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Development and Validation of an Age-Specific Lower Extremity Finite Element Model for Simulating Pedestrian Accidents

Development and Validation of an Age-Specific Lower Extremity Finite Element Model for Simulating... Hindawi Applied Bionics and Biomechanics Volume 2018, Article ID 5906987, 12 pages https://doi.org/10.1155/2018/5906987 Research Article Development and Validation of an Age-Specific Lower Extremity Finite Element Model for Simulating Pedestrian Accidents 1 1,2 1 3 Jing Huang , Yongcheng Long , Yu Yan, and Lin Hu Research Centre of Vehicle and Traffic Safety, State Key Laboratory of Advanced Design and Manufacturing for Vehicle Body, Hunan University, Changsha, China State Key Laboratory of Vehicle NVH and Safety Technology, Chongqing, China College of Automotive and Mechanical Engineering, Changsha University of Science & Technology, Changsha, China Correspondence should be addressed to Jing Huang; huangjing926@hnu.edu.cn Received 3 December 2017; Revised 28 January 2018; Accepted 14 February 2018; Published 21 March 2018 Academic Editor: Jingwen Hu Copyright © 2018 Jing Huang et al. This is an open access article distributed under the Creative Commons Attribution License, which permits unrestricted use, distribution, and reproduction in any medium, provided the original work is properly cited. The objective of the present study is to develop an age-specific lower extremity finite element model for pedestrian accident simulation. Finite element (FE) models have been used as a versatile tool to simulate and understand the pedestrian injury mechanisms and assess injury risk during crashes. However, current computational models only represent certain ages in the population, the age spectrum of the pedestrian victims is very large, and the geometry of anatomical structures and material property of the lower extremities changes with age for adults, which could affect the injury tolerance, especially in at-risk populations such as the elderly. The effects of age on the material mechanical property of bone and soft tissues of the lower extremities as well as the geometry of the long bone were studied. Then an existing 50th percentile male pedestrian lower extremity model was rebuilt to depict lower extremity morphology for 30- to 70-year-old (YO) individuals. A series of PMHS tests were simulated to validate the biofidelity and stability of the created age-specific models and evaluate the lower extremity response. The development of age-specific lower extremity models will lead to an improved understanding of the pedestrian lower extremity injury mechanisms and injury risk prediction for the whole population in vehicle-pedestrian collision accidents. 1. Introduction Pedestrian injuries are preventable; however, success- ful interventions to protect pedestrians and promote safe Pedestrians are road users vulnerable to traffic accidents, who traveling require a better understanding of the injury suffer high injury rate and mortality rate. WHO reported that mechanisms and risk factors for pedestrian crashes. Many more than one-fifth of the people killed on the world’s roads researchers established lower extremity models to study its injury mechanisms in pedestrian collision accidents. Zhang each year are pedestrians [1]; the situation is worse in China, as the road traffic situation in China is more complex than et al. [3, 4] improved and verified the material model of the that of developed countries in Europe and America. The long bone and ligament of the lower extremity based on the THUMS model. Untaroiu et al. [5], using a human body vehicle-pedestrian-mixed traffic in most urban and rural roads presents hidden dangers, inhibiting pedestrian safety. extremity model, combined with an accident reconstruction method, simulated pedestrian lower extremity fractures in The lower extremities and head are the main injury body regions for pedestrians in a vehicle-pedestrian collision acci- collision. Wang et al. [6] used the multibody system and dent, accounting for 31.2% and 32.4%, respectively [2]. Com- finite element model to study the long bone fracture of the pared with head injuries, lower extremity injuries rarely lead lower limb based on two cases of real pedestrian accidents. directly to fatalities; however, they often cause long-term or Meng et al. [7] established a 6-YO child’s lower extremity long bone model and validated this model with a dynamic life-long disabilities. 2 Applied Bionics and Biomechanics load three-point bending test, then discussed the effect of elastic modulus on the injury. Shen et al. [8] established a External periosteum 10-YO child pelvis and lower extremity FE model with growth plates for pedestrian protection. Kong [9] analyzed the pedestrian-vehicle collision accidents in Changsha using statistical methods. Her conclusions noted that children aged 0 to 10 years and middle-aged people over 46 years were Internal periosteum more likely than other age groups to suffer higher fatality rates, with death rates increasing with age for people over Cortical bone 46 YO. In literature [10], the data in the American Pedestrian Injury Causation Study (PCDS) database was analyzed statis- Figure 1: Idealized long bone section. tically; the results showed that the injury risk to the lower extremities of the elderly was 2.44 times higher than that of the young people, and the elderly would suffer more serious injuries and require longer treatment cycles. Hu et al. research results showed that age had an important effect on the injury risk when pedestrians collided with different front-shaped vehicles [11]. All of these research works indi- cate that age is one of the most important factors affecting the injury risk of pedestrian lower extremities; what is more, the effect of age on injury is nonlinear [12–14]. 4 On the other hand, with the continued rapid growth of the elderly population of adults aged 60+ years, which has increased to 210 million (15.5% of the total population by the end of 2014 [15]), the proportion of the elderly in traf- fic accidents increased gradually and was as high as 40% [16], according to the statistics of the Ministry of Public Security. Compared with that of young people, the geom- 0 20 40 60 80 100 etry of anatomical structures and material property of the Section position (%) elderly are quite different [17]. With the further develop- Femur D Femur D P M ment of pedestrian safety research, it is critical to understand Tibia D Tibia D P M the biomechanics change of the human lower extremities with age for pedestrian protection and the establishment Figure 2: Variation of geometric dimensions of long bone cross- of especially vulnerable road user models (such as chil- section (male). dren, the fifth percentile women, obese people, and the elderly) [18]. Table 1: The scaling ratio of long bone cross-section for a 70-YO The finite element model provides a useful tool to assess adult. injury risk and to study the injury biomechanics, while cur- rent models are limited to certain ages in the population. Femur Tibia Section position Therefore, it cannot reflect the difference of pedestrian injury D (%) D (%) D (%) D (%) P M P M at different ages in the accident. The objective of the present 20% 1.3 4.3 5.0 12.9 study is to investigate the geometric changes and material 35% 1.3 9.6 3.7 16.6 property changes with aging for pedestrian lower extremities 50% 1.8 15.2 3.3 15.7 and to develop and validate the age-specific FE models of 65% 2.6 19.4 4.3 14.8 pedestrian lower extremities to accurately model lower extremity morphology and material property for ages between 80% 4.9 19.2 5.3 11.1 30 and 70 years. 2. Materials and Methods are calculated in (1), (2), and (3), respectively, together with external diameter (D ) and internal diameter (D ). P M 2.1. Geometric Changes with Aging. In total, 320 femoral and 99 tibial midshafts derived from individuals aged 21–99 years 2 πD were examined and measured [19–21]; the aim was to deter- TA = , mine the age-related changes in the structure of the human long bone. The geometry of the long bone cross-section nat- πD urally changed with aging, and the change is associated with 2 MA = , marrow cavity area changes. Figure 1 shows the idealized long bone cross-section. The CA = TA − MA 3 total area (TA), medullary area (MA), and cortical area (CA) Increase ratio per decade (%) Applied Bionics and Biomechanics 3 Handle 3D domain Handle Hyper morph Cortical Edge Figure 3: Changes of skeletal mesh with age in finite element model. Table 2: The long bone material properties of lower extremity model. v = 0.33 mm/s Material The young The elderly Femur parameters (30 YO) (70 YO) Density (kg/m ) 2000 2000 Elastic modulus 16.2 13.9 (MPa) Femoral cortical bone Poisson’s ratio 0.3 0.3 Tibia Yield stress 100.22 94.4 Limit strain 0.032 0.019 Density (kg/m ) 2000 2000 Figure 4: Knee ligament simulation model. Elastic modulus 18.3 15.7 (MPa) Tibial cortical bone Poisson’s ratio 0.3 0.3 Yield stress 120.3 113.28 Table 3: The ligament material properties. Limit strain 0.034 0.020 Material The young The elderly Density (kg/m ) 1000 1000 parameters (30 YO) (70 YO) Elastic modulus 752 816.4 Density (kg/m ) 1000 1000 (MPa) Femur Volume modulus cancellous bone Poisson’s ratio 0.45 0.45 4.31 3.5 (GPa) Yield stress 13.25 10.22 C1 34.29 22.13 Limit strain 0.134 0.134 Knee ligament C3 1.54 0.6 Density (kg/m ) 1000 1000 C4 152.85 147.8 Elastic modulus C5 836.42 695.66 752 591.6 (MPa) Tibial Failure strain 0.45 0.263 cancellous bone Poisson’s ratio 0.45 0.45 C1: first Mooney-Rivlin constant; C3: constant scaling of the collagen Yield stress 11.04 8.24 exponential stresses; C4: constant controlling rate of the rise of collagen Limit strain 0.134 0.134 exponential stresses; C5: modulus of straightened collagen fibers. According to the values of TA, MA, and CA of different shortage of anthropometry data related with age, its geomet- ages obtained by Ruff et al. [21], the increase ratio per decade ric change with age is assumed to be the same with the tibia. for D and D of the adult male long bone at 5 cross-sections Taking a 70-YO adult male for example, the basic lower P M can be calculated, as shown in Figure 2; then the D and D extremity model used in this manuscript is derived from an P M of a certain age can be scaled from the basic model via the adult male aging 26 years, according to Figure 2; the scaling corresponding scale ratio. As for the fibula, due to the ratios of the D and D of the femur and tibia are shown P M 4 Applied Bionics and Biomechanics Femur Femur PCL Equivalent muscle LCL ACL Knee joint Patella Fibula MCL Tibia Fibula Meniscus Tibia Foot bone Figure 5: Finite element model of human lower extremity with age characteristics. Table 4: Biomechanical cadaver tests for lower extremities. in Table 1; then the D and D of the 70-YO adult can be P M scaled from the basic model via the corresponding scale ratio Age of the at five cross-sections. Item Cadaver test PMHS (YO) Geometric changes of anatomical structures with aging Thigh and calf Kerrigan et al. [35] 58.5 ± 4.8/9.3 were implemented by changing the long bone cross- Ligaments Dommelen et al. [35, 36] 63 ± 3.3; 53.4 ± 9.9 section—model morphing can be used to generate models of all ages accurately and efficiently [22]. The HyperMorph Knee Bose et al. [37] 53.4 ± 9.9 module of HyperMesh software is used to rebuild an existing The whole lower Kajzer et al. [38] 51 ± 15 50th percentile male adult FE model to obtain the lower extremities extremity models in the full spectrum of ages, as shown in Figure 3. 3D adjust domain was established in 5 cross- sections in long bone meshes firstly, and the control point correlation between age and elastic modulus and ultimate handle of the experimental point was setup in a correspond- stress is developed as ing position based on previous geometric study results. Then Elastic modulus = 473 3+14 99 age − 0 19 age , the location of the control point could be adjusted manually to complete the update of lone bone meshes to get the age- 2 Ultimate stress = 8 94 + 0 13 age − 0 002 age specific models. The material properties of the cortical bone are simulated by the material model of isotropic elastoplastic ( MAT_PIE- 2.2. Material Property Changes with Aging CEWISE_LINEAR_PLASTICITY); as the slope of its stress- strain curve in the plastic stage remains unchanged [24], it 2.2.1. Long Bone. The cancellous bone of the lower extremity can be assumed that the tangent modulus in the plastic stage is a kind of porous structure composed of irregularly will not change with aging. If the strain of meshes rose to the arranged trabecular bones; its mechanical properties are sim- failure strain, the fracture would occur and be simulated by ilar to foamed aluminum. When compressed, there is a sig- mesh deletion. The Cowper-Symonds method is used to sim- nificant elastic phase and the stress is nearly unchanged ulate the effect of strain rate on the material properties, with after the yield point; the limit stress is almost equal to the yield stress scaling equation shown in the yield stress. The material properties of the cancellous bone showed obvious changes with aging because of the 1/p loss of calcification and fibrosis. Its material properties σε, ε = σ ε 1+ , 5 can be simulated by the material model of dynamic elasto- C plastic ( MAT_PLASTIC_KINEMATIC), and the ultimate stress is set to 13.4% according to the research results of where σ ε is the initial yield stress, while ε is the strain rate. literature [23]. C and P are the strain rate transformation parameters. In this The quasistatic compression test data of the cancellous paper, C is 360.5 and P is 3.6. The elastic modulus, ultimate stress, and failure strain of bone from the ages of 16 to 83 years [23] were subjected to quadratic polynomial fitting. The results showed that the the cortical bone of different age groups are obtained by mechanical properties of cancellous bone increased from 20 regression analysis of corresponding test data in literatures to 40 years, while there is a sudden drop after 40 years. The [25, 26] and [27]. The correlation between age and elastic Applied Bionics and Biomechanics 5 (a) (b) Figure 6: Thigh and calf three-point bending test model: (a) thigh and (b) calf. Cylindrical bar Metal cylinder Rotary bearing Knee joint Fixed bearing Sliding bearing Figure 7: Four-point bending test device and finite element model. 400 N 400 N Fixed point Fixed point Fixed point 40 km/h 6.25 kg Marker location 40 km/h Moving ground Fixed ground 6.25 kg (a) (b) Figure 8: Lower extremity bending (a) and shear (b) simulation model. modulus, ultimate stress, and failure strain is developed as bones, as they undergo the same changes with age. The presented in (6), (7), and (8), respectively. material property parameters of the cortical bone and can- cellous bone of the long bones for different ages can be Elastic modulus = 18 01 − 0 059 age, 6 calculated according to the above fitting formulas and corre- sponding scaling coefficients. Table 2 shows the lower Ultimate stress = 130 8 − 0 52 age, extremity long bone material properties of ages 30 years and 70 years for example. Failure strain = 4 23 − 0 033 age Based on the research results of the literature [28, 29], 2.2.2. Ligaments. The ligaments in the knee are connected the change of material property of the cancellous bone and to the bones, which stabilize and restrict the movement of cortical bone is assumed to be the same for lower limb long the knee, including the patellar ligament, meniscofemoral 6 Applied Bionics and Biomechanics 0 10 20 30 40 50 60 0 10 20 30 40 50 60 Displacement (mm) Displacement (mm) Test corridor Elder Test corridor Elder Average Average Younger Younger (a) (b) Figure 9: Force displacement curves of the impactor in three-point bending simulation of the (a) thigh and (b) calf. ligament, medial collateral ligament (MCL), lateral collat- eral ligament (LCL), anterior cruciate ligament (ACL), and posterior cruciate ligament (PCL). The diameter of collagen fibers decreased, while the fiber content increased with aging. For example, the maximum fiber diameter is 180 nm when a man is 15–19 YO and reduced to 110 nm after 60 YO [30, 31]. The change of collagen fibrils will affect the Fracture of femur mechanical properties of the ligament; the ultimate tension, especially, will decline with aging [31]. Figure 10: Femur fracture location in thigh three-point bending In the present study, the ligaments are simulated by the simulation. solid elements to accurately model the geometrical shape of each ligament and their contact with the surrounding tissue. using quadrilateral shell elements. Muscle and skin are mod- The hyperelastic material constitutive model ( MAT_- eled using the solid element and shell element, respectively. SOFT_TISSUE) is used to simulate its mechanical proper- Ligaments are modeled using the solid element and one- ties [32], and the failure of the first-order principal strain dimensional beam element together. The baseline model is is defined for the elements, with the laceration of ligament adjusted according to the pedestrian’s standing posture. simulated by element deletion. Then the previous research results of the geometric changes The experiments of the knee joint ligament carried out by and material property changes with aging are applied to build Woo et al. [33] were simulated; the simulation model is the age-specific lower extremity FE models—including the shown in Figure 4. The ligament properties for different ages adjustment of the material properties and the geometry can be obtained by parameter computational inverse based morphing of the femur, tibia, and fibula, as shown in on the ligament tensile test curves. Figure 5. Then two FE models of the pedestrian lower Table 3 shows the knee joint ligament material properties extremity of typical ages 30 and 70 years are established to of individuals aging 30 to 70 years; for example, C1, C3, C4, investigate the effect of age on injury risk. The selection of and C5 are the parameters of the material model. 30- and 70-YO models was based on a previous recommen- dation that defined a young adult group between 16 and 35 2.3. Development of the Age-Specific Lower Extremity FE YO and elderly group as 66 YO and older [34] . Model. The baseline pedestrian lower extremity model is derived from the Global Human Body Models Consortium 2.4. Model Validation. A series of cadaver test data were used (GHBMC) average male occupant model. The GHBMC is to validate the biofidelity and stability of age-specific pedes- representative of a 50th percentile male adult and was based trian lower extremity FE models, as shown in Table 4. These on medical images of a 26 YO individual. The lower extrem- validation tests were simulated in LS-DYNA software ity model includes the long bone, muscle, ligament, skin, and according to the published test information. other tissues. The cortical bone and cancellous bone of the long bone shaft are modeled using hexahedral elements. 2.4.1. Validation at the Component Level. In Kerrigan’s test The cortical bone covering the long bone ends is modeled [35], the thigh and calf were extracted from PMHS. The Impact force (kN) Impact force (kN) Applied Bionics and Biomechanics 7 Fracture of the fibula at both ends Fracture of the tibia and fibula Fracture of the tibia and fibula (a) (b) Figure 11: Tibia and fibula fracture location in calf three-point bending simulation: (a) the young and (b) the elderly. 2.0 muscle tissues of two ends were removed, and the distal and proximal ends of the femur and tibia were potted in cups and fixed with polyurethane. An impactor driven by a universe machine loaded the thigh and calf at the middle-shaft loca- 1.5 tion at the speed of 1.5 m/s to simulate the loading condition of pedestrian lower extremities in a vehicle-pedestrian colli- sion accident. Then two age-specific FE models of the pedes- 1.0 trian lower extremities of typical ages 30 and 70 YO were used to simulate the same tests with the same experiment set- tings and boundaries. The finite element models are shown in 0.5 Figure 6. Ligament failure caused by lateral bending is a common knee injury for pedestrian during vehicle-pedestrian collision accident. Kerrigan et al. [35] and Bose et al. [37] designed a 0.0 dynamic four-point bending test to estimate knee tolerance. 0 2 4 6 8 101214 Figure 7 illustrates the test principle: isolated knee parts Displacement (mm) were potted in specific cups that rotated around support Simulation age = 30 years Test age = 22–35 years joints during tests. While the distal support connected to Simulation age = 70 years Test age = 60–97 years the tibia was fixed, the proximal support connected to the femur was allowed to move horizontally. The angular speed Figure 12: Ligament ACL displacement force curve comparison of the knee was about 1 deg/ms during tests to simulate the between experiment and simulation. knee-bending load when pedestrian crashed at a speed of 40 Km/h, and the bending moment was measured by a load cell 3. Results and Discussion connected to the femur extension bar. Corresponding simu- lation models were built to perform the same tests in ls- The force displacement curves of the impactor in thigh and dyna, and then the simulation results were compared with calf three-point bending simulations are shown in Figure 9. the test data. Both the simulation results of young (30 YO) and elderly (70 YO) are in the test corridor and consistent with the experimen- 2.4.2. Validation at the Lower Extremity Level. To evaluate tal results, though different from each other. This indicates the whole lower extremity response, 2 loading cases, bending that the response of the young and the elderly is much differ- and shear, were simulated to assess the importance of geo- ent. The impactor force rises slowly initially but is followed metric and material property changes with aging. by a sharp increase. This is because the impactor makes con- According to the tests of Kajzer et al. [38], as shown in tact with the skin, muscle, and other soft tissues first, and Figure 8, the lower extremity was extracted from PMHS on when the femur begins to bend to deform, the force increases. the hip joint and fixed flat on a board to maintain stability. In thigh three-point bending simulation, the femoral The proximal of the femur was fixed with screws, while the fracture occurred in both cases of the young (30 YO) and distal of the femur was fixed with a fixed plate to limit its hor- the elderly (70 YO), as shown in Figure 10. For the elderly, izontal movement. The force sensor would calculate the the femur fracture occurred when the displacement of the bending moment of the knee joint. A force of 400 N was impactor is 40 mm with the impact force 3.2 kN, while the loaded at the hip to simulate the load received by the lower corresponding data of the young is 50 mm and 6.1 kN. extremities when standing. The bending and shear impact In calf bending simulation, both the tibia and fibula are load was conducted at 40 km/h with a 6.25 kg I-shaped impac- fractured, while the fracture locations are different, as shown tor striking the ankle joint and the knee joint, respectively. in Figure 11. The elderly’s fibula was fractured at both ends The impactor was wrapped with a foam of 100 mm × and the middle shaft, but the young’s fibula was only frac- 120 mm × 50 mm at the front. tured in the middle. For the elderly, the fibula fracture Force (kN) 8 Applied Bionics and Biomechanics Rupture of MCL 0 ms 28 ms 40 ms Figure 13: Knee four-point bending simulation process. occurred when the displacement of the impactor is 29 mm, 300 while the young is 36 mm, and the curve showed an obvious decline when the fibula fractured. After the fibula fracture, the impactor continued to load on the tibia, and both the young and the elderly suffered tibia fracture subsequently when the displacement was 50 mm and 42 mm, respectively. The comparison among the ligament displacement force is shown in Figure 12 in terms of results from the simulation of the models and experimental data of Woo et al. [33] in the tensile tests of the femur-ACL-tibia complex. Both the simulation curves for the young and the elderly are well aligned with the experimental results. It is believed that the material property parameters of the ligaments for different ages are reasonable and can reflect the ligament injury at different ages. 0 5 10152025 Figure 13 shows the results of a knee joint four-point Bending angle (°) bending simulation of the young. The medial collateral liga- ment (MCL) is completely ruptured at 28 ms near the tibia Test corridor (Nm) Younger (Nm) junction, which coincides with the test results performed by Elder (Nm) Kerrigan et al. [35] and Bose et al. [37]. The bending-angle-to-bending-moment curves of the Figure 14: Curve of bending angle and bending moment of knees in knee joint are shown in Figure 14. four-point bending simulation. The simulation results of the elderly are in the test corri- dor, while the peak of the young is outside the corridor. This may be due to the ages of the PMHS, as they are between 44 of the elderly is obviously bigger than the young beyond YO and 80 YO—therefore, it is reasonable that the peak of 10 ms, and reached 5 at 20 ms, the similar trend occurred the young (30 YO) is outside the corridor. The simulation in knee joint shear displacement curves. This is because the curves coincide with the test curves before MCL rupture, knee ligament strength of the elderly is much lower than which indicates that the material properties of the ligaments the young and their ligaments usually rupture earlier in the are reasonable. At the beginning, the bending moment same collision condition. For example, the MCL and PCL increases with the bending angle and reaches the maximum ruptured at 11.5 ms and 14.5 m, respectively, in the simula- value when the MCL is about to rupture and then the bend- tion. It was significantly ahead of the results of the young, ing moment decreases sharply. The maximum bending which is 18.5 ms and 20 ms, respectively, and then induced moment of the elderly is about 110 Nm with a bending angle larger knee bending angle and shear displacement. The of 11 kinematics of lower extremity and ligament rupture in a , while the maximum bending moment of the young is about 270 Nm with a bending angle of 18 , far greater than bending test simulation is shown in Figure 16 (take the the elderly. young for example). The time history curves of impactor force, knee joint The time history curve of the impact force and knee joint bending angle, and knee joint shear displacement in lower shear displacement in the lower extremity shear simulation extremity bending simulation are shown in Figure 15. It indi- are shown in Figure 17. It indicates that the simulation results cates that the simulation results of the young and the elderly of the young and the elder show a similar linear shape as that show a linear shape similar to that reported in tests and all are reported in tests and are mostly in the test corridor, except in the test corridor. for the impact force of the young. For the impact force, there is not much difference The peak impact force of the elderly is 5.2 kN, lower than 6.0 kN of the young, and appeared earlier at about 8 ms, while between the simulation results of the young and the elderly. They both reach their maximum value of 4.5 kN at 4 ms. This the knee joint shear displacement of the elderly is obviously may be due to the same kinetic energy of the impactor and larger than that of the young beyond 8 ms. It is possibly the quality of the lower extremity. While the bending angle related to the elderly’s femur fracture occurring at around Bending moment (Nm) Applied Bionics and Biomechanics 9 0 0 0 5 10 15 20 25 30 0 5 10 15 20 25 Time (ms) Time (ms) Test corridor Test corridor Younger Younger Elder Elder (a) (b) −10 0 5 10 15 20 25 30 Time (ms) Test corridor Younger Elder (c) Figure 15: Lower extremity bending simulation results: (a) impact force, (b) knee joint bending angle, and (c) knee joint shear displacement. 8 ms, resulting in the increase of rotation and lateral move- techniques. To evaluate the lower extremity response, a series ment at the fracture point and the decrease of the impact of PMHS tests were simulated to validate the confidence of the force. The detailed injuries of the elderly and the young are models and to assess the importance of geometric and mate- compared in Figure 18. The young only suffered partial rial property changes with aging. The whole age-specificFE femur fracture, while the elder suffered full femur fracture models of pedestrian lower extremity showed numerical sta- and fibula fracture. These injuries were coordinated with that bility, and, in all validation simulations, the response of the of samples 4 and 17 in the PMHS test [38]. young model and the elderly is different from each other. Development of age-specific FE models of the lower extremity will provide valuable tools for understanding variations in 4. Conclusion lower extremity injury patterns due to vehicle-pedestrian col- In the present study, the changes of geometric and mate- lision accidents across populations and in the design of new vehicles with devices for pedestrian protection. rial properties of the lower extremity with aging were studied and age-specific FE models of the lower extremity Further study will involve the sex factor and the geometry for pedestrian-vehicle accident simulation were developed changes of the femoral head/neck and ankle with age. These for 30-YO and 70-YO male pedestrian using morphing would be investigated to establish a pedestrian lower limb Impact force (N) Shear displacement (mm) Bending angle (N) 10 Applied Bionics and Biomechanics Rupture of PCL Rupture of MCL 0 ms 6 ms 12 ms 18.5 ms 20 ms Figure 16: Dynamic simulation of the lower extremity simulation process. 7000 100 0 0 0 5 10 15 20 25 30 010 20 30 40 Time (ms) Time (ms) Test corridor Test corridor Younger Younger Elder Elder (a) (b) Figure 17: Lower extremity shear simulation results: (a) impact force and (b) knee joint shear displacement. Partial fracture of the femur Fracture of the femur Fracture of the fibula (a) (b) Figure 18: Comparison of long bone injury: (a) the young and (b) the elderly. Impact force (N) Shear displacement (mm) Applied Bionics and Biomechanics 11 model with higher biofidelity. 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Development and Validation of an Age-Specific Lower Extremity Finite Element Model for Simulating Pedestrian Accidents

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Copyright © 2018 Jing Huang et al. This is an open access article distributed under the Creative Commons Attribution License, which permits unrestricted use, distribution, and reproduction in any medium, provided the original work is properly cited.
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Hindawi Applied Bionics and Biomechanics Volume 2018, Article ID 5906987, 12 pages https://doi.org/10.1155/2018/5906987 Research Article Development and Validation of an Age-Specific Lower Extremity Finite Element Model for Simulating Pedestrian Accidents 1 1,2 1 3 Jing Huang , Yongcheng Long , Yu Yan, and Lin Hu Research Centre of Vehicle and Traffic Safety, State Key Laboratory of Advanced Design and Manufacturing for Vehicle Body, Hunan University, Changsha, China State Key Laboratory of Vehicle NVH and Safety Technology, Chongqing, China College of Automotive and Mechanical Engineering, Changsha University of Science & Technology, Changsha, China Correspondence should be addressed to Jing Huang; huangjing926@hnu.edu.cn Received 3 December 2017; Revised 28 January 2018; Accepted 14 February 2018; Published 21 March 2018 Academic Editor: Jingwen Hu Copyright © 2018 Jing Huang et al. This is an open access article distributed under the Creative Commons Attribution License, which permits unrestricted use, distribution, and reproduction in any medium, provided the original work is properly cited. The objective of the present study is to develop an age-specific lower extremity finite element model for pedestrian accident simulation. Finite element (FE) models have been used as a versatile tool to simulate and understand the pedestrian injury mechanisms and assess injury risk during crashes. However, current computational models only represent certain ages in the population, the age spectrum of the pedestrian victims is very large, and the geometry of anatomical structures and material property of the lower extremities changes with age for adults, which could affect the injury tolerance, especially in at-risk populations such as the elderly. The effects of age on the material mechanical property of bone and soft tissues of the lower extremities as well as the geometry of the long bone were studied. Then an existing 50th percentile male pedestrian lower extremity model was rebuilt to depict lower extremity morphology for 30- to 70-year-old (YO) individuals. A series of PMHS tests were simulated to validate the biofidelity and stability of the created age-specific models and evaluate the lower extremity response. The development of age-specific lower extremity models will lead to an improved understanding of the pedestrian lower extremity injury mechanisms and injury risk prediction for the whole population in vehicle-pedestrian collision accidents. 1. Introduction Pedestrian injuries are preventable; however, success- ful interventions to protect pedestrians and promote safe Pedestrians are road users vulnerable to traffic accidents, who traveling require a better understanding of the injury suffer high injury rate and mortality rate. WHO reported that mechanisms and risk factors for pedestrian crashes. Many more than one-fifth of the people killed on the world’s roads researchers established lower extremity models to study its injury mechanisms in pedestrian collision accidents. Zhang each year are pedestrians [1]; the situation is worse in China, as the road traffic situation in China is more complex than et al. [3, 4] improved and verified the material model of the that of developed countries in Europe and America. The long bone and ligament of the lower extremity based on the THUMS model. Untaroiu et al. [5], using a human body vehicle-pedestrian-mixed traffic in most urban and rural roads presents hidden dangers, inhibiting pedestrian safety. extremity model, combined with an accident reconstruction method, simulated pedestrian lower extremity fractures in The lower extremities and head are the main injury body regions for pedestrians in a vehicle-pedestrian collision acci- collision. Wang et al. [6] used the multibody system and dent, accounting for 31.2% and 32.4%, respectively [2]. Com- finite element model to study the long bone fracture of the pared with head injuries, lower extremity injuries rarely lead lower limb based on two cases of real pedestrian accidents. directly to fatalities; however, they often cause long-term or Meng et al. [7] established a 6-YO child’s lower extremity long bone model and validated this model with a dynamic life-long disabilities. 2 Applied Bionics and Biomechanics load three-point bending test, then discussed the effect of elastic modulus on the injury. Shen et al. [8] established a External periosteum 10-YO child pelvis and lower extremity FE model with growth plates for pedestrian protection. Kong [9] analyzed the pedestrian-vehicle collision accidents in Changsha using statistical methods. Her conclusions noted that children aged 0 to 10 years and middle-aged people over 46 years were Internal periosteum more likely than other age groups to suffer higher fatality rates, with death rates increasing with age for people over Cortical bone 46 YO. In literature [10], the data in the American Pedestrian Injury Causation Study (PCDS) database was analyzed statis- Figure 1: Idealized long bone section. tically; the results showed that the injury risk to the lower extremities of the elderly was 2.44 times higher than that of the young people, and the elderly would suffer more serious injuries and require longer treatment cycles. Hu et al. research results showed that age had an important effect on the injury risk when pedestrians collided with different front-shaped vehicles [11]. All of these research works indi- cate that age is one of the most important factors affecting the injury risk of pedestrian lower extremities; what is more, the effect of age on injury is nonlinear [12–14]. 4 On the other hand, with the continued rapid growth of the elderly population of adults aged 60+ years, which has increased to 210 million (15.5% of the total population by the end of 2014 [15]), the proportion of the elderly in traf- fic accidents increased gradually and was as high as 40% [16], according to the statistics of the Ministry of Public Security. Compared with that of young people, the geom- 0 20 40 60 80 100 etry of anatomical structures and material property of the Section position (%) elderly are quite different [17]. With the further develop- Femur D Femur D P M ment of pedestrian safety research, it is critical to understand Tibia D Tibia D P M the biomechanics change of the human lower extremities with age for pedestrian protection and the establishment Figure 2: Variation of geometric dimensions of long bone cross- of especially vulnerable road user models (such as chil- section (male). dren, the fifth percentile women, obese people, and the elderly) [18]. Table 1: The scaling ratio of long bone cross-section for a 70-YO The finite element model provides a useful tool to assess adult. injury risk and to study the injury biomechanics, while cur- rent models are limited to certain ages in the population. Femur Tibia Section position Therefore, it cannot reflect the difference of pedestrian injury D (%) D (%) D (%) D (%) P M P M at different ages in the accident. The objective of the present 20% 1.3 4.3 5.0 12.9 study is to investigate the geometric changes and material 35% 1.3 9.6 3.7 16.6 property changes with aging for pedestrian lower extremities 50% 1.8 15.2 3.3 15.7 and to develop and validate the age-specific FE models of 65% 2.6 19.4 4.3 14.8 pedestrian lower extremities to accurately model lower extremity morphology and material property for ages between 80% 4.9 19.2 5.3 11.1 30 and 70 years. 2. Materials and Methods are calculated in (1), (2), and (3), respectively, together with external diameter (D ) and internal diameter (D ). P M 2.1. Geometric Changes with Aging. In total, 320 femoral and 99 tibial midshafts derived from individuals aged 21–99 years 2 πD were examined and measured [19–21]; the aim was to deter- TA = , mine the age-related changes in the structure of the human long bone. The geometry of the long bone cross-section nat- πD urally changed with aging, and the change is associated with 2 MA = , marrow cavity area changes. Figure 1 shows the idealized long bone cross-section. The CA = TA − MA 3 total area (TA), medullary area (MA), and cortical area (CA) Increase ratio per decade (%) Applied Bionics and Biomechanics 3 Handle 3D domain Handle Hyper morph Cortical Edge Figure 3: Changes of skeletal mesh with age in finite element model. Table 2: The long bone material properties of lower extremity model. v = 0.33 mm/s Material The young The elderly Femur parameters (30 YO) (70 YO) Density (kg/m ) 2000 2000 Elastic modulus 16.2 13.9 (MPa) Femoral cortical bone Poisson’s ratio 0.3 0.3 Tibia Yield stress 100.22 94.4 Limit strain 0.032 0.019 Density (kg/m ) 2000 2000 Figure 4: Knee ligament simulation model. Elastic modulus 18.3 15.7 (MPa) Tibial cortical bone Poisson’s ratio 0.3 0.3 Yield stress 120.3 113.28 Table 3: The ligament material properties. Limit strain 0.034 0.020 Material The young The elderly Density (kg/m ) 1000 1000 parameters (30 YO) (70 YO) Elastic modulus 752 816.4 Density (kg/m ) 1000 1000 (MPa) Femur Volume modulus cancellous bone Poisson’s ratio 0.45 0.45 4.31 3.5 (GPa) Yield stress 13.25 10.22 C1 34.29 22.13 Limit strain 0.134 0.134 Knee ligament C3 1.54 0.6 Density (kg/m ) 1000 1000 C4 152.85 147.8 Elastic modulus C5 836.42 695.66 752 591.6 (MPa) Tibial Failure strain 0.45 0.263 cancellous bone Poisson’s ratio 0.45 0.45 C1: first Mooney-Rivlin constant; C3: constant scaling of the collagen Yield stress 11.04 8.24 exponential stresses; C4: constant controlling rate of the rise of collagen Limit strain 0.134 0.134 exponential stresses; C5: modulus of straightened collagen fibers. According to the values of TA, MA, and CA of different shortage of anthropometry data related with age, its geomet- ages obtained by Ruff et al. [21], the increase ratio per decade ric change with age is assumed to be the same with the tibia. for D and D of the adult male long bone at 5 cross-sections Taking a 70-YO adult male for example, the basic lower P M can be calculated, as shown in Figure 2; then the D and D extremity model used in this manuscript is derived from an P M of a certain age can be scaled from the basic model via the adult male aging 26 years, according to Figure 2; the scaling corresponding scale ratio. As for the fibula, due to the ratios of the D and D of the femur and tibia are shown P M 4 Applied Bionics and Biomechanics Femur Femur PCL Equivalent muscle LCL ACL Knee joint Patella Fibula MCL Tibia Fibula Meniscus Tibia Foot bone Figure 5: Finite element model of human lower extremity with age characteristics. Table 4: Biomechanical cadaver tests for lower extremities. in Table 1; then the D and D of the 70-YO adult can be P M scaled from the basic model via the corresponding scale ratio Age of the at five cross-sections. Item Cadaver test PMHS (YO) Geometric changes of anatomical structures with aging Thigh and calf Kerrigan et al. [35] 58.5 ± 4.8/9.3 were implemented by changing the long bone cross- Ligaments Dommelen et al. [35, 36] 63 ± 3.3; 53.4 ± 9.9 section—model morphing can be used to generate models of all ages accurately and efficiently [22]. The HyperMorph Knee Bose et al. [37] 53.4 ± 9.9 module of HyperMesh software is used to rebuild an existing The whole lower Kajzer et al. [38] 51 ± 15 50th percentile male adult FE model to obtain the lower extremities extremity models in the full spectrum of ages, as shown in Figure 3. 3D adjust domain was established in 5 cross- sections in long bone meshes firstly, and the control point correlation between age and elastic modulus and ultimate handle of the experimental point was setup in a correspond- stress is developed as ing position based on previous geometric study results. Then Elastic modulus = 473 3+14 99 age − 0 19 age , the location of the control point could be adjusted manually to complete the update of lone bone meshes to get the age- 2 Ultimate stress = 8 94 + 0 13 age − 0 002 age specific models. The material properties of the cortical bone are simulated by the material model of isotropic elastoplastic ( MAT_PIE- 2.2. Material Property Changes with Aging CEWISE_LINEAR_PLASTICITY); as the slope of its stress- strain curve in the plastic stage remains unchanged [24], it 2.2.1. Long Bone. The cancellous bone of the lower extremity can be assumed that the tangent modulus in the plastic stage is a kind of porous structure composed of irregularly will not change with aging. If the strain of meshes rose to the arranged trabecular bones; its mechanical properties are sim- failure strain, the fracture would occur and be simulated by ilar to foamed aluminum. When compressed, there is a sig- mesh deletion. The Cowper-Symonds method is used to sim- nificant elastic phase and the stress is nearly unchanged ulate the effect of strain rate on the material properties, with after the yield point; the limit stress is almost equal to the yield stress scaling equation shown in the yield stress. The material properties of the cancellous bone showed obvious changes with aging because of the 1/p loss of calcification and fibrosis. Its material properties σε, ε = σ ε 1+ , 5 can be simulated by the material model of dynamic elasto- C plastic ( MAT_PLASTIC_KINEMATIC), and the ultimate stress is set to 13.4% according to the research results of where σ ε is the initial yield stress, while ε is the strain rate. literature [23]. C and P are the strain rate transformation parameters. In this The quasistatic compression test data of the cancellous paper, C is 360.5 and P is 3.6. The elastic modulus, ultimate stress, and failure strain of bone from the ages of 16 to 83 years [23] were subjected to quadratic polynomial fitting. The results showed that the the cortical bone of different age groups are obtained by mechanical properties of cancellous bone increased from 20 regression analysis of corresponding test data in literatures to 40 years, while there is a sudden drop after 40 years. The [25, 26] and [27]. The correlation between age and elastic Applied Bionics and Biomechanics 5 (a) (b) Figure 6: Thigh and calf three-point bending test model: (a) thigh and (b) calf. Cylindrical bar Metal cylinder Rotary bearing Knee joint Fixed bearing Sliding bearing Figure 7: Four-point bending test device and finite element model. 400 N 400 N Fixed point Fixed point Fixed point 40 km/h 6.25 kg Marker location 40 km/h Moving ground Fixed ground 6.25 kg (a) (b) Figure 8: Lower extremity bending (a) and shear (b) simulation model. modulus, ultimate stress, and failure strain is developed as bones, as they undergo the same changes with age. The presented in (6), (7), and (8), respectively. material property parameters of the cortical bone and can- cellous bone of the long bones for different ages can be Elastic modulus = 18 01 − 0 059 age, 6 calculated according to the above fitting formulas and corre- sponding scaling coefficients. Table 2 shows the lower Ultimate stress = 130 8 − 0 52 age, extremity long bone material properties of ages 30 years and 70 years for example. Failure strain = 4 23 − 0 033 age Based on the research results of the literature [28, 29], 2.2.2. Ligaments. The ligaments in the knee are connected the change of material property of the cancellous bone and to the bones, which stabilize and restrict the movement of cortical bone is assumed to be the same for lower limb long the knee, including the patellar ligament, meniscofemoral 6 Applied Bionics and Biomechanics 0 10 20 30 40 50 60 0 10 20 30 40 50 60 Displacement (mm) Displacement (mm) Test corridor Elder Test corridor Elder Average Average Younger Younger (a) (b) Figure 9: Force displacement curves of the impactor in three-point bending simulation of the (a) thigh and (b) calf. ligament, medial collateral ligament (MCL), lateral collat- eral ligament (LCL), anterior cruciate ligament (ACL), and posterior cruciate ligament (PCL). The diameter of collagen fibers decreased, while the fiber content increased with aging. For example, the maximum fiber diameter is 180 nm when a man is 15–19 YO and reduced to 110 nm after 60 YO [30, 31]. The change of collagen fibrils will affect the Fracture of femur mechanical properties of the ligament; the ultimate tension, especially, will decline with aging [31]. Figure 10: Femur fracture location in thigh three-point bending In the present study, the ligaments are simulated by the simulation. solid elements to accurately model the geometrical shape of each ligament and their contact with the surrounding tissue. using quadrilateral shell elements. Muscle and skin are mod- The hyperelastic material constitutive model ( MAT_- eled using the solid element and shell element, respectively. SOFT_TISSUE) is used to simulate its mechanical proper- Ligaments are modeled using the solid element and one- ties [32], and the failure of the first-order principal strain dimensional beam element together. The baseline model is is defined for the elements, with the laceration of ligament adjusted according to the pedestrian’s standing posture. simulated by element deletion. Then the previous research results of the geometric changes The experiments of the knee joint ligament carried out by and material property changes with aging are applied to build Woo et al. [33] were simulated; the simulation model is the age-specific lower extremity FE models—including the shown in Figure 4. The ligament properties for different ages adjustment of the material properties and the geometry can be obtained by parameter computational inverse based morphing of the femur, tibia, and fibula, as shown in on the ligament tensile test curves. Figure 5. Then two FE models of the pedestrian lower Table 3 shows the knee joint ligament material properties extremity of typical ages 30 and 70 years are established to of individuals aging 30 to 70 years; for example, C1, C3, C4, investigate the effect of age on injury risk. The selection of and C5 are the parameters of the material model. 30- and 70-YO models was based on a previous recommen- dation that defined a young adult group between 16 and 35 2.3. Development of the Age-Specific Lower Extremity FE YO and elderly group as 66 YO and older [34] . Model. The baseline pedestrian lower extremity model is derived from the Global Human Body Models Consortium 2.4. Model Validation. A series of cadaver test data were used (GHBMC) average male occupant model. The GHBMC is to validate the biofidelity and stability of age-specific pedes- representative of a 50th percentile male adult and was based trian lower extremity FE models, as shown in Table 4. These on medical images of a 26 YO individual. The lower extrem- validation tests were simulated in LS-DYNA software ity model includes the long bone, muscle, ligament, skin, and according to the published test information. other tissues. The cortical bone and cancellous bone of the long bone shaft are modeled using hexahedral elements. 2.4.1. Validation at the Component Level. In Kerrigan’s test The cortical bone covering the long bone ends is modeled [35], the thigh and calf were extracted from PMHS. The Impact force (kN) Impact force (kN) Applied Bionics and Biomechanics 7 Fracture of the fibula at both ends Fracture of the tibia and fibula Fracture of the tibia and fibula (a) (b) Figure 11: Tibia and fibula fracture location in calf three-point bending simulation: (a) the young and (b) the elderly. 2.0 muscle tissues of two ends were removed, and the distal and proximal ends of the femur and tibia were potted in cups and fixed with polyurethane. An impactor driven by a universe machine loaded the thigh and calf at the middle-shaft loca- 1.5 tion at the speed of 1.5 m/s to simulate the loading condition of pedestrian lower extremities in a vehicle-pedestrian colli- sion accident. Then two age-specific FE models of the pedes- 1.0 trian lower extremities of typical ages 30 and 70 YO were used to simulate the same tests with the same experiment set- tings and boundaries. The finite element models are shown in 0.5 Figure 6. Ligament failure caused by lateral bending is a common knee injury for pedestrian during vehicle-pedestrian collision accident. Kerrigan et al. [35] and Bose et al. [37] designed a 0.0 dynamic four-point bending test to estimate knee tolerance. 0 2 4 6 8 101214 Figure 7 illustrates the test principle: isolated knee parts Displacement (mm) were potted in specific cups that rotated around support Simulation age = 30 years Test age = 22–35 years joints during tests. While the distal support connected to Simulation age = 70 years Test age = 60–97 years the tibia was fixed, the proximal support connected to the femur was allowed to move horizontally. The angular speed Figure 12: Ligament ACL displacement force curve comparison of the knee was about 1 deg/ms during tests to simulate the between experiment and simulation. knee-bending load when pedestrian crashed at a speed of 40 Km/h, and the bending moment was measured by a load cell 3. Results and Discussion connected to the femur extension bar. Corresponding simu- lation models were built to perform the same tests in ls- The force displacement curves of the impactor in thigh and dyna, and then the simulation results were compared with calf three-point bending simulations are shown in Figure 9. the test data. Both the simulation results of young (30 YO) and elderly (70 YO) are in the test corridor and consistent with the experimen- 2.4.2. Validation at the Lower Extremity Level. To evaluate tal results, though different from each other. This indicates the whole lower extremity response, 2 loading cases, bending that the response of the young and the elderly is much differ- and shear, were simulated to assess the importance of geo- ent. The impactor force rises slowly initially but is followed metric and material property changes with aging. by a sharp increase. This is because the impactor makes con- According to the tests of Kajzer et al. [38], as shown in tact with the skin, muscle, and other soft tissues first, and Figure 8, the lower extremity was extracted from PMHS on when the femur begins to bend to deform, the force increases. the hip joint and fixed flat on a board to maintain stability. In thigh three-point bending simulation, the femoral The proximal of the femur was fixed with screws, while the fracture occurred in both cases of the young (30 YO) and distal of the femur was fixed with a fixed plate to limit its hor- the elderly (70 YO), as shown in Figure 10. For the elderly, izontal movement. The force sensor would calculate the the femur fracture occurred when the displacement of the bending moment of the knee joint. A force of 400 N was impactor is 40 mm with the impact force 3.2 kN, while the loaded at the hip to simulate the load received by the lower corresponding data of the young is 50 mm and 6.1 kN. extremities when standing. The bending and shear impact In calf bending simulation, both the tibia and fibula are load was conducted at 40 km/h with a 6.25 kg I-shaped impac- fractured, while the fracture locations are different, as shown tor striking the ankle joint and the knee joint, respectively. in Figure 11. The elderly’s fibula was fractured at both ends The impactor was wrapped with a foam of 100 mm × and the middle shaft, but the young’s fibula was only frac- 120 mm × 50 mm at the front. tured in the middle. For the elderly, the fibula fracture Force (kN) 8 Applied Bionics and Biomechanics Rupture of MCL 0 ms 28 ms 40 ms Figure 13: Knee four-point bending simulation process. occurred when the displacement of the impactor is 29 mm, 300 while the young is 36 mm, and the curve showed an obvious decline when the fibula fractured. After the fibula fracture, the impactor continued to load on the tibia, and both the young and the elderly suffered tibia fracture subsequently when the displacement was 50 mm and 42 mm, respectively. The comparison among the ligament displacement force is shown in Figure 12 in terms of results from the simulation of the models and experimental data of Woo et al. [33] in the tensile tests of the femur-ACL-tibia complex. Both the simulation curves for the young and the elderly are well aligned with the experimental results. It is believed that the material property parameters of the ligaments for different ages are reasonable and can reflect the ligament injury at different ages. 0 5 10152025 Figure 13 shows the results of a knee joint four-point Bending angle (°) bending simulation of the young. The medial collateral liga- ment (MCL) is completely ruptured at 28 ms near the tibia Test corridor (Nm) Younger (Nm) junction, which coincides with the test results performed by Elder (Nm) Kerrigan et al. [35] and Bose et al. [37]. The bending-angle-to-bending-moment curves of the Figure 14: Curve of bending angle and bending moment of knees in knee joint are shown in Figure 14. four-point bending simulation. The simulation results of the elderly are in the test corri- dor, while the peak of the young is outside the corridor. This may be due to the ages of the PMHS, as they are between 44 of the elderly is obviously bigger than the young beyond YO and 80 YO—therefore, it is reasonable that the peak of 10 ms, and reached 5 at 20 ms, the similar trend occurred the young (30 YO) is outside the corridor. The simulation in knee joint shear displacement curves. This is because the curves coincide with the test curves before MCL rupture, knee ligament strength of the elderly is much lower than which indicates that the material properties of the ligaments the young and their ligaments usually rupture earlier in the are reasonable. At the beginning, the bending moment same collision condition. For example, the MCL and PCL increases with the bending angle and reaches the maximum ruptured at 11.5 ms and 14.5 m, respectively, in the simula- value when the MCL is about to rupture and then the bend- tion. It was significantly ahead of the results of the young, ing moment decreases sharply. The maximum bending which is 18.5 ms and 20 ms, respectively, and then induced moment of the elderly is about 110 Nm with a bending angle larger knee bending angle and shear displacement. The of 11 kinematics of lower extremity and ligament rupture in a , while the maximum bending moment of the young is about 270 Nm with a bending angle of 18 , far greater than bending test simulation is shown in Figure 16 (take the the elderly. young for example). The time history curves of impactor force, knee joint The time history curve of the impact force and knee joint bending angle, and knee joint shear displacement in lower shear displacement in the lower extremity shear simulation extremity bending simulation are shown in Figure 15. It indi- are shown in Figure 17. It indicates that the simulation results cates that the simulation results of the young and the elderly of the young and the elder show a similar linear shape as that show a linear shape similar to that reported in tests and all are reported in tests and are mostly in the test corridor, except in the test corridor. for the impact force of the young. For the impact force, there is not much difference The peak impact force of the elderly is 5.2 kN, lower than 6.0 kN of the young, and appeared earlier at about 8 ms, while between the simulation results of the young and the elderly. They both reach their maximum value of 4.5 kN at 4 ms. This the knee joint shear displacement of the elderly is obviously may be due to the same kinetic energy of the impactor and larger than that of the young beyond 8 ms. It is possibly the quality of the lower extremity. While the bending angle related to the elderly’s femur fracture occurring at around Bending moment (Nm) Applied Bionics and Biomechanics 9 0 0 0 5 10 15 20 25 30 0 5 10 15 20 25 Time (ms) Time (ms) Test corridor Test corridor Younger Younger Elder Elder (a) (b) −10 0 5 10 15 20 25 30 Time (ms) Test corridor Younger Elder (c) Figure 15: Lower extremity bending simulation results: (a) impact force, (b) knee joint bending angle, and (c) knee joint shear displacement. 8 ms, resulting in the increase of rotation and lateral move- techniques. To evaluate the lower extremity response, a series ment at the fracture point and the decrease of the impact of PMHS tests were simulated to validate the confidence of the force. The detailed injuries of the elderly and the young are models and to assess the importance of geometric and mate- compared in Figure 18. The young only suffered partial rial property changes with aging. The whole age-specificFE femur fracture, while the elder suffered full femur fracture models of pedestrian lower extremity showed numerical sta- and fibula fracture. These injuries were coordinated with that bility, and, in all validation simulations, the response of the of samples 4 and 17 in the PMHS test [38]. young model and the elderly is different from each other. Development of age-specific FE models of the lower extremity will provide valuable tools for understanding variations in 4. Conclusion lower extremity injury patterns due to vehicle-pedestrian col- In the present study, the changes of geometric and mate- lision accidents across populations and in the design of new vehicles with devices for pedestrian protection. rial properties of the lower extremity with aging were studied and age-specific FE models of the lower extremity Further study will involve the sex factor and the geometry for pedestrian-vehicle accident simulation were developed changes of the femoral head/neck and ankle with age. These for 30-YO and 70-YO male pedestrian using morphing would be investigated to establish a pedestrian lower limb Impact force (N) Shear displacement (mm) Bending angle (N) 10 Applied Bionics and Biomechanics Rupture of PCL Rupture of MCL 0 ms 6 ms 12 ms 18.5 ms 20 ms Figure 16: Dynamic simulation of the lower extremity simulation process. 7000 100 0 0 0 5 10 15 20 25 30 010 20 30 40 Time (ms) Time (ms) Test corridor Test corridor Younger Younger Elder Elder (a) (b) Figure 17: Lower extremity shear simulation results: (a) impact force and (b) knee joint shear displacement. Partial fracture of the femur Fracture of the femur Fracture of the fibula (a) (b) Figure 18: Comparison of long bone injury: (a) the young and (b) the elderly. Impact force (N) Shear displacement (mm) Applied Bionics and Biomechanics 11 model with higher biofidelity. 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