Abstract
Previous biomechanical studies of lateral collateral ligament (LCL) injuries and their surgical repair, reconstruction and rehabilitation have primarily relied on gravity effects with the arm in the varus position. The application of torsional moments to the forearm manually in the laboratory is not reproducible, hence studies to date likely do not represent forces encountered clinically.
The aim of this investigation was to develop a new biomechanical testing model to quantify posterolateral stability of the elbow using an in vitro elbow motion simulator.
Six cadaveric upper extremities were mounted in an elbow motion simulator in the varus position. A threaded screw was then inserted on the dorsal aspect of the proximal ulna and a weight hanger was used to suspend 400g, 600g, and 800g of weight from the screw head to allow torsional moments to be applied to the ulna. An LCL injured (LCLI) model was created by sectioning of the common extensor origin, and the LCL. Ulnohumeral rotation was recorded using an electromagnetic tracking system during simulated active and passive elbow flexion with the forearm pronated and supinated. A repeated measures analysis of variance was performed to compare elbow states (intact, LCLI, and LCLI with 400g, 600g, and 800g of weight).
During active motion, there was a significant difference between different elbow states (P=.001 pronation, P=.0001 supination). Post hoc analysis showed that the addition of weights did not significantly increase the external rotation (ER) of the ulnohumeral articulation (10°±7°, P=.268 400g, 10.5°±7.1°, P=.156 600g, 11°±7.2°, P=.111 800g) compared to the LCLI state (8.4°±6.4°) with the forearm pronated. However, with the forearm supinated, the addition of 800g of weight significantly increased the ER (9.2°±5.9°, P=.038) compared to the LCLI state (5.9°±5.5°) and the addition of 400g and 600g of weights approached significance (8.2°±5.7°, P=.083 400g, 8.7°±5.9°, P=.054 600g).
During passive motion, there was a significant difference between different elbow states (P=.0001 pronation, P=.0001 supination). Post hoc analysis showed that the addition of 600g and 800g but not 400g resulted in a significant increase in ER of the ulnohumeral articulation (9.3°±7.8°, P=.103 400g, 11.2°±6.2°, P=.004 600g, 12.7°±6.8°, P=.006 800g) compared to the LCLI state (3.7°±5.4°) with the forearm pronated. With the forearm supinated, the addition of 400g, 600g, and 800g significantly increased the ER (11.7°±6.7°, P=.031 400g, 13.5°±6.8°, P=.019 600g, 14.9°±6.9°, P=.024 800g) compared to the LCLI state (4.3°±6.6°).
This investigation confirms a novel biomechanical testing model for studying PLRI. Moreover, it demonstrates that the application of even small amounts of torsional moment on the forearm with the arm in the varus position exacerbates the rotational instability seen with the LCL deficient elbow. The effect of torsional loading was significantly worse with the forearm supinated and during passive elbow motion. This new model allows for a more provocative testing of elbow stability after LCL repair or reconstruction. Furthermore, this model will allow for smaller sample sizes to be used while still demonstrating clinically significant differences. Future biomechanical studies evaluating LCL injuries and their repair and rehabilitation should consider using this testing protocol.
