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This paper discusses the ongoing real-world effects on the wearers of restraint systems which are subject to a retractor's failure to lock in a timely manner. Investigation of the ELR performance using both detailed physical examination and inductive methods enables accurate assessment of successful ELR locking at the first opportunity in the crash sequence. Available methods to determine the reliability of the ELR's crash performance are considered and analyzed for assessment of reliability to enable adequate seat belt wearer protection. Corrective measures are analyzed to probe the feasibility of federal safety regulation amendments to mandate a reliability analysis on the propensity for the ELR's failure to lock in a timely manner.

Actual automobile crash scenarios include “wheels-first” landings after the vehicle leaves the road surface and becomes momentarily airborne. These events generate a vertical acceleration vector in a headward direction (+Gz) along the occupant's spinal axis. In this scenario, the vehicle occupant could be in contact with the seat bottom or seat back cushions, or displaced several inches off both the bottom and/or back cushions depending on the effectiveness of the restraint configuration and the dynamics of the vehicle's motion. Military ejection seat researchers have shown that occupant response to +Gz acceleration loading is amplified as a function of the spring-mass damping characteristics of the total system (i.e., the occupant and seat/restraint/cushion subsystems). This amplification phenomenon, commonly known as “dynamic overshoot”, has the propensity to vary widely depending on the built-in controls within a given seat bottom design.

In this paper human volunteer head/neck dynamic response is compared with that of a Hybrid III head and neck. The data base used for the comparison was taken from the extensive Naval Biodynamics Laboratory data base of human and manikin sled runs for various thrust vector directions. Significant differences were identified in the human volunteer and manikin responses for the −X and +Z runs, while an unexpected similarity was observed in the head trajectories for the +Y and −X+Y runs. A detailed analysis of 15 g −X vector exposures is presented. Math model simulations using a linkage model to simulate the response of the beam-like Hybrid III neck to −X sled acceleration profiles are presented, as well as simulations of human runs using a similar linkage model. Improved simulations were obtained by making changes to the linkage model parameters which could presumably be implemented relatively easily on the physical Hybrid III neck.

The Naval Biodynamics Laboratory (NBDL) has collected a database which describes human dynanic responses for −X acceleration exposures as a function of mass distribution variations of the head. Kinematic responses were measured on subjects with, (a) no mass addition; (b) with a helmet and weight-carrier mass addition; (c) and with the helmet and added weights symmetrically located with respect to the mid-sagittal plane of the head. The total mass addition to the head with the weights was approximately 30 percent. The helmet and weights were positioned with reference to the head anatomical coordinate system for each subject, with mass moments of inertia and variations in center of gravity then being determined. This paper compares responses both as a function of a mass distribution parameters and as a model to simulate the observed responses.

Significant kinematic parameters of the head are compared between a Hybrid II head and neck (per Parr 572, Federal Motor Vehicle Safely Standard 208) and human volunteers subjected to the same -Gx sled acceleration profiles. Comparison time profiles between the dummy and human subjects for component's of linear acceleration, velocity, and displacement of the head center of gravity and the first thoracic vertebral body (T1) anatomical origin, as well as components of angular acceleration, velocity, and displacement of the anatomical coordinate systems are presented for 6, 10, and 15G peak sled acceleration in the -Gx environment. Significant differences between the dummy and human volunteers are discussed in regard to peak values of parameters, time latencies, and overall shape agreement in a time window where motion is significant.

The tolerance to abrupt linear deceleration (-Gx) and impact trauma patterns resulting from the use of the Air Force shoulder harness-lap belt restraint were investigated. Eighty-nine deceleration tests were performed with 37 adult male baboons. Peak sled decelerations ranged from 6.5-134 g. The stopping distance varied from 0.5-3.5 ft at 6 in increments. LD50s were calculated to be 102, 103, and 98 g for the 0.5, 2.0, and 3.5 ft stopping distances, respectively. Since the deceleration pulses were similar, the results imply that for the exposure range of these tests, impact lethality is dependent upon magnitude of peak sled deceleration, irrespective of the pulse duration, sled velocity, or stopping distance. At all stopping distances, the primary cause of death was lower brainstem or cervical spinal cord trauma. The pelvic, abdominal, and thoracic injury patterns were significantly different at the various stopping distances.

Head linear and angular accelerations of humans were investigated during exposure to abrupt linear deceleration (-Gx). The 14 subjects were restrained with three different restraints: lap belt only, Air Force shoulder harness-lap belt and air bag plus lap belt. Peak sled decelerations ranged from 7.7-10.3 g. The results indicated that peak head angular and linear resultant accelerations were elevated with the air bag in contrast to the Air Force shoulder harness or lap belt only restraints. However, the peak angular and linear accelerations may have less traumatic consequences than the degree of head-neck hyperextension.

The tolerance to abrupt linear deceleration (-Gx) and the subject interaction with an air bag plus lap belt and air bag only restraint systems were investigated. Twenty adult male baboons comprised the test pool. Peak sled decelerations ranged 8.6-123 g. The results indicated that the tolerance to impact (LD50) utilizing an air bag with or without lap belt was in excess of 120 g. The severest injuries were attributable to the lap belt, and included rupture of the rectus abdominus and quadriceps femoris muscles plus diaphragmatic tearing. There were no significant injuries to subjects restrained with only an air bag. Excellent linear correlations were established between peak lap belt forces and maximum sled deceleration. Comparative evaluation of the air bag restraint with a previously reported lap belt study was made when applicable.

The tolerance to abrupt linear deceleration (- Gx) and the subject response to a lap belt restraint system were investigated. Nineteen adult male baboons comprised the test pool. The effects of impacts of 8.6-40 g were studied, with nonsurvivability used as the index of tolerance. The results indicated that the tolerance to impact (LD50) approximated a 32 g sled deceleration. Lethality was presumed attributable to the secondary impact as the head contacted the floor of the sled. Predominant lethal injuries included avulsion of the atlanto-occipital articulation and dislocation fractures of the cervical vertebrae with resulting transection of the spinal cord. Excellent linear correlations were established between peak lap belt and seat pan forces versus maximum sled deceleration. Likewise, a linear relationship was found between peak head angular accelerations and maximum sled deceleration.