A Thesis
entitled
The Effect of Head Restraint Material Properties, Initial Backset, and T1 Acceleration
Magnitude on the Risk of Whiplash Injury: A Finite Element Study
by
Dhanvin Sunil Desai
Submitted to the Graduate Faculty as partial fulfillment of the requirements for the
Master of Science Degree in Bioengineering
______Dr. Vijay K. Goel, Committee Chair
______Dr. Anand K.Agarwal, Committee Member
______Dr. Scott C. Molitor, Committee Member
______Dr. Patricia R. Komuniechi, Dean College of Graduate Studies
The University of Toledo
August 2013
Copyright 2013, Dhanvin Sunil Desai
This document is copyrighted material. Under copyright law, no parts of this document may be reproduced without the expressed permission of the author. An Abstract of
The Effect of Head Restraint Material Properties, Initial Backset, and T1 Acceleration Magnitude on the Risk of Whiplash Injury: A Finite Element Study
by
Dhanvin Sunil Desai
Submitted to the Graduate Faculty as partial fulfillment of the requirements for the Master of Science Degree in Bioengineering
The University of Toledo August 2013
Whiplash sprains and/or strains occur in 28-53% of motor vehicle collision victims, making it the most common type of injury in these collisions. The costs annually in the United States of whiplash injury are approximately between 4.5 to 8 billion dollars.
Whiplash can have long-term symptoms which can lead to chronic pain. Theories have linked the risk of whiplash injury to facet joints, ligaments, intervertebral discs, vertebral arteries, dorsal root ganglia, and neck muscles. Head restraints were invented in the 1960s to reduce spinal motion by limiting relative motion between the head and thorax. The effectiveness of headrests in reducing injury has been limited to only 13 to 18% reduction in neck injury claims. It was reported that 85% of all whiplash injuries occur during rear- end impacts.
A detailed cervical model was created using a female CT scan. The scan was imported into Mimics. The model was meshed in IA-FEMesh and then imported into
ABAQUS. The headrest model was created in SolidWorks and meshed in 3-Matic. All of the material properties were derived from literature. The bony structures were modeled as linear elastic material models and the discs and ligaments were modeled as non-linear models. This thesis aims to provide a detailed cervical spine finite element model and
iii studies the effect of headrest material properties, initial headrest distance, and T1 acceleration magnitudes on the risk of whiplash injury. The initial goal was to validate the model under static and dynamic loading. Static validation was done in flexion/extension, lateral bending, and axial rotation and comparing the data with cadaver means and standard deviations. Dynamic loading was performed by providing an input acceleration pulse to T1 and comparing the segmental motion with the cadaver data corridors.
A sled dummy test performed at The University of Toledo was used as input variables to the finite element model. An HIII 50th percentile dummy was used for the test. The Y and Z chest accelerations were used as input accelerations to the finite element model. An actuator arm was used to provide the force used to accelerate the sled.
The dummy was seated on a standard seat, which was bolted to the sled. The sled was placed on a wheel system that sat on the sled frame. A brake system was used to provide stopping force.
Four polyurethane material properties were derived from literature and were used as variables for the headrest material properties. The headrest was modeled as a hyperfoam material model in ABAQUS. The initial headrest distance was also varied from 0 mm, 5 mm, 10 mm and 20 mm. The final part was to vary the input T1 acceleration magnitude from 7.5 G, 9.8 G, 13.8 G, and 15 G. The risk of whiplash injury was quantified by measure lower cervical intra-discal pressures, lower cervical peak facet stresses, lower cervical ALL ligament strains, ALAR ligament strains, and lower cervical facet capsular ligament strains. These variables have been defined in literature as the
iv most common sites for cervical spine injury during whiplash. The risk of head trauma was also quantified by measuring the peak head stress during impact.
Results showed a stiffer headrest material decreased the risk of cervical soft tissue injury while increasing the risk of head trauma. As the initial backset distance increased, the risk of cervical soft tissue injury increased and the risk of head trauma increased. As the T1 acceleration magnitude increased, the risk of cervical soft tissue injury increased and the risk of head trauma increased. It can be concluded that these three factors play a role in the risk of whiplash injury. It can also be concluded that a less stiff material should not be used if the goal is to reduce cervical soft tissue injury. The ideal way to reduce the risk of injury would be to make sure the initial headrest distance away from the head is as minimal as possible.
v I would like to dedicate this work to my family and friends. Thank you.
vi Acknowledgements
I would like to acknowledge my advisor Dr. Vijay Goel. His knowledge, experience and guidance have allowed me to reach this point. I would also like to thank my committee members, Dr. Anand Agarwal and Dr. Scott Molitor for their support and mentorship. I would also like to thank Dr. Deniz Erbulut for his assistance with this project. I would like to thank the University of Toledo, Department of Bioengineering, and Koc University for funding this project. I would like to thank all of my lab members, specifically Vivek Palepu for his continued support throughout this project. I would also like to thank Aakash Agarwal, Vikas Kaul and Ata Kiapour for their continued assistance on this project.
vii Table of Contents
Abstract ...... iii
Acknowledgements ...... vii
Table of Contents ...... viii
List of Tables ...... xi
List of Figures ...... xii
List of Abbreviations ...... xx
List of Symbols ...... xxi
1 Introduction ...... 1
1.1 Overview ...... 1
2 Literature Review...... 6
2.1 Overview ...... 6
2.2 Individual Structures ...... 7
2.2.1 Body ...... 7
2.2.2 Pedicles ...... 7
2.2.3 Processes ...... 7
2.2.4 The Laminae ...... 8
2.3 Anatomy of the Cervical Spine ...... 8
2.3.1 Common Sections (C3-C6) ...... 8
viii 2.3.2 Vertebral Section – C1 ...... 10
2.3.3 Vertebral Section – C2 ...... 12
2.3.4 Vertebral Section – C7 ...... 14
2.4 Current Head-Restraint Effectiveness ...... 14
2.5 Head-Restraint Safety Standards ...... 15
2.6 Active Head-Restraint Systems ...... 16
2.7 Whiplash Injury ...... 17
2.8 Types of Whiplash Injury ...... 18
2.9 Whiplash Statistics and Cost ...... 19
2.10 Computer Modeling of Whiplash ...... 20
2.10.1 Multi-Body Modeling ...... 21
2.10.2 Finite Element Modeling ...... 24
2.11 Human Modeling and Testing ...... 24
2.12 In Vivo Testing ...... 25
2.13 In Vitro Testing ...... 26
2.14 Anthropomorphic Model Testing ...... 26
3 Materials and Methods ...... 27
3.1 Model Description ...... 27
3.2 Bony Structures ...... 29
3.3 Intervertebral Discs and Facet Joints ...... 34
3.4 Ligaments ...... 37
3.5 Muscles and Pre-load ...... 39
3.6 Model Validation ...... 40
ix 3.6.1 Static Validation...... 40
3.6.2 Dynamic Validation ...... 41
3.7 Experimental Test Setup ...... 43
3.8 Headrest Material Property Test ...... 44
3.9 Headrest Distance Test ...... 48
3.10 Acceleration Profile Test ...... 50
4 Results ...... 53
4.1 Overview ...... 53
4.2 Validation Results ...... 53
4.3 Test Results ...... 59
5 Discussion ...... 72
5.1 Overview of Results ...... 72
5.2 Limitations and Future Work ...... 76
5.3 Conclusions ...... 76
References ...... 78
A Dynamic Result Curves ...... 87
x List of Tables
3.1 Bony mass properties ...... 33
3.2 Material properties for bony structures ...... 34
3.3 Mooney-Rivlin constants for nucleus pulposus model[85] ...... 35
3.4 Mooney-Rivlin constants for annulus fibrosus model[86]...... 35
3.5 Upper cervical spine (C0-C2) ligament material properties[84] ...... 37
3.6 Lower cervical spine (C2-C7) ligament material properties[84] ...... 38
3.7 Ogden coefficients for soft foam[94] ...... 46
3.8 Prony series constants for soft foam[94] ...... 46
3.9 Ogden coefficients for medium foam 1[95] ...... 46
3.10 Prony series constants for medium foam 1[95]...... 46
3.11 Ogden coefficients for medium foam 2[96] ...... 47
3.12 Prony series constants for medium foam 2[96]...... 47
4.1 Intra-discal pressure under static loading[88] ...... 55
4.2 Facet loads under static loading (C0-C4)...... 55
4.3 Facet loads under static loading (C4-C7)...... 56
xi List of Figures
2-1 Full human spine with various regions[66] ...... 10
2-2 First Cervical Vertebra[66] ...... 12
2-3 Second Cervical Vertebra[66] ...... 13
2-4 Whiplash Injury Mechanism[67] ...... 17
2-5 Various cervical computer models[77] ...... 21
3-1 Modeling the cervical spine ...... 28
3-2 Meshing the cervical spine ...... 28
3-3 Cervical spine model...... 29
3-4 Meshed skull ...... 30
3-5 Meshed C1 vertebra ...... 30
3-6 Meshed C2 vertebra ...... 31
3-7 Meshed C3 vertebra ...... 31
3-8 Meshed C4 vertebra ...... 31
3-9 Meshed C5 vertebra ...... 32
3-10 Meshed C6 vertebra ...... 32
3-11 Meshed C7 vertebra ...... 32
3-12 Meshed T1 vertebra ...... 33
3-13 Meshed C6-C7 disc ...... 35
xii 3-14 Pressure over-closure relationship[87] ...... 36
3-15 C6-C7 disc with uncoverteral clefts ...... 37
3-16 Non-linear stress-strain curve for lower cervical spine ligaments[88] ...... 39
3-17 Flexion/Extension, lateral bending, and axial rotation[92] ...... 40
3-18 Stemper Et. Al. experimental setup[91] ...... 41
3-19 Stemper Et. Al. specimen characteristics[91] ...... 42
3-20 Dynamic validation T1 acceleration input[91] ...... 42
3-21 Loading and boundary conditions for dynamic validation ...... 42
3-22 Sled test setup ...... 43
3-23 Sled unfiltered acceleration data ...... 44
3-24 Sled filtered acceleration data ...... 44
3-25 Foam material property stress-strain curves ...... 48
3-26 Zero mm backset ...... 49
3-27 Five mm backset ...... 49
3-28 Ten mm backset ...... 50
3-29 Twenty mm backset ...... 50
3-30 Peak T1 acceleration of 7.5 G ...... 51
3-31 Peak T1 acceleration of 9.8 G ...... 51
3-32 Peak T1 acceleration of 13.8 G ...... 52
3-33 Peak T1 acceleration of 15 G ...... 52
4-1 Upper cervical spine range of motion at 2 Nm[84] ...... 54
4-2 Lower cervical spine range of motion at 1.5 Nm[90] ...... 54
4-3 C2-C3 relative motion under dynamic loading[91] ...... 57
xiii 4-4 C3-C4 relative motion under dynamic loading[91] ...... 57
4-5 C4-C5 relative motion under dynamic loading[91] ...... 58
4-6 C5-C6 relative motion under dynamic loading[91] ...... 58
4-7 C6-C7 relative motion under dynamic loading[91] ...... 59
4-8 Head-T1 relative motion under dynamic loading[91] ...... 59
4-9 Cervical spine motion during simulated impact ...... 60
4-10 Pressure distribution across C6-C7 nucleus ...... 61
4-11 Peak intra-discal pressure for varying foam properties ...... 63
4-12 Peak facet stress for varying foam properties ...... 63
4-13 Facet capsule strain for varying foam properties ...... 64
4-14 Anterior ligament strain for varying foam properties ...... 64
4-15 Peak intra-discal pressure for varying initial backset ...... 66
4-16 Peak facet stress for varying initial backset ...... 67
4-17 Facet capsule strain for varying initial backset ...... 67
4-18 Anterior ligament strain for varying initial backset ...... 68
4-19 Peak intra-discal pressure for varying T1 acceleration magnitudes ...... 70
4-20 Peak facet stress for varying T1 acceleration magnitudes ...... 70
4-21 Facet capsule strain for varying T1 acceleration magnitudes ...... 71
4-22 Anterior ligament strain for varying T1 acceleration magnitudes ...... 71
5-1 Peak head stress during impact for varying foam material properties ...... 73
5-2 Peak head stress during impact for varying initial backset distance ...... 74
5-3 Peak head stress during impact for varying T1 acceleration magnitudes ...... 74
A-1 Range of motion for model with no headrest, at 7.5 G acceleration ...... 87
xiv A-2 Facet stress for model with no headrest, at 7.5 G acceleration ...... 88
A-3 Intra-discal pressure for model with no headrest, at 7.5 G acceleration ...... 88
A-4 Anterior ligament strain for model with no headrest, at 7.5 G acceleration ...... 89
A-5 Facet capsule strain for model with no headrest, at 7.5 G acceleration ...... 89
A-6 Range of motion for model with soft foam headrest, 10 mm backset and 7.5 G
acceleration ...... 90
A-7 Facet stress for model with soft foam headrest, 10 mm backset and 7.5 G
acceleration ...... 90
A-8 Intra-discal pressure for model with soft foam headrest, 10 mm backset and 7.5 G
acceleration ...... 91
A-9 Anterior ligament strain for model with soft foam headrest, 10 mm backset and
7.5 G acceleration ...... 91
A-10 Facet capsule strain for model with soft foam headrest, 10 mm backset and 7.5 G
acceleration ...... 92
A-11 Range of motion for model with medium foam 1 headrest, 10 mm backset and 7.5
G acceleration ...... 92
A-12 Intra-discal pressure for model with medium foam 1 headrest, 10 mm backset and
7.5 G acceleration ...... 93
A-13 Facet stress for model with medium foam 1 headrest, 10 mm backset and 7.5 G
acceleration ...... 93
A-14 Anterior ligament strain for model with medium foam 1 headrest, 10 mm backset
and 7.5 G acceleration...... 94
xv A-15 Facet capsule strain for model with medium foam 1 headrest, 10 mm backset and
7.5 G acceleration ...... 94
A-16 Range of motion for model with medium foam 2 headrest, 10 mm backset and 7.5
G acceleration ...... 95
A-17 Intra-discal pressure for model with medium foam 2 headrest, 10 mm backset and
7.5 G acceleration ...... 95
A-18 Facet stress for model with medium foam 2 headrest, 10 mm backset and 7.5 G
acceleration ...... 96
A-19 Anterior ligament strain for model with medium foam 2 headrest, 10 mm backset
and 7.5 G acceleration...... 96
A-20 Facet capsule strain for model with medium foam 2 headrest, 10 mm backset and
7.5 G acceleration ...... 97
A-21 Range of motion for model with hard foam headrest, 10 mm backset and 7.5 G
acceleration ...... 97
A-22 Intra-discal pressure for model with hard foam headrest, 10 mm backset and 7.5 G
acceleration ...... 98
A-23 Facet stress for model with hard foam headrest, 10 mm backset and 7.5 G
acceleration ...... 98
A-24 Anterior ligament strain for model with hard foam headrest, 10 mm backset and
7.5 G acceleration ...... 99
A-25 Facet capsule strain for model with hard foam headrest, 10 mm backset and 7.5 G
acceleration ...... 99
xvi A-26 Range of motion for model with medium foam 2 headrest, 0 mm backset and 7.5
G acceleration ...... 100
A-27 Intra-discal pressure for model with medium foam 2 headrest, 0 mm backset and
7.5 G acceleration ...... 100
A-28 Facet stress for model with medium foam 2 headrest, 0 mm backset and 7.5 G
acceleration ...... 101
A-29 Anterior ligament strain for model with medium foam 2 headrest, 0 mm backset
and 7.5 G acceleration...... 101
A-30 Facet capsule strain for model with medium foam 2 headrest, 0 mm backset and
7.5 G acceleration ...... 102
A-31 Range of motion for model with medium foam 2 headrest, 5 mm backset and 7.5
G acceleration ...... 102
A-32 Intra-discal pressure for model with medium foam 2 headrest, 5 mm backset and
7.5 G acceleration ...... 103
A-33 Facet stress for model with medium foam 2 headrest, 5 mm backset and 7.5 G
acceleration ...... 103
A-34 Anterior ligament strain for model with medium foam 2 headrest, 5 mm backset
and 7.5 G acceleration...... 104
A-35 Facet capsule strain for model with medium foam 2 headrest, 5 mm backset and
7.5 G acceleration ...... 104
A-36 Range of motion for model with medium foam 2 headrest, 20 mm backset and 7.5
G acceleration ...... 105
xvii A-37 Intra-discal pressure for model with medium foam 2 headrest, 20 mm backset and
7.5 G acceleration ...... 105
A-38 Facet stress for model with medium foam 2 headrest, 20 mm backset and 7.5 G
acceleration ...... 106
A-39 Anterior ligament strain for model with medium foam 2 headrest, 20 mm backset
and 7.5 G acceleration...... 106
A-40 Facet capsule strain for model with medium foam 2 headrest, 20 mm backset and
7.5 G acceleration ...... 107
A-41 Range of motion for model with medium foam 2 headrest, 10 mm backset and 9.8
G acceleration ...... 107
A-42 Intra-discal pressure for model with medium foam 2 headrest, 10 mm backset and
9.8 G acceleration ...... 108
A-43 Facet stress for model with medium foam 2 headrest, 10 mm backset and 9.8 G
acceleration ...... 108
A-44 Anterior ligament strain for model with medium foam 2 headrest, 10 mm backset
and 9.8 G acceleration...... 109
A-45 Facet capsule strain for model with medium foam 2 headrest, 10 mm backset and
9.8 G acceleration ...... 109
A-46 Range of motion for model with medium foam 2 headrest, 10 mm backset and
13.8 G acceleration ...... 110
A-47 Intra-discal pressure for model with medium foam 2 headrest, 10 mm backset and
13.8 G acceleration ...... 110
xviii A-48 Facet stress for model with medium foam 2 headrest, 10 mm backset and 13.8 G
acceleration ...... 111
A-49 Anterior ligament strain for model with medium foam 2 headrest, 10 mm backset
and 13.8 G acceleration...... 111
A-50 Facet capsule strain for model with medium foam 2 headrest, 10 mm backset and
13.8 G acceleration ...... 112
A-51 Range of motion for model with medium foam 2 headrest, 10 mm backset and 15
G acceleration ...... 112
A-52 Intra-discal pressure for model with medium foam 2 headrest, 10 mm backset and
15 G acceleration ...... 113
A-53 Facet stress for model with medium foam 2 headrest, 10 mm backset and 15 G
acceleration ...... 113
A-54 Anterior ligament strain for model with medium foam 2 headrest, 10 mm backset
and 15 G acceleration...... 114
A-55 Facet capsule strain for model with medium foam 2 headrest, 10 mm backset and
15 G acceleration ...... 114
A-56 Peak stress on head with varying foam properties, at 10 mm backset and 7.5 G
acceleration ...... 115
A-57 Peak stress on head with varying initial backset with medium foam 2 and 7.5 G
acceleration ...... 115
A-58 Peak stress on head with varying acceleration amplitudes with mediuam foam 2
and 10 mm backset...... 116
xix List of Abbreviations
ALL ...... Anterior Longitudinal Ligament CL ...... Capsular Ligament HR ...... Headrest IDP ...... Intra-discal pressure IV-NIC ...... Intervertebral Neck Injury Criterion NDC ...... Neck Displacement Criterion NIC ...... Neck Injury Criteria Nij ...... Normalized Neck Injury Criterion Nkm...... Neck Protection Criterion PMHS………………..Post Mortem Human Subjects ROM ...... Range of Motion
xx List of Symbols
Kg ...... kilo-gram MPa ....Mega-Pascal m ...... Meter mm .....Millimeter N ...... Newton s ...... seconds
xxi Chapter 1
Introduction
1.1 Overview
Whiplash sprains and/or strains occur in 28-53% of motor vehicle collision victims, making it the most common type of injury in these collisions.[1],[2] The costs annually in the United States of whiplash injury are approximately between 4.5 to 8 billion dollars.[3], [4] Whiplash can have long-term symptoms which can lead to chronic pain. Radanov et al.[5] reported that 24% of patients had symptoms one year after the accident and 18% had symptoms 2 year after the accident. It was also found that between
38 to 52% of whiplash injuries occurred in rear end collisions.[1],[6] Theories have linked the risk of whiplash injury to facet joints, ligaments, intervertebral discs, vertebral arteries, dorsal root ganglia, and neck muscles.[7] Studies also state that the underlying pathophysiology whiplash associated disorders remain unknown.[8],[9],[10] Another study reported that approximately 50% of whiplash victims had chronic neck pain 15 years after the accident.[11].[12] It was reported that 85% of all whiplash injuries occur during rear-end impacts.[13]
1 Clinical studies reported that approximately between 54 to 60% of whiplash patients suffered facet joint pain.[14],[15] Other dynamic rear end impact studies have also identified cervical facet joints as a possible source of injury.[16],[17],[18],[19] Pain symptoms caused by injury to these joints are similar to some described whiplash symptoms.[20],[21],[22] Studies have identified mechanoreceptive and nociceptive nerve fibers in facet joint capsules as possible sources of injury.[23],[24] Nociceptive fibers are sensitive to pain while mechanoreceptors produce noxious symptoms with excess tension.[24] Many of the recent investigations have been focusing on the facet joint as the possible source of injury.[25],[26],[27],[28]
In vitro and in vitro studies have reported that the lower cervical spine goes into hyperextension during whiplash mechanism.[29],[30] This hyperextension can result in excessive strains in the anterior soft tissues.[31] Clinical and biomechanical evidence shows anterior longitudinal ligament (ALL) and anterior annulus injures occur during whiplash. MRI studies, surgeries, and autopsies have reported injuries to the
ALL.[32],[33],[34] Cadaver and monkey studies simulating whiplash have also shown ALL tears.[35],[36] These injuries were reported in a cervical spine model in the lower cervical spine due to hyperextension.[37] Other MRI studies have also shown a high prevalence of injurious findings in the ALAR ligaments.[38]
Cervical disk lesions have been consistently documented in radiographic studies in whiplash. Petterson et. al[39] reported in an MRI study that 25% of whiplash patients had cervical disc herniations. Another study reported that 20% of whiplash patients showed severe herniated discs.[40] Another MRI study also reported herniated discs and endplate cartilage injuries in patients.[41] The most common site of disc injury was
2 reported to be C5-C6.[39],[40],[41] Other clinical studies suggest that whiplash injury can accelerate the process of disc degeneration.[42],[43],[44] Watkinson et al reported a significantly greater rate of disc degeneration in whiplash patients for up to 10 years after trauma.[44] Hohl et al reported that 39% of patients without initial signs of disc degeneration had developed signs of disc degeneration in 5 to 10 years.[42] Another study reported that patients who suffered from whiplash underwent fusion 8 years earlier than those who did not suffer from whiplash.[43] Panjabi et al. also reported that the cervical intervertebral discs are at risk for injury due to excessive fiber strain, disc shear strain, and anterior axial deformation.[45]
Head restraints were invented in the 1960s to reduce spinal motion by limiting relative motion between the head and thorax. The effectiveness of headrests in reducing injury has been limited to only 13 to 18% reduction in neck injury claims. [46].[47].[48] It is hypothesized that the restraints are not positioned correctly to constrain physiologic head- neck motions. Studies have also hypothesized that older fixed position head restrains were not placed optimally and may lead to extensive head-neck hyperextension motions.[47],[49] Findings also suggest that the newer generation headrests are usually not positioned correctly.[50],[51],[52] Active head restraints were designed to decrease the distance between the head and headrest directly after a rear end impact. Active head restraints have shown some success in reducing the risk of whiplash injury.[53],[54]
However, these headrests are only available in a limited number of automobiles.[55]
Siegmund et al performed human volunteer testing and reported that head restraint position greatly affected the magnitude and timing of head and thorax kinematics.[56] He showed that increasing the backset resulted in increased segmental
3 motions. Svensson et al. used a rear impact dummy neck attached to a Hybrid III dummy to show that changing the head restraint backset greatly affected the forces and bending moments at the superior and inferior ends of the cervical spine as well as the head accelerations and angles.[57] Stemper Et al. used a finite element model to show the headrest limited head retraction and spinal kinematics on a segmental basis. He reported lower cervical extension magnitudes were decreased with smaller backsets. All though this is a finite element model, the authors only report segmental angles in their findings.[55]
Kitagawa et al. used a finite element model to vary the car seat design parameters.
He showed the greatest change in neck injury criteria was due to the change in the recliner stiffness. The authors did not provide what mathematical model was used to simulate the head restraint foam and they varied the stiffness +30% and -30% from the initial.[58] Another study found that backset had the largest influence on head-neck motion. They reported that the maximum head-torso displacement increased with increasing backset. The study also found that the increased stiffness of the seat-back frame resulted in increased maximum head-torso displacement, while a stiffer lower seat- back cushion combined with a deeper upper seat-back cushion showed a clear reduction of the head-torso displacement.[59] Many studies have reported the effect of increasing acceleration magnitudes on the risk of whiplash injury. Ito Et Al. reported that a rear end collision is most likely to occur at a peak T1acceleration of 5 g and above.[60]
There are various whiplash injury criteria to predict neck injuries and to evaluate the effectiveness of different systems. The Neck Injury Criterion (NIC) is based on the head-T1 relative acceleration and velocity.[61] The Normalized Neck Injury Criterion
4 (Nij) and the Neck Protection Criterion (Nkm) are based on the dynamic loads at the occipital condyles.[62],[63] The Neck Displacement Criterion (NDC)[64] is based on head-
T1 displacement motions and the Intervertebral Neck Injury Criterion (IV-NIC)[65] is based on the normalized rotations of each segmental level.
This thesis aims to address issues in the literature by providing a detailed head to
T1 finite element model. Most literature findings have reported the risk of whiplash injury by a criterion and/or by segmental motion. This thesis aims to provide specific results that may give detailed explanations behind the risk of whiplash injury. A finite element model has several advantages of the cadaver and dummy studies. Those advantages include cost effective testing and detailed results that you may not be able to get from those other studies. The limitations of using a finite model include non-exact material properties and loading conditions as they may be seen during in-vivo conditions.
The background and literature review are provided in chapter 2, the materials and methods are provided in chapter 3, and the results are provided in chapter 4.
1. The hypothesis was that a stiffer material would decrease the chance of
whiplash injury.
2. It was hypothesized that a lower initial headrest distance would decrease
the risk of whiplash injury.
3. It was hypothesized that a lower T1 magnitude would also lower the risk
of whiplash injury.
5 Chapter 2
Literature Review
2.1 Overview
The spine is a column of thirty-three bones grouped into five categories: cervical, thoracic, lumbar, sacral, and coccygeal. The series of bones are mostly flexible and winding in nature, however, the top three sections (the lumbar, the thoracic, and cervical) are considered movable or “true”, while the lower two regions (the sacral and coccygeal) are considered fixed or “false”. The lower two regions are considered false as these sections are fused together to form two bones. In some extremely rare cases, the number of bones in a section will be increased or decreased, with the deficiency being made up in another area. [66]
The individual sections of the spine, the vertebra, have common characteristics.
Most vertebra are made of the viz (which is the anterior section), and the neural arch
(which is the posterior section). These two sections form around an object called the vertebral foramen. The neural arch is made of a pair of pedicles and a pair of laminae.
The arch also supports seven processes: four articular, two transverse, and one spinous.[66]
6
2.2 Individual Structures
2.2.1 Body
The body makes up the gist of the spine as it permeates through most of the vertebral column. The surfaces of the body are flat and rough, and they are attached to the intervertebral fibrocartilages.[66]
2.2.2 Pedicles
The pedicles are two processes which project backward from the upper part of the body. Two concavities, the vertebral notches, exist above and below the pedicles. The vertebral notches form the intervertebral foramina, which facilitates the transmission of spinal vessels and nerves. [66]
2.2.3 Processes
There are a total of seven processes, which are grouped into three sections: the transverse, the spinous, and the articular. Two transverse processes project from both side sides of the point where the laminae joins the pedicle. The transverse processes are responsible for attaching muscles and ligaments.
The spinous process is aimed downward and backward from the junction of the laminae, and it also is responsible for attaching muscles and ligaments. The articular processes project from the intersection of the pedicles and laminae. The inferior surfaces
7 project downward while the superior sections project upward. The articular surfaces are coated with hyaline cartilage.[66]
2.2.4 The Laminae
The Laminae consist of two broad plates aimed backward from the pedicles.
They connect in the middle line posteriorly, which in turn completes the posterior boundary of the vertebral foramen.[66]
2.3 Anatomy of the Cervical Spine
The elements of the cervical spine are the smallest of all of the movable vertebrae.
Vertebra within the cervical spine can be recognized easily as a foramen exists within each transverse process. While the first, second, and seventh vertebrae in the cervical area of the spine have unique characteristics, the third through sixth sections share common properties.
2.3.1 Common Sections (C3-C6)
The human spine is made up of the cervical, thoracic, lumbar, sacrum, and coccyx regions (Figure 2-1). Within the common sections of the cervical spine, the body is smaller and broader on its sides. Both the anterior and posterior surfaces of the body are flat and equal in depth. Both the upper and lower surfaces are concave, however, the upper is concave transversely while the lower is concave backward. The upper surface also projects a lip on both of its sides, while the lower surface has laterally shallow concavities which accept the projecting lips of subjacent vertebra.
8 The spinous process within these sections is separated into two divisions, which are often not equal in size. The articular process are fused together to form the articular pillar, which projects from the junction of the pedicle and lamina. The transverse processes are punctured by the foramen transversarium, which allows the passage of the vertebral artery, the vertebral vein, and the sympathetic nerves in the upper six vertebrae.
The transverse processes within the common sections consist of anterior and posteriors parts. The anterior section of the transverse process is also referred to as the costal process or costal element. The costal element projects from the side of the body, and it is aimed lateralward in front of the foramen. The posterior section of the transverse process projects from the vertebral arch behind the foramen. The costal element ends in the anterior tubercle while the posterior section ends in the posterior tubercle.[66]
The lamiae in the common sections are thinner and narrow above than below.
The vertebral foramen within these sections are large and triangular in form. The pedicles are directed backward and lateralward, and they are attached to the body midway between the upper and lower borders. The pedicles are oriented in this manner so that the superior vertebral notch is as deep as the inferior, however, it is more narrow.[66]
9
Figure 2-1: Full human spine with various regions[66]
2.3.2 Vertebral Section – C1
The first section of the cervical vertebrae (C1) is referred to as the atlas, as it supports the globe of the head (Figure 2-2). The first in unique as it does not contain various sections commonly found in other parts of the spine. C1 lacks a spinous process, and it does not have a body, as the body of the atlas is fused to the next vertebra. This section is ring-like in nature and it contains both a posterior and anterior arch and two lateral masses.[66] The anterior arch’s surface is convex, and its center contains the anterior tubercle, which allows for the attachment of the Longus colli muscles. The arch is concave on its posterior, and it is adorned with a circular facet. The upper and lower
10 borders, respectively, allow for the attachment of the anterior and atlantooccipital membrane and the anterior atlantoaxial ligament.[66] The atlantooccipital connects the anterior arch to the occipital bone.
The posterior arch of C1 ends behind the posterior tubercle and has some similarities to the spinous process. The Recti capitis posteriors minors spawn from the posterior arch as well. The small size of the posterior arch prevents the atlas and skull from interfering because of movement. The posterior section of the arch projects above and behind a rounded edge, which allows for attachment to the posterior atlantooccipital membrane. Behind each superior articular process is a groove that the represents the superior vertical notch. This notch aids in the transmission of the vertebral artery, which winds around the lateral mass.[66]
The lateral masses are the largest and most solid parts of the atlas as they help support the weight of the head. Each mass has a superior and inferior facet. The superior facets are large, oval, concave, and while they approach each other in the front, they diverge behind. These facets are directed in a manner that forms a cup for the corresponding condyle of the occipital bone, which allows for nodding movements of the head. The superior faces are usually subdivided by indentations as well.[66] The inferior articular facets are flat, circular, and slightly convex. These facets articulate with the axis, which allows for rotary movements of the head. A small tubercle exists below the middle margin of each superior facet. These tubercles are attached to the atlantal ligament, which spans across the ring of the atlas, thus dividing the vertebral foramen into two unequal parts. This division creates an anterior region which receives the
11 odontoid process of the across, and a posterior region which transmits the medulla spinalis and its membranes.
This subdivided region of C1 is much larger than what is needed for the medulla spinalis to fit. Because of this excess in size, lateral displacement of the atlas may occur even if the structure is not compressed. The transverse processes within C1 are large, and they are projected lateralward and downward from the lateral masses.[66] These processes are responsible for the attachment of the muscle that help rotate the head. The processes are longer, and their anterior and posterior tubercles are conjoined into one object.
Figure 2-2: First Cervical Vertebra[66]
2.3.3 Vertebral Section – C2
The second cervical section (C2) is referred to as the axis. It forms the pivot upon which the atlas (which holds the head) rotates (Figure 2-3). The most unique characteristic of the axis is the strong odontoid process which rises perpendicularly from the upper surface of the body.[66] The body of S2 is prolonged downward, causing it to
12 overlap the upper and forepart of the third vertebrae within the cervical spine. This overlap creates a longitudinal ridge which separates the two lateral depressions, allowing for the attachment of the Longus colli muscles. Within C2, the stronger odontoid process
(or dens) has a slight neck where it joins the body. A shallow groove is in place for the transverse atlantal ligament, which retains the process in position.[66]
The apex of C2 is pointed, and it is attached to the apical odontoid ligament. The process below the apex is enlarged, and rough surfaces exist on each side to aid the attachment of the alar ligament (which helps connect the process to the occipital bone).
The pedicles, laminae, and the spinous process are all strong and large sections of the second cervical vertebrae. In contrast, the transverse processes are small, and they all end in a single tubercle.[66] The superior anticular surfaces are supported on the pedicles, body, and transverse processes in this section. Behind the anticular are the superior vertebral notches, which are shallow in nature.
Figure 2-3: Second Cervical Vertebra[66]
13 2.3.4 Vertebral Section – C7
The seventh section of the cervical vertebrae (C7) at it contains the vertebra prominens, which is a thick, long spinous process. The process itself extends horizontal in direction and it does not branch off at any point; instead, it ends within a tubercle near the lower end of the ligamentum nuchae. The transverse processes are larger in C7 when compared to other areas of the cervical vertebra. The roots of the transverse process in this section are much thicker on the posterior side, but are much smaller on the anterior.
In rare cases the anterior root becomes large enough to be classified as a separate bone, the cervical rib. The surfaces within C7 have small grooves that hold the eighth spinal nerve.[66]
The size of the foramen transversarium in C7 varies from person to person; in some cases, its size in comparable in all sections of the vertebra, but it usually is smaller on its sides. In rarer cases, it may be double in size compared to other sections of the spine, while in other cases, it may be absent. The vertebral vein or artery sometimes moves down both sides of the foramen transversarium, but in some cases it may only move down one side. However, in most cases both the vertebral vein and artery are situated in front of the transverse processes instead.[66]
2.4 Current Head-Restraint Effectiveness
Correct positioning of a vehicle’s head restraint, or “headrest”, should stop the head from moving backwards during a sudden transmission of force. Unfortunately, there is often space between the head and the headrest, as vehicle occupants often sit with
14 their head tilted forward. If an occupants head is positioned in this fashion, during an impact, the head can jerk back into the head rest, which can result in brain contusions.
Head restrains are not adjusted correctly most of the time injury occurs; commonly, the head rest will be fully retracted. During an impact, the occupant can actually rise out of their seat, especially if the vehicle is struck from behind.[67] If the headrest is too low, and if the occupant’s body rises too far during impact, the head can actually cause the head to flip back over the top of the head restraint. Some modern cars have improved seat design, where the seat back itself is less rigid; in the case of an accident, the seat will give way to the body, which will decrease the force of the head moving backwards.[67]
2.5 Head-Restraint Safety Standards
In 1969, a measure called the Federal Motor Vehicle Safety Standards (FMVSS
202) was implemented standard safety features among all vehicles, which included head restraints. Even with the implementation of FMVSS 202, common symptoms associated with whiplash injury have been continually reported over the years.[68] Importantly,
FMVSS 202 did not place any requirements on head restraint position. Studies since the implementation of FMVSS 202 have shown it to only reduce the prevalence of neck injuries by only 20%. In 2004, the National Highway Traffic Safety Administration updated FMVSS 202 (now called FMVSS 202a).[68] The updated standards aimed to reduce the risk of whiplash injuries by requiring updates to head restraint position and geometry.
The studies used for the basis of FMVSS 202a found that a gap larger than 10 cm between the occupants head the head restraint prior to a read-end collision was
15 commonly associated neck injury symptoms lasting longer than one year.[68]
Computational models have shown that head restraint gaps of only 5-6 cm are sufficient enough to reduce neck motion and loads.[68]
2.6 Active Head-Restraint Systems
Studies using volunteers, test dummies, computer simulations, and cadavers all showed that standard non-active head restraints were not effective in limiting excessive head motions. However, rear impact simulations using test dummies showed that neck extension can be significantly eliminated through proper alteration of head restraint and seatback properties.
Some devices like active head restraints (AHR) and energy absorbing seats aim to alleviate the dangers associated with a gap larger than 10 cm.[68] Studies of active neck injury prevention systems stated that the use of said safety features could reduce whiplash by up to 75%. Active head restraints work by rotating forward based on the momentum of the occupant pressing back into the seat when whiplash occurs. While AHR systems aim to be reactive to an accident, their implementation does effectively address all of the mechanisms associate with whiplash.[68]
While the new standards implemented by FMVSS 202a aimed to improve security standards, their data was not based on biomechanical data. To properly test the effectiveness of AHRs using biomechanical data, a new human model of the neck
(HUMoN) was used for whiplash simulation. This model consisted of a neck specimen mounted on the torso of a rear impact dummy with an anthropometric head.[68] Tests specifically aimed to look for the relation between AHR position and peak neck motion.
16 While tests showed that peak neck motion was significantly reduced by the AHR when compared to non-active restraints, the peaks still exceeded the limits the flexion in section
C1 and the extension of C4-5 in the cervical spine.[68] Tests using the HUMoN model showed potential soft tissue damage even with the use of an AHR. The AHR 3 position
(the maximum gap position for the device) had the greatest potential for extension injury. Moreover, an AHR at position 3 caused the same peak motions in C5-6 and C6-7 allowed by non-active head restraints. Data showed that all other positions of the AHR significantly reduced neck motion, with positions 1 and 5 being the safest.[68]
2.7 Whiplash Injury
A whiplash injury, which can be defined as an acceleration injury, typically occurs when a stationary vehicle with an unaware occupant is struck from behind.
Symptoms related to whiplash include neck pain, dizziness, and headaches, all of which can be reported up to months or years after the accident.[69] Persistent pain after the accident can be from an acute neck tissue injury or an acute organic lesion.[68]
Figure 2-4: Whiplash Injury Mechanism[67]
17 In the United States, the National Highway Traffic Safety Administration estimated that 84% of all neck injuries could be classified as soft tissue injuries. These specific neck injuries are often sub-failure injuries, thus there is not a complete failure of soft tissues. In the past, MRI imaging did not have high enough resolution to identify such injuries. Decreased function and long term pain in the neck can possibly be explained by sub-failure injuries.[69]
Whiplash injury is common in automotive accidents because of the necks response to the sudden deceleration of a vehicle; in only minor rear-end accidents, the neck can be subjected to an extreme increase in acceleration. Soft tissue damage occurs in over 50% of all frontal-impact accidents while similar neck injuries occur in only 16% of all motor vehicle accidents.[70],[71] Other studies show rear impact have a higher incidence rate of soft tissue than any other type of motor vehicle accidents.[72],[73] The upper cervical spine is susceptible to injury because of its relatively low resistance to movement. This weakness is compounded by the weight of the head in comparison to the size and strength of the connective joints in the upper sections of the spine. These factors often lead spinal fractures and the failure of the prominent odontoid process.[70],[73]
2.8 Types of Whiplash Injury
Whiplash injury is classified by both type and severity in order to accurately provide treatment while also making the study of such injuries more easily accessible across multiple fields. Specifically, cervical injuries are described as Type A (involving compression), Type B (involving flexion-extension-distraction), or Type C (involving rotation). Each of these types are further classified by severity levels of I, II, or III.[70],[74]
18 Type A injuries (or Compression injuries) are characterized mostly by bone trauma. Wedge, burst, and teardrop fractures are described by severity levels of I, II, and
III respectively. Each severity level of a Type A injury can be characterized by a larger gap caused by a fracture in the anterior body of the spine.[70]
Type B injuries (or flexion-extension-distraction injuries) can be characterized by tension and rotation of the sagittal plane. Level I Type B cervical injuries are classified by minor sprains, and in rare cases, minor neurological deficits. Level II severity involves severe sprains resulting in ligament and soft tissue damage to the posterior ligaments and intervertebral disc, or fracture in the spinous process. Level III severity describes cases of bilateral dislocation and fracture.[70]
Type C injuries (or rotation injuries) involve lateral bending and axial rotations within the spine. Level I severity involves only a rotation causing single facet fractures.
Level II severity involves separation of the articulate pillar from the vertebra. Level III severity is characterized by unilateral dislocations.[70]
2.9 Whiplash Statistic and Cost
While whiplash rates vary by country, studies released by numerous European countries indicate that there is a relation between an increase in traffic density, and an increase in whiplash related injuries caused by vehicle collisions. This relation has become more evident in recent years where whiplash has become a more common occurrence in automobile accidents. In Japan alone, neck related injury is reported in approximately half of all vehicle based accidents.[69] From April 2000 to March 2001, insurance payout figures for Japan indicated that over one million people were injured in
19 automobile related accidents. Of the one million injured, over half a million injuries were neck related, and 340,000 of those injuries came from rear-end collisions. In total, neck related injuries resulted in costs upward of 317 Billion Yen, while neck injuries caused specifically by rear-end collisions accounted for 60% of that cost.[75] The high cost of neck injuries, specifically those caused by rear-end collisions is not contained. In
Europe, whiplash cost per person is highest in Switzerland, where the average claim is around 35,000 Euro.[68]
On average, more than one million Europeans experience whiplash related injuries each year, accounting for costs upwards of 10 billion Euros. In the United States, the National Highway Traffic Safety Administration reported that whiplash injuries cost approximately 4.5 billion dollars per year.[76] While whiplash typically causes short term injury, the cost per case can vary greatly as lifelong disability is possible. Studies based
Sweden indicates that whiplash injuries account for around 70% of conditions that lead to some kind of disability.[76]
2.10 Computer Modeling of Whiplash
While multi-body dynamics or finite element analysis (or combination of two) can be used to design head and neck models, there are more finite element (FE) models than multi body models. Although FE models require large amounts of computational power, they can provide detailed information about tissue deformations and injury predication.[77] Multi-body models, on the contrary, can contain numerous anatomical details while requiring only moderate amounts of processing power.
20 2.10.1 Multi-Body Modeling
The nature of multi-body models makes them applicable for optimization analysis and parameter variation. The multi-body model contains head and vertebrae reproductions that are comprised of ridged bodies and soft tissue (which includes intervertebral discs, facet joints, ligaments, and muscles), which allows the model to produce biofidelic responses.[77] Multi-body models are capable of predicting whiplash injury as they can simulate displacements of the head with respect to the torso, while also tracking accelerations, intervertebral motions, and force moments.[77] Numerous multi- body models are available, each with its own benefits and limitations.
Figure 2-5: Various cervical computer models[77]
Some of the older cervical spine models are shown in figure 2-5. The model created by Jakobsson et al. is a head-neck system made up of revolute joints which applied resistance to motion according to specified torque versus rotation functions.[77]
The model itself formed the cervical section of the spine, and it was designed to work in the sagittal plane. Only qualitative comparisons of single-impact speed conditions were made using the model, and the time-dependency of the passive applied muscle behavior was not considered during testing.[77]
21 A later model designed by de Jager became an improvement on a previous iteration, the global model, which was limited by its concentration of soft tissue behavioral mechanics located in the intervertebral joints. The improved version was considerably more detailed, containing active muscles, frictionless facet joints, non-linear viscoelastic ligaments, and linear viscoelastic ligaments.[77] The model showed reasonable response in tests and it was validated against front and lateral impacts because of its detailed nature, and because it was comprised of muscles that were formed as straight line elements. The de JaFiger model was later used by Kroonenberg et al. to comprise the head-neck portion of multi-body rear impact full human-body model. Both high and low severity responses recorded by the Kroonenberg et al. model were comparable to past rear-impact sled tests using either volunteers, or human cadavers.
Unfortunately, a lack of experimental testing data at the time did not allow for complete validation of the model. A later model created by Yamazaki et al improved upon de
Jager’s design by changing joint resistance properties and increasing optimization using data from volunteer and sled tests performed by the Japanese Automobile Research institute.[77]
Response rates in multi-body head-and-neck models have improved over the years through a variety of techniques. Linder used mathematical modeling to create to create a rear-impact dummy neck, which was eventually used by Humanetics™ to create their BioRID dummy. The BioRID (which was also labeled as the Anthropometric Test
Device) used two cables in the front and back of the neck as a muscle substitute. Tests using sensitivity analysis showed that combining elastic stiffness and dampening in the muscle substitutes along with non-liner joint stiffness produced better response rates than
22 current neck models available at the time.[77] Results also showed that a neck with only revolute joints and no muscle substitutes were incapable of producing biofidelic responses.
Much like Yamazaki, van der Horst created a model that also improved upon de
Jager’s design. The model incorporated muscles that were able to follow the curvature of the neck, which allowed for more realistic muscle force lines of action.[77] The head-neck design by de Jager was incorporated into a full-body model using the software and analysis software Madymo (MAthematical DYnamic MOdels). During testing, the head- neck model was not isolated, so the entire body was used for rear-impact analysis. Sled tests involving both cadavers and volunteers were simulated through the use of passive muscles.[77] Post-mortem changes in the passive muscle properties of cadavers were studied in a separate simulation. Volunteer tests involved simulations where neither the head restraint nor a seatbelt were present. A separate simulation included the use of both a rigid head restraint and regulation seatbelt. During tests where no safety features were present, there were significant differences between active and passive muscle response, however, simulations where safety features were present showed little to no correlation with volunteer data.[77]
Stemper et al also used Madymo to create a head-neck model, which differed from de Jager’s design through the use of segmented contractile muscles. The model was proved valid through the use three different kind of joint kinematic corridors (global, segmented, and face) which came from modeling different cadaver head-neck complexes where both the skin and muscle structure were still intact.[77] Active muscle effects were not considered in the design of the Stemper et al. model. A detail model created by van
23 Lopik, however, incorporated both passive and active muscle behavior through the program MSC VisualNastran 4D.
2.10.2 Finite Element Modeling
Finite Element models were developed in an attempt to evaluate dynamic response and susceptibility during rear-end collisions[78]. FE models can accurately portray both the geometry and organic properties of both the head and neck, however, these models requires large amounts of processing power.[78] Computer simulations using
FE models can take upwards of 50 hours, and are limited further by the lack accurate material properties. While limitations exist, FE modeling, in comparison to multi-body modeling, can produce a more complete picture of where specific injuries occur during whiplash.[78]
2.11 Human Modeling and Testing
While computational modeling of whiplash can allow for greater insight into both the cause and severity of an injury, actual human testing is needed for the basis of computer models. Testing involving physical models includes actual human volunteers
(also called in vivo), cadavers (also called in vitro), and anthropomorphic test dummies
(ATDs).[78] While testing involving human participants allows for more realistic response rates, experimentation is limited to only low-speed rear impacts to avoid long- term injury. Cadaver testing allows for high-speed rear impact studies, but the stiffer muscle structure within in vivo subjects does not allow for proper muscle activation.
Anthropomorphic test dummies are full sized models of the human body based on in
24 vivo, in vitro, and computational models. While ATDs are anatomically similar to its organic counterparts, they are not as responsive to rear-impact tests as actual humans.[78]
2.12 In Vivo Testing
Human testing over the last 30 years has involved both male and female participants who were subjected to low speed rear-end impacts using either a test sled or vehicle. While tests sometimes involved the use of a head restraint, testing without the use of said restraint were of greater interest because it allowed for an uninhibited motion of the head and neck during impact, which allowed for better insight into the mechanical properties of whiplash.[78] Testing also focused mainly on situations where the subject was looking forward, although, tests where the participant’s head was turned have been performed as well.
During in vivo testing, both electromyographic (EMG) and kinematic response are the main areas of focus. Through the use of accelerometers and reflective targets, kinematic response can be used to track target areas on either the head or neck. Factors like position, velocity, acceleration, center of gravity, and the rotation of the head are all recorded and accounted for.[78] The tracked information is useful, but aspects of the various kinds of intervertebral rotations and motions are most important to understanding whiplash mechanics. While this information can be hard to obtain, advancements in x- ray cinematography have allowed for high-speed tracking of the vertebra during in vitro testing. X-ray cinematography has allowed researchers to capture the motion of the cervical spine as a whole, along with the individual vertebra as well. In conjunction with x-ray cinematography, EMG signals allow for the capture of muscle activation as well.[78]
25 2.13 In Vitro Testing
Human participant research is limited as tests must be performed at low velocities to prevent short term or chronic injury within subjects. As In Vitro testing only involved the use of cadavers, tests can be conducted at high speeds, allowing for the simulation of high acceleration levels associated with whiplash.[78] Tests involve full-body or head- neck cadavers, which are subjected to uniaxial acceleration to the upper thoracic vertebra.[78] While tests of in vitro subjects allow models to be subjected to more realistic impacts, the muscle stiffness associated with cadavers does not allow for natural muscle reaction. In Vitro testing also does not allow for head of the subject to be held in place during impact because of the lack of muscle activation. By removing muscle tissue and supporting the head in its normal posture, and by using systems to replicate muscle properties prior to impact, the limitations of In Vitro testing can be reduced.[79]
2.14 Anthropomorphic Model Testing
Anthropomorphic models (also referred to as “test dummies”) are full body replicas used to simulate the effects of vehicle impacts on human, and population sampling has allowed for the creation of models specifically suited for the study of rear- impact collisions. Models are made to represent the body mass and properties of the average human in order to best replicate muscle and skeletal response during whiplash.
While ATDs can replicate some behavior of whiplash, the construction of the neck joints within the model limits the accuracy of test results. Even though ATDs have trouble simulating the complex motions of the spine during vehicle collisions, their use has led to various safety improvements since their inception.[80]
26 Chapter 3
Materials and Methods
3.1 Model Description
The cervical model (C1-C7) used for this thesis was developed by Vivek Palepu at the University of Toledo. My contribution to the model was the development of the skull, development of the headrest, and assisting in static and dynamic validation of the cervical model. The cervical spine was modeled from CT scans of a 22 year old normal healthy female. The CT scans were imported into Mimics software (Materialise,
Belgium) as DICOM files. Mimics was used for segmentation of the regions of interest.
27
Figure 3-1: Modeling the cervical spine
Once the regions were segmented, 3D geometries were created and cleaned up for meshing purposes. The cervical spine (C7-C1) was meshed in IA-FE Mesh (Iowa, United
States). In this software, STL files were imported from Mimics. Blocks were generated to match to contours of the outer surface of each geometry separately. The software, then automatically creates a volumetric hexahedral mesh of the bone or soft tissue.
Figure 3-2: Meshing the cervical spine
28 The skull mesh was generated in 3-Matic (Materialise, Belgium). The STL file was imported in to the software and an auto mesh was created using the default parameters. The autofix command was used to try to resolve any elements that can cause errors during the analysis. The bones and discs were modeled as 3-D solid continuum elements. The ligaments were modeled as 1-D truss elements. All of the materials properties used for this model were derived from literature. The model consists of 48 shell elements, 1555 truss elements, 16 connector elements, 185392 hexahedral solid elements, and 136280 tetrahedral solid elements. The model has 323291 elements and
257851 nodes in total.
Figure 3-3: Cervical spine model 3.2 Bony Structures
The following pictures provide different views of the meshed bony structures used in this study. As you can see, the skull is made up of all tetrahedral elements. It consists of 109543 elements and 30617 nodes. The C1 vertebra consists of 6351 hexahedral elements and 8446 nodes. The C2 vertebra consists of 27996 hexahedral elements and 32649 nodes. The C3 vertebra consists of 19080 hexahedral elements and
29 22911 nodes. The C4 vertebra consists of 19363 hexahedral elements and 23206 nodes.
The C5 vertebra consists of 25497 hexahedral elements and 30090 nodes. The C6 vertebra consists of 32503 hexahedral elements and 37631 nodes. The C7 vertebra consists of 33091 hexahedral elements and 38634 nodes. The T1 vertebra consists of
6256 hexahedral elements and 7912 nodes.
Figure 3-4: Meshed skull
Figure 3-5: Meshed C1 vertebra
30
Figure 3-6: Meshed C2 vertebra
Figure 3-7: Meshed C3 vertebra
Figure 3-8: Meshed C4 vertebra
31
Figure 3-9: Meshed C5 vertebra
Figure 3-10: Meshed C6 vertebra
Figure 3-11: Meshed C7 vertebra
32
Figure 3-12: Meshed T1 vertebra
The mass properties used for the bony geometry are shown in the following table
(Table 3-1). The densities of the cortical, cancellous, and posterior elements were all scaled to get the proper mass value. The moment of inertia is automatically accounted by geometry and density values.
Table 3.1: Bony mass properties Bone Mass (kg)
Skull[81] 3.5
C1[82] 0.22
C2[82] 0.25
C3[82] 0.24
C4[82] 0.23
C5[82] 0.23
C6[82] 0.24
C7[82] 0.22
T1[82] 10.00
33 All of the bony structures were modeled as linear elastic materials. Because we will not be studying bony fractures, it is alright to model them as reduced integration linear elastic materials. The vertebrae were positioned in such a way that they agreed with literature on normal seating posture. The head was position so it would be in the
Frankfort plane.
Table 3.2: Material properties for bony structures Bone Elastic Modulus (MPa) Poisson’s ratio
Skull-Cortical[83] 15000 0.21
Skull-Cancellous[83] 4500 0.01
Vertebrae-Cortical[84] 10000 0.3
Vertebrae-Cancellous[84] 450 0.25
Vertebrae-Posterior[84] 3500 0.25
3.3 Intervertebral Discs and Facet Joints
The intervertebral discs were modeled as solid 8-node hexahedral elements. Both the nucleus and annulus were modeled as incompressible hybrid elements. The discs are located between the vertebral bodies. The discs were modeled to have annulus fibrosus material and nucleus pulposus material. The volume of the nucleus was considered to be approximately 50 percent of the disc.[84]
34
Figure 3-13: Meshed C6-C7 disc
The nucleus pulposus was modeled as an incompressible hyper-elastic solid material.[85] The annulus fibrosis was modeled as hyper-elastic solid material.[86] A
Mooney-Rivlin formulation was used to model the hyper-elasticity. According to
ABAQUS documentation, the strain energy potential is defined by equation 3.1.
( ̅ ) ( ̅ ) ( ) , [3.1]
Table 3.3: Mooney-Rivlin constants for nucleus pulposus model[85] Parameter Value
C10 0.12
C01 0.03
D1 0
Table 3.4: Mooney-Rivlin constants for annulus fibrosus model[86] Parameter Value
C10 0.7
C01 0.2
35 The annulus fibrosus is made up or a series of fibers embedded in the substance.
This was modeled using tension only truss elements. The fibers were made ±65 degrees with respect to the vertical axis. The annulus fibrosus was made up of eight layers and they had a collagenous fiber content of about 16%. The fiber material properties were determined to be 500 MPa Young’s modulus and 0.4 Poisson’s ratio.[84]
The facet joints were modeled as surface interactions. The initial gaps of the facets were defined from CT scan geometry. The surface interaction was modeled as an exponential pressure increase with zero pressure starting at the initial gap. The pressure was defined as the posterior bone’s Young’s modulus when the facets came in contact.[84]
The contact was modeled as frictionless to simulate synovial fluid behavior. An example is given below.
Figure 3-14: Pressure over-closure relationship[87]
The uncoverteral (Luschka’s) clefts were also modeled in the intervertebral discs.
Elements were removed bilaterally in the outer layer of the annulus fibrosus.
36
Figure 3-15: C6-C7 disc with uncoverteral clefts
3.4 Ligaments
The ligaments in this model were created using 1D truss elements. These truss elements behave as tension only ropes. They were divided in to multiple elements. Each element was given a cross sectional area. These cross sectional areas were specified in such a way that they would match the total in vivo data. All of the ligaments, except the transverse and tectoral membrane, were modeled as hypo-elastic. The coefficients were derived to match uniaxial stress vs. strain curve data so they can have different stiffness during the different regions. Figure 3-16 shows the nonlinear stress-strain curve of the lower cervical spine ligaments.
Table 3.5: Upper cervical spine (C0-C2) ligament material properties[84] Ligament Young’s Modulus (MPa) Poisson’s Cross-
Ratio Sectional
Area (mm2)
Transverse 20 0.3 18
Superior/inferior cruciform 6.0 (<17 %), 10.0 (>17 %) 0.3 5
Alar 5.0 (<17 %), 8.5 (>17 %) 0.3 22
37 Apical 6.0 (<17 %), 10.0 (>17 %) 0.3 5
Accessory 6.0 (<17 %), 10.0 (>17 %) 0.3 5
Nuchal 12.0 (<17 %), 20.0 (>17 %) 0.3 5
Tectoral membrane/ligament 6.3 0.3 5
Anterior Longitudinal 12.0 (<17 %), 20.0 (>17 %) 0.3 6
Posterior C0-C1 capsular 6.0 (<17 %), 10.0 (>17 %) 0.3 5
Anterior C0-C1 capsular 6.0 (<17 %), 10.0 (>17 %) 0.3 5
Posterior C1-C2 capsular 6.0 (<17 %), 10.0 (>17 %) 0.3 5
Anterior C1-C2 capsular 0.2 (<17 %), 1.25 (>17 %) 0.3 5
Table 3.6: Lower cervical spine (C2-C7) ligament material properties[84] Ligament Young’s Modulus (MPa) Poisson’s Cross-
Ratio Sectional
Area (mm2)
Anterior Longitudinal 15.0 (<12 %), 30.0 (>12 %) 0.3 6.1
Ligament
Posterior Longitudinal 10.0 (<12 %), 20.0 (>12 %) 0.3 5.4
Ligament
Capsular Ligament 7.0 (<30 %), 30 (>12 %) 0.3 46.6
Ligamentum Flavum 5.0 (<25 %), 10.0 (>25 %) 0.3 50.1
Interspinous Ligament 4.0 (20-40 %), 8.0 (>40 %) 0.3 13.1
38
Figure 3-16: Non-linear stress-strain curve for lower cervical spine ligaments[88]
3.5 Muscles and Pre-load
No muscles were included in this model. This model tries to predict the behavior in the worst case scenario. Yoganandan Et Al has reported that muscles do not activate until 200 milliseconds. Therefore this model will only try to predict the behavior before
200 ms.[89] A bilateral compressive pre-load was included in this model. The pre load acts to stabilize the spine to reproduce a more physiological biomechanical response. The total magnitude of the compressive force for each functional spinal unit is 73.6 N. The force follows along the line of spinal curvature.[84]
39 3.6 Model Validation
The model was validated under two different loading conditions: static and dynamic. Experimental static data was used from Potluri Et Al. and Faizan Et. Al. and experimental dynamic data was used from Stemper Et Al.[84],[90],[91]
3.6.1 Static Validation
Static validation was done under flexion, extension, left bending, right bending, left rotation, and right rotation. A pre-load was applied in the first step and the respective moments were applied in the second step. A 2 Nm moment was applied in each respective direction to the top segment while the bottom segment was constrained in all directions. The rotation data was compared to that of cadaveric testing from Potluri Et.
Al. and Faizan Et. Al.
Figure 3-17: Flexion/Extension, lateral bending, and axial rotation[92]
40 3.6.2 Dynamic Validation
Dynamic validation was done by simulating the experiment performed by
Stemper Et Al. Five male and five female cadaver specimens were used for the study.
They were isolated at T2 and the skin and musculature were kept intact. They were potted in polymethylmethacrylate (PMMA) at T1. Since the T1 was fixed, thoracic ramping was not accommodated in this experiment. The rear impact was applied using a pendulum- minisled device. The pendulum strike at the posterior face of the fixture created the impact pulse. An accelerometer was attached to the mini-sled to measure posterior to anterior acceleration and this was used as the input to the FE model. The relative motion of the cervical vertebrae in the FE model was compared to that of the experiment.[91]
Figure 3-18: Stemper Et. Al. experimental setup[91]
41
Figure 3-19: Stemper Et. Al. specimen characteristics[91]
Figure 3-20: Dynamic validation T1 acceleration input[91]
Figure 3-21: Loading and boundary conditions for dynamic validation
42
3.7 Experimental Test Setup
A dummy whiplash experiment was performed at the University of Toledo by Dr.
Deniz Erbulut. In this experiment, a car seat was mounted on a sled. A HIII dummy was used for this experiment. Accelerometers were placed on the sled and the head and sternum of the dummy. A pulse simulating low speed rear end impact was applied to the sled with a hydraulic linear actuator. The respective accelerations were recorded using a data acquisition system. Once the data was recorded, they were processed in MATLAB
(Mathworks, Massachusetts, USA) with a low pass Butterworth filter. This was done to remove the high frequency noise and clean up the signal. The figures show effect of the filter on the signal. Tencer Et. Al showed that chest (sternum) acceleration provides a similar profile to that of T1 acceleration.[93]
Figure 3-22: Sled test setup
43
Figure 3-23: Sled unfiltered acceleration data
Figure 3-24: Sled filtered acceleration data
3.8 Headrest Material Property Test
The first experiment was to test for varying material properties of the head rest.
For this, a 3D model of a normal headrest was desined in SOLIDWORKS (Dassault
44 Systems, Massachusetts, USA) from geometrical measurements of the car seat used in the dummy experiment. The headrest was meshed using 3-Matic, which created a solid tetrahedral mesh. The headrest was imported into ABAQUS at a distance of 10 mm away from the posterior apex of the skull. This distance was chosen because it gave good contact and distraction with the skull during the impact. The input acceleration was derived from the dummy experiment. The sled and inferior end of the T1 were given chest-x acceleration profile. The inferior end of T1 was also given a chest-z acceleration profile to simulate thoracic ramping effects.
Material properties of the head restraint were derived from literature. The headrest is made up of polyurethane foam. Four different polyurethane foam properties were considered during this experiment. For the purpose of this experiment, they were called: soft foam, medium foam 1, medium foam 2, and hard foam.[94],[95],[96],[97] An Ogden hyper-foam model was used to define the head restraint material property. The material properties for the soft foam, medium foam 1, and medium foam 2 were defined using coefficients and they were given respective viscoelastic properties. For the hard foam, a curve fitting method was used since no coefficients were provided.[97] Curve fitting was performed using the Plot Digitizer software, and the stress-strain data was then imported into ABAQUS as uniaxial test data. The viscoelasticity was defined using Prony series parameters derived from literature. The hyper-foam Ogden model is given by equation
3.2. The equation for a Prony series expansion is defined by ABAQUS by equation
3.3.[87]
̂ ̂ ̂ ∑ [ (( ) )], [3.2]
45
[3.3] ( ) ∑ ̅ ( ) ,
Table 3.7: Ogden coefficients for soft foam[94] (MPa)
0.0127404 7.281 0
2.7459E-06 -5.7311 0
Table 3.8: Prony series constants for soft foam[94] gi ki τi
0.0973 0 0.30639
0.174 0 1121
0.129 0 1011
Table 3.9: Ogden coefficients for medium foam 1[95] (MPa)
0.018 8 0
0.0012 -2 0.45
Table 3.10: Prony series constants for medium foam 1[95] gi ki τi
0.05 0 0.001
0.05 0 0.003
0.05 0 0.01
46 0.05 0 0.03
0.05 0 0.1
0.05 0 0.3
0.05 0 1
Table 3.11: Ogden coefficients for medium foam 2[96] (MPa)
0.164861 8.88413 0
2.3017E-05 -4.81798 0
Table 3.12: Prony series constants for medium foam 2[96] gi ki τi
0.3003 0 0.010014
0.1997 0 0.1002
47
Figure 3-25: Foam material property stress-strain curves
3.9 Headrest Distance Test
The second experiment performed was to study the effect of varying headrest distance would have on whiplash injury mechanism. The same 7.5 G acceleration profile was used for this test. Medium foam 2 material model was used to define the head rest. It is a limitation that the headrest and T1 were given the same exact acceleration profile, but this will be discussed later. The properties for medium foam 2 gave good results and the output motion seemed to make sense, therefore it was used for this part of the study. The different headrest distances defined during this experiment were 0 mm, 5 mm, 10 mm
48 and 20 mm. The headrest distance was measured as the distance from the apex of the skull to its respective horizontal node on the headrest.
Figure 3-26: Zero mm backset
Figure 3-27: Five mm backset
49
Figure 3-28: Ten mm backset
Figure 3-29: Twenty mm backset
3.10 Acceleration Profile Test
For the last part of this experiment, the goal was test the effect of varying acceleration profiles on whiplash injury mechanism. The medium foam material property was used to simulate the headrest. A headrest distance of 20 mm was used since the peak rotations in the lower cervical spine reached before headrest contact. All of the
50 acceleration profiles were derived from the sled dummy test performed at our lab. The peak chest accelerations in the y-direction varied from 7.5 G, 9.8 G, 13.8 G, and 15 G.
Figure 3-30: Peak T1 acceleration of 7.5 G
Figure 3-31: Peak T1 acceleration of 9.8 G
51
Figure 3-32: Peak T1 acceleration of 13.8 G
Figure 3-33: Peak T1 acceleration of 15 G
52 Chapter 4
Results
4.1 Overview
The results are broken into several categories. The first step of this project was to validate the model under static and dynamic loading. Both loading conditions and data were taken from cadaveric experiments. The study was then broken up into three different tests. The first test was to study the effect of changing head rest material properties. The next test was to change the distance of the headrest away from the skull.
The third test was to study the effect of varying acceleration magnitudes on whiplash injury mechanism.
4.2 Validation Results
The first goal of this project was to validate the FE model under static loading conditions. The data was then compared to cadaver experimental data from Potluri Et.
AL. and Faizan Et. Al. It can be seen that the finite element model flexion plus extension data, lateral bending data, and rotation data falls within the standard deviation of the
53 experimental data. We can conclude that this model behaves accordingly under static loading conditions.
Figure 4-1: Upper cervical spine range of motion at 2 Nm[84]
Figure 4-2: Lower cervical spine range of motion at 1.5 Nm[90]
54 Table 4.1: Intra-discal pressure under static loading[88] Intradiscal Pressure (MPa) Current Cervical Model Oda (1981) Pospiech et al. (1999) Clausen (1996) Loading Mode C3-C4 C4-C5 C5-C6 C6-C7 Various levels C3-C4 C5-C6 C5-C6 Flexion 0.62 0.66 0.54 0.31 0.58 0.32 (0.12-0.43) 0.23 (0.00-0.56) 0.24 Extension 0.38 0.45 0.28 0.21 0.91 0.32 (0.12-0.43) 0.23 (0.00-0.56) 0.02 Left Bending 0.42 0.45 0.27 0.21 0.45 0.16 (0.08-0.31) 0.17 (0.00-0.38) 0.11 Right Bending 0.45 0.33 0.28 0.19 0.45 0.16 (0.08-0.31) 0.17 (0.00-0.38) 0.11 Left Rotation 0.31 0.55 0.39 0.24 0.45 0.25 (0.14-0.36) 0.16 (0.04-0.49) 0.14 Right Rotation 0.41 0.53 0.37 0.26 0.45 0.25 (0.14-0.36) 0.16 (0.04-0.49) 0.14
Table 4.2: Facet loads under static loading (C0-C4) Facet Load (N) C0-C1 Left Facet Right Facet Extension 98.41 100.60 Flexion 51.37 57.47 Left Bending 84.89 0.00 Right Bending 0.00 75.11 Left Rotation 15.24 68.50 Right Rotation 74.02 4.47 C1-C2 Left Facet Right Facet Extension 77.41 67.73 Flexion 34.05 48.16 Left Bending 74.12 0.00 Right Bending 0.00 77.38 Left Rotation 0.00 47.97 Right Rotation 49.90 0.00 C2-C3 Left Facet Right Facet Extension 53.63 77.77 Flexion 0.00 4.07 Left Bending 32.13 0.00 Right Bending 0.00 86.14 Left Rotation 0.00 49.99 Right Rotation 1.37 3.83 C3-C4 Left Facet Right Facet Extension 32.69 21.96 Flexion 0.00 0.00 Left Bending 28.90 0.00 Right Bending 0.00 21.02 Left Rotation 0.00 12.67 Right Rotation 14.64 0.00
55 Table 4.3: Facet loads under static loading (C4-C7) C4-C5 Left Facet Right Facet Extension 33.70 26.56 Flexion 0.00 0.00 Left Bending 38.80 0.00 Right Bending 0.00 32.75 Left Rotation 0.00 13.69 Right Rotation 11.36 0.00 C5-C6 Left Facet Right Facet Extension 47.70 39.90 Flexion 0.00 0.00 Left Bending 39.44 0.00 Right Bending 0.00 40.02 Left Rotation 0.00 27.68 Right Rotation 28.34 0.00 C6-C7 Left Facet Right Facet Extension 17.72 39.60 Flexion 0.00 0.00 Left Bending 24.42 0.00 Right Bending 0.00 28.93 Left Rotation 0.00 15.09 Right Rotation 4.96 0.00
The next step was to validate the model under dynamic loading conditions. The experimental data was derived from Stemper Et. Al.[91] The relative motion of the cervical vertebrae in the FE model was compared to that of the experiment. It can be seen from the graphs that the motion falls mostly within the range. The model does start hit peak extension in all of the vertebral relative motions before the cadaver experiments.
This could be due attributed to the change in posture and material properties. In both the
FE and experiment, the specimen had an initial S-curvature, where there was flexion in the upper and extension in the lower cervical spine. The S-curvature lasted approximately
125 ms in the experiment while it lasted approximately 115 ms in the FE model.[91]
56
Figure 4-3: C2-C3 relative motion under dynamic loading[91]
Figure 4-4: C3-C4 relative motion under dynamic loading[91]
57
Figure 4-5: C4-C5 relative motion under dynamic loading[91]
Figure 4-6: C5-C6 relative motion under dynamic loading[91]
58
Figure 4-7: C6-C7 relative motion under dynamic loading[91]
Figure 4-8: Head-T1 relative motion under dynamic loading[91]
4.3 Test Results
Once the model was determined to be validated under static and dynamic loading, it was used to test the effect of head restraint material properties on whiplash injury. The results used to compare the different cases were intradiscal pressure, facet stress, facet capsule strain, and anterior ligament strains. Figure 4-9 shows the impact with the
59 headrest occurring. The upper cervical spine initially goes in minimal flexion which 0 ms creates a slight “S” shape curvature. The skull then goes into extension and eventually
makes an impact with the headrest.
Figure 4-9: Cervical spine motion during simulated impact
The intra-discal pressures were taken at the centroid of the nucleus elements. The max
stress reading was taken and recorded at each step. An example of the stress distribution
across the nucleus nodes can be seen in figure 4-10. The intra-discal pressure reading was
taken to determine the chance of the disc trauma during rear end whiplash. Maximum
facet stress readings were also taken to compare the risk for facet pinching. Facet strains
and anterior cervical spine ligament strains were also recorded and compared between the
various cases.
60
Figure 4-10: Pressure distribution across C6-C7 nucleus
Four foam material properties were used to study the effect they would have on the chance of injury. All of these properties varied in stiffness values. Results showed as the stiffness of the foam increased, the range of motion increased. The test without a headrest resulted in the highest intra-discal pressures at all levels. The C4-C5 disc had the highest stresses in all of the cases in the lower cervical spine. The C4-C5 disc had an intra-discal pressure of 0.6 MPa. The soft foam also had a C4-C5 intra-discal pressure close to 0.6 MPa. The pressure decreased in all of the discs, as the stiffness of the foam material increased. The hard foam had a C4-C5 intra-discal pressure of approximately 0.5
MPa, which is 0.1 MPa less than the model without a headrest.
Figure 4-12 shows the peak facet stresses taken at the lower cervical spine. As expected, the model without the headrest showed consistently greater peak facet stresses.
The C6-C7 facets showed the greatest peak facet stresses. The model without a headrest showed a peak facet stress of approximately 115 MPa at the C6-C7 level. The soft foam test showed a peak facet stress of approximately 110 MPa at C6-C7. The medium foam 1 had a peak facet stress of approximately 105 MPa and the medium foam 2 had a peak
61 facet stress of approximately 78 MPa at C6-C7. Due to the restriction in motion, the hard foam had the lowest peak facet stress of approximately 50 MPa at C6-C7.
Figure 4-13 shows the peak facet capsule strain in the lower cervical spine when compared to varying headrest material properties. C4-C5 facet capsule experienced the greatest amount of peak strain in all of the cases, followed by C5-C6 and then C6-C7.
The model without a headrest had a peak facet capsule strain of approximately 0.158 mm/mm at C4-C5. The peak facet stress decreases as the stiffness of the headrest increases. The hard foam showed the least amount of facet capsule strain of approximately 0.08 mm/mm at C4-C5.
Figure 4-14 shows the peak anterior ligament strains in the cervical spine when compared with varying headrest material properties. The ALL at C6-C7 and ALAR ligaments showed the greatest peak ligament strains. The ALL is a single ligament but the goal was to try to isolate the different levels and check the areas of interest. The ALL at C6-C7 and ALAR ligaments had the greatest peak ligament strains in the model without a headrest. The ligament strains decrease as the stiffness of the headrest increases. The model without a headrest showed a peak ligament strain of approximately
0.21 mm/mm at ALL 67 and a peak ligament strain of approximately 0.205 mm/mm at
ALAR. The hard foam model showed a peak ligament strain of approximately 0.11 mm/mm at ALL 67 and a peak ligament strain of approximately 0.14 mm/mm at ALAR.
62
Figure 4-11: Peak intra-discal pressure for varying foam properties
Figure 4-12: Peak facet stress for varying foam properties
63
Figure 4-13: Facet capsule strain for varying foam properties
Figure 4-14: Anterior ligament strain for varying foam properties
The next test performed part of this experiment was to study the effect of headrest distance on the chance of whiplash injury. The same parameters were used to study and analyze. Figure 4-15 shows the intra-discal pressures in the discs of the lower cervical spine with varying magnitudes of headrest distance. Again, the C4-C5 nucleus showed
64 the highest intra-discal pressure in the lower cervical spine. The model without a headrest showed the greatest intra-discal pressure of approximately 0.6 MPa. The intra-discal pressure decreased as the initial headrest position moved closer to the skull. The intra- discal pressure where the initial headrest distance was 0 mm was approximately 0.45
MPa at nucleus 45.
Figure 4-16 shows peak facet stress in the lower cervical spine facets with varying initial headrest distance away from the skull. Facets at C6-C7 showed the highest peak facet stresses in all of the cases. The model without a headrest had the highest peak facet stress of approximately 115 MPa at C6-C7. The facet stress decreases as the initial headrest distance away from the skull decreases. The peak facet stress at C6-C7 was approximately 48 MPa when the initial headrest distance was 0 mm.
Figure 4-17 shows peak facet capsular ligament strain in the lower cervical spine with varying initial headrest distance. The capsules at C4-C5 showed the greatest peak capsular strains. The model without the headrest had the greatest strain of approximately
0.158 mm/mm at C4-C5. The strain decreases as the initial headrest distance decreases.
The model with a 0 mm initial headrest distance showed the lowest amount of capsular strain of approximately 0.06 mm/mm at C4-C5.
Figure 4-18 shows peak ligament strain at the ALL and ALAR ligaments with varying initial headrest distance. ALL at C6-C7 had the highest peak ligament strain between the entire ALL. The model without the headrest had the highest peak ligament strain in all of the cases. The ligament strain of the ALL and ALAR ligaments decrease as the initial headrest distance decreases. The model without a headrest showed a peak ligament strain of approximately 0.21 mm/mm at ALL 67 and a peak ligament strain of
65 approximately 0.205 mm/mm at ALAR. The model with a 0 mm initial headrest distance had a peak ALL 67 ligament strain of approximately 0.11 mm/mm and a peak ALAR ligament strain of approximately 0.13 mm/mm.
Figure 4-15: Peak intra-discal pressure for varying initial backset
66
Figure 4-16: Peak facet stress for varying initial backset
Figure 4-17: Facet capsule strain for varying initial backset
67
Figure 4-18: Anterior ligament strain for varying initial backset
The third part of this experiment was to study the effect of varying acceleration magnitudes on the chance of whiplash injury. The discs in the lower cervical spine showed the same trend. The IDP decreases as the peak acceleration magnitude decreases.
The IDP at C6-C7 was approximately 0.84 MPa for 15 G while it was approximately
0.43 MPa for 7.5 G. The IDP at C4-C5 was approximately 0.76 MPa for 15 G while it was approximately 0.55 G for 7.5 G. The IDP at C5-C6 was approximately 0.69 MPa for
15 G while it was approximately 0.51 MPa for 7.5 G.
Figure 4-20 shows the peak facet stress with varying peak acceleration magnitudes. The C6-C7 facet consistently had the greatest peak facet stress in the lower cervical spine for all of the cases. The peak facet stresses in all of the levels decreased as the peak acceleration magnitudes decreased. The C6-C7 peak facet stress was approximately 195 MPa for 15 G while it was approximately 100 MPa for 7.5 G. The C5-
68 C6 peak facet stress was approximately 95 MPa for 15 G while it was approximately 65
MPa for 7.5 G. The C4-C5 peak facet stress was approximately 55 MPa for 15 G while it was approximately 20 MPa for 7.5 G.
Figure 4-21 shows the peak facet capsular ligament strain with varying peak acceleration magnitudes. The face capsular ligament strain at all levels decreased as the peak acceleration magnitudes decreased. The C6-C7 peak facet capsule strain was approximately 0.23 mm/mm for 15 G while it was approximately 0.03 mm/mm for 7.5 G.
The C5-C6 peak facet capsule strain was approximately 0.32 mm/mm for 15 G while it was approximately 0.04 mm/mm for 7.5 G. The C4-C5 peak facet capsule strain was approximately 0.2 mm/mm for 15 G while it was approximately 0.11 mm/mm for 7.5 G.
Figure 4-22 shows the peak ligament strains for the ALL and ALAR ligaments with varying peak acceleration magnitudes. The ALL and ALAR strains decreased as the peak acceleration magnitudes decreased. The ALL 67 peak ligament strain was approximately 0.32 mm/mm for 15 G while it was approximately 0.18 mm/mm for 7.5 G.
The ALAR peak ligament strain was approximately 0.28 mm/mm for 15 G while it was approximately 0.18 mm/mm for 7.5 G.
69
Figure 4-19: Peak intra-discal pressure for varying T1 acceleration magnitudes
Figure 4-20: Peak facet stress for varying T1 acceleration magnitudes
70
Figure 4-21: Facet capsule strain for varying T1 acceleration magnitudes
Figure 4-22: Anterior ligament strain for varying T1 acceleration magnitudes
71 Chapter 5
Discussion
5.1 Overview of Results
The first part of the finite element study was to test the effect headrest material properties on whiplash injury. To study this, four different polyurethane properties were used to simulate the headrest. Acceleration profiles used were recorded during crash test dummy experiments performed at the University of Toledo. Results showed the intra- discal pressure, peak facet stress, ALL ligament strain, ALAR ligament strain, and facet capsule strain decreased at all levels with increasing stiffness of the headrest material.
One can concluded that with increased head restraint stiffness, the risk of cervical spine injury decreases. This can directly be attributed due to the decrease in extension motion with a stiffer headrest material. Figure 5-1 shows the peak head stress during the time of impact. The peak head stress during impact can directly be correlated with the risk of head trauma. Although a stiffer material may prevent cervical spine injury, the chance of a head or concussion injury increases. Although, head injury is not commonly seen in whiplash patients, these results are a way of justification to avoid using a wooden or metal block as a headrest.
72
Figure 5-1: Peak head stress during impact for varying foam material properties
The next part of this finite element study was to test the effect of initial headrest distance on whiplash injury. Results showed the intra-discal pressure, peak facet stress,
ALL ligament strain, ALAR ligament strain, and facet capsule strain decreased at all levels with decreasing initial backset distance. The results showed having the initial headrest distance smaller, decreased the risk for whiplash injury. Figure 5-2 shows the peak head stress during impact for different initial headrest distances. As expected, the chance of head trauma also increases as the initial headrest distance increases. The last part of this thesis was to study the effect of acceleration magnitudes on whiplash injury in the finite element model. Results showed the intra-discal pressure, peak facet stress, ALL ligament strain, ALAR ligament strain, and facet capsule strain increased at all levels with increasing acceleration magnitudes. Results showed as the acceleration magnitude increased, the risk of whiplash injury also increased. Figure 5-3 shows the peak head stress during impact for different acceleration magnitudes. As expected, the chance of
73 head trauma increases as the acceleration magnitude increases due to the severity of the impact.
Figure 5-2: Peak head stress during impact for varying initial backset distance
Figure 5-3: Peak head stress during impact for varying T1 acceleration magnitudes
It is reported that the PMHS sub-catastrophic failure strain of facet capsular ligaments in isolated models was from 35 ± 21% to 65 ± 74%.[98],[99] Other PMHS and
74 cadaveric models have reported peak facet capsular ligament strains from 29 to
40%.[100],[101],[102],[103]. Fice Et. Al. reported similar facet capsular ligament strains for his finite element model at 7 G, 10 G, 12, G, and 14 G.[94] His model did not reach sub- traumatic failure region in the lower cervical spine for accelerations up to 16 G. The facet capsular ligament strains did not reach failure strains for this finite element model but they would reach if the acceleration magnitudes were higher. It can also be seen that the
IDP was greater in certain models than what is reported in various cadaveric studies.
Stemper Et. Al. performed a study using varying backset distances. He reported that potentially injury would occur prior to headrest contact if the initial backset was
60mm posterior to the head.[55] The study did not report any stress or strain values and reported the chance for injury in terms of segmental motion. It is a limitation of this thesis that different acceleration profiles were not used for T1 and the headrest.
Therefore, no conclusion can be made about the actual backset distance needed to reduce the risk of injury. It can only be reported that as the initial backset distance increases, the risk of cervical soft tissue injury increases.
Results of this study shows there is a big reduction in the IDP, facet stress, and ligament strains from the medium foam 1 to medium foam 2. This can be attributed to the big different in the stiffness curves of the foam. The difference between no headrest, soft foam and medium foam 1 is minimal because of the difference in the stiffness curves is minimal. Results also show a relatively linear relationship between the varying headrest distance to peak IDP, facet stress, and ligament strains. There is also a linear relationship between the peak acceleration amplitude to the peak IDP, facet stress, and ligament strains.
75 It can also be seen from the results that peak IDP values, facet stresses, and ligament strains were not directly correlated with the range of motion. The dynamic graphs are shown in Appendix A. Although C5-C6 consistently showed the greatest motion in extension, C4-C5 consistently showed the greatest IDP and greatest facet capsular strain, while C6-C7 showed the greatest facet stress and ALL strain. But it is also seen that as the range of motion of the different segments increased, the IDP, facet stress and ligament strains values all increased. C2-C3 capsular facet also showed the greatest ligament strain, which was seen at approximately the same time as peak flexion.
Since the focus of this study was on lower cervical spine injury, this was not used as a comparison variable.
5.2 Limitations and Future Work
The limitations of this model include the lack of muscle behavior in the model.
Since muscles are said not to be activated before 200 ms, they were not modeled for this current study. Another limitation was that the headrest and T1 were given the same exact acceleration profiles. The model was also limited because there was not a T1 angle profile given to the model. The model also did not have a brain model incorporated and the weight of the brain was simulated by distributing it across the skull. Future would include the incorporation of muscles in the model. Another future study should be the addition and validation of a brain model to study concussions and head trauma.
5.3 Conclusions
The author successfully validated the cervical spine model under static and dynamic loading. The validated model was then used to test the effect of whiplash injury.
76 Results showed a stiffer material would decrease the risk for cervical injury but would increase the risk for head injury. Results also showed the risk of cervical injury decreased with a lesser T1 acceleration magnitude and with a smaller initial headrest distance. The author would not recommend the usage of any low stiffness polyurethane foams for car head restraints.
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86 Appendix A
Dynamic Result Curves
Figure A-1: Range of motion for model with no headrest, at 7.5 G acceleration
87
Figure A-2: Facet stress for model with no headrest, at 7.5 G acceleration
Figure A-3: Intra-discal pressure for model with no headrest, at 7.5 G acceleration
88
Figure A-4: Anterior ligament strain for model with no headrest, at 7.5 G acceleration
Figure A-5: Facet capsule strain for model with no headrest, at 7.5 G acceleration
89
Figure A-6: Range of motion for model with soft foam headrest, 10 mm backset and 7.5 G acceleration
Figure A-7: Facet stress for model with soft foam headrest, 10 mm backset and 7.5 G acceleration
90
Figure A-8: Intra-discal pressure for model with soft foam headrest, 10 mm backset and 7.5 G acceleration
Figure A-9: Anterior ligament strain for model with soft foam headrest, 10 mm backset and 7.5 G acceleration
91
Figure A-10: Facet capsule strain for model with soft foam headrest, 10 mm backset and 7.5 G acceleration
Figure A-11: Range of motion for model with medium foam 1 headrest, 10 mm backset and 7.5 G acceleration
92
Figure A-12: Intra-discal pressure for model with medium foam 1 headrest, 10 mm backset and 7.5 G acceleration
Figure A-13: Facet stress for model with medium foam 1 headrest, 10 mm backset and 7.5 G acceleration
93
Figure A-14: Anterior ligament strain for model with medium foam 1 headrest, 10 mm backset and 7.5 G acceleration
Figure A-15: Facet capsule strain for model with medium foam 1 headrest, 10 mm backset and 7.5 G acceleration
94
Figure A-16: Range of motion for model with medium foam 2 headrest, 10 mm backset and 7.5 G acceleration
Figure A-17: Intra-discal pressure for model with medium foam 2 headrest, 10 mm backset and 7.5 G acceleration
95
Figure A-18: Facet stress for model with medium foam 2 headrest, 10 mm backset and 7.5 G acceleration
Figure A-19: Anterior ligament strain for model with medium foam 2 headrest, 10 mm backset and 7.5 G acceleration
96
Figure A-20: Facet capsule strain for model with medium foam 2 headrest, 10 mm backset and 7.5 G acceleration
Figure A-21: Range of motion for model with hard foam headrest, 10 mm backset and 7.5 G acceleration
97
Figure A-22: Intra-discal pressure for model with hard foam headrest, 10 mm backset and 7.5 G acceleration
Figure A-23: Facet stress for model with hard foam headrest, 10 mm backset and 7.5 G acceleration
98
Figure A-24: Anterior ligament strain for model with hard foam headrest, 10 mm backset and 7.5 G acceleration
Figure A-25: Facet capsule strain for model with hard foam headrest, 10 mm backset and 7.5 G acceleration
99
Figure A-26: Range of motion for model with medium foam 2 headrest, 0 mm backset and 7.5 G acceleration
Figure A-27: Intra-discal pressure for model with medium foam 2 headrest, 0 mm backset and 7.5 G acceleration
100
Figure A-28: Facet stress for model with medium foam 2 headrest, 0 mm backset and 7.5 G acceleration
Figure A-29: Anterior ligament strain for model with medium foam 2 headrest, 0 mm backset and 7.5 G acceleration
101
Figure A-30: Facet capsule strain for model with medium foam 2 headrest, 0 mm backset and 7.5 G acceleration
Figure A-31: Range of motion for model with medium foam 2 headrest, 5 mm backset and 7.5 G acceleration
102
Figure A-32: Intra-discal pressure for model with medium foam 2 headrest, 5 mm backset and 7.5 G acceleration
Figure A-33: Facet stress for model with medium foam 2 headrest, 5 mm backset and 7.5 G acceleration
103
Figure A-34: Anterior ligament strain for model with medium foam 2 headrest, 5 mm backset and 7.5 G acceleration
Figure A-35: Facet capsule strain for model with medium foam 2 headrest, 5 mm backset and 7.5 G acceleration
104
Figure A-36: Range of motion for model with medium foam 2 headrest, 20 mm backset and 7.5 G acceleration
Figure A-37: Intra-discal pressure for model with medium foam 2 headrest, 20 mm backset and 7.5 G acceleration
105
Figure A-38: Facet stress for model with medium foam 2 headrest, 20 mm backset and 7.5 G acceleration
Figure A-39: Anterior ligament strain for model with medium foam 2 headrest, 20 mm backset and 7.5 G acceleration
106
Figure A-40: Facet capsule strain for model with medium foam 2 headrest, 20 mm backset and 7.5 G acceleration
Figure A-41: Range of motion for model with medium foam 2 headrest, 10 mm backset and 9.8 G acceleration
107
Figure A-42: Intra-discal pressure for model with medium foam 2 headrest, 10 mm backset and 9.8 G acceleration
Figure A-43: Facet stress for model with medium foam 2 headrest, 10 mm backset and 9.8 G acceleration
108
Figure A-44: Anterior ligament strain for model with medium foam 2 headrest, 10 mm backset and 9.8 G acceleration
Figure A-45: Facet capsule strain for model with medium foam 2 headrest, 10 mm backset and 9.8 G acceleration
109
Figure A-46: Range of motion for model with medium foam 2 headrest, 10 mm backset and 13.8 G acceleration
Figure A-47: Intra-discal pressure for model with medium foam 2 headrest, 10 mm backset and 13.8 G acceleration
110
Figure A-48: Facet stress for model with medium foam 2 headrest, 10 mm backset and 13.8 G acceleration
Figure A-49: Anterior ligament strain for model with medium foam 2 headrest, 10 mm backset and 13.8 G acceleration
111
Figure A-50: Facet capsule strain for model with medium foam 2 headrest, 10 mm backset and 13.8 G acceleration
Figure A-51: Range of motion for model with medium foam 2 headrest, 10 mm backset and 15 G acceleration
112
Figure A-52: Intra-discal pressure for model with medium foam 2 headrest, 10 mm backset and 15 G acceleration
Figure A-53: Facet stress for model with medium foam 2 headrest, 10 mm backset and 15 G acceleration
113
Figure A-54: Anterior ligament strain for model with medium foam 2 headrest, 10 mm backset and 15 G acceleration
Figure A-55: Facet capsule strain for model with medium foam 2 headrest, 10 mm backset and 15 G acceleration
114
Figure A-56: Peak stress on head with varying foam properties, at 10 mm backset and 7.5 G acceleration
Figure A-57: Peak stress on head with varying initial backset with medium foam 2 and 7.5 G acceleration
115
Figure A-58: Peak stress on head with varying acceleration amplitudes with medium foam 2 and 10 mm backset
116