine
vertebrae (21). Pullout results were comparable to data