Passengers under age 8 must be secured in an appropriate child restraint system as covered by the Child Passenger Protection Act. When riding in a truck with only a front seat equipped with safety belts, a child under age 8 must be secured in an appropriate child restraint system.
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- Driver Restraint System
- Children under age 8 and under 57” must be properly secured in a child restraint system. All other children age 8 to 16 or over 57″ must be secured in a properly fitted seat belt. Children under age 2, and under 40 lbs. Or under 40”, must be properly secured in a rear-facing child restraint system.
- This report examines factors associated with passenger vehicle driver restraint use status in fatal crashes. Given that seat belt use is an important countermeasure against injury or death in vehicle crashes, the.
Abstract
Restrained driver and right-front passenger kinematics and injury outcome in frontal collisions are compared using FARS data and human cadaver sled tests. The FARS data indicate that a frontal airbag may provide greater benefit for a passenger than for a driver. The thoracic injuries sustained by passenger subjects restrained by a force-limited, pretensioned belt and airbag are evaluated, and kinematics are compared to driver-side subjects. The injury-predictive ability of existing thoracic injury criteria is evaluated for passenger-side occupants. Driver and passenger kinematic differences are identified and the implications are discussed. The chest acceleration of the passenger-side subjects exhibited a bimodal profile with an initial (and global) maximum before the subject loaded the airbag. A second acceleration peak occurred as the subject loaded both the belt and the airbag. A similarly restrained driver-side subject loaded the belt and airbag concurrently at the time of peak chest acceleration and therefore did not exhibit this biomodal chest acceleration.
Non-drivers represent a significant portion of the fatalities and severe injuries that occur in automobile collisions. According to the Fatality Analysis Reporting System (FARS) for 1998, drivers accounted for 14,895, or approximately 74% of all passenger-vehicle frontal impact fatalities, while other occupants, mostly in the right-front seat position (Figure 1), accounted for 5,248 fatalities. This uneven fatality distribution is due largely to the higher crash exposure of drivers - the Texas Transportation Institute [Christiansen 1990], for example, estimated for urban work trips an average of only about 1.2 persons in each vehicle - though impact direction, offset, intrusion, and population age, size, and gender differences also influence the relative fatality risk.
Frontal impact fatalities by position in vehicle.
found right-front passengers (RFPs) 15% more likely to be injured in a collision than drivers, a conclusion that Evans (1991) attributed to a preponderance of female RFPs. To account for variations in occupant and collision characteristics, used the double pair comparison method [] and concluded that unbelted drivers and RFPs have approximately the same probability of sustaining a fatal injury for the same set of collision and occupant parameters. This finding does not account, however, for differences in restraint interaction since only unrestrained occupants, and no airbag-equipped vehicles, were included in the study.
In addition to collision exposure and occupant factors, the greater potential for RFPs to be out of position and the inherent differences in the driver and RFP environments and restraint systems may contribute to a skewed injury distribution. For example, the relative proximity of the steering wheel limits the distance available for a driver’s forward excursion in a frontal impact. In contrast, a RFP generally has more distance available over which to reach a common velocity with the decelerating vehicle in a frontal impact. Despite the potential benefit of this additional space, belt effectiveness in vehicles without airbags has generally been found to be slightly lower for RFPs than for drivers, possibly due in part to unequal injury potential for unbelted drivers and RFPs. In a study of 1.5 million driver and RFPs involved in a collision, found that seat belt use reduced driver fatalities by 63 percent to 67 percent, while RFP fatalities were reduced by only 53 percent to 55 percent. Viano (1995) also found a disparity in restraint belt effectiveness by seated position: 39 percent effectiveness for RFPs compared to 42 percent for drivers (no airbag). For restraint systems including both a lap-shoulder belt and an airbag, however, Viano concluded that overall restraint system effectiveness would be approximately equal for drivers and RFPs (47%). In a May 1999 report to Congress, the National Highway Traffic Safety Administration (NHTSA) estimated that, for frontal impacts, passenger airbags are approximately as effective (32 percent) for all RFPs age 13 and older as driver airbags (31 percent) are for all drivers.
Unfortunately, the paucity of RFP airbag deployments in the field precludes a double pair comparison of head or thoracic injury risk for belted drivers and RFPs when an airbag deploys. Currently in the National Automotive Sampling System (NASS) through 1998), for example, there are only 493 matched pairs of driver and adult RFPs involved in a frontal impact wherein both frontal airbags deployed, at least one occupant received chest injuries, and both occupants were belted. Within these 493 cases, there are only 10 drivers and 15 RFPs who sustained an AIS 3+ chest injury. Current field data can, however, be used for a limited comparison of driver and RFP airbags. All frontal impacts in 1998 FARS were analyzed to evaluate the driver-to-passenger fatality (DPF) ratios for various restraint configurations (Table 1). Though the greater potential for driver-side offset collisions and intrusion influences the number of fatalities and the restraint system interaction, a comparison of fatality ratios provides a means for reducing the influence of collision factors. The DPF ratio for belted males decreases from 9.7:1 when the occupant’s frontal airbag deploys to 5.7:1 when a frontal airbag is not present (i.e., there were nearly 10 driver fatalities with an airbag for every RFP fatality with an airbag, but when there was no airbag, there were only about 6 driver fatalities for every RFP fatality). Similarly, the DPF ratio for unbelted males decreases from 8.5:1 with an airbag to 6.3:1 with no airbag. DPF ratios for female occupants exhibit similar trends, as do those for all occupants. These ratios indicate that a frontal airbag may provide more fatality-reducing benefit for belted RFPs than for belted drivers in frontal collisions. Similarly, an airbag may provide greater benefit for unbelted RFPs than for unbelted drivers, though, as would be expected, the effect is not as pronounced (the difference in ratios is not as pronounced).
Table 1
1998 FARS Data (Front Seat Occupants, Frontal Impacts)
| Restraint Configuration | ||||
|---|---|---|---|---|
| Airbag L/S Belt | Air bag No Belt | No Airbag L/S Belt | No Airbag No Belt | |
| Driver Fatalities | ||||
| Male | 494 | 665 | 1274 | 4498 |
| Female | 376 | 275 | 784 | 1039 |
| Total | 870 | 940 | 2058 | 5537 |
| RFP Fatalities | ||||
| Male | 51 | 78 | 223 | 711 |
| Female | 109 | 73 | 432 | 576 |
| Total | 160 | 151 | 655 | 1287 |
| DPF Ratios | ||||
| Male/Male | 9.7:1 | 8.5:1 | 5.7:1 | 6.3:1 |
| Female/Female | 3.4:1 | 3.8:1 | 1.8:1 | 1.8:1 |
| Total/Total | 5.4:1 | 6.2:1 | 3.1:1 | 4.3:1 |
One purpose of this paper is to evaluate the thoracic injuries sustained by RFPs restrained by a force-limited, pretensioned belt and airbag, and to compare the kinematics with existing tests of similarly restrained driver-side subjects. A secondary purpose is to evaluate the injury-predictive ability of existing thoracic injury criteria when applied to these subjects. These tests are illustrative, particularly for the evaluation of injury criteria, because the complex loading of the steering wheel on the anterior thorax is not present, and because, to date, cadaveric sled-test evaluation of thoracic injury criteria has been done using only driver-side subjects.
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BACKGROUND
Many cadaver tests have been done with different restraint conditions and subjects seated in the driver position (DP) (e.g., Kuppa and Eppinger 1998). Unfortunately, few tests have been performed using RFP cadavers, and the authors know of no tests that have been performed using belted cadavers with an airbag in the RFP position. As a result, a direct comparison of driver and RFP kinematics, injury predictors, and injury outcome in this restraint environment is not possible. Voigt and Lange (1971) performed a series of simulated frontal impact reverse-acceleration sled tests to evaluate instrument panel designs. These tests included 14 RFP cadavers, but all subjects were unbelted without airbags and no corresponding tests were performed with subjects in the DP. In 1998, Berg, et al. presented a sled test series that included two unbelted cadavers in the DP and a single unbelted RFP cadaver. Airbags were used and the sled was subjected to a simulated 50 km/h change-in-velocity (ΔV) frontal impact with a nominal 11g acceleration pulse. The utility of these tests for thoracic injury evaluation is limited, however, because none of the cadavers received thoracic injuries, possibly due to the low acceleration pulse.
THORACIC INJURY CRITERIA
The use of ATDs to evaluate human injury risk in automobile collisions is limited because many biomechanical mechanisms of injury (e.g., stress distribution and magnitude within the tissue or internal organ displacement) generally cannot be measured. As a result of this limitation, measurable kinematic and kinetic parameters must be used to predict the likelihood of injury. These engineering parameters, correlated with observed injury in cadavers or animal subjects, are referred to as injury criteria. In the thoracic injury literature, there has been no general agreement on which injury criteria are the best indicators of thoracic injury risk in a collision. Peak chest acceleration, peak chest compression, maximum viscous response (VCmax), and other parameters have been correlated to the presence of injury. Comprehensive reviews of the background, development, and critical values of these criteria have been published by Cavanaugh (1993), Kuppa and Eppinger (1998), and Prasad (1999).
Morgan, et al. (1994) evaluated a series of 63 driver-side cadaver tests and developed a combined model (CM, see equation 2 of Morgan, et al.) for the prediction of thoracic injury. This CM includes the peak chest compression (measured, using chestbands, at any of five locations on the anterior thorax), the 3-ms peak resultant acceleration measured at the first thoracic vertebra (T1), and the age of the cadaver at death, and was found to correlate with injury better than either peak chest compression or peak T1 acceleration alone. Further, these authors found that specific characteristics of the restraint system influenced the probability of injury for a given magnitude of CM. Specifically, they found that there is a greater probability of injury with a “belt-like” restraint than with an “airbag-like” restraint for the same measured value of CM (see Figure 16 and Figure 17 of Morgan, et al.). Based on this finding, these authors recommended separate injury criteria for “belt-like” restraints and “airbag-like” restraints. The recommended “belt-like” criterion includes measured compression terms only (no acceleration or viscous terms), while the “airbag-like” criterion includes both compression and acceleration terms (see equation 4 and equation 5 of Morgan, et al.). With these separate injury criteria, the authors were able to predict injury correctly in 48 of 52 tests.
Recently, the combined thoracic criterion CTI has been proposed as a desirable criterion for the evaluation of combined belt-and-airbag restraint systems [Eppinger, et al. 1999]. Similar to the CM of Morgan, et al. (1994), CTI is based on the finding that a two-parameter logistic regression model using the T1 peak acceleration and the maximum chest compression correlated well with the probability of an AIS 3 or higher injury (defined by the number and distribution of rib fractures sustained by the cadaver). Kuppa and Eppinger (1998) stated that a combined acceleration and compression model “is able to take into account the differences in thoracic response due to belt and airbag loading.” The findings of Morgan, et al. (1994) seem to indicate that the mechanism by which the combined acceleration-compression model is able to differentiate restraint loading is in the relative contributions of the compression and acceleration terms: a dominant maximum compression term indicates “belt-like” loading while a relatively more dominant acceleration term indicates “airbag-like” loading.
In 2000, Hassan and Nusholtz revisited the development of CTI. They identified several confounding issues associated with the tests, data collection, and statistical model used to develop CTI. They also discussed the appropriateness of using statistical analyses of sled tests for the development of thoracic injury criteria. Cases of massive rib fractures (greater than 10 fractures) with relatively little measured chest compression (8 percent to 16 percent) were identified as deleterious to the dataset. Examination of these tests reveals that, depending on the specific anthropometry of the cadaver subject, interaction with the lower steering wheel rim and airbag may create a complex loading environment on the ribs, and chest deformation measured at discrete locations in the anatomically transverse plane may not be a good indicator of injury risk - particularly when injury risk is defined almost exclusively by the number and location of rib fractures. evaluated a series of driver-side, unbelted cadaver sled tests and identified rib fractures on the anatomically inferior ribs as the result of steering wheel loading, even in the presence of an airbag. Inspection of test number UVA96 from Kuppa and Eppinger (1998) verifies that rib fractures can result from steering wheel and airbag loading, and that these fractures can occur at areas remote to the chestbands (Figure 2). This relatively young and large (58 years old at death, 97 kg) subject sustained 14 rib fractures (7 displaced) despite a measured peak chest compression of only 8 percent1. Because individual cadaver anthropometry (overall size in addition to relative segment sizes) varies among subjects, the interaction with the steering wheel and airbag can be difficult to quantify. Testing with subjects in the right-front seating position may ameliorate some of these confounding issues and allow for a more direct analysis of the combined restraint loading on the thorax.
https://imdiubabo.tistory.com/18. Steering wheel, airbag, and anterior thorax interaction (test UVA96 from Kuppa and Eppinger (1998)) and resulting rib fracture pattern.
PASSENGER SLED TEST METHODOLOGY
ATDs and cadavers were seated in the RFP position of a reinforced contemporary mid-size vehicle buck mounted on a deceleration sled (Figure 3). A hydraulic decelerator (model 931-4000, Via Systems, Salinas, California) was used to shape the sled acceleration pulse. A nominal impact velocity of 48 km/h and an approximately trapezoidal vehicle acceleration pulse (18 g peak, 105 ms duration) were used for all tests. All subjects were restrained with a three-point belt system (including a buckle pretensioner and nominal 3.5-kN force-limiting retractor), the original-equipment instrument panel (IP), and an airbag.
Passenger test environment compared to typical driver test environment.
Tension gauges measured belt loads at three locations: the upper shoulder belt (between the d-ring and subject’s shoulder), the lower shoulder belt (between the subject’s hip and the buckle), and the outer lap belt (between the subject’s hip and the lap belt mounting location on the vehicle buck).
Photographic documentation included digital still images taken before and after the test, and high-speed (1000 frames/sec) color digital video taken with imagers (Kodak RO, Rochester, New York) positioned off-board on the driver and RFP sides. Tests also included an on-board 16-mm high-speed film camera (Photosonics 1-B, Burbank, California). For redundancy, black-and-white digital video (Kodak Ektapro, Rochester, New York) was taken in all tests. Sled impact velocity was measured using a proximity sensor and a redundant optical gate.
ATD TESTS
A fiftieth percentile male Hybrid III ATD was used for all tests. Three ATD tests were performed. The ATDs were fitted with an instrumentation package that included tri-axial accelerometer arrays at the head and chest centers of gravity, pelvis, and externally at the upper spine, a uniaxial accelerometer on the sternum, 6-axis upper and lower neck load transducers, uniaxial femur load transducers, and a single rotary potentiometer that measured deflection at one point on the sternal plate (“chest slider”). Supplemental instrumentation included an array of string potentiometers in the chest (Hagedorn and Pritz 1993), two 40-gauge chestbands (Eppinger 1989) positioned externally at the nominal level of the second and fifth ribs, and two magnetohydrodynamic (MHD) angular rate sensor arrays mounted internally at the head center-of-gravity and externally at the upper spine. Occupant positioning guidelines from the Ford Motor Company (1999) and NCAP were used.
CADAVER TESTS
Four cadaver tests were performed. Occupant positioning guidelines from Ford Motor Company (1999) and NCAP were used, though the final seat and subject positions were dictated by several factors, including the cadaver’s H-point location, knee-to-IP distance, chest-to-IP distance, and clearance between the head and windshield roof header. The dominating factor was the head clearance, since head strikes against the windshield header have been found to generate artifactual thoracic acceleration peaks. In no test, however, was the H-point fore/aft position more than 6.2 cm away from the center position (Table 2).
Table 2
| Test No. | Subject | H-Point (mm rear of OEM Center) | Chest to I.P. (mm) | Knee to I.P. (mm) |
|---|---|---|---|---|
| 576 | ATD | 1 | 470 | 99 |
| 577 | cadaver | 62 | 510 | 129 |
| 578 | cadaver | 3 | 479 | 107 |
| 579 | cadaver | 4 | 470 | 116 |
| 580 | cadaver | 60 | 522 | 128 |
Cadavers were preserved until the time of testing by refrigeration or freezing. The cadavers were obtained through either the Virginia State Anatomical Board or the Anatomical Board of the State of Texas. Age, weight, height, and cause-of-death criteria were used to screen the cadavers such that subjects within the target population were used for testing (Table 3). Cadavers were screened for infectious diseases, including HIV and hepatitis A, B, and C. Subjects having an extended period of convalescence prior to death were excluded from the study. Pre-test radiographs were taken to identify existing pathology; if abnormal skeletal conditions were found, the specimen was removed from the test population. All test procedures were approved by the University of Virginia institutional review board prior to testing.
Table 3
| Test No. | Age/Gender | Mass (kg) | Ht. (cm) | Cause of Death | Preservation Method | Chest Depth (cm) | |
|---|---|---|---|---|---|---|---|
| 4th rib | 8th rib | ||||||
| 577 | 57/M | 70 | 174 | Pneumonia | Refrigerated | 21.1 | 23.3 |
| 578 | 69/F | 53 | 155 | CVA* | Frozen | 20.8 | 21.7 |
| 579 | 72/F | 59 | 156 | MI* | Frozen | 23.8 | 25.2 |
| 580 | 57/M | 57 | 177 | ACI* | Frozen | 23.1 | 24.2 |
MI: Myocardial Infarction
Pulmonary and cardiovascular systems were pressurized to typical in vivo levels (approximately 10 kPa) immediately before testing.
Cadaver instrumentation included tri-axial accelerometer arrays mounted on the posterior pelvis and at the levels of the twelfth and the eighth or ninth thoracic vertebrae (T12 and T8/9). The acceleration measured at T8/9 was considered to be analogous to the acceleration at the chest center-of-gravity (CG). Arrays of three orthogonal accelerometers and MHD sensors were mounted on the posterior surface of the head and posteriorly on T1, both approximately on the midsagittal plane. A uniaxial accelerometer was mounted on the body of the sternum, immediately below the manubrium. Where applicable, digital data were filtered according to SAE recommended practice J211. Chest deflections were measured using two 40-gauge chest bands installed externally at the levels of the fourth and eighth ribs, measured laterally. In selected tests, a pressure transducer was inserted into the aortic arch to verify pressure stability during sled run-up and to measure intra-aortic pressure during the impact.
RESULTS
In all tests, the instrumentation functioned properly and well-controlled, repeatable response was achieved. The airbag and buckle pretensioner deployed properly in all tests except test 579. In this test, a broken circuit between the capacitive discharge squib ignitor and the pretensioner prevented pretensioner activation.
ATD TESTS
Three ATD tests were performed, but the repeatability was sufficient to justify presentation of a single test. The peak chest acceleration of 26g occurred at 67 ms. The maximum chest deflection, as measured by the chest slider, was 29 mm and occurred at 107 ms (Figure 4, Table 4). The chest acceleration exhibited a bimodal profile, with an initial maximum occurring at approximately the time of peak shoulder belt force (Figure 4), before the subject’s torso loaded the airbag (Figure 5). After this first peak, the acceleration decreased to a local minimum at approximately 93 ms. This decrease was accompanied by a decrease in lap belt force, though there was not a corresponding decrease in shoulder belt force (Figure 4). The chest acceleration subsequently increased to a second peak as the subject loaded the airbag.
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ATD resultant chest acceleration, chest compression (from chest slider), and restraint belt loading time histories.
RFP ATD at times of peak resultant chest acceleration compared to similarly restrained driver ATD.
Table 4
| Test | 571 ATD | 577 CAD | 578 CAD | 579 CAD | 580 CAD |
|---|---|---|---|---|---|
| Chest CG Resultant Acceleration, g (time, ms) | 26 (67) | N/A | N/A | N/A | N/A |
| T8/9 Resultant Acceleration, g (unscaled) (time, ms) | N/A | 42 (57) | 48 (60) | 43 (62) | 46 (62) |
| T8/9 Resultant Acceleration, g (scaled to 50th male) | N/A | 41 | 43 | 40 | 42 |
| T1 Resultant Acceleration, g (unscaled) (time, ms) | N/A | 34 (87) | 30 (72) | 33 (96) | 35 (91) |
| T1 Resultant Acceleration, g (scaled to 50th male) (time, ms) | N/A | 33 | 27 | 30 | 32 |
| Chest compression measured by slider, mm (time, ms) | 29 (107) | N/A | N/A | N/A | N/A |
| Upper chestband peak compression, mm [%](time, ms) | 55 [20*] (113) | 50 [24*] (113) | 52 [25*] (101) | 88 [37*] (104) | 67 [29*] (104) |
| Lower chestband peak compression, mm [%] (time, ms) | 48 [18*] (74) | 19 [8*] (64) | 32 [15*] (60) | 43 [17*] (98) | 40 [17*] (60) |
| Upper chestband peak sternal compression, mm [%](time, ms) | 46 [16*] (114) | 48 [23*] (112) | 52 [25*] (101) | 81 [34*] (105) | 66 [28*] (104) |
| Upper sternal peak compression velocity, m/s (time, ms) | 2.7 (23) | 2.7 (24) | 2.6 (14) | 3.7 (75) | 2.7 (12) |
| Upper sternal VCmax, m/s (time, ms) | 0.2 (63) | 0.2 (66) | 0.3 (93) | 1.0 (75) | 0.3 (52) |
| CTI (using slider) | 0.57 | N/A | N/A | N/A | N/A |
In contrast to this response, a similarly restrained subject on the driver side loads the belt and relatively more proximate airbag concurrently at the time of peak chest acceleration, even when a belt pretensioner is used (Figure 5). For the same force limit level, the driver-side subject’s belt force saturates slightly later than the passenger-side subject’s because the driver interacts with the airbag earlier. In addition, since the driver-side subject loads both the belt and the airbag at the time of peak chest acceleration, the chest acceleration does not exhibit the distinct two-peak profile measured by the passenger ATD [Shaw, et al. 2000].
CADAVER TESTS
Kinematic Data
The cadaver acceleration responses were normalized to the standard 50th percentile male mass (78 kg) using the scaling procedure described by Eppinger, et al. (1984).
As was the case with the ATD tests, the passenger subjects exhibited a different response than that typically exhibited by driver subjects: the maximum scaled passenger T8/9 acceleration (considered to be analogous to chest CG acceleration) of approximately 40g occurred before the cadaver loaded the airbag. The T8/9 acceleration profile exhibited two distinct local maxima; the first occurred at 55 ms to 65 ms, approximately the time of shoulder belt force saturation (Figure 6b). A subsequent decrease in passenger chest acceleration (local minimum) occurred at approximately 75 ms. The shoulder belt force did not exhibit a decrease in magnitude corresponding to the decrease in T8/9 acceleration. Similar to the ATD tests, however, a corresponding decrease was observed in lap belt force (Figure 7). The second chest acceleration peak, which was of lesser magnitude than the first peak, occurred as the subject was loaded by both the belt and the airbag (Figure 8). Published driver-side cadaver tests performed with a comparable restraint configuration and impact speed [Shaw, et al. 2000] did not indicate as pronounced a bimodal chest acceleration response (Figure 6a). The driver subjects loaded the airbag earlier than the passenger subjects, and the peak chest acceleration occurred as the subject loaded the belt and airbag simultaneously.
Resultant T8/9 and T1 acceleration time histories for passenger-side cadaver tests compared to driver-side cadaver tests in the literature (all scaled to 50th percentile male mass).
Belt force time histories for passenger-side cadaver tests.
Cadaver position at time of T8/9 resultant acceleration maxima – (a) driver-side subject and passenger-side subject at (b) time of initial peak and (c) time of second peak.
The resultant acceleration measured at T1 in the passenger-side tests did not exhibit the distinct bimodal profile measured at T8/9 (Figure 6d). The T1 acceleration maximum also occurred later than that at T8/9 - between 72 ms and 96 ms after impact.
The maximum normalized chest compression, as measured by the upper chestband, was lowest in test 577 (24% of initial chest depth) and highest in test 579 (37%) and occurred between 101 and 113 ms after impact (Table 4). The subject was loaded by both the belt and the airbag at the time of peak chest compression (Figure 8).
Injury Outcome
Post-test radiographs were taken and detailed necropsies were performed to quantify thoracic injury outcome. Injuries were coded according to the Abbreviated Injury Scale (AIS) (1998). Rib fractures were the most common injury noted, though a sternum fracture and superficial liver lacerations were also found. Due to postmortem autolysis, however, the liver lacerations may be artifactual. No other soft tissue injuries were observed (Table 5). The costal pleura was not lacerated in any test.
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Table 5
| Test | Thoracic Injuries | AIS | MAIS |
|---|---|---|---|
| 577 | None | 0 | 0 |
| 578 | 4 left rib fx. (L 4, 5, 6, 8) | 3 | |
| 2 superficial liver lac. | 2 | 3 | |
| 579* | 7 left rib fx. (2 - 8), 4 right rib fx. (2 – 5) | 4 | |
| Sternum fx. | 2 | 4 | |
| 580 | None | 0 | 0 |
The subject in test 579 had eleven rib fractures, seven more than any other subject. This is likely due to several factors: a) the cadaver was the oldest in the set and one of the least massive, b) the cadaver was female, and c) due to a mechanical malfunction, the buckle pretensioner did not activate in this test.
DISCUSSION
As a result of the observed differences in driver and passenger occupant thoracic loading mechanisms, and hence the resulting kinematics, existing kinematic injury predictors and injury reference values that have been developed using subjects seated exclusively on the driver side may need to be modified in order to apply to subjects seated on the passenger side. Although the low number of passenger tests performed to date precludes a statistical analysis of existing injury criteria in the RFP environment, specific observations about the predictive ability of injury criteria for this set of tests will be discussed.
The acceleration measured at the chest CG is an indicator of the force acting on the thorax and as such has been used to describe thoracic injury tolerance. Peak acceleration alone does not, however, completely describe the distribution of the force on the thorax (i.e., a point force of high magnitude and a distributed force of low local magnitude can result in the same measured peak acceleration at the spine). Individual rib loading depends on the distribution of force applied on the thorax. concluded that ribs typically fracture in bending. In other words, a rib fracture is dependent on the moment applied to the rib. Therefore, since the applied moment on each rib depends on the distribution of the force applied to the anterior thorax, peak chest acceleration alone is not a universally accurate predictor of rib fractures. The acceleration term in a combined compression-acceleration model may, however, have injury predictive value due to its value as an indicator of load distribution on the anterior thorax: for a given level of peak thoracic compression, higher chest acceleration indicates a more distributed load (such as would occur if an airbag restraint was present) (Figure 9).
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Schematic depiction of two different thoracic loading conditions: a) belt loading, b) combined loading. Sternal compression is approximately equal for each condition (compression in spring “B”), as is Cmax (compression in spring “C”). The superimposed airbag loading, however, results in greater total force on the anterior thorax (represented by increased compression in spring “A”) and a higher acceleration at the spine or chest CG.
The peak chest acceleration (T8/9) measured in these passenger tests does not, however, indicate the relative loading of the belt and airbag because, in both the ATD and cadaver tests, the peak acceleration occurred before the subject was loaded by the airbag. Further, this maximum does not appear to differentiate injury and no-injury tests (Figure 10, upper plot). All cadaver tests resulted in peak scaled chest (T8/9) acceleration between 39g and 43g, below the 50th percentile male injury reference value (IRV) of 60g, and no relationship between the peak T8/9 or T1 acceleration and the injury outcome is observed (see also Table 4 and Table 5).
Peak chest compression and the first and second acceleration peaks for passenger subjects. The numbers near each data point indicate the number of rib fractures sustained by that cadaver.
The magnitude of the second peak in passenger chest acceleration occurs as the subject loads both the belt and the airbag. Therefore, it is analogous to the peak acceleration measured using driver subjects. This peak relates better with the presence of injury than does the magnitude of the first peak (Figure 10, lower plot). Though none of these values approaches the 50th percentile male IRV, test 579 (11 rib fractures) resulted in the highest second peak in scaled chest acceleration, indicating that this peak may have some injury-predictive value.
Differences in the acceleration (both magnitude and shape of the time history) measured at T8/9 and at T1 were observed in these passenger-side tests. The acceleration measured at T8/9 was found to represent the dummy chest CG bimodal response while the acceleration measured at T1 did not (compare Figure 4 with Figure 6b and Figure 6d). This is an indication that thoracic injury criteria developed using acceleration measured at T1 may not be appropriate for use in dummy tests where the chest CG acceleration is used. In addition, by exhibiting the bimodal profile, T8/9 acceleration provided a better description of the restraint loading on the thorax than did T1 acceleration.
Peak normalized chest compression measured by the upper chestband was found to relate to injury outcome. All tests except 579 were near the injury reference value of 28% [Kuppa 1999], with test 579 yielding a value of 37%. The subject in this test received numerous bilateral rib fractures, a sternal fracture, and a separation of the sterno-clavicular junction. Therefore, it is possible that the large peak chest compression measured in this test is a result of decreased thoracic stiffness due to injury rather than the cause of the injury.
The maximum sternal viscous response, VCmax, (measured externally using the upper chestband) was highest for test 579, reaching the injury reference value of 1.0 m/s at 75 ms. The maximum sternal compression velocity of 3.7 m/s also occurred in test 579 – the only test in which the compression velocity was above the minimum value (3.0 m/s) recommended for application of VCmax as an injury predictor [Lau and Viano 1986]. Similar to the maximum compression, the high sternal compression velocity and VCmax measured in this test may be the result of injury.
The probability of AIS 3+ chest injury in these tests may be estimated using the response of the ATD and existing probability functions (Table 6). Depending on the specific criterion chosen, thoracic injury probability for this passenger-side test environment ranges from 1.5% using CTI to 18.1% using peak chest acceleration alone. One of the three cadavers tested (test 579 excluded due to pretensioner failure) sustained AIS3+ chest injuries. The under-prediction of injury outcome is expected, given the low number of tests, the age and gender makeup of the cadaver subjects, and the injury tolerance differences between living and dead subjects [, Chapon, et al. 1980, Foret-Bruno 1978]. The probability range resulting from this pastiche of criteria indicates, however, that the source of the injury criterion (e.g., blunt impactor tests or sled tests), the tested restraint condition, and the specific type of injury being predicted must all be considered in the interpretation of dummy measures.
Table 6
Thoracic Injury Probability Using Three Different Criteria*.
| ATD measure | Value | Probability of AIS 3+ chest injury |
|---|---|---|
| Peak chest acceleration | 26 g | 0.181 |
| CTI | 0.57 | 0.015 |
| Maximum chest compression | 29 mm | 0.088 |
CONCLUSIONS
Field database cases are currently insufficient for a double pair comparison analysis of driver and RFP airbag effectiveness. Therefore, a fatality ratio comparison methodology was presented which provides a useful means of comparing driver and RFP restraint systems. This comparison indicates that an airbag may provide greater fatality-preventing benefit for RFPs than for drivers.
The four cadaver tests presented here are insufficient for a full evaluation of the efficacy and biomechanical basis of existing injury criteria. The right-front passenger position does, however, provide a desirable environment for evaluating thoracic injury criteria. The complex loading created by the steering wheel in a driver-side test can be avoided and restraint system effects can be evaluated directly.
In these tests, subjects (both cadavers and ATDs) seated in the right-front passenger position exhibited two distinct peaks in chest acceleration when restrained by a pretensioned, force-limited belt and airbag system. The first and global peak occurred under primarily belt loading. The second peak occurred as the subject was loaded by both the belt and airbag. Because the maximum chest acceleration occurred before the passenger-side subject interacted with the airbag, this first peak is not useful for evaluating the relative contributions of belt and airbag loading on the anterior thorax. The second peak does, however, occur during combined loading and therefore may have some injury-predictive value, though more testing is needed to evaluate this hypothesis.
This bimodal chest acceleration response was not observed for similarly restrained subjects seated in the driver position. In published driver-side tests, the airbag module was closer to the subject and as a result the driver-side subjects engaged the airbag approximately 10 ms earlier than the passenger subjects, precluding an initial belt-generated peak in chest acceleration.
Both the ATD chest CG acceleration and the cadaver T8/9 acceleration exhibited the bimodal profile, while the acceleration measured at T1 did not. The T8/9 acceleration provided a better description of the restraint system loading on the thorax than the T1 acceleration. Further, the T8/9 acceleration peak was approximately 25% higher than the peak acceleration at T1 (see Figure 6). These differences in measured acceleration raise questions about the validity of using T1 acceleration to develop injury criteria which are then applied using ATD chest CG acceleration to evaluate restraint system performance.
(Presenter: Richard Kent)
Driver Restraint System
Jack Richman: I’m a physician and I always have a struggle with cadaver tests because I examine live people and I’ve also looked at cadavers and there’s really no comparison in terms of how a cadaver will respond versus a live person even though we try to make those comparisons. I think this is significant. Tissues in cadavers are brittle, muscles tear. There’s no response to injury. When a person is struck, the nervous system kicks in. There is no nervous system in a cadaver to kick in. The bones are dehydrated and brittle; they will fracture easily, and yet we try to make a lot of conclusions about live people by using cadavers. I think in a lot of these papers I’m very anxious that we don’t try and draw too many conclusions between cadavers and live people because the response of tissues which is non existent in cadavers and response of tissues in humans which are very responsive are different.
R. Kent: Your comments are certainly well taken and I agree with you that the injury outcome is different, but testing with human cadavers regardless of the injury outcome does give us a much better idea of the overall kinematics of the collision, specifically as you get away from really controlled restraint conditions. If you test with a lap belt only and an airbag, the dummy response and the cadaver response are very different and one would assume the cadaver would be more human-like, but you’re certainly right with your comments. Thank you.
Jeff Crandall: If I can just expand on that comment. Granted, there are differences between cadavers and living persons. However, some of the statements such as bones being brittle and these sort of things, I think the bones for a cadaveric population are very representative of what you have in a living population. There’s minimal changes post mortem in that area. In terms of some of the muscular effects, some of the passive responses, there’s also not that great a change. Obviously, when you’re talking about active responses and some of the neurological reactions, that indeed is a major factor. However, there’s also a subset for which cadavers are representative. There’s a response time, for example in a crash beyond which many of these muscles can’t react. So there is a certain population of which cadavers are representative.
Jack Richman: There are differences in the cadavers such as in the older female where she may be osteoporotic where she would probably have fractures even if she were alive and therefore there may be some correlations in that. But at the same time I think what we do, as long as we look at the other variables of the age and sex of the cadaver which was done here, then I would agree with your first comments and also how long it’s been since the person died versus whatever because there is muscle change over time – so that if the cadavers are reasonably fresh, I’ll agree with you on the muscle, but if the cadaver is not, then it’s different. Ati radeon 3000 windows 10.
J. Crandall: And I think that point is a good one and one that Murray Mackay brought up this morning. What we all have to be striving toward is to make something that is age and gender dependent.
Jim Pywell: Nice paper. I know that we evaluated dummies and cadavers at essentially mid point position for the seat. People don’t sit like that and I’m wondering whether there’s an opportunity for us to reconsider positioning in terms of how we test. Kodak easyshare printer dock series 3 drivers for windows 7. In addition, most foreign manufacturers put their wheel less proximate to the occupant. I think we’re going to see US manufacturers move the wheel less proximate as well. My question is do both those things tend to get us to evaluate systems consistently between driver and passengers, as I think we will be opening up those spaces between the driver and the wheel and the restraint system?
R. Kent: That’s an excellent comment. I didn’t really get the question out of that. Let me address it this way. Both of those tests, the driver and the passenger, were in the mid position and I think if you move the occupants back, you start to exacerbate these differences because then you get an even longer belt loading before you get airbag loading, particularly on the passenger side. I think that’s a very valid point.
ACKNOWLEDGMENTS
This research was sponsored by the Alliance of Automobile Manufacturers (AAM), though the views expressed herein are the authors’ and are not necessarily those of AAM. The authors gratefully acknowledge the critical reviews provided by Shashi Kuppa, Priya Prasad and Guy Nusholtz.
Footnotes
1This test was removed from the dataset for the development of CTI, see Eppinger et al., 1999.
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