How To Understand The Physiology Of The Head And Cervical Spine

Whiplash injuries in-vitro studies Narayan Yoganandan, PhD, Brian D. Stemper, PhD, Frank A. Pintar, PhD, and Raj D. Rao^, MD Department of Neurosurgery ^Department of Orthopedic Surgery Medical College of Wisconsin Milwaukee, WI Correspondence: Narayan Yoganandan, Ph.D. Professor & Chair, Biomed Eng Department of Neurosurgery Medical College of Wisconsin 9200 West Wisconsin Avenue Milwaukee, WI 53226 Tel: 414-384-3453 Fax: 414-384-3493 e-mail: yoga@mcw.edu

Whiplash injuries In-vitro studies Page 1 I. INTRODUCTION II. INTACT PMHS STUDIES III. ISOALTED PMHS HEAD-NECK COMPLEX STUDIES A. SPECIMEN PREPARATION, MOUNTING, INSTRUMENTATION, AND LOADING B. DATA ANALYSIS C. OVERALL HEAD-NECK KINEMATICS D. SEGMENTAL KINEMATICS E. LOCAL COMPONENT KINEMATICS IV. ANATOMIC STUDIES OF CERVICAL FACET JOINT V. DIFFERENCES IN KINEMATICS BETWEEN PHYSIOLOGIC AND POSTEROANTERIOR ACCELERATION LOADINGS VI. APPLICATION OF IN VITRO STUDIES TO COMPUTATIONAL MODELING A. DEVELOPMENT OF KINEMATIC CORRIDORS BASED ON IN VITRO STUDIES B. VALIDATION OF A COMPUTATIONAL MODEL C. EFFECTS OF INITIAL SPINAL POSTURE D. EFFECTS OF THORACIC RAMPING E. ANTERIOR LONGITUDINAL LIGAMENT KINEMATICS F. REFLEX MUSCLE CONTRACTION IN UNAWARE OCCUPANTS G. MUSCLE CONTRACTION IN AWARE OCCUPANTS VII. MECHANISMS OF INJURY VIII. SUMMARY

Whiplash injuries In-vitro studies Page 2 I. INTRODUCTION To understand the biomechanics of trauma, in vivo, in vitro, physical, and computational models are used. In vivo models include experimental animals and human volunteers. In vitro models primarily consist of post mortem human subject (PMHS) investigations. Physical models include anthropomorphic test devices, widely known as crash dummies. Computational models include occupant kinematics and stress analysis-based finite element analysis research. Human volunteer studies allow monitoring of physiological responses to the external insult at subinjury thresholds. In contrast, in vitro studies are suitable to understand the biomechanics from subinjury to injury producing load or acceleration levels although physiological evaluations are not incorporated into the experimental design. In vitro studies provide data for the design, development, and validation of biofidelic physical models. In addition, results from carefully designed in vitro experiments serve as validation and verification data for computational models so that parametric simulations can be conducted to better understand the intrinsic biomechanics of structural components potentially involved in injury and/or pain producing processes. Consequently, with a primary focus on in vitro biomechanical studies, this chapter presents kinetic and anatomical responses of the head and cervical spine, with a focus on applications in clinical, epidemiological, and safety-engineering aspects of whiplash injuries. Computational outcomes are also presented. Because it is widely acknowledged that whiplash injury is predominantly rear impact motor vehicle-related 131, the chapter focuses on this loading mode, i.e., posteroanterior acceleration loading. II. INTACT PMHS STUDIES During 1967-1971, studies were conducted using two embalmed intact PMHS 69,70. One PMHS was previously used in windshield, shoulder harness, and steering wheel evaluation programs, and the other was used to determine the force-deflection characteristics of the chest 68. Because x-rays from previous tests revealed head/facial fractures with no evidence of neck damage, the two PMHS were determined to be suitable to investigate rear impact biomechanics. Using a horizontal accelerator, tests were conducted at 14.4 and 24 km/h change in velocity (ΔV, a commonly used measure of impact severity in crashworthiness research). The simulation sequence consisted of two series, with and without head support. The first test was with a rigid seatback at 14.4 km/h, and remaining tests were conducted at 24 km/h. The degree of seatback

Whiplash injuries In-vitro studies Page 3 rigidity was incrementally increased, and the seatback was fully rigid for the final run. The authors stated: x-rays were taken after any simulation in which damage to the cervical spine was suspected to have occurred. X-rays showed minor damage between C3 and C4 vertebrae in one and no damage in the other PMHS. Acknowledging that soft tissue trauma identification is only inferential from x-rays, these studies did not identify and document the specific component(s) sustaining trauma. It should however be noted that imaging technologies such as cryomicrotomy were not available at the time of this research. These studies have formed a basis for specifying human tolerance to injury in rear impact 32. In 1972, sled tests were conducted using 21 unembalmed PMHS torsos 16. The specimens were isolated at T10 and mounted on a flat rigid plate. In all but one, the head was transected at the base of the skull and a spherical wooden surrogate was inserted. The preparations were divided into no head restraint and simulated head restraint groups. Tests were conducted from 19.1 to 25.6 km/h ΔV with head restraint and between 25.0 and 25.6 km/h ΔV without head restraint. In all tests, a steel backrest with foam upholstery supported the torso, and the backrest was level with the shoulder. In tests with head restraint, increased translation of the thorax, initiating relative motion between the head and thorax resulted in slight forward neck flexion. No injuries occurred in this group. In tests without head restraint, torso acceleration lasting 30 to 40 msec led to rotation of the head while the neck was in extension. Timing of peak head rotation and deceleration were coincident. Autopsy procedures identified disc injuries followed by ruptures of anterior and posterior longitudinal ligaments, joint capsules, ligamentum flavum, and posterior vertebral body and spinous process fractures. In a later study, 49 sled tests were conducted from 6 to 15 km/h ΔV using two female and four male PMHS 35. Initial head angles ranged from -45 to +45 0. The horizontal distance between the head and head restraint (backset) ranged from zero to 16 cm. A high-speed video camera was used to obtain sagittal plane motions of landmarks attached to two cervical vertebral bodies. Rearward head rotation initiated 60 to 100 msec after impact. After 100 to 160 msec, the head began to rotate forward, after the shoulders had translated forward. Rearward head rotation was minimized with zero backset, and was maximized with 16 cm backset; other studies using computational modeling and dummy tests have shown similar results 30,102. Head to mid-cervical

Whiplash injuries In-vitro studies Page 4 spine motion initiating between 50 and 80 msec following impact resulted in upper neck flexion. Peak upper neck flexion up to 45 0 occurred between 100 and 130 msec. Although no injury information was provided, the importance of avoiding extreme relative motion between the head, neck, and torso to minimize neck trauma was emphasized. Several studies were reported in 2000. One study subjected four PMHS to 19 rear impacts to investigate effects of ΔV and head restraint on head and neck kinetics at 11 and 16 km/h ΔV 11. The study identified ramping up and rearward T1 rotation. According to the authors, the first phase of kinematics included lower and upper spine (C5-T1, OC-C2) initial flexion, and middle spine (C2-C3) extension with the head remaining horizontal/stationary. This phase is termed as the retraction phase. The second phase was characterized by extension of the entire spine (C2- T1) as the head began to rotate rearward. The spinal column continued to extend during the third phase. In head restraint tests, the head contacted the restraint toward the end of the first phase. Neck kinetics increased with ΔV. Injuries were not reported in the original article. Another study subjected six PMHS to 26 rear impacts, measured intervertebral kinematics, and reported similar upper (C1-C3) and lower (C5-C6) spine curvatures during initial stages 27,61, although, a later study by the same group of authors reported the absence of S-curvature 113. Peak facet capsule strains occurred prior to maximum head extension and prior to head restraint contact 27,61. Injuries were discovered during autopsy in four specimens at the C4-C5 to C7-T1 facet capsules, discs, muscles, and thyroid cartilage. A fracture to the lower part of C6 occurred in one specimen. A significant majority of injuries occur in low-speed motor vehicle crashes and the most common impacting vector is from the rear, inducing single cycle postero-anterior acceleration to the occupant through the seat 4,10,12,14,19,47,65,75,91,93,111. Injury identifications are not reliable from models using repeated applications of external loading to the same specimen. This is because repeated acceleration input alters the initial preloaded status of the specimen, and previously applied loading may have induced unknown injury or altered spinal strength as components of the complex cervical spine are nonlinear and heterogeneous. Above all, repeated loading does not occur in real world motor vehicle-induced trauma. By definition, posteroanterior

Whiplash injuries In-vitro studies Page 5 acceleration-induced injury is soft tissue related; fractures are uncommon. To document injuries from this type of loading, in vitro studies cannot rely on traditional imaging techniques such as x-ray and computed tomography (CT) as references are only inferential. Because of the in vitro nature of the experimental model coupled with tissue types involved in injury (e.g., facet joint distraction without lateral mass fracture), resolution and other factors, magnetic resonance imaging is not the most appropriate modality to identify trauma. Autopsy is also not the most efficacious tool because techniques for soft tissue injury identifications are imprecise. These reasons may have precluded earlier researchers from identifying soft tissue trauma to specific components of the head-neck complex. In 2000, using unembalmed intact PMHS and cryomicrotomy, soft tissue injuries were identified and documented due to single application of rear impact posteroanterior acceleration 128,142. Using a whole-body deceleration sled adopted in frontal and side impact PMHS and dummy research 129,130,133, male and female PMHS were screened for the absence of spinal trauma, seated in the normal driving posture with the Frankfort plane horizontal, and subjected to 15 or 25 km/h ΔV rear impacts. Kinematic data were recorded to determine head-neck forces and moments in the three-dimensional (3-D) anatomical plane, posttest x-rays and CT images were obtained, and cervical spines were subjected to cryomicrotomy. The rigid seat had a cushion and typical automotive seat geometry. A worst-case scenario was simulated by allowing unconstrained motion of the head-neck. Structural abnormalities included stretch/tear of the ligamentum flavum, annulus disruption, anterior longitudinal ligament rupture, and facet joint compromise with capsular ligament tear (Figure 1). These results indicated that single application of posteroanterior acceleration loading can induce soft tissue-related alterations not identifiable using routine clinical imaging modalities. Identified and documented abnormalities supported clinical postulates regarding the location of soft tissue injuries and offered an explanation for headache and neck pain, the two most common complaints in whiplash patients 12,14,47,48,55,66,79,80,91,111. To the best knowledge of the authors, this continues to be the only study documenting trauma using realistic real-world posteroanterior acceleration loading environment, recognizing that variables such as seat and head restraint, occupant anthropometry, demographics, and positioning

Whiplash injuries In-vitro studies Page 6 may play a role in injury outcome. Identification and documentation of kinetics (kinematics) and injuries have formed a basis along with clinical findings regarding the potential involvement of spinal components to pursue focused investigations using other in vitro models. Because of the 3-D nature of the head-neck complex, and because injuries can occur due to excessive motions, other in vitro models have been used to determine segmental and local component kinematics. III. ISOLATED HEAD-NECK COMPLEX STUDIES Using T1 acceleration output from intact PMHS sled experiments as a basis for driving the PMHS intact head-neck complex model 20,101,103-105,132, the overall head-neck, level-by-level intersegmental, and local component motions; forces and bending moments; and acceleration metrics were determined to offer additional insights into the biomechanics of load transfer and explain mechanisms of injury. Results from these studies have also served as validation and verification data for computational models to better understand the intrinsic biomechanics of structural components due to posteroanterior acceleration loading. A. SPECIMEN PREPARATION, MOUNTING, INSTRUMENTATION, AND LOADING Isolated PMHS intact head-neck complexes from both genders, and free from Hepatitis A, B, and C, and HIV, were screened for musculoskeletal and head trauma. Specimen preparation included isolation at T2 and removal of the esophagus and trachea, leaving intact the head and ligamentous spine with skin and musculature. Retroreflective targets were inserted into the lateral region of each lateral mass, anterior regions of each vertebral body (C2-C7), and mastoid process. Smaller targets were placed at the four sagittal plane corners of each facet joint (Figure 2). Posteroanterior acceleration loading was delivered using a calibrated mini-sled pendulum 141. The specimen, fixed at the thoracic end with the head unconstrained, was mounted on the minisled and oriented similar to PMHS tests by maintaining the Frankfort plane horizontal and orienting occipital condyles superior to the T1 body (Figure 2). T1 was anteriorly oriented 25 o. A six-axis load cell was placed at the distal end to determine 3-D forces and moments. Accelerometers on the mini-sled and head were used to determine input ΔV and head kinematics. Load cell and accelerometer data were acquired at 12,500 Hz and filtered and processed according to the Society of Automotive Engineers SAE J211 specifications. Two tests were initially conducted at 2.1 km/h ΔV to determine baseline kinematics. The acceleration matrix

Whiplash injuries In-vitro studies Page 7 (Figure 3) was such that each higher ΔV test (4.6, 6.6, 9.3, and 12.4 km/h) was preceded by the lowest ΔV test and radiography. Testing terminated upon the detection of trauma, defined as joint instability on gross inspection or palpation by clinical personnel, x-rays, or drastic/unexpected biomechanical signal changes. Sequential images from 1,000 frames/sec high-speed digital video cameras were used to statistically analyze kinematics and correlate with accelerations and 3-D forces and moments. B. DATA ANALYSIS High-speed digital images were temporally analyzed to determine the overall spinal responses. Kinematics were processed using the principles of continuous motion analysis in the sagittal plane 134,140. Intervertebral rotations (Figure 2) at various segmental levels of the head-neck complex were determined using retroreflective target data. The overall angular motion was defined as the motion of the head with respect to the base. Segmental rotations of one spinal vertebra with respect to its adjacent vertebra were computed. Facet joint motions were defined as anterior and posterior translations along (shear, tangential motion) and perpendicular to(distraction or compression, normal motion) the joint line in ventral and dorsal joint regions (Figure 4). Analysis of temporal force and moment metrics revealed patterns of load transmission in the head-neck complex. C. OVERALL HEAD-NECK KINEMATICS Because the cervical spine has a natural lordotic curvature and supports the eccentrically located head at its upper regions, it is important to describe curvatures of the column in relation to the cranium. Regional or overall changes in the signature(s) of the head-neck complex from the initial lordotic or unloaded curvature can explain the mechanism of load transfer and perhaps offer explanations to the mechanism of injury. For example, flexion predisposes posterior headneck structures to distraction-mediated mechanism of load transfer, and dorsal structure stretch may explain injury or pain attributed to the local region. Overall head-neck kinematics were analyzed using the above described intact head-neck complex tests. During the initial stages of acceleration input, a transient decoupling of the head occurred with respect to the neck exhibiting a lag of the cranium. During the loading phase, the

Whiplash injuries In-vitro studies Page 8 head and upper cervical spine responded with local flexion concomitant with head lag while the lower column was under local extension. This established a reverse or S-curvature of the headneck complex (Figure 5) 104. Although transient, the biphasic S-curvature is nonphysiologic 20. Dynamic alterations of the occiput to C2 complex may impart potentially adverse forces to related neural structures, with subsequent development of a neuropathic pain process. Excessive flexion of the posterior upper cervical regions can be correlated to headaches 20. With continuing application of the acceleration, inertia of the head caught up with intervertebral deformations. The entire head-neck complex was subsequently under a mono-phase extension mode with C- curvature 135,136,138. Formation of the transient nonphysiologic curvature represented translation of the neck with respect to the head before the cervical column attained a stable extension signature. The observed S- and C-curvatures and phase differences between the spine and head in the intact head-neck complex model have also been reported in studies using isolated ligamentous columns supporting an artificial head 40,86. The initial transient S- and ultimate C-curvatures of the head-neck complex were independent of gender and ΔV, although the transient curvature shifted inferiorly with increasing ΔV 104. These evaluations indicate that the human head-neck structure behaves similarly without external insult level or gender bias. The former identification is helpful in the design of safety-engineering systems. For example, female and male dummy designs can be based on similar overall kinematic responses. However, the latter identification fails to explain female gender-bias to injury, reported in clinical and epidemiological literature 6,12,14,18,42,43,47,48,51,55,66,79,80,90,91,93,100,110-112,118. D. SEGMENTAL KINEMATICS It is well known that motion and injury are interrelated. Segmental kinematics refer to motion of one vertebra with respect to the adjacent vertebra. Acknowledging that motions have been a hallmark for diagnosing spinal instability for over six decades 56, determinations of segmental motions due to posteroanterior acceleration loading may assist in treatment. Temporal sagittal plane segmental angulations were determined using the above described intact head-neck complex tests. Flexion and extension rotations were computed during maximum S-

Whiplash injuries In-vitro studies Page 9 curvature, defined as the time of peak C2 C3 flexion (Figure 5). The mean time of attainment of peak S-curvature decreased with increasing ΔV. The earliest time of occurrence was 68 ms for 12.4 km/h ΔV, and the latest was 100 ms for 2.1 km/h ΔV (Table 1). The mean time of attainment of maximum S-curvature for the lowest ΔV coincided with the end of loading phase. Segmental angulations increased caudally, and this phenomenon was true at all ΔV and for both genders. At all ΔV, the C2 C3 segment responded with greatest flexion, and the C6 C7 segment responded with greatest extension. Extension and flexion angles for both genders increased with increasing ΔV (Figure 6). While no discernable pattern for the effect of ΔV on angulations was apparent at the C2 C3 segment, the C3 C4 segment transitioned from extension to flexion with increasing ΔV. Female head-neck complexes responded with significantly greater (p<0.05) angulations at C2-C3 and C4 to C7 levels than male specimens (Figure 6) 104. Acknowledging similar anatomical and structural developments in the human cervical spine for both genders, these kinematics-based results partly explain gender bias in rear impact-induced injury. Increased female segmental motions offer a biomechanical explanation for higher incidence of whiplash-related complaints to this population 104. An implication of this finding is that females cannot be viewed as scaled-down males for injury tolerance and crashworthiness evaluations. Thus, while female and male dummies may be based on similar overall kinematic responses, from a segmental perspective, differing designs or use of different evaluation criteria may need to be explored to accommodate the gender bias. E. LOCAL COMPONENT KINEMATICS Local component motions focus on specific soft tissues such as intervertebral discs, facet joints, and ligaments. Tension/compression and shear motions of joint structures can be associated with stretch of the connective soft tissue components. Excessive stretch may result in sub- or catastrophic tissue failure. While the latter may result in acute instability, the former may lead to chronic pathology, an important outcome in whiplash patients 107. Consequently, determinations of local component motions are valuable for a better understanding of the biomechanics and mechanisms of injury. Localized sagittal plane facet joint kinematics were analyzed using the above described intact head-neck complex tests. Lower cervical facet joints demonstrated varying magnitudes of

Whiplash injuries In-vitro studies Page 10 localized compression and shear. While the anterior- and posterior-most regions of the joint responded with similar anteroposterior shear motions along the joint (Figure 7), the posteriormost joint region responded with compression and the anterior-most region responded with distraction (Figure 8). This may account for region-dependent injury mechanisms, i.e., distraction in the anterior and pinching in the posterior regions. The combination of shear/tangential and axial/normal motions, i.e., shear plus distraction mechanism in the ventral and shear plus compression mechanism in the dorsal region results in joint capsule stretch. During the time of S-curvature, facet joints demonstrated statistically different (p<0.05) regiondependent behavior, with compression in the dorsal and distraction in the ventral regions. Shear motion across the joint was not different between regions, although shear magnitudes were significantly (p<0.05) greater than distraction magnitudes. In addition, shear motion was greater from C4-C6 levels in females 103. Nonphysiologic kinematic responses may induce stresses in lower facet joints, resulting in possible compromise sufficient to elicit neuropathic or nociceptive pain. Injury to the ventral region stems from tensile failure of the joint capsule. Injury to the dorsal region stems from pinching of the capsule or synovial fold and contact between subchondral bone of superior and inferior processes. While cartilage surrounding the articular process of the lateral mass is devoid of nerve endings, the subchondral bone underlying the cartilage is innervated. The relatively softer cartilage may expose adjacent subchondral bones in dorsal regions of the lateral mass to impinge, eliciting pain. This analysis offers an explanation to facet joint-induced neck pain, reported in rear impact motor vehicle epidemiology literature 14,47,79,100,118. Because excess spinal motion is biomechanically related to abnormalities, and because lower cervical facet joints sustain greater motion in female specimens, this population is susceptible to trauma. Other studies have reported similar kinematics. Using intact PMHS, maximum facet joint capsule strains occurred prior to maximum head extension and head restraint contact 27. This finding may partly explain the lack of effectiveness of head restraints in reducing rear impactinduced injuries 52,82. Another study using OC-T1 columns supporting an artificial head reinforced the compression/pinching injury theory discussed above by demonstrating that facet

Whiplash injuries In-vitro studies Page 11 joint compression exceeded physiologic limits under 3.5-8.0 g posteroanterior acceleration 87. Using C1-T1 columns fixed at the two ends and subjected to quasistatic loading, another study showed that shear stiffness decreases in the presence of axial compression 127. This finding may imply a role for thoracic ramping and uncoiling of the dorsal spine in the mechanism of load transfer and injury. Quasi-static isolated motion segment tests have shown that pretorque increases facet capsule strains 126. Initially rotated head position before the application of rear impact-induced posteroanterior acceleration loading has been associated with greater injury susceptibility in clinical and epidemiological literature 25,31,91,111. These metrics have been used to explain whiplash injury biomechanics. III. ANATOMIC STUDIES OF CERVICAL FACET JOINT To determine level and gender dependency on facet joint morphology in the cervical spine, an in vitro anatomic study subjected unembalmed intact PMHS cervical spinal columns to cryomicrotomy. Following x-ray and CT, specimens were sectioned in the sagittal plane at 20- to 40 µm. Facet joint width, and cartilage thickness and gap were extracted from sequential sections. Width was defined as the maximum distance along the joint at its midlevel. The cartilage on the superior and inferior facet surfaces was divided into five equal segments, and the thickness was measured at 0, 25, 50, 75, and 100% of the length of the cartilage (Figure 9). The cartilage gap was defined as the distance from the ventral- or dorsal-most region of the joint to the location where the cartilage began to appear. These measurements were obtained on the superior and inferior surfaces for each joint (Table 2). Data were grouped into upper (C1 C2, UCS) and lower (C3 C7, LCS) cervical spine regions based on the interaction outcome from factorial analysis of variance. Interactions were detected between gender and location of the facet joint gap, i.e., dorsal versus ventral and between upper versus lower spinal regions. The gap was significantly lower (p<0.05) in UCS than LCS (Figure 10). In addition, the gap at the ventral and dorsal regions was significantly (p<0.05) lower in UCS than in LCS (Figure 11). The gap in the dorsal region was lower in males than females. Facet cartilage was thickest in the middle region (Figure 12) and UCS (Figure 13). The overall mean thickness was significantly lower (p<0.05) in females than males in UCS (Figure 11). In contrast, for the facet joint width that decreased from the rostral to caudal direction (Figure 14), significant differences (p<0.05) were apparent only between UCS and LCS.

Whiplash injuries In-vitro studies Page 12 Since bilateral facet joints contribute to intervertebral rotations, variations in geometry affect spinal kinematics. Significantly greater facet joint width in the upper than the lower spine may have implications in the mechanics of the human neck due to posteroanterior acceleration loading. Increased facet joint geometry results in increased translations 71. The initial position of the head affects segmental motions of the cervical spinal column; lower cervical joint motions were smaller in human volunteers oriented in an initial chin-in/out position than in the normal position 89. Since the soft cartilage is devoid of nerve endings, its direct role is likely minimal in the pain process. However, the gap, or the lack of cartilage cover at the ends of the joint, accentuates the role of this component on subchondral bone mechanics in trauma. If the facet joint opposition is decreased leading to contact with adjacent processes, due to the pinching mechanism reported from intact human-head neck complex studies, lack of cartilage cover at the ends reduces the cushion normally provided to the bone. This phenomenon occurs at the ends due to gap. Human volunteer and clinical investigations have implicated the facet joint as a source of pain 3,7. As indicated, epidemiological studies have reported that females are more vulnerable to whiplash-associated disorders such as chronic neck pain; complaints before and after changes in the legal system were higher in females than males 14. The finding that the cartilage cover is less extensive in females may predispose the female subchondral bone to more direct forces than the male bone, particularly at lower cervical levels. Human volunteer rear impact studies have reported impingement of facet joints in rear impact 54. Posterior compression of the facet joint exposes the dorsal region of the apophyseal anatomy to additional compressive forces. Because dorsal impingement occurs during lower cervical extension in rear impact-induced posteroanterior acceleration loading, and because the dorsal cartilage gap is larger in females, bone-to-bone contact is more likely in this population. The bony contact between the two adjacent facet surfaces may be a source of abnormality that leads to dysfunction in females. These anatomic findings together with kinematic results from head-neck complex tests have been successful in supporting theories and clinical and epidemiological observations, i.e., headache and neck pain due to single rear impact-induced posteroanterior acceleration loading. Other in

Whiplash injuries In-vitro studies Page 13 vitro studies examining facet joint anatomy have focused on nerve fibers in the facet joint capsule and synovial fold. Supporting the facet joint impingement theory, nociceptors have been identified in synovial folds, and nerve fibers immunoreactive for Substance P have been identified in lumbar facet joints 9,36. Nerve fibers in lumbar synovial folds are primarily involved in vasoregulation and are of minimal importance in nociception due to limited immunoreactivity for substance P 41. However, recent literature has identified nervous tissue immunoreactive for substance P and calcitonin gene-related peptide (CGRP) in cervical facet joint synovial folds and capsules 49,53. This finding suggests that cervical synovial folds may play a role in the initiation and modulation of facet joint-mediated pain. Specifically, neuropeptides may have long-lasting hyperalgesic effects 73, wherein Substance P is involved in pain initiation, while CGRP is involved in the persistence of pain 92. IV. DIFFERENCES IN KINEMATICS BETWEEN PHYSIOLOGIC AND POSTEROANTERIOR ACCELERATION LOADINGS Localized spinal component kinematics during the acceleration phase have offered explanations for facet-joint mediated neck injuries in whiplash patients 20,87,103,137. However, the magnitude of deformation to induce injury is yet to be quantified. A complicating factor for the facet joint injury theory is that segmental angulations due to rear impact-induced posteroanterior acceleration loading do not exceed normal physiological ranges of motion 50,84,104. A study tested the hypothesis that lower cervical facet joints sustain greater motions due to posteroanterior acceleration loading than due to physiologic loading. Unembalmed human cadaver osteoligamentous cervical columns were subjected to physiological loading 101. In general, specimen preparation, data acquisition, and analysis followed similar procedures described for the intact head-neck complex studies. For applying physiologic loading, the inferior end of the specimen was attached to the load frame through a six-axis load cell, and an arch was attached to the superior edge of the top fixative to transform superior inferior loads applied using an electrohydraulic piston into physiological extension plus compression. Because of the eccentric placement of the head with respect to the spinal axis, compression combined with moment always acts on the human neck. Kinematics due to posteroanterior acceleration loading obtained from intact head-neck complex tests, described in section III were compared to physiological kinematics. Shear and distraction motions were computed in posterior and anterior

Whiplash injuries In-vitro studies Page 14 facet joint regions. Results from physiological and posteroanterior acceleration loadings were limited to the loading phase with the latter confined to the period from the initiation of impact to peak S-curvature. Facet joint shear and distraction in the anterior and posterior regions were plotted as a function of segmental extension. The slope represented the rate, RSM for rate of shear motion and RDM for rate of distraction motion. Strong linear correlations existed for RSM under physiological and posteroanterior acceleration (R 2 = 0.82 and 0.80) loadings, although differences were not statistically significant. Linear correlations for RDM in the anterior region for physiological and posteroanterior acceleration loadings (R 2 = 0.40 and 0.49) were weaker than RSM correlations and with no statistical difference. Linear correlation for RDM in the posterior region demonstrated statistical differences (p<0.05), with very low correlation for posteroanterior acceleration loading. Segmental and facet joint motion patterns were different between loading modes (Figure 15). For identical magnitudes of segmental extension, compared to physiological loading, posteroanterior acceleration loading resulted in markedly increased facet joint motion. Under segmental extension, C4-C5 anterior and posterior facet joint regions sustained approximately uniform posteriorly directed shear motion. RSM magnitudes were significantly greater (p<0.05) during posteroanterior acceleration than physiological loading in the anterior and posterior joint regions (Figure 16, top). Facet joints exhibited distraction in the anterior and compression in the posterior joint regions during segmental extension. RDM was significantly greater (p<0.05) due to posteroanterior acceleration than physiological loading in the anterior region. In the posterior joint region, RDM was not significantly different during both loadings, although physiological joint motions were generally greater (Figure 16, bottom). The finding that the kinematics of lower facet joints are fundamentally different between posteroanterior acceleration loading and normal physiological extension may explain the increased vulnerability of facet joints to injury. V. APPLICATION OF IN VITRO STUDIES TO COMPUTATIONAL MODELING Computational models developed in conjunction with experiments permit parametric studies and the determination of intrinsic tissue-level responses such as stresses. In addition, computational

Whiplash injuries In-vitro studies Page 15 models have the unique feature of absolute reproducibility and repeatability. Analysis of intrinsic biomechanics enhances the understanding of the behavior of structural components potentially involved in injury and/or pain producing processes. The following sections describe certain advances made in the area. A. DEVELOPMENT OF KINEMATIC CORRIDORS BASED ON IN VITRO STUDIES Using results from the described in vitro intact head-neck complex tests and adopting standard normalization procedures, mean plus/minus one standard deviation kinematic corridors were developed 133. Responses included head translation with respect to neck base, overall head-neck rotation (Figure 17), segmental rotation (Figure 18), and local facet joint component motion (Figure 19) as a function of spinal level and ΔV. Fiftieth percentile male head mass was used to scale individual head-neck complex specimen data. Corridors forming a dataset of motion responses serve as important data for validation and verification of computational models. B. VALIDATION OF A COMPUTATIONAL MODEL Using the described experimental results as a primary basis, a recently developed computational model was used for the analysis of intrinsic biomechanics. The head-neck model included rigid body representations for the head and seven cervical and T1 vertebrae 106. Mass and inertial properties of vertebrae and soft tissues were lumped into each body. To approximate the head and vertebral geometry, surfaces were created using three- and four-node finite elements. The rigid bodies were interconnected using discrete nonlinear viscoelastic elements to represent discs, facet joints, ligaments, and muscles. Nonlinear elastic contact interaction was defined between adjacent vertebrae. The normal lordotic curvature was incorporated using x-ray data from 48 human volunteers 117. Anterior and posterior longitudinal ligaments, ligamentum flavum, interspinous ligament, facet joint capsular ligaments, ligamentum nuchae, alar ligament, transverse ligament and tectorial membrane were simulated using Kelvin restraints. Leveldependent, nonlinear and viscoelastic ligament and disc material properties were based on literature and facet joints were simulated with high compressive stiffness 106. Musculature was simulated using 136 Hill muscle elements, accounting for 16 flexor and extensor muscle groups, and incorporating active and passive properties 28. The model was exercised under the same loading and boundary conditions used in intact head-neck complex tests. Parameters used in the

Whiplash injuries In-vitro studies Page 16 validation process included overall head-neck motions, level-dependent segmental angulations, and local facet joint kinematics. Responses of the computational model were within in vitro experimental corridors (Figures 17-19) for the first 100 ms for the overall head-neck extension, segmental angulation, and facet joint motion. Model responses also agreed with human volunteer vertebral and head-t1 angulations, and isolated cervical spine anterior longitudinal ligament elongations 107. The muscle contraction scheme was validated using rear impact tests on human volunteers with precontracted neck muscles 109. Although other models have included full-body 119 or head-neck complex geometries 122, validation has been typically limited to head and thorax kinematics. To the best knowledge of the authors, this is the most comprehensively validated and verified computational model for investigating the intrinsic biomechanics due to posteroanterior acceleration loading. Because of the high degree of confidence, the model was used in parametric studies to study effects of posture, thoracic ramping, anterior ligament distortion, and awareness. C. EFFECTS OF INITIAL SPINAL POSTURE The computational model was used to quantify local facet joint kinematics in terms of capsular ligament distractions for normal lordotic, kyphotic, and straightened spinal postures. The three curvatures were simulated using human volunteer x-rays 117. Boundary conditions in the three models were identical with the exception of T1 orientation, altered to accommodate each posture. Facet joint capsular ligament elongations were determined as a function of posture, level, and anatomic region. C2 to C7 elongations in the ventral, lateral, dorsal, and medial joint regions were determined (Figure 20). Elongations were investigated at the time of peak S- curvature. The transient nonphysiologic S-curvature lasted for 113, 110, and 119 ms for the normal, straight, and kyphotic postures. The normal posture responded with greatest elongations in the dorsal region at the C2 C3 level. However, from C3 to C7 levels, greatest elongations occurred in the lateral region. This finding is consistent with segmental angulations, wherein the C2 C3 segment demonstrated flexion, and C3 to C7 levels responded with extension. Segmental extension stretches ventral structures, while segmental flexion stretches dorsal structures.

Whiplash injuries In-vitro studies Page 17 In general, the kyphotic and straight postures responded with the same regional dependence as the normal posture (Figure 20). Abnormal postures increased C2 C3 capsular ligament dorsal elongations. The kyphotic posture responded with the largest increase in dorsal ligament elongation magnitude. Abnormal postures also increased ligament elongations in the lateral anatomic region, specifically at lower cervical levels (Figure 21). The straight posture resulted in increased elongations in the lateral region at C5 C6 and C6 C7 levels. In contrast, the kyphotic posture responded with increased elongations from C4 to C7 levels with the largest increase at the C5 C6 level. Clinical findings support these results, i.e., abnormal spinal curvature negatively influences spinal kinematics and injury. In a follow-up study investigating 146 patients five years after sustaining soft tissue injuries in automotive collisions, patients with kyphosis and no other degenerative changes at the time of injury had a significantly higher incidence of disc degeneration in the lower cervical spine than patients with normal curvature at the time of trauma 48. In most cases, disc degeneration developed only at one level. While degeneration at a specific level may not always lead to noxious response, it alters loading patterns in the cervical column, may influence degeneration at adjacent levels, results in osteophyte formation, accelerates progressive changes in curvature, or leads to spinal instability 23,33,74,96. Studies have shown that spondylosis at one level pre-disposes rostral levels to early instability 22. D. EFFECTS OF THORACIC RAMPING Upward sliding of the occupant in the seat and/or straightening of dorsal spine results in thoracic ramping. The base of the neck responds to the ramping action by displacing superiorly and rotating into extension. This added input at the base of the neck has been theorized to affect kinematics and the mechanism of injury 11,127. Magnitudes of thoracic motions are influenced by factors such as seatback angle, occupant anthropometry, spinal curvature, seatback properties, and ΔV. Effects of ramping on injury kinematics were studied using the described computational model. Ramping was applied as vertical displacement and extension of T1. Motion magnitudes were controlled by modulating T1 vertical acceleration and angular stiffness. The baseline vertical acceleration pulse was 40 ms in duration and 1.8 g peak. It was developed based on the acceleration required to induce the average magnitude of vertical T1 displacement

Whiplash injuries In-vitro studies Page 18 108. Vertical and horizontal T1 accelerations were initiated simultaneously. Baseline T1 extension angular stiffness was similar to C6-C7 sagittal plane bending stiffness. The T1 vertebra was unconstrained in the sagittal plane at the initiation of horizontal acceleration and began extension immediately. Peak T1 vertical displacement and extension rotation times were based on literature 27,98. Parameterization was used to investigate ramping effects by independently varying T1 vertical displacement, Group A, and T1 extension, Group B. In Group A, T1 angular stiffness remained constant, resulting in uniform maximum extension of approximately 16 o, occurring approximately at 220 ms. Meanwhile, T1 vertical acceleration was altered to increase and decrease vertical displacements (Figure 22). In Group B, T1 vertical acceleration was held constant, resulting in consistent temporal vertical displacements, with peak values of approximately 32 mm, occurring approximately at 210 ms. Meanwhile, T1 angular stiffness was altered to increase and decrease extension (Figure 23). Ranges of T1 vertical displacement and extension rotation encompassed the range of values reported in intact PMHS and human volunteer tests 108. C2 to T1 segmental angulations were analyzed during maximum S-curvature and at the end of the retraction phase. In Group A, extension and vertical displacement of T1 began immediately following initiation of posteroanterior acceleration loading and reached peak values at approximately 200 ms. Increasing vertical T1 displacement had minimal effects on segmental angulations during maximum S-curvature and at the end of the retraction phase. In all cases, segmental angulations decreased by less than one-degree between all T1 vertical displacements. In general, increasing T1 vertical displacement had greater effects on segmental angulations during maximum S- curvature than at the end of the retraction phase. However, the magnitude of this effect was minimal. Although vertical T1 displacement magnitudes increased by a factor of 13 during maximum S-curvature and 70 at the end of the retraction phase between the lowest and highest accelerations, C4 to T1 segmental angulations were altered by an average of only 0.6 o during maximum S-curvature and 0.1 o at the end of the retraction phase. Increasing T1 vertical displacement generally decreased lower cervical segmental angulations.

Whiplash injuries In-vitro studies Page 19 In Group B, maximum extensions ranged from 9 to 52 o. Maximum T1 extension (52 o ) occurred at 300 ms. Increasing T1 extension had a greater effect on kinematics during maximum S- curvature than at the end of the retraction phase. Segmental angulations were altered by 1.1 o (mean) during maximum S-curvature and 0.8 o (mean) at the end of the retraction phase. In general, increasing T1 extension decreased segmental angulations during maximum S-curvature and at the end of the retraction phase. Although thoracic ramping may not significantly influence kinematics due to posteroanterior acceleration loading, coupled motions may play a role in the biomechanics of load transfer and mechanism of injury 143. Clinical and experimental studies have identified cervical facet joints as a source of pain 8,20,27,101,103. Sagittal plane orientation of facet joints is such that compressive forces from vertical thoracic displacement result in posterior joint shear. This motion, in addition to the shearing motion resulting from retraction 103 may increase the likelihood of capsular ligament failure by decreasing joint stiffness; a similar hypothesis was advanced using in vitro quasi-static testing of C1-T1 columns 127. It should be noted that effects of thoracic ramping are modulated by variables including gender, posture, occupant awareness, ΔV, and anthropometry and demographics. E. ANTERIOR LONGITUDINAL LIGAMENT KINEMATICS The computational model was used to determine anterior longitudinal ligament stretch during the retraction phase. This was based on the clinical observation that the ligament sustains injuries in motor vehicle crashes 26,44,45,125. As indicated before, anterior longitudinal ligament injuries occurred in PMHS subjected to single rear impact acceleration 128. The computational model was subjected to 8.6, 10.8, and 13 km/h ΔV. Ligament distraction was measured between anterior extents of opposing body surfaces and compared between the three ΔV and quasi-static catastrophic failure magnitudes. Distractions at all levels increased with increasing ΔV. Maximum distractions occurred during the retraction phase in the lower spine, particularly at C5- C7 levels (Figure 24). Distractions at these levels were 52 and 28% of the anterior longitudinal ligament failure distractions 72. Peak distractions at C2 C4 levels occurred after the spine had transitioned into extension. The greatest distraction occurred at the C2 C3 level for 2.4 m/s ΔV and at the C3 C4 level for 3.0 and 3.6 m/s ΔV (Figure 25). Although peak distractions were

Whiplash injuries In-vitro studies Page 20 approximately uniform across all levels, C3-C6 distractions approached failure values. In particular, maximum C3 C4 distraction at higher ΔV was within one standard deviation limits, implying susceptibility. However, this occurred during the extension phase, which may be limited depending on occupant anthropometry, seating, backset, and head restraint design and effectiveness. Although C4-C6 distractions were below ultimate failure, these distractions may exceed subcatastrophic thresholds. The hypothesis that a nociceptive response may be induced by excess distraction of facet joint capsular ligaments, although catastrophic ligament failure may not occur, is also applicable to the anterior ligament due to the presence of nociceptors 57,58,67. Distraction magnitudes (Figure 25) exceeding subcatastrophic injury thresholds may result in hyperalgesia, and these findings demonstrate that the anterior ligament may be susceptible to trauma due to posteroanterior acceleration loading. Anterior ligament trauma has implications in spinal instability. Catastrophic injury can result in acute dysfunction while subcatastrophic injury may be chronic. Anterior ligament injuries correlate with extension instability 95. Injuries sustained due to posteroanterior acceleration loading typically show no neurological abnormalities 2. Due to laxity or local tear of the anterior longitudinal ligament, outer layers of the annular fibers may have decreased stiffness at the segmental level under subsequent physiological loads. Because the ligament is intimate to annular fibers, presuming all components of the segment are normal prior to injury, subcatastrophic ligament failure may lead to in chronic changes within the disc. The added hypermobility may lead to early spine degeneration and long-term segmental instability stemming from the anterior column 29. F. REFLEX MUSCLE CONTRACTION IN UNAWARE OCCUPANTS Due to posteroanterior acceleration loading, reflex muscle contraction in unaware occupants occurs through a centrally generated response or a localized muscle stretch reflex 13,59,116. Muscle contraction stiffens the head-neck complex, and may decrease spinal motions and injury likelihood 13,54,59. Although not conclusive, some studies have theorized that injurious motions occur during the retraction phase before attaining considerable muscle contraction 20,35,61,83,86,104.