Response Corridors of the Human Head-Neck Complex in Rear Impact

Abstract

In recent years, anatomically accurate dummies were developed to produce a more biofidelic response to rear impacts. The detailed dummy geometry permits more comprehensive kinematic validation, which is typically accomplished through response corridors developed using human volunteers and cadavers. In the present investigation, ten intact head-neck cadaver specimens were subjected to 1.8 and 2.6 m/s rear impacts. Response corridors were developed for overall head to T1 angulation, head retraction, and segmental angulations from C2–C3 through C6–C7 levels. The corridors were developed to emphasize the detailed validation of dummy response during initial stages of whiplash loading.

Experiments using human volunteers and cadavers are typically used to evaluate biofidelity of dummies [Mertz and Patrick, 1971; Eppinger et al., 1984; Roberts et al., 1990; Maltese et al., 2002]. For example, to design and develop dummies for side impacts (WoldSID, EuroSID II, EuroSID IIre, etc.) intact human cadaver sled tests were conducted under varying conditions (e.g., low- and high-velocity impacts) [Maltese et al., 2002]. The widely used and federally regulated frontal impact Hybrid III dummy was developed based on human cadaver experiments. Most commonly, kinematic and kinetic responses of dummies are compared to experimental corridors developed from mean and standard deviation of volunteer and cadaver responses [Yoganandan et al., 2002]. Dummy responses falling within the corridors are deemed valid (representative of the normal human response) for the given loading vector. Corridors are obtained under experimental conditions modeling the actual automotive impact or inertial loading applied to the occupant while minimizing variables to facilitate reproducibility. Although numerous channels of experimental data are available (e.g., displacements, accelerations, forces, moments), dummy response is typically validated with respect to the responses that most accurately and comprehensively characterize the injury mechanism for the given loading vector [Mertz and Patrick, 1971; Eppinger et al., 1984; Roberts et al., 1990; Maltese et al., 2002].

In the 1970s, the Hybrid III dummy was developed as a human surrogate for the design and evaluation of automotive safety systems for frontal impacts [Foster et al., 1977]. Although the Hybrid III dummy remains the most commonly used, when tested in rear impact (whiplash), it was repeatedly shown to have a response different from human volunteers. In particular, it was reported that the Hybrid III response was too stiff relative to cadaver specimens in rear impact [Cappon et al., 2000] or that the response was closer to that of human volunteers with initially tensed neck musculature [Kim et al., 2003]. In the 1990s, the RID and TRID-II necks were developed and incorporated with Hybrid III elements to produce a more biofidelic response of the head-neck complex in rear impact [Svensson and Lovsund, 1992; Thunnissen et al., 1996]. The dummy necks included segmented geometry interconnected by hinge joints and deformable elements. More recently, the BioRID series (I through P3) and RID2 were developed, incorporating advanced cervical geometry along with unique thorax and abdomen designs [Davidsson et al., 1998; Davidsson et al., 1999; Cappon et al., 2001]. The overall, head, and T1 kinematic responses of these dummies were shown to be more biofidelic than the Hybrid III during rear impacts [Linder et al., 2000; Siegmund et al., 2001].

A more biofidelic kinematic response of the head-neck complex in rear impact was attained in the newer generation of dummies through the incorporation of more accurate cervical geometry, e.g., segmented spinal column [Siegmund et al., 2001]. The human cervical spine sustains three unique kinematic phases in whiplash. The S-curvature phase is initiated by the lag of the head behind the thorax due to its inertia (retraction). The C-curvature phase follows and is characterized by overall extension of the head and neck. The final phase is rebound, the time during which the head and neck move into flexion after making contact with the head restraint. The three kinematic phases are characterized by unique segmental responses at each spinal level. Due to a paucity of detailed segmental data, validation of dummy segmental responses in rear impact has not been attained. In the present protocol, segmental validation of dummy responses in rear impact was deemed important for two reasons. First, the biphasic S-curvature demonstrates level-dependent segmental angulations and has been hypothesized to be the time during which whiplash injury occurs [Kaneoka et al., 1999; Deng et al., 2000; Yoganandan et al., 2001; Stemper et al., 2003]. Secondly, whiplash injury is manifested by the failure of local soft tissue components [Barnsley et al., 1994]. For a dummy to be useful in the design and testing of automotive safety systems, it must be capable of accurately reproducing the human response to rear impact inertial loading. To achieve this, a dummy should be designed and validated based on segmental kinematics in addition to overall kinematics and kinetics. Therefore, the purpose of the present investigation was to develop rear impact kinematic corridors to be compared to the response of existing and future dummies for the validation of overall head-neck angulation and retraction and segmental angulations.

METHODS

Five male and five female cadaver specimens were selected based on lack of appreciable spinal degeneration and gender/age matching criteria. All specimens were screened for HIV and Hepatitis A, B, and C prior to testing according to accepted procedures. Specimen characteristics are presented in Table 1.

Table 1.

Specimen characteristics.

	Age (yrs)	Height (cm)	Body weight (kg)
Male	61.8 ± 13.1	173.5 ± 2.8	75.0 ± 15.9
Female	55.2 ± 17.2	167.2 ± 2.2	54.6 ± 14.1

Head-neck specimens were isolated at the T2 level, leaving skin and musculature of the head and neck intact. The esophagus and trachea were removed to facilitate the potting procedure. Specimens were potted in polymethylmethacrylate (PMMA) at the level of T1, which was oriented 25 deg anteriorly to simulate normal driving posture [Stemper et al., 2003]. Thoracic fixation eliminated ramping motion of T1. This limitation is covered in the discussion. Skin and musculature of the neck were attached to the PMMA base to simulate passive muscle resistance to flexion/extension motion of the head-neck complex. A small portion of skin and musculature was removed from the right lateral side of the specimen to expose the spinal column. Photoreflective targets were placed on vertebral bodies and lateral masses of C2 through C7 vertebrae. Two targets were also placed along the lateral side of C1 and the Frankfort plane of the skull. Spinal targets were used to measure sagittal plane angular motions of the cervical vertebrae and head. Segmental angulations were calculated as the angle of one vertebra with respect to the immediately adjacent vertebra (Figure 1). Retraction of the head relative to T1 was computed as the horizontal displacement of the head relative to T1.

Figure 1.

Definition of segmental angulation.

Rear impact loading was applied using a pendulum-minisled device (Figure 2). Specimens were mounted to the minisled and oriented such that the occipital condyles were directly superior to the T1 vertebral body and the Frankfort plane was maintained horizontal using a fixation device.

Figure 2.

Experimental test setup.

The pendulum strike at the posterior edge of the minisled created the impact pulse, and the minisled was allowed to accelerate in the anterior direction. An accelerometer attached to the minisled measured the posterior to anterior acceleration (Figure 3), which was integrated to obtain change in velocity (ΔV). Testing was conducted at 1.8 and 2.6 m/s impact velocities, which were consistent with the range of velocities at which the risk of whiplash injury is greatest [Temming and Zobel, 2000]. After each test, specimens were visually and radiographically inspected for injury, defined as dislocation of the facet joint or intervertebral disc, ligament or endplate failure, or bony disruption. Head-neck kinematics were obtained using a high-resolution digital camera at 1,000 frames per second. Data were filtered according to SAE specifications.

Figure 3.

Minisled acceleration for 2.6 m/s impact velocity.

Specimen responses were normalized with respect to standard 50^th percentile anthropometry [Eppinger et al., 1984; Maltese et al., 2002]. Mass scaling was accomplished through comparison of the head mass of each specimen to the head mass of the 50^th percentile male, using the λ_m scaling factor. Equation 1 was applied to segmental, overall, and retraction responses of the individual specimens.

$${\theta }_{\text{s}}={\lambda }_{\text{m}}^{1/3}{\theta }_{\text{i}}$$ (1)

The mean scaled responses plus and minus one standard deviation were used to develop experimental corridors for segmental angulations at C2–C3, C3–C4, C4–C5, C5–C6, and C6–C7 levels, overall head to T1 angulation, and retraction of the head relative to T1. Corridors were developed for each impact velocity.

RESULTS

Specimens sustained cervical S-curvature, characterized by flexion in upper and extension in lower cervical segments, during the active loading phase. The S-curvature lasted approximately 125 ms and was immediately followed by overall extension of the head and cervical spine (C-curvature). No specimen sustained injury.

Validation corridors for segmental angulations from C2–C3 through C6–C7 levels are presented in figures 4 through 8 for 1.8 m/s and figures 9 through 13 for 2.6 m/s. Segmental angulations demonstrated level-dependent behavior during S-curvature. During this time, the C2–C3 and C3–C4 levels sustained mean flexion angulation while the C4–C5 through C6–C7 levels sustained mean extension. After 125 ms, all levels demonstrated extension angulation. Magnitudes of segmental angulation were greater for 2.6 m/s impact velocity than 1.8 m/s velocity.

Figure 4.

C2–C3 segmental angulation at 1.8 m/s.

Figure 8.

C6–C7 segmental angulation at 1.8 m/s.

Figure 9.

C2–C3 segmental angulation at 2.6 m/s.

Figure 13.

C6–C7 segmental angulation at 2.6 m/s.

Extension angulation of the head was delayed approximately 50 ms after the initiation of T1 acceleration (Figures 14 and 15). After 50 ms, the head moved into extension relative to the T1 vertebra. Mean extension angles at the end of the loading phase (100 ms) were 23.6 deg for 2.6 m/s and 19.1 deg for 1.8 m/s.

Figure 14.

Head angulation relative to T1 at 1.8 m/s.

Figure 15.

Head angulation relative to T1 at 2.6 m/s.

During the initial stages of loading, the head translated posteriorly relative to T1 (retraction) for both impact velocities (Figures 16 and 17). This posterior translation was the result of T1 accelerating anteriorly while the head remained stationary due to its inertia. At 125 ms, the mean posterior head retraction was 89.8 mm for the 2.6 m/s impact velocity and 80.7 mm for the 1.8 m/s impact velocity.

Figure 16.

Head retraction relative to T1 at 1.8 m/s.

Figure 17.

Head retraction relative to T1 at 2.6 m/s.

Posterior retraction of the head relative to T1 occurred in two stages: translation prior to appreciable head rotation (chin-in) and translation coupled with extension rotation. The initial stage of retraction (up to 5 deg head rotation) lasted 31 ms for the 2.6 m/s impact velocity, and 32 ms for the 1.8 m/s impact velocity.

DISCUSSION

Corridors presented in this paper demonstrate the mean kinematic response of 50^th percentile subjects exposed to increasing magnitudes of rear impact loading. Although specimen anthropometry was not representative of the 50^th percentile, responses were scaled to account for this biological variation and represent the 50^th percentile population using an accepted scaling method [Eppinger et al., 1984; Maltese et al., 2002]. Mass scaling of kinematic responses was performed because the corridors were developed to validate rear impact dummies currently used in crashworthiness research, and these dummies typically represent the 50^th percentile geometry. Data presented in this study can also be mass-scaled to develop kinematic corridors to validate the 5^th percentile female and 95^th percentile male dummies also used in rear impact. The scaling technique implemented in the present investigation, based on the mass of the head, was deemed more acceptable than scaling based on other anthropomorphic measurements because spinal kinematics under inertial loading are influenced in large part by the mass of the head [Yoganandan and Pintar, 1997].

In 1992, Penning hypothesized that hyper-translation of the head relative to T1 prior to head rotation may induce injurious loading on the spinal soft tissues. Using lateral radiographs of 10 outpatients with neck complaints, the experimental limit of head translation relative to a fixed T1 vertebra was determined to be 26 mm [Penning, 1992]. Specimens in the present study demonstrated an average of 28.6 mm posterior translation of the head relative to T1 (chin-in) prior to appreciable extension rotation (5 deg) of the head for the 2.6 m/s impact velocity. The fact that this value exceeded the cited limit for the relatively low-impact velocities used in this study underscores the importance of the retraction phase in whiplash injury. It is likely that greater impact velocities result in greater magnitudes of retraction. Therefore, because of the possible importance of head retraction in the whiplash injury mechanism, this measure should be of high priority for validation of dummy kinematics in rear impact.

Kinematic response corridors presented in this paper accounted for the first 125 ms after the initiation of anterior T1 acceleration, encompassing the retraction phase and a portion of the extension phase. The rebound phase, which occurs after contact with the head restraint, is neglected. The timing of head restraint contact depends on the positioning of the head restraint, e.g., height and backset, relative to the head. Head restraint positioning, in turn, is largely dependent upon the curvature of the neck, which differs between dummy models and automotive occupants. Although improperly positioned head restraints may induce hyperextension-type injuries in rear impact, whiplash injuries affecting the soft tissues of the spine and, in particular, the facet joints, most commonly occur during the retraction phase, as the spinal column is subjected to the S-curvature [Ono and Kaneoka, 1997; Cusick et al., 2001; Stemper et al., 2003]. In order to eliminate the variability associated with head restraint positioning and focus validation efforts during the time of injury, kinematic corridors presented in this study are limited to the first 125 ms after the initiation of T1 horizontal acceleration.

A limitation of the present analysis is the fixation of the thoracic vertebrae. In rear impacts, the thoracic interaction with the seatback tends to straighten the thoracic kyphosis and create a ‘ramping’ effect at the cervicothoracic junction, wherein the upper thorax translates superiorly while rotating posteriorly. This phenomenon has been reproduced experimentally using human volunteers and full body cadaver specimens subjected to rear impacts [McConnell et al., 1995; Deng et al., 2000; Yoganandan et al., 2000]. The magnitude of these motions depends on numerous factors including the stiffness of the seatback, initial curvature of the thoracic spine, angle of the seatback, magnitude of rear impact, awareness of the occupant, and occupant posture. Experimental results quantify this variability, where mean T1 rotation at 100 ms for a seatback angle of 20 deg was twice the magnitude of that for a seatback angle of 0 deg [Deng et al., 2000]. The degree to which the thoracic ramping affects cervical spinal kinematics is unclear. In a previous study investigating rear impact kinematics using an intact head-neck cadaver model, it was demonstrated that head to T1 angle for specimens with constrained thorax was not markedly different from data obtained using full body human cadavers and human volunteers subjected to similar rear impact velocities [Stemper et al., 2003]. Therefore, in the present investigation, T1 was constrained against ramping to focus the analysis on sagittal plane kinematics of the cervical spine in rear impact.

The kinematic corridors presented in this manuscript represent a unique focus on the development of experimental corridors for the comprehensive validation of dummy response in rear impacts. Detailed cervical geometry employed in recently developed rear impact dummies permits accurate analysis and validation of dummy response in this loading vector. This data can be used in conjunction with head and thoracic kinematic and kinetic data from human volunteer and full body cadaver experiments to develop and validate biofidelic dummies for rear impact.

CONCLUSIONS

Detailed kinematic response corridors were developed for the validation of dummies in automotive rear impacts. The kinematic responses of intact head-neck complexes were mass-scaled to develop mean and standard deviation response corridors for segmental and head to T1 angulations and head retraction relative to T1 at two impact velocities. Development of the corridors focused on the proposed whiplash injury mechanism (head retraction and cervical S-curvature) and time at which injury most likely occurs. These response corridors were developed to validate existing rear impact dummies (BioRID, RID, TRID) and aid in the development of new dummies.

Figure 5.

C3–C4 segmental angulation at 1.8 m/s.

Figure 6.

C4–C5 segmental angulation at 1.8 m/s.

Figure 7.

C5–C6 segmental angulation at 1.8 m/s.

Figure 10.

C3–C4 segmental angulation at 2.6 m/s.

Figure 11.

C4–C5 segmental angulation at 2.6 m/s.

Figure 12.

C5–C6 segmental angulation at 2.6 m/s.