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Children riding on the bed over the cab in campers can be injured in forward force collisions from striking the glazing material and/or being ejected through the opening. The two types of glazing commonly used are acrylic and laminated. A comparison of the performance of the two types of glazing in simulated forward force collisions at velocities up to 30 mph showed the acrylic material to pose threats of neck and back injury and the laminated material to result in lacerations. Ejections occurred with the acrylic that were not present with the laminated windshields when correct glazing techniques were used. With poor installation procedures, ejections occurred in both types of glazing materials. It is concluded that the best way to avoid injury is to prevent the child from riding in the over-the-cab bunk. If the child does ride there, his body axis should be positioned at an angle to the longitudinal axis of the vehicle.

A review of the existing mathematical models of a car occupant in a rear-end crash reveals that existing models inadequately describe the kinematics of the occupant and cannot demonstrate the injury mechanisms involved. Most models concentrate on head and neck motion and have neglected to study the interaction of the occupant with the seat back, seat cushion, and restraint systems. Major deficiencies are the inability to simulate the torso sliding up the seat back and the absence of the thoracic and lumbar spine as deformable, load transmitting members. The paper shows the results of a 78 degree-of-freedom model of the spine, head, and pelvis which has already been validated in +Gz and -Gx acceleration directions. It considers automotive-type restraint systems, seat back, and seat cushions, and the torso is free to slide up the seat back.

A methodical investigation and measurement of human dynamic response to impact acceleration is being conducted as a Joint Army-Navy-Wayne State University investigation. Details of the experimental design were presented at the Twelfth Stapp Car Crash Conference in October 1968. Linear accelerations are being measured on the top of the head, at the mouth, and at the base of the neck. Angular velocity is also being measured at the base of the neck and at the mouth. A redundant photographic system is being used for validation. All data are collected in computer compatible format and data processing is by digital computer. Selected data in a stage of interim analysis on 18 representative human runs of the 236 human runs completed to date are presented. Review of the data indicates that peak accelerations measured at the mouth are higher than previous estimates.

It is shown that the response of the occiput of a cadaver to sinusoidal vibration input to the frontal bone corresponds closely to that of a simple damped spring-mass system having a natural frequency equal to the first mode frequency of the skull, 0.17 damping factor. The first and third bending mode of the skull occurred near 300 and 900 Hz for both the cadaver preparation with silicon gel filled cranial cavity and the live human head. A second mode was found near 600 Hz in the live human. Head acceleration levels at which opposite pole pressure reached near —1 atm were 170 g and 500–600 g in the human cadaver and live monkey head, respectively, which values are roughly inversely proportional to major intracranial diameters. A method is derived for comparing the impact response of a simple system to a general shaped pulse to that of the cadaver head.

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.