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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.

Digital human modeling (DHM) progress worldwide will be much faster and cohesive if the diverse community now developing simulations has a global blueprint for DHM, and is able to work together efficiently. DHM developers and users can save time by building on each other's work. This paper highlights a panel discussion on DHM goals and strategic plans for the next decade to begin formulating the international blueprint. Four subjects are chosen as the starting points: (1) moving DHM into the public safety and internet arenas, (2) role of DHM in computer assisted surgery and automotive safety, (3) DHM in defense applications, and (4) DHM to improve workplace ergonomics.

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.

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.

Heidelberg-type side impact sled tests were conducted using SID side impact dummies. These tests were run under similar conditions to a series of cadaveric sled tests funded by the Centers for Disease Control in the same lab. Tests included 6.7 and 9 m/s (15 and 20 mph) unpadded and 9 m/s padded tests. The following padding was used at the thorax: ARSAN, ARCEL, ARPAK, ARPRO, DYTHERM, 103 and 159 kPa (15 and 23 psi) crush strength paper honeycomb, and an expanded polystyrene. In all padded tests the dummy Thoracic Trauma Index, TTI(d) was below the value of 85 set by federal rulemaking (49 CFR, Part 571 et al., 1990). In contrast, cadavers in 9 m/s sled tests did not tolerate ARSAN 601 (MAIS 5) and 23 psi (159 kPa) paper honeycomb (MAIS 5), and 20 psi (138 kPa) Verticel™ honeycomb (MAIS 4), but tolerated 15 psi (103 kPa) paper honeycomb (average thoracic MAIS 2.3 in six tests).

In the automotive seating industry measurements of lumbar support prominence and height are performed to assess their effects on occupant comfort. This project investigated measurement methods for lumbar support prominence and height in an occupied seat. Fifteen participants provided subjective responses of their perceived lumbar support prominence and height utilizing specifically developed visual analog scales. Also, pressure measurements were taken while the participants were seated. The recently developed H-point manikin II was utilized as a standardized sitter. Specifically, the lumbar support prominence (LSP) measure was used for the prominence measures. Pressure mat readings with the seated manikin was used for lumbar support height determination and prominence correlations. With both manikin and participants in the seat, the lumbar support was digitized through the rear of the seat.

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.

Twelve unembalmed human cadavers were tested for lower abdominal injury tolerance and mechanical response. The impacts were in an anterior-to-posterior direction and the level of impact was primarily in the lower abdomen at the L3 level of the lumbar spine. The impactor mass was either 32 kg or 64 kg. The impactor face was a 25 mm diameter aluminum bar, with the long axis of the bar parallel to the width of the cadaver body. In this paper, mechanical response is presented in terms of force-time and penetration-time histories, and force vs. abdominal penetration cross-plots. Injury tolerance is described in terms of post-impact necropsy findings and AIS ratings. Based on our studies, the lower abdomen of the unembalmed human cadaver is much less stiff than is suggested by previous research, and the stiffness is velocity and mass dependent, as is suggested by the correlation coefficients presented in this paper. Force-time history and force-penetration response corridors are presented.

Facial impact experiments were conducted on eleven unembalmed human cadavers. A 32 kg or 64 kg impactor with a 25 mm diameter, rigid, cylindrical contact surface was oriented in the left-right direction relative to the face and contacted the nose at the elevation of the infraorbital margins. The impactor was propelled toward the race along an anterior-to-posterior path, with contact velocities ranging from 10 to 26 km/h. Accelerometers mounted on the impactor and the occiput provided data for analyzing the dynamics of the impacts. While the threshold for nasal bone fractures was not determined, it appears that a peak force of about 3 kN (filtered 180 Hz) is a representative threshold for more severe fracture patterns. A preliminary dynamic force vs penetration response specification for the above mode of loading is offered.

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.

Computational human body models, especially detailed finite element models are suitable for investigation of human body kinematic responses and injury mechanism. A real-world lateral vehicle-tree impact accident was reconstructed by using finite element method according to the accident description in the CIREN database. At first, a baseline vehicle FE model was modified and validated according to the NCAP lateral impact test. The interaction between the car and the tree in the accident was simulated using LS-Dyna software. Parameters that affect the simulation results, such as the initial pre-crash speed, impact direction, and the initial impact location on the vehicle, were analyzed. The parameters were determined by matching the simulated vehicle body deformations and kinematics to the accident reports.

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.

Lateral impact testing has been performed on the shoulder, thorax, abdomen and pelvis of human cadavers by several investigators. The impacts have either been whole body impacts in sled tests or pendulum type impacts to the separate regions. Based on the forces produced in these tests and the accompanying injury, initial recommendations can be made on force-tolerance and padding tolerance to the various regions of the human body in side impact. The pelvis has the highest force tolerance, followed by the shoulder, abdomen and thorax. Padding crush strength tolerance based on these forces and estimated contact areas are presented. This information is of practical importance to engineers who design door interior trim for side impact safety.

Seventeen Heidelberg type cadaveric side impact sled tests, two sled-to-sled tests, and forty-four pendulum tests have been conducted at Wayne State University, to determine human responses and tolerances in lateral collisions. This paper describes the development of a simplified finite element model of a human occupant in a side impact configuration to simulate those cadaveric experiments. The twelve ribs were modeled by shell elements. The visceral contents were modeled as an elastic solid accompanied by an array of discrete dampers. Bone condition factors were obtained after autopsy to provide material properties for the model. The major parameters used for comparison are contact forces at the level of shoulder, thorax, abdomen and pelvis, lateral accelerations of ribs 4 and 8 and of T12, thoracic compression and injury functions V*C, TTI and ASA.