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A BIOMECHANICAL COMPARISON BETWEEN A BIOLOGICAL
INTERVERTEBRAL DISC AND SYNTHETIC INTERVERTEBRAL DISC
IMPLANTS UNDER COMPLEX LOADING: AN IN VITRO STUDY
A Thesis
Presented to
The Graduate Faculty of The University of Akron
In Partial Fulfillment
of the Requirements for the Degree
Master of Science
Snehal Chokhandre
August, 2007
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A BIOMECHANICAL COMPARISON BETWEEN A BIOLOGICAL
INTERVERTEBRAL DISC AND SYNTHETIC INTERVERTEBRAL DISC
IMPLANTS UNDER COMPLEX LOADING: AN IN VITRO STUDY
Snehal Chokhandre
Thesis
Approved: Accepted:
____________________________ ____________________________
Advisor Dean of the College
Dr. Glen O. Njus Dr. George K. Haritos
____________________________ ____________________________
Committee Member Dean of the Graduate School
Dr.Stanley Rittgers Dr. George R. Newkome
____________________________ ____________________________
Department Chair Date
Dr. Daniel B. Sheffer
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ABSTRACT
This study aimed at evaluating the intervertebral disc implants for structural
properties in comparison with the biological intervertebral disc. We tried to understand
the load response of a cadaveric intervertebral disc structure under a physiologic complex
loading compared to its replacement and also the response of the cadaveric disc structure
to the current test standards for intervertebral implants.
Four cadaveric disc structures and four elastomeric intervertebral disc implants
(E-d) (Theken Disc, Akron, OH) were tested under modified ISO testing specifications
for replacements and also under single axis and coupled loads. The complex loading
included a combination of flexion-extension (6º,-3º), left and right lateral bending (2º,-
2º), axial rotation (2º,-2º) and axial compression (900-1700N). When tested under
modified ISO loading, the flexion-extension stiffness and axial rotation stiffness values
were found to be significantly different (p=0.0002 and p=0.0027, respectively) and no
significant difference was found between lateral bending stiffness values (p=0.9304).
When the two groups were tested under single axis loading, there was a
significant difference in the axial compression stiffness and axial rotation stiffness values
(p= 0.0067 and p=0.0027, respectively) and no significant difference was seen in the
flexion-extension stiffness and lateral bending stiffness values (p= 0.1092 and p=0.1348,
respectively). Under coupled loading of flexion-extension and lateral bending there was a
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significant difference in the lateral bending stiffness values between the two groups
(p=0.0485) but no significant difference was seen in the flexion-extension stiffness values
(p=0.1197).
Fatigue characteristics of the cadaveric intervertebral disc structures, E-d and a
pseudo Charite´ which was designed and fabricated similar to the Charite´ intervertebral
disc (Depuy Spine, Inc), were determined and compared. All the discs were fatigued
under modified ISO testing specifications. Stiffness values for single axis loadings and
disc heights were used for comparison and failure was characterized by a decrease in disc
height. The decrease in disc height at the given loading was considerably higher for the
cadaveric specimens and all the cadaveric disc structures failed due to fractures in the
vertebral bodies.
The study also aimed at evaluating the current testing standards for the
intervertebral disc implants as we put forth the argument that the actual biological
structure (intervertebral disc structure) itself would not survive the testing specifications
which its replacement is supposed to bear without failure.
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TABLE OF CONTENTS
LIST OF TABLES………………………………………………………………….
LIST OF FIGURES…………………………………………………………………
CHAPTER
I. INTRODUCTION………………………………………………….……..
1.1 Overview……….…………...…………………………………….
1.2 Objectives of the study……………………………………….......
1.3 Hypothesis…………………...……………………………………
II. LITERATURE REVIEW…...……….……………………………………
2.1 Structure and function of spine …………………………………..
2.2 Mechanical loading of spine………………………….........……..
2.2.1 Mechanical function……………………………………….
2.2.2 Mechanical damage……………………………………….
2.3 Lower back pain………………………………………………….
2.4 Intervertebral disc..………………………….…………………….
2.4.1 Structure and function of intervertebral disc……………...
2.4.2 Intervertebral disc degeneration………………………….
2.4.3 Association with pain……………………………………..
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2.5 Treatment options for lower back pain generated due to damaged
or degenerated disc …………………………………………………
2.6 Disc Arthroplasty: Artificial intervertebral disc……………………
2.7 Loads on the biological disc and the need to mimic them in the
replacements………………………………………………………
III. METHODOLOGY……………….………………...……………………..
3.1 Specimens ……………………………..…………………………
3.2 Sample size determination…………………….………….……….
3.3 Sample preparation…..…………………………...…..……………
3.3.1 Cadaveric specimen……………………………….……….
3.3.2 Intervertebral disc implants………………………..………
3.4 Biomechanical testing…………………………………………….. 3.4.1 Cadaveric specimens …………………………………….
3.4.2 Intervertebral disc implants…….………………………… 3.5 Testing protocol………….…………………………………….…..
3.6 Data acquisition…………………………………………………… 3.7 Data analysis……………………………………………………….
3.8 Statistical analysis………………………………………………… IV. RESULTS………………………...………………...……………………..
4.1 Single and multi axial testing ……………………………………..
4.2 Fatigue test…………………………………………………………
4.3 Failure……………………………………………………………..
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V. DISCUSSION
5.1 Overview……….…………..……..……………………………….
5.2 Single and multi axial testing……………………………………...
5.3 Fatigue characteristics comparisons……………………………….
5.4 Limitations of the study…………………………………………..
5.5 Future work……………………………………………………….
BIBLIOGRAPHY…………………………………………………..…….……….
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LIST OF TABLES
Table
2.1
2.2
2.3
4.1
4.2
4.3
4.4
4.5
4.6
4.7
4.8
4.9
4.10
4.11
4.12
Proposed designs for intervertebral disc prosthesis……………………….
Loads on spine……………………………………………………………..
Testing parameters…………………………………………………………
Statistical results…………………………………………………………...
Axial stiffness comparison under fatigue………………………………….
Flexion-extension (FE) stiffness comparison under fatigue……………….
Lateral bending (LB) stiffness comparison under fatigue…………………
Axial rotation (AR) stiffness comparison under fatigue………...…………
Shear forces comparison………………………………………..………….
Disc height comparison under fatigue……………………………………..
Summary of biomechanical testing of cadaveric disc L1L2 (October 30th
,2006) ……………………….……………………………………………...
Summary of biomechanical testing of cadaveric disc L1L2 (January 10 th,
2007)……………………………………………………………………….
Summary of biomechanical testing of cadaveric disc L3L4 (February 20th
,2007)………………………………………………………………………..
Summary of biomechanical testing of cadaveric disc L2L3 (March 19th,
2007)………………………………………………………………………..
Summary of biomechanical testing of synthetic disc E-d1 (November 6th,
2006) …………………….……………...………………………………….
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4.13
4.14
4.15
4.16
Summary of biomechanical testing of synthetic disc E-d2 (April 10th
,2007)…………...……. ……………………………………………….….
Summary of biomechanical testing of synthetic disc E-d3 (April 10th,
2007)…………. ….…. …………………………………………….…….
Summary of biomechanical testing of synthetic disc E-d4 (April 10th,
2007)…………. ……. …………………………………………….…….
Summary of biomechanical testing of pseudo- Charite´ (April 10th, 2007)
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LIST OF FIGURES
Figure
1.1
2.1
2.2
2.3
2.4
2.5
3.1
3.2
3.3
3.4
3.5
3.6
3.7
3.8
3.9
3.10
3.11
A three dimensional coordinate system (according to ISO 2631)…………
Human spine ………………………………………………………………
Motion segment ….……………………………………………………….
Spinal ligaments…...………………………………………………………
Spinal cord and nerve roots……………………………………………….
Intervertebral disc………………………………………………………….
Intact motion segment (Top view)……………………………..………….
Intact motion segment (Side view)………………………………………...
Intact posterior region of the motion segment……………………………..
After removal of the posterior region……………………………………...
Motion segment in alginate…………………….………………………….
An alginate negative…………………….…………………………………
Alginate positive ……………………………..……………………………
Final specimen of 5.1 cm…………………………………………………..
Specimen with one potted vertebral body and Steinman pin in othervertebral body……………………….…………………………………..…
Specimen with both vertebral bodies potted in PMMA and ready fortesting ………………………………………………………………..……
E-d (Theken Disc, Akron, OH) and pseudo Charite´ …………………….
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3.12
3.13
3.14
3.15
3.16
3.17
3.18
4.1
4.2
4.3
4.4
4.5
4.6
4.7
4.8
4.9
4.10
4.11
4.12
Cadaveric testing specimen……………………………………………….
Multi Axial Endura Tech Testing system………………………………….
Temperature monitoring system…………………………………………...
Implant testing……………………………………………………………..
Loading pattern for modified ISO…………………………………………
Data acquisition software………………………………………………….
Typical graphs obtained from one of the data sets………………………...
Load – displacement graphs for Cadaveric Discs and Synthetic Implants..
Axial Stiffness Comparison………………………………………………..
Flexion –Extension Angle vs. Moment graphs for Cadaveric Discs and
Synthetic Implants E-d for Flexion-Extension with Static Compression………..
Flexion- Extension Stiffness Comparison…………………………………
Flexion –Extension Angle vs. Moment graphs for Cadaveric Discs and
Synthetic Implants E-d for Flexion-Extension with Dynamic Compression……..
Flexion-Extension Stiffness Comparison………………………………….
Lateral Bending Angle vs. Moment graphs for Cadaveric Discs and Synthetic
Implants E-d for Lateral Bending with Static Compression……………………..
Lateral Bending Stiffness Comparison…………………………………….
Flexion-Extension Angle vs. Moment graphs for the coupled loading of Flexion-Extension, Lateral Bending with Dynamic Compression……….
Lateral Bending Angle vs. Moment graphs for the coupled loading of
Flexion-Extension, Lateral Bending with Dynamic Compression……….
Flexion-Extension and Lateral Bending Stiffness comparisons………….
Axial Rotation Angle vs. Torque graphs for Axial Rotation with Static
Compression………………………………………………………………..
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4.13
4.14
4.15
4.16
4.17
4.18
Axial Rotation Stiffness Comparison…………………………………..….
Flexion-Extension Angle vs. Moment graphs for Modified ISO…………..
Lateral Bending Angle vs. Moment graphs for Modified ISO……..….…..
Axial Rotation Angle vs. Moment graphs for Modified ISO……………..
Stiffness Comparisons (Modified ISO)…………………………..………..
Radiographs to assess damage to the cadaveric disc………………………
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CHAPTER I
INTRODUCTION
This chapter provides an overview of the rationale behind conducting this study
with a brief description of the in vitro technique used followed by the goals, objectives
and the hypothesis.
1.1 Overview
Lower back pain continues to remain a major reason of morbidity around the
world. With an average incidence rate of 60%, about 2% of the work force in US has
compensable back injuries every year [23].Among the back related injuries, the lumbar
case increased from 39 to 47% in 2003. The annual cost of low back disability in US has
been estimated about 50 billion dollars [23, 17].
The healthcare expenditure on US is
very high and the prevalence in similar in men and women. About 1% of the US
population faces chronic disability due to back pain and 1% suffer temporary disability.
Few of those who are disabled for more than six months have a chance of returning to
work and after two years of disability, their chances of reemployment almost vanish [23].
Even though almost every anatomical structure in the lower back has been
implicated, the intervertebral disc has been associated with it the most, mainly because of
the age –related deterioration. The changes in the intervertebral disc generally precede or
coincide, with other degenerative changes in the spine [8].
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The degeneration of the intervertebral discs is a natural process. With increasing
age, the discs lose their elasticity, flexibility and shock absorbing characteristics.
Intervertebral discs act in part as cushions that provide shock absorption between the
vertebrae. Degeneration may cause a lot of symptoms such as back pain, nerve root
pathology, spinal cord compression, etc. As the discs age, they shrink and therefore the
disc height reduces and the space available for nerve roots and spinal cord also reduces
[31]. This compression of spinal nerves may result in pain, loss of muscle control and in
extreme cases, even paralysis. The damaged disc may also lead to segmental instability.
Besides age related deterioration, trauma can also lead to disc damage.
When the conservative measures fail and a damaged disc or segmental instability
directly damages the neural elements or threatens to do so, operative intervention
becomes necessary. Decompression involves removal of the disc material that
compresses the spinal nerve causing pain and sensory changes in the affected nerve. Even
though it is effective in relieving pain caused by herniated disc, it is not enough to restore
the nucleus to its original load sharing capacity [35, 19].
A standard protocol to alleviate pain from degenerated or herniated discs (may be
due to unnatural mechanical loadings) is the disc removal and fusion of the two adjacent
vertebrae. However, fusion impairs the normal motion and even though it may be
considered a standard treatment in many instances, a number of issues develop.
The loss of mobility may result in stiffness and loss of functional capacity. The
increased stresses on the adjacent non-fused areas due to the transfer of stresses from the
fused segments may lead to adjacent segment degeneration and/or remodeling [31, 32].
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An alternative to spinal fusion is total disc replacement (spinal arthroplasty),
which intends to maintain the motion of the operative level by removing the damaged
disc and limits the adjacent segment breakdown. This is a current focal point in the
European and North American surgical and industrial areas. However, the focus has
been on using current technology and techniques on the replacement. Very little to no
current technology/ techniques, have been used to study what is being replaced. The vast
majority of this information is decades old [26].
As the intervertebral disc is a complex anatomic and functional structure, the
development of an artificial intervertebral disc which would be an efficient, durable and
reliable structure, is a challenge. The disc function is difficult to reproduce and the choice
of materials that will bear the loads is also of important consequence as the strains
supported by spine are different from those of peripheral joints. The complex strains
supported by the intervertebral disc make the implant development even more difficult.
For e.g., as the spine undergoes 100 million flexion cycles during lifetime (not including
the slight motion during breathing which is about 6 million a year), the optimal life for
the implant is found to be 30 million cycles and a minimum of 10 million cycles [26].
Biochemical problems, difficulties in fitting the implant and the size and weight of the
implant make the design, development and surgical techniques a challenge. Except a
couple of artificial intervertebral disc designs, which are still not being extensively used,
most of the designs are under clinical trials [31].
In determining the durability of the implant, the information regarding the loads
that it will be exposed to needs to be known. Different loading conditions and motion
controls must be considered too. Further, the current testing protocols given by American
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society for testing and materials (ASTM) and the International Standards Organization
(ISO) differ [15]. Also, the loads that spine undergoes during daily activities are still not
completely known, much less understood.
As the motion preservation devices are gaining popularity, their preclinical
mechanical testing is of prime importance to predict their in vivo safety and efficacy.
This study aimed at evaluating the intervertebral disc implants for structural properties in
comparison with the biological intervertebral disc. We tried to understand the load
response of a cadaveric intervertebral disc structure under a physiologic complex loading
compared to its replacement and also the response of the cadaveric disc to the current test
standards for intervertebral implants.
We tested the cadaveric disc structures and the synthetic intervertebral disc
implants under the testing specifications provided by the ISO for artificial intervertebral
discs. The complex loading included a combination of flexion-extension (6º,-3º), left and
right lateral bending (2º,-2º), axial rotation (2º,-2º) and axial compression (900-1700N).
Figure 1.1 gives an illustration of the coordinate system and the type of loads the
specimens were subjected to.
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Figure 1.1 A three dimensional coordinate system (according to ISO 2631) [16]
1.2 Objectives of the study
1. To make a biomechanical comparison of the performance of a cadaveric disc
structure to that of the artificial intervertebral disc under ISO testing standards.
2. To test the human cadaveric intervertebral discs under complex loads (modified
ISO) in vitro.
3. To compare the fatigue characteristics of the cadaveric disc structure, the
elastomeric intervertebral disc implant (Theken Disc, Akron, OH) and a Pseudo
Charite´ (C-p) designed like Charite´ artificial intervertebral disc (DePuy Spine,
Inc.)
SaggitalFrontal
Transverse
SaggitalFrontal
Transverse
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1.3 Hypothesis
Null hypothesis (H0)
1) There are no statistical differences with respect to structural stiffness values
between human cadaveric discs compared to the elastomeric intervertebral disc
implant (Theken Disc, Akron, OH) with respect to transverse, saggital and frontal
planes.
Alternate hypothesis (H1)
1) There are statistical differences with respect to structural stiffness values between
human cadaveric discs compared to the elastomeric intervertebral disc implant
(Theken Disc, Akron, OH) with respect to transverse, saggital and frontal planes.
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CHAPTER II
LITERATURE REVIEW
2.1 Structure and function of spine
The human spine is a complex mechanical structure that is flexible, suitably
strong and stable. Regardless of the activity there is always some type of load on the
spine. The spine protects the spinal cord and nerve roots, provides structural support
and bears the loads of head, shoulders and the upper body [34, 12]. It keeps the upper
body weight balanced evenly on pelvis. This puts minimal workload on the muscles
to maintain the upright posture [34, 12].
In humans, the spine is composed of 24 vertebrae, sacral bones and the coccyx
(Figure 2.1). The spinal column provides body's main upright support. The spine
forms three curves. The cervical spine (at the neck), is curved slightly anteriorly. The
thoracic spine (at the middle back) is curved posteriorly. The lower back (lumbar
spine) curves slightly anteriorly. The normal curve of the neck and lower back is
called lordosis while the normal curve of the thoracic spine is called kyphosis. An
increased lordosis or kyphosis may be associated with a congenital problem,
neuromuscular problem, poor posture, osteoporosis, obesity etc [39, 27,39].
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Figure 2.1 Human Spine [37] Figure 2.2 Motion Segment [37]
The term motion segment is that structure formed by two adjacent vertebrae
associated ligaments, and intervertebral disc [Figure 2.2]. The intervertebral discs act like
shock absorbers to suppress impulsive forces transmitted by activities like jumping, they
increase the time taken by the forces to reach the lower extremities while decreasing the
force [36,37]. The bony projections at the back of the vertebra form the vertebral arch
which consists of two pedicles and two laminae. The spinal canal contains the spinal cord
and under each pedicle spinal nerves exit the spinal cord and pass through the foramina to
branch out to the body. The spinous process, two transverse processes, two superior
facets, and two inferior facets arise from the vertebral arch [36,37]. The facet joints resist
axial rotation and the ligaments also work as motion limiters. The intersegmental
ligaments include the ligamentum flavum, interspinous and intertransverse ligaments.
The intersegmental ligaments that hold many vertebrae together are the anterior and
posterior longitudinal ligaments, and the supraspinous ligaments [36, 39, 38,39].
Intervertebral Disc
Vertebral Body
Facet Joint
Pedicle
Nerve Root
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Figure 2.3 Spinal Ligaments [37] Figure 2.4 Spinal Cord and Nerve Roots [37]
The spinal cord runs within the spinal canal from the brainstem to the first lumbar
vertebra and then the cord fibers separate to form cauda equina and then branch out to
legs and feet. The spinal cord works as a messenger between the brain and the rest of the
body by carrying motor messages from the brain to the body and relays sensory messages
from the body to the brain with the help of spinal nerves. Any damage could lead to a
sensory and motor function loss [36, 37, 38]. As the facet joints and disc allow the
movement between the two vertebrae and the foramens provides area for the nerve roots
to branch out from the spinal cord, any instability or excessive movement may lead to
irritation or pinching of the nerve roots. This generally results in pain and may increase
pressure on facet joints leading to inflammation and muscle spasms as the muscles try to
stop the segment movement. Instability leads to faster degeneration of the spine [34, 27]
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2.2 Mechanical loading of spine
Irrespective of the activity there is always some kind of load on the spine. The
combination and magnitude of loads change with activity. The following section aims at
discussing the mechanical function and damage due to inappropriate loading on the spine.
2.2.1 Mechanical function
The intervertebral discs resist compressive forces mainly in upright posture, the
facet joints protect the discs from torsion and excess shear and the ligaments prevent
excessive bending [2,28] Nearly 20% of the compressive load falls on the facet joints but
this can go up to 70% if the intervertebral discs are degenerated and narrowed [40]. In the
lumbar region, The facet joints resist horizontal forces acting perpendicular to their
articular surfaces and limit the range of axial rotation. They can resist shearing forces of
about 2kN. [2] The posterior intervertebral ligaments can resist 100N (posterior
longitudinal ligament) to nearly 1kN for facet joint capsular ligaments.[18] The fibers in
the interspinous and capsular ligaments are oriented and sized accordingly to resist
forward bending movement. [2]
2.2.2 Mechanical damage
To investigate the failure mechanisms, dynamic loading tests are required so that
the elastic limit can be detected by loading the specimen rapidly and the non-reversible
deformation is real and not a viscous creep. The vertebral body is the first structure to fail
in compression. Failure occurs at lower loads in fatigue and damage is mostly at the end
plate. Compressive fatigue damage is more common due to the micro fractures. Fatigue
damage accumulates rapidly if spine is exposed to mechanical vibrations. Rapid
compressive forces can cause a burst fracture. Fracture occurs in a single loading cycle
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with forces of approximately 2kN or in a response to cyclic forces between 380N to
760N [2].
Facet joints are not well oriented to resist compressive forces. With the disc
narrowing of 1-3mm the loading on articular surfaces increases. In the lumbar spine the
damage generally occurs in 1-3º of movement and the damage probably occurs in
subchondral bone behind the articular surface [2, 11].
The intervertebral disc and the facet joints resist extension and axial rotation and
thus do not affect the ligaments much. The interspinous ligament is the first to rupture in
hyperflexion. A 2º further extension is required to damage apophysial joint capsular
ligament. [2]. Disc damage cannot be caused directly by compressive loading as the
compressive loads have been found to affect the adjacent vertebral body. An
understanding of the mechanical loading of the lumbar spinal structures is important as
abnormal mechanical loading is one of the reasons of damaged spinal structures which
lead to lower back pain
2.3 Lower back pain
Lower back pain can occur due to several reasons. With increase in age, the bone
strength decreases, muscles weaken and discs lose flexibility and the ability to cushion
the vertebrae. A sprain, strain or spasm in muscles or ligament can cause back pain. Over
compressed spine can lead to disc rupture and pinching of nerve roots resulting in pain.
Degenerative conditions like arthritis, osteoporosis, congenital abnormalities, obesity,
smoking, weight gain during pregnancy, stress, poor physical condition, posture
inappropriate for the activity being performed, and poor sleeping position also may
contribute to low back pain [35, 38].
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Several conditions can cause lower back pain. A herniated, or ruptured disc can
push into the space containing the spinal cord or the nerve root and results in pain. In the
cauda equina syndrome, the disc material is pushed into spinal canal and compresses a
bundle of sacral and lumbar nerve roots and can result in permanent neurological
damage. [35]. If the ruptured disc presses the sciatic nerve which carries nerve fiber to
legs, knee and foot it may cause low back pain combined with pain in legs. A severe
condition can lead to loss of motor control loss in the legs. Spinal degeneration may
result in narrowing of spinal canal. Spinal stenosis, osteoporosis, skeletal irregularities,
spondylitis (severe inflammation of spinal joints) all lead to lower back pain in varying
degrees. Inarguably, of all the anatomical structures in the lumbar spine associated with
the lower back pain, intervertebral disc has been associated with it the most and mainly
due to age related deterioration and damage due to various reasons.
2.4 Intervertebral Disc
Intervertebral discs act as shock absorbers and to a certain extent motion limiters.
The following section discusses the structure and function of the intervertebral disc, the
possible causes and effects of degeneration and its association with pain.
2.4.1 Structure and function of intervertebral disc
The intervertebral disc is a mixed structure consisting of peripheral collagenous
bands called annulus fibrosus made of 15-20 concentric layers of alternating oblique
fibers [26,12]. The central core is made up of mucopolysacharide gel and proteoglycans
and is called nucleus pulposus [26,12]. This core is very hydrophilic and generates a
tension on the annulus even when no external loading is present [26,12]. The preloading
enhances the resistance to external forces and helps in dividing the compressive forces.
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The disc allows movement along and around three main axes. Thus, the center of rotation
is constantly modified along two axes simultaneously. The intervertebral disc provides a
major part of its stability on its own [Figure 2.5]. The collagen fiber arrangement in the
annulus creates an efficient system to control and restrict rotation [26].
Figure 2.5 Intervertebral Disc
2.4.2 Intervertebral disc degeneration
Degeneration of the intervertebral disc has been shown to be positively correlated
with respect to age in humans. There is generally a gradual decrease in the hydrophyllic
proteoglycans and the associated water content [8,4]. The nucleus dehydrates and shrinks
resulting in load changes on the annulus. Radial tears may occur in the annulus. And if
natural healing does not occur, the proteoglycan material of the nucleus may migrate
from the center to the periphery through the tear. [8,4] This may lead to further
delamination of the annulus and may results in back pain due to simulation of
sinuvertebral nerve. The nucleus may transgress all the layers of annulus resulting in a
herniation. This may mechanically deform the nerve root and result in a radicular pain
[8,4]. Due to accumulation of metabolic products, the pH changes with change in
immunoglobulins and prostaglandins, which results in back pain. Degenerated discs
result in compromised stability and increased motion between vertebrae [4].
Nucleus
Pulposus
Lamellae
Annulus Fibrosus
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The causes of disc degeneration are not well understood and also the clinical
variability makes it difficult to assess the risk for development of severe or earlier onset
of disc degeneration. Male discs have been found to degenerate earlier than their age
matched female counterparts. [8,26] Other factors that may lead to disc degeneration
include smoking, exposure to whole body vibrations and heavy lifetime occupational and
leisure physical loading [8,26] A genetic predisposition has also been implicated. Since
the disc is not the only mobile structure of the functional unit , secondary osteo- arthrosic
modifications of the facet joints influence disc degeneration and vice versa. Furthermore,
the origin of pain in the functional unit is ill understood and is more complex than
peripheral joints [8, 26].
2.4.3 Association with pain
The exact role disc degeneration plays in the occurrence of low back pain is
unclear .The reason why the intervertebral disc is considered the axial pain generator may
be because the posterior portion of the annulus fibrosus is innervated by fibers of the
sinuvertebral nerve which is a branch of the dorsal root ganglion. Irritation of this nerve
is thought to be one reason of axial back pain. Sensory information from the lumbar
intervertebral discs is conducted to other spinal levels through the paravertebral
sympathetic trunks [21] Therefore, decompression of the nerve root is unlikely to reduce
low back pain symptoms. Thus, removal of the disc and denervation of the annulus is
more likely to reduce discogenic pain than decompression alone. Further, disc material
has been shown to be a direct source of chemically irritative substances such as
phospholipase A2, prostaglandin E, substance P, and lactic acid [21, 8] Disc herniation in
a non degenerated disc can be due to abnormal loads, for e.g. due to trauma.
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2.5 Treatment options for lower back pain generated due to damaged or degenerated disc
Pain can occur at any stage of the degeneration process, from annular tear to
complete disc degeneration. Depending on the severity of the pain and the reason, several
treatment options have been tried over the years. Anti-convulsants are used to treat
certain type of nerve pain. Some antidepressants like amitriptyline and desipramine have
been known to dull pain signals [35, 19]. Spinal manipulations, acupuncture, traction,
transcutaneous electrical stimulation have all been used to treat lower back pain. When
the conservative measures fail and segmental instability directly damages the neural
elements or threatens to do so, operative intervention becomes necessary [19].
Decompression involves removal of the disc material that compresses the spinal nerve
causing pain and sensory changes in the affected nerve. Even though it is effective in
relieving pain caused by herniated disc, it is not enough to restore the nucleus to its
original load sharing capacity [35, 19, 13].
Fusion (arthrodesis) involves eliminating the motion between two or more
vertebrae by using a bone grafting procedure and internal fixation system. [4, 31] This
procedure is performed when there is a gross instability of motion segment resulting from
a traumatic injury or degeneration [4, 31]. The technique is also used to abolish motion at
a painful but stable, degenerated disc. Fusion thus eliminates instability and helps in
relieving pain. Even though spinal fusion is extensively used, several problems are
associated with it. The loss of mobility from long segment fusions may result in stiffness
and loss of functional capacity. The transfer of stress from the fused areas to the
bordering non-fused areas may result in adjacent segment degeneration. In many patients
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with multiple levels of mild degenerative disease, surgery to correct all degenerative
segments would often be too extensive and disabling [8, 31]. Disc arthroplasty
alternatives are designed to preserve motion segments, reduce the risk of facet damage,
and limit associated adjacent segment degeneration.
2.6 Disc Arthroplasty: Artificial intervertebral disc
Disc arthroplasty offers an alternative treatment, that of alleviating pain while
preserving the physiological motion. This involves replacing the entire damaged disc
(nucleus and annulus) or the damaged nucleus depending on the need. Nucleus
replacements aim to restore the disc height and return the annular fibers to their natural
length, which would facilitate normal load distribution. The theoretical advantages of
disc arthroplasty are prevention of adjacent segment disease, protection of neural element
by restoring disc height and shorter recovery time as patients would not recuperative
period for fusion maturation [4, 26, 14, 13].Total disc replacements would be used when
removal of all the sources of pain (including nucleus and annulus) becomes necessary
and healing is not possible in any way. Physiologic motion is complex and the prosthesis
should approximate the size and motion of a physiologic disc to avoid distraction or
overloading of the facet joints.
The indications for total disc replacement are similar to fusion, including back
and leg pain which is unresponsive to appropriate attempts at non-operative treatments,
radiographic evidence of disc degeneration with varying degree of disc space collapse
[13,22]. Other indications involve post laminotomy/ discotomy syndromes. It should be
avoided in patients with osteoporosis as a weaker bone would fail fast and anything
greater than grade I spondylolisthesis because an unstable segments cannot be held in
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place with an artificial disc. Further, patients with significant canal stenosis or neural
compressive disease or pain related to significant scarring from previous surgery should
not be treated by disc replacement [13,22].
The first disc arthroplasty attempt was performed with a steel ball endoprosthesis
by Fernstorm in the late 1950s. Since then several implants have been designed. Few of
them have been tested in animal models and even fewer reached the clinical trial stage
[14]. Further, there are several problems associated with the implant design, like,
presence of three joints at each level, the ability of the intervertebral disc to provide a
complex combination of mobility and stability, the changing center of rotation at a given
motion segment with changing position of the motion segment [6,14].
Some prosthesis strive to reproduce the viscoelastic properties of the disc and are
made of silicones or polymers and some aim at reproducing the range of motion (i.e.
motion characteristics ) and are made of metal or polyethylene couples. Some attempt to
combine both principles. Table 2.1 shows some of the designs that have been proposed
over the years [26].
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Table 2.1 Proposed designs for intervertebral disc prosthesis
Year Researcher Intervertebral disc prosthesis design
1950
1973
1975
1978
1987
1991
1994
1995
2000
2000
Nechemson
Stubstad et al.
Froning
Weber
Downey
Pisharodi
Baumgartner
Beer and Beer
Bryan and Kunzler
Gauchet
Liquid Silicone rubber
Reinforced elastic polymer disc
Discoid bladder like implant
Polyethylene structures with ceramic ovoid core
Cushion made of silicon with inner core of fluid
Hollow bag containing springs
Elastic beads replacing nucleus
Disc shaped screwed plates joined by springs
Two threaded hollow half cylinders
Two round plates and an intermediate deformable body
Cervical and lumbar disc arthroplasty has reached the stage of clinical trials in
United States. The spinal arthroplasty techniques replace damaged, painful and
incompetent intervertebral discs with a prosthesis designed to restore normal disc height,
and function. Artificial disc replacement is considered experimental by the Food and
Drug Administration (FDA), but is becoming an increasingly more common intervention
for patients. While artificial intervertebral discs have been used internationally for over
10 years, only two devices, Charité (DePuy Spine, Inc.) and ProDisc-L (Synthes Spine)
have received approval. Other devices are currently under investigation in this country as
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part of the FDA process of approval [26,31,32]. The structural, functional and pathogenic
factors make the development of an artificial disc a challenge. Complex strains,
biochemical problems and surgical difficulties further demand an implant that is
biocompatible and can withstand the long term complex mechanical demands. Currently
there are four different types of artificial discs that are undergoing either clinical, in-vivo
or in-vitro evaluation. The different types are the following.
Composite: e.g. Link SB Charite® disc, Prodisc®
Hydraulic: e.g. PDN® Prosthetic Disc Nucleus
Elastic: e.g. Acroflex® Disc
Mechanical: e.g. Maverick ® Disc prosthesis
Link SB Charite® disc Prodisc®
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2.7 Loads on the biological disc and need to mimic them in the replacements
The spine undergoes approximately 100 million movements during an average
lifetime. The optimal lifespan of the existing implants is only 30 million movements with
a recommended minimum of 10 million. The implant should be able to withstand the
kind of loads the biological disc has to bear without failing [25,18]
So far the biological disc has been studied in isolated loading conditions, mainly
compression. But there is always some kind of coupled load on the disc. Studies have not
focused much on how the disc performs under complex loads. A few studies have
focused on multidirectional properties of the disc [25,18]
To determine the durability of the implant, the type of the load the implant will be
exposed to, should be known. Different loading conditions and motion controls must also
be considered. So far for testing intervertebral disc implants,the test parameters are
debatable. The parameters should be physiologically accurate. Daily loads on spine can
be summarized as follows: [15]
Table 2.2 Loads on spine [15]
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The testing criteria provided by ASTM and ISO are considerably different (Table
1) and the motion paths generated by both are also different. The ISO parameters
combine all the loading conditions together resulting in a cross- shear motion profile [15].
Table 2.3 Testing parameters [15]
This study aims at understanding how the biological disc would perform under
complex loading conditions under which an artificial intervertebral disc implant is tested.
This would give us an insight on weather the biological disc can withstand the test
criteria set up for the artificial disc implants and also how it performs under complex or
coupled loads.
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CHAPTER III
METHODOLOGY
Introduction
The objective of this study was to make a biomechanical comparison of the
human cadaveric disc and the synthetic intervertebral disc implants under multi axial
loading. The study also aimed at evaluating the performance of cadaveric disc under the
testing specifications prescribed for the replacement of the human disc by the
International Standards Organization (ISO). The following content gives detailed
information of how the study was conducted.
3.1 Specimens
Human cadaveric specimens
Standard radiographic tests were performed on the human cadaveric specimens to
ensure they were normal and did not have any fractures or any such structural damage.
Radiographs were obtained in mid saggital and oblique planes and were verified by an
orthopedic surgeon to see if they were normal.
Intervertebral disc implants
Elastomeric intervertebral disc implants by Theken Disc, Akron, OH (E-d) were
used for the biomechanical comparison with the cadaveric disc structures under multi
axial loading. For fatigue testing, an artificial intervertebral disc was prepared by
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Theken Disc, Akron, OH which was based on the dimensions determined from
the literature about the Charite´ artificial intervertebral disc implant (DePuy Spine, Inc).
We named it pseudo Charite´.
3.2 Sample size determination
To determine the sample size of cadaveric specimens and elastomeric
intervertebral disc implant E-d (Theken Disc, Akron, OH) required for test, the following
formula was used.
( ) [ ]( )2
])[1(2,
2/2 ν ν α δ σ
P t t n
−+×≥
Where, n : number of replications / sample size
σ : true standard deviation
δ : smallest true difference to be detected
ν : degrees of freedom of sample standard deviation
α : significance level
P : desired probability that a difference will be found to be significant
Power of the test
T : two tailed t-table value.
Values for above equation were: standard deviation of 10%; significance level of
0.05; power of the test was taken as 90%; and, smallest true difference to be detected,
30%. The sample size was found out to be 4.
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3.3 Sample preparation
The cadaveric specimens and the pseudo Charite´ needed preparation prior
testing. The elastomeric implants E-d did not need any fixture adaptations or preparation
before testing. The following section discusses the sample preparation done for the study.
3.3.1 Cadaveric specimens
For the test, fresh cadaveric specimens were used. They were kept hermetically
frozen at -20 ºC and cool thawed prior to any work or biomechanical testing. Studies
have shown that freezing and thawing at room temperature have little effect on the
biomechanical behavior of the disc [16]. Radiographic tests were performed on the
human cadaveric specimens to ensure they were normal and did not have any fractures or
any such structural damage. Radiographs were obtained in mid saggital and oblique
planes and were verified by an orthopedic surgeon to see if they were normal. The age,
gender, weight, size, cause of death and other information about the donors were not
available to this author.
The samples were dissected from the intact spines. All the muscle mass was
removed (Figure 3.1). In each cadaveric specimen, all the posterior elements including
the lamina, facet joints and pedicles were removed, leaving the anterior longitudinal
ligament and the posterior longitudinal ligament intact (Figure 3.2). All tissue work was
conducted by suitably trained personnel.
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Figure 3.1 Intact motion segment Figure 3.2 Intact motion segment(Top View) (Side View)
Figure 3.3 Intact posterior region Figure 3.4 After removal of theof the motion segment posterior region
One objective was to keep the intervertebral disc, anterior and posterior
longitudinal ligaments and some part of the adjacent vertebral bodies intact and
maintaining the height of the specimen at 5.1 cm. To achieve this, parallel cuts were
made in the superior and inferior vertebral bodies in the transverse plane. The removal
was symmetric with respect to the centroidal transverse plane through the disc.
Considerable portions of the vertebral bodies were removed such that the remaining
region could be used to hold the specimen in place and the intervertebral disc takes most
of the loads while testing. To achieve this, an intricate series of steps were taken. First an
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impression of the spinal segment (after removing all the posterior parts) using alginate
dental impression material was taken (Figure 3.5 and 3.6). The next step was that of a
positive made by pouring alginate into the negative impression (Figure 3.7). This positive
was then used to get another negative though this time the negative was of
polymethylmethacrylate (PMMA). It was this negative that was actually used to rigidly
hold the cadaveric specimen and appropriate cuts were made. To cut the specimen such
that the distance of the transverse planes in which the cuts were 2.55 cm on both sides
from the transverse plane passing through the center of the disc. First, the height of the
intervertebral disc was determined by prodding the specimen with a hypodermic needle
without causing any damage to the tissue, such that the intersection of the disc and
vertebral bodies was determined. Once the height of the disc was established, another
hypodermic needle was placed exactly halfway. From this centrally placed needle, 2.55
cm distance on both the sides was determined and marked. Appropriate cuts in the
vertebral bodies in the transverse plane were made using a Milwaukee saw (Figure 3.8).
Figure 3.5 Motion segment in alginate Figure 3.6 Alginate negative
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Figure 3.7 Alginate positive Figure 3.8 Final specimen of 5.1 cm
For fixation to the testing fixture each specimen was potted in PMMA. The
vertebral bodies were embedded in PMMA using a circular potting construct. The
specimen was placed such that the center of the specimen and the construct are
overlapping. To get the center of the specimen in transverse plane, a circle was assumed
on the disc and its center was determined [Figure 3.9]. The center typically lied a
centimeter from the posterior longitudinal ligament. The center of both the vertebral
bodies was determined in a similar fashion.
The base plate of the potting construct had a particular hole pattern in which
small brass rods were placed before PMMA was poured in the assembly, so as to get a
similar hole pattern in the PMMA which held the vertebral body. This was done to place
and secure the potted specimen in the testing chamber which has a similar hole pattern.
Machining was done on the holes to get the required hole diameter.
Attention was given so that the specimens were not torqued at odd angles during
the preparatory phase. Steinman pins were drilled in the vertebral bodies for proper load
transmission in bending and torsion while testing (Figure 3.9). Once the specimen was
placed in the construct and secured ,a PMMA powder/monomer mixture was poured in
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the potting construct and allowed to set. Both vertebral bodies were potted in similar
fashion (Figure 3.10).
Figure 3.9 Specimen with one potted vertebral body and Steinman pins in the other
vertebral body
Figure 3.10 Specimen with both vertebral bodies potted in PMMA and ready for testing
3.3. 2 Intervertebral disc implants:
For comparison under multi axis by loading with human cadaveric disc structures,
four elastomeric disc implants by Theken Disc, Akron, OH, were used. The specimens
did not need any fixture adaptation prior to testing as testing fixtures were already
available (Figure 3.11).
For fatigue testing, an artificial intervertebral disc was prepared by Theken Disc,
Akron, OH which was based on the dimensions determined from the literature about the
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Charite´ artificial intervertebral disc implant (DePuy Spine, Inc). As the implant consists
of two cobalt chrome end plates and an intervening ultrahigh molecular weight
polyethylene sliding core, the two end plates were potted in PMMA following the
protocol used for cadaveric discs. The cobalt chrome end plates were circular, therefore
we did not need to approximate the center. Steinman pins were inserted in both the end
plates before they were potted in PMMA (Figure 3.11).
Figure 3.11 E-d (Theken Disc, Akron, OH) and pseudo Charite´
3.4 Biomechanical testing
Biomechanical testing of all the specimens was done at Theken Disc, Akron, OH.
The specimens used and the testing equipment used are described in the following
section.
3.4.1 Cadaveric Specimen
The testing chamber was built such that a temperature probe could be inserted to
maintain the desired temperature. The potted cadaveric specimens were fixed in the
testing chamber and kept moist by immersion in Ringers solution and maintained at 37 ûC
throughout the experiment. (Figure 3.12)
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Posterior Chamber Anterior Chamber Specimen in Testing
Figure 3.12 Cadaveric Testing Specimens
The testing set up would then be placed in a four independent axis prototype servo
Pneumatic EnduraTEC Systems Corp. spine simulator (Figure. 3.13) with an engaged
active temperature control system with the help of which the temperature was maintained
at 37ûC throughout the cadaveric specimen testing (Figure 3.14).
The spine simulator has a six axis load cell which measures the transmitted forces.
The spine simulator was capable of simulating all the loading components separately or
in any combination. This included flexion- extension moments, left and right lateral
bending moments, left and right torsion and axial compression. Shear forces could be
measured by the system even though they could not be controlled.
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Once the specimen was fixed in the testing chamber, it was placed in the spine
simulator such that the anterior region of the specimen was aligned with the flexion
simulator. The testing chamber was locked in the X-Y plane using a set of clamps and
screws. This was done once it was placed horizontally in the simulator, with minimal tilts
in bending planes. Before locking, the bending moments were brought to a very low
value. (Figure 3.8)
Figure 3.13 Multi-Axial EnduraTEC Testing System
Figure 3.14 Temperature monitoring system
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3.4.2 Intervertebral disc implants
The disc implants were also tested in the same system but they did not need
temperature or moisture control. The alignment and placement was done in a manner
similar to the cadaveric specimens (Figure 3.15).
Figure 3.15 Implant testing
3.5 Testing protocol
Spinal loading in vivo is complex and not very well understood. Further the
magnitudes vary from individual to individual depending on the type and extent of
activity, age etc. The objective was to evaluate the cadaveric disc structure under the
testing specifications for the disc replacement while also evaluating the structural
characteristics of the cadaveric discs and the synthetic implants.
The specimens were tested in all the individual, two axis and multi axis loading
conditions (per the ISO testing specifications). All tests were started with an axial
compression of 100 – 600 N at 0.25 Hz. This was followed by a combination of coupled
loads as follows:
a) Flexion – extension (4û, -2û) at 0.25Hz with static compression (600 N).
b) Flexion –extension (4û, -2û) at 0.5Hz with cyclic compression (300 – 700 N).
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c) Lateral bending (-2û,-2û) at 0.25Hz with static compression (600 N).
d) Lateral bending (-2û, -2û) at 0.75 Hz, flexion –extension (4û, -2û) at 0.25 Hz with
dynamic (cyclic) compression (600 N).
e) Torsion (-3û- 3û) with static compression at 0.25 Hz.
In the end the specimens were tested under a modified ISO complex loading
specification (Figure 3.16): flexion – extension (6û,-3û) at 0.5 Hz, lateral bending (-2û,-
2û) at 0.75Hz, torsion (-2û,-2û) at 0.75, cyclic compression (900 N – 1700 N). After each
two axes and /or multi axis loading test, the specimens were again tested under axial
compression (100N – 600N). The specimens were tested for 50±5 cycles for each
loading, so that the first few cycles could be taken as preconditioning and the results
would not be affected by the initial viscoelasticity of the specimen. The number of cycles
was large enough to get the appropriate data but not enough to damage the specimen. For
data analysis the last five cycles were typically chosen.
The cadaveric specimens were fatigued under the ISO testing specifications until
failure. Elastomeric intervertebral disc implant (Theken Disc, Akron, OH) (E-d) were
also fatigued in similar fashion. An artificial intervertebral disc (pseudo-Charite´) was
prepared by Theken Disc, Akron, OH which was based on the dimensions determined
from the literature about the Charite´ artificial intervertebral disc implant (DePuy Spine,
Inc).
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Figure 3.16 Loading pattern for modified ISO
3.6 Data acquisition
The spine simulator system was interfaced with a data acquisition and motion
measurement system (Figure 3.17) which was capable of determining the six motion
components for the three dimensional relative motion between the two vertebral bodies.
This included the three translation and three rotational motions. The relative motion
between the two bodies was transformed to the local coordinate system and the errors
were recorded throughout the test. The data acquisition system collected the data in real
time. Along with the loads, it also measures and collected the moments, shear forces,
tilts, temperature and errors.
Test 7 - ISO modified
-4.0
0.0
4.0
8.0
axl rot
lateral bending
flex exten
normalised force
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Figure 3.17 Data Acquisition Software
3.7 Data Analysis
Offline analysis was conducted using Microsoft Excel®. To calculate the axial
stiffness, flexion- extension stiffness, lateral bending stiffness and axial rotation stiffness
values, linear region of the last five loading and unloading cycles of the load vs.
displacement graph, flexion-extension angle vs. flexion -extension moment graph, lateral
bending angle vs. lateral bending moment graph and axial rotation vs. torque graph
respectively , were regressed (Figure 18).
Figure 3.18 Typical graphs obtained from one of the data sets
Axial Comp. - L1 - L2 - Jan. 10, 2007
0
200
400
600
800
9.4 9.6 9.8 10.0
Displacement ( mm )
F o r c e ( N )
Lat Bend- L1 - L2 - Oct. 30, 2006
-6
-3
0
3
6
-6.0 -3.0 0.0 3.0
Lateral Angle ( Deg. )
L a t e r a l M o m .
( N - m )
Rotat ion & Stat ic Axial Loading- L1 -
L2 - Jan. 10, 2007
-20
-10
0
1020
-4.0 -2.0 0.0 2.0 4.0
Torque Angle ( ° )
T o r q u e M o m . (
N -
m )
Flex.-ExDC. - E-d1- Nov. 2, 2006
-4
0
4
8
-3.0 0.0 3.0 6.0
Flex.-Ext. Angle ( ° )
F l e x . -
E x t . M o m .
( N -
m )
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Shear forces, effect of one loading on the other and errors were also evaluated in
the offline analysis. To determine failure the disc height was recorded and monitored.
Failure was characterized by a constantly reducing disc height. After the fatigue tests, to
evaluate the failure of the cadaveric specimens, radiographs were taken.
3.8 Statistical Analysis
The statistical model assumed for this study was independent t-test. The model is
Yij = µ + τi + εij
Where: Yij: dependent variable (output)
µ : Underlying mean of all groups
τi : treatment
εij : Random error
In this study, ‘i’ is 2. In order to prevent biasing effects due to loading, the testing
order within each specimen was randomized. The disc type was taken as the independent
variables and construct stiffness was taken as the dependant variable. Statistical analysis
using SAS software package (SAS Institute, Cary, NC.) was performed. The means and
standard deviations for each of the variables were also found.
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CHAPTER IV
RESULTS
4.1 Single and multi-axial testing
Following the methodology described in the earlier chapter, four specimens from
each group i.e. cadaveric intervertebral disc structures and E-d, synthetic disc implants,
were tested. The results of the tests were obtained for both the groups and were
compared.
Initial Axial Compression
All the specimens were tested under cyclic axial compression (100N to 600N, at
0.25 Hz). All specimens were tested for 50 cycles of this loading. The load vs.
displacement graphs were plotted and the axial stiffness values were calculated by
regressing the last five loading and unloading cycles for all the specimens (Figure 4.1).
The specimen heights were recorded throughout the test and the changes were evaluated.
Axial stiffness comparisons of the two groups with four specimens each were made. The
average axial stiffness values and standard deviations were calculated (Figure 4.2). The
average axial stiffness for cadaveric discs was 1490± 212 N/mm and that for synthetic
implants (E-d) was 2454±320 N/mm.
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Figure 4.1 Load – Displacement graphs for Cadaveric Discs and Synthetic Implants E-d
Axial Compression (Implants E-d)
Displacement (mm)
8.6 8.8 9.0 9.2 9.4 9.6 9.8
L o a d ( N )
0
100
200
300
400
500
600
700
E-d1
E-d2
E-d3
E-d4
Axial Compression (Cadaveric Specimens)
Displacement (mm)
7.5 8.0 8.5 9.0 9.5 10.0 10.5
L o a d ( N )
0
100
200
300
400
500
600
700
L1L2_Oct30 '06
L1L2_Jan10 '07
L2L3 Mar 19th '07
L3L4_Feb20th '07
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Figure 4.2 Axial Stiffness Comparison
Flexion – Extension with Static Compression
The specimens were tested under Flexion- Extension (4û,-2û) at 0.25 Hz. and a
static compression of 500 N. The Flexion- Extension moments were recorded and the
graphs of the angles vs. moments were plotted (Figure 4.3).The flexion-extension
stiffness was calculated by regressing the last five cycles of the graphs for all the eight
specimens. The flexion-extension stiffness values were calculated for all specimens and
the means and standard deviations were calculated for both the groups. The mean flexion-
extension stiffness for cadaveric specimens was 2.44 ±0.71 Nm/º and that of synthetic
implants (E-d) was 1.77±0.09 Nm/º (Figure 4.4).Throughout this test, the axial rotation
and lateral bending angles were maintained at 0º.
Axial Compression
A x i a l C o m p r e s s i o n S t i f n e s s ( N / m m
)
0
500
1000
1500
2000
2500
3000
Cadaveric specimens
Implants (E-d)
n=4
n=4
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Figure 4.3. Flexion –Extension Angle vs. Moment graphs for Cadaveric Discs and Synthetic
Implants E-d for Flexion-Extension with Static Compression
Flexion-Extension, Static Compression (Implants E-d )
Flexion-Extension angle (°)
-3 -2 -1 0 1 2 3 4 5
F l e x i o n - E x t e n s i o n m o m e n t ( N m
)
-6
-4
-2
0
2
4
6
8
10
E-d1
E-d2
E-d3
E-d4
Flexion Extension , Static Compression (Cadaveric specimens)
Flexion -Extension angle (°)
-3 -2 -1 0 1 2 3 4 5
F l e x i o n - E x t e n s i o n m o m e n t ( N m )
-30
-20
-10
0
10
20
L1L2 Oct 30th '06
L1L2 Jan 10th '07
L2L3 Mar 19th '07L3L4 Feb 20th '07
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Figure 4.4 Flexion-Extension Stiffness Comparison
Flexion- Extension with Dynamic Compression
All the specimens were tested under Flexion-Extension (4û,-2û) at 0.25 Hz and a
dynamic compression of 300 to 700 N. Like the earlier test, the stiffness values were
calculated by regressing the last five loading and unloading cycles of the angle – moment
graphs (Figure 4.5). The flexion- extension stiffness values were calculated like earlier
stage and the means and standard deviations were calculated for both the groups. The
mean flexion-extension stiffness for cadaveric specimens was 2.58±0.64 Nm/ degree and
that of synthetic implants (E-d) was 1.78±0.03 Nm/degree (Figure 4.6). The axial rotation
and lateral bending angles were maintained at 0º.
Flexion-Extension , Static Compression
F l e x i o n - E x t e n s i o n S t i f n e s s ( N m / ° )
0.0
0.5
1.0
1.5
2.0
2.5
3.0
3.5
n=4
n=4
Cadaveric specimens
Implants (E-d)
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Figure 4.5. Flexion –Extension Angle vs. Moment graphs for Cadaveric Discs and Synthetic
Implants E-d for Flexion-Extension with Dynamic Compression
Flexion-Extension, Dynamic Compression (Cadaveric Specimens)
Flexion-Extension angle (°)
-3 -2 -1 0 1 2 3 4 5
F l e x i o n - E x t e n s i o n m o m e n t ( N m )
-10
-5
0
5
10
15
20
L1L2 Oct 30th '06
L1L2 Jan 10th '07
L2L3 Mar 19th '07
L3L4 Feb 20th '07
Flexion-Extension, Dynamic compression (Implants E-d)
Flexion-Extension angle (°)
-2 -1 0 1 2 3 4 5
F l e x i o n - E x t e n s i o n m o m
e n t ( N m )
-6
-4
-2
0
2
4
6
8
E-d1
E-d2
E-d3
E-d4
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Figure 4.6 Flexion-Extension Stiffness Comparison
Lateral Bending with Static Compression
All the specimens (cadaveric and E-d) were tested under left-right lateral bending
(2û,-2û) at 0.25 Hz and a static axial compression of 500N. The lateral bending stiffness
values were calculated from the lateral bending angle vs lateral bending moment graphs
for all the specimens (Figure 4.7). The stiffness values were calculated for both the
groups and the means and standard deviation values were also calculated for both the
groups. The mean lateral bending stiffness for cadaveric specimens was 2.48±0.37 Nm/ º
and that of synthetic implants (E-d) was 2.9±0.4 Nm/ º (Figure 4.8). The flexion-
extension and axial rotation angles were maintained at 0º during this test.
Flexion-Extension ,Dynamic Compression
F l e x i o n - E x t e n s i o n s t i f n e s s ( N m / °
)
0.0
0.5
1.0
1.5
2.0
2.5
3.0
3.5
n=4 Cadaveric specimens
Implants (E-d)
n=4
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Figure 4.7 Lateral Bending Angle vs. Moment graphs for Cadaveric Discs and Synthetic
Implants E-d for Lateral Bending with Static Compression
Lateral Bending, Static Compression (Implants E-d)
Lateral Bending angle (°)
-3 -2 -1 0 1 2 3
B e n d i n g m o m e n t ( N m )
-8
-6
-4
-2
0
2
4
6
8
10
12
E-d1
E-d2
E-d3
E-d4
Lateral Bending ,Static Compression (Cadaveric specimens)
Lateral Bending angle (°)
-3 -2 -1 0 1 2 3
L a t e r a l B e n d i n g m o m e n t ( N m )
-8
-6
-4
-2
0
2
4
6
8
10
L1L2 Oct 30th '06
L1L2 Jan 10th '07
L2L2 Mar 19th '07
L3L4 Feb 20th '07
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Figure 4.8 Lateral Bending Stiffness Comparison
Flexion- Extension, Lateral Bending with Dynamic Compression
All the specimens were tested under a coupled loading of flexion-extension
(4û- 2û) at 0.25 Hz, lateral bending (2û- 2û) at 0.75 Hz and a dynamic axial
compression of 300 to 700N (Figure 4.9 and 4.10). The flexion-extension stiffness
and lateral bending stiffness values were calculated like earlier stages. The flexion –
extension stiffness and lateral bending stiffness values were calculated for both the
groups and the means and standard deviations were obtained. The mean flexion-
extension stiffness for the cadaveric specimens was found to be 2.52±0.61 Nm/º and
that of synthetic implants (E-d) was found to be 1.9±0.2 Nm/º. The mean lateral
bending stiffness for the cadaveric specimens was found to be 2.57±0.45 Nm/º and
that of synthetic implants (E-d) was found to be 3.4±0.5 Nm/º (Figure 4.11). The
axial rotation angle was maintained at 0º for this test.
Lateral Bending ,Static Compression
L a t e r a l B e n d i n g s t i f n e s s ( N m / ° )
0
1
2
3
4
n=4
Cadaveric specimens
Implants (E-d)n=4
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Figure 4.9 Flexion Extension Angle vs. Moment graphs for the coupled loading of
Flexion-Extension, Lateral Bending with Dynamic Compression
Flexion-Extension (Implants E-d)
Flexion-Extension angle (°)
-3 -2 -1 0 1 2 3 4 5
F l e x i o n - E x t e n s i o n m o m e n t
( N m )
-6
-4
-2
0
2
4
6
8
10
E-d1
E-d2
E-d3
E-d4
Flexion-Extension (Cadaveric specimens)
Flexion -Extension angle (°)
-3 -2 -1 0 1 2 3 4 5
F l e x i o n - E x t e n s i o n m o m e n t ( N m
)
-10
-5
0
5
10
15
L1L2 Oct 30th '06
L1L2 Jan 10th '07
L2L3 Mar 19th '07
L3L4 Feb 20th '07
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Figure 4.10 Lateral Bending Angle vs. Moment graphs for the coupled loading of
Flexion-Extension, Lateral Bending with Dynamic Compression
Lateral Bending (Cadaveric Specimens)
Lateral Bending angle (°)
-3 -2 -1 0 1 2 3
L a t e r a l B e n d i n g m o m e n t ( N m )
-6
-4
-2
0
2
4
6
8
10
12
L1L2 Oct 30th '06L1L2 Jan 10th '07
L2L3 Mar 19th '07
L3L4 Feb 20 '07
Lateral Bending(Implants E-d)
Lateral Bending angle (°)
-3 -2 -1 0 1 2 3
L a t e r a l B e n d i n g m o m e n t (
N m )
-8
-6
-4
-2
0
2
4
6
8
10
12
E-d1
E-d2
E-d3
E-d4
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Figure 4.11 Flexion- Extension and Lateral Bending Stiffness Comparison
Axial rotation with static compression
The specimens were tested under axial rotation (3û,-3û) at 0.25 Hz and a static
compression of 500N and keeping the flexion-extension and lateral bending angles at 0û.
The axial rotation or torsional stiffness was calculated by regressing the last five loading
and unloading cycles of the angle – torque graphs (Figure 4.12). The stiffness values
were calculated for both the groups and the means and standard deviations were
calculated. The mean axial rotation stiffness for the cadaveric specimens was found to be
2.87±1.04 Nm/ º and that of synthetic implants (E-d) was found to be 1.4±0.13 Nm/º
(Figure 4.13). The flexion-extension and lateral bending angles were maintained at 0º for
this test.
Flexion-Extension, Lateral Bending with Dynamic Compression
S t i f f n e s s ( N m / ° )
0
1
2
3
4
5
Cadaveric specimens
Implants (E-d)
Flexion-Extension Lateral Bending
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Figure 4.12 Axial Rotation Angle vs. Torque graphs for Axial Rotation with StaticCompression
Axial Rotation, Static Compression (Implants E-d)
Axial Rotation angle (°)
-4 -3 -2 -1 0 1 2 3 4
A x i a l T o r q u e ( N
m )
-6
-4
-2
0
2
4
6
E-d1
E-d2
E-d3
E-d4
Axial Rotation, Static Compression (Cadaveric specimens)
Axial Roation angle (°)
-4 -3 -2 -1 0 1 2 3 4
A x i a l T o r q u e ( N m )
-15
-10
-5
0
5
10
15
L1L2 Oct 30th '06
L1L2 Jan 10th '07L2L3 Mar 19th '07
L3L4 Feb 20th '07
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Figure 4.13 Axial Rotation Stiffness Comparison
Testing under modified ISO specifications
After every test mentioned above, an axial compression test was performed to
make sure there was no structural damage. The test under ISO specifications was
performed at the end of the experiment when it was established that the specimens
were not damaged in any way after the earlier tests and no changes in the axial
compression values were observed.
The complex loading was composed of flexion extension (6û,-3û) at 0.5 Hz, lateral
bending (2û,-2û) at 0.75 Hz, axial rotation (2û,-2û) at 0.75 Hz and a dynamic axial
compression of 900 to 1700N (Figures 4.14, 4.15 and 4.16).
Axial Rotation ,Static Compression
A x i a l R o t a t i o n S t i f f n e s s ( N m / ° )
0
1
2
3
4
5
n=4
n=4
Cadaveric Specimens
Implants (E-d)
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The mean flexion-extension stiffness for the cadaveric specimens was found to
be 2.64±0.23 Nm/degree and that of synthetic implants (E-d) was found to be
1.4±0.17 Nm/degree. The mean lateral bending stiffness for the cadaveric specimens
was found to be 2.47±0.87 Nm/degree and that of synthetic implants (E-d) was found
to be 2.4±0.08 Nm/degree. The mean axial rotation stiffness for the cadaveric
specimens was found to be 2.93±0.69 Nm/degree and that of synthetic implants (E-d)
was found to be 1.2±0.07 Nm/degree.(Figure 4.17)
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4.14 Flexion Extension Angle vs. Moment graphs for Modified ISO
Flexion-Extension (Implants E-d)
Flexion -Extension angle (°)
-4 -2 0 2 4 6 8
F l e x i o n - E x t e n s i o n m o m
e n t ( N m )
-6
-4
-2
0
2
4
6
8
E-d1
E-d2
E-d3
E-d4
Flexion Extension, ISO (Cadaveric Specimens )
Flexion -Extension angle (°)-4 -2 0 2 4 6 8
F l e x i o n - E x t e n s i o n m o m e n t ( N m
)
-30
-20
-10
0
10
20L1L2 Oct 30th '06
L1L2 Jan 10th '07
L2L3 Mar 19th '07
L3L4 Feb 20th '07
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Figure 4.15 Lateral Bending Angle vs. Moment graphs for Modified ISO
Lateral Bending (Implants E-d)
Lateral Bending angle (°)-3 -2 -1 0 1 2 3
l a t e r a l B e n d i n g m o m e n t ( N m )
-8
-6
-4
-2
0
2
4
6
8
E-d1
E-d2
E-d3
E-d4
Lateral Bending ,ISO(Cadaveric Specimens)
Lateral Bending angle (°)
-3 -2 -1 0 1 2 3
L a t e r a l B e n d i n g m o m e n t ( N m )
-15
-10
-5
0
5
10
15L1L2 Oct 30th, 06
L1L2 Jan 10th '07
L2L3 Mar 19th '07
L3L4 Feb 20th '07
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Figure 4.16 Axial Rotation Angle vs. Torque graphs for Modified ISO
Axial Rotation (Implants)
Axial Roation angle (°)
-3 -2 -1 0 1 2 3
A x i a l T o r q u e ( N
m )
-4
-3
-2
-1
0
1
2
3
4
E-d1
E-d2
E-d3
E-d4
Axial Rotation,ISO (Cadaveric specimens)
Axial Rotation angle (°)
-3 -2 -1 0 1 2 3
A x i a l T o r q u e ( N m )
-8
-6
-4
-2
0
2
4
6
8
L1L2 Oct 30th '06
L1L2 Jan 10th '07
L2L3 Mar 19th '07L3L4 Feb 20th '07
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Figure 4.17 Stiffness Comparisons (Modified ISO)
An independent t-test was performed using SAS software (SAS Institute, Cary,
NC) on the results to evaluate whether the two groups were significantly different. For
the test the synthetic and cadaveric disc types were taken as the independent variable and
the axial stiffness as the dependent variable. The significance level of the test was taken
as 0.05 (Table 4.1). All the results were also checked for possible outliers. No outliers
were detected.
ISO
S t i f f n e s s ( N m / ° )
0
1
2
3
4
Cadaveric Specimens
Implants (E-d)
Flexion-Extension Lateral Bending Axial Rotation
n=4
n=4
n=4
n=4
n=4
n=4
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Table 4.1 Statistical Results
Loading p-value Result
AC 0.0067 S
FE,AC 0.1092 NSFE,DC 0.0431 S
LB,AC 0.1348 NS
FE ( FE,LB,DC) 0.1197 NS
LB ( FE,LB,DC) 0.0485 S
AR,SC 0.0027 S
FE ( Modified ISO) 0.0002 S
LB ( Modified ISO) 0.9304 NS
AR ( Modified ISO) 0.0027 S
S- Significantly different, NS – Not significantly different
4.2 Fatigue test
Fatigue comparisons were made between three structures: the cadaveric disc
structures, the Theken synthetic disc (E-d) and a pseudo Charite´ made by Theken Disc,
Akron, OH using the information available in literature about the Charite´ intervertebral
disc implant by Depuy Spine Inc. The disc structures were fatigued under complex ISO
loading specifications. Failure was characterized by a constant decrease in the height of
the specimen.
One sample from each type was taken and fatigued under the modified ISO
complex loading. The axial stiffness before and after the fatigue test were determined and
compared. The cadaveric discs were fatigued till they failed and failure was
characterized by a constant decrease in disc height. Changes in axial compression
stiffness, flexion-extension stiffness, lateral bending stiffness and axial rotation stiffness
before and after the test were evaluated for the three discs (Table 4.2- 4.5).
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Table 4.2 Axial stiffness comparison under fatigue
Disc Structures
Initial AxialStiffness
(N/mm)After Fatigue (N/mm) % Difference
L1L2 Tested-Jan10th (Cad.) 1791 1065 (1500 cycles ) 41%
E-d 3 (ThekenDisc) 2188 2079 (1500 cycles ) 5%
Pseudo- Charite 5783 5431 (1500 cycles) 8%
Table 4.3 Flexion-Extension (FE) Stiffness comparison under fatigue
Table 4.4 Lateral Bending (LB) Stiffness comparison under fatigue
SpecimenLB Stiffness
Initial(Nm/º)
LB Stiffness
Final(Nm/º)
L1L2 Jan 10th 2.86 3.61
E-d 3 3.55 2.50
Pseudo- Charite 2.04 2.06
SpecimenFE Stiffness
Initial(Nm/º)
FE StiffnessFinal
(Nm/º)
L1L2 Jan 10th 3.21 2.71
E-d 3 1.80 1.60
Pseudo-Charite 1.23 1.24
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Table 4.5 Axial Rotation (AR) Stiffness comparison under fatigue
AR Stiffness AR Stiffness
Initial FinalSpecimen
(Nm/º) (Nm/º)
L1L2 Jan 10th 4.03 2.43
E-d 3 1.50 1.30
Pseudo- Charite 0.03 0.04
The maximum anterior-posterior (AP) shear force and lateral shear force values
were determined from the axial compression tests before and after the fatigue tests were
conducted on the specimens. A comparison of the shear forces before and after the
fatigue test is as given below (Table 4.6).
Table 4.6 Shear forces comparison
Specimen Max.AP Shear
Initial (N)
Max.AP Shear
Final (N)
L1L2 Jan 10th 14.6 -60.8
E-d 3 -0.5 -5.0
Pseudo- Charite´ -3.0 -5.4
SpecimenMax.Lateral
Shear
Initial (N)
Max.LateralShear
Final (N)
L1L2 Jan 10th -5.7 1.5
E-d 3 -4.1 -3.6
Pseudo- Charite´ 1.0 0.2
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The fatigue failure comparison between the three disc types was based on the
change in disc height before and after the fatigue test. Change in height was recorded
after the fatigue test was run on all the three specimens. The values are taken from the
axial compression tests before and after the fatigue tests. Only one specimen was
compared from each type.
Table 4.7 Disc height comparison under fatigue
The synthetic discs E-d have been tested under ISO fatigue test protocol for 7M
cycles without failure at Theken Disc, Akron, OH. The Charite´ disc is already
undergoing clinical trials and has been approved by the FDA after clearing the testing
criteria.
Specimen Disp.(mm)
Initial
Disp. (mm)
Post Fatigue
Change
in height
(mm)
L1L2 Jan 10th 9.75
4.574 (1500
cycles) 5.17
E-d 39.6
9.41 (1500cycles) 0.22
Pseudo-Charite
17.317.23 (1500
cycles)0.10
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4.3 Failure
The ISO standards expect the implants to last for 10 million cycles of the complex
loading without failure. One objective of the study was to evaluate whether the actual
biological disc which is going to be replaced by the implant, would last under the testing
specifications. The failure of the biological disc structure was characterized by a
consistent decrease in specimen height. To determine the kind of failure the structure had
undergone, radiographs were taken and a bony failure was detected, which characterized
the structural failure.
Figure 4.18 Radiographs to access damage to the cadaveric disc
Fracture
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4.4 Result tables
Table 4.8 Summary of biomechanical testing of Cadaveric disc L1L2 (October. 30th,2006)
Oct 30th and 31st , 2006 L1-L2: Human
Axial Flex-Ext LateralAxialRot.
Stable Stable
Stiff. Stiff. Stiff. Stiff. Max. Min. δ
(N/mm) (Nm/°) (Nm/°) (Nm/°) ( mm ) ( mm ) ( mm )
AC 1428.65 / / / 9.23 8.88 0.35
FESC 1.83 / / 8.70 8.60 0.11
FEDC 1.70 / / 8.62 8.51 0.10
LBSC / 2.24 / 8.47 8.39 0.07LBFEC 1.67 2.22 / 8.48 8.34 0.15
ARSC / / 2.36 8.41 8.22 0.18
2.51 2.04 1.95 2.90 2.69 0.21
2.69 3.06 2.19 2.13 1.92 0.21
2.65 2.97 1.90 1.28 1.14 0.14ISOb
2.87 2.96 1.92 1.00 0.84 0.16
AC 1674.74 / / / 1.72 1.35 0.38
* Highlighted row values in ISOb were used for calculations.
* For Table 4.8 to 4.15
AC- Axial Compression
FESC- Flexion-Extension with Static Compression
FEDC- Flexion- Extension with Dynamic Compression
LBSC- Lateral Bending with Static Compression
LBFEC- Lateral Bending, Flexion-Extension with Dynamic Compression
ARSC- Axial Rotation with Static Compression
ISOb- Modified ISO
AC- Axial Compression
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Table 4.9 Summary of biomechanical testing of Cadaveric disc L1L2 ( January 10th,2007)
Jan. 10, 2007 L1-L2: Human
Axial Flx-Ext Lateral AxialRot.
Stable Stable
Stiff. Stiff. Stiff. Stiff. Max. Min. δ
(N/mm) (Nm/°) (Nm/°) (Nm/°) ( mm ) ( mm ) ( mm )
AC 1791.38 / / / 9.75 9.48 0.28
FESC 3.21 / / 9.11 9.05 0.06
FEC 3.20 / / 9.48 9.36 0.11
LBSC / 2.86 / 9.43 9.29 0.14
LBFEC 3.11 3.14 / 9.51 9.27 0.23
TSC / / 4.03 9.34 9.24 0.11
FELBTC 2.82 2.99 4.05 9.46 9.23 0.23
2.56 2.88 3.29 6.63 5.98 0.65
2.65 3.22 2.72 5.31 4.68 0.63ISOb 17
2.71 3.61 2.43 4.11 3.56 0.55
AC 1065.38 / / / 4.57 4.13 0.45
* Highlighted row values in ISOb were used for calculations.
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Table 4.10 Summary of biomechanical testing of Cadaveric disc L3L4 ( February 20th
,2007)
Feb 20th and
21st, 2007 L3-L4: Human
Axial Flx-Ext Lateral
Axial
Rot. Stable Stable
Stiff. Stiff. Stiff. Stiff. Max. Min. δ
(N/mm) (Nm/°) (Nm/°) (Nm/°) ( mm ) ( mm ) ( mm )
AC. 1294 / / / 8.18 7.85 0.33
FESC / 1.85 / / 8.04 7.8 0.24
FE DC 2272 2.61 / / 7.97 7.8 0.17
LB SC / 2.72 / 7.91 7.53 0.39
FE LB DC 2.56 2.71 / 8.07 7.76 0.31AR SC 3.4 / / 8.45 8.17 0.28
ISOb 2.98 3.46 2.97 5.37 5.11 0.25
2.73 3.35 2.73 5.03 4.75 0.28
2.53 3.19 2.56 4.64 4.32 0.32
3.07 3.21 2.41 4.28 3.94 0.34
2.83 3.06 2.32 3.97 3.6 0.37
2.59 3.03 2.05 3.5 2.85 0.66
2.65 3.21 1.91 3.07 2.25 0.82
AC 809 / / / 3.6 3.17 0.43
* Highlighted row values in ISOb were used for calculations.
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Table 4.11 Summary of biomechanical testing of Cadaveric disc L2L3 (March 19th,2007)
March 19th , 2007 L2 L3 Human
Axial Flx-Ext Lateral Axial Rot Stable Stable
Stiff. Stiff. Stiff. Stiff. Max. Min. d
(N/mm) (Nm/°) (Nm/°) (Nm/°) ( mm ) ( mm ) ( mm )
AC 1449.00 / / / 10.20 9.87 0.33
FESC 2.90 / / 10.00 9.80 0.19
FEDC 2.82 / / 9.84 9.68 0.16
LBDC / 2.11 / 9.86 9.44 0.41
FELBDC 2.76 2.20 / 9.85 9.64 0.22
ARSC / / 1.70 10.35 10.06 0.29
Axial Comp. 11 / / / 9.90 9.64 0.26
2.50 1.49 3.43 9.06 8.00 1.06
2.43 1.19 3.00 9.03 7.80 1.23
2.19 1.01 2.70 8.95 7.73 1.22ISOb 18
2.31 1.03 2.60 8.83 7.60 1.23
* Highlighted row values in ISOb were used for calculations.
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Table 4.12 Summary of biomechanical testing of Synthetic disc E-d1 (November 6th,2006)
Nov 06,2006 E-d1
Axial Flx-Ext LateralAxialRot. Stable Stable
Stiff. Stiff. Stiff. Stiff. Max. Min. δ
(N/mm) (Nm/°) (Nm/°) (Nm/°) ( mm ) ( mm ) ( mm )
AC 1964.59 / / / 8.96 8.66 0.3
FE SC 1.65 / / 8.75 8.65 0.1
FE DC 1.75 / / 8.78 8.62 0.15
LB SC / 2.68 / 8.74 8.24 0.5
FE LBDC 1.69 2.83 / 8.81 8.59 0.22
AR SC / / 1.25 8.64 8.62 0.02
ISOb 1.2 2.36 1.14 8.2 7.85 0.35
Table 4.13 Summary of biomechanical testing of Synthetic disc E-d2 ( April 10th
, 2007)
April 10th and 11th,
2007 E-d2
Axial Flx-Ext Lateral
Axial
Rot. Stable Stable
Stiff. Stiff. Stiff. Stiff. Max. Min. δ
(N/mm) (Nm/°) (Nm/°) (Nm/°) ( mm ) ( mm ) ( mm )
AC 2207 / / / 9.59 9.37 0.22
FE SC 1.76 / / 9.44 9.37 0.07
FE DC 1.74 / / 9.46 9.34 0.12
LB SC / 2.72 / 7.91 7.52 0.39
FE LB DC 1.85 3.71 / 9.53 9.28 0.26
AR SC / / 1.45 9.66 9.63 0.03
ISOb 1.43 2.36 1.25 9.55 8.39 1.16
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Table 4.14 Summary of biomechanical testing of Synthetic disc E-d3 (April 10th
, 2007)
April 10th and 11th,
2007 E-d3
Axial Flx-Ext Lateral
Axial
Rot. Stable Stable
Stiff. Stiff. Stiff. Stiff. Max. Min. δ
(N/mm) (Nm/°) (Nm/°) (Nm/°) ( mm ) ( mm ) ( mm )
AC 2721.36 / / / 9.57 9.52 0.05
FE SC 1.86 / / 9.52 9.43 0.09
FE DC 1.8 / / 9.51 9.39 0.12
LB SC / 2.86 / 9.6 9.1 0.5
FE LB DC 2.23 3.13 / 9.63 9.36 0.27
AR SC / / 1.41 9.42 9.4 0.02
ISOb 2.18 3.33 1.47 9.6 9.35 0.24
AC 1.66 2.59 1.29 8.89 8.53 0.36
Table 4.14 Summary of biomechanical testing of Synthetic disc E-d3 ( April 10th
, 2007)
April 10th and 11th,
2007 E-d4
Axial Flx-Ext LateralAxialRot. Stable Stable
Stiff. Stiff. Stiff. Stiff. Max. Min. δ
(N/mm) (Nm/°) (Nm/°) (Nm/°) ( mm ) ( mm ) ( mm )
AC 2187.93 / / / 9.64 9.42 0.23
FE SC 1.8 / / 9.45 9.45 0
FE DC 1.78 / / 9.48 9.34 0.14
LB SC / 3.55 / 9.55 9.05 0.49
FE LB DC 2 3.85 / 9.57 9.3 0.27AR SC / / 1.5 9.39 9.37 0.02
ISOb 1.57 2.51 1.29 8.82 8.4 0.43
AC / / / 9.42 9.18 0.24
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Table 4.15: Summary of biomechanical testing of pseudo- Charite ( February 20th, 2007)
Feb 20th 2007 Pseudo Charite
Axial Flx-Ext Lateral
Axial
Rot. Stable Stable
Stiff. Stiff. Stiff. Stiff. Max. Min. δ
(N/mm) (Nm/°) (Nm/°) (Nm/°) ( mm ) ( mm ) ( mm )
AC 19928.5 / / / 8.33 8.27 0.06
FE SC 0.92 / / 8.36 8.23 0.13
FE DC 0.96 / / 8.35 8.23 0.12
LB SC / / 0.52 8.37 8.25 0.12
FE LB DC 0.92 0.48 / 8.43 8.18 0.25
AR SC / / 0.03 8.29 8.29 0
ISOb 0.9 0.43 0.04 8.4 8.08 0.32
AC / / / 8.3 8.25 0.05
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CHAPTER V
DISCUSSION
5.1 Overview
An intervertebral disc implant is expected to allow normal physiologic motion
and permit the spine to regulate its motion instead of spinal motion adapting to the
implant. The implant should address the in vivo loading conditions, and it should resist
wear and material delamination [30].
This study aims at evaluating two implants against the biological disc structure
that they are designed to replace. It is an attempt to assess whether the structural
properties of one of the implants (E-d) are comparable to the actual cadaveric disc
structures so as to get an idea of how well the implant could mimic the structural
properties of an actual disc structure. The comparison is also made between the cadaveric
disc structure and two implant designs on the basis of fatigue characteristics, which
essentially focuses on finding out the durability of the implant against the biological
structure under the recommended loading. The study also aims at evaluating the ISO
testing specifications for the intervertebral disc implants. We believe that they are so
demanding that even the actual biological disc structure would fail under such loadings.
Thus, expecting an implant to pass these tests would perhaps lead to more robust designs
but it could also lead to possible rejection of some good implant designs that may not last
as long under the test environment as the ISO standards require.
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5.2 Single and multi-axial tests
The objective of these tests was to make a comparison between the groups under
complex modified ISO loading. We started the test with single axis loading and then
introduced one type of load at every step of the testing protocol. The structural properties
of the two groups were compared for all the single and multi-axis tests.
Axial compression
The cadaveric discs and the implants E-d were tested under a cyclic compression
of 100-600N at 0.25 Hz. The axial stiffness of the two groups was significantly different
(p =0.0067). The average value for the cadaveric disc structures was found to be 1490±
212 N/mm and that of the implant was 2454±320 N/mm. The higher value of the mean
axial stiffness for the implants might be because one of the implants was mistakenly
tested under a smaller range of axial compression which led to a higher value of axial
stiffness for that specimen.
Flexion-Extension loading, lateral bending loading and axial rotation loading with static
compression
When the specimens were tested under flexion-extension (4º,-2º), lateral bending
(2º,2º) and axial rotation (2º, 2º) individually with a static compression of 500N, there
was no significant difference detected between the flexion-extension stiffness values for
the flexion-extension loading with static compression (p =0.1092) and lateral bending
stiffness values for the lateral bending with static compression test (p=0.1348) for the two
groups. There was a significant difference in the axial rotation stiffness values for the
axial rotation with static compression test (p=0.0027). The mean flexion-extension
stiffness for cadaveric specimens was 2.44 ±0.71 Nm/º and that of synthetic implants (E-
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d) was 1.77±0.09 Nm/º. The mean lateral bending stiffness for cadaveric specimens was
2.48±0.37 Nm/ º and that of synthetic implants (E-d) was 2.9±0.4 Nm/º. The mean axial
rotation stiffness for the cadaveric specimens was found to be 2.8±1.04 Nm/º and that of
synthetic implants (E-d) was found to be 1.4±0.13 Nm/º. The loading characteristics for
all the three loadings were similar for the two groups.
Therefore, based on these results we conclude that the two disc types behave
similarly under flexion-extension and lateral bending and there behavior is different in
axial rotation. In other words, the implants mimic the structural properties of the
cadaveric discs in flexion extension and lateral bending.
Flexion –extension loading with dynamic compression
The flexion-extension stiffness values were significantly different for the two
groups (p=0.0431). The mean flexion-extension stiffness for cadaveric specimens was
2.58±0.64 Nm/º and that of synthetic implants (E-d) was 1.78±0.03 Nm/º. Based on the
results we conclude that the implants could not mimic the structural properties of the
cadaveric discs under the given loading.
Flexion- extension, lateral bending with dynamic compression
In the coupled loading with flexion – extension (4º,-2º), lateral bending (2º,-2º)
and axial compression (300-700N), there was no significant difference in the flexion –
extension stiffness values between the two groups (p=0.1197). However, there was a
significant difference in the lateral bending stiffness values (p= 0.0485). Under a coupled
load, the two disc types behave differently.
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Complex modified ISO loading
The complex loading was consisted of flexion extension (6˚,-3˚) at 0.5 Hz, lateral
bending (2˚,-2˚) at 0.75 Hz, axial rotation (2˚,-2˚) at 0.75 Hz and a dynamic axial
compression of 900 to 1700N. Under the modified ISO loading specifications the flexion
–extension stiffness values for the two groups were determined to be significantly
different (p = 0.0002). The axial rotation stiffness values for the two groups were also
significantly different (p= 0.0027). The lateral bending stiffness values were comparable
for the two groups. The loading characteristics for flexion-extension were a little skewed
but that could not have affected the stiffness value determination as the slope of the linear
region of the loading and unloading cycles was considered for stiffness value
determination. The mean flexion-extension stiffness for the cadaveric specimens was
found to be 2.64±0.23 Nm/º and that of synthetic implants (E-d) was found to be
1.4±0.17 Nm/º. The mean lateral bending stiffness for the cadaveric specimens was found
to be 2.47±0.87 Nm/º and that of the synthetic implants (E-d) was found to be 2.4±0.08
Nm/º. The mean axial rotation stiffness for the cadaveric specimens was found to be
2.93±0.69 Nm/º and that of synthetic implants (E-d) was found to be 1.2±0.07 Nm/º.
Thus, the two groups have different structural properties under complex modified ISO
loading. From the nature of the graphs for different loadings it can be seen that all the
implants behaved fairly similarly as expected while the cadaveric specimens showed
variability. In certain cases where there was a sudden change in the moment values,
which could be attributed to a protruded bone spur.
Therefore, on the basis of the results that we obtained from this study, we
conclude that there are statistical differences with respect to structural stiffness values
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between the two groups with respect to transverse and saggital planes. No significant
difference was found in the stiffness values in the frontal planes when tested under ISO
specifications. On this basis, we reject the null hypothesis.
5.3 Fatigue characteristics comparison
One specimen from each of the three groups was tested under fatigue to
compare the durability of the three disc structures under the given modified ISO loading
specification. The axial stiffness and disc height dropped considerably for the biological
disc after nearly 1500 cycles of complex modified ISO loading. However, the stiffness
values and disc heights for the implants (both E-d and pseudo Charite´) remained
comparatively unaffected at similar loading cycles. This implies that the cadaveric discs
were weaker compared to the implants and could not withstand the loading conditions.
For testing the cadaveric specimens large portions of vertebral bodies were removed
because the testing equipments could only accommodate specimens with a height less
than 5.1 mm. This modification accelerated the early failure of the cadaveric disc
structure as the bony region failed first. Using intact vertebral bodies with larger test
grips would perhaps allow the specimens last longer than the average 1500 cycles at
which all the cadaveric specimens failed.
Another possible reason that might have been instrumental in an early failure of
the cadaveric disc structure might be the testing standards. We do not have enough
knowledge of the type of physiological loads the spine has to withstand but perhaps they
are too extreme as compared to the actual physiological loadings. If the implants pass this
test, they may be robust designs. But it also raises a question about the ISO testing
standards because if the biological structure failed at a very small percentage of the
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expected life (i.e. number of cycles) of the implant under testing environment, then
perhaps the testing specifications are too demanding and this may lead to rejection of
some good designs which may not last for the specified time under the testing conditions.
Failure was characterized by decreased height of the specimen. The cadaveric
disc structures failed very early and even though the intervertebral disc material did not
seem to fail, the disc structure collapsed due to the bony structure failure. A similar
failure was observed in all the cadaveric specimens. The bony vertebral bodies could not
bear the extreme loadings and experienced fracture due to fatigue.
Since other biological structures like facet joints, ligaments etc. would be present
in real life situations and the loads would be distributed among all the spinal structures,
the assumption that the intervertebral disc should bear such high loads as specified by the
ISO is questionable. Another question that this study raised is do we really need such a
robust implant? If other biological structures begin to deteriorate with age, would the
normal implant change the load distribution in other spinal structures? Testing the
cadaveric disc structures under ISO specifications gave us an insight into how much and
what kinds of loads (i.e. individually or in a combination) the intervertebral disc structure
can bear. Comparison with the available in vivo loading information can provide a better
understanding of the loads the biological structure can withstand.
The pseudo Charite´ proved superior to the other two disc structures in both axial
stiffness and axial shift comparison. A possible reason for this is that the design of the
pseudo Charite´ is different compared to the other two disc structures which share a
similar basic design. The difference in the material properties and the different design
might have contributed to the more robust structure of the pseudo Charite´.
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5.4 Limitations of the study
The results of this study should be considered in light of the following limitations:
• The limited number of cadaveric specimens was an important issue. Increasing
the number of specimens and including more segments from different donors
would give us a better understanding of the performance of the cadaveric discs
as the variability in the properties was high among the small group that was
studied.
• As discs from different vertebral levels were used, there could be a difference in
the segment material properties, even though the author does not have any
references to support this speculation. If segments from different levels could be
tested and compared, we might be able to understand if the results were affected
due to variation in material properties.
• The information regarding the age of the donors was not available to the author.
The age of the specimens could possibly have an effect on the results. Since
with specimens taken from old donors, there is a possibility that degeneration
had already set in.
• The information regarding the bone mineral density was not available to the
author. If any of the cadaveric specimens was osteoporotic, it might have
affected the results. Testing the specimens for bone mineral density could help
us understand the results better.
• The radiographs shown after were not taken immediately at failure as we used
the drop in disc height as our criteria and observed it until we were sure that the
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discs had failed. Therefore, we could not determine the specific bony region
where the failure might possibly have set in.
• We could not allow the specimen height of more than 5.1 cm due to the design
of the testing equipment and, thus, had to remove a considerable part of the
bones. This might have caused the cadaveric specimens to fail at a lower range
of cycles.
5.5 Future work
These limitations should be considered when designing the studies in the future,
and a few suggestions are:
• Designing 3D finite element models for the biological intervertebral disc
structures so that and the implants can facilitate the study of the performance of
these discs without actually performing the experiments. This would also aid in
changing the loading conditions and finding their effect on the different disc
designs. The effects due to variability of properties in cadaveric specimens
could be eliminated using a 3D model.
• A prior knowledge of the age, bone mineral density etc. would help in selecting
the appropriate cadaveric specimens for testing.
• The results obtained would be more reliable if discs from a single vertebral level
would be used.
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