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The two main motivations for Wayne State University (WSU) and Henry Ford Hospital (HFH) researchers to develop numerical human surrogates are advanced computing technology and a high-speed x-ray imaging device not available just a decade ago. This paper summarizes the capabilities and limitations of detailed component models of the human body, from head to foot, developed at WSU over the last decade (Zhang et al. 2001, Yang et al. 1998, Shah et al. 2001, Iwamoto et al. 2000, Lee et al. 2001 and Beillas et al. 2001). All of these models were validated against global response data obtained from relevant high-speed cadaveric tests. Additionally, some models were also validated against local kinematics of bones or soft tissues obtained using the high-speed x-ray system. All of these models have been scaled to conform to the key dimensions of a 50th percentile male.

Unembalmed cadavers restrained with a three point harness were exposed to a deceleration environment of 20, 30 and 40 mph BEV.* Injuries were tabulated from detailed autopsies. The results Indicate an AIS-1 injury at 25.5 mph, an AIS-2 injury at 31.5 mph and an AIS-3 injury at 34.5 mph. The AIS-3 injury level is recommended as the maximum acceptable injury. The cadavers sustained the same types of injury that have been reported in medical literature including bruises, abrasions, lacerations, fractures and viscera ruptures, but injury severities were greater in the cadavers than in living humans at a given collision severity. Also, there is a wide spread in the degree of injury between cadavers due to differences in age and physical condition. The threshold of cadaver rib fracture is 30 mph and the threshold of cadaver vertebral fracture is between 30 and 40 mph for the environment utilized. More numerous and severe abdominal injuries were observed.

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 combined program of accident investigation, staged collisions, and simulated collisions involving three-point harnessed occupants in frontal force collisions has provided a means of correlating injury with forces and/or other physical parameters associated with the injuries. With a strict screening to ensure complete data on each accident, 128 cases involving 169 occupants at barrier equivalent velocities from 2-53 mph were compared with the results from 11 staged collisions and 72 simulated collisions. There were 14 rib cage injuries ranging from single sternum fracture to seven rib fractures at velocities of 10-53 mph at injury levels of AIS 2 and 3. A single AIS 4 injury was the most serious injury and consisted of a ruptured spleen. The most serious brain injury was an AIS 2. Two cervical vertebra fractures were found, one of which was a 12-year-old male and the other a 76-year-old female. Only 14 occupants had AIS 3 injuries.

Five different restraint systems-mandatory harness, airbag + 20% lap belt usage, airbag, passive three point harness, and torso and knee bar-are analyzed for fatality and injury reduction, benefit/cost ratio, and cost-effectiveness. The mandatory harness is superior to the others in all comparisons with approximately 100,000 lives saved over the first 10 years which is about twice as many as would be saved by the other systems. A major advantage of the mandatory harness is that practically all of the vehicles are equipped while the other systems will require 10 years for complete installation.

Spinal kinematics of the living human volunteers undergoing -Gx impact acceleration are described along with the experimental procedures followed to acquire such data. There were 4 male and 3 female volunteers who were subjected to impacts in the tensed and relaxed mode from 2 - 8 g, in 1-g increments. Their lower extremities were tightly clamped to the impact seat and the pelvis was restrained by a lapbelt. The biodynamic response of the living spine is quite similar to that of the cadaveric spine, particularly in terms of T1 displacement, acceleration at T1 and flexural resistance. Female volunteers tend to withdraw from the test program at lower g-levels than males due to transient neck pain.

A six-axis femoral load cell was designed, fabricated, calibrated and tested in an ATD upper leg. It has linear characteristics with adequate output and dynamic response. Its output compared well with that of load cells mounted behind a knee bolster which was impacted by an ATD instrumented with this load cell. Although the axial load was found to be well below the present limit of FMVSS208, the bending moments were found to be high. A conservative tolerance limit for axial load should be considered and cadaver data are required.

Conventional 30 mil HPR laminated and wide-zone monolithic tempered windshields are compared on a safety performance basis from the stand-points of occupant injuries from frontal force collisions and injury or loss of control from breakage from high speed external impact of stones. All experiments were conducted with the windshields installed by conventional methods in an automobile. Occupant injury potential as measured by the Severity Index for brain damage at a 30 mph barrier impact simulation was approximately two times as high for the tempered as for the laminated windshields, although only one tempered windshield exceeded the recommended maximum value of 1,000. Severe lacerations resulted in all impacts in which the tempered glass broke. Less severe lacerations were found for the laminated windshield impacts at comparable speeds.

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.

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.

Unembalmed cadavers were exposed to −Gx acceleration while restrained by applying a frontal load to the chest. A pre-deployed non-venting production air cushion mounted on a non-collapsible horizontal steering column provided the distributed load. The sled deceleration pulse was determined from a series of Part 572 dummy runs in which the HIC, chest acceleration and knee loads were at but not in excess of the limits specified in the current FMVSS 208. A total of six cadavers have been tested. In three of the runs, there were severe neck injuries of the type which have not been observed previously in belted tests. They include complete severance of the cord, complete avulsion of the odontoid process, atlanto-occipital separation with ring fracture. This study does not claim to establish the injury potential of air bags but uses the air bag to provide a uniform restraining load to the chest to investigate the mechanism of neck injuries.

A new laminated automobile windshield called Triplex “Ten-Twenty,” fabricated from two thermally stressed glass plies of 2.3 mm soda-lime float glass laminated with a 0.76 mm HPR polyvinyl butyral interlayer, has been biomechanically evaluated by Triplex Safety Glass Co., Ltd., using a dropping headform and a skull impactor, and by Wayne State University, using a 50th percentile anthropomorphic dummy on the WHAM III sled test facility. The results of these evaluations at velocities up to 60 km/h are expressed in terms of Gadd index, head injury criterion, and various laceration scales including the new Triplex laceration index (TLI). Some details are also given of other properties of the windshield. The results of the evaluations indicate that the Ten-Twenty windshield offers a reduction of about two units on the TLI scale equivalent to one of the following: 1. A 99% reduction in the number of cuts when the length and depth of cuts remain unaltered. 2.

Knee impacts along the femoral axis of unembalmed male cadavers and Part 572 dummies were made with rigid pendulum impactors at Wayne State University. The dummy exhibited significantly higher knee impact forces than the cadaver subjects. This difference of response is shown to be due to differences of effective leg mass and knee padding. The dummy with its heavy rigid metal skeleton is not like its human counterpart, where the majority of the leg weight is composed of loosely coupled flesh. The knee impacts of the dummy subjects showed that the dummy femur transducer force was consistently less than the corresponding dummy knee impact force by a constant ratio of 0.8. We recommend that the “skeletal” weight of the Part 572 dummy leg should be substantially reduced, with the weight difference being added to a properly simulated leg flesh. Also, the simulated flesh covering of the knee should be modified to reduce the peak force resulting from rigid body impacts.

An improved windshield with a special, thin, plastic inner surface attached to the inner surface of a three layer windshield similar to those used in the United States minimizes lacerations from occupant impact to the windshield during a collision. The plastic coats the sharp edges of the broken glass preventing or minimizing laceration. It was evaluated by comparing its laceration performance with that of a standard windshield in simulated barrier crashes at velocities up to 65 km/h. No lacerations resulted from impact to the Securiflex windshield at Barrier Equivalent Velocities up to 65 km/h. Substantial laceration resulted at velocities above 20 km/h with the standard windshield. It is concluded that the Securiflex windshield essentially eliminates lacerations in the particular vehicle involved at velocities up to at least 65 km/h.

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 biodynamic model of the spine simulated the action of spinal musculature on the head, vertebral bodies and pelvis in the midsagittal plane. Muscle was treated as a force generator whose contractile force was dependant on muscle stretch, stretch rate and neural delay time. Eight model runs were conducted with and without muscle, simulating +Gz and -Gx impact acceleration. The model predicted that spinal musculature was incapable of affecting overall spinal column kinematics. However, as a result of muscle contraction, significantly higher local axial forces were predicted in the discs and facets than were predicted when muscle was absent.

The paper gives an evaluation of the performance of lap and shoulder belt restraint systems currently being used in American-built automobiles. Comparisons are made of the response characteristics of a volunteer, an anthropometric dummy, and a cadaver when subjected to identical collision environments while wearing a three or four point torso restraint system as occupants of the right front seat. Simulated frontal force barrier collisions in a modified automobile provided the realistic environment for the restraint system performance study. Human tolerances, interior vehicle geometry, and the interaction of the restrained occupant with the vehicle during the collision are reported in detail.

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.

The biodynamic response of cadaver torsos subjected to -Gx impact acceleration is discussed in this paper, with particular emphasis on the response of the vertebral column. The existence of an axial force along the spine and its manifestation as a load on the seat pan are reported. Spinal curvature appears to be an important factor in the generation of this spine load. In anthropometric dummies, the spine load does not exist. Details of the testing and results are given, and the development of a mathematical model is shown.

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.