Development and Validation of an Active Muscle Simplified Finite Element Human Body Model in a Standing Posture

Validation and application of a finite element model simulating failure thresholds of skin during blunt puncture with varying impactor geometries

Injuries caused by knives or other sharp tools such as scissors and screwdrivers are common in violent crimes and self-defense acts. The force thresholds of skin have been quantified based on the puncture instrument to assess degree of force in forensic cases, but limited studies have investigated blunt instruments and the effect of skin thickness. A finite element (FE) computational model was developed to simulate blunt puncture of skin. Curve fitting and manual optimization were performed to obtain Ogden material coefficients. The model was validated with experimental force-time curves for spherical impactors of diameter 3, 5, and 8 mm into thin, average, and thick skin at slow and fast loading rates (n = 18 total conditions), resulting in an average CORA score of 0.725. The average maximum principal stress at the time of experimental failure was 57.3 MPa with a coefficient of variance of 0.18, and the median value of 54.8 MPa was selected as the failure criterion. The validated model was applied to load seven spherical impactors, five Hex screwdrivers, and three Torx screwdrivers into skin with thicknesses ranging from 2 to 3 mm. Increased skin thickness resulted in greater force, displacement, and strain energy at failure. Cross-sectional area of the impactor and failure thresholds of skin expressed a linear relationship for normalized force (R2 ≥ 0.88), displacement (R2 ≥ 0.77), and normalized strain energy (R2 ≥ 0.92). The validated FE model may be used to determine the force required to penetrate skin with a case-specific blunt instrument.

Effect of individual spinal muscle activities on upright posture using a human body finite element model

The occurrence of diseases characterized by irregular spinal alignment, such as kyphosis, lordosis, scoliosis, and dropped head syndrome (DHS) is increasing, particularly among older adults. DHS is characterized by an excessive forward tilt of the head and neck, causing the head to droop. Although it is believed that muscle activity plays a role in both the onset and treatment of DHS, the underlying mechanisms remain unclear. To elucidate the mechanism, we used a human body finite element model, which included the erector spinae muscle group, and a muscle controller with fixed legs for spinal posture stabilization. The model replicated muscle activation levels during the maintenance of an upright posture under gravity, similar to those obtained from experimental data. Parametric simulations to investigate the effect of each spinal muscle impairment on upright posture with and without compensatory activities of the other muscles suggest that trunk extensors; the multifidus L1-S and longissimus thoracis muscles, and hip flexors; psoas major and iliacus muscles play an integral role in maintaining an upright posture. These findings support the results of a rehabilitation study that reported that exercises targeting the trunk, psoas muscles, and cervical extensors could improve global spinal alignment and clinical outcomes in DHS.

Injury Risk Predictions in Lunar Terrain Vehicle (LTV) Extravehicular Activities (EVAs): A Pilot Study

Extravehicular activities will play a crucial role in lunar exploration on upcoming Artemis missions and may involve astronauts operating a lunar terrain vehicle (LTV) in a standing posture. This study assessed kinematic response and injury risks using an active muscle human body model (HBM) restrained in an upright posture on the LTV by simulating dynamic acceleration pulses related to lunar surface irregularities. Linear accelerations and rotational displacements of 5 lunar obstacles (3 craters; 2 rocks) over 5 slope inclinations were applied across 25 simulations. All body injury metrics were below NASA's injury tolerance limits, but compressive forces were highest in the lumbar (250-550N lumbar, tolerance: 5300N) and lower extremity (190-700N tibia, tolerance: 1350N) regions. There was a strong association between the magnitudes of body injury metrics and LTV resultant linear acceleration (ρ = 0.70-0.81). There was substantial upper body motion, with maximum forward excursion reaching 375 mm for the head and 260 mm for the chest. Our findings suggest driving a lunar rover in an upright posture for these scenarios is a low severity impact presenting low body injury risks. Injury metrics increased along the load path, from the lower body (highest metrics) to the upper body (lowest metrics). While upper body injury metrics were low, increased body motion could potentially pose a risk of injury from flail and occupant interaction with the surrounding vehicle, suit, and restraint hardware.

Effect of Active Muscles on Astronaut Kinematics and Injury Risk for Piloted Lunar Landing and Launch While Standing

While astronauts may pilot future lunar landers in a standing posture, the response of the human body under lunar launch and landing-related dynamic loading conditions is not well understood. It is important to consider the effects of active muscles under these loading conditions as muscles stabilize posture while standing. In the present study, astronaut response for a piloted lunar mission in a standing posture was simulated using an active human body model (HBM) with a closed-loop joint-angle based proportional integral derivative controller muscle activation strategy and compared with a passive HBM to understand the effects of active muscles on astronaut body kinematics and injury risk. While head, neck, and lumbar spine injury risk were relatively unaffected by active muscles, the lower extremity injury risk and the head and arm kinematics were significantly changed. Active muscle prevented knee-buckling and spinal slouching and lowered tibia injury risk in the active vs. passive model (revised tibia index: 0.02-0.40 vs. 0.01-0.58; acceptable tolerance: 0.43). Head displacement was higher in the active vs. passive model (11.6 vs. 9.0 cm forward, 6.3 vs. 7.0 cm backward, 7.9 vs. 7.3 cm downward, 3.7 vs. 2.4 cm lateral). Lower arm movement was seen with the active vs. passive model (23 vs. 35 cm backward, 12 vs. 20 cm downward). Overall simulations suggest that the passive model may overpredict injury risk in astronauts for spaceflight loading conditions, which can be improved using the model with active musculature.