Summary Statement

Spinal flexibility in bending and axial torque has been shown to exhibit very modest changes with advancing disc degeneration. This study is the first to address the possible relationship in pure anterior shear and no clear relationship was observed.

Purpose: A long spinal fusion across the thoracolumbar region is sometimes applied in scoliosis. Adjacent level degeneration below these constructs has been documented. Treatment with an artificial disc replacement below the fusion has been proposed to prevent degeneration there. There is currently little data detailing the expected biomechanics of this situation. The objective of this study was to evaluate range of motion (ROM) and helical axis of motion (HAM) changes due to one- and two-level Maverick total disc replacement adjacent to a long spinal fusion.

Method: A multidirectional flexibility testing protocol with compressive follower preload was used to test seven human cadaveric spine specimens (T8-S1). A continuous pure moment ±5.0 Nm was applied in flexion-extension (FE), lateral bending (LB) and axial rotation (AR), with a compressive follower preload of 400 N. The motion of each vertebra was monitored with an optoelectronic camera system. The test was completed for the intact condition and after each surgical technique:

T8-L4 fusion and facet capsulotomy at L4–L5 and L5-S1;

Maverick total disc replacement and fusion with the CD Horizon system was performed. Repeated measures ANOVA was used to analyze changes in ROM and HAM of the L4–L5 and L5-S1 segments.

Results: Following L4-L5 Maverick replacement, L4-L5 ROMs tended to decrease slightly (on average from 6.2°±2.8° to 5.1°±3.8° in FE, 1.1°±1.1° to 0.9°±0.5° in LB and 1.3°±0.9° to 1.0°±0.6° in AR). With two-level Maverick implantation, L5-S1 ROMs tended to increase slightly in FE (from 6.6°±2.6° to 7.1°±3.9°), and to decrease slightly in LB (from 1.5°±0.9° to 1.0°±0.3°) and AR (from 1.5°±1.5° to 1.1°±0.6°), compared to the fused condition. As a trend, HAM location shifted posteriorly in FE and AR, and inferiorly in LB following Maverick replacement. However, neither ROM nor HAM at these two segments showed any significant change due to the implantation of one-or two-level Maverick total disc replacement in any of the three directions.

Conclusion: The present results suggested that lower lumbar segments with Maverick disc replacement exhibited intact-like kinematics in both extent and quality of motion.

Purpose: The average age of people suffering spinal cord injuries in many countries is shifting toward an older population, with a disproportionate number occurring in the spondylotic cervical spine. These injuries are typically due to low energy impacts, such as a fall from standing height. Since a stenotic spinal canal (a common feature of a spondylotic cervical spine) can cause myelopathy when the spine is flexed or extended, traumatic flexion or extension likely causes the injury during the low energy impact. However, this injury mechanism has not been observed experimentally.

Method: To better understand this injury mechanism an in-vitro study, using six whole cervical porcine spines, was conducted. The following techniques were combined to directly observe spinal cord compression in a stenotic spine during physiologic and super-physiologic motion:

A radio-opaque surrogate cord, with material properties matched to in-vivo specimens, replaced the real spinal cord.

Results: Initial results show that a stenotic occlusion that removes all extra space in the canal in the neutral posture, without compressing the cord, can lead to spinal cord compression within physiologic ranges of flexion and extension. The spinal cord can also be compressed during slightly super-physiologic flexion and extension with only 25% canal occlusion. Physiologic loads and motions in the same spines did not cause cord compression when canal occlusion was 0%.

Conclusion: These results support the hypothesis that cervical spinal canal stenosis increases the risk of spinal cord injury because spinal cord compression was observed during motions and loads that would be safe for a non-stenotic spine. These results are limited primarily due to the use of a porcine spine. However, this new stenosis model and experimental technique will be applied to in-vitro human spine specimens in future work.

Evidence suggests that femoral neck fractures initiate in the superolateral cortex, where it is significantly thinner in older than younger individuals (Mayhew, et al. Lancet 2005). Thus, we sought to determine the relative time-course of crack initiation and propagation during a simulated hip fracture.

Four unembalmed frozen, human cadaveric specimens (mean age = 78 yrs) were loaded to failure in sideways fall configuration at a rate of 100 mm/sec using a materials testing system. Images of the fracture were captured with two high-speed video cameras at a resolution of 384x384 pixels, and sample rate of 9,111 Hz (frames/second).

Test A: The load-displacement (L-D) curve had three distinct peaks: at the first peak (4390 N), the head and neck rotated slightly. At the second peak (4607 N), a visible local compressive fracture appeared in the superior cortex of the proximal neck. At the third peak (3582 N), a neck-spanning tensile failure occurred in the inferior neck. Test B: At the first and second peak loads (1714 N and 3040 N) fluid was released from the posterior then superior and inferior surfaces. The third peak load (3361 N) corresponded to a local compressive failure in the lateral superior neck, followed by a neck-spanning tensile failure medially. Test C: The L-D curve was linear until ultimate load (3038 N). A compressive crack first appeared on the anterior-superior surface of the neck cortex, then fractured in the inferior neck. Test D: The L-D curve was linear until ultimate load. A small local crack appeared in the superior cortex of the proximal neck at ultimate load (3841 N).

We found that during ex vivo simulations of hip fracture, the femur failed initially in the superior cortex of the neck, and then failed in the inferior cortex. This is the first study to demonstrate, with high speed video data, the location of crack initiation and its propagation. These preliminary data support the hypothesis of Mayhew et al. (Mayhew, et al. Lancet 2005) in terms of fracture development and could relate to clinically relevant fracture types.

Information regarding the axes of motion or centers of rotation of the normal cervical spine are necessary to evaluate the similarity of the motion allowed by cervical total disc replacement designs to the natural cervical spine. However, little data has been presented previously regarding the three-dimensional axes of motion of the cervical spine for the three primary motions of flexion/extension, lateral bending and axial rotation. The objective of this study was to measure the three-dimensional axes of motion (Helical axis of Motion) in the natural sub-axial cervical spine using ex-vivo human cadaveric cervical spines.

To measure the Helical Axes of Motion (HAM) for the sub-axial cervical spine under flexion/extension, lateral bending and axial torsion moments and evaluate the effect of a physiologic axial preload on the axes locations and orientations.

This study demonstrated the feasibility of calculating the HAM in the cervical spine using an ex-vivo experimental protocol.

The HAM is a three-dimensional analogue to the two-dimensional center of rotation. The data presented here can be used to evaluate the similarity of the motion allowed by total disc replacement designs to the natural cervical spine. They can also be applied for the characterization of spinal trauma, pathology, instability or surgical devices.

The orientation and locations of the HAMs for axial torsion loading are presented in Figure 1. In flexion/extension the HAM penetrated the sagittal plane near the posterior aspect of the vertebral body and near the cranial endplate. The lateral bending results were similar to the axial torsion results. The addition of axial preload had little effect on the position and orientation of the HAM.

Sub-axial (level C2-C7) cadaveric cervical spine functional spinal units (n=7) were subjected to pure moments of 1 Nm. Specimens were tested with and without axial preloads of 200 N. Vertebral kinematics were measured using an optoelectronic motion analysis system. These data are particularly applicable to the evaluation and design of “motion-retaining” devices such as total disc replacements, facet joint replacement systems or flexible stabilization systems.

Please contact author for figures and diagrams.

We performed a biomechanical study on human cadaver spines to determine the effect of three different interbody cage designs, with and without posterior instrumentation, on the three-dimensional flexibility of the spine. Six lumbar functional spinal units for each cage type were subjected to multidirectional flexibility testing in four different configurations: intact, with interbody cages from a posterior approach, with additional posterior instrumentation, and with cross-bracing. The tests involved the application of flexion and extension, bilateral axial rotation and bilateral lateral bending pure moments. The relative movements between the vertebrae were recorded by an optoelectronic camera system.

We found no significant difference in the stabilising potential of the three cage designs. The cages used alone significantly decreased the intervertebral movement in flexion and lateral bending, but no stabilisation was achieved in either extension or axial rotation. For all types of cage, the greatest stabilisation in flexion and extension and lateral bending was achieved by the addition of posterior transpedicular instrumentation. The addition of cross-bracing to the posterior instrumentation had a stabilising effect on axial rotation. The bone density of the adjacent vertebral bodies was a significant factor for stabilisation in flexion and extension and in lateral bending.