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A PC based interactive program was developed to simulate the unfolding and deploying process of a driver side airbag in the sagittal plane. The airbag was represented by a series of nodes. The maximum allowable stretch was less or equal to one between any two nodes. We assumed that the airbag unfolding was pivoted about folded points. After the completion of the unfolding process the airbag would begin to deploy. During the deploying process, two parameters were used to determine the nodal priority of the inflation. The first parameter was the distance between the instantaneous and final positions of a node. Nodes with longer distances to travel will have to move faster. We also considered the distance between the current nodal position and the gas inlet location. For a node closer to the gas inlet, we assumed that the deploying speed was faster. A graphical procedure was used to calculate the area of the airbag.

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.

Lower extremity injuries in frontal automotive crashes usually occur with footwell intrusion where both the knee and foot are constrained. In order to identify factors associated with tibial shaft injury, a series of numerical simulations were conducted using a finite element model of the whole human body. These simulations demonstrated that tibial mid-shaft injuries in frontal crashes could be caused by an abrupt change in velocity and a high rate of footwell intrusion.

A nonlinear foam was added to a previously created three-dimensional finite element model of the Hybrid III dummy chest which consisted of six steel ribs, rib damping material, the sternum, a spine box and a pendulum. Two standard calibration pendulum impact tests for a Hybrid III dummy chest were used to validate the new model. An explicit finite element analysis code PAM-CRASH was utilized to simulate the dynamic process. At impact velocities of 6.7 m/s and 4.3 m/s, the force and deflection time history as well as the force-deflection plots showed good agreement between model predictions and calibration data. Peak strains also agreed well with experimental data.

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 mechanisms of knee fracture were studied experimentally using cadaveric knees and analytically by computer simulation. Ten 90 degree flexed knees were impacted frontally by a 20 kg pendulum with a rigid surface, a 450 psi (3.103 MPa) crush strength and a 100 psi (0.689 MPa) crush strength aluminum honeycomb padding and a 50 psi (0.345 MPa) crush strength paper honeycomb padding at a velocity of about five m/s. During rigid surface impact, a patella fracture and a split condylar fracture were observed. The split condylar fracture was generated by the patella pushing the condyles apart, based on a finite element model using the maximum principal stress as the injury criterion. In the case of the 450 psi aluminum honeycomb padding, the split condylar fracture still occurred, but no patella fractures were observed because the honeycomb provided a more uniform distribution of patella load. No bony fractures in the knee area occurred for impacts with a 50 psi paper honeycomb padding.

Previously, a 10-year-old (YO) pediatric thorax finite element model (FEM) was developed and verified against child chest stiffness data measured from clinical cardiopulmonary resuscitation (CPR). However, the CPR experiments were performed at relatively low speeds, with a maximum loading rate of 250 mm/s. Studies showed that the biomechanical responses of human thorax exhibited rate sensitive characteristics. As such, the studies of dynamic responses of the pediatric thorax FEM are needed. Experimental pediatric cadaver data in frontal pendulum impacts and diagonal belt dynamic loading tests were used for dynamic validation. Thoracic force-deflection curves between test and simulation were compared. Strains predicted by the FEM and the injuries observed in the cadaver tests were also compared for injury assessment and analysis. This study helped to further improve the 10 YO pediatric thorax FEM.

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.

A finite element model of the human lower extremity has been developed to predict lower extremity injuries in full frontal and offset frontal impact. The model included 30bones from femur to toes. Each bone was modeled using crushable solid elements for the orbicular bone and damageable shell elements for the cortical bone. The models of the long bones for the lower extremities were validated against data obtained from quasi-static 3-pointbending tests by Yamada (1970). The ankle, knee and hip joints were modeled as bone-to-bone contacts and included major ligaments and tendons. The ankle model was validated against data obtained from quasi-staticdorsiflexion, inversion and eversion tests by Petit et al. (1996) and against data obtained from dynamic impactcadaveric tests by Kitagawa et al. (1998). The possibility of using this model to predict injuries was discussed.

An understanding of stiffness characteristics of different body regions, such as thorax, abdomen and pelvis of ES-2re and SID-IIs dummies under controlled laboratory test conditions is essential for development of both compatible performance targets for countermeasures and occupant protection strategies to meet the recently updated FMVSS214, LINCAP and IIHS Dynamic Side Impact Test requirements. The primary purpose of this study is to determine the transfer functions between the ES-2re and SID-IIs dummies for different body regions under identical test conditions using flat rigid wall sled tests. The experimental set-up consists of a flat rigid wall with five instrumented load-wall plates aligned with dummy’s shoulder, thorax, abdomen, pelvis and femur/knee impacting a stationary dummy seated on a rigid low friction seat at a pre-determined velocity.

The main purpose of this study was to determine the impact responses of the different body regions (shoulder, thorax, abdomen and pelvis/leg) of the ES-2re and SID-IIs dummies using rigid wall impacts under different initial test conditions. The experimental set-up consisted of a flat rigid wall with five instrumented load-wall plates aligned with dummy’s shoulder, thorax, abdomen, pelvis and knee impacting a stationary dummy seated on a rigid seat at a pre-determined velocity. The relative location and orientation of the load-wall plates was adjusted relative to the body regions of the ES-2re and SID-IIs dummies respectively.

The main purpose of this study was to determine the impact responses of the different body regions (shoulder, thorax, abdomen and pelvis/leg) of the ES-2re and SID-IIs dummies using rigid wall impacts under different initial test conditions. The experimental set-up consisted of a flat rigid wall with five instrumented load-wall plates aligned with dummy’s shoulder, thorax, abdomen, pelvis and knee impacting a stationary dummy seated on a rigid seat at a pre-determined velocity. The relative location and orientation of the load-wall plates was adjusted relative to the body regions of the ES-2re and SID-IIs dummies respectively.

Aortic injuries during blunt thoracic impacts can lead to life threatening hemorrhagic shock and potential exsanguination. Experimental approaches designed to study the mechanism of aortic rupture such as the testing of cadavers is not only expensive and time consuming, but has also been relatively unsuccessful. The objective of this study was to develop a computer model and to use it to predict modes of loading that are most likely to produce aortic ruptures. Previously, a 3D finite element model of the human thorax was developed and validated against data obtained from lateral pendulum tests. The model included a detailed description of the heart, lungs, rib cage, sternum, spine, diaphragm, major blood vessels and intercostal muscles. However, the aorta was modeled as a hollow tube using shell elements with no fluid within, and its material properties were assumed to be linear and isotropic.

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.

Currently, three-dimensional finite element models of the human body have been developed for frequently injured anatomical regions such as the brain, chest, extremities and pelvis. While a few models of the human body include the abdomen, these models have tended to oversimplify the complexity of the abdominal region. As the first step in understanding abdominal injuries via numerical methods, a 3D finite element model of a 50th percentile male human abdomen (WSUHAM) has been developed and validated against experimental data obtained from two sets of side impact tests and a series of frontal impact tests. The model includes a detailed representation of the liver, spleen, kidneys, spine, skin and major blood vessels.

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.