Annals of Biomedical Engineering (Ó 2015) DOI: 10.1007/s10439-015-1255-1
Measurement of Head Impact Due to Standing Fall in Adults Using Anthropomorphic Test Dummies MARZIEH HAJIAGHAMEMAR,1 MORTEZA SEIDI,1 JAMES R. FERGUSON,2 and VINCENT CACCESE1 1
Advanced Biomechanics Laboratory for Injury Reduction and Rehabilitation, Mechanical Engineering Department, University of Maine, Orono, ME, USA; and 2Alba-Technic LLC, Winthrop, ME, USA (Received 9 September 2014; accepted 14 January 2015) Associate Editor Stefan M Duma oversaw the review of this article.
Abstract—The kinematics and kinetics of head impact due to a standing fall onto a hard surface are summarized. Head injury due to impact from falls represents a significant problem, especially for older individuals. When the head is left unprotected during a fall, the impact severity can be high enough to cause significant injury or even death. To ascertain the range of head impact parameters, the dynamic response was captured for the pedestrian version of the 5th percentile female and 50th percentile male Hybrid III anthropomorphic test dummies as they were dropped from a standing position with different initial postures. Five scenarios of falls were considered including backward falls with/without hip flexion, forward falls with/without knee flexion and lateral falls. The results show that the head impact parameters are dependent on the fall scenario. A wide range of impact parameters was observed in 107 trials. The 95% prediction interval for the peak translational acceleration, peak angular acceleration, peak force, impact translational velocity and peak angular velocity are 146–502 g, 8.8–43.3 krad/s2, 3.9–24.5 kN, 2.02– 7.41 m/s, and 12.9–70.3 rad/s, respectively. Keywords—Head impact, Falls, Head injury criteria, Acceleration, Head protection device.
INTRODUCTION Falls are the leading cause of nonfatal unintentional injuries and in 2008 alone more than 8.5 million people were treated in emergency departments for fall-related events.19 Among the most devastating outcomes of falling is traumatic brain injury (TBI) that can result in serious and even fatal outcomes. TBI is a contributing factor to nearly a third (30.5%) of all injury-related
Address correspondence to Marzieh Hajiaghamemar, Advanced Biomechanics Laboratory for Injury Reduction and Rehabilitation, Mechanical Engineering Department, University of Maine, Orono, ME, USA. Electronic mail: [email protected]
deaths in the United States.7 Different types of head injury may occur due to a severe fall impact. Head injury can be predicted using a variety of measures that depend on injury type. The following parameters can be used for the assessment of head injury potential: peak resultant translational acceleration of the head center of gravity,2 rotational velocity and rotational acceleration,21 peak of impact force on the skull,1,12,25 impact duration using measures such as the head injury criteria (HIC15),2,22,24 and brain response using measures such as pressure, strain and strain rate of the brain tissue.3 Outside of actual human testing, which is not viable for severe head impact fall research, alternative methods used to assess human fall response by physical or numerical means must be interpreted with caution. The Hybrid III anthropomorphic test dummy (ATD) plays an important role in the ability to produce accurate fall simulation data without the use of human subjects. Mertz compared the response of the 50% male Hybrid III head to cadaver data16 and showed that the Hybrid III head is viable for impact comparisons with the cadaver test data. Fall studies using ATDs have been conducted by multiple researchers. Ibrahim and Margulies13 used ATDs to investigate low-heights falls (1, 2, and 3 ft) in toddlers and infants and found that the toddler is likely less vulnerable to skull fracture but more vulnerable to neurological impairment compared with the infant. Heller et al.11 dropped 1- and 3-year old Hybrid III dummies onto carpet and concrete surfaces from various heights where the impact velocities varied from 3.8 to 5.5 m/s to investigate the head impact response due to falls. They reported that serious head injuries may occur in 1- and 3-year-old children even from a short distance fall. In 2007, Nagata and Ohno18 conducted a study on elderly falls using a test dummy
Ó 2015 Biomedical Engineering Society
HAJIAGHAMEMAR et al.
that represented an average Japanese male and found that the rotations of ankle and hip joints can affect the severity of head injuries by decreasing the head impact velocity and by extending the impact time. Another study was performed to ascertain head impact values for falls from standing and kneeling positions using a 50th percentile male Hybrid-III and reported the average peak acceleration magnitude of 302 g, and a HIC15 of 1487.4 Authors in another work investigated the influence of the body on the effective mass involved during fall-related head impacts using Hybrid-III ATDs and reported that the effective mass can be estimated as the mass of head plus half of the neck mass.23 Some studies attempted to numerically simulate real life fall accidents using a multibody model numerical analysis program MADYMOTM.6,20 O’Riordain et al.20 investigated four cases of falls in persons ranging from 11 to 76 years old with injuries ranging from contusion to skull fracture. Their results showed a maximum acceleration ranging from 243 to 435 g, while the angular accelerations ranged from 17.6 to 43.5 krad/s2. Doorly et al.6 investigated 10 real life cases where falls occurred with the subjects standing; the calculations resulted in the peak translational acceleration between 189 and 456 g, angular acceleration of 7.4–49.2 krad/s2 and HIC15 ranging from 511 to 5951. It was observed that the multibody model gave good representations of the real life cases, however, when the case was complex, the results were not ideal and showed a high variance due to head contact characteristics and initial conditions. In another study, Hamel et al.9 used pedestrian multibody models to simulate backward falls from a standing position. They simulated free falls and blocked falls and studied the influence of impact velocity (3.5–6.5 m/s) and impact surface. They concluded that the mechanism of skull fracture in the case of a backward fall from a standing height is influenced by impact velocity, impact surface, cortical thickness and rigidity. In addition to ATD tests and numerical simulations, many researchers conducted cadaveric studies to investigate the response of the human head to impact. Zhang et al.26 used a controlled drop of the cadaver head instrumented with an accelerometer array onto padding surfaces of 40D and 90D durometer. They reported angular accelerations vs. impact translational velocities ranging from an average of 9 krad/s2 at an impact velocity of 2.5 m/s to 44 krad/s2 at an impact velocity of 5.5 m/s based on their linear curve fit for the 90D surface. Hardy et al.10 investigated the response of the human head to impact using cadavers at impact velocities between 2.9 and 3.9 m/s. When using a CFC 1000 filter, their results of the peak translational acceleration and HIC15 for the unprotected occipital impact tests ranged from 153 to 408 g and 372 to 2540. Since there is an anatomical equivalency between cadaver skull bone and
in vivo human, cadaveric tests are usually used to determine the tolerance of the head to skull fracture.25 Yoganandan et al.25 investigated levels of force required to fracture the skulls of 12 unembalmed cadavers using a hemispherical impactor with a 48 mm radius. They performed impact tests on various locations of the skull at a rate of 7.1–8.0 m/s. They noted failure loads ranging from 4.5 to 14.1 kN (11.9 ± 0.9 kN). Also, Allsop et al.1 carried out temporo-parietal impact tests on 31 unembalmed cadaver heads with two types of flat rigid impactors, one circular and 2.54 cm in diameter, the other a rectangular plate 5 9 10 cm. Fracture force for the small circular plate ranged between 2.5 and 10.0 kN while for rectangular plates it ranged between 5.8 and 17.0 kN. They concluded that the skull’s peak force threshold increases with greater surface area of the impacting force and fracture is more a function of stress (force/area) than force alone. Thus, an impact with a flat ground surface—more like a case of falling—requires higher forces to cause fracture as demonstrated in the rectangular plate. In both of the above studies, the impactors hit the stationary cadaver heads, while in the fall impact case a moving head strikes the ground. There are some cadaveric studies that more closely represent falling cases where the head is not stationary. Hodgson and Thomas12 dropped embalmed human male cadaver heads at impact velocities ranging from 1.6 to 4.7 m/s to assess rear, side and front head impacts and reported 5.6–17.8 kN (10 ± 4.9 kN) for fracture threshold. The aim of the present study was to quantify the kinematics and kinetics of head impact due to a standing fall onto a hard surface. Since it is not possible to assess the response of humans falling onto hard surfaces in the laboratory due to safety concerns, the pedestrian versions of a Hybrid III 5th percentile female and 50th percentile male ATD were used. Several fall scenarios were initiated from a frontal, sideward and backward orientation and the responses of the ATDs were quantified. The longterm agenda of this research team is to develop an informed design of non-stigmatizing head protection for fall-related head impact.
MATERIALS AND METHODS ATDs were used to quantify the kinetics and kinematics of head impact in adults falling from a standing position. To this end, a HumaneticsTM Hybrid III 5th percentile female ATD with 1.5 m height and 49 kg mass and a 50th percentile male ATD with 1.75 m height and 78 kg mass were employed. The hip of the standard ATD was replaced with a movable pedestrian style hip. Ball joints at the femur to pelvis connection of the pedestrian style hip allowed hip rotation, abduction/adduction and flexion/extension to simulate
Head Impact Due to Standing Fall in Adults
various human postures. The initial orientation of the ATD for forward, backward and sideward falls of the female ATD is shown in Fig. 1a; these orientations were the same for the male ATD. A force plate model AMTI LG6-4-2000 was used to measure impact force and an auxiliary platform was laid out so that only the head of the ATD hit the force plate. Several trials of backward falls were also performed using the laboratory floor as the impact surface to assess the influence of the force plate on the impact response. The head of the ATD was instrumented with an accelerometer array consisting of four tri-axial accelerometers (PCB, Depew, NY) with a measuring range of ±500 and ±5000 g in each direction, for female and male ATD, respectively. The accelerometer array mounted inside the head form is illustrated in Fig. 1b. Additionally, three ARS PRO-8K angular rate sensors (DTS, Seal Beach, CA) were mounted on perpendicular faces of the accelerometer at the center of gravity of the heads to directly capture the angular velocity. Angular acceleration for the head center of gravity was calculated from nine of the twelve acceleration signals by the method described by Kang et al.14 using the following equations: x_ x ¼
aBackz aCGz aCGy aTopx þ 2d1 2d3
ð1Þ
x_ y ¼
aTopy aCGx aCGz aSidez þ 2d3 2d2
ð2Þ
aCGx aBackx aSidex aCGy ; 2d1 2d2
ð3Þ
x_ z ¼
where d1 = 47.36 mm, d2 = 26.40 mm for both ATDs and d3 = 77.08 mm, d3 = 88.05 mm for the female and male ATDs respectively. A motion capture system from ViconTM recorded the whole body movement at 200 Hz using 10 cameras distributed about the laboratory. A total of 35 markers based on the plug-in gait full body model from ViconTM were attached to the ATD. Two additional markers were added to the pelvis to supplement the model data when the body obstructed some of the markers from the cameras. The ATDs were kept in a standing posture using an electro-mechanical release mechanism with a mass of 0.072 kg that was triggered by the Vicon MX giganet. Also, a digital video camera was used to record the experiments. Data of the force plate and accelerometers were collected at a sampling rate of 20 kHz using USB1608FS-Plus DAQ (Measuring Computing, Norton, MA). The data were filtered using a CFC1000 (Chanel Frequency Class) filter that is a fourth order Butterworth filter with a 1650 Hz cutoff frequency. Data were collected and processed by a PC using a NI
LabVIEW software (National Instrument, Texas, USA) with help of MCC ULx library for NI LabVIEW, then analyzed using MATLABÒ. Five scenarios of falls were studied as illustrated in Fig. 2. Case 1 (BK-NHB) was a backward fall where the hip joints did not rotate before the head hits the ground; this happens when a person does not react quickly to the fall. The motion of this case is similar to the dynamics of a single inverted pendulum. Case 2 (BK-HB) was a backward fall with hip flexion which the hip hits the ground first followed by the head impact. The dynamics of this case is similar to the behavior of a double inverted pendulum. Case 3 (FW-KB) was a forward fall with the knee flexion during which the knees hit the ground first followed by the head. Case 4 (FW-NKB) was a forward fall with the knee joints fixed. Case 5 (LAT) was a sideward fall from a standing position where the shoulder typically contacted the ground first and then the head. In running the LAT experiments it was observed for the male ATD that the head did not make contact with the ground in this orientation due to the size of the ATD shoulder, therefore, the arm of the male ATD was removed in the LAT fall case. Although there are numerous head injury indicators, the HIC15 was used as the injury predictor for the purpose of this paper since it is a common measure used in many published head impact studies and is one that includes the influence of impact duration. The HIC15 measure was calculated using the formula: ( 2:5 ) 1 t 2 aðtÞdt HIC15 ¼ max ðt2 t1 Þ ; ð4Þ t 2 t 1 t1 where aðtÞ is the resultant translational acceleration as a function of time and t1 and t2 define the time interval that maximizes the HIC15. The HIC15 criterion was computed from the acceleration signature measured in g’s during impact and the maximum time duration of (t2 2 t1) is limited to a maximum value of 15 ms. The time interval (t2 2 t1) is referred to as the HIC15 duration and is selected to maximize the HIC15 value while the impact duration is defined as the time difference between the leading and trailing edges of the primary translational acceleration pulse. The HIC15 was first proposed by Versace24 and then modified by NHTSA 1998.15 According to Mertz et al.,17 a HIC15 of 700 represents a 5% risk of a severe injury. Chinn et al.5 found that of all the parameters analyzed in their study of motorcycle helmets, the HIC15 provided the most reliable head injury severity prediction with a correlation coefficient of r = 0.8 and the HIC15 of 1000 and 1500 correspond to a moderate and serious head injury, respectively.
HAJIAGHAMEMAR et al.
FIGURE 1. Experimental setup, (a) test configuration using the Hybrid III 5th percentile female dummy, (b) the configuration of accelerometers and angular rate sensors.
RESULTS For the five fall scenarios of the male and female ATD, a total of 107 trials were conducted. Different ATD response kinematics were observed. In the first case (BK-NHB), the hip joints did not rotate. Shortly after the pelvic struck (within 25–45 ms), the torso contacted the ground while in the second case (BKHB) a 350–450 ms time interval was observed between pelvic and torso contact to the floor. The head impact occurred within 20 ms after the torso contacted the floor for both backward scenarios while this time interval was 5 ms for the forward scenarios. Sideward fall scenarios showed different kinematics, the pelvic struck the ground followed by the arm/shoulder and then the head. The time intervals between arm/shoulder contact and head impact were observed to be approximately 40 ms for the female ATD and 30 ms for the male ATD. A summary of measurements of the female and male ATD including results of the peak translational acceleration, peak angular acceleration, impact translational velocity, peak angular velocity, impact duration and peak force are given Tables 1 and 2. Here and throughout this paper, the peak translational acceleration and peak angular acceleration denote the maximum of translational and angular acceleration signals
after impact occurred. The impact translational velocity represents the maximum of translational velocity signal immediately prior to impact and the peak angular velocity denotes the maximum angular velocity signal either immediately prior to impact or after impact, whichever is greater. The first row in the table gives the number of trials conducted for each type of fall. The number after the plus sign shows the number of trials from a preliminary study on a backward fall of the Hybrid III 5th percentile female ATD that did not include acquisition of angular acceleration and angular velocity.8 The summary provides the mean of the peak values, standard deviation, and range of the peak values (given in parentheses) for the kinetic and kinematic measures. As seen in Tables 1 and 2, the peak translational acceleration of the male ATD was on average greater than the female ATD due to the higher height and mass of the male dummy. On the other hand, angular accelerations were similar for both ATDs except during lateral falls. This is attributed to the change in dynamic response caused by the shoulder of the male dummy being disassembled and removed during the lateral testing. A typical translational and angular acceleration magnitude time-history truncated at a time of 3.5 ms after floor impact is shown in Fig. 3 for the various fall conditions for both the female and male ATD. The
Head Impact Due to Standing Fall in Adults
FIGURE 2. The scenarios of standing fall. TABLE 1. Summary of results for female ATD. Condition/measure Number of trials Trans. acc. (g) Ang. acc. (krad/s2) Ang. vel. (rad/s) Trans. vel. (m/s) Impact duration (ms) Peak force (kN) HIC15 HIC15 duration (ms)
BK-NHB
BK-HB
FW-KB
FW-NKB
LAT
12 + 7 368 ± 31 (308–411) 28.8 ± 3.4 (24.4–35.2) 48.7 ± 4.5 (37.5–54.2) 4.80 ± 0.22 (4.37–5.19) 2.5 ± 0.2 (2.3–3.0) 14.7 ± 1.1 (12.4–16.4) 2173 ± 298 (1652–2678) 1.4 ± 0.1 (1.3–1.6)
15 + 14 289 ± 47 (157–361) 23.6 ± 4.5 (13.5–33.8) 32.7 ± 5.7 (20.5–45.5) 3.78 ± 0.53 (2.49–4.84) 2.5 ± 0.3 (2.2–3.4) 11.5 ± 2.1 (6.0–14.4) 1218 ± 425 (331–2156) 1.4 ± 0.1 (1.2–1.8)
7 243 ± 26 (204–278) 28.9 ± 5.9 (19.8–35.9) 44.1 ± 5.7 (34.3–52.8) 4.39 ± 0.63 (3.53–5.40) 3.9 ± 0.7 (3.0–4.8) 9.8 ± 1.4 (7.5–11.6) 888 ± 207 (682–1312) 1.8 ± 0.3 (1.3–2.3)
5 319 ± 26 (293–354) 29.7 ± 4.0 (23.6–34.6) 46.8 ± 3.1 (41.7–49.9) 5.81 ± 0.28 (5.55–6.28) 4.1 ± 0.5 (3.4–4.7) 12.7 ± 1.4 (11.0–14.6) 1846 ± 213 (1608–2111) 1.8 ± 0.1 (1.6–2.0)
11 206 ± 79 (64–305) 22.2 ± 10.6 (4.5–37.8) 30.1 ± 10.1 (17.2–43.8) 2.90 ± 0.92 (1.37–4.42) 3.1 ± 0.7 (2.1–4.8) 7.6 ± 3.0 (2.6–11.6) 598 ± 417 (41–1431) 1.5 ± 0.3 (1.3–2.3)
HAJIAGHAMEMAR et al. TABLE 2. Summary of results for male ATD. Condition/measure Number of trials Trans. acc. (g) Ang. acc. (krad/s2) Ang.vel. (rad/s) Trans. vel. (m/s) Impact duration (ms) Peak force (kN) HIC15 HIC15 duration (ms)
BK-NHB
BK-HB
FW-KB
FW-NKB
LAT
10 451 ± 38 (390–524) 29.2 ± 5.8 (19.6–41.0) 58.9 ± 5.7 (48.3–67.6) 6.75 ± 0.27 (6.31–7.11) 3.0 ± 0.4 (2.6–4.0) 22.8 ± 2.1 (19.3–26.6) 4142 ± 536 (3182–5046) 1.6 ± 0.1 (1.4–1.8)
9 295 ± 89 (198–447) 19.3 ± 9.6 (10.7–38.2) 34.3 ± 17.5 (17.0–64.4) 4.85 ± 1.33 (3.29–6.80) 3.4 ± 0.4 (3.0–4.0) 14.9 ± 4.6 (9.4–22.2) 1826 ± 1342 (530–4234) 1.7 ± 0.1 (1.6–1.8)
5 402 ± 69 (311–489) 23.8 ± 15.6 (8.9–46.2) 37.5 ± 24.3 (16.0–75.8) 6.46 ± 1.29 (4.43–7.90) 3.2 ± 0.3 (2.8–3.5) 20.3 ± 3.7 (15.4–25.6) 3286 ± 1136 (1711–4691) 1.7 ± 0.2 (1.6–1.9)
5 439 ± 118 (294–564) 21.1 ± 13.7 (7.0–40.2) 32.1 ± 16.8 (15.6–53.6) 6.68 ± 1.20 (4.82–7.64) 3.1 ± 0.5 (2.6–3.8) 21.6 ± 6.1 (14.2–29.6) 4072 ± 2247 (1353–6435) 1.6 ± 0.1 (1.4–1.8)
7 355 ± 49 (264–418) 34.3 ± 5.8 (30.4–46.9) 56.4 ± 1.6 (54.1–58.6) 5.12 ± 0.66 (3.91–6.10) 2.9 ± 0.3 (2.4–3.3) 17.1 ± 2.2 (12.6–19.0) 2128 ± 585 (996–2745) 1.5 ± 0.2 (1.2–1.7)
black boxes in the figure represent the start and end times used for the HIC15 computation. Correlation between impact parameters like impact translational velocity and peak translational acceleration, peak angular velocity and acceleration, impact translational velocity and peak force; and peak translational and angular acceleration were established and the 95 and 99% prediction intervals were computed (Figs. 4, 5, 6, and 7). In addition to testing on the force plate, several trials of impact onto the laboratory floor were conducted to ascertain whether testing the head impact on the force plate can be used to simulate fall-related head impact onto a common commercial building surface. Peak translational accelerations of head impact onto the force plate were 289 ± 47 g for the second scenario (BK-HB) of the female ATD. This was compared to the same case striking the concrete floor covered with a layer of thin carpeting with no padding that resulted in peak translational accelerations of 311 ± 9 g. Accordingly, test results showed that impact onto the force plate are in the same range as an impact onto a building surface such as carpeted concrete floor.
DISCUSSION The response of the female and male ATDs was quantified to describe the kinematics and kinetics of different scenarios of falls from standing focusing on head impact. The results show that the head impact parameters were dependent on the fall direction and type (Tables 1, 2). The BK-NHB case was the most severe case among all fall scenarios studied herein. The highest mean value of peak translational acceleration (368 g for female and 451 g for male), impact force
(14.7 kN for female and 22.8 kN for male), HIC15 value (2173 for female and 4142 for male) and peak angular velocity (48.7 rad/s for female and 58.9 rad/s for male) was observed for this type of fall. For a lateral fall, the mean values of peak translational acceleration (206 g for female and 355 g for male) and HIC15 (598 for female and 2128 for male) were relatively lower compared to the backward and forward cases. The mean value of impact duration and HIC15 duration for all types of falls were in similar ranges but slightly longer impact duration was observed for forward fall cases. In addition to the study of the effect of fall scenarios on head impact, the bounds of the head impact parameters that can arise from a standing fall were assessed. For this reason, all data were grouped together and the mean and standard deviation of peak translational acceleration, peak angular acceleration, impact force, impact translational velocity, peak angular velocity and HIC15 are 324 ± 89 g, 26.0 ± 8.6 krad/s2, 14.2 ± 5.1 kN, 4.71 ± 1.34 m/s, 41.6 ± 14.3–60 rad/s, and 1982 ± 1290, respectively. Comparisons with Previous Studies A major task in the development of a methodology to evaluate fall protection devices is to assess the bounds of the impact parameters that occur during an unprotected fall. Also, establishing correlations between the impact parameters can be used to improve numerical analysis of real life fall incidents and to develop simplified experimental procedures for testing of head protection devices. To that end, correlation between impact parameters were established and the 95 and 99% prediction intervals were also plotted (Figs. 4, 5, 6, and 7) to provide bounds for estimations of the head impact parameters due to fall. Prior
Head Impact Due to Standing Fall in Adults
(c) 600
Female ATD BK-NHB BK-HB FW-KB FW-NKB LAT
400
Translational Acceleration (g)
Translational Acceleration (g)
(a) 500
300 200 100
Male ATD BK-NHB BK-HB FW-KB FW-NKB LAT
500 400 300 200 100 0
0 0
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(d)
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BK-NHB BK-HB FW-KB FW-NKB LAT
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(b)
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FIGURE 3. Head impact time histories for various fall conditions. (a) Translational acceleration for female ATD, (b) Translational acceleration for male ATD, (c) Angular acceleration for female ATD, (d) Angular acceleration for male ATD.
methods employed in other studies include numerical simulations using the MADYMOTM program,6,20 cadaver tests,10,12,26 and ATD testing4 were also compared to the current ATD study to verify the bounds of the head impact response due to a standing fall (Figs. 4, 5, 6, and 7). The impact translational velocity vs. the peak translational acceleration and the peak angular acceleration vs. the peak angular velocity are shown in Fig. 4. Plotted in this figure are the 95% (long dash) and 99% (short dash) prediction interval bounds and linear least square fit (solid line) of the ATD data. The linear fit passing through the origin resulted in a coefficient of determination of 0.73 and 0.57 for the results shown in Figs. 4a and 4b. Results from MADYMOTM reconstruction6,20 are presented in both the translational and angular plots as shown in Figs. 4a and 4b, respectively. Results of cadaver studies conducted by Zhang et al.26 and Hodgson and Thomas12 are also presented but only in the translational plot (Fig. 4a), since angular velocity at impact was not reported in these studies. Zhang et al.26 dropped the cadaver heads onto 40D and 90D durometer but only
the data of the harder surface (90D) was used in the comparison. The numerical model and cadaver test results fit within the 99% prediction interval for peak translational acceleration and with the exception of 3 outliers the same can be said for angular acceleration. The numerical models tended to compute a lower peak translational acceleration than the ATD and cadaver test results. Angular acceleration computed in the numerical models tended to be greater than the values obtained in the ATD tests. Peak translational acceleration vs. the HIC15 value is presented in Fig. 5 where a power law fit of the ATD test resulted in a coefficient of determination of 0.97 indicating a strong correlation with exponent 2.42. This result was as expected due to the nature of the HIC15 computation. The current ATD test results are compared to MADYMOTM simulations,6,20 cadaver test10 and another ATD study.4 The MADYMOTM simulations6,20 tended toward higher HIC15 values that indicate a longer impact duration in the computation (Fig. 5). The cadaver studies by Hardy10 tended toward lower HIC15 values indicating shorter impact
HAJIAGHAMEMAR et al. 35
700
Backward Fall-ATD Forward Fall-ATD Lateral Fall-ATD Zhang-Cadaver Hodgson-Cadaver Doorly-MADYMO O'Riordain-MADYMO
Translational Acceleration (g)
600 500
y = 67.967x R² = 0.7287
Backward Fall-ATD Forward Fall-ATD Lateral fall-ATD Hodgson-Cadaver Doorly-MADYMO O'Riordain-MADYMO
30
Impact Force (kN)
(a)
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25
y = 3.0455x R² = 0.8127
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Backward Fall-ATD y = 0.6091x Forward Fall-ATD R² = 0.5727 Lateral Fall-ATD Doorly-MADYMO O'Riordain-MADYMO
50
60 2
40 30 20 10 0 0
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Angular Velocity (rad/s) FIGURE 4. Impact translational velocity vs. peak translational acceleration, linear regression line (solid line), 95% (long dash) and 99% (short dash) prediction bounds for ATD tests compared to studies by Zhang,26 Hodgson,12 Doorly,6 and O’Riordian.20 (a) Translational, (b) Angular.
Backward Fall-ATD Forward Fall-ATD Lateral Fall-ATD Hardy-Cadaver Doorly-MADYMO O'Riordain-MADYMO Caccese-ATD
6000 5000 4000
HIC
Angular Acceleration (krad/s )
2
Angular Acceleration (krad/s )
(b)
FIGURE 6. Peak impact force vs. impact translational velocity, linear regression line (solid line), 95% (long dash) and 99% (short dash) prediction bounds for ATD compared to Hodgson,12 Doorly,6 and O’Riordian.20
y = 0.0014x2.4218 R² = 0.971
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Translational Acceleration (g) FIGURE 5. HIC15 vs. peak translational acceleration and power regression curve for ATD compared to Hardy,10 Doorly,6 O’Riordian,20 and Caccese.4
Backward Fall-ATD Forward Fall-ATD Lateral Fall-ATD Zhang-Cadaver Doorly-MADYMO O'Riordain-MADYMO
50 40 30 20 10 0 0
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Translational Acceleration (g) FIGURE 7. Peak translational acceleration vs. peak angular acceleration for ATD compared to Zhang,26 Doorly,6 and O’Riordian.20
durations. The other ATD study also indicates similar impact duration and HIC15 value4 (Fig. 5). Peak of impact force is often used to predict skull fracture since a high level of force can cause failure of the cranial bone. The peak impact force vs. the impact velocity for the ATD tests, cadaver tests by Hodgson12 and numerical analyzes by Doorly6 and O’Riordain20 are shown in Fig. 6. Also plotted in the figure are the 99 and 95% prediction intervals for the ATD test data. It was observed on average that the ATD test data resulted in higher levels of impact force at an equivalent impact velocity. In the current study, the head impact forces ranged from 2.60 to 16.37 kN for female ATD tests and from 12.6 to 29.6 kN for the male ATD tests. Impact forces for occipital impacts due to
Head Impact Due to Standing Fall in Adults
backward falls and frontal impacts due to forward falls were 69% and 46% higher than lateral impacts for female ATD and 22 and 12% higher than lateral impacts for male ATD. The impact forces in many of our ATD tests were above the fracture thresholds1,12,25 (Fig. 6). Therefore, based on the ATD standing fall tests, probability of skull fracture is high when the fall is unprotected and occurs onto a hard surface, especially in the FW-NKB and the BK-NHB fall scenarios. The relationship between peak angular and translational acceleration is important to several types of injury predictions and depends greatly on impact orientation and location that are highly variable in real life cases. The peak angular vs. translational acceleration for the ATD testing as well as the cadaver testing by Zhang26 and the numerical analyzes by Doorly6 and O’Riordian20 are shown in Fig. 7. Little correlation was observed in the results, as expected, with a coefficient of determination of 0.42, 0.37, and 0.54 for the backward, forward and lateral ATD scenarios. The numerical analysis studies performed for various fall scenarios showed a similar distribution. The cadaver study by Zhang26 showed a stronger correlation with a coefficient of determination of 0.90 since the impact parameters were controlled in the study. This result indicates that when conditions are controlled, as they are in drop tower testing, there is typically a defined relationship between the peak angular and translational acceleration. In interpretation of the results it is noted that a limitation of this study is a lesser biofidelity of the torso and neck compared to the head for fall study that may affect the results. The authors in another work4 studied the effect of neck stiffness by comparing a standard (70–80 Shore A durometer) and lower stiffness (35 Shore A durometer) neck in head impacts using Hybrid-III head and neck assembly. The results showed that the influences of neck stiffness were less than 10% for rear and frontal impacts and less than 17% for lateral impacts for both translational and angular accelerations.4 In addition, as mentioned in the method section, the arm of the male ATD was removed in the LAT fall case. This condition affects the results especially on head rotational movement as the ATD hits the shoulder instead of arm. Also, it should be noted that the results of this study do not include the human tissue response (like skull and brain tissues) to an impact and actual human reaction to a fall. This study relies on the ATD design to simulate the human response and therefore should be used with caution.
CONCLUSION This paper presents the response of a 5th percentile female and 50th percentile male pedestrian ATD sub-
ject to standing falls. Conditions of forward, backward, and sideways falls were studied. Test results were compared to cadaver, numerical analyzes and other ATD studies. The female ATD generally resulted in less severe impact attributed to the lower drop height and subsequently lower impact velocity. Mean peak forces were also in a range that indicates potential for head injury including skull fracture. Among all the fall cases, the unprotected BK-NHB fall had the highest mean value of peak translational acceleration, impact force and HIC15 value and was the fall scenario with the potential for the most severe injury. Correlation between impact parameters were established and the 95 and 99% prediction intervals were obtained to provide bounds for estimation of the head impact parameters. The results presented can be used to improve numerical analysis of real life fall incidents and to develop simplified experimental procedures for testing of head protection devices. Although there are major differences in methodology, when ATD testing was compared to cadaver testing and numerical analyzes, the numerical and physical studies lay within the 99% prediction bounds of the 107 ATD tests reported.
ACKNOWLEDGMENTS The authors gratefully acknowledge support from the Maine Technology Institute under Grant Number MTAF-3001 for funding of the laboratory facility, the National Institutes of Health/National Institute on Aging under grant number NIH 5R44AGO33936-03, the George and Caterina Sakellaris Graduate Fellowship and the National Science Foundation under Award No. 1417120.
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Allsop, D. L., T. R. Perl, and C. Y. Warner. Force/ deflection and fracture characteristics of the temporoparietal region of the human head. In: Proceeding of 35th Stapp Car Crash Conference. Warrendale, PA: Society of Automative Engineers, 1991, pp. 269–278. 2 Auer, C., M. Schonpflug, G. Beier, and W. Eisenmenger. An analysis of brain injuries in real world pedestrian traffic accidents by computer simulation reconstruction. Zurich: Proceeding of International Society of Biomechanics XVIIIth Congress, pp. 680–686, 2001. 3 Bain, A. C., and D. F. Meaney. Tissue-level thresholds for axonal damage in an experimental model of central nervous system white matter injury. J. Biomech. Eng. 122:615–622, 2000. 4 Caccese, V., J. Ferguson, J. Lloyd, M. Edgecomb, M. Seidi, and M. Hajiaghamemar. Response of an impact test apparatus for fall protective headgear testing using a Hybrid-III head/neck assembly. Exp Tech. 2014. doi: 10.1111/ext.12079.
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Chinn, B., B. Canaple, S. E. Derler, D. Doyl, D. Otte, E. Schuller, and R. Willinger. COST 327: Motorcycle Safety Helmets. Belgium: European Commission, Directorate General for Energy and Transport, 2001. 6 Doorly, M. C., J. P. Phillips, and M. D. Gilchrist. Reconstructing real life accidents towards establishing criteria for traumatic head impact injuries. In: Proceeding of IUTAM Symposium on Impact Biomechanics: From Fundamental Insights to Applications. Netherlands: Springer, 2005, pp. 81–90. 7 Faul, M., L. Xu, M. M. Wald, and V. G. Coronado. Traumatic brain injury in the United States: emergency department visits, hospitalizations and deaths 2002–2006. Atlanta, GA: Centers for Disease Control and Prevention, National Center for Injury Prevention and Control, 2010. 8 Hajiaghamemar M., M. Seidi, V. Caccese, and M. Shahinpoor. Characteristics of falling impact of head using a test dummy. In: Proceeding of ASME International Mechanical Engineering Congress and Exposition, San Diego, CA: American Society of Mechanical Engineers, 2013. 9 Hamel, A., M. Llari, M. Piercecchi-Marti, P. Adalian, G. Leonettiand, and L. Thollon. Effects of fall conditions and biological variability on the mechanism of skull fractures caused by falls. Int. J. Leg. Med. 127:111–118, 2013. 10 Hardy, W. N., M. J. Mason, C. D. Foster, C. S. Shah, J. M. Kopacz, K. H. Yang, A. I. King, J. Bishop, M. Bey, W. Anderst, and S. Tashman. A study of the response of the human cadaver head to impact. Stapp Car Crash J. 51:17– 80, 2007. 11 Heller, M. F., J. George, G. T. Yamaguchi, J. C. McGowan, and M. T. Prange. Linear Head accelerations resulting from short falls into the occiput in children. In: Annual Meeting of the American Society of Biomechanics, State College, PA: American Society of Biomechanics, 2009. 12 Hodgson, V. R. and L. M. Thomas. Breaking strength of the human skull vs. impact surface curvature. US Department of Transportation. NHTSA Rep. DOT/HS-800-583, 1971. 13 Ibrahim, N. G., and S. S. Margulies. Biomechanics of the toddler head during low-height falls: an anthropomorphic dummy analysis. J. Neurosurg. Pediatr. 6:57–68, 2010.
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Kang, Y. S., K. Moorhouse, and J. H. Bolte. Measurement of six degrees of freedom head kinematics in impact conditions employing six accelerometers and three angular rate sensors (6ax configuration). J. Biomech. Eng. Trans. ASME 133:111007, 2011. 15 Kleinberger, M., E. Sun, R. Eppinger, S. Kuppa, and R. Saul. Development of improved injury criteria for the assessment of advanced automotive restraint systems. Washington, DC: National Highway Traffic Safety Administration, 1998. 16 Mertz H. Biofidelity of the Hybrid III head. No. 85124, SAE Technical Paper, 1985. 17 Mertz, H. J., P. Prasad, and N. L. Irwin. Injury risk curves for children and adults in frontal and rear collisions. No. 973318. SAE Technical Paper, 1997. 18 Nagata, H., and H. Ohno. Analysis of backward falls caused by accelerated floor movements using a dummy. Ind. Health 45:462–466, 2007. 19 National Safety Council. Injury FactsÒ. Itasca, IL, 2011. 20 O’Riordain, K., P. M. Thomas, J. P. Phillips, and M. D. Gilchrist. Reconstruction of real world head injury accidents resulting from falls using multibody dynamics. Clin. Biomech. 18:590–600, 2003. 21 Rowson, S., S. M. Duma, J. G. Beckwith, J. J. Chu, R. M. Greenwald, and M. M. Crisco. Rotational head kinematics in football impacts: an injury risk function for concussion. Ann. Biomed. Eng. 40:1–13, 2012. 22 Schmitt, K. U., P. F. Niederer, M. H. Muser, and F. Walz. Trauma Biomechanics (3rd ed.). Zurich: Springer, pp. 63– 89, 2010. 23 Seidi, M., M. Hajiaghamemar, and V. Caccese. Evaluation of effective mass during head impact due to standing falls. Int. J. Crashworthiness, 2014. doi:10.1080/13588265.2014.983261. 24 Versace, J. A review of the severity index. No. 710881. SAE Technical Paper, 1971. 25 Yoganandan, N., F. A. Pintar, A. Sances, Jr., P. R. Walsh, C. L. Ewing, D. J. Thomas, and R. G. Snyder. Biomechanics of skull fracture. J. Neurotrauma 12:659–668, 1995. 26 Zhang, J., N. Yoganandan, and F. A. Pintar. Dynamic biomechanics of the human head in lateral impacts. Ann. Adv. Automot. Med. 53:249–256, 2009.
Measurement of Head Impact Due to Standing Fall in Adults Using Anthropomorphic Test Dummies MARZIEH HAJIAGHAMEMAR,1 MORTEZA SEIDI,1 JAMES R. FERGUSON,2 and VINCENT CACCESE1 1
Advanced Biomechanics Laboratory for Injury Reduction and Rehabilitation, Mechanical Engineering Department, University of Maine, Orono, ME, USA; and 2Alba-Technic LLC, Winthrop, ME, USA (Received 9 September 2014; accepted 14 January 2015) Associate Editor Stefan M Duma oversaw the review of this article.
Abstract—The kinematics and kinetics of head impact due to a standing fall onto a hard surface are summarized. Head injury due to impact from falls represents a significant problem, especially for older individuals. When the head is left unprotected during a fall, the impact severity can be high enough to cause significant injury or even death. To ascertain the range of head impact parameters, the dynamic response was captured for the pedestrian version of the 5th percentile female and 50th percentile male Hybrid III anthropomorphic test dummies as they were dropped from a standing position with different initial postures. Five scenarios of falls were considered including backward falls with/without hip flexion, forward falls with/without knee flexion and lateral falls. The results show that the head impact parameters are dependent on the fall scenario. A wide range of impact parameters was observed in 107 trials. The 95% prediction interval for the peak translational acceleration, peak angular acceleration, peak force, impact translational velocity and peak angular velocity are 146–502 g, 8.8–43.3 krad/s2, 3.9–24.5 kN, 2.02– 7.41 m/s, and 12.9–70.3 rad/s, respectively. Keywords—Head impact, Falls, Head injury criteria, Acceleration, Head protection device.
INTRODUCTION Falls are the leading cause of nonfatal unintentional injuries and in 2008 alone more than 8.5 million people were treated in emergency departments for fall-related events.19 Among the most devastating outcomes of falling is traumatic brain injury (TBI) that can result in serious and even fatal outcomes. TBI is a contributing factor to nearly a third (30.5%) of all injury-related
Address correspondence to Marzieh Hajiaghamemar, Advanced Biomechanics Laboratory for Injury Reduction and Rehabilitation, Mechanical Engineering Department, University of Maine, Orono, ME, USA. Electronic mail: [email protected]
deaths in the United States.7 Different types of head injury may occur due to a severe fall impact. Head injury can be predicted using a variety of measures that depend on injury type. The following parameters can be used for the assessment of head injury potential: peak resultant translational acceleration of the head center of gravity,2 rotational velocity and rotational acceleration,21 peak of impact force on the skull,1,12,25 impact duration using measures such as the head injury criteria (HIC15),2,22,24 and brain response using measures such as pressure, strain and strain rate of the brain tissue.3 Outside of actual human testing, which is not viable for severe head impact fall research, alternative methods used to assess human fall response by physical or numerical means must be interpreted with caution. The Hybrid III anthropomorphic test dummy (ATD) plays an important role in the ability to produce accurate fall simulation data without the use of human subjects. Mertz compared the response of the 50% male Hybrid III head to cadaver data16 and showed that the Hybrid III head is viable for impact comparisons with the cadaver test data. Fall studies using ATDs have been conducted by multiple researchers. Ibrahim and Margulies13 used ATDs to investigate low-heights falls (1, 2, and 3 ft) in toddlers and infants and found that the toddler is likely less vulnerable to skull fracture but more vulnerable to neurological impairment compared with the infant. Heller et al.11 dropped 1- and 3-year old Hybrid III dummies onto carpet and concrete surfaces from various heights where the impact velocities varied from 3.8 to 5.5 m/s to investigate the head impact response due to falls. They reported that serious head injuries may occur in 1- and 3-year-old children even from a short distance fall. In 2007, Nagata and Ohno18 conducted a study on elderly falls using a test dummy
Ó 2015 Biomedical Engineering Society
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that represented an average Japanese male and found that the rotations of ankle and hip joints can affect the severity of head injuries by decreasing the head impact velocity and by extending the impact time. Another study was performed to ascertain head impact values for falls from standing and kneeling positions using a 50th percentile male Hybrid-III and reported the average peak acceleration magnitude of 302 g, and a HIC15 of 1487.4 Authors in another work investigated the influence of the body on the effective mass involved during fall-related head impacts using Hybrid-III ATDs and reported that the effective mass can be estimated as the mass of head plus half of the neck mass.23 Some studies attempted to numerically simulate real life fall accidents using a multibody model numerical analysis program MADYMOTM.6,20 O’Riordain et al.20 investigated four cases of falls in persons ranging from 11 to 76 years old with injuries ranging from contusion to skull fracture. Their results showed a maximum acceleration ranging from 243 to 435 g, while the angular accelerations ranged from 17.6 to 43.5 krad/s2. Doorly et al.6 investigated 10 real life cases where falls occurred with the subjects standing; the calculations resulted in the peak translational acceleration between 189 and 456 g, angular acceleration of 7.4–49.2 krad/s2 and HIC15 ranging from 511 to 5951. It was observed that the multibody model gave good representations of the real life cases, however, when the case was complex, the results were not ideal and showed a high variance due to head contact characteristics and initial conditions. In another study, Hamel et al.9 used pedestrian multibody models to simulate backward falls from a standing position. They simulated free falls and blocked falls and studied the influence of impact velocity (3.5–6.5 m/s) and impact surface. They concluded that the mechanism of skull fracture in the case of a backward fall from a standing height is influenced by impact velocity, impact surface, cortical thickness and rigidity. In addition to ATD tests and numerical simulations, many researchers conducted cadaveric studies to investigate the response of the human head to impact. Zhang et al.26 used a controlled drop of the cadaver head instrumented with an accelerometer array onto padding surfaces of 40D and 90D durometer. They reported angular accelerations vs. impact translational velocities ranging from an average of 9 krad/s2 at an impact velocity of 2.5 m/s to 44 krad/s2 at an impact velocity of 5.5 m/s based on their linear curve fit for the 90D surface. Hardy et al.10 investigated the response of the human head to impact using cadavers at impact velocities between 2.9 and 3.9 m/s. When using a CFC 1000 filter, their results of the peak translational acceleration and HIC15 for the unprotected occipital impact tests ranged from 153 to 408 g and 372 to 2540. Since there is an anatomical equivalency between cadaver skull bone and
in vivo human, cadaveric tests are usually used to determine the tolerance of the head to skull fracture.25 Yoganandan et al.25 investigated levels of force required to fracture the skulls of 12 unembalmed cadavers using a hemispherical impactor with a 48 mm radius. They performed impact tests on various locations of the skull at a rate of 7.1–8.0 m/s. They noted failure loads ranging from 4.5 to 14.1 kN (11.9 ± 0.9 kN). Also, Allsop et al.1 carried out temporo-parietal impact tests on 31 unembalmed cadaver heads with two types of flat rigid impactors, one circular and 2.54 cm in diameter, the other a rectangular plate 5 9 10 cm. Fracture force for the small circular plate ranged between 2.5 and 10.0 kN while for rectangular plates it ranged between 5.8 and 17.0 kN. They concluded that the skull’s peak force threshold increases with greater surface area of the impacting force and fracture is more a function of stress (force/area) than force alone. Thus, an impact with a flat ground surface—more like a case of falling—requires higher forces to cause fracture as demonstrated in the rectangular plate. In both of the above studies, the impactors hit the stationary cadaver heads, while in the fall impact case a moving head strikes the ground. There are some cadaveric studies that more closely represent falling cases where the head is not stationary. Hodgson and Thomas12 dropped embalmed human male cadaver heads at impact velocities ranging from 1.6 to 4.7 m/s to assess rear, side and front head impacts and reported 5.6–17.8 kN (10 ± 4.9 kN) for fracture threshold. The aim of the present study was to quantify the kinematics and kinetics of head impact due to a standing fall onto a hard surface. Since it is not possible to assess the response of humans falling onto hard surfaces in the laboratory due to safety concerns, the pedestrian versions of a Hybrid III 5th percentile female and 50th percentile male ATD were used. Several fall scenarios were initiated from a frontal, sideward and backward orientation and the responses of the ATDs were quantified. The longterm agenda of this research team is to develop an informed design of non-stigmatizing head protection for fall-related head impact.
MATERIALS AND METHODS ATDs were used to quantify the kinetics and kinematics of head impact in adults falling from a standing position. To this end, a HumaneticsTM Hybrid III 5th percentile female ATD with 1.5 m height and 49 kg mass and a 50th percentile male ATD with 1.75 m height and 78 kg mass were employed. The hip of the standard ATD was replaced with a movable pedestrian style hip. Ball joints at the femur to pelvis connection of the pedestrian style hip allowed hip rotation, abduction/adduction and flexion/extension to simulate
Head Impact Due to Standing Fall in Adults
various human postures. The initial orientation of the ATD for forward, backward and sideward falls of the female ATD is shown in Fig. 1a; these orientations were the same for the male ATD. A force plate model AMTI LG6-4-2000 was used to measure impact force and an auxiliary platform was laid out so that only the head of the ATD hit the force plate. Several trials of backward falls were also performed using the laboratory floor as the impact surface to assess the influence of the force plate on the impact response. The head of the ATD was instrumented with an accelerometer array consisting of four tri-axial accelerometers (PCB, Depew, NY) with a measuring range of ±500 and ±5000 g in each direction, for female and male ATD, respectively. The accelerometer array mounted inside the head form is illustrated in Fig. 1b. Additionally, three ARS PRO-8K angular rate sensors (DTS, Seal Beach, CA) were mounted on perpendicular faces of the accelerometer at the center of gravity of the heads to directly capture the angular velocity. Angular acceleration for the head center of gravity was calculated from nine of the twelve acceleration signals by the method described by Kang et al.14 using the following equations: x_ x ¼
aBackz aCGz aCGy aTopx þ 2d1 2d3
ð1Þ
x_ y ¼
aTopy aCGx aCGz aSidez þ 2d3 2d2
ð2Þ
aCGx aBackx aSidex aCGy ; 2d1 2d2
ð3Þ
x_ z ¼
where d1 = 47.36 mm, d2 = 26.40 mm for both ATDs and d3 = 77.08 mm, d3 = 88.05 mm for the female and male ATDs respectively. A motion capture system from ViconTM recorded the whole body movement at 200 Hz using 10 cameras distributed about the laboratory. A total of 35 markers based on the plug-in gait full body model from ViconTM were attached to the ATD. Two additional markers were added to the pelvis to supplement the model data when the body obstructed some of the markers from the cameras. The ATDs were kept in a standing posture using an electro-mechanical release mechanism with a mass of 0.072 kg that was triggered by the Vicon MX giganet. Also, a digital video camera was used to record the experiments. Data of the force plate and accelerometers were collected at a sampling rate of 20 kHz using USB1608FS-Plus DAQ (Measuring Computing, Norton, MA). The data were filtered using a CFC1000 (Chanel Frequency Class) filter that is a fourth order Butterworth filter with a 1650 Hz cutoff frequency. Data were collected and processed by a PC using a NI
LabVIEW software (National Instrument, Texas, USA) with help of MCC ULx library for NI LabVIEW, then analyzed using MATLABÒ. Five scenarios of falls were studied as illustrated in Fig. 2. Case 1 (BK-NHB) was a backward fall where the hip joints did not rotate before the head hits the ground; this happens when a person does not react quickly to the fall. The motion of this case is similar to the dynamics of a single inverted pendulum. Case 2 (BK-HB) was a backward fall with hip flexion which the hip hits the ground first followed by the head impact. The dynamics of this case is similar to the behavior of a double inverted pendulum. Case 3 (FW-KB) was a forward fall with the knee flexion during which the knees hit the ground first followed by the head. Case 4 (FW-NKB) was a forward fall with the knee joints fixed. Case 5 (LAT) was a sideward fall from a standing position where the shoulder typically contacted the ground first and then the head. In running the LAT experiments it was observed for the male ATD that the head did not make contact with the ground in this orientation due to the size of the ATD shoulder, therefore, the arm of the male ATD was removed in the LAT fall case. Although there are numerous head injury indicators, the HIC15 was used as the injury predictor for the purpose of this paper since it is a common measure used in many published head impact studies and is one that includes the influence of impact duration. The HIC15 measure was calculated using the formula: ( 2:5 ) 1 t 2 aðtÞdt HIC15 ¼ max ðt2 t1 Þ ; ð4Þ t 2 t 1 t1 where aðtÞ is the resultant translational acceleration as a function of time and t1 and t2 define the time interval that maximizes the HIC15. The HIC15 criterion was computed from the acceleration signature measured in g’s during impact and the maximum time duration of (t2 2 t1) is limited to a maximum value of 15 ms. The time interval (t2 2 t1) is referred to as the HIC15 duration and is selected to maximize the HIC15 value while the impact duration is defined as the time difference between the leading and trailing edges of the primary translational acceleration pulse. The HIC15 was first proposed by Versace24 and then modified by NHTSA 1998.15 According to Mertz et al.,17 a HIC15 of 700 represents a 5% risk of a severe injury. Chinn et al.5 found that of all the parameters analyzed in their study of motorcycle helmets, the HIC15 provided the most reliable head injury severity prediction with a correlation coefficient of r = 0.8 and the HIC15 of 1000 and 1500 correspond to a moderate and serious head injury, respectively.
HAJIAGHAMEMAR et al.
FIGURE 1. Experimental setup, (a) test configuration using the Hybrid III 5th percentile female dummy, (b) the configuration of accelerometers and angular rate sensors.
RESULTS For the five fall scenarios of the male and female ATD, a total of 107 trials were conducted. Different ATD response kinematics were observed. In the first case (BK-NHB), the hip joints did not rotate. Shortly after the pelvic struck (within 25–45 ms), the torso contacted the ground while in the second case (BKHB) a 350–450 ms time interval was observed between pelvic and torso contact to the floor. The head impact occurred within 20 ms after the torso contacted the floor for both backward scenarios while this time interval was 5 ms for the forward scenarios. Sideward fall scenarios showed different kinematics, the pelvic struck the ground followed by the arm/shoulder and then the head. The time intervals between arm/shoulder contact and head impact were observed to be approximately 40 ms for the female ATD and 30 ms for the male ATD. A summary of measurements of the female and male ATD including results of the peak translational acceleration, peak angular acceleration, impact translational velocity, peak angular velocity, impact duration and peak force are given Tables 1 and 2. Here and throughout this paper, the peak translational acceleration and peak angular acceleration denote the maximum of translational and angular acceleration signals
after impact occurred. The impact translational velocity represents the maximum of translational velocity signal immediately prior to impact and the peak angular velocity denotes the maximum angular velocity signal either immediately prior to impact or after impact, whichever is greater. The first row in the table gives the number of trials conducted for each type of fall. The number after the plus sign shows the number of trials from a preliminary study on a backward fall of the Hybrid III 5th percentile female ATD that did not include acquisition of angular acceleration and angular velocity.8 The summary provides the mean of the peak values, standard deviation, and range of the peak values (given in parentheses) for the kinetic and kinematic measures. As seen in Tables 1 and 2, the peak translational acceleration of the male ATD was on average greater than the female ATD due to the higher height and mass of the male dummy. On the other hand, angular accelerations were similar for both ATDs except during lateral falls. This is attributed to the change in dynamic response caused by the shoulder of the male dummy being disassembled and removed during the lateral testing. A typical translational and angular acceleration magnitude time-history truncated at a time of 3.5 ms after floor impact is shown in Fig. 3 for the various fall conditions for both the female and male ATD. The
Head Impact Due to Standing Fall in Adults
FIGURE 2. The scenarios of standing fall. TABLE 1. Summary of results for female ATD. Condition/measure Number of trials Trans. acc. (g) Ang. acc. (krad/s2) Ang. vel. (rad/s) Trans. vel. (m/s) Impact duration (ms) Peak force (kN) HIC15 HIC15 duration (ms)
BK-NHB
BK-HB
FW-KB
FW-NKB
LAT
12 + 7 368 ± 31 (308–411) 28.8 ± 3.4 (24.4–35.2) 48.7 ± 4.5 (37.5–54.2) 4.80 ± 0.22 (4.37–5.19) 2.5 ± 0.2 (2.3–3.0) 14.7 ± 1.1 (12.4–16.4) 2173 ± 298 (1652–2678) 1.4 ± 0.1 (1.3–1.6)
15 + 14 289 ± 47 (157–361) 23.6 ± 4.5 (13.5–33.8) 32.7 ± 5.7 (20.5–45.5) 3.78 ± 0.53 (2.49–4.84) 2.5 ± 0.3 (2.2–3.4) 11.5 ± 2.1 (6.0–14.4) 1218 ± 425 (331–2156) 1.4 ± 0.1 (1.2–1.8)
7 243 ± 26 (204–278) 28.9 ± 5.9 (19.8–35.9) 44.1 ± 5.7 (34.3–52.8) 4.39 ± 0.63 (3.53–5.40) 3.9 ± 0.7 (3.0–4.8) 9.8 ± 1.4 (7.5–11.6) 888 ± 207 (682–1312) 1.8 ± 0.3 (1.3–2.3)
5 319 ± 26 (293–354) 29.7 ± 4.0 (23.6–34.6) 46.8 ± 3.1 (41.7–49.9) 5.81 ± 0.28 (5.55–6.28) 4.1 ± 0.5 (3.4–4.7) 12.7 ± 1.4 (11.0–14.6) 1846 ± 213 (1608–2111) 1.8 ± 0.1 (1.6–2.0)
11 206 ± 79 (64–305) 22.2 ± 10.6 (4.5–37.8) 30.1 ± 10.1 (17.2–43.8) 2.90 ± 0.92 (1.37–4.42) 3.1 ± 0.7 (2.1–4.8) 7.6 ± 3.0 (2.6–11.6) 598 ± 417 (41–1431) 1.5 ± 0.3 (1.3–2.3)
HAJIAGHAMEMAR et al. TABLE 2. Summary of results for male ATD. Condition/measure Number of trials Trans. acc. (g) Ang. acc. (krad/s2) Ang.vel. (rad/s) Trans. vel. (m/s) Impact duration (ms) Peak force (kN) HIC15 HIC15 duration (ms)
BK-NHB
BK-HB
FW-KB
FW-NKB
LAT
10 451 ± 38 (390–524) 29.2 ± 5.8 (19.6–41.0) 58.9 ± 5.7 (48.3–67.6) 6.75 ± 0.27 (6.31–7.11) 3.0 ± 0.4 (2.6–4.0) 22.8 ± 2.1 (19.3–26.6) 4142 ± 536 (3182–5046) 1.6 ± 0.1 (1.4–1.8)
9 295 ± 89 (198–447) 19.3 ± 9.6 (10.7–38.2) 34.3 ± 17.5 (17.0–64.4) 4.85 ± 1.33 (3.29–6.80) 3.4 ± 0.4 (3.0–4.0) 14.9 ± 4.6 (9.4–22.2) 1826 ± 1342 (530–4234) 1.7 ± 0.1 (1.6–1.8)
5 402 ± 69 (311–489) 23.8 ± 15.6 (8.9–46.2) 37.5 ± 24.3 (16.0–75.8) 6.46 ± 1.29 (4.43–7.90) 3.2 ± 0.3 (2.8–3.5) 20.3 ± 3.7 (15.4–25.6) 3286 ± 1136 (1711–4691) 1.7 ± 0.2 (1.6–1.9)
5 439 ± 118 (294–564) 21.1 ± 13.7 (7.0–40.2) 32.1 ± 16.8 (15.6–53.6) 6.68 ± 1.20 (4.82–7.64) 3.1 ± 0.5 (2.6–3.8) 21.6 ± 6.1 (14.2–29.6) 4072 ± 2247 (1353–6435) 1.6 ± 0.1 (1.4–1.8)
7 355 ± 49 (264–418) 34.3 ± 5.8 (30.4–46.9) 56.4 ± 1.6 (54.1–58.6) 5.12 ± 0.66 (3.91–6.10) 2.9 ± 0.3 (2.4–3.3) 17.1 ± 2.2 (12.6–19.0) 2128 ± 585 (996–2745) 1.5 ± 0.2 (1.2–1.7)
black boxes in the figure represent the start and end times used for the HIC15 computation. Correlation between impact parameters like impact translational velocity and peak translational acceleration, peak angular velocity and acceleration, impact translational velocity and peak force; and peak translational and angular acceleration were established and the 95 and 99% prediction intervals were computed (Figs. 4, 5, 6, and 7). In addition to testing on the force plate, several trials of impact onto the laboratory floor were conducted to ascertain whether testing the head impact on the force plate can be used to simulate fall-related head impact onto a common commercial building surface. Peak translational accelerations of head impact onto the force plate were 289 ± 47 g for the second scenario (BK-HB) of the female ATD. This was compared to the same case striking the concrete floor covered with a layer of thin carpeting with no padding that resulted in peak translational accelerations of 311 ± 9 g. Accordingly, test results showed that impact onto the force plate are in the same range as an impact onto a building surface such as carpeted concrete floor.
DISCUSSION The response of the female and male ATDs was quantified to describe the kinematics and kinetics of different scenarios of falls from standing focusing on head impact. The results show that the head impact parameters were dependent on the fall direction and type (Tables 1, 2). The BK-NHB case was the most severe case among all fall scenarios studied herein. The highest mean value of peak translational acceleration (368 g for female and 451 g for male), impact force
(14.7 kN for female and 22.8 kN for male), HIC15 value (2173 for female and 4142 for male) and peak angular velocity (48.7 rad/s for female and 58.9 rad/s for male) was observed for this type of fall. For a lateral fall, the mean values of peak translational acceleration (206 g for female and 355 g for male) and HIC15 (598 for female and 2128 for male) were relatively lower compared to the backward and forward cases. The mean value of impact duration and HIC15 duration for all types of falls were in similar ranges but slightly longer impact duration was observed for forward fall cases. In addition to the study of the effect of fall scenarios on head impact, the bounds of the head impact parameters that can arise from a standing fall were assessed. For this reason, all data were grouped together and the mean and standard deviation of peak translational acceleration, peak angular acceleration, impact force, impact translational velocity, peak angular velocity and HIC15 are 324 ± 89 g, 26.0 ± 8.6 krad/s2, 14.2 ± 5.1 kN, 4.71 ± 1.34 m/s, 41.6 ± 14.3–60 rad/s, and 1982 ± 1290, respectively. Comparisons with Previous Studies A major task in the development of a methodology to evaluate fall protection devices is to assess the bounds of the impact parameters that occur during an unprotected fall. Also, establishing correlations between the impact parameters can be used to improve numerical analysis of real life fall incidents and to develop simplified experimental procedures for testing of head protection devices. To that end, correlation between impact parameters were established and the 95 and 99% prediction intervals were also plotted (Figs. 4, 5, 6, and 7) to provide bounds for estimations of the head impact parameters due to fall. Prior
Head Impact Due to Standing Fall in Adults
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FIGURE 3. Head impact time histories for various fall conditions. (a) Translational acceleration for female ATD, (b) Translational acceleration for male ATD, (c) Angular acceleration for female ATD, (d) Angular acceleration for male ATD.
methods employed in other studies include numerical simulations using the MADYMOTM program,6,20 cadaver tests,10,12,26 and ATD testing4 were also compared to the current ATD study to verify the bounds of the head impact response due to a standing fall (Figs. 4, 5, 6, and 7). The impact translational velocity vs. the peak translational acceleration and the peak angular acceleration vs. the peak angular velocity are shown in Fig. 4. Plotted in this figure are the 95% (long dash) and 99% (short dash) prediction interval bounds and linear least square fit (solid line) of the ATD data. The linear fit passing through the origin resulted in a coefficient of determination of 0.73 and 0.57 for the results shown in Figs. 4a and 4b. Results from MADYMOTM reconstruction6,20 are presented in both the translational and angular plots as shown in Figs. 4a and 4b, respectively. Results of cadaver studies conducted by Zhang et al.26 and Hodgson and Thomas12 are also presented but only in the translational plot (Fig. 4a), since angular velocity at impact was not reported in these studies. Zhang et al.26 dropped the cadaver heads onto 40D and 90D durometer but only
the data of the harder surface (90D) was used in the comparison. The numerical model and cadaver test results fit within the 99% prediction interval for peak translational acceleration and with the exception of 3 outliers the same can be said for angular acceleration. The numerical models tended to compute a lower peak translational acceleration than the ATD and cadaver test results. Angular acceleration computed in the numerical models tended to be greater than the values obtained in the ATD tests. Peak translational acceleration vs. the HIC15 value is presented in Fig. 5 where a power law fit of the ATD test resulted in a coefficient of determination of 0.97 indicating a strong correlation with exponent 2.42. This result was as expected due to the nature of the HIC15 computation. The current ATD test results are compared to MADYMOTM simulations,6,20 cadaver test10 and another ATD study.4 The MADYMOTM simulations6,20 tended toward higher HIC15 values that indicate a longer impact duration in the computation (Fig. 5). The cadaver studies by Hardy10 tended toward lower HIC15 values indicating shorter impact
HAJIAGHAMEMAR et al. 35
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Angular Velocity (rad/s) FIGURE 4. Impact translational velocity vs. peak translational acceleration, linear regression line (solid line), 95% (long dash) and 99% (short dash) prediction bounds for ATD tests compared to studies by Zhang,26 Hodgson,12 Doorly,6 and O’Riordian.20 (a) Translational, (b) Angular.
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FIGURE 6. Peak impact force vs. impact translational velocity, linear regression line (solid line), 95% (long dash) and 99% (short dash) prediction bounds for ATD compared to Hodgson,12 Doorly,6 and O’Riordian.20
y = 0.0014x2.4218 R² = 0.971
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Backward Fall-ATD Forward Fall-ATD Lateral Fall-ATD Zhang-Cadaver Doorly-MADYMO O'Riordain-MADYMO
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durations. The other ATD study also indicates similar impact duration and HIC15 value4 (Fig. 5). Peak of impact force is often used to predict skull fracture since a high level of force can cause failure of the cranial bone. The peak impact force vs. the impact velocity for the ATD tests, cadaver tests by Hodgson12 and numerical analyzes by Doorly6 and O’Riordain20 are shown in Fig. 6. Also plotted in the figure are the 99 and 95% prediction intervals for the ATD test data. It was observed on average that the ATD test data resulted in higher levels of impact force at an equivalent impact velocity. In the current study, the head impact forces ranged from 2.60 to 16.37 kN for female ATD tests and from 12.6 to 29.6 kN for the male ATD tests. Impact forces for occipital impacts due to
Head Impact Due to Standing Fall in Adults
backward falls and frontal impacts due to forward falls were 69% and 46% higher than lateral impacts for female ATD and 22 and 12% higher than lateral impacts for male ATD. The impact forces in many of our ATD tests were above the fracture thresholds1,12,25 (Fig. 6). Therefore, based on the ATD standing fall tests, probability of skull fracture is high when the fall is unprotected and occurs onto a hard surface, especially in the FW-NKB and the BK-NHB fall scenarios. The relationship between peak angular and translational acceleration is important to several types of injury predictions and depends greatly on impact orientation and location that are highly variable in real life cases. The peak angular vs. translational acceleration for the ATD testing as well as the cadaver testing by Zhang26 and the numerical analyzes by Doorly6 and O’Riordian20 are shown in Fig. 7. Little correlation was observed in the results, as expected, with a coefficient of determination of 0.42, 0.37, and 0.54 for the backward, forward and lateral ATD scenarios. The numerical analysis studies performed for various fall scenarios showed a similar distribution. The cadaver study by Zhang26 showed a stronger correlation with a coefficient of determination of 0.90 since the impact parameters were controlled in the study. This result indicates that when conditions are controlled, as they are in drop tower testing, there is typically a defined relationship between the peak angular and translational acceleration. In interpretation of the results it is noted that a limitation of this study is a lesser biofidelity of the torso and neck compared to the head for fall study that may affect the results. The authors in another work4 studied the effect of neck stiffness by comparing a standard (70–80 Shore A durometer) and lower stiffness (35 Shore A durometer) neck in head impacts using Hybrid-III head and neck assembly. The results showed that the influences of neck stiffness were less than 10% for rear and frontal impacts and less than 17% for lateral impacts for both translational and angular accelerations.4 In addition, as mentioned in the method section, the arm of the male ATD was removed in the LAT fall case. This condition affects the results especially on head rotational movement as the ATD hits the shoulder instead of arm. Also, it should be noted that the results of this study do not include the human tissue response (like skull and brain tissues) to an impact and actual human reaction to a fall. This study relies on the ATD design to simulate the human response and therefore should be used with caution.
CONCLUSION This paper presents the response of a 5th percentile female and 50th percentile male pedestrian ATD sub-
ject to standing falls. Conditions of forward, backward, and sideways falls were studied. Test results were compared to cadaver, numerical analyzes and other ATD studies. The female ATD generally resulted in less severe impact attributed to the lower drop height and subsequently lower impact velocity. Mean peak forces were also in a range that indicates potential for head injury including skull fracture. Among all the fall cases, the unprotected BK-NHB fall had the highest mean value of peak translational acceleration, impact force and HIC15 value and was the fall scenario with the potential for the most severe injury. Correlation between impact parameters were established and the 95 and 99% prediction intervals were obtained to provide bounds for estimation of the head impact parameters. The results presented can be used to improve numerical analysis of real life fall incidents and to develop simplified experimental procedures for testing of head protection devices. Although there are major differences in methodology, when ATD testing was compared to cadaver testing and numerical analyzes, the numerical and physical studies lay within the 99% prediction bounds of the 107 ATD tests reported.
ACKNOWLEDGMENTS The authors gratefully acknowledge support from the Maine Technology Institute under Grant Number MTAF-3001 for funding of the laboratory facility, the National Institutes of Health/National Institute on Aging under grant number NIH 5R44AGO33936-03, the George and Caterina Sakellaris Graduate Fellowship and the National Science Foundation under Award No. 1417120.
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