1 Clinical Biomechanics 2 (25) Influence of thoracic ramping on whiplash kinematics Brian D. Stemper a,b, *, Narayan Yoganandan a,b, Raj D. Rao c, Frank A. Pintar a,b a Department of Neurosurgery, Medical College of Wisconsin, 92 W. Wisconsin Ave., Milwaukee, WI53226, USA b Department of Veterans Affairs Medical Center, 5 W. National Ave Research 151, Milwaukee, WI53295, USA c Department of Orthopaedic Surgery, 92 W. Wisconsin Ave., Milwaukee, WI53226, USA Received 22 November 24; accepted 2 June 25 Abstract Background. Some experimental whiplash investigations using human volunteers and full-body cadavers reported thoracic ramping, characterized by superior translation and extension rotation of the cervico-thoracic junction. The effect of this phenomenon on cervical spinal kinematics has not been quantitatively determined. Methods. A comprehensively validated computational model exercised in 2.7 m/s rear impact was used to determine effects of superior translation and extension rotation of T1 on cervical segmental kinematics during the retraction phase. Findings. In general, thoracic ramping had a minimal effect on cervical intervertebral kinematics during retraction. Interpretation. Results of the present study demonstrated that magnitude of thoracic ramping plays a minimal role in the whiplash injury mechanism due to decreased effect on cervical segmental kinematics. Ó 25 Elsevier Ltd. All rights reserved. Keywords: Biomechanics; Computer model; Whiplash; Thoracic ramping; Parametric study; Cervical spine 1. Introduction Whiplash injuries commonly occur in low-velocity rear-end automotive impacts. These impacts result in abnormal inertial loading of the head neck complex for occupants of the struck vehicle that is abruptly accelerated anteriorly (Cusick et al., 21; Luan et al., 2; Siegmund et al., 23). Interaction of the occupant with the seatback results in three kinematically coupled motions at the base of the cervical spine that load the head neck complex; each may affect spinal soft tissue injury. The primary motion is anteriorly directed acceleration, commonly used to quantify rear-impact severity. However, secondary thoracic motions include superiorly directed acceleration and extension rotation of the upper * Corresponding author. Address: Department of Neurosurgery, Medical College of Wisconsin, 92 W. Wisconsin Ave., Milwaukee, WI 53226, USA. address: (B.D. Stemper). thoracic spine and were previously referred to as thoracic ramping (Davidsson et al., 1998; Deng et al., 2; Luan et al., 2; McConnell et al., 1993). Because no contact-induced loads are placed on the head neck complex during initial stages, the head remains stationary due to inertia as the thorax displaces anteriorly. This motion results in relative retraction of the head, forcing the cervical spine to compensate with S-shaped curvature characterized by extension in lower and flexion in upper segments (Cusick et al., 21; Deng et al., 2; Grauer et al., 1997; Panjabi et al., 1998; Stemper et al., 23). Experimental research using human volunteers and cadavers has often cited retraction as the time at which whiplash injury occurs (Cusick et al., 21; Deng et al., 2; Grauer et al., 1997; Penning, 1992; Stemper et al., 23). Injury mechanisms involving excess cervical soft tissue distortion were hypothesized to occur during this period (Bostrom et al., 1996; Brault et al., 2; Cusick et al., 21; Pearson et al., 24; Stemper et al., 24a) /$ - see front matter Ó 25 Elsevier Ltd. All rights reserved. doi:1.116/j.clinbiomech
2 12 B.D. Stemper et al. / Clinical Biomechanics 2 (25) Secondary thoracic kinematics may influence the whiplash injury mechanism. Experimental studies using human volunteers and full-body cadavers described ramping motion to result from upward sliding of the occupant in the seat and/or straightening of thoracic kyphosis (Bertholon et al., 2; Davidsson et al., 1998; Deng et al., 2; McConnell et al., 1993; Ono et al., 1997; Siegmund et al., 1997; van den Kroonenberg et al., 1998). The base of the cervical spine responds to thoracic motion by displacing superiorly and rotating into extension. Motion magnitudes are influenced by numerous factors including seatback angle, occupant anthropometry, impact magnitude, initial thoracic curvature, and seatback material properties. The degree to which thoracic ramping affects cervical spinal kinematics and, more importantly, the whiplash injury mechanism remains unclear. A limited number of experimental studies quantitatively reported thoracic ramping (Table 1). In general, magnitudes of thoracic vertical displacement and extension angulation have widely varied. Deng et al. (2) tested human cadavers in rear impact using and 2 seatback angles and reported initial thoracic curvature affected cervical kinematics. Extension angulation of T1 was greater with 2 than seatbacks. However, magnitude of thoracic vertical acceleration was not affected by seatback angle. Using a rigid seat with 24 seatback, Yoganandan et al. (2b) subjected one male and four female cadavers to single rear impacts at 4.1 or 6.9 m/s. Magnitude of vertical thoracic acceleration varied between 2.4 and 79.8 m/s 2 ; higher velocity tests resulted in greater vertical acceleration. Siegmund et al. (1997) subjected human volunteers to 1.1 and 2.2 m/s vehicle-to-vehicle rear impacts. Higher velocities resulted in greater vertical displacement of the cervicothoracic junction. These studies quantitatively described thoracic ramping. However, its effect on cervical kinematics and the whiplash injury mechanism has not been systematically explored. The purpose of this investigation was to quantify thoracic ramping effects on cervical kinematics in whiplash. Parametric analysis and a comprehensively validated computational model were used to independently account for effects of vertical displacement and extension rotation of T1 on cervical segmental kinematics during the retraction phase of whiplash. Vertical thoracic displacement and extension rotation were varied within the literature range. Segmental kinematics describe level-by-level soft tissue distortion (Stemper et al., 24a). Because motion and injury are related, increasing segmental angulations leads to excess tissue distortion, eventually resulting in catastrophic or subcatastrophic failure that may be responsible for whiplash injuries. 2. Methods A head neck computational model was used to investigate effects of thoracic ramping on cervical kinematics in whiplash (Stemper et al., 24b). The model was exercised using MADYMO computer modeling software (TNO Automotive, Delft, The Netherlands) and consisted of the skull, seven cervical vertebrae, and first thoracic vertebra represented as rigid bodies characterizing mass and inertial properties of bone and surrounding soft tissues. The model approximated 5th percentile male geometry. Soft tissue spinal components were incorporated using discrete elements encompassing tissue-specific degrees of freedom. Nonlinear, viscoelastic, and level-dependent material properties were obtained from literature: ligament tension (Pintar, 1986; Yoganandan et al., 2a), intervertebral disc shear (Moroney et al., 1988), tension (Pintar et al., 1986), and compression (Eberlein et al., 1999), facet joint compression (van der Horst, 22), OC-C1 flexion/extension (Goel et al., 1988), and C1 C2 axial rotation (Goel et al., 1988). Maximum static values for ligament and intervertebral disc stiffness are presented in Table 2. Gravitational acceleration (9.81 m/s 2 ) was included in all simulations. Sagittal plane linear displacement of T1 was controlled using horizontal and vertical acceleration inputs. Sagittal plane rotation of T1 was controlled using a nonlinear stiffness characteristic, similar to lower cervical angular stiffness (C6 C7). The T1 vertebra was constrained against linear and angular motions outside the sagittal plane. Initial conditions were controlled specifically to represent the normal population subjected to unexpected rear impact. Cervical posture was shown to affect Table 1 Thoracic ramping kinematics from experimental investigations implementing human volunteers (v) and full-body human cadaver (c) subjects Investigator Sled/vehicle impact severity (m/s) T1 vertical acceleration (g) T1 vertical displacement (mm) Bertholon et al., 2 (c) Davidsson et al., 1998 (c) Deng et al., 2 (c) Ono et al., 1997 (v) 1.1, 1.7, <3 Siegmund et al., 1997 (v) 1.1, van den Kroonenberg et al., 1998 (v) T1 Angulation ( )
3 B.D. Stemper et al. / Clinical Biomechanics 2 (25) Table 2 Static linear material property values for model soft tissue components Structure Component Value Intervertebral disc Anterior posterior shear 62. N/mm Posterior anterior shear 5. N/mm Lateral shear 73. N/mm Tension 64. N/mm Compression a 82.3 N/mm Anterior longitudinal Tension a 14.7 N/mm b ligament Posterior longitudinal 16.4 N/mm a ligament Ligamentum flavum 16.3 N/mm a Interspinous ligament 5.1 N/mm a Facet joint capsular ligament 23.3 N/mm a Passive muscle Linear spring constant.334 N/mm Strain asymptote.7 a Maximum stiffness for piecewise linear approximation. b Mean stiffness across cervical levels. head neck kinematics under a variety of dynamic loading conditions (Liu and Dai, 1989; Maiman et al., 22; Portnoy et al., 1972; Stemper et al., 25). Therefore, initial cervical orientation reflected mean lordotic curvature of 48 normal human volunteers (Takeshima et al., 22), T1 was given an initial anterior orientation of 25, and occipital condyles were positioned directly superior to the T1 vertebral body (Stemper et al., 23, 24b). The Frankfort plane was oriented horizontally prior to whiplash acceleration (Cusick et al., 21; van den Kroonenberg et al., 1998; Yoganandan et al., 2b). Active muscle contraction was not included. Passive muscle response of 32 muscles was approximated using Hill-type elements. Muscles were assigned localized attachments at most vertebral levels between origin and insertion points (van der Horst, 22). Nonlinear force length relationship was modeled according to Deng and Goldsmith (1987), wherein passive neck muscle forces were increased by a factor of ten to account for baseline muscle tone (Table 2). The model was subjected to rear-impact loading applied to T1 as a postero-anteriorly directed acceleration trace. Pulse shape was modeled after horizontal thoracic acceleration measured in full-body cadavers (Yoganandan et al., 2b). The pulse was integrated over the active phase to obtain T1 change in velocity (2.7 m/s). The T1 pulse had maximum acceleration of 4.2 g and pulse width of 125 ms. Speed change of T1 is not necessarily equal to automobile or seatback speed change. However, rear-impact magnitude fit within the range of impact severities at which epidemiological investigations reported whiplash injuries (Temming and Zobel, 2). Overall, segmental, and localized kinematics of the model were validated previously against experimental rear-impact data (Stemper et al., 24b). Model response (e.g., head angulation and displacement relative to T1, level-by-level segmental angulations, and localized region-dependent facet joint motions) fit within mean and standard deviation corridors during the retraction phase. To attain kinematic validation, material properties were altered within literature variation, and spinal alignment was changed from the original MADYMO model (Happee et al., 1998; van den Kroonenberg et al., 1997). Because previous model validation did not include thoracic ramping, the baseline model (described below) was subjected to secondary validation. Segmental angulations were compared to maximum segmental data from full-body cadavers subjected to similar rear-impact velocities (< 3. m/s) with 2 seatback and headrest (Deng et al., 2). Full-body cadavers demonstrated a minimum of 3 T1 extension. Segmental extension magnitudes from the baseline model (C3 C4 through C5 C6) obtained at the end of the retraction phase fit within experimental standard deviations (Fig. 1). Maximum flexion at C2 C3 also fit within experimental standard deviations. The C6 C7 and C7 T1 levels were not included in this comparison as experimental data were not available. Thoracic ramping was applied as vertical displacement and extension angulation of T1. These motions were controlled by modulating T1 vertical acceleration and angular stiffness. The baseline vertical acceleration pulse was based on literature (Siegmund et al., 1997) and consisted of 17.5 m/s 2 maximum vertical acceleration with approximately 4 ms pulse duration. Vertical acceleration was initiated simultaneously with initiation of horizontal T1 acceleration. Baseline T1 angular stiffness was similar to lower cervical segmental stiffness (C6 C7). The T1 vertebra was unconstrained in the sagittal plane at the initiation of horizontal T1 acceleration and began extension angulation immediately. Timing of vertical displacement and extension angulation was consistent with previous investigations (Deng et al., 2; Siegmund et al., 1997). Segmental angle (deg) Maximum angulations from Dengetal., 2 Model data from maximum S-curvature Model data from end of the retraction phase c2c3 c3c4 c4c5 c5c6 Fig. 1. Secondary validation of the model relative to maximum segmental angulations measured in full-body cadaver experiments (darker bars). Computer model segmental angulations were obtained at maximum S-curvature (C2 C3) or at the end of the retraction phase (C3 C4 to C5 C6).
4 122 B.D. Stemper et al. / Clinical Biomechanics 2 (25) Vertical T1 displacement (m).4.2 (a) Time of maximum S-curvature. End of the retraction phase Time (ms) 1 T1 extension angulation (deg) (b) Time of maximum S-curvature. End of the retraction phase Time (ms) Fig. 2. Thoracic ramping motions for (a) Group 1 and (b) Group 2 simulations. Times of kinematic measurement are indicated. Baseline traces are indicated in bold. The effect of thoracic ramping on cervical kinematics was investigated in two steps: independently varying magnitude of T1 vertical displacement (Group 1) and T1 extension angulation (Group 2). In Group 1 simulations, T1 angular stiffness was constant, resulting in uniform maximum extension values of approximately 15.9 (occurring at ms). Meanwhile, T1 vertical acceleration was altered within the literature range to increase and decrease vertical displacement (Fig. 2a). In Group 2 simulations, T1 vertical acceleration was constant, resulting in equal temporal vertical displacements, with uniform maximum values of approximately 32. mm (occurring at ms). Meanwhile, T1 sagittal plane angular stiffness was altered to increase and decrease extension angulation (Fig. 2b). Maximum extension angulations were within the literature range. Segmental angulations, defined as the sagittal plane angle of one vertebra with respect to the immediately adjacent vertebra, were analyzed from C2 C3 through C7 T1 during maximum S-curvature and at the end of the retraction phase. Maximum S-curvature was defined as the time at which C2 C3 sustained maximum flexion. The end of the retraction phase was defined as the time that the C2 C3 segment returned to its initial sagittal plane angle and began to rotate into extension. 3. Results The cervical spine demonstrated retraction and extension phases of whiplash (Fig. 3a). During the retraction phase, the cervical spine sustained S-curvature as upper
5 B.D. Stemper et al. / Clinical Biomechanics 2 (25) Fig. 3. (a) Representative initial position, retraction, and extension phases of whiplash demonstrated by the baseline computational model. Initial position of the T1 vertebral body is indicated by the x z coordinate system in each figure and (b) segmental angulations of the baseline model. Table 3 Thoracic ramping kinematics for simulations with constant T1 angular stiffness and altered T1 vertical displacement (Group 1 simulations) Max. vertical accel. (m/s 2 ) Values at maximum S-curvature Values at end of retraction phase Vertical displ. (mm) T1 angulation ( ) Time (ms) Vertical displ. (mm) T1 angulation ( ) Time (ms) segments demonstrated flexion and lower segments demonstrated extension (Fig. 3b). Flexion at C2 C3 lasted approximately 15 ms, after which all cervical segments demonstrated extension. Group 1 simulations applied constant T1 extension while increasing T1 vertical acceleration, resulting in maximum displacements of mm. Six simulations were performed in Group 1. Extension angulation and vertical displacement of T1 began immediately following initiation of anterior acceleration and reached maximum values at approximately 2 ms. T1 kinematics for Group 1 simulations are provided in Table 3. Increasing vertical T1 displacement had a minimal effect on segmental angulations during maximum S-curvature (Fig. 4a) and at the end of the retraction phase (Fig. 4b). In all cases, segmental angulations were decreased by less than.9 from lowest to highest T1 vertical displacements. In general, increasing T1 vertical displacement had a greater effect on segmental angulations during maximum S-curvature than at the end of the retraction phase. However, magnitude of this effect was minimal. Although vertical T1 displacement magnitudes increased by a factor of 12.9 during maximum S-curvature
6 124 B.D. Stemper et al. / Clinical Biomechanics 2 (25) Segmental angle (deg) C7-T1 C6-C7 C5-C6 C4-C5 C3-C4 C2-C3 Segmental angle (deg) C7-T1 C6-C7 C5-C6 C4-C5 C3-C4 C2-C (a) T1 vertical displacement at maximum S-curvature (mm) Segmental angle (deg) C6-C7 C5-C6 C7-T1 C4-C5 C3-C4 C2-C (b) T1 vertical displacement at the end of the retraction phase (mm) Fig. 4. Cervical segmental angulations for Group 1 simulations (a) at maximum cervical S-curvature and (b) at the end of the retraction phase. (a) Segmental angle (deg) (b) T1 angle at maximum s-curvature (deg) C6-C7 C5-C6 C7-T1 C4-C5 C3-C4 C2-C T1 angle at the end of the retraction phase (deg) Fig. 5. Cervical segmental angulations for Group 2 simulations (a) at maximum cervical S-curvature and (b) at the end of the retraction phase. Table 4 Thoracic ramping kinematics for simulations with constant T1 vertical acceleration and altered T1 extension angulation (Group 2 simulations) T1 angular stiffness Values at maximum S-curvature Values at end of retraction phase Vertical displ. (mm) T1 angulation ( ) Time (ms) Vertical displ. (mm) T1 angulation ( ) Time (ms) 2.k a k k k k a k = Baseline T1 extension stiffness. and 7.4 at the end of the retraction phase between lowest and highest accelerations, lower cervical segmental angulations (C4 C5 through C7 T1) were altered by an average of only.62 during maximum S-curvature and.5 at the end of the retraction phase. Increasing T1 vertical displacement generally decreased lower cervical segmental angulations. Group 2 simulations applied constant temporal T1 vertical displacements while varying T1 sagittal plane angular stiffness, resulting in maximum extension magnitudes of Maximum extension of T1 (52.3 ) occurred at 3 ms. A total of five simulations were performed in Group 2. T1 kinematics for Group 2 simulations are provided in Table 4. Increasing T1 extension had a greater effect on cervical kinematics during maximum S-curvature (Fig. 5a) than at the end of the retraction phase (Fig. 5b). Segmental angulations were altered by mean 1.1 (SD.1) during maximum S-curvature and mean.8 (SD.5) at the end of the retraction phase. In general, increasing T1 extension decreased cervical segmental angulations during maximum S-curvature and at the end of the retraction phase. 4. Discussion The present study quantified effects of thoracic ramping on kinematics of the cervical spine in whiplash using parametric analysis and a comprehensively validated computational model. Thoracic ramping was applied
7 B.D. Stemper et al. / Clinical Biomechanics 2 (25) Maximum T1 extension (deg) Present Model Literature Maximum T1 vertical displacement (mm) Present Model Literature Fig. 6. Comparison of thoracic ramping motions from the present model to literature variation. Literature values were derived from studies presented in Table 1. by independently modulating vertical thoracic displacement (Group 1) and thoracic extension (Group 2) to encompass the wide range of values reported in literature (Fig. 6). Thoracic ramping had a minimal effect on cervical kinematics both during the retraction phase (maximum S-curvature) and at the end of the retraction phase, altering segmental angulations by less than 1.5 in all cases. Although slight increases in segmental angulation occurred at C2 C3 and C3 C4, and in one case segmental extension at C4 C5, increasing magnitudes of ramping generally decreased segmental angulations, with a greater effect during maximum S-curvature than at the end of the retraction phase. However, due to small variations in level-by-level segmental angulations (e.g., spinal tissues undergo approximately equal deformation magnitudes), magnitude of thoracic ramping plays a minimal role in whiplash injury mechanisms involving excess soft tissue distortion. Maximum segmental angulations obtained during the retraction phase from the present model were compared to segmental active range of motion (RoM) data from two studies using asymptomatic volunteers (Table 5). In general, mean experimental RoM values were similar, with a maximum difference of 12.9% at the C5 C6 level. Maximum computer model retraction phase angulations Table 5 Comparison of model angulations to literature range of motion values ( ) C2 C3 C3 C4 C4 C5 C5 C6 C6 C7 Johnson 7.2 (.9) a 7.8 (1.1) 9.8 (1.2) 1.5 (1.3) 8.2 (1.2) et al., 1977 Ordway 6.5 (4.8) a 8.3 (5.1) 9.5 (3.9) 9.3 (3.8) 8.3 (5.8) et al., 1999 Present.21 a model b a Flexion angulation, values in parentheses indicate standard deviations. b Maximum retraction phase angulations. were well below experimental RoM values with the exception of C5 C6, which approached standard deviations, and C6 C7, which was within or exceeded standard deviations. This finding, however, is not indicative of injury at those levels. Experimental studies reported active voluntary RoM, which is likely well below the threshold of injury. Maximum segmental extension approaching voluntary limits and nearing the injurious range is consistent with previous findings. For example, following exposure of full-body cadavers to rear impacts, Yoganandan et al. (21) used cryomicrotomy to identify spinal ligament and facet joint failure concentrated primarily in lower cervical levels. Likewise, in an investigation of patients with chronic neck pain after whiplash, Barnsley et al. (1995) reported posterior neck pain could be attributed to lower cervical facet joint injuries. Clinical and epidemiological studies support this conclusion, often citing posterior neck pain as the most common complaint for whiplash patients (Cassidy et al., 2; Hildingsson and Toolanen, 199; Karlsborg et al., 1997; Mayou and Bryant, 1996; Norris and Watt, 1983; Radanov et al., 1995; Sturzenegger et al., 1994). While it is difficult to determine tissue failure and impossible to diagnose pain using a computer model, present results demonstrated a trend of lower cervical motions approaching injurious limits. However, magnitude of thoracic ramping had a minimal effect on these motions. Present findings are contrary to previous investigations asserting that thoracic ramping plays a significant role in whiplash (Kettler et al., 24). While the present study implemented a head neck model with overall, segmental, and localized rear-impact kinematics validated against experiments at three impact velocities, the previous investigation was limited in scope. For example, the conclusion that T1 angulation affects whiplash biomechanics was based on head accelerations from side impact testing. Although side-impacts can result in soft-tissue neck injuries, the impact direction differs from the traditional whiplash definition (rear impact), and biomechanics of the head neck complex are markedly different. The previous study also implemented a dummy neck (solid non-segmented column) to assess only global biomechanics. In contrast, the present model used a segmented cervical column exercised in rear impact to investigate effects of T1 extension and vertical displacement on segmental angulations during retraction. Level-by-level segmental angulations provide a more detailed description of temporal spinal kinematics and effects of thoracic ramping. Although thoracic ramping may not significantly influence cervical kinematics in whiplash, coupled thoracic motions may still play a role in the whiplash injury mechanism. Several previous clinical and experimental investigations implicated cervical facet joints in whiplash injury (Barnsley et al., 1995; Cusick et al., 21;
8 126 B.D. Stemper et al. / Clinical Biomechanics 2 (25) Deng et al., 2; Stemper et al., 24a). Sagittal plane orientation of cervical facet joints (45 ) is such that compressive forces resulting from vertical thoracic displacement may result in posterior joint shearing. This motion, in addition to the shearing motion resulting from retraction (Stemper et al., 24a), may increase likelihood of capsular ligament tensile failure by decreasing joint stiffness. A similar assertion was made using quasi-static testing and isolated cervical spine specimens (Yang et al., 1997). Although thoracic ramping may contribute to the whiplash injury mechanism, magnitude of its effect may be modulated by numerous variables including gender, cervical posture, occupant awareness, impact severity, and occupant anthropometry and age. Timing of the whiplash injury mechanism is important in determining effects of kinematic factors such as thoracic ramping. Clinical and experimental investigations implicated the retraction phase as the time of whiplash injury based on abnormal spinal loading. Retraction phase injury theories include facet joint ligament injury, nerve root injury from pressure gradients in the spinal canal, and eccentric muscle contraction injuries to the anterior neck muscles (Aldman, 1986; Bostrom et al., 1996; Brault et al., 2; Cusick et al., 21; Deng et al., 2; Grauer et al., 1997; Panjabi et al., 1998; Penning, 1992; Stemper et al., 23). Retraction occurs early in the whiplash event. Results of the present investigation demonstrated that thoracic ramping has a minimal effect on injury mechanisms that occur during this time. Frankfort plane angle of mean 8 (SD 5 ) was associated with normal posture in human volunteers (Vasavada et al., 21). To demonstrate model sensitivity to initial orientation changes, a simulation was performed wherein the Frankfort plane was extended 8 from the baseline model. Results demonstrated minimal variation in segmental angulations during the retraction phase. In particular, time of maximum S-curvature occurred four msec later, and the end of the retraction phase occurred six msec earlier. For C2 C3 through C7 T1, mean segmental angulations were altered by.92 at maximum S-curvature and.11 at the end of the retraction phase. Decreased sensitivity to initial head orientation may be attributed to flexion/extension occurring primarily at the atlanto-occipital joint. Therefore, small changes in orientation of the Frankfort plane prior to whiplash loading do not markedly affect initial spinal orientation. Limitations of the present model include absence of active muscle contraction and head restraint. In response to unexpected thoracic acceleration, neck muscles reflexively contract to stiffen the head neck complex and may decrease injurious motions (Siegmund and Brault, 2). However, reflex neck muscle contraction in the unaware occupant is characterized by inherent delays, and muscles may not generate sufficient forces to alter spinal kinematics during the retraction phase (Stemper et al., in press). The present focus on the retraction phase, ending at a mean time of 143 ms, reduces the necessity of modeling reflex muscle contraction. Other studies demonstrated that muscles affect head neck response during frontal and lateral impacts (Brolin et al., 25; de Jager et al., 1996; van der Horst et al., 1997; Williams and Belytschko, 1983). However, injuries resulting from frontal and side impacts occur late in the event, giving muscles sufficient time to react and generate forces. Similarly, analysis of kinematics during the retraction phase also reduces the necessity for a head restraint, as the head contacts the head restraint just prior to peak retraction (Davidsson et al., 1998). Therefore, peak retraction magnitude is a function of the position of the head relative to the head restraint before impact. However, results of the present investigation demonstrated that thoracic ramping had a greater effect on kinematics earlier in the event (prior to the time of peak retraction). Because of this, the model is sufficient to delineate effects of thoracic ramping on cervical kinematics in whiplash, as this effect decreases from the time of maximum S-curvature to the end of the retraction phase. 5. Conclusions A comprehensively validated head neck computational model demonstrated that increasing magnitudes of vertical acceleration and extension rotation of T1 have a minimal effect on spinal kinematics during the whiplash retraction phase, the time most commonly associated with soft tissue injury. Acknowledgements This study was supported in part by PHS CDC Grant R49CCR , and the Department of Veterans Affairs Medical Research. 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