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Objective: Researchers have found a variety of uses for the Hybrid III (HIII) dummy that fall beyond the scope of its original purpose as an automotive crash test dummy. Some of these expanded roles for the HIII introduce situations that were not envisioned in the dummy's original design parameters, such as a relatively rapid succession of tests or outdoor testing scenarios where temperature is not easily controlled. This study investigates how the axial compressive stiffness of the HIII lumbar spine component is affected by the duration of the time interval between tests. Further, it measures the effect of temperature on the compressive stiffness of the lumbar spine through a range of temperatures relevant to indoor and outdoor testing., Methods: High-rate axial compression tests were run on a 50 th percentile male HIII lumbar component in a materials testing machine. To characterize the effects of tests recovery intervals, between-test recovery was varied from 2 hours to 1 minute. To quantify temperature effects, environmental temperature conditions of 12.5°, 25°, and 37.5 °C were tested., Results: During repeated compressive loading, the force levels decreased consistently across long and short rest intervals. Even after 2 hours of rest between tests, full viscoelastic recovery was not observed. Temperature effects were pronounced, resulting in compressive force differences of 261% over the range of 12.5° to 37.5 °C. Compared to the stiffness of the lumbar at 25 °C, the stiffness at 37.5 °C fell by 40%; at 12.5 °C, the stiffness more than doubled, increasing by 115%., Conclusions: A modest decrease in temperature can be sufficient to dramatically change the response and repeatability of the lumbar HIII component in compressive loading. The large magnitude of the temperature effect has severe implications in its ability to overwhelm the contributions of targeted test variables. These findings highlight the importance of controlling, monitoring and reporting temperature conditions during HIII testing, even in indoor laboratory environments.

The Hybrid III (HIII) dummy is one of the most widely used anthropomorphic test devices (ATDs) in the world, and researchers have found a variety of uses for it outside of its original purpose as an automotive crash test dummy. These expanded roles have introduced situations outside the dummy's original design parameters, where a number of tests must be run in relatively rapid succession or where it may not be possible to control the temperature of the test environment., Objective: This study has 2 aims. The first is to determine how the duration of the time interval between tests affects the axial compression performance of the HIII neck. The second is to quantify the effect of temperature on the neck's compressive stiffness through a range of temperatures relevant to indoor or outdoor testing., Methods: To characterize the effects of different test conditions, a series of high-rate axial compressive tests was run on a 50th percentile male HIII neck component in a materials testing machine. Between-test recovery intervals were varied from 2 h to 1 min, and temperature conditions of 0, 12.5, 25, and 37.5 °C were tested., Results: Though the duration of the recovery interval had little impact on the recorded force (less than 1%), the component did exhibit considerable strain creep over the course of the test. Temperature had a strong influence on the compressive stiffness of the component. Compared to the stiffness at 25 °C (near room temperature), the stiffness of the neck at 37.5 °C fell by 15%; at 0 °C, the stiffness more than doubled., Conclusions: This study demonstrates that though the duration of the recovery interval between tests has a small influence on neck stiffness, temperature effects should not be overlooked because they influence neck compressive stiffness considerably. The relationship between recorded force and temperature is well represented by exponential decay models. These findings highlight the importance of monitoring and controlling for temperature effects during all HIII testing.

Objective: Traumatic injuries are the leading cause of death of children aged 1-19 in the United States and are principally caused by motor vehicle collisions, with the head being the primary region injured. The neck, though not commonly injured, governs head kinematics and thus influences head injury. Vehicle improvements necessary to reduce these injuries are evaluated using anthropomorphic testing devices (ATDs). Current pediatric ATD head and neck properties were established by scaling adult properties using the size differences between adults and children. Due to the limitations of pediatric biomechanical research, computational models are the only available methods that combine all existing data to produce injury-relevant biofidelity specifications for ATDs. The purpose of this study is to provide the first frontal impact biofidelity corridors for neck flexion response of 6- and 10-year-olds using validated computational models, which are compared to the Hybrid III (HIII) ATD neck responses and the Mertz flexion corridors., Methods: Our virtual 6- and 10-year-old head and neck multibody models incorporate pediatric biomechanical properties obtained from pediatric cadaveric and radiological studies, include the effect of passive and active musculature, and are validated with data including pediatric volunteer 3 g dynamic frontal impact responses. We simulate ATD pendulum tests-used to calibrate HIII neck bending stiffness-to compare the pediatric model and HIII ATD neck bending stiffness and to compare the model flexion bending responses with the Mertz scaled neck flexion corridors. Additionally, pediatric response corridors for pendulum calibration tests and high-speed (15 g) frontal impacts are estimated through uncertainty analyses on primary model variables, with response corridors calculated from the average ± SD response over 650 simulations., Results and Conclusions: The models are less stiff in dynamic anterioposterior bending than the ATDs; the secant stiffness of the 6- and 10-year-old models is 53 and 67 percent less than that of the HIII ATDs. The ATDs exhibit nonlinear stiffening and the models demonstrate nonlinear softening. Consequently, the models do not remain within the Mertz scaled flexion bending corridors. The more compliant model necks suggest an increased potential for head impact via larger head excursions. The pediatric anterioposterior bending corridors developed in this study are extensible to any frontal loading condition through calculation and sensitivity analysis. The corridors presented in this study are the first based on pediatric cadaveric data and provide the basis for future, more biofidelic, designs of 6- and 10-year-old ATD necks.

Objective: Traumatic brain injury (TBI) is a major public health issue, affecting millions of people annually. Anthropomorphic test devices (ATDs) and finite element models (FEMs) provide a means of understanding factors leading to TBI, potentially reducing the occurrence. Thus, there is a need to ensure that these tools accurately model humans. For example, the Hybrid III was not based on 3-dimensional human head shape data. The objective of this study is to produce average head and skull contours for an average U.S. male that can be used for ATDs and FEMs., Methods: Computed tomography (CT) scans of adult male heads were obtained from a database provided by the University of Virginia Center for Applied Biomechanics. An orthographic viewer was used to extract head and skull contours from the CT scans. Landmarks were measured graphically using HyperMesh (Altair, HyperWorks). To determine the head occipital condyle (OC) centroid, surface meshes of the OCs were made and the centroid of the surfaces was calculated. The Hybrid III contour was obtained using a MicroScribe Digitizer (Solution Technologies, Inc., Oella, MD). Comparisons of the average male and ATD contours were performed using 2 methods: (1) the midsagittal and midcoronal ATD contours relative to the OC centroid were compared to the corresponding 1 SD range of the average male contours; (2) the ATD sagittal contour was translated relative to the average male sagittal contour to minimize the area between the 2 contours., Results: Average male head and skull contours were created. Landmark measurements were made for the dorsum sellae, nasion skin, nasion bone, infraorbital foramen, and external auditory meatus, all relative to the OC centroid. The Hybrid III midsagittal contour was outside the 1 SD range for 15.2 percent of the average male head contour but only by a maximum distance of 1.5 mm, whereas the Hybrid III midcoronal head contour was outside the 1 SD range for 12.2 percent of the average male head contour by a maximum distance of 2 mm. Minimization of the area between the midsagittal contours resulted in only 2.3 mm of translation, corroborating the good correlation between the contours established by initial comparison., Conclusions: Three-dimensional average male head and skull contours were created and measurements of landmark locations were made. It was found that the 50th percentile male Hybrid III corresponds well to the average male head contour and validated its 3D shape. Average adult head and skull contours and landmark data are available for public research use at http://biomechanics.pratt.duke.edu/data .

Objective: During dynamic injury scenarios, such as motor vehicle crashes, neck biomechanics contribute to head excursion and acceleration, influencing head injuries. One important tool in understanding head and neck dynamics is computational modeling. However, realistic and stable muscle activations for major muscles are required to realize meaningful kinematic responses. The objective was to determine cervical muscle activation states for 6-year-old, 10-year-old, and adult 50th percentile male computational head and neck models. Currently, pediatric models including muscle activations are unable to maintain the head in an equilibrium position, forcing models to begin from nonphysiologic conditions. Recent work has realized a stationary initial geometry and cervical muscle activations by first optimizing responses against gravity. Accordingly, our goal was to apply these methods to Duke University's head-neck model validated using living muscle response and pediatric cadaveric data., Methods: Activation schemes maintaining an upright, stable head for 22 muscle pairs were found using LS-OPT. Two optimization problems were investigated: a relaxed state, which minimized muscle fatigue, and a tensed activation state, which maximized total muscle force. The model's biofidelity was evaluated by the kinematic response to gravitational and frontal impact loading conditions. Model sensitivity and uncertainty analyses were performed to assess important parameters for pediatric muscle response. Sensitivity analysis was conducted using multiple activation time histories. These included constant activations and an optimal muscle activation time history, which varied the activation level of flexor and extensor groups, and activation initiation and termination times., Results: Relaxed muscle activations decreased with increasing age, maintaining upright posture primarily through extensor activation. Tensed musculature maintained upright posture through coactivation of flexors and extensors, producing up to 32 times the force of the relaxed state. Without muscle activation, the models fell into flexion due to gravitational loading. Relaxed musculature produced 28.6-35.8 N of force to the head, whereas tensed musculature produced 450-1023 N. Pediatric model stiffnesses were most sensitive to muscle physiological cross-sectional area., Conclusions: Though muscular loads were not large enough to cause vertebral compressive failure, they would provide a prestressed state that could protect the vertebrae during tensile loading but might exacerbate risk during compressive loading. For example, in the 10-year-old, a load of 602 N was produced, though estimated compressive failure tolerance is only 2.8 kN. Including muscles and time-variant activation schemes is vital for producing biofidelic models because both vary by age. The pediatric activations developed represent physiologically appropriate sets of initial conditions and are based on validated adult cadaveric data.

The limited availability of pediatric biomechanical impact response data presents a significant challenge to the development of child dummies. In the absence of these data, the development of the current generation of child dummies has been driven by scaling of the biomechanical response requirements of the existing adult test dummies. Recently published pediatric blunt thoracic impact response data provide a unique opportunity to evaluate the efficacy of these scaling methodologies. However, the published data include several processing anomalies and nonphysical features. These features are corrected by minimizing instrumentation and processing error to improve the fidelity of the individual force-deflection responses. Using these data, biomechanical impact response corridors are calculated for a 3-year-old child and a 6-year-old child. These calculated corridors differ from both the originally published postmortem human subject (PMHS) corridors and the impact response requirements of the current child dummies. Furthermore, the response of the Hybrid III 3-year-old test dummy in the same impact condition shows a similar deflection but a significantly higher force than the 3-year-old corridor. The response of the Hybrid III 6-year-old dummy, on the other hand, correlates well with the calculated 6-year-old corridor. The newly developed 3-year-old and 6-year-old blunt thoracic impact response corridors can be used to define data-driven impact response requirements as an alternative to scaling-driven requirements.

Objective: The neck injury index, NII, developed in ISO 13232 (2005) as a testing and evaluation procedure for assessing the risk of injury to the AO/C1/C2 region of the cervical spine in motorcycle riders is reevaluated using an existing postmortem human subjects (PMHS) data set and resulting in a reformulated NII criterion applicable to PMHS tests., Methods: A recent series of 36 PMHS head/neck component tests was used to examine the risk of neck injury in frontal impacts and to assess the predictive capability of NII for impacts of various orientations. Using force and moment load cell PMHS experimental data, injury risk was assessed using NII evaluated with the ISO 13232-5 algorithms., Results: The injury risk predictions are compared with the injury outcomes from the head/neck PMHS. The NII criterion underestimated the injury incidence of the PMHS experimental group. The average predicted risk of injuries for the experimental injury tests based on NII across the MAIS levels was 0.7 percent, though there were 11 AIS 3+ injuries observed in the actual testing (30.6%). Using the experimental injury outcomes and the experimental force and moment time histories, the normalizing coefficients from NII are reevaluated to minimize the difference between NII risk assessment and the experimental injury outcome in the least squares (L(2)) basis. This reanalysis is compared with existing human and PMHS neck injury criteria., Conclusions: By reanalyzing the NII formulation using an existing PMHS injury data set with known forces and moments and known injury outcomes, a new NII(PMHS) is developed that uses PMHS loads to predict injury. This reformulation removes the dependency of the original NII formulation on the forces and moments from motorcyclist anthropomorphic test device (MATD) experiments and simulations yet retains the advantages of the multi-axial neck injury criterion.

Objective: Three postmortem human surrogates (PMHS) were positioned and rigidly mounted through the spine to a tabletop test fixture for the purpose of characterizing thoracic response to diagonal belt loading with well-defined boundary conditions., Methods: These PMHS were mounted to a stationary apparatus that supported the spine and shoulders in a configuration comparable to that seen in a 48 km/h automobile sled test at the time of maximum chest deformation. A belt restraint was positioned across the anterior torso with attachments at D-ring and buckle locations based on the geometry of a mid-sized sedan. The belt was attached to a trolley driven by a hydraulic ram linked to a universal test machine. Ramp and hold experiments were conducted at rates of 0.5, 0.9, and 1.2 m/s and hold times of 60 s. Ramp-hold displacement waveforms of up to 20 percent of the chest depth were applied to the chest while the resulting belt loads and spinal reaction loads were recorded. These data were used to identify parameters in a seven-parameter thoracic structural model mathematically analogous to a viscoelastic material model. A final test with 40 percent deflection was performed at the completion of the loading sequence., Results: Model fits to ramps of different magnitudes indicated that the assumption of temporal linearity was reasonable over the range of inputs in this study. In agreement with previous studies, the spatial (force-deflection) response was only slightly nonlinear, indicating that a fully linear model would be reasonable up to the deflection levels used here., Conclusions: Pronounced variability in the instantaneous elastic behavior was observed among the three test subjects, whereas the relaxation behavior exhibited less variability.

Objective: This article assesses the position-dependent injury tolerance of the hip in the frontal direction based on testing of eight postmortem human subjects., Methods: For each subject, the left and right hemipelvis complex was axially loaded using a previously developed test configuration. Six positions were defined from a seated femur neutral condition, combining flexed, neutral, and extended femur positions with abducted, neutral, and adducted positions., Results: Axial injury tolerances based on peak force were found to be 6,850 +/- 840 N in the extended, neutral position and 4,080 +/- 830 N in the flexed, neutral position. From the flexed neutral orientation, the peak axial force increased 18% for 20 degrees abduction and decreased 6% for 20 degrees adduction. From the extended, neutral orientation, the peak axial force decreased 4% for 20 degrees abduction and decreased 3% for 20 degrees adduction. However, as there is evidence that increases in loading may occur after the initiation of fracture, the magnitude of the peak force is likely related to the extent of injury, not to the initial tolerance. Using the axial femur force at the initiation of fracture (assessed with acoustic crack sensors) as a potentially more relevant indicator of injury may lower the existing injury criteria. This fracture initiation force varied by position from 3,010 +/- 560 N in the flexed, neutral position to 5,470 N in the extended, abducted position. Further, there was a large position-dependent variation in the ratio of fracture initiation force to the peak axial force. The initiation of fracture was 83% of the peak axial force in the extended, abducted position, but the ratio was 34% in the extended, adducted position., Conclusions: This may have significant implications for the development of pelvic injury criteria by automobile designers attempting to mitigate pelvis injuries.

Thoracic deformation under an applied load is an established indicator of injury risk, but the force required to achieve an injurious level of deformation currently is not understood adequately. This article evaluates how two potentially important factors, loading condition and muscle tensing, affect the structural response of the dynamically loaded thorax. Structural models of two human cadaver thoraxes and two porcine thoraxes were used to quantify the effects. The human cadavers, which represent anthropometric extremes, were subjected to anterior loading from (1) a 5.1-cm-wide belt oriented diagonally (i.e., seatbelt-like loading), (2) a 15.2-cm-diameter rigid hub, and (3) a 20.3-cm-wide belt oriented laterally (i.e., a distributed load). A structural model having the mathematical formulation of a quasilinear viscoelastic material model was used to model the elastic and viscous response, with ramp-hold tests used to determine the model coefficients. The effect of thoracic musculature was assessed using similar ramp-hold tests on the porcine subjects, each with and without forced muscle contraction. Even maximally contracted thoracic musculature is shown to have a minimal effect on the response, with similar elastic and viscous characteristics exhibited by each subject regardless of muscle tone. The elastic response is shown to be approximately a factor of three stiffer for diagonal belt loading and for this distributed loading condition than for the hub loading, indicating that the response is influenced most by the particular anatomical structures that are engaged and, secondarily, by the area of load application. Specifically, shoulder involvement is shown to have a strong influence. The force relaxation is found to be pronounced, but insensitive to the loading condition, with long-time force relaxation coefficients (G( infinity )) in the range of 0.1 to 0.3. The findings of this study provide restraint-specific guidelines for the force-deflection characteristics of both physical and computational thoracic models.