Search Results

This paper describes the design and development of the Hybrid III 10-year-old crash test dummy. The size of the dummy was chosen to fill the gap between the Hybrid III 6-year-old and the Hybrid III small adult female dummy which is also about the size of a 13-year-old teenager. Characteristic dimensions and segment weights of the dummy are based on the anthropometry of the average 10-year-old. Biofidelity response guidelines for forehead, sternum and knee impacts and for fore/aft neck bending are scaled from the midsize adult male biofidelity guidelines taking into account the effects of differences in size, mass and material properties due to the age difference. The dummy is similar in construction to the other Hybrid III dummies except it has an adjustable lumbar spine which allows the dummy to slouch and its neck structure is aligned with its thoracic spine. Data are given showing the responses of the prototype dummy relative to its biofidelity guidelines.

A family of adult and child size dummies was developed under the direction of two task groups of the SAE Mechanical Human Simulation Subcommittee of the Human Biomechanics and Simulation Standards Committee. These new child size dummies represent fiftieth percentile children who are 6 months, 12 months, 18 months, 3 years, and 6 years old. The sizes and total body weights of the dummies were based on detailed anthropometry studies of children of these ages. The techniques used to establish the segment masses and the resulting design goals are detailed. Appropriate impact response requirements were scaled from the biofidelity response requirements of the Hybrid III, taking into account the differences in size, mass and elastic modulus of bone between adults and children. The techniques used to establish the biomechanical impact response requirements for the child dummies are discussed and the resulting biomechanical impact response requirements are given.

The development of the Head Injury Risk Curve (HIRC) and the Skull Fracture Risk Curves (SFRC) that were proposed by Prasad and Mertz for the adult driving population are reviewed, and the problem with using the Maximum Likelihood method to analyze the cadaver 15ms HIC data is discussed. The cadaver data base for skull fracture is expanded by including non-fracture 15ms HIC values for a number of cadavers which had skull fractures at higher impact severities and by including cadaver test results of Ono and Tarriere. This expanded data set was analyzed using the Mertz/Weber method and a new method called “Certainty Grouping”. An updated version of the Skull Fracture Risk Curve (SFRC) is proposed. The efficacy of this revised curve is demonstrated by comparing its predictions to the results of simulated fracture impacts using a finite element model of the head. The HIRC was not changed since no additional brain damage data were analyzed.

This paper summarizes the rationale used to specify the geometric, inertial and impact response requirements for a small adult female dummy and a large adult male dummy with impact biofidelity and measurement capacity comparable to the Hybrid III dummy, the most advanced midsize adult male dummy. Body segment lengths and weights for these two dummies were based on the latest anthropometry studies for the extremes of the U.S.A. adult population. Other characteristic body segment dimensions were calculated from geometric and mass scaling relationships that assured that each body segment had the same mass density as the corresponding body segment of the Hybrid III dummy. The biomechanical impact response requirements for the head, neck, chest and knee of the Hybrid III dummy were scaled to give corresponding biomechanical impact response requirements for each dummy.

This paper describes a study to determine the influence of preimpact vehicle braking on the positions and postures of unrestrained, children in the front seat at the time of collision. Anesthetized baboons were used as child surrogates. The unrestrained animals were placed in various initial sitting, kneeling, and standing positions typically assumed by children while traveling in automobiles. Tests were conducted with various front seat positions and seat covering materials. Measurements were made of pertinent vehicle dynamics and surrogate kinematics during the hard braking event. For each initial condition evaluated, a photosequence is given showing typical positions and postures of the surrogate during the braking event.

This paper summarizes the results of tests conducted with anesthetized animals that were exposed to a wide range of passenger inflatable restraint cushion forces for a variety of impact sled - simulated accident conditions. The test configurations and inflatable restraint system concepts were selected to produce a broad spectrum of injury types and severities to the major organs of the head, neck and torso of the animals. These data were needed to interpret the significance of the responses of an instrumented child dummy that was being used to evaluate child injury potential of the passenger inflatable restraint system being developed by General Motors Corporation. Injuries ranging from no injury to fatal were observed for the head, neck and abdomen regions. Thoracic injuries ranged from no injury to critical, survival uncertain.

A primary concern in the development of a passenger inflatable restraint system is the possibility that a child could be in the path of the deploying cushion either due to initial position at the time of an accident or due to precrash braking accompanying an accident. Previous studies by General Motors and Volvo have indicated that serious injuries to children are possible if the cushion/child interaction forces are not controlled by system design. This paper describes an instrumented child dummy which was developed to provide measurements of the various cushion/child interaction forces. An analysis is given describing the types of injuries which could be associated with the various types of interaction forces. These results were used to develop appropriate dummy instrumentation for indicating the severity of the cushion/child interaction. A description of the modifications made to an existing three-year-old child dummy are described.

An analysis is presented that was used to interpret the significance of response measurements made with a specially instrumented, 3-year-old child dummy that was used to evaluate child injury potential of the second-generation, passenger inflatable restraint system that was being developed by General Motors Corporation. Anesthetized animals and a specially instrumented child dummy, both 3-year-old child surrogates, were exposed to similar inflating-cushion, simulated collision environments. The exposure environments were chosen to produce a wide spectrum of animal injury types and severities, and a corresponding broad range of child dummy responses. For a given exposure environment, the animal injury severity ratings for the head, neck, thorax and abdomen are paired with dummy response values corresponding to these body regions.

A viscoelastic neck structure that responds to impact environments in a manner similar to the human neck is described. The neck structure consists of four ball-jointed segments and one pin-connected “nodding” segment with viscoelastic resistive elements inserted between segments that provide bending resistance as well as the required energy dissipation. Primary emphasis was placed on developing appropriate flexion and extension responses with secondary emphasis placed on axial, lateral, and rotational characteristics. The methods used to design the resistance elements for the neck structure are discussed. Three variations of the resistive elements have been developed that meet the response characteristics based on the data of Mertz and Patrick. However, no single resistive element has satisfied the flexion and extension characteristics simultaneously, but such an element appears to be feasible.

Human volunteers were subjected to static and dynamic environments which produced noninjurious neck responses for neck extension and flexion. Cadavers were used to extend this data into the injury region. Analysis of the data from volunteer and cadaver experiments indicates that equivalent moment at the occipital condyles is the critical injury parameter in extension and in flexion. Static voluntary levels of 17.5 ft lb in extension and 26 ft lb in flexion were attained. A maximum dynamic value of 35 ft lb in extension was reached without injury. In hyperflexion, the chin-chest reaction changes the loading condition at the occipital condyles which resulted in a maximum equivalent moment of 65 ft lb without injury. Noninjurious neck shear and axial forces of 190 lb and 250 lb are recommended based on the static strength data obtained on the volunteers. Neck response envelopes for performance of mechanical necks are given for the extension and flexion modes of the neck.

Human tolerance to knee, chest, and head impacts based upon skeletal fracture of cadavers is reported. The results are based upon unrestrained cadaver impacts in a normal seated position in simulated frontal force accidents at velocities between 10 and 20 mph and stopping distances of 6-8 in. The head target was covered with 15/16 in. of padding. No skull or facial fractures were observed at loads up to 2640 lb. Extensive facial fractures and a linear skull fracture occurred during the application of the maximum head force of 4350 lb. The chest target was 6 in. in diameter with 15/16 in.of padding. The padding was rolled over the edge of the target to minimize localized high force areas on the ribs. A 1/8 in. diameter rod was inserted through the chest and fastened through a ball joint and flange to the soft tissue at the sternum.

The kinematics of rear-end collisions based on published acceleration pulses of actual car-to-car collisions (10 and 23 mph) were reproduced on a crash simulator using anthropomorphic dummies, human cadavers, and a volunteer. Comparison of the responses of subjects without head support were based on the reactions developed at the base of the skull (occipital condyles). The cadavers gave responses which were representative of persons unaware of an impending collision. The responses of both dummies used were not comparable with those of the cadavers or volunteer, or to each other. An index based on voluntary human tolerance limits to statically applied head loads was developed and used to determine the severity of the simulations for the unsupported head cases. Results indicated that head torque rather than neck shear or axial forces is the major factor in producing neck injury.

A comparison is presented of the forces, accelerations, and kinematics of an anthropomorphic dummy for identical sled impacts for unrestrained, lap belted, and lap and diagonal chest restrained conditions. Biaxial accelerometers were mounted in the head, chest, and on the proximal end of the femur to obtain the accelerations during the impacts. Seat belt load cells were put in series with the belts at each anchor point. Biaxial load cells were positioned to be impacted by the head, chest, and each knee for the unrestrained condition and by the head and chest for the lap belted configuration. For the lap and diagonal chest restrained condition these load cells were not used. Impacts of 10 and 20 miles per hour were made with sled stopping distance of 4 and 9 inches, respectively. At 20 miles per hour the head struck with a force of 1580 pounds in the unrestrained mode, 600 pounds with the lap belt, and did not hit with the lap and shoulder harness.