small female-specific biomechanical corridors ......ircobi conference – graz (austria) september...

IRCOBI Conference – Graz (Austria) September 2004 137

SMALL FEMALE-SPECIFIC BIOMECHANICAL CORRIDORS IN SIDE IMPACTS

Narayan Yoganandan, Frank A. Pintar, Thomas A. Gennarelli Department of Neurosurgery Medical College of Wisconsin

Milwaukee, WI, USA

ABSTRACT

Biomechanical response corridors are necessary to evaluate the performance (e.g., biofidelity) of dummies. In side impacts, while corridors are available for the mid-size male, data applicable to the small size female are not reported. Because the small female anthropometry is gaining increasing worldwide attention, this study was undertaken to develop corridors. Sled tests were conducted using post mortem human subjects at a velocity of 6.7 m/s. Three chestbands were used to compute deflection-time histories at the axilla, xyphoid process, and tenth rib levels. Triaxial accelerometers were fixed to the upper and lower spine and sacrum to record acceleration-time histories. Specimens contacted the load wall with varying initial conditions (rigid and padded; flat wall and offset) from which impact forces to the thoracic, abdominal, and pelvic regions were obtained using load cell data. Adopting signal processing and mass scaling methods, corridors were derived for forces, accelerations, and chest deflections at three levels for all initial conditions. All time history corridors were expressed as mean plus/minus one standard deviation. Effects of using gender-specific tests for the derivation of corridors are discussed. These results will assist in the assessment of anthropomorphic test devices with small female anthropometry. Key Words: Biomechanics, corridors, side impact, sled tests, injury metrics BIOMECHANICAL RESPONSES ARE DETERMINED using experimental models such as post mortem human subjects (PMHS) and anthropomorphic test devices (dummies) for application in crashworthiness research. Although PMHS do not permit direct physiological measurements, because of their injury-producing capabilities and anatomical similarities with the in vivo human, impact and injury responses from this model are used in the design and development of dummies. While the Hybrid III dummy has undergone extensive evaluation and is accepted as the federally regulated frontal dummy in the United States and widely used elsewhere, side impact dummies are being continuously updated and redesigned (CFR 2000; Yoganandan et al., 2002). For example, the mid-size male WorldSID is under development, and results from recent modifications of the EuroSID (e.g., ES-2re) are being evaluated for its efficacy in crashworthiness research (Kuppa et al., 2003). Performance evaluations are possible because of the availability of time-varying force, acceleration, and deformation responses of various regions of the human body in PMHS sled tests (Maltese et al., 2002). Our group derived such response corridors (mean and plus/minus one standard deviation) in the cited reference for the 50th percentile (mid-size) male. The small female anthropometry has attracted considerable attention recently due to lessons learned from epidemiological, anatomical, and biomechanical studies. For example, under similar frontal airbag deployment exposures, the small female dummy encounters higher impact loads than the mid-size male dummy (Melvin et al., 1993). Out-of-position loading scenarios are particularly critical to small size females. Response corridors applicable to small females in side impacts are needed to evaluate the performance (e.g., biofidelity) of dummies (SID IIs) of this anthropometry. The present study, undertaken due to a lack of small female-specific responses, is designed to develop corridors by subjecting PMHS to simulated side sled impact tests under varying initial conditions.


METHODS Unembalmed PMHS were procured, medical records evaluated, and screened for HIV, and Hepatitis A, B, and C. Anthropomorphic data and pretest x-rays were obtained according established procedures (Pintar et al., 1997). Specimens were dressed in tight-fitting leotards, and a mask covered the head/face. Prepared subjects were placed on a Teflon-coated bench seat (1.3-m long) fixed to the platform of a deceleration sled, configured with an impacting load wall, to simulate side impact. Figure 1 shows the schematic of the load wall and sled buck. Four plates (upper plate for measuring contact forces with the mid-thorax, middle plate for the abdomen, lower plate for the pelvis, and extremity plate for the lower extremities) were used in the load wall design. The configuration of the load wall was such that it was flat or had a 100-mm pelvic or thoracic offset, achieved by moving the lower or upper plate by the predetermined distance. The vertical height of the upper edge of the thoracic plate was set at 400 mm to prevent shoulder contact. This dimension was chosen to represent the average windowsill height of passenger cars. Thus, data obtained reflects real-world environment. The initial positioning was such that the Frankfort plane was horizontal, legs were stretched parallel to the midsagittal plane, and normal curvature and alignment of the thoracolumbar spine were maintained without any initial torso rotation. The specimen contacted the initially configured load wall (padded, rigid, or offset) without any significant changes in the anatomical interrelationships between the various body segments. The entire pelvis up to the level of the iliac crest of the PMHS contacted the pelvic load plate. The abdominal load plate was exposed to the lower regions of the ribcage. The thoracic load plate engaged the section of the middle ribcage. The test matrix included rigid and padded impacts (10-cm thick, LC200, compressive stiffness 103 kPa). PMHS pressurization methods are given (Pintar et al., 1997).

Figure 1: Schematic of the side impact buck with the thoracic (T), abdominal (A), pelvis (P), and extremity (L) plates to measure impact forces from the dummies.

Instrumentation consisted of an accelerometer on the sled to obtain the change in

velocity. Tri-axial accelerometers were fixed to the upper spine (T1 region), lower spine (T12 region), and sacrum. The Cartesian coordinate system of reference was adopted. The positive x-acceleration was along the posterior-anterior direction, positive y-axis acceleration was along the left-right axis, and positive z-axis acceleration was along the superior-inferior direction. To record medial-lateral accelerations of the struck-side ribcage, uni-axial accelerometers were fixed to the left side of ribs four and eight and sternum. Eleven load cells (two each in the thorax and abdomen, four in the pelvis, and three in the extremity)


were used to record the dynamic forces. Three chestbands were fixed at the level of the axilla (upper), xyphoid process (middle), and tenth rib (lower) to measure deformation-time (contours) histories during impact. All data were gathered according to SAE J211 specifications using digital data acquisition systems. Forces were sub-sampled at 3200 Hz, chestband signals were filtered at class 600, and deformation contours were computed at 250 one-millisecond intervals. On each contour, three locations were selected. Starting at the spine (defined as zero percent of contour circumference) and following the clockwise contour path, locations were identified at 20, 25, and 30 percent of the circumference. A line was drawn between the sternum (one-half of contour circumferences) and spine on each contour, and the three identified locations were projected onto the sternum-spine line. The distance was measured between each point on the contour of the left side of the thorax and the projected sternum-spine line. The resulting three measurements (at 20, 25 and 30 percent circumference) were averaged to obtain the mean deflection. The process was repeated at all time intervals to generate a left-half chest deflection-time plot for each chestband signal. Normalized chest deflections were obtained by dividing the above-determined deflection by the one-half depth of the specimen’s chest.

Because the study was aimed at determining corridors for the small size female

anthropometry, data were mass scaled to 46.3 kg (Eppinger et al., 1978; Eppinger et al., 1984; CFR 2000). The first 20 tests in table 1 were previously analyzed to derive injury criteria (Kuppa et al., 2003). Briefly, normalized forces were obtained by scaling the measured forces to the two-thirds power of the ratio of the masses (between the small size female and actual PMHS), acceleration to the minus one-third power of the ratio, and deflections to the cube root of the ratio. The following procedure was used to establish time-zero. Contact with the load plate was determined by identifying the first data point in the force-time plot at which the load exceeded 0.2 kN. Decrementing along the curve indicated the time at which the force crossed the null value. This time was associated with time-zero for all sensor data. To quantify the time of occurrence of a particular signal with respect to time-zero, a characteristic time was determined for each signal. Decrementing along the force- or deflection-time plot indicated the time at which the force or deflection decreased to 20% of the peak value. The interval between time-zero and this time was the characteristic time. Acceleration signals also underwent the same process with the exception that its integrated signal was used instead of the original acceleration-time history. The average characteristic time for each signal group (e.g., all sensors in the same chestband) quantified the mean time at which signals in the group occurred with respect to time-zero. Signals in each group were aligned using principles of cumulative variance (Maltese et al., 2002). The mean response at each time step was computed along with plus/minus one standard deviation for forces, deflections, and acceleration data under each initial condition.

RESULTS Of the 27 specimens, four male specimens were used in each condition shown in table 1 (except two in the thoracic offset). Four females specimens were tested in rigid flat wall, three in padded flat wall, and two in thoracic offset condition. No female specimens were used in pelvic offset tests. Considered for analyses were forces in the thorax, abdomen, and pelvis; deflections from the upper, middle, and lower chest bands; and accelerations at the upper and lower spine and pelvis. Time-history-based corridors expressed as mean plus/minus one standard deviation were grouped based on the type of initial condition. All tests were conducted at a change in velocity of 6.7 m/s. Figures 2-6 show corridors for the rigid and padded flat wall, thoracic offset, and rigid and padded pelvic offset tests. A comparison of the peak mean force, deflection, and acceleration data for all initial conditions is shown in table 2 and the time of attainment of these variables are provided in table 3.


Figure 2: Time-history-based corridors, expressed as mean plus/minus one standard deviation, for the rigid flat wall condition.

Thorax

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Figure 3: Time-history-based corridors, expressed as mean plus/minus one standard deviation, for the padded flat wall condition.

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Figure 4: Time-history-based corridors, expressed as mean plus/minus one standard deviation, for the rigid thoracic offset flat wall condition.

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Figure 5: Time-history-based corridors, expressed as mean plus/minus one standard deviation, for the rigid pelvic offset condition.

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Figure 6: Time-history-based corridors, expressed as mean plus/minus one standard deviation, for the padded pelvic offset condition.

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Table 1: Summary of data.

Age Gender Weight Height Force (kN) Acceleration (g) Chest deflection

Condition (yrs) (kg) (m) Thorax Abdomen Pelvis T1 T12 Pelvis Upper Rib Middle Rib Lower Rib

Rigid flat 73 M 89 1.80 6.2 2.1 3.9 47.0 47.0 ----- 0.63 0.31 0.41

Rigid flat 27 M 72 1.75 6.8 2.8 8.4 63.0 59.0 76.0 0.30 0.32 0.21

Rigid flat 55 M 76 1.72 5.1 5.3 3.4 49.0 56.0 61.0 0.43 0.37 -

Padded flat 70 M 71 1.73 3.4 2.0 5.4 24.0 71.0 33.0 0.05 0.20 0.19

Padded flat 56 M 64 1.80 - 1.6 4.5 25.0 22.0 80.0 0.18 0.33 0.51

Rigid pelvic offset 78 M 90 1.66 4.0 0.6 12.4 75.0 35.0 51.0 0.15 0.15 0.12


Padded flat 72 M 66 1.60 3.8 2.1 4.0 51.1 35.3 38.5 0.52 0.46 0.36

Padded pelvic offset 59 M 73 1.77 3.0 0.6 7.3 36.1 33.1 31.4 0.24 0.16 0.21


Padded flat 75 F 42 1.60 2.5 1.5 5.4 57.6 63.7 47.1 0.32 0.34 0.48

Rigid flat 67 F 74 1.74 4.9 3.1 5.9 53.8 64.2 41.4 0.39 0.39 0.34

Rigid flat 86 M 67 1.70 5.3 2.3 8.3 49.2 52.4 75.9 0.47 0.36 0.31

Padded flat 79 M 53 1.66 3.3 1.6 5.2 38.1 43.7 35.7 0.28 0.28 0.32

Rigid flat 45 F 63 1.74 4.3 2.4 8.4 59.9 56.2 80.2 0.31 0.26 0.25

Thoracic offset 46 F 69 1.54 8.8 0.5 7.4 65.1 50.8 47.6 0.34 0.22 0.25

Thoracic offset 39 M 66 1.86 8.4 0.7 9.3 56.3 67.7 117.0 0.41 0.55 0.42

Rigid flat 56 F 64 1.64 3.4 1.8 6.7 50.6 58.7 37.2 0.53 0.43 0.42

Padded flat 54 F 61 1.72 2.7 1.4 5.2 41.4 34.2 34.3 0.44 0.46 0.46

Rigid flat 73 F 50 1.54 3.0 2.0 8.4 71.0 80.6 61.7 0.25 0.37 0.25

Padded flat 58 F 48 1.56 2.5 1.9 4.7 41.8 54.7 111.2 0.38 0.30 0.30



Thoracic offset 79 M 74 1.76 7.0 0.9 6.1 65.9 41.2 36.8 0.53 0.48 0.21

Thoracic offset 58 F 48 1.63 7.2 0.3 7.8 128.3 76.0 74.8 0.41 0.58 0.15

Padded pelvic offset 57 M 70 1.75 3.6 1.0 7.4 30.0 43.7 32.3 - - -

Rigid pelvic offset 76 M 73 1.76 4.0 0.2 12.2 53.8 41.8 72.7 - - -


Forces in all three regions were higher for rigid flat wall tests than padded tests although the magnitude of difference was the least for the abdomen (Table 2). This was also true for pelvic offset tests (abdomen exception). Both flat wall and pelvic offset tests resulted in the highest forces in the pelvic region (padded and rigid condition), followed by the thoracic and minimal forces in abdominal regions. Padded condition resulted in wider pulse durations (rise time to peak force) than rigid condition, and the time sequence of attainment of peak forces was pelvis, followed by abdomen and thorax (Table 3). These results were true for rigid and padded flat wall and pelvic offset tests. As expected, thoracic offset tests showed a leading tendency, i.e., higher magnitude (Table 2) and earlier time of occurrence of peak force for the thoracic region than the pelvis, and abdominal forces were relatively low (Table 3).

The highest accelerations occurred in the lower spine with the exception of thoracic

and rigid pelvic offset tests wherein the upper spine responded with the highest acceleration (Table 2). Sacrum accelerations were comparable to upper spine accelerations in all tests. Similar to the case of forces, the time of attainment of peak accelerations increased with the introduction of padding (Table 3). With the exception of thoracic offset tests, pelvic accelerations preceded upper spine accelerations.

Table 2: Mean peak force, acceleration, and chest deflection data.

Condition Thorax

(N) Abdomen

(N) Pelvis

(N)

Upper spine

(g)

Lower spine

(g) Sacrum

(g)

Upper chest (mm)

Middle chest (mm)

Lower chest (mm)

Rigid flat 3505 1432 4844 54.02 58.95 58.67 55.82 47.22 41.73

Padded flat 2589 1418 4483 31.01 58.50 36.40 43.27 50.01 55.70

Thoracic offset 6191 167 5890 64.45 53.79 66.93 59.04 65.82 32.55

Rigid pelvic offset 2964 161 8228 61.19 40.21 61.70 31.22 22.66 25.28

Padded pelvic offset 2463 514 5634 31.33 40.63 33.78 35.93 32.59 41.63

Table 3: Time of attainment (sec) of mean peak thoracic, abdominal and pelvic forces; upper and lower spine and sacrum accelerations; and upper, middle, and lower chest

deflections.

Condition Thorax Abdomen Pelvis Upper spine

Lower spine Sacrum

Upper chest

Middle chest

Lower chest

Rigid flat 0.0272 0.0228 0.0188 0.0209 0.0191 0.0175 0.0272 0.0272 0.0272

Padded flat 0.0331 0.0294 0.0206 0.0294 0.0238 0.0256 0.0447 0.0450 0.0459

Thoracic offset 0.0147 0.0300 0.0238 0.0178 0.0147 0.0238 0.0266 0.0309 0.0272

Rigid pelvic offset 0.0397 0.0106 0.0128 0.0366 0.0306 0.0134 0.0491 0.0491 0.0403

Padded pelvic offset 0.0447 0.0331 0.0269 0.0394 0.0359 0.0206 0.0606 0.0563 0.0531

The upper region of the chest sustained the highest magnitude of deflections compared to the lower and middle regions in rigid flat wall and rigid pelvic offset tests, while padded flat wall and pelvic offset tests produced the highest chest deflections in the lower region (Table 2). In contrast, the middle chestband recorded the highest deflections in thoracic offset tests. As before, padded tests increased the time of peak deflections in all cases (Table 3). Deflections corridors were expressed using the one-half chest width of the SID IIs dummy (160 mm). Therefore, if a dummy with different chest anthropometrics (e.g., BioSID) needs to be evaluated for its performance, these deflection corridors should be scaled to match chest dimensions of other dummies (175.4 mm for the BioSID). DISCUSSION Data from male and female specimens were used to derive corridors although responses reported in this study are applicable to the female gender. The procedure of


using both genders for derivation of corridors has been traditionally adopted in impact biomechanics. Mass scaling procedures were used to normalize data with respect to specific size (e.g., mid-size male) and derive mean responses from PMHS tests because of variations in subject mass. The method is based on the assumption that the material properties, i.e., modulus of elasticity and density are equal between the tested subject and the intended dummy size. Using both genders and mass-scaling procedures, corridors developed for the mid-size male anthropometrics (mass 75 kg) are being used to assess the biodifelity of dummies such as ES-2 and ES-2re (Rhule et al., 2002; Yoganandan et al., 2002; Kuppa et al., 2003). The impetus of the present study was subsequent to these recent research activities.

To evaluate the biomechanical performance of side impact countermeasures, the need to develop a second generation side impact dummy was recognized in 1993 (Daniel et al., 1995). Because frontal impact tests using the 5th percentile female dummy indicated higher loads than the 50th percentile male dummy, second generation dummy development efforts were focused on these anthropometrics and gender, leading to the design of the SID IIs dummy for side impact applications. Because of these developments and the focus of motor vehicle anthropometrics on the small size for this gender for frontal impacts (US federal regulations 208), the present study developed corridors for this group of the population (CFR 2000).

Corridors have been obtained using the following methods in frontal, rear, and side impact biomechanics literature focusing on the head, neck, and chest (Mertz and Patrick 1971; Lobdell et al., 1973; Yoganandan et al., 1997; Bolte et al., 2003; Stemper et al., 2003). One method involves plotting all responses from all specimens and drawing hypothetical curves that represent the outer and inner boundaries of the entire range in response. This methodology houses responses from each specimen inside the corridor, and hence, all “trajectories” will be within the corridor. “Eyeball” averaging of the curves is a second method. This has been used in early biomechanical studies, for example, during the development of neck responses from whiplash loading and chest impact responses applicable for frontal crashes (Mertz and Patrick 1971; Lobdell et al., 1973). The third is a more objective method wherein responses from individual specimens are scaled to a specific mass, in the present study to the small female, and deriving a mean response using the statistical principles of averaging and deviation techniques (Eppinger 1976). Deviations from the mean response can be used to illustrate the corridor in the form of mean and plus/minus a percentage of standard deviation curves. As reported by Yoganandan and Pintar, the Society of Automotive Enginneers in 1994 used plus/minus 0.5 standard deviation limits for characterizing the head impact response. (Yoganandan and Pintar 2003). However, it is more common to use one standard deviation. This process, adopted in the present study, has also been used for developing corridors in representing shoulder response in lateral and neck response in rear impacts (Mertz and Patrick 1971; Bolte et al., 2003; Stemper et al., 2003). Because corridors presented in figures 2-6 in the present study show one standard deviation limits in addition to the mean curve, response data from all specimens at all time intervals do not fall entirely within this corridor limit. To emphasize and ensure that plots are readable, only mean and plus/minus one standard deviations are presented in this paper.

A comparison of the male corridors derived in the previous study by Maltese et al.

with the corridors derived in the present study shows variations (more than 50 percent in some cases) in the force, acceleration, and deflections under different initial conditions. Differences arise because data were normalized using the body weight of the mid-size male in the Maltese et al. study in contrast to the 5th percentile weight for the corridors presented in the present investigation.

As expected, rigid flat wall test resulted in higher forces than padded flat wall tests,

pelvic offset tests produced greater loads in the pelvis than the thorax or abdomen, and


thoracic offset tests resulted in the highest forces in the thorax (Table 1). Forces in abdominal regions were lower than thoracic or pelvic forces (Figures 2-6) in all tests. These findings are secondary to the initial condition, i.e., presence of padding and leading position of a particular component in offset tests. Padding always increased the time duration and decreased forces and accelerations in all flat wall tests. Although peak deflections showed the same tendency in time duration, magnitudes were not consistent presumably because of the continued contact of the torso-pelvic region with the load wall.

It is true that when injuries occur, there is frequently a drop in load or alteration in

kinematics. However, this is true at the local level where injury occurs. For example, fracture of a spinal segment, whether a vertebral body or intervertebral disc, is always associated with a decrease in force (Yoganandan et al., 1989). Similarly, fracture of a rib is also always associated with a drop in force and a concomitant increase in deflection, usually bending (Yoganandan and Pintar 1998). However, corridors derived in this study represent the overall response of the chest to side impact. Localized trauma does not affect the overall response. This is depicted in figure 7, wherein data from all specimens are plotted for the padded low-speed flat wall initial condition. As can be seen from the plots, specimens with widely varying injury severity (AIS 4 and AIS 1) have almost similar magnitudes of force. Furthermore, there is no single dip that can be attributed to the range in the injury severity. There is also no appreciable change in the pattern of the curves although the injury severities in the entire spectrum ranged from none (AIS 0) to AIS 4, indicating the relative insensitivity of local trauma on the overall response. Similar responses were found in other conditions.

Figure 7: Thorax force histories on an individual specimen basis for the padded flat wall condition. Legend indicates the injury severity according to the AIS 1990 score. The order of plots in the legend is arranged in the ascending sequence of the peak force.

Using data provided in table 1, it is possible to assess the effect of parameters such as age, height, seated height, and body weight. For example, while there was no difference between males and females in the age parameter for the rigid flat wall condition (mean age 60 years), for the other initial conditions, the average age was seven years higher in males than females. The mean seated height was seven centimeters higher in males than females for the thoracic offset and eight centimeters higher in males than females for the padded flat wall conditions. The average overall height of the specimens was higher in males than females (seven centimeters for the padded and eight centimeters for the rigid flat wall conditions). For all conditions, the mean body weight was higher for male subjects (13 kg)

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than female subjects. Thus, male specimens had greater anthropometric measurements than female specimens. However, an examination of the age parameter suggests the following. It is well known that increasing age negatively influences bone quality and biomechanical strength. This is true for both genders. Because the age of the specimens used in this study did not show significant differences (fall within the same decade), variations in this parameter may have little effect on the outcome, i.e., corridors. It is difficult to derive conclusions with similar levels of confidence using other parameters such as seated height because of lack of literature information. Despite these limitations, the present effort of providing corridors is a good first step in side impact research.

Because specimens from both genders were tested under rigid and padded flat wall

and thoracic offset conditions, it was possible to compare data on the following basis. Magnitudes of mean peak variables (force, acceleration, deflection) for the 5th percentile female anthropometry derived in this study using both female and male specimens (Figures 2-6) when compared with data obtained using only female specimens (both sets scaled to 5th

percentile mass) revealed variations in biomechanical parameters. The percentage change was expressed as the normalized difference in the variable (force, deflection, or acceleration) between the computation using both genders and the computation using the female only gender; normalization was accomplished using the female-only data. Figure 8 illustrates differences in scaled peak biomechanical variables. Because each response starts at time zero (defined earlier) and ends approximately at the same time for each test condition, differences are maximized at the peak amplitude of each parameter. Forces varied from -21 to +13 percent, peak accelerations from -34 to +21 percent, and peak deflection from -26 to +12 percent. Change depended on initial condition. These analyses indicate that corridors obtained in the present study using both genders (for mass scaling) may need appropriate modulation for the female gender, and, to achieve this objective, additional sled tests using female PMHS are necessary. However, the present corridors may be a used as a first approximation in the assessment of performance of side impact dummies. Because pelvic offset tests were not conducted using females, it is not possible to estimate variations in the three biomechanical parameters for this condition. The present findings underscore the need to pursue further testing and analysis for improvements in crashworthiness evaluations.

A logical extension of the analysis aimed at evaluating the efficacy of mid-size (50th percentile) male corridors shows similar differences. This issue is not the subject matter of the present study, and hence, not explored further. Therefore, it may be necessary to revisit the earlier analysis and derive corridors using male-only data (Maltese et al., 2002). This ongoing exercise will be the subject of another presentation.

As can be seen from figure 8, in general rigid flat wall tests demonstrated the highest

differences in mean peak forces (based on both genders versus the female-only analysis). Variations in forces were similar and within ten percent between padded and offset tests. Although spinal accelerations (with the exception of the upper acceleration in the offset condition) showed similar magnitudes of variations (within -15 percent), the polarity reversed in the lower spine for padded flat wall condition. Pelvic accelerations also showed a similar tendency, i.e., increased for rigid wall and offset and decreased for padded flat wall conditions. Regarding differences in mean peak deflections, thoracic offset tests showed increases (up to 12 percent), and flat wall tests demonstrated decreases up to 26 percent. These analyses clearly underscore non-uniform variations in biomechanical parameters.


Figure 8: Percentage changes in magnitudes of mean peak force (top), acceleration (middle), and chest deflection (bottom) for three types of initial conditions (see legend) between data obtained using only female specimens, and male and female specimens. See text for details.

Although samples are limited, observed variations may be explained by gender-

based anatomical, physiological, and biomechanical considerations. For example, the pelvic anatomy of females is different from males (Agur and Lee 1991). The ilio-pubic ramus is thinner and smaller in females than males (Cesari and Ramet 1982). The cross-sectional area of the pelvis is also smaller in the female (Haffner 1995). Furthermore, mass and geometrical distributions of various body segments are different between two genders (CFR 2000). While the mass of the pelvis of the 5th percentile female is 61 percent of the mid-size male, the lateral-lateral distance between the bilateral greater trochanters is 94 percent of the mid-size male anthropometry (NHTSA 1983). From a mechanical perspective, advancing age decreases bone strength, and with reference to the present study, this includes vertebrae, pelvis, and thoracic rib cage. A majority of injuries sustained by

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specimens in this study were to skeletal structures. While studies quantifying the bone mineral content and its relation to the strength of spine and hip are available, paucity of data exist regarding rib strength, a common source of skeletal trauma in side impacts (Yoganandan et al., 1988; Yoganandan and Pintar 1998). Dorsal spine vertebrae and ribs have significant cancellous components. Although the cortical bone does not show significant age or gender preference in strength characteristics, mechanical properties of the cancellous bone in the musculoskeletal system demonstrate age and gender bias (Yamada 1970; Riggs et al., 1981). This is because the human trabecular architecture undergoes changes due to aging in both genders. However, the female gender is more susceptible to reduced bone mineral because of physiological processes such as decrease in the osteoblastic activity and postmenopausal stage of life. In fact, Cesari and Ramet grouped side impact responses based on gender and suggested different thresholds for males and females (Cesari and Ramet 1980). Statistical analysis was limited in this previous study due to sample size constraints. The present study also has the same limitation. Female specimens should, therefore, be used to develop corridors for this population group to include accurate representations of these multifaceted factors. Further research is needed that may lead, in particular, to check existing and/or develop new scaling methodologies for bio corridors and injury criteria between varies sizes and gender. Data provided in this paper, however, are a good first step, and the results are useful as a first estimate in the evaluation of the performance of small female dummies. CONCLUSIONS

This study was designed to develop side impact corridors applicable to the small female anthropometry. Sled tests were conducted using post mortem human subjects at a velocity of 6.7 m/s. Adopting signal processing and mass scaling methods, small female-specific corridors (expressed as mean plus/minus one standard deviation) were derived for forces, accelerations, and chest deflections at three levels for initial conditions that included padded and rigid and flat wall and offset conditions. Reasons for using gender-specific tests for the derivation of corridors are discussed based on anatomical, physiological and biomechanical considerations. These time-history corridors will assist in the assessment of anthropomorphic test devices used in crashworthiness evaluations. ACKNOWLEDGMENTS This study was supported in part DOT NHTSA DTNH22-03-H-04117 and VA Medical Research. The assistance of Drs. Peter G. Martin, Shashi Kuppa, and Rolf H. Eppinger are acknowledged. REFERENCES Agur, A. M. R. and Lee, M. L. Grant's Atlas of Anatomy. (1991), Wiiliams and Wilkins. Bolte, J., Hines, M., Herriott, R., McFadden, J. and Donnelly, B. “Shoulder impact and injury

due to lateral and oblique loading.” Stapp Car Crash J 47: (2003) 35-53. Cesari, D. and Ramet, M. Evaluation of pelvic fracture tolerance in side impact. Stapp Car

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Yamada, H. Strength of Biological Materials. (1970) Baltimore, MD, Williams & Wilkins. Yoganandan, N., Myklebust, J. B., Wilson, C. R., Cusick, J. F. and Sances, A., Jr.

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Yoganandan, N. and Pintar, F. “Biomechanics of temporo-parietal skull fracture.” Clinical Biomechanics 18(1): (2003) 1-27.

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small female-specific biomechanical corridors ......ircobi conference – graz (austria) september...

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