Biomechanical Analyses of Head Acceleration during Contact Sports and Ankle Joint Complex Injury

Abstract

Athletes are frequently exposed to conditions that can result in bodily injuries, such as concussions and ankle sprains. This research is divided into two separate biomechanical studies: the design and evaluation of a device to simulate football helmet collisions and its parameters, and a cadaveric study to investigate the effects of ankle sprains. An impactor system was designed and built along with a computational model to simulate football helmet collisions, test helmet designs, and evaluate the influence of different parameters on head acceleration. Peak head accelerations of 20g to 60g, and rise times were targeted based on measures reported in the literature. A computational model of the impactor system was developed in Adams to determine design parameters such as neck stiffness and dampening. A pendulum impactor was constructed to achieve various impact energies by changing the release angle of the arm reaching up to 110 J. A dummy’s neck was designed as a single degree-of-freedom hinge joint with variable stiffness. Peak head accelerations agreed within 6% of literature reported accelerations. Similar to previous studies, the head reaches its peak acceleration in 10 to 12 ms. Neck stiffness did not affect the head peak acceleration during the 20 ms following impact. Moreover, the computational model revealed that adding a dampening of 1.75 N.s/mm to the neck results in 20 g decrease in the head peak acceleration. A second study aimed both to characterize the effects of collateral ankle ligament injuries using nine cadaveric ankle joints and to quantify the contribution of lower leg muscle forces to the ankle joint kinematics. Intact ankles were tested, then anterior talofibular (ATFL) and calcaneofibular (CFL) ligaments were sequentially resected to simulate two grades of ankle injury. The tibialis anterior (TA) and extensor digitorum longus (EDL) tendons were loaded with static weights and distributed based on their physiological cross sectional area. Weak TA and Weak EDL configuration were simulated by reducing their respective muscle loads by 50%. The ankle was moved from full dorsiflexion to plantarflexion using a stepper motor attached to Achilles tendon. The effect of muscle configuration (Weak TA and Weak EDL) and injuries (ΔATFL and ΔATFL-CFL) compared to baseline (intact ankle with physiological load set) on the measured ankle joint kinematics was determined using a repeated measures ANOVA. The weaker EDL demonstrated 1° to 3° higher inversion than physiological loads through the cycle significantly. After the simulated ATFL-CFL injury, the trials demonstrated a 2° increase in inversion with respect to intact ankle kinematics in dorsiflexion. The required Achilles’ load increased by up to 24% in plantarflexion after the injury indicating a significant reduction in efficiency. Based on the higher inversion results in weaker muscle set, strengthening the EDL might help to prevent the hyper inversion injuries.