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This study examines the static pressure distribution under both width belts in the shoulder and the pelvis of 15 volunteer subjects. The subjects applied the belt loads to themselves through a lever and pulley system. The configuration of the belts simulated the typical arrangement of a six-point belted upright-seated racing driver. The pressure distribution between the belt and the volunteer's body was determined and recorded with Tek-Scan pressure sensing grids. The paper presents the results of the measurements by comparing the actual area of significant loading beneath the two widths and materials of both lap and shoulder belts. In, general, there no significant increase in loaded area for the wider belts.

The heads of auto racing drivers and military pilots are usually not supported so that neck fatigue and injury can be a serious problem. A new Head And Neck Support (HANS) is being developed to reduce head motions and neck loads. The biomechanical performance of HANS has been evaluated by crash victim modeling with CAL 3-D and by impact sled testing with a Hybrid III dummy. Modeling and testing were conducted at 30 and 35 mph BEV and with acceleration directions from the front, right front, and right lateral. The model and test results show that head motions, neck loading, and the potential for neck injury are all significantly reduced with HANS compared to without HANS.

Studies have shown that dummies can be used to study various issues relating to an unrestrained driver's interaction with the steering system in frontal crashes. However, current dummies have limitations in simulation of car occupants and to assess the spectrum of injury types and mechanisms. Human cadaver subjects were used to study abdominal injury and “severe” steering wheel deformation as part of an evaluation of energy absorbing steering systems. A predominant factor influencing abdominal injury in these tests was the impact location of the lower rim, injury being associated with the rim aligned 50 mm below the xiphoid. The dummies developed approximately twice the impact force than the cadaver subjects in these severe tests with a noncompressible column, in part due to the chest of the dummies “bottoming” out on a rigid spine.

This paper builds on previous research on the development of a head displacement model for restrained occupants in frontal collisions. Physical and mathematical simulations were performed utilizing the 5th percentile female and 50th percentile male Hybrid III dummies to measure the effect of occupant size, seat belt system design and crash severity on resultant head displacement of occupants in frontal collisions. Sled and simulation accelerations ranged from 10 g to 20 g with delta-V's from 6.6 m/s to 10.0 m/s. Results indicate a difference between the 5th percentile female and 50th percentile male dummies. Preliminary assessment of head displacement as a function of occupant kinetic energy demonstrated good correlation.

A three-dimensional finite element model of a human neck has been developed in an effort to study the mechanics of cervical spine while subjected to impacts. The neck geometry was obtained from MRI scans of a 50th percentile male volunteer. This model, consisting of the vertebrae from C1 through T1 including the intervertebral discs and posterior elements, was constructed primarily of 8-node brick elements. The vertebrae were modeled using linear elastic-plastic materials, while the intervertebral discs were modeled using linear viscoelastic materials. Sliding interfaces were defined to simulate the motion of synovial facet joints. Anterior and posterior longitudinal ligaments, facet joint capsular ligaments, alar ligaments, transverse ligaments, and anterior and posterior atlanto-occipital membranes were modeled as nonlinear bar elements or as tension-only membrane elements. A previously developed head and brain model was also incorporated.

In a frontal automotive crash, the driver's foot is usually stepping on the brake pedal as an instinctive response to avoid a collision. The tensile force generated in the Achilles tendon produces a compressive preload on the tibia. If there is intrusion of the toe board after the crash, an additional external force is applied to the driver's foot. A series of dynamic impact tests using human cadaveric specimens was conducted to investigate the combined effect of muscle preloading and external force. A constant tendon force was applied to the calcaneus while an external impact force was applied to the forefoot by a rigid pendulum. Preloading the tibia significantly increased the tibial axial force and the combination of these forces resulted in five tibial pylon fractures out of sixteen specimens.

A finite element human model has been developed to simulate occupant behavior and to estimate injuries in real-world car crashes. The model represents an average adult male of the US population in a driving posture. Physical geometry, mechanical characteristics and joint structures were replicated as precise as possible. The total number of nodes and materials is around 67,000 and 1,000 respectively. Each part of the model was not only validated against human test data in the literature but also for realistic loading conditions. Additional tests were newly conducted to reproduce realistic loading to human subjects. A data set obtained in human volunteer tests was used for validating the neck part. The head-neck kinematics and responses in low-speed rear impacts were compared between the measured and calculated results. The validity of the lower extremity part was examined by comparing the tibia force in a foot impact between the test data and simulation results.

Recent action by some racecar sanctioning bodies making head/neck restraint use mandatory for competitors has resulted in a number of methods attempting to provide head/neck restraint. This paper evaluates the performance of a number of commercially available head/neck restraint systems using a stock car seating configuration and a realistic stock car crash pulse. The tests were conducted at an impact angle of 30 degrees to the right, with a midsize male Hybrid III anthropomorphic test device (ATD) modified for racecar crash testing. A six-point latch and link racing harness restrained the ATD. The goal of the tests was to examine the performance of the head/neck restraint without the influence of the seat or steering wheel. Three head/neck restraint systems were tested using a sled pulse with a 35 mph (56 km/h) velocity change and 50G peak deceleration. Three tests with three samples of each system were performed to assess repeatability.

This paper analyzed the mechanisms of injury in high speed, right-lateral impacts of stock car auto racing, and interaction of the occupant and the seat system for the purpose of reducing the risk of injury, primarily rib fractures. Many safety improvements have been made to stock car racing recently, including the Head and Neck Support devices (HANS®), the 6-point restraint harnesses, and the implementation of the SAFER Barrier. These improvements have contributed greatly to mitigating injury during the race crash event. However, there is still potential to improve the seat structure and the understanding of the interaction between the driver and the seat in the continuation of making racing safety improvements. This is particularly true in the case of right-lateral impacts where the primary interaction is between the seat supports and the driver and where the chest is the primary region of injury.

Auto racing has been in vogue from the time automobiles were first built. With the dawn of modern cars came higher engine capacities; the speeds involved in these races and crashes increased as well. However, the advent of passive restraint systems such as the helmet, HANS (Head and Neck Support device), multi-point harness system, roll cage, side and frontal crush zones, racing seats, fire retardant suits, and soft-wall technology, have greatly improved the survivability of the drivers in high-speed racing crashes. Three left lateral crashes from Begeman and Melvin (2002), Case #LAS12, #IND14 and #99TX were used as inputs to the Wayne State Human Body Model (WSHBM) in a simulated racing buck. Twelve simulations with delta-v, six-point harness and shoulder pad as design variables were analyzed for the average maximum principal strain (AMPS) in the aorta. The average AMPS for the high-speed crashes were 0.1551±0.0172 while the average maximum pressure was 110.50±4.25 kPa.

The primary purpose of this research was to generate a “linked set” of data between collision severity, occupant weight and collision-induced seat belt markings to assist in reconstruction of motor vehicle accidents. The secondary purpose was to establish a preliminary threshold of belt load to produce known collision-induced seat belt markings. Sled tests were performed utilizing Hybrid III 5th and 50th percentile crash test dummies. Sled accelerations ranged from 10.0 g to 23.6 g and Delta-V’s from 6.4 m/s to 11.3 m/s. Post-test inspections and photographs taken of the seat belts documented collision-induced markings on components such as the D-Ring, latch plate, webbing and stitching. Belt loads were analyzed to establish preliminary thresholds for the production of observable markings.

A humerus provisional reference value (PRV) based on human surrogate data was developed to help evaluate upper arm injury potential. The proposed PRV is based on humerus bone bending moments generated by testing pairs of cadaver arms to fracture in three-point bending on an Instron testing machine in either lateral-medial (L-M) or anterior-posterior (A-P) loading, at 218 mm/s and 0.635 mm/s loading rates. The results were then normalized and scaled to 50th and 5th percentile sized occupants. The normalized average L-M bending moment at failure test result was 6 percent more than the normalized average A-P bending moment. The normalized average L-M shear force at failure was 23 percent higher than the normalized average A-P shear force. The faster rate of loading resulted in a higher average bending moment overall - 8 percent in the L-M and 14 percent in the A-P loading directions.

At the 2002 MSEC, we presented a paper on the sled test evaluation of racecar head/neck restraint performance (Melvin, et al. 2002). Some individuals objected to the 3 msec clip filtering procedures used to eliminate artifactual spikes in the neck tension data for the HANS® device. As a result, we are presenting the same test data with the spikes left in the neck force data to reassure those individuals that these spikes did not significantly affect the results and conclusions of our original paper. In addition we will add new insights into understanding head/neck restraint performance gained during two more years of testing such systems. This paper re-evaluates the performance of three commercially available head/neck restraint systems using a stock car seating configuration and a realistic stock car crash pulse. The tests were conducted at an impact angle of 30 degrees to the right, with a midsize male Hybrid III anthropomorphic test device (ATD) modified for racecar crash testing.

A total of eight low-speed, rear-end impact tests using two Post Mortem Human Subjects (PMHS) in a seated posture are reported. These tests were conducted using a HYGE-style mini-sled. Two test conditions were employed: 8 kph without a headrestraint or 16 kph with a headrestraint. Upper-body kinematics were captured for each test using a combination of transducers and high-speed video. A 3-2-2-2-accelerometer package was used to measure the generalized 3D kinematics of both the head and pelvis. An angular rate sensor and two single-axis linear accelerometers were used to measure angular speed, angular acceleration, and linear acceleration of T1 in the sagittal plane. Two high-speed video cameras were used to track targets rigidly attached to the head, T1, and pelvis. The cervical spine kinematics were captured with a high-speed, biplane x-ray system by tracking radiopaque markers implanted into each cervical vertebra.

Finite element models of human body segments have been developed in recent years. Numerical simulation could be helpful when understanding injury mechanisms and to make injury assessments. In the lower leg injury research in NISSAN, a finite element model of the human ankle/foot is under development. The mesh for the bony part was taken from the original model developed by Beaugonin et al., but was revised by adding soft tissue to reproduce realistic responses. Damping effect in a high speed contact was taken into account by modeling skin and fat in the sole of the foot. The plantar aponeurosis tendon was modeled by nonlinear bar elements connecting the phalanges to the calcaneus. The rigid body connection, which was defined at the toe in the original model for simplicity, was removed and the transverse ligaments were added instead in order to bind the metatarsals and the phalanges. These tendons and ligaments were expected to reproduce a realistic response in compression.

Although various automobile accident surveys showed between 20 to 30% of lower extremity injuries involved the foot or ankle, there is little information in the existing literature on the the injury mechanisms of ankle injuries for automobile occupants involved in frontal impacts. This study addresses the injury to ankles involving dorsiflexion caused by impact loading to the bottom of the foot. Types of injuries include malleolus fractures and ligament avulsions and ruptures. Nine pair of cadaver and two Hybrid 3 lower limbs were impacted on the bottom of the foot with a 16 kg pneumatically propelled linear impactor. A horizontally oriented bar struck the foot 62 mm distally of the ankle joint with velocities between 3 and 8 m/s. The proximal end of the tibia/fibula was fixed to a rigid support through a triaxial load cell. Load cells on the foot and impactor along with high-speed photography provided the response data of the foot and ankle.

The purposes of this study were to measure the relative linear and angular displacements of each pair of adjacent cervical vertebrae and to compute changes in distance between two adjacent facet joint landmarks during low posterior- anterior (+Gx) acceleration without significant hyperextension of the head. A total of twenty-six low speed rear-end impacts were conducted using six postmortem human specimens. Each cadaver was instrumented with two to three neck targets embedded in each cervical vertebra and nine accelerometers on the head. Sequential x-ray images were collected and analyzed. Two seatback orientations were studied. In the global coordinate system, the head, the cervical vertebrae, and the first or second thoracic vertebra (T1 or T2) were in extension during rear-end impacts. The head showed less extension in comparison with the cervical spine.

Previous studies have hypothesized that the shoulder may be used to absorb some impact energy and reduce chest injury due to side impacts. Before this hypothesis can be tested, a good understanding of the injury mechanisms and the kinematics of the shoulder is critical for occupant protection in side impact. However, existing crash dummies and numerical models are not designed to reproduce the kinematics and kinetics of the human shoulder. The purpose of this study was to develop a finite element model of the human shoulder in order to achieve a deeper understanding of the injury mechanisms and the kinematics of the shoulder in side impact. Basic anthropometric data of the human shoulder used to develop the skeletal and muscular portions of this model were taken from commercial data packages. The shoulder model included three bones (the humerus, scapula and clavicle) and major ligaments and muscles around the shoulder.

The Lower Limb Model for Safety (LLMS) is a finite element model of the lower limb developed mainly for safety applications. It is based on a detailed description of the lower limb anatomy derived from CT and MRI scans collected on a subject close to a 50th percentile male. The main anatomical structures from ankle to hip (excluding the hip) were all modeled with deformable elements. The modeling of the foot and ankle region was based on a previous model Beillas et al. (1999) that has been modified. The global validation of the LLMS focused on the response of the isolated lower leg to axial loading, the response of the isolated knee to frontal and lateral impact, and the interaction of the whole model with a Hybrid III model in a sled environment, for a total of nine different set-ups. In order to better characterize the axial behavior of the lower leg, experiments conducted on cadaveric tibia and foot were reanalyzed and experimental corridors were proposed.

Posterior translation of the tibia with respect to the femur can stretch the posterior cruciate ligament (PCL). Fifteen millimeters of relative displacement between the femur and tibia is known as the Injury Assessment Reference Value (IARV) for the PCL injury. Since the anterior protuberance of the tibial plateau can be the first site of contact when the knee is flexed, the knee bolster is generally designed with an inclined surface so as not to directly load the projection in frontal crashes. It should be noted, however, that the initial flexion angle of the occupant knee can vary among individuals and the knee flexion angle can change due to the occupant motion. The behavior of the tibial protuberance related to the knee flexion angle has not been described yet. The instantaneous angle of the knee joint at the timing of restraining the knee should be known to manage the geometry and functions of knee restraint devices.