A Biomechanical Evaluation of Lumbar Facet Replacement Systems

Transcription

1 A Thesis Entitled A Biomechanical Evaluation of Lumbar Facet Replacement Systems By Miranda N. Shaw Submitted as partial fulfillment of the requirements for the Master of Science in Bioengineering Adviser: Vijay K. Goel, Ph.D. Graduate School The University of Toledo August 2005

2 The University of Toledo College of Engineering I HEREBY RECOMMEND THAT THE THESIS PREPARED UNDER MY SUPERVISION BY: Miranda N. Shaw ENTITLED: A Biomechanical Evaluation of Lumbar Facet Replacement Systems BE ACCEPTED IN PARTIAL FULFILLMENT OF THE REQUIREMENTS FOR THE DEGREE OF: Master of Science in Bioengineering Thesis Advisor: Vijay K. Goel, Ph.D. Recommendation concurred by: Ozan Akkus, Ph.D. Ashok Biyani, M.D. Committee On Final Examination Mohamed Samir Hefzy, Ph.D. Dean, College of Engineering ii

3 Acknowledgment Two years of hardships, research, and wonderful learning experiences have come to conclusion with the writing of this thesis. There are several individuals that have helped pave the path for a successful completion that I would like to extend my gratitude to. A wonderful leader, researcher, teacher, and person, it is my pleasure to express my extreme appreciation to my advisor, Dr. Vijay K. Goel. He has supported this work in incredible ways and I will always consider him part of the foundation of my future success. I would also like to thank my committee members for offering learning and research assistance throughout the process. I extend a special thanks to Dr. Koichi Sairyo for his kindness, helpfulness, and superb learning experiences. Special thanks go to those who have kept me sane during the long nights and rough times. Srilakshmi Vishnubhotla has been through it all sitting beside me and will always share great memories. My colleagues for creating a wonderful learning environment, my family for the unending support, and the bioengineering department for the extended help, I thank you all. When it was least expected and most needed, a friendly shoulder was always available from Michael Metzger; thank you. Lastly, I would like to thank Medicine Lodge, Inc. and Facet Solutions, Inc. for partially funding the cadaveric research. iii

4 Abstract of A Biomechanical Evaluation of Lumbar Facet Replacement Systems Miranda N. Shaw Submitted as partial fulfillment of the requirements for the Master of Science in Bioengineering The University of Toledo August 2005 A three-dimensional, nonlinear, experimentally validated L3-S1 ligamentous lumbar finite element model was used to determine the effectiveness of artificial facets on the stability of the lumbar spine. Several models were generated to evaluate different artificial facet designs across L4-L5 including capping with and without screws, pedicle screw based designs with sliding articulating surfaces, and a modified universal joint. The pedicle screw based designs were also placed in a destabilized with wide laminectomy model. For comparison purposes, a rigid screw and rod model and wide laminectomy at L4-L5 model were also created and analyzed. The model predictions were compared with a cadaveric study in which a pedicle screw based artificial facet with sliding articulating surfaces was implanted at L4-L5. All FE models and cadaver specimens were evaluated in extension, flexion, bending, and rotation without follower load and 6 N-m applied moment, and extension and flexion with 400N follower load and iv

5 6 N-m moment for angular motions. Facet loads and stresses were also determined for all FE models. Results show that motion increases with artificial facets as compared to intact; dramatically in flexion with a wide laminectomy. FE model results are in agreement with cadaveric results, further validating the FE models. Facet loads generally decreased for most levels in all loading modes compared to intact. Implant stresses were less than the yield stress of titanium and pedicle screw stresses were comparable to screw stresses experienced in fusion cases. Among the designs modeled, pedicle screw based artificial facets performed optimally. v

6 Table of Contents Acknowledgement...iii Abstract... iv Table of Contents... vi List of Figures and Tables... x Chapter I Introduction... 1 Chapter Overview... 1 Significance of Back Pain... 1 Lumbar Spine Anatomy... 1 LBP and Treatments... 4 Scope of Study... 6 Chapter II Literature Review... 7 Chapter Overview... 7 Aging Spine... 7 Spinal Disorders... 8 Osteoporosis... 8 Osteophytes... 8 Disc Degeneration... 8 Facet Joint Osteoarthritis... 9 Facet Hypertrophy... 9 Spinal Stenosis Facet Tropism Treatment Options Decompression Fusion Artificial Discs Artificial Facets and Indications Role of Facet Joint Facet Joint Replacement Systems Facet Capping Cartilage Replacement Articulating Surface Designs vi

7 Posterior Element Replacement Ball & Socket Facet Replacement Systems Summary Chapter III - Materials and Methods Chapter Overview Finite Element Analysis Intact Finite Element Model Geometric Model Bony Element Modeling Apophyseal (Facet) Joint Intervertebral disc Ligaments Material Property Definitions Boundary Conditions and Loading Modes Finite Element Model Validation Artificial Facet Joint Designs Destabilized Spine with Wide Laminectomy Artificial Facet Caps Pedicle Screw Based Artificial Facet Pedicle Screw Based Artificial Facet with Wide Laminectomy Universal Joint Artificial Facet Rigid Screw & Rod System Finite Element Model Data Analysis In Vitro Artificial Facet Joint Study In Vitro Data Analysis Artificial Facet FE Model Validation Chapter IV Results Chapter Overview Cadaveric Study Results Finite Element Model Results Angular Motion Facet Loads Peak von Mises Stresses Chapter V Discussion Chapter Overview Discussion Range of Motion Facet Loads Maximum Implant Stress Conclusions vii

8 Study Limitations Future work References Appendix A Functional Anatomy of the Spine Introduction Vertebrae Posterior Elements Facet Joints Intervertebral Disc Nucleus Pulposus Annulus Fibrosus Ligaments Nerve Pathways Low Back Muscles Appendix B Biomechanics of the Lumbar Spine Axial Compression Flexion Extension Axial Rotation Lateral Bending Range of Motion Shear Appendix C - Finite Element Model Validation Appendix D Modified Universal Joint Artificial Facet Introduction Materials and Methods Discussion Appendix E Facet Replacement System Facet Load Components Introduction Results Appendix F Artificial Facet Disc Stresses Introduction Disc Stress Results Discussion viii

9 Conclusion Appendix G Facet Joint Replacement Patents and Patent Applications Appendix H - Publications by the Author Related to Work ix

10 List of Figures and Tables Figure 1.1: (A) Lateral view of a lumbar motion segment showing anatomical features. (B) Transverse view of a lumbar vertebra...3 Figure 1.2: Ligaments of the lumbar spine...3 Figure 2.1: Superior facet replacement cap showing placement of securing screws in U.S. Pat. No Figure 2.2: Inferior facet replacement cap with the superior screw placement shown from U.S. Pat. No Figure 2.3: Superior and inferior facet caps with pedicle screw support from U.S. Published Patent Application No. 2005/ Figure 2.4: Articular cartilage replacement system from U.S. Pat. No Figure 2.5: An transverse slice of the complete facet joint replacement system depicted in U.S. Pat No Figure 2.6: A posterior view of the plate supporting the inferior facet replacement from U.S. Pat No Figure 2.7: Lateral view of the artificial facet proposed in U.S. Pat. No Figure 2.8: Posterior view of the superior and inferior components attached to the respective vertebra using pedicle screws shown in U.S. Pat. No Figure 2.9: A prosthesis designed to replace multilevel facet joints in U.S. Pat. No Figure 2.10: Posterior view of the device used to replace two levels of facet joints...26 Figure 2.11: Total posterior element replacement prosthesis proposed in U.S. Pat. No x

11 Figure 2.12: (A) Translaminar screw and inferior facet replacement head as depicted by Berry. (B) Superior facet replacement screw is shown describing articulation with the inferior facet replacement where motion is restricted with a tether. (C) A sheath may also be used to restrict motion between the superior and inferior facet replacement...28 Figure 2.13: The inferior and superior facets are replaced with a ball and socket design...28 Figure 2.14: A replacement system for the posterior elements proposed by Berry...29 Figure 2.15: A ball and socket artificial facet joint proposed by Serhan et al...30 Figure 2.16: A sliding articular surface facet replacement design with a novel ligament concept to support the joint in tension...30 Figure 3.1: (A) Intact L3-L5 finite element mesh. (B) Intact L3-S1 finite element mesh...36 Figure 3.2: A midsagittal cross-section of L3-S1 lumbar spine indicating important anatomical features...37 Table 3.1: Material property definitions and element types for the intact L3-S1 finite element spine model. For bilateral structures (facet joints and capsular ligaments), the total number of elements is listed...40 Table 3.2: Comparison of intact L3-L5 and L3-S1 finite element predictions and results from Schultz et al [56]. FE model predictions fall within one standard deviation of in vitro results...41 Figure 3.3: (A) Ligamentous L1-S1 spinal motion segment used to determine bending moments. (B) Comparison of the experimental and FE results are in agreement for flexion and extension. (C) Experimental and FE results are compared in left and right bending. (D) Results from cadaveric studies and FE are compared in left and right rotation...42 Table 3.3: Material properties of titanium used in all artificial facet models...43 Figure 3.4: The intact model was modified with a wide laminectomy at L4-L Figure 3.5: (A) L3-S1 FE model with artificial facet caps at L4-L5. (B) L3-S1 FE model with artificial facet caps secured with screws at L4-L Figure 3.6: The pedicle screw based artificial facet design with a 3mm thick stem connecting the metal facets to the pedicle screw (3mm)...46 Figure 3.7: The second design utilizing pedicle screws with a connecting stem thickness of 5mm (5mm)...47 xi

12 Figure 3.8: The pedicle screw design was modified such that the connecting stems increased in width around the pedicle screws and were tied to the screws and bony pedicles (Support)...47 Figure 3.9: The pedicle screw based artificial facet design with a 3mm thick stem connecting the artificial facets to the pedicle screws in the destabilized model (3mmL)...48 Figure 3.10: The 5mm wide laminectomy pedicle screw based design at L4-L5 (5mmL)...48 Figure 3.11: The wide laminectomy model with the Support pedicle screw based design at L4-L5 (SupportL)...49 Figure 3.12: A universal joint was modified to act as an artificial facet joint...50 Figure 3.13: A rigid screw and rod system across the L4-L5 motion segment...51 Figure 3.14: DXA scan L1-L4 of two specimens used for the cadaveric study...54 Figure 3.15: (A) CT slice of lumbar spine with Beekley CT spots in the transverse processes. (B) CT phantom image of the entire lumbar specimen displaying the CT spots in the L4 and L5 spinous processes...55 Figure 3.16: (A) OptoTrak measuring system. (B) Spine testing setup with LED s and tracking cameras...55 Figure 3.17: Destabilized spine with removed L4-L5 facets circled...56 Figure 3.18: A sketch of an implant placed at L4-L5 similar to the implant used in the cadaveric study...56 Figure 3.19: Specimen displaying the six loading modes, the follower load application, LED s for motion tracking, and the global coordinate system...57 Figure 3.20: Results for the in vitro study and artificial facet FE models are in agreement for general trends in motion at 6 N-m in (A) Flexion/Extension; (B) Bending; (C) Rotation...60 Figure 3.21: Results for the in vitro study and artificial facet FE models are in agreement for general trends in motion with 400N follower load and 6 N-m applied moment in extension...61 Figure 4.1: Results of average relative motion across the L4-L5 motion segment for the in vitro study for intact (Intact), destabilized without facets (Destab), and with artificial facets (Implant) for 0 to 6 N-m applied moment: (A) Extension/Flexion; (B) Left/Right Bending; (C) Left/Right Rotation...66 Figure 4.2: Comparison of relative motion across L4-L5 with and without follower load for 0 to 6 N-m extension/flexion moment...66 xii

13 Figure 4.3: Relative motions at all levels of the lumbar spine for the in vitro study with 6 N-m applied moment: (A) Extension; (B) Flexion; (C) Bending; (D) Rotation...68 Figure 4.4: Relative motions at all levels of the lumbar spine for the in vitro study with 400N follower load and 6 N-m applied moment: (A) Extension; (B) Flexion...69 Table 4.1: Percent changes in relative motion at 6 N-m applied moment for the cadaveric study in each loading mode. A positive percent change indicates an increase in relative motion where as a negative percent change indicates a decrease in relative motion...70 Table 4.2: Percent changes in relative motion for 400N follower load and 6 N-m applied moment for the cadaveric study in extension and flexion. No data was available for the implant undergoing flexion with FL. A positive percent change indicates an increase in relative motion where as a negative percent change indicates a decrease in relative motion...70 Figure 4.5: Relative motions at all levels of the lumbar spine for the FE study with 6 N-m applied moment: (A) Extension; (B) Flexion; (C) Bending; (D) Rotation...77 Figure 4.6: Relative motions at all levels of the lumbar spine for the FE study with 400N follower load and 6 N-m applied moment: (A) Extension; (B) Flexion...78 Table 4.3: Percent changes in relative motion at 6 N-m applied moment for the FE models in each loading mode. A positive percent change indicates an increase in relative motion where as a negative percent change indicates a decrease in relative motion...79 Table 4.4: Percent changes in relative motion with 400N follower load and 6 N-m applied moment for the FE models in each loading mode. A positive percent change indicates an increase in relative motion where as a negative percent change indicates a decrease in relative motion...80 Figure 4.7: Relative motions at all levels of the lumbar spine with wide laminectomy for the FE study with 6 N-m applied moment: (A) Extension; (B) Flexion; (C) Bending; (D) Rotation...82 Figure 4.8: All levels relative motion of the wide laminectomy lumbar spine for the FE study with 400N follower load and 6 N-m applied moment: (A) Extension; (B) Flexion...83 Table 4.5: Percent changes in relative motion at 6 N-m applied moment for the wide laminectomy FE models in each loading mode. A positive percent change xiii

14 indicates an increase in relative motion where as a negative percent change indicates a decrease in relative motion...84 Table 4.6: Percent changes in relative motion with 400N follower load and 6 N-m applied moment for the wide laminectomy FE models in each loading mode. A positive percent change indicates an increase in relative motion where as a negative percent change indicates a decrease in relative motion...85 Figure 4.9: Facet loads at all levels of the lumbar spine for the FE study with 6 N-m applied moment: (A) Extension; (B) Bending; (C) Rotation...90 Figure 4.10: Facet loads at all levels of the lumbar spine for the FE study with 400N follower load and 6 N-m applied moment: (A) Extension; (B) Flexion...91 Table 4.7: Percent changes in total facet loads for 6 N-m applied moment for the artificial facet FE models in each loading mode. A positive percent change indicates an increase in facet loads where as a negative percent change indicates a decrease in facet loads...92 Table 4.8: Percent changes in total facet loads for 400N follower load and 6 N- m applied moment for the artificial facet FE models in each loading mode. A positive percent change indicates an increase in facet loads where as a negative percent change indicates a decrease in facet loads...92 Figure 4.11: Facet loads at all levels of the wide laminectomy lumbar spine for the FE study with 6 N-m applied moment: (A) Extension; (B) Bending; (C) Rotation...94 Figure 4.12: Facet loads at all levels of the wide laminectomy lumbar spine for the FE study with 400N follower load and 6 N-m applied moment: (A) Extension; (B) Flexion...95 Table 4.9: Percent changes in total facet loads for 6 N-m applied moment for the artificial facet FE models with wide laminectomy in each loading mode. A positive percent change indicates an increase in facet loads where as a negative percent change indicates a decrease in facet loads...96 Table 4.10: Percent changes in total facet loads with 400N follower load and 6 N-m applied moment for the artificial facet wide laminectomy FE models in each loading mode. A positive percent change indicates an increase in facet loads where as a negative percent change indicates a decrease in facet loads...96 Figure 4.13: Stress plots for the inferior and superior artificial facet caps at L4- L5 with 6 N-m bending moment in all loading modes Figure 4.14: Stress plots for the inferior and superior artificial facet caps at L4- L5 with 400N follower load and 6 N-m bending moment in all loading modes xiv

15 Figure 4.15: Stress plots for the inferior and superior artificial facet caps with screws at L4-L5 with 6 N-m bending moment in all loading modes Figure 4.16: Stress plots for the inferior and superior artificial facet caps with screws at L4-L5 with 400N follower load and 6 N-m bending moment in all loading modes Table 4.11: Peak von Mises stress (MPa) in inferior artificial facet cap designs for 6 N-m applied moment Table 4.12: Peak von Mises stress (MPa) in superior artificial facet cap designs for 6 N-m applied moment Table 4.13: Peak von Mises stress (MPa) in inferior and superior artificial facet cap designs for 6 N-m moment and 400N follower load Table 4.14: Peak von Mises stress (MPa) in inferior and superior artificial facet cap screws for 6 N-m applied moment Table 4.15: Peak von Mises stress (MPa) in inferior and superior artificial facet cap screws for 400N follower load and 6 N-m applied moment Table 4.16: Peak von Mises stress (MPa) in bony pedicle of L5 for intact and artificial facet cap cases for 6 N-m moment Table 4.17: Peak von Mises stress (MPa) in bony pedicle of L5 for intact and artificial facet cap cases for 6 N-m moment and 400N follower load Figure 4.17: Stress plots for the 3mm stem pedicle screw based artificial facets at L4-L5 with 6 N-m bending moment in all loading modes Figure 4.18: Stress plots for the 3mm stem pedicle screw based artificial facets at L4-L5 with 400N follower load and 6 N-m bending moment in all loading modes Figure 4.19: Stress plots for the 5mm stem pedicle screw based artificial facets at L4-L5 with 6 N-m bending moment in all loading modes Figure 4.20: Stress plots for the 5mm stem pedicle screw based artificial facets at L4-L5 with 400N follower load and 6 N-m bending moment in all loading modes Figure 4.21: Stress plots for the supported stem pedicle screw based artificial facets at L4-L5 with 6 N-m bending moment in all loading modes Figure 4.22: Stress plots for the supported stem pedicle screw based artificial facets at L4-L5 with 400N follower load and 6 N-m bending moment in all loading modes xv

16 Table 4.18: Peak von Mises stress (MPa) in inferior pedicle screw based facet designs for 6 N-m applied moment Table 4.19: Peak von Mises stress (MPa) in superior pedicle screw based facet designs for 6 N-m applied moment Table 4.20: Peak von Mises stress (MPa) in inferior and superior pedicle screw based facet designs for 6 N-m moment and 400N follower load Table 4.21: Peak von Mises stress (MPa) in L4 pedicle screw for facet designs and a rigid system with 6 N-m applied moment Table 4.22: Peak von Mises stress (MPa) in L5 pedicle screw for facet designs and a rigid system with 6 N-m applied moment Table 4.23: Peak von Mises stress (MPa) in L4 and L5 pedicle screw for facet designs and a rigid system with 400N follower load and 6 N-m applied moment Figure 4.23: Stress plots for the 3mm stem pedicle screw based artificial facets at L4-L5 across a wide laminectomy with 6 N-m bending moment in all loading modes Figure 4.24: Stress plots for the 3mm stem pedicle screw based artificial facets at L4-L5 across a wide laminectomy with 400N follower load and 6 N-m bending moment in all loading modes Figure 4.25: Stress plots for the 5mm stem pedicle screw based artificial facets at L4-L5 across a wide laminectomy with 6 N-m bending moment in all loading modes Figure 4.26: Stress plots for the 5mm stem pedicle screw based artificial facets at L4-L5 across a wide laminectomy with 400N follower load and 6 N-m bending moment in all loading modes Figure 4.27: Stress plots for the supported stem pedicle screw based artificial facets at L4-L5 across a wide laminectomy with 6 N-m bending moment in all loading modes Figure 4.28: Stress plots for the supported stem pedicle screw based artificial facets at L4-L5 across a wide laminectomy with 400N follower load and 6 N-m bending moment in all loading modes Table 4.24: Peak von Mises stress (MPa) in inferior pedicle screw based facet designs with wide laminectomy for 6 N-m applied moment Table 4.25: Peak von Mises stress (MPa) in superior pedicle screw based facet designs with wide laminectomy for 6 N-m applied moment xvi

17 Table 4.26: Peak von Mises stress (MPa) in inferior and superior pedicle screw based facet designs with wide laminectomy for 6 N-m moment and 400N follower load Table 4.27: Peak von Mises stress (MPa) in L4 pedicle screw for facet designs with wide laminectomy and a rigid system with 6 N-m applied moment Table 4.28: Peak von Mises stress (MPa) in L5 pedicle screw for facet designs with wide laminectomy and a rigid system with 6 N-m applied moment Table 4.29: Peak von Mises stress (MPa) in L4 and L5 pedicle screw for facet designs with wide laminectomy and a rigid system with 400N follower load and 6 N-m applied moment Figure A.1: The spinal column displaying the cervical, thoracic, lumbar, sacral, and cocygeal regions [53] Figure A.2: Anatomic regions on lumbar vertebrae [3] Figure A.3: Annular fibers and the fiber orientation in the intervertebral disc [3] Figure A.4: The seven major ligaments of the lumbar spine [3] Figure A.5: Neural anatomy of the lumbar spine [3] Table B.1: Range of motion for lumbar segments in all rotation modes [53] Figure C.1: (A) Experimental FSU to determine strains in the posterior instrumentation. (B) Specimen setup using the MTS machine Figure C.2: Results of measured moment versus applied moment in a cadaveric study Figure C.3: FE model simulation of the strain experiment Table C.1: Material properties simulated in the FE model for the disc inserts Figure C.4: Results of measured moment versus applied moment in a FE study Figure C.5: FE predicted flexion motion for L3 and L4 vertebrae with 400N compression compared to Schultz [56] and Dooris [53] Figure C.6: FE predicted extension motion for L3 and L4 vertebrae with 400N compression compared to Schultz [56] and Dooris [53] Figure C.7: FE predicted lateral bending motion for L3 and L4 vertebrae with 400N compression compared to Schultz [56] and Dooris [53] xvii

18 Figure C.8: FE predicted axial rotation motion for L3 and L4 vertebrae with 400N compression compared to Schultz [56] and Dooris [53] Figure D.1: Modified universal joint across the L4-L5 motion segment Figure E.1: Facet load components in the lateral, anterior-posterior shear, and vertical directions and the percent changes as compared to intact for artificial facet FE models in extension, bending, and rotation for 6 N-m applied moment across L3-L4 right and left facet joints. A positive percent change indicates an increase in facet loads where as a negative percent change indicates a decrease in facet loads Figure E.2: Facet load components in the lateral, anterior-posterior shear, and vertical directions and the percent changes as compared to intact for artificial facet FE models in extension, bending, and rotation for 6 N-m applied moment across L4-L5 right and left facet joints. A positive percent change indicates an increase in facet loads where as a negative percent change indicates a decrease in facet loads Figure E.3: Facet load components in the lateral, anterior-posterior shear, and vertical directions and the percent changes as compared to intact for artificial facet FE models in extension, bending, and rotation for 6 N-m applied moment across L5-S1 right and left facet joints. A positive percent change indicates an increase in facet loads where as a negative percent change indicates a decrease in facet loads Figure E.4: Facet load components in the lateral, anterior-posterior shear, and vertical directions and the percent changes as compared to intact for artificial facet FE models in extension, bending, and rotation for 400N follower load and 6 N-m applied moment across L3-L4 right and left facet joints. A positive percent change indicates an increase in facet loads where as a negative percent change indicates a decrease in facet loads Figure E.5: Facet load components in the lateral, anterior-posterior shear, and vertical directions and the percent changes as compared to intact for artificial facet FE models in extension, bending, and rotation for 400N follower load and 6 N-m applied moment across L4-L5 right and left facet joints. A positive percent change indicates an increase in facet loads where as a negative percent change indicates a decrease in facet loads Figure E.6: Facet load components in the lateral, anterior-posterior shear, and vertical directions and the percent changes as compared to intact for artificial facet FE models in extension, bending, and rotation for 400N follower load and 6 N-m applied moment across L5-S1 right and left facet joints. A positive percent change indicates an increase in facet loads where as a negative percent change indicates a decrease in facet loads xviii

19 Figure E.7: Facet load components in the lateral, anterior-posterior shear, and vertical directions and the percent changes as compared to intact for artificial facet FE models with wide laminectomy in extension, bending, and rotation for 6 N-m applied moment across L3-L4 right and left facet joints. A positive percent change indicates an increase in facet loads where as a negative percent change indicates a decrease in facet loads Figure E.8: Facet load components in the lateral, anterior-posterior shear, and vertical directions and the percent changes as compared to intact for artificial facet FE models with wide laminectomy in extension, bending, and rotation for 6 N-m applied moment across L4-L5 right and left facet joints. A positive percent change indicates an increase in facet loads where as a negative percent change indicates a decrease in facet loads Figure E.9: Facet load components in the lateral, anterior-posterior shear, and vertical directions and the percent changes as compared to intact for artificial facet FE models with wide laminectomy in extension, bending, and rotation for 6 N-m applied moment across L5-S1 right and left facet joints. A positive percent change indicates an increase in facet loads where as a negative percent change indicates a decrease in facet loads Figure E.10: Facet load components in the lateral, anterior-posterior shear, and vertical directions and the percent changes as compared to intact for artificial facet FE models with wide laminectomy in extension, bending, and rotation for 400N follower load and 6 N-m applied moment across L3-L4 right and left facet joints. A positive percent change indicates an increase in facet loads where as a negative percent change indicates a decrease in facet loads Figure E.11: Facet load components in the lateral, anterior-posterior shear, and vertical directions and the percent changes as compared to intact for artificial facet FE models with wide laminectomy in extension, bending, and rotation for 400N follower load and 6 N-m applied moment across L4-L5 right and left facet joints. A positive percent change indicates an increase in facet loads where as a negative percent change indicates a decrease in facet loads Figure E.12: Facet load components in the lateral, anterior-posterior shear, and vertical directions and the percent changes as compared to intact for artificial facet FE models with wide laminectomy in extension, bending, and rotation for 400N follower load and 6 N-m applied moment across L5-S1 right and left facet joints. A positive percent change indicates an increase in facet loads where as a negative percent change indicates a decrease in facet loads Figure F.1: L3-L4, L4-L5, and L5-S1 intervertebral disc stress contours (MPa) of the nucleus pulposus and annulus fibrosus for each artificial facet and intact FE model with 6 N-m applied moment in extension xix

20 Figure F.2: L3-L4, L4-L5, and L5-S1 intervertebral disc stress contours (MPa) of the nucleus pulposus and annulus fibrosus for each artificial facet and intact FE model with 6 N-m applied moment in flexion Figure F.3: L3-L4, L4-L5, and L5-S1 intervertebral disc stress contours (MPa) of the nucleus pulposus and annulus fibrosus for each artificial facet and intact FE model with 6 N-m applied moment in lateral bending Figure F.4: L3-L4, L4-L5, and L5-S1 intervertebral disc stress contours (MPa) of the nucleus pulposus and annulus fibrosus for each artificial facet and intact FE model with 6 N-m applied moment in axial rotation Figure F.5: L3-L4, L4-L5, and L5-S1 intervertebral disc stress contours (MPa) of the nucleus pulposus and annulus fibrosus for each artificial facet and intact FE model with 400N follower load and 6 N-m applied moment in extension Figure F.6: L3-L4, L4-L5, and L5-S1 intervertebral disc stress contours (MPa) of the nucleus pulposus and annulus fibrosus for each artificial facet and intact FE model with 400N follower load and 6 N-m applied moment in flexion Table F.1: L3-L4, L4-L5, and L5-S1 intervertebral disc peak von Mises stress values (MPa) and percent changes as compared to intact of the nucleus pulposus for each artificial facet and intact FE model with 6 N-m applied moment in extension, flexion, bending, and rotation. A positive percent change indicates an increase in stresses where as a negative percent change indicates a decrease in stresses as compared to intact Table F.2: L3-L4, L4-L5, and L5-S1 intervertebral disc peak von Mises stress values (MPa) and percent changes as compared to intact of the nucleus pulposus for each artificial facet and intact FE model with 400N follower load and 6 N-m applied moment in extension and flexion. A positive percent change indicates an increase in stresses where as a negative percent change indicates a decrease in stresses as compared to intact Table F.3: L3-L4, L4-L5, and L5-S1 intervertebral disc peak von Mises stress values (MPa) and percent changes as compared to intact of the annulus fibrosus for each artificial facet and intact FE model with 6 N-m applied moment in extension, flexion, bending, and rotation. A positive percent change indicates an increase in stresses where as a negative percent change indicates a decrease in stresses as compared to intact Table F.4: L3-L4, L4-L5, and L5-S1 intervertebral disc peak von Mises stress values (MPa) and percent changes as compared to intact of the annulus fibrosus for each artificial facet and intact FE model with 400N follower load and 6 N-m applied moment in extension and flexion. A positive percent change indicates an increase in stresses where as a negative percent change indicates a decrease in stresses as compared to intact xx

21 Figure F.7: L3-L4, L4-L5, and L5-S1 intervertebral disc stress contours (MPa) of the nucleus pulposus and annulus fibrosus for each artificial facet with wide laminectomy and intact FE model with 6 N-m applied moment in extension Figure F.8: L3-L4, L4-L5, and L5-S1 intervertebral disc stress contours (MPa) of the nucleus pulposus and annulus fibrosus for each artificial facet with wide laminectomy and intact FE model with 6 N-m applied moment in flexion Figure F.9: L3-L4, L4-L5, and L5-S1 intervertebral disc stress contours (MPa) of the nucleus pulposus and annulus fibrosus for each artificial facet with wide laminectomy and intact FE model with 6 N-m applied moment in lateral bending Figure F.10: L3-L4, L4-L5, and L5-S1 intervertebral disc stress contours (MPa) of the nucleus pulposus and annulus fibrosus for each artificial facet with wide laminectomy and intact FE model with 6 N-m applied moment in axial rotation Figure F.11: L3-L4, L4-L5, and L5-S1 intervertebral disc stress contours (MPa) of the nucleus pulposus and annulus fibrosus for each artificial facet with wide laminectomy and intact FE model with 400N follower load and 6 N-m applied moment in extension Figure F.12: L3-L4, L4-L5, and L5-S1 intervertebral disc stress contours (MPa) of the nucleus pulposus and annulus fibrosus for each artificial facet with wide laminectomy and intact FE model with 400N follower load and 6 N-m applied moment in flexion Table F.5: L3-L4, L4-L5, and L5-S1 intervertebral disc peak von Mises stress values (MPa) and percent changes as compared to intact of the nucleus pulposus for each artificial facet with wide laminectomy and intact FE model with 6 N-m applied moment in extension, flexion, bending, and rotation. A positive percent change indicates an increase in stresses where as a negative percent change indicates a decrease in stresses as compared to intact Table F.6: L3-L4, L4-L5, and L5-S1 intervertebral disc peak von Mises stress values (MPa) and percent changes as compared to intact of the nucleus pulposus for each artificial facet with wide laminectomy and intact FE model with 400N follower load and 6 N-m applied moment in extension and flexion. A positive percent change indicates an increase in stresses where as a negative percent change indicates a decrease in stresses as compared to intact Table F.7: L3-L4, L4-L5, and L5-S1 intervertebral disc peak von Mises stress values (MPa) and percent changes as compared to intact of the annulus fibrosus for each artificial facet with wide laminectomy and intact FE model with 6 N-m applied moment in extension, flexion, bending, and rotation. A positive percent change indicates an increase in stresses where as a negative percent change indicates a decrease in stresses as compared to intact xxi

22 Table F.8: L3-L4, L4-L5, and L5-S1 intervertebral disc peak von Mises stress values (MPa) and percent changes as compared to intact of the annulus fibrosus for each artificial facet with wide laminectomy and intact FE model with 400N follower load and 6 N-m applied moment in extension and flexion. A positive percent change indicates an increase in stresses where as a negative percent change indicates a decrease in stresses as compared to intact xxii

23 Chapter I Introduction Chapter Overview The following chapter discusses the importance of studying the lumbar spine. The anatomy of the lumbar spine will be briefly discussed. Low back diseases will be introduced with a description of current treatments. Lastly, the scope of the current study is defined. Significance of Back Pain Low back pain is a debilitating disease that may strike a person of any age. Chronic back pain is the number one reason for healthcare expenditures in the United States [1]. Most commonly cited in worker s compensation claims, back pain is also the most frequent cause of employee absenteeism [1]. In 1998 alone, $90 billion was spent on back pain in the United States [1]. It is easy to understand the importance and need for development and advancement of new low back pain treatment techniques. Lumbar Spinal Anatomy The spinal column consists of five main regions: cervical, thoracic, lumbar, sacral, and cocygeal. Each region contains several bony vertebrae separated by flexible intervertebral discs. Two vertebrae connected with an intervertebral disc and associated 1

24 2 joints and ligaments are defined as a motion segment. All motion segments act together to support the upper body and protect the spinal cord. The lumbar spinal region consists of five vertebrae known as L1 through L5. The anterior region of the spinal column contains the vertebral body and intervertebral discs, while the posterior region consists of the bilateral facet joint and posterior elements including two pedicles, laminae, and transverse processes and the spinous process (Figures 1.1A and B). The facet joint is made of two bony protrusions from the inferior lamina and superior pedicle. The facet joint protrusions are enclosed in a connective tissue capsule to create a synovial joint with cartilage protecting the articulating (sliding) surfaces of the joint. The facet joints support and restrict movements in flexion, extension, lateral bending, and axial rotation. The intervertebral disc is composed of two sections; the annulus fibrosus and nucleus pulposus. The annulus fibrosus consists of several layers oriented in alternating directions from one layer to the next, providing strength in tension and enclosure for the nucleus pulposus. The nucleus is a gel-like substance composed of water, collagen, and proteoglycans. The water content of the nucleus is very high and creates a hydrostatic pressure to maintain the shape and flexibility of the disc and resist compression. Seven ligaments attach bone to bone in the spinal column (Figure 1.2). The anterior and posterior longitudinal ligaments are on the anterior and posterior sides of the vertebral body, respectively. The posterior ligaments include the intertransverse, interspinous,

25 3 supraspinous, ligamentum flavum, and capsular ligaments. The capsular ligaments surround the facet joint and provide enclosure and support to the joint. A more detailed description of the lumbar spine anatomy is in Appendix A. A B Figure 1.1: (A) Lateral view of a lumbar motion segment showing anatomical features [12]. (B) Transverse view of a lumbar vertebra [12]. Figure 1.2: Ligaments of the lumbar spine [13].

26 4 LBP and Treatments The cause of low back pain is patient specific and highly variable and may be due to several different disease states. Disc degeneration disease (DDD) is the most common disabling spinal disease [2]. However, disc degeneration has been proven to be the source of pain in a very small percentage of patients [3]. A degenerated disc decreases in hydration and height, possibly pinching exiting nerve roots to result in pain. The annular fibers undergo some fatigue failure and decomposition begins, resulting in immunological and inflammatory responses [3]. Spinal stenosis, the narrowing of passageways, has potential to cause back pain. Patients with spinal stenosis often find pain decreases when the spine is in a slightly flexed position, relieving pressure within the lumbar canal [3]. Spinal stenosis is often treated by decompression laminectomy or facetectomy surgery [4]. In a laminectomy procedure, the spinous process and lamina are removed at the level of stenosis along with associated ligaments [3, 5]. The facet joints may also be partially or completely removed in a facetectomy [6]. Facet joints may also be the cause of low back pain. The facet joints are innervated by the medial branch of the posterior primary ramus of the exiting spinal nerve [7, 8]. The joint capsule is also fully innervated and may lead to further pain [7]. It has been suggested that a cortisone injection to the facet reduces pain in some patients and can be used as a diagnostic tool to determine other spinal diseases [3,7,8]. Facet hypertrophy may cause nerve root irritation due to projection of the facet into the spinal canal, thereby

27 5 pinching nerves [3]. Controversy exists whether severe facet tropism (asymmetry of the bilateral facet joints) causes pain. Several authors have claimed a relationship between tropism and disc degeneration so that pain would be felt, while others find no relationship or pain indication [9,10,11]. When noninvasive procedures such as medication and exercise fail to relieve LBP, surgical intervention is considered. Most commonly, spinal fusion will be performed to treat chronic, debilitating low back pain. Fusion involves immobilizing a motion segment using bone grafts or mechanical devices such as cages as a replacement for the natural intervertebral disc. However, fusion is only 50% to 70% successful in clinical trials [1]. While fusion relieves back pain early, pain is likely to return due to facet hypertrophy, facet arthropathy, spinal stenosis, osteophyte formation, and posterior muscular debilitation, as well as disc degeneration at adjacent levels to the fusion site [1]. Thus, young patients are likely to have additional spine surgery due to fusion disease. Therefore, non-fusion techniques are currently being investigated and often used for treatment options. Artificial discs have recently become available and allow for motion at the diseased site, but many patients are not candidates since contraindications include arthritic, absent, or deformed facets and severe facet hypertrophy. In cases of severe facet tropism, facet hypertrophy, arthritic or degenerated facet joints, spinal stenosis, after laminectomy and facetectomy surgeries, and in addition to artificial discs, facet joint replacement may be a viable option.

28 6 Scope of Study The purpose of this study is to evaluate the biomechanical effectiveness of various artificial facet designs in the lumbar spine. The hypothesis is that artificial facets will restore motion of a destabilized lumbar spine to that of intact while maintaining other biomechanical factors such as stress and load distribution similar to that of intact. An experimentally validated, three dimensional, ligamentous finite element model of the L3-S1 spine was used to determine range of motion and facet loads throughout the lumbar spine and implant stresses as a result of artificial facets placed at the L4-L5 motion segment. Designs such as capping the facets, using pedicle screws with sliding articulating surfaces for the facet joint, and placing a modified universal joint across the facet joint was analyzed in models with a bilateral facetectomy and wide laminectomy and facetectomy. Stress distribution, load transfer, and rotational characteristics predicted by the model were compared to the intact spine, a destabilized spine with a wide laminectomy, and a rigid screw and rod system commonly used for fusion surgeries. A preliminary, in vitro study using two fresh, ligamentous L2-S1 spines were also evaluated with a pedicle screw based artificial facet design with sliding articulating surfaces replacing the L4-L5 facet joint. Relative motions of the intact, destabilized condition by removal of the L4-L5 facet joint and implanted conditions were compared to the finite element model predicted motions.

29 Chapter II Literature Review Chapter Overview An explanation of an aging spine will be given discussing different spinal disorders such as spinal degenerative changes, disc degeneration, spinal stenosis, facet hypertrophy, and facet tropism. Treatment options for many spinal diseases such as laminectomies and facetectomies, fusion, and artificial discs will be discussed. Although literature concerning artificial facets lacks possible indications and reasons the facet joint must be considered when treating LBP patients is given. The chapter is concluded with a brief overview of current patents for artificial facets. Aging Spine As the spine ages, it is inevitable degenerative changes will occur. Degeneration often starts with biochemical alterations, then micro-structural, and finally gross structural changes of the spine [14]. Between the third and fifth decades of life, major degenerative changes occur such as degeneration of the intervertebral discs and facet joints. There are several diseases brought about by aging including, but not limited to, osteoporosis, degenerative disc disease, facet degeneration, facet tropism, facet osteoarthritis, and spinal stenosis. 7

30 8 Spinal Disorders Osteoporosis The degenerative changes that occur throughout the spinal column as one ages affect many different structures. Osteoporosis is the weakening of bone, a result of an increase in porosity of trabecular bone, occurring most commonly in the vertebral body. Most often a disease inflicted upon women, men can also lose up to 30% of bone density and women up to 50%, starting around age 40 [15]. The loss in bone density reduces the vertebral body to withstand compression loads and vertebral fractures become prevalent. Persons with a decreased bone density of 0.05 g/cm 3 have a 99% chance of vertebral fractures [15]. Osteophytes Although osteophytes themselves do not cause pain, osteophyte formation is an indicator for the development of osteoarthritis [8]. Osteophytes occur among women and men alike and seem to increase in frequency with heavy physical activity levels [8]. Osteophytes not only occur on the vertebral body, but also in the facet joints [16,17]. Disc Degeneration There are several reasons and causes for disc degeneration. Degeneration occurs after cellular activity producing extracellular matrix decreases or stops [14]. After the cellular changes begin, gross anatomic changes appear. Often the annulus begins to tear radially and separate from the vertebral body, as well as display more macroscopic mechanical degradation [14,15]. The disc also thins as one ages due to a loss of water content [15, 18] causing a decrease in disc height. The material properties of the disc nucleus pulposus shift from fluid like to solid like [15]. It has been found that degenerative

31 9 changes in one disc affect adjacent levels adversely. In Kim et al it was shown with an FE model that intradiscal pressures increased in the disc immediately above the degenerated level suggesting further degeneration will occur along the spinal column [19]. Loss of disc height, irregular end plates, sclerosis of the disc, and osteophyte formation, as well as facet joint osteoarthritis is often the result of disc degeneration [18]. Facet Joint Osteoarthritis Facet joints are adversely affected by degenerative changes. Facet joints with osteoarthritis exhibit changes in swelling, stiffness, deformity, instability, a decreased range of motion, and a change in load-bearing [16]. These changes to the joints may cause significant changes to segmental motion throughout the spinal column [16] and changes in stress distribution and load sharing [8]. Osteophytes have also been found during early stages of facet joint osteoarthritis and tend to decrease spinal motion [16,17]. The loading path increases through adjacent discs due to facet joint destruction, which may accelerate disc degeneration [17]. Facet joint degeneration changes are almost always associated with disc degeneration [8,17,20,21]. Facet Hypertrophy Spinal pain may also be caused by hypertorphic facet joints. The superior facet often becomes hypertrophied and leads to nerve root irritation [3]. The facet may project medially and pinch the nerve root between the facet and the disc [3]. The facet may also project cephald and pinch the nerve root between it and the pedicle [3]. Inferior and superior facet hypertrophy commonly induces spinal stenosis [4].

32 10 Spinal Stenosis Spinal stenosis is the narrowing of the lumbar spinal canal that leads to compression of neural roots [4,22]. Spinal stenosis often occurs during the fifth and seventh decades of life [4]. There are several mechanisms which cause spinal stenosis. Degenerated discs collapse the intervertebral foramen thereby compressing the exiting nerve [4]. Both inferior and superior facet hypertrophy cause central and lateral stenosis, narrowing of the lumbar spinal canal and intervertebral foramen, respectively [4]. Significant pain begins in the low back and radiates down the buttocks and is worsened by walking, exercising, or standing for persons suffering from spinal stenosis [4,22]. Pain relief is often felt by sitting, leaning forward, or squatting [4,22]. Non-operative treatments seem to offer no abatement of the symptoms [4,23] and some type of surgical intervention usually prevails. Facet Tropism Facet tropism is the asymmetry of the bilateral facet joint angles in the lumbar and lumbosacral regions [10,11,24,25]. The role of facet tropism is not well understood. Several investigations have been completed to determine if asymmetry has a role in disc degeneration or failure, even though 23% of patients without low back pain exhibit asymmetry [26]. Ahmed et al found no correlation between asymmetry and disc failure noting that regardless of facet orientation, axial rotation was refrained by the facet joints. However, if a coupled motion of axial rotation and flexion occurs, disc failure is probable with great tropism [27]. Farfan et al discovered there is a correlation between asymmetrical joints and disc pathology, specifically stating a correlation between the sides of disc prolapsed and the more obliquely directed facet [26]. Facet tropism was

33 11 determined by Park et al to be a possible key factor in far lateral lumbar disc herniation [24], further advocating Farfan s suggestion that disc degeneration occurs on the side of greater tropism. In Karacan s more recent study, disc herniation and asymmetry were correlated [25]. If disc herniation is more likely with facet tropism, disc degeneration may be as well. However, several conflicting reports have been published stating there is a correlation [11] and there is not a correlation [9,10] between disc degeneration and facet tropism. Treatment Options There are several treatment options for low back pain sufferers. While many surgical options are available and will be discussed, those alternatives should only be exercised in a worst case condition. Non-invasive practices like exercises and physical therapy as well as medications may alleviate pain. Several spinal braces are available for different disorders that may successfully treat some patients. Epidural steroid injections are an additional treatment, but are not found effective for distinct pathologies such as spinal stenosis or disc herniation [3]. When non-invasive procedures fail to relieve pain, surgical intervention occurs. Decompression Decompression surgery is often performed to relieve pain due to nerve compression from spinal stenosis, facet hypertrophy, disc degeneration, or other disease situations [3]. There are several different categories of decompression surgeries. Only a laminectomy and facetectomy surgery will be discussed. In a laminectomy surgery, the lamina, spinous process, and all associated ligaments are potentially removed to decompress the

34 12 nerve root while leaving the facet joint intact. Different degrees of a laminectomy are performed from only a small portion of the lamina being removed to a wide laminectomy, in which all the spinous process and ligaments are removed along with the entire lamina, while retaining the facets [5]. A facetectomy removes only the facet joint [5]. A laminectomy and facetectomy can be performed together to remove the entire group of posterior elements. Several biomechanical investigations have been performed on spine stability after decompression surgery. Several investigators have found an increased risk of slippage after surgery [6,28]; however, Mariconda et al did not detect any slippage after a clinical trial [29]. Johnsson et al subjectively assessed patients following decompression surgery and grouped them based on postoperative results of good (no symptoms or slight residual pain, but clearly improved walking capacity) and poor (unchanged or increased pain after surgery) [6]. Slippage was found in both groups and 20% of all patients; mean slippage was 4.4 ± 3.3mm in the good group and 5.75 ± 3.6mm in the poor group [6]. A general trend of decreased disc height was also found in all groups [6]. In Lee s study, all patients with preoperative slippage had further postoperative slippage and 3.7% of all decompression patients developed postoperative slippage [28]. Postoperative movements in forward bending and axial rotation are discouraged after a decompression surgery due to a decrease in spinal stability [30]. In a finite element spine model, Zander et al determined a facetectomy increases motion in rotation and annular disc stresses [30]. After a wide laminectomy, motion increased in flexion, but little

35 13 change in bending and extension was noted [30]. In vitro biomechanical studies have found range of motion increases significantly in flexion and axial rotation due to facetectomies [31,32]. Abumi et al determined the range of motion for different degrees of facetectomies and reported motion increases in rotation with increasing degrees of facetectomy such that a complete bilateral facetectomy resulted in the greatest range of motion [31]. Flexion also resulted in increased range of motion when a facetectomy was performed at any degree of bone removal [31]. Pintar et al studied the motion of functional spinal units under compression and flexion. As the compression load was increased, instability increased when a facetectomy was performed [32]. When the posterior ligaments were removed, further instability was noted [32]. In finite element studies of decompression surgeries, it was determined that stresses increase greatly in the annulus and small loads increase displacement by large degrees with a facetectomy greater than 50% [33,34]. Lee et al compared different laminectomy and facetectomy condition motions and found angular motion increases in axial rotation, while removal of facets and posterior elements has the least effect in lateral bending [33]. When a bilateral laminectomy and bilateral facetectomy was examined, Lee et al reported motion increased considerably in flexion and extension [33]. Motion increased greatly in rotation when only the facet joint was removed indicating the importance of the facets in limiting axial rotation [33]. However, a degenerated disc results in less angular rotation than a normal disc [33]. Annular stresses increase significantly in rotation as well as in flexion and extension [33]. Teo et al also found an increase in flexibility of a lumbar motion segment when different degrees of facetectomies were performed in a finite

36 14 element model [34]. Flexibility increased significantly by 30% as compared to intact when a complete bilateral facetectomy was performed and an anterior shear load was applied [34]. Facet loads also decreased as the amount of facet removed was increased [34]. Such increases in instability of the lumbar spine after decompression surgeries such as laminectomies and facetectomies indicate the need for further intervention. Following wide laminectomies or complete facetectomies, many authors suggest the need for fusion to reduce the resultant instability [5,31,33,34]. Fusion Spine fusion, also known as arthrodesis, is frequently used to decrease motion between spine segments in the hope that pain will be alleviated [3]. Fusion is often completed in late stages of disc degeneration [18] and post-laminectomy or facetectomy surgeries due to spinal instability. Fusion is performed by placing a bone graft or interbody spinal fusion system such as BAK or PLIF cages in replacement of the disc nucleus pulposus. Fusion often results in adverse effects on adjacent segments to the fusion site. Excessive motion, degenerational changes, and spinal stenosis have been noted at adjacent levels [3]. Nagata et al found motion changes occur not only at the adjacent levels, but also at more distal levels when one or more segments are fused [35]. In a canine model, T6-T13 was fused and motion increased 43% in flexion and 34% in extension [35]. Facet loads also increased at other segments, maximally at the adjacent segment [35]. As the number of immobilized segments increased, the facet loads at the joint immediately adjacent to the

37 15 fusion site increased [35]. Ligament loading also changes considerably due to fusion. The posterior ligaments experience stress shielding after fusion which may lead to further back pain [36]. Okuda et al found that patients with segmental fusion that developed adjacent segment degeneration also had facet tropism and lamina horizontalization [37]. Artificial Discs Preserving spinal motion while relieving pain is desired for every patient; however, fusion eliminates motion at the diseased level and often causes future pain. Thus, artificial discs are being developed to restore natural motion to a diseased level. Few discs are currently on market, but preliminary results appear promising. In some short term studies, pain was relieved and the implants were found relatively safe [18]. Many complications with artificial discs have already risen. Ooij et al completed a clinical trial in which anterior migration of the disc, degeneration at other levels, subsidence of the prosthesis, facet joint arthrosis, and polyethylene wear occurred in several patients [38]. Development of facet joint hypertrophy may be increased resulting in spinal stenosis with the use of an artificial disc [39]. Guyer et al states that pain from the facet joint will not be addressed by an artificial disc, thus any facet joint disease should be considered a contraindication of artificial discs [39]. Dooris et al completed a finite element study on artificial discs and determined placing the disc more posteriorly increased the range of motion in flexion up to 44% more than intact depending on the amount of annulus retained during surgery [40]. In extension, the motion increased up to 40% more than intact when little annulus was retained [40]. Facet loads predicted by the FE model increased 150% over the intact loads when the disc was placed more anteriorly,

38 16 clearly showing the importance of proper disc placement and the imperativeness of facet joints [40]. Indications for an artificial disc are similar to fusion indications, but several contraindications exist such as arthritic, absent, or deformed facets and severe facet hypertrophy [41]. Little is understood about the role of facet joints with artificial discs. It has been found that artificial discs increase the range of motion [40,42]. Such increases in motion most likely place increased stress and strain on the facet joint possibly resulting in painful joints if the facet joints are degenerated or otherwise diseased [42]. Artificial Facets and Indications To restore normal function at a diseased site, artificial facets may be an alternative to fusion or other surgeries for treatment of severe facet tropism, facet hypertrophy, arthritic or degenerated facet joints, spinal stenosis, after laminectomy and facetectomy surgeries, and in addition to artificial discs. Facet replacements must restore normal motion in all modes of flexion, extension, lateral bending, and axial rotation. They must perform well under shear and torsional loads and be able to bear 20 to 30% of the physiological load [2]. The prosthesis also should be easy to place in all patients and fix to the bone well to reduce the risk of loosening. While the facet replacement market is miniscule today and all implants are in research and design stages, it has been estimated in the year 2010 to exceed $90 million with the average price per implant to be $7000 [1].

39 17 Role of Facet Joint While the intervertebral disc is principally involved in stabilizing the spine and supporting load, facet joints share the functions. Facet joint function is to guide vertebral body motion and resist compression, rotation, and shear. Facet joints share in supporting 10-15% of the load [43] and experience large stresses when in contact [44]. Such large loads and stresses give rise to many of the facet diseases. The facet joint is an important anatomical consideration and probable pain generator. When restoring normal function to patients with diseased spinal motion segments, the facet joint cannot be overlooked. Facet Joint Replacement Systems Several designs for facet joint replacements are currently being designed and investigated. While these designs are highly confidential, several patents have been awarded. A comprehensive examination of these proposed devices follows. Facet Capping In U.S. Patent Number awarded in 1996, an artificial facet joint was invented by Fitz that involves capping the diseased, damaged, or painful facet joint [43]. There are two components: a superior component caps the inferior facet and an inferior component caps the superior facet. Both of the components completely cover the articulating surface with a highly polished chrome or high density polyethylene. The superior component is a hollow cone in shape, designed to simply cover the inferior facet. The cap securely fastens to the inferior facet with a small screw at the tip of the facet. The inferior component, a hollow cone in shape, covers the distal portion of the superior facet joint. There are two small screws fastening the cap onto the superior facet and lamina. The

40 18 design shown in Figures 2.1 and 2.2 detail the placement of the screws [43]. The caps are constructed of stainless steel, unalloyed titanium, or a titanium-aluminum alloy. The interior surface of both prostheses is coated with a porous inner surface to facilitate bony ingrowth. This particular design will be manufactured in several shapes and sizes to replace all cervical, thoracic, and lumbar articular processes. Figure 2.1: Superior facet replacement cap showing placement of securing screws in U.S. Pat. No [43]. Figure 2.2: Inferior facet replacement cap with the superior screw placement shown from U.S. Pat. No [43].

41 19 While the author believes this design will reduce or eliminate pain in the joint, many others such as Goble et al and Serhan et al suggest that caps will not eliminate pain, only cover up the pain-producing area [41,45]. The capping construct is discouraged by surgeons due to the small size of the prostheses and difficulty in screw placement. Another disadvantage of the capping design is a distinct possibility of failure due to mechanical loosening. Evidence of mechanical loosening has been shown when capping articular bone ends such as femoral heads which eventually leads to avascular necrosis of the bony support structure [41,45]. Lee recently created another capping device that utilizes the pedicle for support in U.S. Published Patent Application No. 2005/ [2]. The prosthesis shown in Figure 2.3 is composed of a superior component conical in shape and slides over a surgically prepared inferior facet [2]. The inferior component is a cupped shape prosthesis with an interior surface roughly triangular in shape to fit over a surgically prepared surface. The superior and inferior components are supported by a pedicle screw and tabs that bend around the joint. While this prosthesis does use pedicle screws for extra support, it retains a portion of the diseased or painful facet joint. Once again, this design may not reduce or eliminate the patient s pain. A similar cap design will be evaluated further in the current study.

42 20 Figure 2.3: Superior and inferior facet caps with pedicle screw support from U.S. Published Patent Application No. 2005/ [2]. Cartilage Replacement A prosthesis supported by the lamina and spinous process that only replaces the cartilage of the facet joint has been suggested by Martin in U.S. Pat. No awarded in 2000 [46]. The implant replaces one or both facets in the joint, unilateral or bilateral. The articular surface of the facet is replaced by a blade having the same shape, position, and orientation of the natural facet. Support plates extend from the articular surface to anchor against the lamina and spinous process. It can be anchored in several ways such as screws, hooks, or clamps. The implant is constructed of stainless steel, titanium, or titanium alloy. The interior side of the implant is coated with hydroxyl-apatite. The patent art shown in Figure 2.4 displays the proposed artificial joint [46]. Note that the inferior and superior facet is intact, but the cartilaginous articulating surface has been removed.

43 21 Figure 2.4: Articular cartilage replacement system from U.S. Pat. No [46]. The Martin design has disadvantages in that it is supported by the lamina. The lamina is highly variable among persons and very complex. To reproduce the positioning of the implant consistently may be very difficult [41,45]. Also, the prosthesis has a strong indication it will not eliminate pain at the facet joint since it only replaces the cartilage covering the inferior and superior facets [41,45]. Articulating Surface Designs The designs that have been reviewed leave the inferior and superior facets intact by removing the cartilaginous lining or capping the bone. It has been strongly suggested that the bony structures may be causing the joint pain; therefore, the entire facet should be removed and replaced. U.S. Pat. No by Reiley awarded in 2003 does just that concept by removing the entire inferior and superior facet and replacing them with sliding articulating devices [47]. The inferior facet is replaced by a spherical component attached to a plate. The plate surrounds the posterior elements of the spine including attaching directly to the spinous process via screws shown by #515 in Figure 2.5 [47].

44 22 The entire plate construct is shown in Figure 2.6 displaying the member #515 and the pedicle screws that are used to attach the plate to the spine [47]. The superior facet is replaced by a cup member to articulate with the head of the inferior replacement. The superior replacement is supported by screws or stems that pass through the pedicle. The entire implant is constructed of biocompatible metal or other common medical materials. Figure 2.5: An transverse slice of the complete facet joint replacement system depicted in U.S. Pat No [47]. Figure 2.6: A posterior view of the plate supporting the inferior facet replacement from U.S. Pat No [47].

45 23 The design removes a large amount of natural bone, i.e. the lamina is completely removed as in a wide decompressive laminectomy, which may be disadvantageous. Although laminectomies are performed to manage spinal stenosis, it has been found to increase spondylolisthesis and instability [6,28,47]. Thus, it may be ascertained this design could result in spondylolisthesis. Also, the design requires the removal of several ligaments such as the supraspinous and interspinous to place the inferior facet plate at the site of facet replacement as well as the level above. This ligament removal has great potential to decrease spinal stiffness in several motions, especially in flexion [30,31,32,33,34]. However, the prosthesis has an advantage over previous designs in that it completely removes the painful facet joint. In 2002, Goble et al in U.S. Pat. No replaced the entire facet structure as well using a pedicle screw based design with sliding articulating surfaces [41]. Shown in Figures 2.7 and 2.8, the superior component replaces the inferior facet and attaches to the pedicle of the inferior facet vertebra [41]. The inferior component replaces the superior facet and attaches to the pedicle of the superior facet vertebra. Both components comprise of a facet body for articulation and a flange with a hole for a pedicle screw to pass through. When the flange is press fitted and perpendicular to the bone of the pedicle, the facet implant body is positioned at the same orientation of the natural facet. A similar design to the Goble patent will be discussed in detail in the current study.

46 24 Figure 2.7: Lateral view of the artificial facet proposed in U.S. Pat. No [41]. Figure 2.8: Posterior view of the superior and inferior components attached to the respective vertebra using pedicle screws shown in U.S. Pat. No [41]. U.S. Pat. No has several advantages over the designs already discussed. The prosthesis completely replaces the diseased or painful facet joint. The design also reduces error in placement due to the tight press fit against the pedicle bone. Another advantage is the prosthesis has porous bone coatings located at bone interfaces to promote bone ingrowth to improve the strength and stability [41]. However, a potential disadvantage of the design is the inferior replacement component may begin twirling about the pedicle screw axis after long term loosening [45].

47 25 Goble et al expanded the prosthesis in U.S. Pat. No to include a prosthesis for multiple facet joint replacement in U.S. Pat. No awarded in 2003 [48]. In this patent, the inferior facet body as discussed above is connected with a bridge that passes from the right to the left facet replacement. A flange then attaches the inferior facet replacement to the superior facet replacement. The right superior facet replacement is connected to the left replacement by a bridge. This design is shown in Figure 2.9 [48]. The replacement is securely attached to the pedicle with a tight press fit against the pedicle bone using pedicle screws. The most inferior facet to be replaced, the superior facet, is replaced with the design discussed above, Figure 2.10 [48]. Figure 2.9: A prosthesis designed to replace multilevel facet joints in U.S. Pat. No [48].

48 26 Figure 2.10: Posterior view of the device used to replace two levels of facet joints [48]. While this design utilizes a similar design as in U.S. Pat. No and has many advantages, a major disadvantage is that several posterior ligaments must be removed for proper placement. These ligaments include the capsular, supraspinous, and intertransverse and when removed result in instability [30,31,32,33,34]. Another disadvantage is that the facet joints at different levels are highly variable. However, the inventors have proposed creating a kit with several different sizes of multiple facet joint prostheses [48]. Posterior Element Replacement In extreme cases of disease or trauma, the facet joints may not be the only compromised structures [49]. Other posterior elements such as the transverse process, spinous process, pedicle, and lamina may also need replaced. In this circumstance, Fallin et al have proposed an entire posterior element replacement in U.S. Pat. No [49]. The invention provides several different replacement options with the most extreme case being that all posterior elements must be replaced (Figure 2.11) [49]. The prosthesis is a

49 27 plate that contains surfaces for the superior and inferior facet and prosthetic spinous and transverse processes. Once all the posterior elements have been removed from the vertebral body, it securely attaches to the body via prosthetic pedicles and pedicle screws. In Figure 2.11, member #150 notes holes in the prosthetic transverse and spinous process for possible reattachment of soft tissues [49] to prevent instability. Figure 2.11: Total posterior element replacement prosthesis proposed in U.S. Pat. No [49]. Ball & Socket Facet Replacement Systems A new concept for replacing facet joints was developed in U.S. Published Patent Application No. 2005/ by Berry involving a ball and socket joint [50]. The invention completely removes the inferior facet replacing it with a ball shaped head attached to a stem. The stem screws onto a translaminar fastener (Figure 2.12A). The superior facet is also excised and replaced by a screw with a cup shaped head. The superior facet replacement screws into the lamina and pedicle of the vertebra in such a way that the inferior head articulates with the cupped superior head. The motion of the head and cup may be restricted with the use of a tether or flexible sheath (Figure

50 B,C) [50]. The complete setup is shown in Figure 2.13 [50]. The invention also claims a replacement system for the spinous process, facet joints, and lamina posterior elements. The lamina and spinous process are replaced by a plate that screws into the pedicle with the inferior facet replacement stem described above attached (Figure 2.14) [50]. The inferior facet head articulates with the superior cup as previously explained. A flexible spinous process bumper fits over the prosthetic spinous process of the plate and extends between the spinous processes of the adjacent vertebrae. A B C Figure 2.12: (A) Translaminar screw and inferior facet replacement head as depicted by Berry. (B) Superior facet replacement screw is shown describing articulation with the inferior facet replacement where motion is restricted with a tether. (C) A sheath may also be used to restrict motion between the superior and inferior facet replacement [50]. Figure 2.13: The inferior and superior facets are replaced with a ball and socket design [50].

51 29 Figure 2.14: A replacement system for the posterior elements proposed by Berry [50]. Another ball and socket design has been proposed in U.S. Published Patent Application No. 2005/ by Serhan et al [45]. In this design, the inferior and superior facet articulation surfaces are convex and concave, respectively, to create a ball and socket joint. The superior and inferior components are attached to the vertebral body with pedicle screws. A unique feature of this prosthesis is the pedicle screw has a proximal groove and set screw so that the superior and inferior replacements may be adjusted to the proper distance by the surgeon (Figure 2.15) [45]. The present invention also includes a facet ligament shown in Figure 2.16 as member #109 that attaches either to the superior and inferior vertebrae or the superior and inferior prosthetic facet components to support the facet joint in tension as well as in compression [45]. Figure 2.16 also depicts another method for replacing the facet joint using sliding articulating surfaces for the inferior and superior facet with the components attaching to the pedicle with pedicles screws [45].

52 30 Figure 2.15: A ball and socket artificial facet joint proposed by Serhan et al [45]. Figure 2.16: A sliding articular surface facet replacement design with a novel ligament concept to support the joint in tension [45]. All U.S. patent applications and awarded patents can be found on the United States Patent and Trademark Office website ( Links to the patents on the website are included in Appendix G for further review and referencing.

53 31 Summary Low back pain is a significant problem. While several non-surgical and surgical solutions exist, all fall short of restoring normal motion to the diseased segment. Decompression surgery increases the instability, fusion decreases motion at the fused site but increases the risk of degeneration at adjacent levels, and artificial discs increase motion and have several contraindications. If normal function is to be restored to the diseased spine, artificial facets may be a viable treatment option. Several patents have been awarded for artificial facet designs including capping and sliding articulating surfaces. Each of these designs must be thoroughly evaluated to determine the biomechanical effectiveness before clinical use. These evaluations can be performed using finite element models of the lumbar spine and cadaveric studies.

54 Chapter III Materials and Methods Chapter Overview The materials and methods for designing a facet joint replacement system will be described in detail in the chapter. The reason for completing a finite element study has been described along with the intact L3-S1 spine model used throughout the study. A full description of the intact FE model has been given along with validation of the model. The intact model was modified for artificial facet joints at the L4-L5 motion segment and is described in detailed below. The capping method, pedicle screw based designs with articulating surfaces, and a modified universal joint design is fully explained. The chapter is concluded with an in vitro study to further validate the FE models. Finite Element Analysis A theoretical analysis of biological systems can readily be performed utilizing finite element analysis (FE) techniques. An FE model can efficiently compute biomechanical parameters in investigations where material properties vary, geometries are irregular, and several loading modes are necessary, as compared to closed-form solutions. Many parameters that are difficult or impractical to obtain in physical models, such as stress and strain fields, are solved during finite element simulations. Therefore, FE has become 32

55 33 an invaluable tool in the design and analysis of orthopaedic implants to examine tissue mechanical reactions to instrumentation and the load characteristics of the instruments. The following investigation will compare L3-S1 lumbar spine FE models implanted with different designs for facet joint replacement systems. These designs involve capping the L4-L5 inferior and superior facets, an L4-L5 pedicle screw construct, and a modified universal joint replacing the L4-L5 facets. Analysis will compare the predicted loads, stresses, and displacements of the intact spine to destabilized spine, a rigid system, and the artificial facet implanted models. Intact Finite Element Model Geometric Model The intact ligamentous, nonlinear, L3-S1 lumbar FE spine model (Figure 3.1B) used for this study was created from an intact ligamentous L3-L5 lumbar FE spine model (Figure 3.1A) developed by 51 [51]. The L3-S1 intact model consisted of 31,054 elements, 38,664 nodes, and was symmetric about the midsagittal plane (Figure 3.2), while the original L3-L5 intact model had 13,339 elements, 16,240 nodes, and was also symmetric about the midsagittal plane. Both model dimensions were obtained from computer tomography (CT) scans (transverse sections of 1.5 mm thickness) of a healthy, deformity free cadaveric spine. A lordotic curve of approximately 8º was simulated at the L3-L5 level. The model was constructed and analyzed with the commercial software package ABAQUS 6.4. The intact L3-L5 model was validated by 51 [51] and the intact L3-S1 model validation is discussed later and included further in Appendix C.

56 34 The mesh in the L3-S1 model was refined from the original L3-L5 mesh. The finite element method is an approximation such that as meshes change, numerical results change. As the number of elements in a body increases, the error decreases until the error approaches zero and an optimal mesh is realized. The L3-L5 model mesh was not optimized. In Table 3.2, the motions of the L3-L5 model are given as well as a comparison of the L3-S1 motions. The L3-S1 motions strongly correlated with the L3- L5 motions. Therefore, the L3-S1 refined mesh was optimal. Bony Element Modeling The vertebral bodies and posterior bone was defined as a cancellous bone core surrounded by a 0.5mm thick cortical bone shell. Thus, the material properties of bone were varied along the respective regions (Table 3.1). All of the bony elements were constructed with a three-dimensional hexagonal element (C3D8). A C3D8 element is defined using eight nodes with each node possessing three degrees of freedom. Apophyseal (Facet) Joint Anatomically, the inferior and superior facets contain a thin cartilaginous layer lining the articular surface. This thin layer was simulated with three-dimensional gap contact elements (GAPUNI). These elements transfer force along a single direction as a specified gap closes between nodes. An initial gap of 0.5mm was assumed based on CT images of cadaveric specimens. The thin cartilaginous layer was simulated with ABAQUS s softened contact parameter which adjusts force transfer between nodes exponentially, depending on the gap size. Upon full closure of the gap, the stiffness of the joint assumed the same stiffness as the surrounding bone. The lumbar facet joints

57 35 were oriented at an inclination of 72 from horizontal, determined from several CT images. Intervertebral Disc The intervertebral disc was modeled as a composite of a solid matrix with embedded fibers in concentric rings around a pseudofluid nucleus. Seven concentric rings of ground substance about the nucleus each contained two evenly spaced layers of fibers (plus one ground substance ring with one layer of fibers) oriented at ± 30 to the horizontal. The fibers were defined via the REBAR command. It was assumed an overall collagenous fiber content of 16% of the annular volume that was distributed amongst the seven layers [52]. The fiber thickness and stiffness increased in the radial direction and is shown in Table 3.1. A no compression option was defined for the annulus fibers so that the fibers resist tension only. Since physiologically the nucleus is fluid filled, the hydrostatic properties were simulated with hexagonal C3D8 elements assigned low stiffness (1MPa) values and near incompressibility (Poisson s ratio of ). Ligaments All seven major ligaments were represented in the intact spine model. These ligaments were as follows: anterior longitudinal, posterior longitudinal, intertransverse, ligamentum flavum, interspinous, supraspinous, and capsular. The ligaments were modeled as three dimensional, two node truss elements (T3D2) and assigned nonlinear material properties such that at initially low strains, the ligaments exhibited low stiffness, but as the strains increase, the ligament stiffness increased. This material property was simulated using the hypoelastic material designation which allows axial stiffness to be a function of axial strain. The material properties as well as cross-sectional areas of the ligaments are given

58 36 in Table 3.1 and were chosen based on reported literature [53]. The model s ligament elements were aligned in the direction of anatomical fiber orientation. All ligaments in the model were assumed to be unstressed at rest even though the ligamentum flavum and longitudinal ligaments experienced a pre-stress. The magnitude of the pre-stress was directly related to the geometrical configuration of the region. Thus, great variability existed among specimens such that precise values were difficult to achieve [55]. A B Figure 3.1: (A) Intact L3-L5 finite element mesh. (B) Intact L3-S1 finite element mesh.

59 37 Annulus Fibrosus L3 Superior Facet Pedicle L3/4 Disc Nucleus Pulposus L4 Inferior Facet L4/5 Disc Lamina L5 L5/S1 Disc Spinous Process S1 Figure 3.2: A midsagittal cross-section of L3-S1 lumbar spine indicating important anatomical features. Material Property Definitions Material properties defined in the above mentioned model are summarized in Table 3.1. All material properties were chosen after careful review of published literature. The model material properties were assumed to be isotropic and homogenous [54,55]. The ligament material properties were nonlinear and hypoelastic; therefore, a Young s modulus was assigned to a respective strain value. Boundary Conditions and Loading Modes Boundary conditions are an essential mechanism of finite element modeling. The inferior most nodes of S1 were fully constrained in all directions allowing unconstrained motion of all superior elements. These nodes included the S1 vertebral body and posterior elements of S1. Compressive loads were equally distributed amongst the superior most

60 38 nodes of the L3 vertebral body simulating physiological spinal loading conditions. The compressive loads were applied normal to the vertebral body throughout analysis, thus the load acted as a follower load. Due to individual nodes having no rotational degrees of freedom, moment loading was applied to stiff crossed beams rigidly attached to the superior most nodes of the L3 vertebral body. Finite Element Model Validation The model used for this study, L3-S1, was developed using a previously validated L3-L5 model. The L3-L5 model (Figure 3.1A) was validated by comparing numerical predictions of load-displacement behavior, ligament strains, and disc bulge with reported values in the literature. Load-displacement FE predictions were similar to in vitro values reported by Schultz et al shown in Table 3.2 [56]. Numerical displacement values were found by recording the absolute position of two nodes coordinates on opposite sides of each vertebral body, parallel with the plane of motion. After a pure moment was applied, the deflection of each node coordinate was measured and the appropriate angle calculated. Each measurement was taken in the frontal plane for lateral bending, transverse plane for axial rotation, and sagittal plane for flexion and extension. Predicted facet contact forces were compared with Yang and King [57], Shirazi-Adl [58], and Kim [59]. Disc bulge predictions measured at mid-disc height were compared to previous experimental and analytical results. Ligamentous strain predictions were compared with in vitro trends reported by Panjabi et al [53].

61 39 Validation of the load-displacement behaviors predicted by the L3-S1 FE model was also performed. Experimental load-displacement behaviors were determined using fresh, ligamentous, L1-S1 cadaveric spines (Figure 3.3A). The intact L3-S1 FE model validation was performed using specimens from a previous study; not the cadaveric study presented in this document. Another study was used to validate the model since the current cadaveric study described in detail below only involved two specimens and the study used to validate the model included five spines to allow for the use of statistical analysis. The specimens were thawed to room temperature and all soft tissue was removed, leaving the ligaments intact. Pure flexion, extension, left and right bending, and left and right axial rotation moments were applied and measured by an OptoTrak measuring system. Flexion and extension loading modes shown in Figure 3.3B, left and right bending in Figure 3.3C, and left and right rotation in Figure 3.3D display that the model predictions fall within one standard deviation of the experimental values. Additionally, the mesh density of the model in this study was increased and the current density represents an optimal value. Further validation of the L3-S1 FE model is explained in Appendix C.

62 40 Element Set Number of Elements ABAQUS Element Library Type Modulus of Elasticity (MPa) Poisson s Ratio, ν Cross-Sectional Area (mm 2 ) Bony Regions Vertebral Cortical Bone 3312 C3D Vertebral Cancellous Bone C3D Posterior Cortical Bone 3632 C3D Posterior Cancellous Bone 1834 C3D Intervertebral Disc Annulus (Ground Substance) 5376 C3D Annulus Fibers 2685 REBAR Nucleus Pulposus 1920 C3D Joints Apophyseal Joints 216 GAPUNI Softened, Ligaments Anterior Longitudinal 216 T3D Posterior Longitudinal 144 T3D Intertransverse 30 T3D Ligamentum Flavum 21 T3D Interspinous 21 T3D Supraspinous 9 T3D Capsular 84 T3D Table 3.1: Material property definitions and element types for the intact L3-S1 finite element spine model. For bilateral structures (facet joints and capsular ligaments), the total number of elements is listed [53].

63 41 Rotation ( ) from 4.7 Nm Moment N Compression L3/5 FE Model Predictions L3/S1 FE Model Predictions in-vitro Results Schultz et al. Flexion L3-L4: 3.29 L4-L5: 3.36 L3-L4: 3.20 L4-L5: 3.32 L5-S1: ± 1.86 Extension L3-L4: 1.84 L4-L5: 1.62 L3-L4: 1.67 L4-L5: 1.40 L5-S1: ± 0.98 Right Lateral Bending L3-L4: 2.33 L4-L5: 2.31 L3-L4: 2.32 L4-L5: 2.13 L5-S1: ± 1.63 Left Lateral Bending L3-L4: 2.33 L4-L5: 2.31 L3-L4: 2.32 L4-L5: 2.13 L5-S1: ± 1.47 Right Axial Rotation L3-L4: 1.28 L4-L5: 1.25 L3-L4: 1.34 L4-L5: 1.20 L5-S1: ± 0.33 Rotation ( ) from 10.6 Nm Moment N Compression Flexion L3-L4: 5.32 L4-L5: 5.08 L3-L4: 5.19 L4-L5: 5.00 L5-S1: ± 1.00 Extension L3-L4: 3.45 L4-L5: 3.35 L3-L4: 3.83 L4-L5: 3.80 L5-S1: ± 1.02 Right Lateral Bending L3-L4: 5.13 L4-L5: 5.18 L3-L4: 5.15 L4-L5: 4.91 L5-S1: ± 1.22 Left Lateral Bending L3-L4: 5.13 L4-L5: 5.18 L3-L4: 5.15 L4-L5: 4.91 L5-S1: ± 0.79 Right Axial Rotation L3-L4: 2.98 L4-L5: 2.75 L3-L4: 3.17 L4-L5: 2.97 L5-S1: ± 0.67 Table 3.2: Comparison of intact L3-L5 and L3-S1 finite element predictions and results from Schultz et al [56]. FE model predictions fall within one standard deviation of in vitro results.

64 42 A B Ang Displacement (Degs) Ang Displacement (Degs) CADAVERIC AND FEM COMPARISON Intact Right Rotation (Cadaveric) Intact Right Rotation (FEM) Intact Left Rotation (Cadaveric) Intact Left Rotation (FEM) L3-S1 L3-L4 L4-L5 L5-S1 Spinal Segment C CADAVERIC AND FEM COMPARISON Intact Right Bending (Cadaveric) Intact Right Bending (FEM) Intact Left Bending (Cadaveric) Intact Left Bending (FEM) L3-S1 L3-L4 L4-L5 L5-S1 Spinal Segment D Figure 3.3: (A) Ligamentous L1-S1 spinal motion segment used to determine bending moments. (B) Comparison of the experimental and FE results are in agreement for flexion and extension. (C) Experimental and FE results are compared in left and right bending. (D) Results from cadaveric studies and FE are compared in left and right rotation.

65 43 Artificial Facet Joint Designs Three different facet joint replacement systems were designed and simulated in an FE spine model. All designs were first modeled as Pro/Engineer CAD drawings and precise geometries were imported into ABAQUS 6.4 for meshing and analysis. Any screw geometries used in the designs were simplified by removing threads and angular ends. All metal implant components were modeled as titanium using the material properties given in Table 3.3. All models were stressed with a 6 N-m bending moment applied at the superior most surface of L3 in flexion, extension, lateral bending, and axial rotation. A 400N follower load and 6 N-m bending moment in flexion and extension was also applied to all models at the superior most surface of L3. The results from the FE model were then compared with results obtained from an in vitro study. Titanium Material Properties Young's Modulus 110 Gpa Poisson's Ratio 0.3 Table 3.3: Material properties of titanium used in all artificial facet models. Destabilized Spine with Wide Laminectomy The intact L3-S1 FE model was destabilized by performing a wide laminectomy. To model the wide laminectomy, the L4 spinous process, L3-L4 and L4-L5 interspinous and supraspinous ligaments and ligamentum flavum, and most of the L4 lamina was removed (Figure 3.4). The L4-L5 facet joint and capsular ligaments remained intact. The model was created to understand the biomechanical effects of a wide laminectomy and to determine if artificial facets would be effective in restoring stability and natural motion to the destabilized lumbar spine.

66 44 Figure 3.4: The intact model was modified with a wide laminectomy at L4-L5. Artificial Facet Caps The intact L3-S1 FE model was modified at the L4-L5 facet joint to place artificial facet caps. Partial ligamentum flavum and capsular ligaments were removed at L4-L5 as well as 2mm of the outer layer of each bony L4 inferior and L5 superior facet. 2mm thick titanium caps were placed over the reduced facets. The caps were modeled using eight node hexagonal elements. In one model, the metal caps were tied to the bony facets thereby simulating perfect bone osteointegration (Figure 3.5A). In another model, the caps were tied to the bony facets and further secured with screws 2mm in diameter and 8mm in length (Figure 3.5B). Two superior screws were tied to the L5 pedicle while one inferior screw was tied to the inferior L4 facet and lamina. The screws were modeled with hexagonal elements. In both models the GAPUNI elements were modeled as titanium.

67 45 L4 Inferior Facet L5 Superior Facet A L5 Superior Facet L4 Inferior Facet B Inferior Facet Screws Superior Facet Screws Figure 3.5: (A) L3-S1 FE model with artificial facet caps at L4-L5. (B) L3-S1 FE model with artificial facet caps secured with screws at L4-L5. Pedicle Screw Based Artificial Facet The pedicle screw based artificial facet utilizes pedicle screws to attach replacement facets in the L3-S1 FE spine model. The L4-L5 facet joint and capsular ligaments were removed simulating a facetectomy surgery. Pedicle screws 6.5mm in diameter and 55mm in length modeled with tetrahedral elements were inserted at L4 and L5 two-thirds into the vertebral body. Titanium sliding articulating surfaces representing the facets replaced the natural facets and attached to stems that surrounded and tied to the pedicle screws. In one model (3mm), the artificial facet stem was 3mm thick (Figure 3.6). The stems were then increased to 5mm thick shown in Figure 3.7 for the first design (5mm)

68 46 modification. In a third model and modification, the stems were 3mm thick and increased in width around the pedicle screws for further support (Figure 3.8). The increased stem width surrounded the pedicle screw and was tied to the screw as well as to the surrounding pedicle bone to simulate a press fit against the bony pedicle. Further reference to the model with increased supported stems around the pedicle screws will be known as support. In all models, the GAPUNI elements were modeled as titanium. L4 Inferior Facet Pedicle Screw L5 Superior Facet Figure 3.6: The pedicle screw based artificial facet design with a 3mm thick stem connecting the metal facets to the pedicle screw (3mm).

69 47 L4 Inferior Facet Pedicle Screw L5 Superior Facet Figure 3.7: The second design utilizing pedicle screws with a connecting stem thickness of 5mm (5mm). L4 Inferior Facet Pedicle Screw L5 Superior Facet Figure 3.8: The pedicle screw design was modified such that the connecting stems increased in width around the pedicle screws and were tied to the screws and bony pedicles (Support).

70 48 Pedicle Screw Based Artificial Facet with Wide Laminectomy The pedicle screw based artificial facet with wide laminectomy models use the destabilized model with artificial facets placed across L4-L5. The artificial facet systems used were the same as in the facetectomy models with pedicle screw based artificial facet designs described above. Figures 3.9, 10, and 11 show the 3mm (3mmL), 5mm (5mmL), and support (SupportL) artificial facet designs, respectively, in the wide laminectomy model. Figure 3.9: The pedicle screw based artificial facet design with a 3mm thick stem connecting the artificial facets to the pedicle screws in the destabilized model (3mmL). Figure 3.10: The 5mm wide laminectomy pedicle screw based design at L4-L5 (5mmL).

71 49 Figure 3.11: The wide laminectomy model with the Support pedicle screw based design at L4-L5 (SupportL). Universal Joint Artificial Facet A universal joint was modified to become a facet joint replacement system. Pedicle screws 6.5mm in diameter and 55mm in length were placed at L4 and L5. The pedicle screws were inserted two-thirds into the vertebral body through the pedicles. A standard universal joint was modified to rotate about the mid axis. The joint was tied to the pedicle screws across the L4-L5 motion segment (Figure 3.12). The model is further discussed in Appendix E.

72 50 Rotational Joint Pedicle Screws Bending Joint Figure 3.12: A universal joint was modified to act as an artificial facet joint. Rigid Screw & Rod System For comparison purposes, a rigid screw and rod system was modeled. Pedicle screws 6.5mm in diameter and 55mm in length meshed with tetrahedral elements were placed at L4 and L5, two-thirds into the vertebral body and tied to the bony pedicles and body. A rod 6.5mm in diameter was rigidly tied to the head of each pedicle screw, traversing the L4-L5 motion segment. Figure 3.13 shows the rigid system across the L4-L5 motion segment. A facetectomy was performed at L4-L5 removing the inferior L4 facet, L5 superior facet, and capsular ligament.

73 51 Pedicle Screws Rod Figure 3.13: A rigid screw and rod system across the L4-L5 motion segment. Finite Element Model Data Analysis Several biomechanical parameters were evaluated to determine the effectiveness of the artificial facet designs. Angular motions at all spinal levels were found for a 6 N-m bending moment in flexion, extension, bending, and rotation and a 400N follower load and 6 N-m bending moment in flexion and extension. Total facet loads at L3-L4, L4-L5, and L5-S1 were found for all models and facet load components were computed. Angular motions and facet loads were compared to the intact and destabilized models. Maximum von Mises stresses were also determined for each implant and all screws. Stresses were compared to the changes in bone stresses in the intact model and pedicle screw stresses in the rigid screw and rod system. In Vitro Artificial Facet Joint Study For further validation of the artificial facet finite element models and understanding of artificial facet design, a cadaveric study was completed on two specimens. The

74 52 preliminary study of one design of facet replacement systems, a pedicle screw based design most similar to the Support design described in the FE model section above, was performed on only two specimens. While the FE models have similar pedicle screw based designs, the exact cadaveric study implant was not modeled. Therefore, only similar trends were used to validate the FE models. Six cases were performed including intact without follower load (Intact NFL) and with follower load (Intact FL), destabilized by removing the L4-L5 facet joint without and with follower load (Destab NFL and Destab FL), and artificial facet implanted at L4-L5 without and with follower load (Implant NFL and Implant FL). A pure bending moment was applied up to 6 N-m in all loading modes including flexion/extension, left and right lateral bending, and left and right axial rotation. A complete discussion of the testing follows. Two fresh, frozen cadaver spines were obtained and tested for any deformities or disease. Figure 3.14 shows DXA scans taken to determine bone mineral density of the two specimens and also displaying severe scoliosis in one spine. The T-score for the bone density of the spines were -0.9 and The osteoporotic spine was used to test implant performance in a severe case. CT slices at 0.5mm were also obtained of each specimen to ensure there was little degeneration, fusion, or other deformities with 4mm Beekley Spots TM in the L4 and L5 transverse and spinous processes. A CT image showing the CT spots in the transverse processes is shown in Figure 3.15A while the CT spots in the spinous processes are shown by the CT scan image in Figure 3.15B. Both specimens were thawed to room temperature and cleaned of all soft tissue leaving the discs and ligaments intact. Transverse cuts were made at the L1/L2 disc and approximately S2 to

75 53 create a motion segment from L2 to S1. Each specimen was potted to create a loading frame and base for kinematics testing. S2 was secured into Bondo TM (a 2-part epoxy resin) using three wood screws and four bolts to anchor onto the loading frame. The L2 vertebra was potted with one quarter inch rod passing through the medial-lateral plane of the body and another rod passing through the anterior-posterior plane of the body. The rods were secured with Bondo TM around the vertebral body. The spines were refrozen until the day of testing. On the day of testing, the specimens were thawed for 10 hours to room temperature. Once completely thawed, the spine was firmly bolted into a customized spine simulator in the field of measurement of an OptoTrak active marker optical measuring system (Figure 3.16A). The OptoTrak system tracks light-emitting diodes rigidly attached to each vertebral body (Figure 3.16B). The positions of the diodes were transformed into angular rotations in reference to the base or S1. The rotations of the L4-L5 motion segment were plotted with respect to the applied moment. Pure moments were applied using a system of arms, pulleys, and weights to create quasistatic loading conditions. Moments of 1.5, 3.0, and 6.0 N-m were generated and applied in all six degrees of freedom (DOF) in flexion/extension (F/E), right and left lateral bending (RB,LB) and right and left axial rotation (RR,LR). To reduce the viscoelastic effects, the spines were maximally ranged in all directions before data collection began. The specimens were allowed to stabilize to minimize creep for 30 seconds between each loading condition.

76 54 Each specimen was tested in the intact condition (Intact), destabilized condition with removed L4-L5 facet joints (Destab) shown in Figure 3.17, and with artificial facet implants at the L4-L5 level (Implant). A sketch similar to the implants used in the study is shown in Figure In each testing condition, pure moments were applied and rotations measured in flexion, extension, bending, and rotation. A total follower load of 400N was applied at L2. Intact, destabilized, and implant conditions were tested in flexion and extension with the follower load. For simplification, all cases without follower load will be recognized as NFL while conditions with follower load will be FL. Figure 3.19 shows an intact specimen with all LED s attached and the instrumentation needed to apply a follower load of 400N to L2. Throughout the experiments, all specimens were kept moist with saline. The test with follower load in flexion was not completed at 6 N-m due to apparatus failure and great concern for specimen failure; therefore, no data was presented for flexion FL. Figure 3.14: DXA scan L1-L4 of two specimens used for the cadaveric study.

77 55 A B Figure 3.15: (A) CT slice of lumbar spine with Beekley CT spots in the transverse processes. (B) CT phantom image of the entire lumbar specimen displaying the CT spots in the L4 and L5 spinous processes. A B Figure 3.16: (A) OptoTrak measuring system. (B) Spine testing setup with LED s and tracking cameras.

78 56 Figure 3.17: Destabilized spine with removed L4-L5 facets circled. L4 Inferior Facet Implant Pedicle Screws L5 Superior Facet Implant Figure 3.18: A sketch of an implant placed at L4-L5 similar to the implant used in the cadaveric study.

79 57 Torsion Follower Load Flex/Ext x y Lateral Bending z Figure 3.19: Specimen displaying the six loading modes, the follower load application, LED s for motion tracking, and the global coordinate system. In Vitro Data Analysis Rotational displacement plots were generated for each loading step (1.5, 3.0, and 6.0 N- m) with and without the 400N follower load for the L4-L5 motion segment in all loading modes. For each condition, intact (Intact), destabilized (Destab), and with L4-L5 artificial facet implants (Implant), relative motion was plotted for the L4-L5 motion segment. In the plots, extension, left bending, and left rotation were negative while the paired moments, flexion, right bending, and right rotation, respectively, were shown as positive. Maximum angular motion for each motion segment was also plotted for all loading modes with and without follower load and testing conditions; intact, destabilized, and implanted.

80 58 Artificial Facet FE Model Validation The artificial facet pedicle screw based finite element models used in the study were validated using the in vitro study described above. Since the modeled implants were similar to the implant used in the cadaveric study, but not the same, only trends were analyzed. Also, a statistical analysis was not performed since only two specimens were used in the study. The intact cadaveric and FE model data was presented along with the implanted condition so that trends in motion would be clearly understood. The relative motion trends of L3-S1, L4-S1, L3-L4, L4-L5, and L5-S1 were validated by comparing the cadaveric and FE model data for intact and pedicle screw based artificial facet conditions. As shown in Figure 3.20A, B, and C, trends between cadaveric and FE models were similar without follower load. In flexion and extension the cadaveric implanted condition resulted in an increase in motion as compared to the cadaveric intact condition. The same trend occurred in the FE models in extension, but little change was found in flexion. In bending, the cadaveric study showed an increase in motion from intact to implanted condition while the FE models showed little change between the intact and implanted models. Motion increased in rotation in both the cadaveric study and FE study with implants as compared to intact conditions. Thus, the FE models resulted in similar trends as the cadaveric study, further validating the effectiveness of the finite element method in determining the biomechanical effect of artificial facet implants.

81 59 When a follower load of 400N was applied to the specimens and FE models, trends were in agreement in extension. Flexion data was unavailable for the implanted cadaveric condition, thus validation was not completed. Motion increased in extension in the cadaveric study for the implanted condition as compared to intact. The FE models have similar trends such that motion increased with implants at most levels (motion decreased slightly at L5-S1) as compared to intact. Therefore, the artificial facet FE models performed similarly to the cadaveric study with follower load, providing further validation that FE models predict motions in good agreement with in vitro studies. Angular Displacement (Degs) Cadaveric and FEM Comparison of Artificial Facet Models in Flexion/Extension at 6 N-m L3-S1 L4-S1 L3-L4 L4-L5 L5-S1 Intact Flex (Cadaveric) Implant Flex (Cadaveric) Intact Flex (FEM) 3mm Flex (FEM) 5mm Flex (FEM) Support Flex (FEM) Intact Ext (Cadaveric) Implant Ext (Cadaveric) Intact Ext (FEM) 3mm Ext (FEM) 5mm Ext (FEM) Support Ext (FEM) Spinal Segment A

82 60 Cadaveric and FEM Comparison of Artificial Facet Models in Bending at 6 N-m Angular Displacement (Degs) Intact (Cadaveric) Implant (Cadaveric) Intact (FEM) 3mm (FEM) 5mm (FEM) Support (FEM) 0.00 L3-S1 L4-S1 L3-L4 L4-L5 L5-S1 Spinal Segment B Cadaveric and FEM Comparison of Artificial Facet Models in Rotation at 6 N-m Angular Displacement (Degs) Intact (Cadaveric) Implant (Cadaveric) Intact (FEM) 3mm (FEM) 5mm (FEM) Support (FEM) L3-S1 L4-S1 L3-L4 L4-L5 L5-S1 Spinal Segment C Figure 3.20: Results for the in vitro study and artificial facet FE models are in agreement for general trends in motion at 6 N-m in (A) Flexion/Extension; (B) Bending; (C) Rotation.

83 61 Cadaveric and FEM Comparison of Artificial Facet Models in Extension with 400N FL and 6 N-m Moment L3-S1 L4-S1 L3-L4 L4-L5 L5-S Angular Displacement (Degs) Intact Ext (Cadaveric) Implant Ext (Cadaveric) Intact Ext (FEM) 3mm Ext (FEM) 5mm Ext (FEM) Support Ext (FEM) Spinal Segment Figure 3.21: Results for the in vitro study and artificial facet FE models are in agreement for general trends in motion with 400N follower load and 6 N-m applied moment in extension.

84 Chapter IV Results Chapter Overview Results from the finite element method and cadaveric study will be presented in the following chapter. The cadaveric study results will include motions for all conditions in extension, flexion, left and right bending, and left and right rotation loading modes for the implanted motion segment without and with follower load. The FE data will then be presented for the capping and pedicle screw based designs in facetectomy and wide laminectomy models. Relative motion, facet loads, and implant stresses will be given comparing the designs. Cadaveric Study Results Relative motions were determined using two cadaver L2-S1 lumbar segments. Pure moments were applied to the specimens from 0 to 6 N-m, with and without follower load of 400N. Three conditions were measured: intact, destabilized by removal of the L4-L5 facet joint and capsular ligaments, and implanted with pedicle screw based sliding articulating surface artificial facets across the L4-L5 level. The average relative motion of the L4-L5 segment for each condition was plotted as well as average maximum values at all segments. Data for left and right bending was averaged for loading mode bending and left and right rotation data was averaged for loading mode rotation. 62

85 63 Data at 6 N-m with follower load in flexion was not reported since the experiment was not conducted due to concern of specimen and apparatus failure. Motion increased for the intact, destabilized, and implanted cases as applied moment was increased. Figures 4.1A, B, and C show the motion at L4-L5 for each condition with no follower load (NFL). In all loading modes, destabilization decreased the stability of the lumbar spine. Motion increased compared to intact with the implant at L4-L5, but was less than the destabilized condition at moments less than 3 N-m. Upon application of 3 N-m, the implanted motion increased in all loading modes but flexion and right bending. However, motion increased in flexion at application of 6 N-m moment. The addition of a 400N follower load (FL) (Figure 4.2) decreased motion for the intact, destabilized, and implanted cases as compared to no follower load cases. With follower load, intact motion was greatest while motion decreased with the implant in extension. Motion was less than intact in flexion with the implant, but was greater than the destabilized case. Relative motions for all levels with respect to S1 at 6 N-m applied moment are shown in Figures 4.3A, B, C, and D. The destabilized and implanted conditions increased motion at all levels from intact in extension, except at L4-L5 where destabilized decreased 16.3% (implant increased 7.5%). The implant decreased motion at all levels when compared to destabilization except at L4-L5. In flexion, motion increased over intact with destabilization and implantation, at L4-L5 by 29.6% and 50.4% (destabilized and implant, respectively), and at all levels but L3-L4 where implantation decreased motion. All levels increased in relative motion in bending for both the destabilized and implanted

86 64 condition as compared to intact. The implant did increase instability from destabilization at all levels in bending except at L3-L4. At L4-L5, motion increased in bending 25.1% for destabilized and 21.0% for the implanted conditions. Rotation caused an increase in motion at all levels with destabilization and implantation from intact. The greatest increase in motion at L4-L5 occurred in rotation where destabilization increased 40.4% and the implant increased 64.1%. The implant resulted in greater flexibility at all levels compared to intact and destabilized, but decreased motion as compared to destabilized at L5-S1 in rotation. Percent changes in the relative motion of each level and case is given in Table 4.1. With an applied follower load of 400N and moment of 6 N-m, motion in extension decreased for the destabilized condition as compared to the intact case at all levels except L2-S1 and L4-L5 (motion increased 5.7%) shown in Figure 4.4A. Relative motion decreased with the implant in extension FL; 21.8% at L4-L5. In flexion (Figure 4.4B), destabilization decreased motion at all levels 60.5% at L4-L5. The implanted case was not presented due to testing time constraints and load application fixture failure. Percent changes in relative motion for all levels as compared to intact are given in Table 4.2.

87 65 Extension/Flexion at L4/5 for 6 N-m Rotational Angle (Degrees) Intact 8 Destab Implant Moment (N-m) A Left/Right Bending at L4/5 for 6 N-m Rotational Angle (Degrees) Intact Destab Implant Moment (N-m) B

88 66 Left/Right Rotation at L4/5 for 6 N-m Rotational Angle (Degrees) Intact 5 Destab 4 Implant Moment (N-m) C Figure 4.1: Results of average relative motion across the L4-L5 motion segment for the in vitro study for intact (Intact), destabilized without facets (Destab), and with artificial facets (Implant) for 0 to 6 N-m applied moment: (A) Extension/Flexion; (B) Left/Right Bending; (C) Left/Right Rotation. Rotational Angle (Degrees) Extension/Flexion at L4/5 for 6 N-m Intact NFL Destab NFL Implant NFL Intact FL Destab FL Implant FL Moment (N-m) Figure 4.2: Comparison of relative motion across L4-L5 with and without follower load for 0 to 6 N-m extension/flexion moment.

89 67 Relative Motion for 6 N-m Extension Angular Motion (Degrees) L2-S1 L3-S1 L4-S1 L3-L4 L4-L5 L5-S1 Intact Destab Implant A Relative Motion for 6 N-m Flexion Angular Motion (Degrees) L2-S1 L3-S1 L4-S1 L3-L4 L4-L5 L5-S1 Intact Destab Implant B

90 68 Relative Motion for 6 N-m Bending Angular Motion (Degrees) L2-S1 L3-S1 L4-S1 L3-L4 L4-L5 L5-S1 Intact Destab Implant C Relative Motion for 6 N-m Rotation Angular Motion (Degrees) L2-S1 L3-S1 L4-S1 L3-L4 L4-L5 L5-S1 Intact Destab Implant Figure 4.3: Relative motions at all levels of the lumbar spine for the in vitro study with 6 N-m applied moment: (A) Extension; (B) Flexion; (C) Bending; (D) Rotation. D

91 69 Relative Motion for 400N FL and 6 N-m Extension Angular Motion (Degrees) L2-S1 L3-S1 L4-S1 L3-L4 L4-L5 L5-S1 Intact Destab Implant A Relative Motion for 400N FL and 6 N-m Flexion 20.0 Angular Motion (Degrees) L2-S1 L3-S1 L4-S1 L3-L4 L4-L5 L5-S1 Intact Destab Figure 4.4: Relative motions at all levels of the lumbar spine for the in vitro study with 400N follower load and 6 N-m applied moment: (A) Extension; (B) Flexion. B

92 70 Table 4.1: Percent changes in relative motion at 6 N-m applied moment for the cadaveric study in each loading mode. A positive percent change indicates an increase in relative motion where as a negative percent change indicates a decrease in relative motion. % Change in Relative Motion in Cadaveric Study for 6 N-m Moment L2-S1 L3-S1 L4-S1 L3-L4 L4-L5 L5-S1 Intact Destab Extension Implant Flexion Bending Rotation Destab Destab Destab Implant Implant Implant Table 4.2: Percent changes in relative motion for 400N follower load and 6 N-m applied moment for the cadaveric study in extension and flexion. No data was available for the implant undergoing flexion with FL. A positive percent change indicates an increase in relative motion where as a negative percent change indicates a decrease in relative motion. % Change in Relative Motion in Cadaveric Study for 400N FL and 6 N-m Moment L2-S1 L3-S1 L4-S1 L3-L4 L4-L5 L5-S1 Intact Destab Extension Implant Flexion Destab Implant Finite Element Model Results Angular Motion Relative motions predicted by the artificial facet finite element models across all segments are presented and compared to intact, destabilized by wide laminectomy

93 71 condition, and rigid system. Two loading conditions were simulated for the FE models: 6 N-m bending moment shown in Figures 4.5A, B, C, and D and 400N follower load with 6 N-m bending moment shown in Figures 4.6A and B. Percent changes in relative motion as compared to intact for each artificial facet FE model are given in Table 4.3 for 6 N-m applied moment and Table 4.4 for 400N follower load and 6 N-m applied moment. Motions were also determined for the effectiveness of an artificial facet in stabilizing the spine with a wide laminectomy. The wide laminectomy model results are shown in Figures 4.7A, B, C, and D for 6 N-m bending moment and Figures 4.8A and B for 400N follower load and 6 N-m moment. Percent changes from intact in relative motion as a result of the wide laminectomy procedure are given in Table 4.5 for 6 N-m and Table 4.6 for 400N follower load and 6 N-m moment loading conditions. Extension, flexion, lateral bending, and axial rotation were simulated. The resultant motions are presented below for a rigid system, destabilized, and different artificial facet designs. Rigid System The rigid screw and rod system placed across the L4-L5 motion segment greatly decreased motion at L4-L5 in all loading modes with and without follower load. In extension and flexion, motion increased at the L3-L4 motion level with and without follower load, but was similar to intact in bending and rotation. The L3-S1 and L4-S1 segment motion decreased in all modes while L5-S1 segment was similar to intact motion. Artificial Facet Caps Artificial facet caps placed at L4-L5 increased motion in all loading modes and compression application at the L4-L5 segment by 23.5, 3.8, 1.0, and 17.1% with no follower load for extension, flexion, bending, and rotation, respectively, and 36.8% and

94 72 2.0% with follower load for extension and flexion. At all other levels in extension and rotation with and without follower load, motion increased, but was similar to intact in flexion and bending. Artificial Facet Caps with Screws When screws were used to further secure the artificial facet caps at L4-L5, angular motion results were similar to caps without screws. Relative motion of L4-L5 increased from intact by 23.8%/ 3.8% in extension/flexion NFL and 37.3%/2.0% in extension/flexion FL, and 1.0% and 17.4% in bending and rotation, respectively. All other motion segments experienced an increase in motion in extension and rotation, but were similar to intact in flexion and bending regardless of follower load application. 3mm Pedicle Screw Based Artificial Facet A pedicle screw based design with 3mm thick stems resulted in an increase in motion in extension and rotation, but decrease in flexion and bending at L4-L5. Motion increased from intact by 34.1% and 35.8% with no follower load in extension and rotation and 56.5% and 0.1% in extension and flexion with follower load. A decrease in motion at L4-L5 resulted from flexion NFL and bending moments of 0.2% and 2.9%, respectively. The L3-S1 and L4-S1 segments increased in motion in extension and rotation regardless of follower load. In extension with and without follower load and rotation, L3-L4 segment was similar to intact motion. L5-S1 motion segment decreased in relative motion in extension (NFL and FL), but was similar to intact in rotation. Flexion and bending resulted in motions similar to intact at all levels.

95 73 5mm Pedicle Screw Based Artificial Facet Increasing the stem thickness to 5mm on the pedicle screw based artificial facet did not reflect any changes in motion trends across all segments as compared to the 3mm design. Motion increased in extension NFL and FL by 33.0% and 56.4% and in rotation by 35.7% at L4-L5, while motion decreased by 0.2% and 2.9% in flexion and bending, respectively. Motions also increased in extension (NFL and FL) and rotation at the L3- S1 and L4-S1 segments. In all other loading modes and motion segments, the relative motion was very similar to intact except at L5-S1 in extension with and without follower load where motion decreased slightly. Pedicle Screw Based Artificial Facet with Pedicle Screw Support Motion resulting from an increase in the stem width around the pedicle screw and the implant being tied to the bony pedicle was closer to intact than other designs. At L4-L5, motion increased by 5.7% and 12.9% without follower load and 7.9% with follower load in extension and rotation, respectively and 0.1% in flexion FL. A decrease by 0.2% and 3.5% in flexion and bending occurred at L4-L5. Motion also increased in rotation at the L3-S1, L4-S1, and L5-S1 segments. L5-S1 decreased in motion in extension regardless of compression load application. All other levels in all loading modes were similar to intact. Destabilized Destabilization of the lumbar spine by performing a wide laminectomy at L4-L5 increased motion in all loading modes at L4-L5. Motion increased by 37.2% in extension, 96.6% in flexion, 0.7% in bending, 39.5% in rotation, and 59.7% and 85.2% in extension and flexion with follower load. All other levels also increased in motion in extension and

96 74 flexion regardless of load application, except at L5-S1 in extension NFL in which motion was similar to intact. A wide laminectomy did not affect motion in lateral bending at any level. Motion increased at levels L3-S1 and L4-S1 in rotation, but was similar to intact in rotation at L3-L4 and L5-S1. 3mm Pedicle Screw Based Artificial Facet with Wide Laminectomy When the 3mm artificial facet pedicle screw based design was placed in the destabilized spine (3mmL), motion increased at L4-L5 in all loading modes but bending in which motion was similar to intact. Motion increased by 35.2, 112.5, and 39.5%, in extension, flexion, and rotation, respectively with NFL, and 58.1% and 94.3% in extension and flexion with FL. L3-S1 and L4-S1 increased in motion as compared to intact for all loading modes irrespective of follower load. L3-L4 motion was similar to intact in extension (NFL and FL), bending, and rotation, but increased in flexion (NFL and FL). L5-S1 motion was similar to intact in all loading modes but decreased in motion slightly in extension with follower load. 5mm Pedicle Screw Based Artificial Facet with Wide Laminectomy The 5mm pedicle screw based design did not reflect any changes in motion trends across all segments as compared to the 3mm design. Motion increased in all loading modes but bending where a decrease in motion occurred. Motion increased by 34.2, 112.5, and 37.7% in extension, flexion, and rotation NFL, respectively. When a 400N follower load was applied, motion increased in extension and flexion by 56.3% and 94.3%. Motion decreased in bending by 3.4% at L4-L5.

97 75 Pedicle Screw Based Artificial Facet with Pedicle Screw Support with Wide Laminectomy The additional support in the final pedicle screw based design led to increased in motion at L4-L5 in all loading modes but bending. Motion increased by 8.0, 112.5, and 16.5% in extension, flexion, and rotation without follower load and 11.3% and 94.3% in extension and flexion with follower load. However, motion decreased in bending by 58.1%. At all other levels, motion was similar to intact in extension with and without follower load except at L5-S1 where motion decreased with FL. Flexion increased motion at all levels NFL and FL except at L5-S1 with FL where motion was similar to intact. Rotation resulted in increased in motion at L3-S1 and L4-S1, but was similar to intact at L3-L4 and L5-S1. Motion decreased in bending at all levels. Relative Motion for 6 N-m Extension Angular Motion (Degrees) L3-S1 L4-S1 L3-L4 L4-L5 L5-S1 Intact Rigid Cap Cap Screw mm mm Support A

98 76 Relative Motion for 6 N-m Flexion 12.0 Angular Motion (Degrees) L3-S1 L4-S1 L3-L4 L4-L5 L5-S1 Intact Rigid Cap Cap Screw mm mm Support B Relative Motion for 6 N-m Bending Angular Motion (Degrees) L3-S1 L4-S1 L3-L4 L4-L5 L5-S1 Intact Rigid Cap Cap Screw mm mm Support C

99 77 Relative Motion for 6 N-m Rotation 8.0 Angular Motion (Degrees) L3-S1 L4-S1 L3-L4 L4-L5 L5-S1 Intact Rigid Cap Cap Screw mm mm Support Figure 4.5: Relative motions at all levels of the lumbar spine for the FE study with 6 N- m applied moment: (A) Extension; (B) Flexion; (C) Bending; (D) Rotation. D

100 78 Relative Motion for 400N FL and 6 N-m Extension 7.0 Angular Motion (Degrees) L3-S1 L4-S1 L3-L4 L4-L5 L5-S1 Intact Rigid Cap Cap Screw mm mm Support A Relative Motion for 400N FL and 6 N-m Flexion 14.0 Angular Motion (Degrees) L3-S1 L4-S1 L3-L4 L4-L5 L5-S1 Intact Rigid Cap Cap Screw mm mm Support B Figure 4.6: Relative motions at all levels of the lumbar spine for the FE study with 400N follower load and 6 N-m applied moment: (A) Extension; (B) Flexion.

101 79 % Change in Relative Motion for FE Models Compared to Intact for 6 N-m Moment L3-S1 L4-S1 L3-L4 L4-L5 L5-S1 Intact Rigid Cap Cap Screw Extension 3mm mm Support Rigid Cap Flexion Cap Screw mm mm Support Rigid Cap Bending Cap Screw mm mm Support Rigid Cap Rotation Cap Screw mm mm Support Table 4.3: Percent changes in relative motion at 6 N-m applied moment for the FE models in each loading mode. A positive percent change indicates an increase in relative motion where as a negative percent change indicates a decrease in relative motion.

102 80 % Change in Relative Motion for FE Models Compared to Intact for 400N and 6 N-m Moment L3-S1 L4-S1 L3-L4 L4-L5 L5-S1 Intact Rigid Cap Extension Cap Screw mm mm Support Rigid Cap Flexion Cap Screw mm mm Support Table 4.4: Percent changes in relative motion with 400N follower load and 6 N-m applied moment for the FE models in each loading mode. A positive percent change indicates an increase in relative motion where as a negative percent change indicates a decrease in relative motion. Relative Motion for 6 N-m Extension Angular Motion (Degrees) L3-S1 L4-S1 L3-L4 L4-L5 L5-S1 Intact DestabL Rigid mmL mmL SupportL A

103 81 Relative Motion for 6 N-m Flexion Angular Motion (Degrees) L3-S1 L4-S1 L3-L4 L4-L5 L5-S1 Intact DestabL Rigid mmL mmL SupportL B Relative Motion for 6 N-m Bending Angular Motion (Degrees) L3-S1 L4-S1 L3-L4 L4-L5 L5-S1 Intact DestabL Rigid mmL mmL SupportL C

104 82 Relative Motion for 6 N-m Rotation 7.0 Angular Motion (Degrees) L3-S1 L4-S1 L3-L4 L4-L5 L5-S1 Intact DestabL Rigid mmL mmL SupportL D Figure 4.7: Relative motions at all levels of the lumbar spine with wide laminectomy for the FE study with 6 N-m applied moment: (A) Extension; (B) Flexion; (C) Bending; (D) Rotation.

105 83 Relative Motion for 400N FL and 6 N-m Extension 7.0 Angular Motion (Degrees) L3-S1 L4-S1 L3-L4 L4-L5 L5-S1 Intact Rigid DestabL mmL mmL SupportL A Relative Motion for 400N FL and 6 N-m Flexion Angular Motion (Degrees) L3-S1 L4-S1 L3-L4 L4-L5 L5-S1 Intact Rigid DestabL mmL mmL SupportL B Figure 4.8: All levels relative motion of the wide laminectomy lumbar spine for the FE study with 400N follower load and 6 N-m applied moment: (A) Extension; (B) Flexion.

106 84 Table 4.5: Percent changes in relative motion at 6 N-m applied moment for the wide laminectomy FE models in each loading mode. A positive percent change indicates an increase in relative motion where as a negative percent change indicates a decrease in relative motion. % Change in Relative Motion for FE Models Compared to Intact for 6 N-m Moment L3-S1 L4-S1 L3-L4 L4-L5 L5-S1 Intact DestabL Rigid Extension 3mmL mmL SupportL DestabL Rigid Flexion 3mmL mmL SupportL DestabL Rigid Bending 3mmL mmL SupportL DestabL Rigid Rotation 3mmL mmL SupportL

107 85 % Change in Relative Motion for FE Models Compared to Intact for 400N FL and 6 N-m Moment L3-S1 L4-S1 L3-L4 L4-L5 L5-S1 Intact DestabL Rigid Extension 3mmL mmL SupportL DestabL Rigid Flexion 3mmL mmL SupportL Table 4.6: Percent changes in relative motion with 400N follower load and 6 N-m applied moment for the wide laminectomy FE models in each loading mode. A positive percent change indicates an increase in relative motion where as a negative percent change indicates a decrease in relative motion. Facet Loads Total facet loads were determined for each model at 6 N-m moment and compared to the intact case. At L4-L5, loads increased up to 14% and maximally decreased by 26% in rotation and extension, respectively. Figures 4.9A, B, and C show total facet loads compared to intact for all artificial facet designs and include a table of data values at 6 N- m bending moments of extension, bending, and rotation. Flexion facet loads are equal to zero at all levels with no applied follower load, thus flexion data is not included. Facet loads for 6 N-m extension and flexion moments with 400N follower load are shown in Figures 4.10A and B along with data values. Tabular data of percent changes in facet loads as compared to intact for each artificial facet model are shown in Table 4.7 for 6 N- m moment and Table N follower load and 6 N-m moment. Total facet loads for the destabilized models are also included and shown in Figures 4.11A, B, C, and D for 6

108 86 N-m bending moment and Figures 1.12A and B for 400N follower load and 6 N-m bending moment. Percent change in facet loads for artificial facets in the wide laminectomy model as compared to intact are included in Table 4.9 for 6 N-m moment and Table 4.10 for 400N follower load and 6 N-m applied moment. Facet load components for vertical, lateral, and shear loads were calculated and are included in Appendix F for further reference. Artificial Facet Caps Facet loads decrease in extension and bending loading modes with artificial facet caps at the L4-L5 motion segment by 13.8% and 12.1% with no follower load and 14.8% with follower load in extension. However, facet loads increased as compared to intact in rotation by 14.2%. At the L3-L4 and L5-S1 levels, facet loads decreased in extension, bending, and rotation without follower load as well as with follower load in extension and flexion (L5-S1 only). Artificial Facet Caps with Screws When screws were added to further secure the caps, facet loads were similar to loads with caps only. In extension and bending, facet loads across L4-L5 decreased by 14.0% and 12.6% with no follower load and 14.8% with follower load in extension. Rotation resulted in increased facet loads by 14.0% at L4-L5. At other levels, facet loads decreased in extension, flexion, bending, and rotation regardless of follower load application. 3mm Pedicle Screw Based Artificial Facet A 3mm stem pedicle screw based design decreased facet loads in all loading modes regardless of follower load application by 25.7, 25.1, 1.1, and 28.0%, for extension,

109 87 bending, rotation NFL and extension FL at L4-L5. Facet loads across L3-L4 in extension and bending increased with no follower load as well as with follower load in extension, but was similar to intact in rotation. In all motions, facet loads were similar to intact at L5-S1 but in flexion with follower load where facet loads decreased. 5mm Pedicle Screw Based Artificial Facet Increasing the thickness of the facet replacement stem to 5mm did not result in changes in the facet load transfer trends from the 3mm design. Facet loads at L4-L5 decreased in extension NFL by 24.6%, 23.9% in bending, and 0.9% in rotation and 27.4% in extension with follower load. All other trends were similar to the 3mm thick stem design. Pedicle Screw Based Artificial Facet with Pedicle Screw Support Facet loads changed from intact less than other designs when the stem around the pedicle screw was increased in width and the implant was tied to the bony pedicle. Loads decreased by 2.7% in extension, 4.9% in bending, and 5.2% with extension FL, but increased in rotation by 13.8%. Facet loads also increased at L3-L4 in extension and bending and rotation despite of follower load application. Flexion FL resulted in a decrease in facet loads across L5-S1 while extension, bending, and rotation facet loads were similar to intact. Destabilized By performing a wide laminectomy at L4-L5, but leaving the L4-L5 facets intact, facet loads decreased in all loading modes regardless of follower load application. Loads decreased by 20.3, 25.2, and 22.0% in extension, bending, and rotation, respectively at L4-L5 with no follower load. With follower load, facet loads decreased by 11.1% in

110 88 extension at L4-L5. At L3-L4 and L5-S1, facet loads were similar to intact, but increased at L5-S1 in flexion with follower load. 3mm Pedicle Screw Based Artificial Facet with Wide Laminectomy By placing a 3mm stem artificial facet pedicle screw based design across the L4-L5 motion segment with a wide laminectomy, facet loads decreased at L4-L5. Facet loads decreased by 26.5, 25.5, and 2.4% in extension, bending, and rotation NFL, respectively, and 28.8% in extension FL. At L3-L4, facet loads increased in extension (NFL and FL) and bending, but were similar to intact in rotation. At L5-S1, facet loads were similar to intact for all loading modes, but decreased in flexion with follower load. 5mm Pedicle Screw Based Artificial Facet with Wide Laminectomy Increasing the stem thickness to 5mm from 3mm did not change the facet loads across L4-L5 greatly. Facet loads decreased by 25.5% in extension NFL, 24.4% in bending, 2.3% in rotation, and 28.1% in extension with FL. All other levels exhibited the same results as the 3mm design. Pedicle Screw Based Artificial Facet with Pedicle Screw Support with Wide Laminectomy Facet loads at L4-L5 decreased with the support design with a wide laminectomy for all loading modes but rotation. Facet loads decreased by 4.4% and 6.1% in extension and bending NFL and 6.9% in extension with FL. An increase in facet loads at L4-L5 of 11.7% was found in rotation. L3-L4 facet loads increased in extension, with and without follower load, and bending but were similar to intact in rotation. L5-S1 facet loads were similar to intact in all loading modes, but decreased in flexion with follower load.

111 Facet Loads for 6 N-m Extension Facet Load (N) L3/L4 L4/L5 L5/S1 Intact Cap Cap Screw mm mm Support A 50.0 Facet Loads for 6 N-m Bending Facet Load (N) L3/L4 L4/L5 L5/S1 Intact Cap Cap Screw mm mm Support B

112 Facet Loads for 6 N-m Rotation Facet Load (N) L3/L4 L4/L5 L5/S1 Intact Cap Cap Screw mm mm Support C Figure 4.9: Facet loads at all levels of the lumbar spine for the FE study with 6 N-m applied moment: (A) Extension; (B) Bending; (C) Rotation Facet Loads for 400N FL and 6 N-m Extension Facet Load (N) L3/L4 L4/L5 L5/S1 Intact Cap Cap Screw mm mm Support A

113 Facet Loads for 400N FL and 6 N-m Flexion Facet Load (N) L3/L4 L4/L5 L5/S1 Intact Cap Cap Screw mm mm Support B Figure 4.10: Facet loads at all levels of the lumbar spine for the FE study with 400N follower load and 6 N-m applied moment: (A) Extension; (B) Flexion.

114 92 Table 4.7: Percent changes in total facet loads for 6 N-m applied moment for the artificial facet FE models in each loading mode. A positive percent change indicates an increase in facet loads where as a negative percent change indicates a decrease in facet loads. % Change in Total Facet Loads for FE Models Compared to Intact for 6 N-m Moment L3/L4 L4/L5 L5/S1 Intact Cap Cap Screw Extension 3mm mm Support Cap Cap Screw Bending 3mm mm Support Cap Cap Screw Rotation 3mm mm Support Table 4.8: Percent changes in total facet loads for 400N follower load and 6 N-m applied moment for the artificial facet FE models in each loading mode. A positive percent change indicates an increase in facet loads where as a negative percent change indicates a decrease in facet loads. % Change in Total Facet Loads for FE Models Compared to Intact for 400N FL and 6 N-m Moment L3/L4 L4/L5 L5/S1 Intact Cap Cap Screw Extension 3mm mm Support Cap Cap Screw Flexion 3mm mm Support

115 93 Facet Loads for 6 N-m Extension Facet Load (N) L3/L4 L4/L5 L5/S1 Intact DestabL mmL mmL SupportL A Facet Loads for 6 N-m Bending Facet Load (N) L3/L4 L4/L5 L5/S1 Intact DestabL mmL mmL SupportL B

116 94 Facet Loads for 6 N-m Rotation Facet Load (N) L3/L4 L4/L5 L5/S1 Intact DestabL mmL mmL SupportL C Figure 4.11: Facet loads at all levels of the wide laminectomy lumbar spine for the FE study with 6 N-m applied moment: (A) Extension; (B) Bending; (C) Rotation. Facet Loads for 400N FL and 6 N-m Extension Facet Load (N) L3/L4 L4/L5 L5/S1 Intact DestabL mmL mmL SupportL A

117 95 Facet Loads for 400N FL and 6 N-m Flexion Facet Load (N) L3/L4 L4/L5 L5/S1 Intact DestabL mmL mmL SupportL B Figure 4.12: Facet loads at all levels of the wide laminectomy lumbar spine for the FE study with 400N follower load and 6 N-m applied moment: (A) Extension; (B) Flexion.

118 96 Table 4.9: Percent changes in total facet loads for 6 N-m applied moment for the artificial facet FE models with wide laminectomy in each loading mode. A positive percent change indicates an increase in facet loads where as a negative percent change indicates a decrease in facet loads. % Change in Total Facet Loads for FE Models Compared to Intact for 6 N-m Moment L3/L4 L4/L5 L5/S1 Intact DestabL Extension 3mmL mmL SupportL DestabL Bending 3mmL mmL SupportL DestabL Rotation 3mmL mmL SupportL % Change in Total Facet Loads for FE Models Compared to Intact for 400N FL and 6 N-m Moment L3/L4 L4/L5 L5/S1 Intact DestabL Extension 3mmL mmL SupportL DestabL Flexion 3mmL mmL SupportL Table 4.10: Percent changes in total facet loads with 400N follower load and 6 N-m applied moment for the artificial facet wide laminectomy FE models in each loading mode. A positive percent change indicates an increase in facet loads where as a negative percent change indicates a decrease in facet loads.

119 97 Peak von Mises Stresses Peak von Mises stresses were determined for each implant and screw used in the designs. The stress values were compared to the tensile and yield strength of titanium to ensure of implant failure. The yield strength of titanium with a Modulus of Elasticity of 110GPa and Poisson s ratio of 0.30 is 760MPa and the tensile strength is 790MPa [60]. The tensile strength of titanium may be directly compared to the von Mises stresses in the implants. A direct comparison can be completed since von Mises stress is used to estimate yield criteria for ductile materials and is calculated by combining principle stresses in all three directions. Stress plots for each implant is shown and tables of the peak von Mises stress found in each component is included for the models. Artificial Facet Caps Maximum stresses in the inferior facet cap occurred in extension NFL of 45.4MPa and FL 39.5MPa. The superior cap experienced high stresses in rotation NFL of 98.3MPa and 32.5MPa in extension FL. A maximum factor of safety of 7.7 was determined for the cap design since high stresses were experienced in rotation. Stress plots of the caps are shown in Figure 4.13 with no follower load and 4.14 with follower load. All maximum stresses are given in Tables 4.11, 12, and 13 for all loading modes. Artificial Facet Caps with Screws The inferior facet cap (with screws) experienced the largest von Mises stresses of 89.0MPa and 63.6MPa in extension NFL and FL, respectively. The superior facet cap had a maximum stress of 80.6MPa and 51.0MPa in extension NFL and with follower load. The maximum stress in the inferior cap yielded a factor of safety of 8.5. Tables 4.11, 12, and 13 compare the peak stresses of both capping designs. Maximum stress

120 98 values for each screw used to secure the caps are given in Tables 4.14 and 15. A von Mises stress of 31.5MPa was found in extension NFL and FL in the inferior facet cap screw. The most anteriorly placed superior cap screw (Superior Screw A) in the spine model had a peak stress of 64.3MPa in rotation and 48.1MPa with follower load in extension while the most posteriorly positioned superior facet cap screw (Superior Screw P) only experienced 18.5MPa in rotation and 14.9MPa in extension with follower load. Since screws were inserted into the L5 bony pedicle, stresses in the pedicle bone were also determined. Tables 4.16 and 17 summarize the stresses found in the pedicles for intact and the capping with screws design. A maximum stress in intact occurred in rotation NF of 14.8MPa and in extension FL of 17.5MPa. When screws were inserted into the pedicle, stresses increased greatest in extension NFL to 32.9MPa and 33.6MPa in extension with follower load. Stress contours of the caps and screws are shown in Figures 4.15 and 16. Rigid Screw and Rod System A rigid system model simulating a fusion surgery was used to determine the maximum von Mises stresses in pedicle screws frequently found in surgical use. Thus, the rigid system stresses could be directly compared to the stresses in the pedicle screws used in the artificial facet designs. In this way, the rigid system was used as a gold standard. The stresses in the L4 and L5 pedicle screws are shown in Tables 4.21, 22, and 23. A maximum von Mises stress in the L4 rigid screw system pedicle screw occurred in rotation NFL of 69.0MPa and 58.1MPa in extension FL. In the L5 pedicle screw, a peak stress of 59.5MPa was found in rotation and 65.5MPa in flexion FL.

121 99 3mm Pedicle Screw Based Artificial Facet Stresses in the artificial facet joint increased from the capping design greatly. A maximum stress was located in the inferior facet component of 144.5MPa in rotation and 146.9MPa in extension with follower load shown in stress contours in Figures 4.17 and 18. The superior facet component stresses were less at 69.1MPa in rotation and 36.1MPa in extension with follower load. The factor of safety of the design was 5.2 due to high stresses in the inferior facet component. The maximum pedicle screw stresses were less than the rigid system in all loading modes, except extension FL for L4 where stress increased. In rotation maximum von Mises stress was at 63.6MPa and 55.8MPa for L4 and L5 pedicle screws, respectively. The maximum L5 facet pedicle screw stress was only 16.9MPa in extension FL, but the L4 pedicle screw maximum stress was 74.7MPa (28.7% increase from rigid) in extension with follower load. Tables 4.18 through 4.23 give further information on maximum stresses in the implant and pedicle screws. 5mm Pedicle Screw Based Artificial Facet The peak von Mises stress in the 5mm inferior facet component was 149.4MPa in extension NFL and 156.4MPa in extension FL. The superior facet component stresses were less at 133.6MPa in rotation NFL and 104.7MPa in extension with follower load. The factor of safety decreased to 4.9 when a 5mm stem was present on the inferior component. Maximum pedicle screw stresses at L4 were 60.8MPa and 76.3MPa, a 31.3% increase from rigid, in rotation and extension FL, respectively. The L5 pedicle screw stresses were less at 45.0MPa in rotation and in extension FL at 13.0MPa. Stress plots of the pedicle screws and artificial facets are shown in Figures 4.19 and 20 and Tables 4.18 to 23 have maximum stress values in the implants and pedicle screws.

122 100 Pedicle Screw Based Artificial Facet with Pedicle Screw Support Peak von Mises stresses in the inferior facet component were 179.8MP and 267.0MPa due to motion in extension NFL and FL, respectively, for the Support design. The superior component experienced less stress at 139.1MPa with rotation and 62.4MPa in extension FL. Due to the very high stress in the inferior component with follower load, the factor of safety was only 2.8. While stresses were high in the components, stress decreased greatly in the pedicle screws. The L4 pedicle screw experienced only 6.8MPa in rotation and 7.3MPa in extension with follower load, maximally. The L5 pedicle screw had a maximum stress of 6.1MPa in extension NFL and 7.3MPa in flexion with follower load. Figures 4.21 and 22 show the stress contours of the screws and implants and Tables 4.18 to 23 give peak stress values for the implants and pedicle screws. 3mm Pedicle Screw Based Artificial Facet with Wide Laminectomy Stresses in the 3mm artificial facet joint increased when the spine was destabilized. A maximum stress was located in the inferior facet component of 171.8MPa in rotation and 196.5MPa in extension with follower load shown in the stress contours in Figures 4.23 and 24. The superior facet component maximum stresses were much less at 86.6MPa in rotation and 48.6MPa in extension with follower load. The factor of safety of the design decreased due to the destabilization. The factor of safety was 5.2, but decreased to 3.9 due to higher stresses in the inferior component. The pedicle screw stresses were less than the rigid system in all loading modes but extension (NFL and FL). In extension NFL, maximum von Mises stress in the L4 pedicle screw was 63.1MPa (22.0% increase from rigid) and with follower load stress increased to 74.3MPa (28.0% increase from rigid). Stress in the L5 pedicle screw with the 3mm design was less than the rigid system

123 101 stresses for all loading modes. Maximum stress occurred in rotation (56.0MPa) and in extension FL (19.4MPa). Tables 4.24 through 4.29 give information on maximum stresses in the implant and pedicle screws in the wide laminectomy models. 5mm Pedicle Screw Based Artificial Facet with Wide Laminectomy The peak von Mises stress in the 5mm inferior facet component used to stabilize the wide laminectomy was 185.8MPa in extension NFL and 235.9MPa in extension FL. The superior facet component maximum stresses were less than the inferior component at 159.5MPa in rotation NFL and 104.0MPa in extension with follower load. The factor of safety decreased from the 3mm design further to 3.2 due to stresses induced by extension with follower load in the inferior facet component. Pedicle screw stresses resulting from the 5mm artificial facet design were less than the rigid system in all loading modes but extension NFL and FL. Pedicle screw stresses at L4 were 63.7MPa, an increase from rigid by 23.2% and 65.2MPa an increase of 12.2% from rigid in extension NFL and FL, respectively. While the L5 pedicle screw stresses were less at 56.1MPa in rotation and in extension FL at 20.7MPa. Stress plots of the pedicle screws and artificial facets are shown in Figures 4.25 and 26 and Tables 4.24 to 29 have maximum stress values in the implants and pedicle screws. Pedicle Screw Based Artificial Facet with Pedicle Screw Support with Wide Laminectomy Peak von Mises stresses in the inferior facet component were 223.3MP and 259.3MPa due to motion in extension NFL and FL, respectively with the Support design across L4- L5. The superior component experienced less maximum stress at 145.3MPa with rotation and 62.4MPa in extension FL. Due to the very high stress in the inferior component with

124 102 follower load, the factor of safety was only 2.9, slightly greater than previously reported. Since stress was high in the components, the pedicle screws were greatly unloaded. The L4 pedicle screw experienced 10.0MPa in extension NFL and 13.2MPa in extension with follower load. The L5 pedicle screw showed stresses of 11.1MPa in extension NFL and 13.2MPa in extension with follower load. Figures 4.27 and 28 show the stress contours of the screws and implants and Tables 4.24 to 29 give peak stress values for the implants and pedicle screws. 0N + 6 N-m Ext 0N + 6 N-m Flex 0N + 6 N-m RB 0N + 6 N-m RR Figure 4.13: Stress plots for the inferior and superior artificial facet caps at L4-L5 with 6 N-m bending moment in all loading modes.

125 N + 6 N-m Ext 400N + 6 N-m Flex Figure 4.14: Stress plots for the inferior and superior artificial facet caps at L4-L5 with 400N follower load and 6 N-m bending moment in all loading modes. 0N + 6 N-m Ext 0N + 6 N-m Flex 0N + 6 N-m RB 0N + 6 N-m RR Figure 4.15: Stress plots for the inferior and superior artificial facet caps with screws at L4-L5 with 6 N-m bending moment in all loading modes.

126 N + 6 N-m Ext 400N + 6 N-m Flex Figure 4.16: Stress plots for the inferior and superior artificial facet caps with screws at L4-L5 with 400N follower load and 6 N-m bending moment in all loading modes. Table 4.11: Peak von Mises stress (MPa) in inferior artificial facet cap designs for 6 N- m applied moment. Peak von Mises Stress (MPa) in Inferior Cap 6 N-m Moment Extension Flexion Bending Rotation Cap Cap Screw Table 4.12: Peak von Mises stress (MPa) in superior artificial facet cap designs for 6 N- m applied moment. Peak von Mises Stress (MPa) in Superior Cap 6 N-m Moment Extension Flexion Bending Rotation Cap Cap Screw Peak von Mises Stress (MPa) 400N Compression & 6 N-m Moment Inferior Cap Superior Cap Extension Flexion Extension Flexion Cap Cap Screw Table 4.13: Peak von Mises stress (MPa) in inferior and superior artificial facet cap designs for 6 N-m moment and 400N follower load.

127 105 Peak von Mises Stress (MPa) in Cap Screws 6 N-m Moment Extension Flexion Bending Rotation Inferior Screw Superior Screw A Superior Screw P Table 4.14: Peak von Mises stress (MPa) in inferior and superior artificial facet cap screws for 6 N-m applied moment. Peak von Mises Stress (MPa) in Cap Screws 400N Compression & 6 N-m Moment Extension Flexion Inferior Screw Superior Screw A Superior Screw P Table 4.15: Peak von Mises stress (MPa) in inferior and superior artificial facet cap screws for 400N follower load and 6 N-m applied moment. Table 4.16: Peak von Mises stress (MPa) in bony pedicle of L5 for intact and artificial facet cap cases for 6 N-m moment. Peak von Mises Stress (MPa) in L5 Pedicle 6 N-m Moment Extension Flexion Bending Rotation Intact Cap Screw Peak von Mises Stress (MPa) in L5 Pedicle 400N Compression & 6 N-m Moment Extension Flexion Intact Cap Screw Table 4.17: Peak von Mises stress (MPa) in bony pedicle of L5 for intact and artificial facet cap cases for 6 N-m moment and 400N follower load.

128 106 0N + 6 N-m Ext 0N + 6 N-m Flex 0N + 6 N-m RB 0N + 6 N-m RR Figure 4.17: Stress plots for the 3mm stem pedicle screw based artificial facets at L4-L5 with 6 N-m bending moment in all loading modes. 400N + 6 N-m Ext 400N + 6 N-m Flex Figure 4.18: Stress plots for the 3mm stem pedicle screw based artificial facets at L4-L5 with 400N follower load and 6 N-m bending moment in all loading modes.

129 107 0N + 6 N-m Ext 0N + 6 N-m Flex 0N + 6 N-m RB 0N + 6 N-m RR Figure 4.19: Stress plots for the 5mm stem pedicle screw based artificial facets at L4-L5 with 6 N-m bending moment in all loading modes. 400N + 6 N-m Ext 400N + 6 N-m Flex Figure 4.20: Stress plots for the 5mm stem pedicle screw based artificial facets at L4-L5 with 400N follower load and 6 N-m bending moment in all loading modes.

130 108 0N + 6 N-m Ext 0N + 6 N-m Flex 0N + 6 N-m RB 0N + 6 N-m RR Figure 4.21: Stress plots for the supported stem pedicle screw based artificial facets at L4-L5 with 6 N-m bending moment in all loading modes. 400N + 6 N-m Ext 400N + 6 N-m Flex Figure 4.22: Stress plots for the supported stem pedicle screw based artificial facets at L4-L5 with 400N follower load and 6 N-m bending moment in all loading modes.

131 109 Table 4.18: Peak von Mises stress (MPa) in inferior pedicle screw based facet designs for 6 N-m applied moment. Peak von Mises Stress (MPa) in L4 Inferior Facet 6 N-m Moment Extension Flexion Bending Rotation 3mm mm Support Table 4.19: Peak von Mises stress (MPa) in superior pedicle screw based facet designs for 6 N-m applied moment. Peak von Mises Stress (MPa) in L5 Superior Facet 6 N-m Moment Extension Flexion Bending Rotation 3mm mm Support Peak von Mises Stress (MPa) 400N Compression & 6 N-m Moment Inferior Facet Superior Facet Extension Flexion Extension Flexion 3mm mm Support Table 4.20: Peak von Mises stress (MPa) in inferior and superior pedicle screw based facet designs for 6 N-m moment and 400N follower load. Peak von Mises Stress (MPa) in L4 Pedicle Screw 6 N-m Moment Extension Flexion Bending Rotation Rigid mm mm Support Table 4.21: Peak von Mises stress (MPa) in L4 pedicle screw for facet designs and a rigid system with 6 N-m applied moment.

132 110 Peak von Mises Stress (MPa) in L5 Pedicle Screw 6 N-m Moment Extension Flexion Bending Rotation Rigid mm mm Support Table 4.22: Peak von Mises stress (MPa) in L5 pedicle screw for facet designs and a rigid system with 6 N-m applied moment. Peak von Mises Stress (MPa) 400N Compression & 6 N-m Moment L4 Pedicle Screw L5 Pedicle Screw Extension Flexion Extension Flexion Rigid mm mm Support Table 4.23: Peak von Mises stress (MPa) in L4 and L5 pedicle screw for facet designs and a rigid system with 400N follower load and 6 N-m applied moment.

133 111 0N + 6 N-m Ext 0N + 6 N-m Flex 0N + 6 N-m RB 0N + 6 N-m RR Figure 4.23: Stress plots for the 3mm stem pedicle screw based artificial facets at L4-L5 across a wide laminectomy with 6 N-m bending moment in all loading modes. 400N + 6 N-m Ext 400N + 6 N-m Flex Figure 4.24: Stress plots for the 3mm stem pedicle screw based artificial facets at L4-L5 across a wide laminectomy with 400N follower load and 6 N-m bending moment in all loading modes.

134 112 0N + 6 N-m Ext 0N + 6 N-m Flex 0N + 6 N-m RB 0N + 6 N-m RR Figure 4.25: Stress plots for the 5mm stem pedicle screw based artificial facets at L4-L5 across a wide laminectomy with 6 N-m bending moment in all loading modes. 400N + 6 N-m Ext 400N + 6 N-m Flex Figure 4.26: Stress plots for the 5mm stem pedicle screw based artificial facets at L4-L5 across a wide laminectomy with 400N follower load and 6 N-m bending moment in all loading modes.

135 113 0N + 6 N-m Ext 0N + 6 N-m Flex 0N + 6 N-m RB 0N + 6 N-m RR Figure 4.27: Stress plots for the supported stem pedicle screw based artificial facets at L4-L5 across a wide laminectomy with 6 N-m bending moment in all loading modes. 400N + 6 N-m Ext 400N + 6 N-m Flex Figure 4.28: Stress plots for the supported stem pedicle screw based artificial facets at L4-L5 across a wide laminectomy with 400N follower load and 6 N-m bending moment in all loading modes.

136 114 Peak von Mises Stress (MPa) in L4 Inferior Facet 6 N-m Moment Extension Flexion Bending Rotation 3mmL mmL SupportL Table 4.24: Peak von Mises stress (MPa) in inferior pedicle screw based facet designs with wide laminectomy for 6 N-m applied moment. Peak von Mises Stress (MPa) in L5 Superior Facet 6 N-m Moment Extension Flexion Bending Rotation 3mmL mmL SupportL Table 4.25: Peak von Mises stress (MPa) in superior pedicle screw based facet designs with wide laminectomy for 6 N-m applied moment. Peak von Mises Stress (MPa) 400N Compression & 6 N-m Moment Inferior Facet Superior Facet Extension Flexion Extension Flexion 3mmL mmL SupportL Table 4.26: Peak von Mises stress (MPa) in inferior and superior pedicle screw based facet designs with wide laminectomy for 6 N-m moment and 400N follower load. Peak von Mises Stress (MPa) in L4 Pedicle Screw 6 N-m Moment Extension Flexion Bending Rotation Rigid mmL mmL SupportL Table 4.27: Peak von Mises stress (MPa) in L4 pedicle screw for facet designs with wide laminectomy and a rigid system with 6 N-m applied moment.

137 115 Peak von Mises Stress (MPa) in L5 Pedicle Screw 6 N-m Moment Extension Flexion Bending Rotation Rigid mmL mmL SupportL Table 4.28: Peak von Mises stress (MPa) in L5 pedicle screw for facet designs with wide laminectomy and a rigid system with 6 N-m applied moment. Table 4.29: Peak von Mises stress (MPa) in L4 and L5 pedicle screw for facet designs with wide laminectomy and a rigid system with 400N follower load and 6 N-m applied moment. Peak von Mises Stress (MPa) 400N Compression & 6 N-m Moment L4 Pedicle Screw L5 Pedicle Screw Extension Flexion Extension Flexion Rigid mmL mmL SupportL

138 Chapter V Discussion Chapter Overview The following chapter includes a discussion of all results previously reported. Literature lacks pertaining to the design and efficacy of facet replacement systems; however, the discussion pertains to the artificial facet study at hand and previous studies on the biomechanical results of performing facetectomy and laminectomy on the lumbar spine. All reported results for the cadaveric study and FE simulations will be reviewed for general trends. A description of the limitations of the study is included. Lastly, proposals for future work to evaluate the need and biomechanical issues involved in designing artificial facets are included. Discussion Range of Motion Biomechanical Studies It has been reported that unilateral and bilateral total facetectomy in cadaveric studies results in instability in axial rotation and flexion [31]. FE studies are in agreement that 75% resection of the facet joints increases the instability of the lumbar spine [34], especially in rotation and flexion [30]. In decompression surgeries in which the 116

139 117 supraspinous and interspinous ligaments are resected, Pintar et al and Zander et al found motion increases, thereby suggesting posterior ligaments should be preserved [30,32]. In Vitro Study Destabilization of the cadaveric spines at L4-L5 by a facetectomy increased motion in all loading modes but extension no follower load and flexion with follower load. Maximum change in motion at L4-L5 occurred in flexion with follower load of 60.9% less than intact. Extension NFL was the loading mode resulting in the greatest change in motion for the lumbar segment, followed by flexion FL, rotation, bending, flexion NFL, and extension FL. Implanted artificial facets changed motion in flexion NFL greatest, then rotation, bending, extension NFL, and finally extension FL. The implant increased motion in all loading modes but extension FL at L4-L5. The maximum increase in motion at L4-L5 was in rotation by 64.1%, possibly due to implant misalignment, capsular ligament removal, or implant bending. In general, the implant caused an increase in motion at all segments. However, application of a follower load decreased motion with the implant in extension as compared to intact. FE Study Motion of the lumbar spine at L4-L5 increased in all loading modes due to cap designs. Pedicle screw based designs increased L4-L5 motion in all modes as well but flexion NFL and bending in which slight decreases occurred. Maximum change in motion at L4- L5 with the cap design was in extension with follower load (37.3% increase due to cap screw design). The pedicle screw based designs changed motion greatest at L4-L5 in

140 118 extension FL where a maximum increase of 56.5% due to the 3mm design occurred. The support design resulted in motion more similar to intact in all loading modes but bending. The lumbar segment motion with cap designs was affected most in extension with follower load, followed by extension, rotation, flexion NFL, flexion FL, and bending. Pedicle screw based designs resulted in the greatest change in motion in extension FL, followed closely by rotation, extension NFL, bending, flexion FL, and flexion NFL. Motion increased due to the wide laminectomy procedure at L4-L5 in all loading modes. The pedicle screw implants also increased motion at L4-L5 in all loading modes but bending (Support decreased motion by 58.1%) more than the destabilized condition. Therefore, the current designs of artificial facets are not suggested for use to stabilize wide laminectomy procedures. The increase is likely due to the removal of posterior ligaments (supraspinous, interspinous, and capsular). Maximum changes in motion at L4-L5 occur in flexion NFL by 96.6% increase in the destabilized spine and 112.5% increase with the artificial facet pedicle screw based designs. Flexion NFL resulted in the greatest motion changes in all the levels of the lumbar spine followed by flexion FL, extension FL, rotation, extension NFL, and bending (except the support design) for the destabilized and artificial facet models. Range of Motion Summary The cadaver and FE study agree that motion increases due to implantation of artificial facet systems. The increase is most likely due to removal of the capsular ligaments due to the total facetectomy, further confirming the results of Abumi et al, Teo et al, and

141 119 Zander et al [30,31,34]. The greater instability of the lumbar spine for the wide laminectomy models in the FE study was expected, due to removal of posterior ligaments (supraspinous, interspinous, and capsular) [30,32]; therefore, some type of instrumentation is needed to induce stability. When artificial facets were used as the stabilizing instrumentation, normal motion was not restored to the destabilized by wide laminectomy spine, especially in flexion. The increase in spinal motion may result in abnormal loading patterns and cause pain. Due to motion increases, disc stresses are likely to increase, further leading to disc degeneration and patient pain [19,80]. See Appendix F for further discussion on disc stresses due to artificial facet implantation. Facet Loads Biomechanical Studies Facet joint degeneration results in spinal instability [16] and increased loading of the intervertebral disc [17]. Increased disc loading due to facet load decreases often causes disc degeneration at the diseased level [17]. FE studies performed by Kim et al confirm in vitro investigations that degenerated discs correspond with decreased facet loads [19]. One level disc degeneration may also result in changes in adjacent levels shown in an FE study [19], which is a clinically observed event [80]. Destabilization of the lumbar spine by facetectomy or laminectomy in FE models hardly affect facet loads [30,34]. FE Study Artificial facets across the L4-L5 motion segment decreased facet load transfer across the joint in all loading modes but rotation where both cap designs and the support design

142 120 increased facet loads. Maximum decreases in facet loads from intact at L4-L5 occur in extension with follower load. Caps decrease loads by 14.8% and the 3mm pedicle screw based design decreased facet loads by 28.0%. However, the support design increased facet loads maximally in rotation by 13.8%. Facet loads were changed the most in extension NFL with the cap designs followed by extension FL, rotation, bending, and flexion FL. The pedicle screw based designs experienced maximal facet load changes as compared to intact in bending, followed by extension NFL, extension FL, flexion FL, and rotation. Destabilization by wide laminectomy changed facet loads from intact greatest in bending, followed by rotation, extension NFL, extension FL, and flexion FL. The pedicle screw based artificial facets resulted in similar trends in that bending resulted in the maximal change from intact in facet loads followed by extension NFL, extension FL, flexion FL, and rotation. Maximum change in loads transferred across L4-L5 due to destabilization was 25.2% decrease in bending. The pedicle screw based designs resulted in a similar maximum decrease in facet loads at L4-L5 of 28.8% in extension with follower load as a result of the 3mm design. The support design increased facet loads at L4-L5 by 11.7% in rotation, maximally. In general, destabilization decreased facet loads in all bending moments. Implantation of the pedicle screw designs also decreased facet loads in all loading modes, although the support design increased load transfer in rotation.

143 121 Facet Loads Summary Facet loads decreased due to artificial facet implantation, regardless of implant used or follower load application. Follower load application in extension resulted in further decrease in facet loads from intact for the facetectomy and wide laminectomy models across the cap and pedicle screw based designs at L4-L5. Therefore, L4-L5 disc loading must increase to properly transfer the follower load through the spinal segment. Increased loads in the disc result in disc degeneration [17,19]. Facet loads did not change due to destabilization in all loading modes confirmed by Teo et al and Zander et al [30,34]. Due to the likely increase in disc loading as a result of artificial facet implants, disc degeneration at the implanted and adjacent levels is probable [19,80]. Maximum Implant Stress Since a rigid screw and rod system has been used in surgery for several years, it has been used as a gold standard in this study to evaluate stresses in the pedicle screws. All loading modes and models led to stresses less than in the rigid system for the L4 and L5 screw, but in extension NFL and FL in which L4 pedicle screw stresses increased by 15% and 30%, respectively in facetectomy models and 23% and 28% in wide laminectomy models. By placing caps with screws in the L5 pedicle, stresses in the bone surrounding the screws increased from that of intact. Stresses in extension due to the cap screws in the L5 pedicle increased 600% more than that of intact. The artificial facet implants experience stresses well below the yield strength of titanium. The factor of safety was lowest for the support pedicle screw based design at 2.8 due to large stresses in the inferior component, while all other designs had factor of safeties of 4.9 or greater. When

144 122 the implants were used to stabilize the wide laminectomy, the factor of safety decreased for each design but the support which was 2.9, equivalent to the lowest facet of safety. Even though the artificial facet pedicle screw stresses were comparable to the rigid system pedicle screw stresses, the implant may still fail at the bone-screw interface. The rigid screw and rod system was designed to aid in the support and completion of fusion, but pedicle screw loosening has been a significant problem during the fusion process [75,76]. Artificial facet pedicle screws will be implanted and stressed the entire life span of a patient; therefore, pull-out due to screw loosening is an important design consideration. A coating to improve the bone-implant contact such as Hydroxyapatite coating may reduce the risk of pull-out due to screw loosening. It has been shown in short-term studies that Hydroxyapatite coating improves bone-to-implant interaction such that pull-out resistance was increased [75,76]. Conclusions In vitro cadaveric and finite element studies are complementary investigation techniques in the analysis of lumbar spine biomechanics. Neither technique describes the full extent of complex biomechanical issues individually. While cadaveric studies are excellent in determining range of motion due to instrumentation and surgical techniques, several limitations exist and other parameters such as facet loads and stresses are impractical to obtain. Finite element models supplement the in vitro studies to determine such important biomechanical factors and are validated by the cadaveric data. However, finite element models also have several limitations such as inability to model variation and

145 123 material property differences in specimens. The finite element models used in this study were validated with previous in vitro investigations performed and the preliminary artificial facet in vitro study. Once validated, the finite element models were an invaluable tool in understanding facet loads and implant stresses due to facet joint replacement systems. Artificial facets increased motion in both the in vitro and finite element study. The motion increase was most likely due to partial ligamentum flavum and capsular ligament removal. Stability to the lumbar spine was not restored when artificial facets were used across a wide laminectomy. Designs having a capsular ligament surrounding the joint may help to reduce the increased motion. Facet loads decreased in all loading modes with all implant designs at L4-L5 and adjacent levels. The decrease in facet loads may result in accelerated disc degeneration at the implanted and adjacent levels [17,19] and future surgical intervention for young patients. All implant stresses were well below the actual yield strength of titanium; therefore, the implants are unlikely to mechanically fail. Several design modifications may be made to the implants to increase the factor of safety and efficacy to restore normal spinal biomechanics. The designs presented in the study were reviewed by orthopedic experts to determine the optimum design for implantation considering surgery simplification, current surgical knowledge, and patient safety. The cap design with screws was considered a poor design based on small screw size and difficult screw placement, potentially undermining the safety of the patient. The cap design without screws was also inadequate. Several

146 124 studies have been completed on mechanical loosening when capping articular bone ends which eventually lead to avascular necrosis [41,45]. If the caps loosen, nearby structures and spinal nerves may be damaged. Contraindications for the cap designs include any degree of laminectomy resulting in cap implant use highly selective. The pedicle screw based designs are more welcoming to surgeons generally due to a previous knowledge base on pedicle screw placement. However, the orientation of the facet system may be in jeopardy due to pedicle screw placement variability. During implantation, the support system may be better than others since it will be forced against the pedicle increasing the likelihood of correct joint orientation as well as range of motion. This difference was shown in the FE models since the support system resulted in motion and loads most similar to intact due to the tie contact between the implant and pedicle bone. Facet joint replacement systems require thorough biomechanical investigation before human implantation. It is desirable that artificial facets restore normal biomechanical parameters to the destabilized spine. However, the implants evaluated in the current study did not restore relative motion or facet loads to that of intact. Thus, it is believed accelerated degeneration of spinal components will occur due to implantation. Several investigations and design modifications must be completed to improve the effectiveness of facet joint replacement systems. Study Limitations Several limitations exist in the described cadaveric and finite element study. For the cadaveric study, only two specimens were used making any statistical analysis

147 125 meaningless. Also, the FE results are compared to the cadaveric study using a similar implant. While the modeled implant and the cadaveric study implant are very similar, some variation does exist. Implant size and exact placement in the cadaveric study were not modeled. Although the FE models have been validated with in vitro data, biological variability such as degeneration and osteophytes cannot be modeled. However, the cadaveric results were in good agreement with the finite element results regardless of specimen variability. The finite element model is also isotropic and homogenous, falling short of simulating a true physiological situation. Apparent limitations exist with finite element modeling such as contact simulations, geometric modeling, and material property definitions. Future Work While no facet replacement system is currently available on the market, it is in the foreseeable future. Several biomechanical studies are necessary to fully evaluate the effectiveness and safety of artificial facets. The study discussed in this paper only addresses two different facet replacement systems even though there are several more potential designs. With increasing knowledge and interest in facet replacement, several more novel concepts will be patented and require thorough investigation. A ball and socket design was introduced and patented when the current study was near completion, thus the design could not be included in the biomechanical study. It is highly suggested a ball and socket design be investigated using cadaveric and finite element

148 126 studies. Thus, the range of motion, stress distribution, and load transfer for the ball and socket design may be compared to the modeled designs. The current models were analyzed as if placed during surgery at the natural orientation of the facet joint. However, it is highly unlikely that the replacement joint will be placed at exactly the correct orientation for every surgery. Thus, a study is needed to evaluate biomechanical parameters as a result of orientation. Facet joint orientation is unknown if it is of extreme importance to the function of the spine or not [10], therefore, it is unknown if changing the orientation of the replacement systems will greatly change the range of motion, load transfer, and stresses throughout the spinal column making the need for a facet orientation study imperative. Even though titanium is used frequently for orthopedic implants, it is certainly not the only material available. Studies involving different materials would be helpful to further complete the design process. These materials could be, but not limited to, medical grade stainless steel, polymers, or ceramics. For articulating surfaces, different materials may be in contact than the titanium on titanium in the current study. Thus, new designs may include ultra-high molecular weight polyethylene, ceramic, or other polymers articulating with metal or other materials. Similar articulating surfaces have been used in total hip and knee implants and wear has been shown, therefore the full extent of wear due to lining the implants must be studied as well utilizing cyclic tests.

149 127 One patent was introduced that involved multilevel artificial facets. A study involving more than one level with artificial facets is needed to evaluate the effectiveness of the concept. Different facet replacement designs should be used to compare the designs and determine the most appropriate for multilevel surgeries. Screw pull-out is a significant concern in rigid screw and rod systems used in the stabilization of fusion. Pedicle screw loosening during the fusion process often results in non-union or loss of correction [75,76]. Using cadaver specimens, a pull-out strength test and cyclic loading experiment should be performed on implanted artificial facets to fully understand implant life. Fusion surgeries are common today for young and old patients alike. However, it has been found that fusion causes adjacent level degeneration including facet joint degeneration which may lead to further surgical intervention in many patients [37]. Instead of fusing the degenerated adjacent levels, artificial facets may be a consideration. A finite element study including a fusion site above or below the placement of artificial facets may be helpful to understand biomechanical factors involved. Since severe facet degeneration is a contraindication for total disc arthroplasty systems, facet replacements may be a conjunctive market [1]. If future uses for artificial facets are with total disc arthroplasties, a study is needed to investigate the biomechanical considerations involved. Both cadaveric and finite element studies would be useful in evaluation of artificial facets at the level of artificial disc placement and at adjacent levels.

150 References 1. Viscogliosi AG, JJ Viscogliosi, MR Viscogliosi. Beyond total disc, the future of spine surgery. Viscogliosi Bros., LLC. New York Lee CK. U.S. Published Patent Application Number A1, Feb 24, White, AA and MM Panjabi. Clinical Biomechanics of the Spine. 2 nd ed. Philadelphia: J. B. Lippincott Company, Sheehan JM, CI Shaffrey, JA Jane. Degenerative lumbar stenosis: The neurosurgical perspective. Clin Orthop 2001; 384: Wiltse LL, WH Kirkaldy-Willis, GWD McIvor. The treatment of spinal stenosis. Clin Orthop 1976; 115: Johnsson KE, S Willner, K Johnsson. Postoperative instability after decompression for lumbar spinal stenosis. Spine 1986; 11(2): Chen TY. The controversy over diagnosis and treatment of facet pain. J Chin Med Assoc 2005; 68(2): Borenstein D. Does osteoarthritis of the lumbar spine cause chronic low back pain? Current Pain and Headache Reports 2004; 8: Vanharanta H, T Floyd, DD Ohnmeiss, SH Hochschuler, RD Guyer. The relationship of facet tropism to degenerative disc disease. Spine 1993; 18(8):

151 Boden SD, KD Riew, K Yamaguchi, TP Branch, D Schellinger, SW Wiesel. Orientation of the lumbar facet joints: association with degenerative disc disease. J Bone Joint Surg 1996; 78A(3): Noren R, J Trafimow, GBJ Andersson, MS Huckman. The role of facet joint tropism and facet angle in disc degeneration. Spine 1991; 16(5): Benoist M. Natural history of the aging spine. Eur Spine J 2003; 12(Suppl 2):S Ferguson SJ, T Steffen. Biomechanics of the aging spine. Eur Spine J 2003; 12(Suppl. 2):S97-S Fujiwara A, TH Lim, HS An, N Tanaka, CH Jeon, GBJ Andersson, VM Haughton. The effect of disc degeneration and facet joint osteoarthritis on the segmental flexibility of the lumbar spine. Spine 2000; 25(23): Gries NC, U Berlemann, RJ Moore, B Vernon-Roberts. Early histological changes in lower lumbar discs and facet joints and their correlation. Eur Spine J 2000; 9: Gamradt SC, JF Wang. Lumbar disc arthroplasty. Spine J 2005; 5: Kim YE, VK Goel, JN Weinstei, TH Lim. Effect of disc degeneration at one level on the adjacent level in axial mode. Spine 1991; 16(3): Fujiwara A, K Tamai, M Yamato, HS An, H Yoshida, K Saotome, A Kurihashi. The relationship between facet joint osteoarthritis and disc degeneration of the lumbar spine: an MRI study. Eur Spine J 1999; 8:

152 Moore RJ, TN Crotti, OL Osti, RD Fraser, B Vernon-Roberts. Osteoarthritis of the facet joints resulting from anular rim lesions in sheep lumbar discs. Spine 1999; 24(6): Garfin SR, HN Herkowitz, S Mirkovic. Spinal stenosis. J Bone Joint Surg 1999; 81A(4): Johnsson KE, I Rosen, A Uden. The natural course of lumbar spinal stenosis. Clin Orthop 1992; 279: Park JB, H Chang, KW Kim, SJ Park. Facet tropism: a comparison between far lateral and posterolateral lumbar disc herniations. Spine 2001; 26(6): Karacan I, T Aydin, Z Sahin, M Cidem, H Koyuncu, I Aktas, M Uludag. Facet angles in lumbar disc herniation: their relation to anthropometric features. Spine 2004; 29(10): Farfan HF, JD Sullican. The relation of facet orientation to intervertebral disc failure. Candian J Surg 1967; 10: Ahmed AM, NA Duncan, DL Burke. The effect of facet geometry on the axial torque-rotation response of lumbar motion segments. Spine 1990; 15(5): Lee CK. Lumbar spinal instability (olisthesis) after extensive posterior spinal decompression. Spine 1983; 8(4): Mariconda M, R Fava, A Gatto, C Longo, C Milano. Unilateral laminectomy for bilateral decompression of lumbar spinal stenosis: a prospective comparative study with conservatively treated patients. J Spinal Disord & Techniques 2002; 15(1):39-46.

153 Zander T, A Rohlmann, C Klockner, G Bergmann. Influence of graded facetectomy and laminectomy on spinal biomechanics. Eur Spine J 2003; 12: Abumi K, MM Panjabi, KM Kramer, J Duranceau, T Oxlan, JJ Crisco. Biomechanical evaluation of lumbar spinal stability after graded facetectomies. Spine 1990: 15(11): Pintar FA, JF Cusick, N Yoganandan, J Reinartz, M Mahesh. The biomechanics of lumbar facetectomy under compression-flexion. Spine 1992; 17(7): Lee KK, EC Teo. Effects of laminectomy and facetectomy on the stability of the lumbar motion segment. Med Eng Physics 2004; 26: Teo EC, KK Lee, QHW Ng, K Yang. The biomechanics of lumbar graded facetectomy under anterior-shear load. IEEE Trans Biomed Eng 2004; 51(3): Nagata H, MJ Schendel, EE Transfeldt, JL Lewis. The effects of immobilization of long segments of the spine on the adjacent and distal facet force and lumbosacral motion. Spine 1993; 18(16): Kotani Y, BW Cunningham, A Cappuccino, K Kaneda, PC McAfee. The effects of spinal fixation and destabilization on the biomechanical and histologic properties of spinal ligaments: an in vivo study. Spine 1998; 23(6): Okuda S, M Iwasaki, A Miyauchi, H Aono, M Morita, T Yamamoto. Risk factors for adjacent segment degeneration after PLIF. Spine 2004; 29(14): van Ooij A, FC Oner, AJ Verbout. Complications of artificial disc replacement. J Spinal Disord Techniques 2003; 16(4):

154 Guyer RD, DD Ohnmeiss. Intervertebral disc prostheses. Spine 2003; 28(15S):S15-S Dooris AP, VK Goel, NM Grosland, LG Gilbertson, DG Wilder. Load-sharing between anterior and posterior elements in a lumbar motion segment implanted with an artificial disc. Spine 2001; 26(6):E Goble ME, TW Fallin, RW Hoy. U.S. Patent Number , June 17, Cinotti G, T David, F Postacchini. Results of disc prosthesis after a minimum follow-up period of 2 years. Spine 1996; 21(8): Fitz WR. U.S. Patent Number , Nov 5, Sharma M, NA Langrana, J Rodriguez. Role of ligaments and facets in lumbar spinal stability. Spine 1995; 20(8): Serhan HA, R Labrom, JR Hawkins, W Dunbar, A Borgstrom, SD Kwak, A Dooris, MA Slivka. U.S. Published Patent Application Number A1, Mar 10, Martin JR. U.S. Patent Number , Oct 17, Reiley MA. U.S. Patent Number , Aug 26, Goble EM, TW Fallin, RW Hoy. U.S. Patent Number , May 20, Fallin TW, EM Goble, RW Hoy. U.S. Patent Number B1, Jul 16, Berry BM. U.S. Published Patent Application Number A1, Feb 10, W. Clinically relevant Biomechanical Parameters in the lumbar spine- a theoretical and in vitro experimental investigation. in Bioengineering. 1995, University of Iowa.

155 Shirazi-Adl A, AM Ahmed, SC Shrivatsava. A finite element study of a lumbar motion segment subjected to pure sagittal plane motions. J Biomech 1986; Dooris, Andrew P: Effect of Implanted Artificial Discs on the Lumbar Spine An Experimental and Analytical Investigation (PhD Dissertation). 54. Dupuis PR et al. Radilogical diagnosis of degenerative lumbar spine instability. Spine 1985; 10: Grosland, Nicole M: Spinal Adaptation in Response to Interbody Fusion Systems: A Theoretical Investigation (Ph.D. in Biomedical Engineering). 56. Schultz AB et al. Mechanical properties of human spine motion segment part I: responses in flexion, extension, lateral bending, and torsion. J Biomech Eng 1979; 101: Yang KH, AI King. Mechanism of facet load transmission as a hypothesis for low-back pain. Spine 1984; 9(6): Shirazi-adl. Finite-element evaluation of contact loads on facets of an L2-L3 lumbar segment in complex loads. Spine 1991; 16(5): Kim, YE (Thesis): Analytical investigation of ligamentous lumbar spine biomechanics. Biomedical Engineering Department: University of Iowa, Iowa City, IA , William D. Materials Science and Engineering: An Introduction. 5 th edition. New York: John Wiley & Sons, Inc. 2000; p Okuda S, M Iwasaki, A Miyauchi, H Aono, M Morita, T Yamamoto. Spine 2004; 29(14): Van De Graff, Kent M. Human Anatomy. Boston: McGraw Hill, 2002.

156 Goel VK, YE Kim, TM Lim. An analytical investigation into the biomechanics of spinal instrumentation. Spine 1988; 13: Agur, Anne MR and Ming J Lee. Grant s Atlas of Anatomy. 10 th ed. Philadelphia: Lippincott Williams & Wilkins, Panjabi MM et al. Human lumbar vertebrae: Quantitative three-dimensional anatomy. Spine 1992; 17(3): Holekamp S. A Parametric Biomechanical Study of Intervertebral Sealant Devices for the Lumbar Spinal Disc University of Toledo, Ohio. 67. Hutton WC, AM Adams. Can the lumbar spine be crushed by heavy lifting? Spine 1982; 7: Nachemson A. Lumbar interdiscal pressure. Acta Orthop Scand 1960; 96: Farfan HF et al. The effects of torsion on the lumbar intervertebral joints. J Bone Joint Surgery 1970; 52A: Adams MA, WC Hutton, JRR Stott. The resistance to flexion of the lumbar intervertebral joint. Spine 1980; 5: Tencer AF, AM Ahmed, DL Burke. Some static mechanical properties of the lumbar intervertebral joint, intact and injured. J Biomech Eng 1982; 104: Sonoda T. Studies on the strength for compression, tension, and torsion of the human vertebral column. J Kyoto Pref. Med Univ. 1962; 71: Vadapalli, Sasidhar. Stability imparted by a posterior lumbar interbody fusion cage following surgery - a biomechanical evaluation. Master s Thesis, Bioengineering, University of Toledo, Toledo, Ohio, 2004.

157 Khandha, Ashutosh. A parametric evaluation of the biomechanics of lower lumbar spine after fusion surgery. Master s Thesis, Bioengineering, University of Toledo, Toledo, Ohio, Bengt S, C Olerud, C Johansson, S Larsson. Improved bone-screw interface with hydroxyapatite coating: an in vivo study of loaded pedicle screws in sheep. Spine 2001; 26(24): Sanden B, C Olerud, S Larsson. Hydroxyapatite coating enhances fixation of loaded pedicle screws: a mechanical in vivo study in sheep. Eur Spine J 2001; 10(4): Lotz JC. Animal models of intervertebral disc degeneration. Spine 2004; 29(23): Stokes IAF, JC Iatridis. Mechanical conditions that accelerate intervertebral disc degeneration: overload versus immobilization. Spine 2004; 29(23): Edwards WT, NR Ordway, Y Zheng, G McCullen, Z Han, HA Yuan. Peak stresses observed in the posterior lateral annulus. Spine 2001; 26(16): Natarajan RN, JR Williams, GBJ Andersson. Recent advances in analytical modeling of lumbar disc degeneration. Spine 2004; 29(23):

158 Appendix A Functional Anatomy of the Spine Introduction The spine has four major functions: to support the head and upper extremities while permitting freedom of movement, to enable bipedalism, to provide attachment for various muscles, ribs, and visceral organs, and to protect the spinal cord and permit passage of the spinal nerves [62]. The spine is composed of a series of bony vertebrae separated by spongy discs, all supported by ligaments, muscles, tendons, fascia, and other connective tissue. The vertebral column is composed of individual vertebrae divided into five regions: 7 cervical, 12 thoracic, 5 lumbar, 5 fused sacral, and 4 or 5 fused cocygeal [62] (Figure A.1). The function of the spine is facilitated by each region being curved in the sagittal plane [62] as can be seen in Figure A.1. The cervical and lumbar regions are lordotic (bowing dorsally) and the thoracic is kyphotic (bowing ventrally) [53]. The wedge like shape of the intervertebral discs and the shape of the vertebrae help to produce the curve [53]. 136

159 137 Cervical Thoracic Lumbar Cocygeal Sacrum Figure A.1: The spinal column displaying the cervical, thoracic, lumbar, sacral, and cocygeal regions [53]. Vertebrae Spine stiffness is created by the bony vertebrae. The vertebrae not only provide structural stiffness, but also attachment points for connective tissues. The anterior region of the vertebral body is approximately cylindrical in shape; composed of a hard outer cortical bone shell surrounding a porous cancellous bone core. The stiffness, or elastic modulus, of cortical bone is very high (E > 10 GPa), where as the cancellous bone core is less stiff (E < 1 GPa) [63]. The vertebral endplates at the superior and inferior most surfaces of the vertebrae are concave in shape. The endplates are the strongest region and composed of cortical bone with layers of cartilage to aid attachment of the intervertebral disc. The lumbar vertebrae (Figure A.2) are the largest and strongest of the spinal column. The lumbar vertebrae are larger in the lateral direction than in the anterior-posterior diameter and are slightly thicker on the anterior side to provide a portion of the lordotic

160 138 curve [64]. The foramina of the lumbar vertebrae are triangular in shape, smaller than the cervical vertebrae, but larger than thoracic [64]. The spinous processes are thick and broad in the lumbar region and project downward. The fifth lumbar vertebra varies considerably amongst individuals in size and shape of the processes and articulations [53]. Vertebral foramen Body Pedicle Pedicle Body Superior articular process Spinous process Lamina Transverse process Transverse process Superior articular process Spinous process Inferior vertebral notch Inferior articular process Figure A.2: Anatomic regions on lumbar vertebrae [3]. Posterior Elements The posterior elements include all structures posterior to the intervertebral foramen such as the pedicle, lamina, spinous process, transverse processes, and articular processes (inferior and superior facet). The pedicles are dense, posteriorly projecting regions that support the facet joint surfaces. Surrounding the spinal canal are the laminae, broad dense bony regions that complete the foramen. The spinous process projects posteriorly from the junction of the lamina. Ligaments, muscles, and other connective tissue attaches directly to the spinous process. The two transverse processes project laterally from the lamina and pedicle intersection. The transverse processes are also attachment sites for

161 139 the ligaments, muscles, and other connective tissues. The superior articular process projects from the superior side of the pedicle with a cartilaginous layer directed posteriorly. The inferior articular process projects inferiorly from the inferior side of the pedicle with a cartilaginous layer facing anteriorly. Facet Joints There are four facet, or articular, surfaces including two superior and two inferior surfaces on each vertebra. Each facet surface engages a neighboring vertebra s facet surface through a diarthroidal joint. The joint consists of cartilaginous articular surfaces, synovium, capsular ligaments, and nerve filaments [53]. The facet joint is important in spinal biomechanics by largely supporting the spine in extension, rotation, and load transfer. The orientation of the facets is crucial to the motion of the spine and varies between the cervical, thoracic, and lumbar regions. In the lumbar region, the facets are oriented vertically and approximately 45 from the mid-sagittal plane [3]. Intervertebral Disc Intervertebral discs allow for a flexible connection between each of the vertebral bodies. The disc is unique in that it is deformable to allow bending, twisting, flexing, and extending of the spinal column. Each disc varies in size, in proportion to the neighboring vertebral bodies, and can contribute to one-fifth to one-third of the total spine s length [3,53]. The lumbar disc shape is wedge like in the sagittal plane with thickness greatest anteriorly and decreasing posteriorly to result in a lordotic curve. In the horizontal plane, the shape of the disc is similar to the vertebral body and often described as elliptical with

162 140 major and minor axes [65] which change considerably throughout the spinal column. The lumbar disc is most often characteristically described as oblate. An adult disc is avascular and composed of a fluidic central nucleus pulposus and surrounding annulus fibrosus. The disc attaches to the vertebral bodies via a hyaline cartilage endplate [53]. Nucleus Pulposus The nucleus is a mucoprotein gel composed of a large concentration of proteoglycans and short amino acid chains [3,53]. The nucleus water content is 70-90% in healthy discs [3]. Since the nucleus is enclosed by the annulus and endplates, the large water content creates a large hydrostatic pressure [53]. In the lumbar region, the nucleus occupies 30-50% of the total cross-sectional area of the disc [3]. Annulus Fibrosus The annulus fibrosus (Figure A.3) forms the outer boundary of the intervertebral disc and is composed of fibrous tissue in concentric laminated bands [3]. Approximately twelve layers of fibers encircle the nucleus with each fiber oriented at ± 30 from the horizontal alternating direction from one layer to the next [53]. The fiber layers and fibers increase in thickness while the collagen content decreases as one moves radially outward. Annular fibers respond mechanically with an increase in resistance to an increase in axial load [53].

163 141 Nucleus pulposus Annulus laminates Annulus fibers Figure A.3: Annular fibers and the fiber orientation in the intervertebral disc [3]. Ligaments Ligaments are effective in carrying loads along the direction in which the fibers it is composed of are oriented [3]. Thus, the spinal ligaments provide resistance to tensile forces, but buckle under compression [3]. Ligaments have several functions. These functions are as follows: 1) ligaments allow physiological motion with minimum energy expenditure, 2) they protect the spinal cord by restricting motion within defined limits, 3) ligaments provide stability to the spine, and 4) they protect the spinal cord from traumatic motions in which high loads are applied at fast speeds [3]. There are seven major ligaments in the lumbar spine shown in Figure A.4: anterior longitudinal, posterior longitudinal, supraspinous, interspinous, ligamentum flavum, intertransverse, and capsular. The anterior longitudinal ligament traverses the anterior surface of the entire spine. The anterior longitudinal primarily restricts motion in

164 142 extension. The posterior longitudinal ligament traverses the posterior surface of the entire spine, lining the vertebral foramen. The supraspinous ligament runs along the posterior edge of the spinous process providing stability in flexion [3]. The interspinous ligament is attached to adjacent spinous processes in the sagittal plane and also helps resist flexion. The ligamentum flavum is the most elastic ligament, helps protect the spinal cord, and connects to adjacent laminae [3]. The intertransverse ligament attaches to neighboring transverse processes and restricts motion in bending and axial rotation. The facet capsular ligaments surround the facet joint and the fibers are oriented perpendicular to the facet surface helping to provide stability in flexion [3]. Intertransverse Ligament Ligamentum Flavum Posterior Longitudnal Ligament Facet Capsular Ligament Interspinous Ligament Anterior Longitudnal Ligament Supraspinous Ligament Figure A.4: The seven major ligaments of the lumbar spine [3]. Nerve Pathways A distinct function of the spinal column is to protect the spinal cord. The spinal cord is continuous with the brain and extends through the vertebral canal of the entire spinal

165 143 column. The spinal cord conducts information from the brain to muscles and organs throughout the entire body [62]. Damage to the spinal cord or any nerve tracts may bring pain or disability to the innervated area. Due to the extreme complexity of the nerve passages shown in Figure A.5, only a brief overview is described below. The spinal cord proceeds through the spinal canal and is composed of three membranes for protection: the dura mater, pia mater, and arachnoid [62]. Dorsal and ventral nerve roots exit the dural mater sheath into the intervertebral foramen. The dorsal and ventral roots combine to form spinal nerves, then continue onto peripheral nerves [64]. The path the nerves exit the spinal canal can decrease in size due to intervertebral disc degeneration, inflammation of surrounding tissues, or trauma. Referred pain can be caused by neural root compression or stretch. Free nerve endings have been found within the outer half of the annulus, the anterior longitudinal ligament, and the facet joints [53]. Therefore, these structures can cause pain. Figure A.5: Neural anatomy of the lumbar spine [3].

166 144 Low Back Muscles Low back muscles provide stability and control to the spinal column. Without muscles, the spine would buckle under an axial load as small as 20N [53]. The muscles of the back are symmetrically arranged left/right and in opposing flexion/extension groups supporting and controlling the spine within the range of motion imposed by the ligaments and joints. Muscles directly related to spinal column control are divided into two large groups depending upon location and insertion point: postvertebral and prevertebral. The postvertebral group is primarily located anterior or near the body and the prevertebral group is located posterior to the vertebral body [3]. The prevertebral muscles include the four sets of abdominal muscles: the external oblique, internal oblique, transverse abdominus, and rectus abdominus [3]. The postvertebral muscles are a more complex set and include the deep, intermediate, and superficial groups. The deep muscles are short and connect adjacent spinous processes. The intermediate muscles connect the transverse process with the superiorly adjacent spinous process. The superficial muscles are collectively called the erector spinae [3]. Because muscles produce force via tension, one group will be more active than the other (left/right of pre/post- vertebral) in a single motion. However, controlled motion requires the activation of both sets of muscles. Even upright posture involves a number of both anteriorly and posteriorly located muscles in small, controlled contractions [3].

167 Appendix B Biomechanics of the Lumbar Spine It is difficult to comprehensively describe the behavior of the spine to forces and moments, the topic of spine biomechanics. Recently, several techniques have been developed to quantify the spine s functional response to loading. These techniques include optoelectronic tracking, finite element modeling, advanced radiography, in vitro analysis, MRI, detailed clinical reporting, and biological characterization [53]. However, there are many questions about how the spine works that remain unanswered. Spine biomechanics involves many natural loading conditions and is extremely complex. The material presented below has been simplified by discussing each spine loading condition separately. The loading conditions presented include compression, flexion, extension, axial rotation or torsion, and lateral bending. These conditions will be considered on a functional spinal unit (FSU) or motion segment in the lumbar region described by Dupuis et al [66]. An FSU includes two adjacent vertebrae, the disc, and ligaments between them. Axial Compression The spine is loaded along its axis, perpendicular to the intervertebral discs in axial compression. The axial load is a result of upper body weight and posterior muscle 145

168 146 tension. Compression forces may be much greater than body weight due to normal activities such as lifting which may cause the load to increase around 6000 to 9000N [53, 67]. Even upright posture can increase the loads on the spine dramatically. The displacement of the spine due to compression loads can be quantified by a stiffness coefficient. The compressive stiffness coefficient varies considerably due to several factors such as soft tissue degeneration, segment level and size, and bone density [53]. Axial compression is resisted by the disc and facet joints depending on the segment angulation and disc hydration. Most of the compressive load in the lumbar spine is resisted by the intervertebral disc due to the orientation of the facet surface being almost parallel to the axial force. However, if the spine is extended and the disc is compressed, the facets may share one six of the load. According to Nachemson, the lumbar facets may carry as much as 18% of the axial compressive load [68]. The facets may carry much more of the load when combined loading and posture is considered as well as when disc degeneration occurs. The disc changes due to compressive loads by the annulus decreasing in height and bulging radially. The bulging causes the laminar fibers to become circumferentially oriented. Compression also places the nucleus under hydrostatic pressure. Significant axial pressure may result in vertebral body endplate failure.

169 147 Flexion The largest range of motion of the lumbar spine occurs in the sagittal plane including flexion and extension rotations. A moment induces a forward rotation of the spine in flexion causing a stiffness reported to be 0.8 N-m/deg [56]. The stiffness is caused by several factors. The anterior portion of the disc compresses and the posterior portion undergoes tension to create a wedge. All the ligaments other than the anterior longitudinal are stretched. The facets slide past each other in flexion while the capsular ligaments are placed in tension [69]. Spinal failure for a lumbar FSU may occur when moments are as large as 50 N-m [70]; however, physiologic flexion moments may be as large as 100N. Such large moments are balanced by muscular contraction to convert some moment into axial compression. Extension Moments that produce a backward bending in extension are limited by the facet joint. Loads are shared among the facets, disc, and ligaments and vary considerably due to anatomic factors, degenerative conditions, size, and measurement technique. Tencer et al reported extension stiffness values of approximately 0.74 /N-m with posterior elements and 1.1 /N-m without [53,71]. Schultz [56] and Markolf [53] found the lumbar FSU to rotate 1.3 /N-m and 0.5 /N-m, respectively.

170 148 Axial Rotation Torsion or axial rotation is the result of a moment that causes the spine to rotate about its longitudinal axis. Axial rotation is unique in that it produces rotation in the same plane as the intervertebral disc. In the lumbar region, the facet joint limits the range of motion. One facet is under compression while the other facet experiences tension due to the capsular ligaments. The rotational load is shared by the ligaments, approximately 10% of the spinal stiffness, and the disc, about 45%. Lumbar discs may yield at loads up to 45 N-m in axial rotation [72]. Lateral Bending In lateral bending, an applied moment causes the spine to bend to the left or right side of the body. Several structures in the FSU resists lateral bending including the intertransverse ligaments, capsular ligaments, and lateral portions of the annulus. Lateral bending in the lumbar spine also produces a coupled moment of flexion and a slight axial rotation caused by the orientation of the facet joints [56]. Thus, spine stiffness decreases in lateral bending if the facet joints are removed. If the posterior elements are removed, lateral bending stiffness decreases from 2.07 N-m/deg to 1.15 N-m/deg according to Tencer et al [71]. Schultz [56] and Markolf [53] found varying rotational stiffness for the intact lumbar spine of 1.22 and 2.6 N-m/deg, respectively. Range of Motion The range of motion of the spine is controlled by several structures. A ligament in tension limits displacements between two separating elements. Facets and spinous

171 149 processes restrict motion by contacting surfaces. Endplates are restricted in motion due to strains in the intervertebral disc. Typical ranges of motion (ROM) for the lumbar spine are listed in Table B.1 [53]. Range of Motion (degrees) Spinal Level Flexion Extension Lateral Bending Axial Rotation L1-L L2-L L3-L L4-L L5-S Table B.1: Range of motion for lumbar segments in all rotation modes [53]. Shear Shear forces may occur in the spine in the plane of the disc. Normal activities such as bending and lifting can produce shear as high as 750N [53]. The facets and intervertebral disc resist shear in the lumbar spine. Since the facets are oriented oblique to the midsagittal plane and vertical, they share shear load. The facets can support between one half and two thirds of the shear force [53]. A disc that is degenerated or undergone creep contributes little to resisting shear such that the facets may have to bear most of the load [53].

172 Appendix C Finite Element Model Validation Further validation of the model used in this study was completed with posterior instrumentation using a fresh, cadaveric L2-L3 functional spine unit (FSU). The specimen was thawed to room temperature. The L2-L3 intervertebral disc was removed and replaced by an instrumented silicone disc. The ligaments, tissue, and L2-L3 facet joint were also removed from the FSU. The endplates were leveled with polyester material to provide a flat surface area and the superior endplate of L2 was fitted with a rigid polymer plate to resist indentation. Posterior instrumentation with mounted strain gages, pedicle screws and 5mm rods were used to reassemble the unit (Figure C.1A). The disc space was sized to allow insertion of discs labeled air gap, silicone, and polyurethane, in order of increasing stiffness. The FSU was placed on the load cell of an MTS 850 Bionix (MTS, Eden Prairie MN) testing machine and a controlled compressive load was applied to the top plate using a 5.0 mm diameter pin mounted to the MTS actuator (Figure C.1B). The load was applied two inches anterior to the plane of the rods, first medially and then offset to the right [73]. Output was recorded by a voltmeter upon stabilization. Compression loads were applied from 50N to 300N in steps of 50N with corresponding moments in the sagittal plane of 2.54 to N-m. The force, displacement, and time were also recorded during the 150

173 151 testing. Instrumentation bending moment was calculated and plotted against the applied moment. Results are shown in Figure C.2 for load applied medially [74]. This study was replicated with the finite element model. The intact L3-L5 FE model was modified such that L5 vertebral body and the L4-L5 disc was removed, the L3-L4 disc was removed and replaced with models of the inserts used in the cadaveric study, and all posterior elements were removed leaving only the pedicle intact. Pedicle screws were placed into the pedicles and vertebral bodies with 5mm rods as shown in Figure C.3. Interactions were defined at the surfaces in contact at the bone-pedicle screw interface and between the pedicle screw and rod. The model was constrained in all degrees of motion at the inferior most surface of L4 vertebra. A compressive force normal to the superior surface of L3 was applied to simulate the cadaveric MTS compression. Two different discs were modeled, silicone and polyurethane, and material properties simulated are listed in Table C.1. Resulting bending moments were measured at the center of the pedicle screw rod approximately at the location of the strain sensors in the cadaveric study. Results are shown in Figure C.4. The FE results are in strong correlation with the cadaveric study further substantiating the validity of the model. Rotational moments were also validated for the FE model. Predicted motions were compared in rotation to values reported in the literature. Validation of rotational moment is show in Figures C.5 through C.8.

174 152 A B Figure C.1: (A) Experimental FSU to determine strains in the posterior instrumentation. (B) Specimen setup using the MTS machine. Figure C.2: Results of measured moment versus applied moment in a cadaveric study.

175 153 Figure C.3: FE model simulation of the strain experiment. Polyurethane Insert Silicone Insert PMMA Posterior Instrumentation Tensile Modulus (psi) 600, 1750, x x 107 Poisson s Ratio Table C.1: Material properties simulated in the FE model for the disc inserts. Figure C.4: Results of measured moment versus applied moment in a FE study.

176 154 Rotation (deg) Dooris Shultz L3 L4 L3-Fluid L4-Fluid Flexion Moment (N-m) Figure C.5: FE predicted flexion motion for L3 and L4 vertebrae with 400N compression compared to Schultz [56] and Dooris [53]. Rotation (deg) Dooris Shultz L3 L4 L3-fluid L4-fluid Extension Moment (N-m) Figure C.6: FE predicted extension motion for L3 and L4 vertebrae with 400N compression compared to Schultz [56] and Dooris [53].

177 Lateral Bending 12 Rotation (deg) Dooris Shultz L3 L4 L3-Fluid L4-Fluid Moment (N-m) Figure C.7: FE predicted lateral bending motion for L3 and L4 vertebrae with 400N compression compared to Schultz [56] and Dooris [53]. Rotation (deg) Dooris Shultz L3 L4 L3-fluid L4-fluid Axial Rotation Moment (N-m) Figure C.8: FE predicted axial rotation motion for L3 and L4 vertebrae with 400N compression compared to Schultz [56] and Dooris [53].

178 Appendix D Modified Universal Joint Artificial Facet Introduction A novel design was created by the author and inserted into the L3-S1 nonlinear, ligamentous, validated finite element model. The design for the artificial facet involved a modified universal joint shown in Figure D.1 that secures to L4 and L5 pedicle screws. Rotational Joint Pedicle Screws Bending Joint Figure D.1: Modified universal joint across the L4-L5 motion segment. Materials and Methods A universal joint was modified to become a facet joint replacement system. Pedicle screws 6.5mm in diameter and 55mm in length were placed at L4 and L5. The pedicle 156

179 157 screws were inserted two-thirds into the vertebral body. A standard universal joint was modified to rotate about the mid axis, shown in Figure D.1. The universal joint was tied to the pedicle screws across the L4-L5 motion segment. The joint allowed for six degrees of freedom. Flexion and extension and left and right bending occurred due to sliding contact around pins (designated by the Bending Joint in Figure E.1) while left and right rotation occurred at the rotational joint ( Rotational Joint shown in Figure E.1). Due to the complexity of the joint, it was not able to be modeled properly in ABAQUS 6.4. Discussion To allow for bending and rotation, sliding contacts must be modeled at each bending and rotational joint. When all contacts were defined, the middle block, shown in dark green in Figure D.1, was not properly constrained and was left floating during analysis. Thus, only one sliding contact could be defined at a time, limiting the device to only pure moments instead of the full six degrees of freedom. Since the pedicle screws were inherently placed at an angle with respect to the applied moments, the joint could not be positioned such that each contact would be at the most efficient angle to allow for full motion at one sliding contact surface. Therefore, the motion of the joint was highly restricted and mechanical binding occurred, restricting the motion of the joint and the spine. Due to modeling constraints, the effectiveness of the modified universal joint was not correctly determined and the modeling was ended. When the finite element software will be upgraded to allow sliding contacts to constrain floating mechanical parts, the universal joint may be properly modeled.

180 Appendix E Facet Replacement System Facet Load Components Introduction Facet loads are determined using the centroidal stress (S11) experienced by the GAPUNI elements in between the inferior and superior facet surfaces. Since the area of each GAPUNI element is one millimeter squared, the centroidal stress value reported by ABAQUS is equal to the force transferred across the joint. The GAPUNI elements are oriented such that the force determined from this method is perpendicular to the facet contact. Thus, the total facet load may be broken into components using the direction cosines of the GAPUNI elements. These components are oriented in the medial-lateral, anterior-posterior, and vertical directions with respect to the spine. Thus, medial-lateral (Lateral) is parallel to the frontal plane, anterior-posterior (AP Shear) is parallel to the midsagittal plane, and vertical (Vertical) runs superiorly and inferiorly along the spine. The vertical component of facet loads represents the compression load transferred through the joint shared by the disc and the AP shear component represents the force involved in transverse shear. The lateral components act against each other on the right and left sides of the joint, thus the lateral components cancel when combined into the total facet load. The lateral components reported are not represented by a positive or negative value in the tables since the direction is relative to viewer position. 158

181 159 Results Facet load components for 6 N-m applied moment at L3-L4 are given in Table E.1, at L4- L5 in Table E.2, and L5-S1 in Table E.3 and for 400N follower load and 6 N-m applied moment at L3-L4 facet load components are given in Table E.4, L4-L5 in Table E.5, and L5-S1 in Table E.6. The component loads were also determined for the wide laminectomy models and are given in Tables E.7 to 12.

182 Extension Bending Rotation Facet Load Components (N) across L3/L4 and Percent Changes Compared to Intact at L3/L4 for 6 N-m Applied Moment L3/L4 Facet Loads (N) Right Left Lateral % Change AP Shear % Change Vertical % Change Lateral % Change AP Shear % Change Vertical % Change Intact Cap Cap Screw mm mm Support Intact Cap Cap Screw mm mm Support Intact Cap Cap Screw mm mm Support Table E.1: Facet load components in the lateral, anterior-posterior shear, and vertical directions and the percent change as compared to intact for artificial facet FE models in extension, bending, and rotation for 6 N-m applied moment across L3-L4 right and left facet joints. A positive percent change indicates an increase in facet loads where as a negative percent change indicates a decrease in facet loads. 160

183 Extension Bending Rotation Facet Load Components (N) across L4/L5 and Percent Changes Compared to Intact at L4/L5 for 6 N-m Applied Moment L4/L5 Facet Loads (N) Right Left Lateral % Change AP Shear % Change Vertical % Change Lateral % Change AP Shear % Change Vertical % Change Intact Cap Cap Screw mm mm Support Intact Cap Cap Screw mm mm Support Intact Cap Cap Screw mm mm Support Table E.2: Facet load components in the lateral, anterior-posterior shear, and vertical directions and the percent change as compared to intact for artificial facet FE models in extension, bending, and rotation for 6 N-m applied moment across L4-L5 right and left facet joints. A positive percent change indicates an increase in facet loads where as a negative percent change indicates a decrease in facet loads. 161

184 Extension Bending Rotation Facet Load Components (N) across L5/S1 and Percent Changes Compared to Intact at L5/S1 for 6 N-m Applied Moment L5/S1 Facet Loads (N) Right Left Lateral % Change AP Shear % Change Vertical % Change Lateral % Change AP Shear % Change Vertical % Change Intact Cap Cap Screw mm mm Support Intact Cap Cap Screw mm mm Support Intact Cap Cap Screw mm mm Support Table E.3: Facet load components in the lateral, anterior-posterior shear, and vertical directions and the percent change as compared to intact for artificial facet FE models in extension, bending, and rotation for 6 N-m applied moment across L5-S1 right and left facet joints. A positive percent change indicates an increase in facet loads where as a negative percent change indicates a decrease in facet loads. 162

185 Extension Flexion Facet Load Components (N) across L3/L4 and Percent Changes Compared to Intact at L3/L4 for 400N FL and 6 N-m Applied Moment L3/L4 Facet Loads (N) Right Left Lateral % Change AP Shear % Change Vertical % Change Lateral % Change AP Shear % Change Vertical % Change Intact Cap Cap Screw mm mm Support Intact Cap Cap Screw mm mm Support Table E.4: Facet load components in the lateral, anterior-posterior shear, and vertical directions and the percent change as compared to intact for artificial facet FE models in extension, bending, and rotation for 400N follower load and 6 N-m applied moment across L3-L4 right and left facet joints. A positive percent change indicates an increase in facet loads where as a negative percent change indicates a decrease in facet loads. 163

186 Extension Flexion Facet Load Components (N) across L4/L5 and Percent Changes Compared to Intact at L4/L5 for 400N FL and 6 N-m Applied Moment L4/L5 Facet Loads (N) Right Left Lateral % Change AP Shear % Change Vertical % Change Lateral % Change AP Shear % Change Vertical % Change Intact Cap Cap Screw mm mm Support Intact Cap Cap Screw mm mm Support Table E.5: Facet load components in the lateral, anterior-posterior shear, and vertical directions and the percent change as compared to intact for artificial facet FE models in extension, bending, and rotation for 400N follower load and 6 N-m applied moment across L4-L5 right and left facet joints. A positive percent change indicates an increase in facet loads where as a negative percent change indicates a decrease in facet loads. 164

187 Extension Flexion Facet Load Components (N) across L5/S1 and Percent Changes Compared to Intact at L5/S1 for 400N FL and 6 N-m Applied Moment L5/S1 Facet Loads (N) Right Left Lateral % Change AP Shear % Change Vertical % Change Lateral % Change AP Shear % Change Vertical % Change Intact Cap Cap Screw mm mm Support Intact Cap Cap Screw mm mm Support Table E.6: Facet load components in the lateral, anterior-posterior shear, and vertical directions and the percent change as compared to intact for artificial facet FE models in extension, bending, and rotation for 400N follower load and 6 N-m applied moment across L5-S1 right and left facet joints. A positive percent change indicates an increase in facet loads where as a negative percent change indicates a decrease in facet loads. 165

188 Facet Load Components (N) across L3/L4 and Percent Changes Compared to Intact at L3/L4 for 6 N-m Applied Moment for Wide Laminectomy Models L3/L4 Facet Loads (N) Right Left Lateral % Change AP Shear % Change Vertical % Change Lateral % Change AP Shear % Change Vertical % Change Intact DestabL Extension 3mmL mmL SupportL Intact DestabL Bending 3mmL mmL SupportL Intact DestabL Rotation 3mmL mmL SupportL Table E.7: Facet load components in the lateral, anterior-posterior shear, and vertical directions and the percent change as compared to intact for artificial facet FE models with wide laminectomy in extension, bending, and rotation for 6 N-m applied moment across L3-L4 right and left facet joints. A positive percent change indicates an increase in facet loads where as a negative percent change indicates a decrease in facet loads. 166

189 Extension Bending Rotation Facet Load Components (N) across L4/L5 and Percent Changes Compared to Intact at L4/L5 for 6 N-m Applied Moment for Wide Laminectomy Models L4/L5 Facet Loads (N) Right Left Lateral % Change AP Shear % Change Vertical % Change Lateral % Change AP Shear % Change Vertical % Change Intact DestabL mmL mmL SupportL Intact DestabL mmL mmL SupportL Intact DestabL mmL mmL SupportL Table E.8: Facet load components in the lateral, anterior-posterior shear, and vertical directions and the percent change as compared to intact for artificial facet FE models with wide laminectomy in extension, bending, and rotation for 6 N-m applied moment across L4-L5 right and left facet joints. A positive percent change indicates an increase in facet loads where as a negative percent change indicates a decrease in facet loads. 167

190 Extension Bending Rotation Facet Load Components (N) across L5/S1 and Percent Changes Compared to Intact at L5/S1 for 6 N-m Applied Moment for Wide Laminectomy Models L5/S1 Facet Loads (N) Right Left Lateral % Change AP Shear % Change Vertical % Change Lateral % Change AP Shear % Change Vertical % Change Intact DestabL mmL mmL SupportL Intact DestabL mmL mmL SupportL Intact DestabL mmL mmL SupportL Table E.9: Facet load components in the lateral, anterior-posterior shear, and vertical directions and the percent change as compared to intact for artificial facet FE models with wide laminectomy in extension, bending, and rotation for 6 N-m applied moment across L5-S1 right and left facet joints. A positive percent change indicates an increase in facet loads where as a negative percent change indicates a decrease in facet loads. 168

191 Extension Flexion Facet Load Components (N) across L3/L4 and Percent Changes Compared to Intact at L3/L4 for 400N FL and 6 N-m Applied Moment for Wide Laminectomy Models L3/L4 Facet Loads (N) Right Left Lateral % Change AP Shear % Change Vertical % Change Lateral % Change AP Shear % Change Vertical % Change Intact DestabL mmL mmL SupportL Intact DestabL mmL mmL SupportL Table E.10: Facet load components in the lateral, anterior-posterior shear, and vertical directions and the percent change as compared to intact for artificial facet FE models with wide laminectomy in extension, bending, and rotation for 400N follower load and 6 N-m applied moment across L3-L4 right and left facet joints. A positive percent change indicates an increase in facet loads where as a negative percent change indicates a decrease in facet loads. 169

192 Extension Flexion Facet Load Components (N) across L4/L5 and Percent Changes Compared to Intact at L4/L5 for 400N FL and 6 N-m Applied Moment for Wide Laminectomy Models L4/L5 Facet Loads (N) Right Left Lateral % Change AP Shear % Change Vertical % Change Lateral % Change AP Shear % Change Vertical % Change Intact DestabL mmL mmL SupportL Intact DestabL mmL mmL SupportL Table E.11: Facet load components in the lateral, anterior-posterior shear, and vertical directions and the percent change as compared to intact for artificial facet FE models with wide laminectomy in extension, bending, and rotation for 400N follower load and 6 N-m applied moment across L4-L5 right and left facet joints. A positive percent change indicates an increase in facet loads where as a negative percent change indicates a decrease in facet loads. 170

193 Facet Load Components (N) across L5/S1 and Percent Changes Compared to Intact at L5/S1 for 400N FL and 6 N-m Applied Moment for Wide Laminectomy Models L5/S1 Facet Loads (N) Right Left Lateral % Change AP Shear % Change Vertical % Change Lateral % Change AP Shear % Change Vertical % Change Intact DestabL Extension 3mmL mmL SupportL Intact DestabL Flexion 3mmL mmL SupportL Table E.12: Facet load components in the lateral, anterior-posterior shear, and vertical directions and the percent change as compared to intact for artificial facet FE models with wide laminectomy in extension, bending, and rotation for 400N follower load and 6 N-m applied moment across L5-S1 right and left facet joints. A positive percent change indicates an increase in facet loads where as a negative percent change indicates a decrease in facet loads. 171

194 Appendix F Artificial Facet Disc Stresses Introduction Intervertebral disc stresses are believed to be important in the determination and prediction of disc degeneration and failure. Compression on the intervertebral disc results in redistribution of stress within the disc, inducing various biological responses [77]. If the compressive load is great enough to result in the nucleus pressure being larger than the swelling pressure, the volume of the nucleus decreases due to fluid leakage [77]. Delamination, the separation of annular layers, is thought to increase interlaminar shear stresses in the intervertebral disc [78]. Delamination occurs in the presence of high interlaminar stresses, thus, degeneration and disc failure is probable [78]. Disc failure has been associated with activities such as frequent bending and twisting, heavy physical work, and vibrational exposure, according to current literature [78]. All of the listed activities result in higher than normal disc stresses [78] suggesting increased disc stresses, as compared to intact, may cause disc degeneration. Normal estimated stress magnitudes found within the disc are 0.1 to 0.3 MPa, but increase to as high as 1.0 to 3.0 MPa under very large loading conditions [78]. With a pressure transducer inserted through a cadaver disc, Edwards et al was able to quantify 172

195 173 disc stresses in flexion and extension with 1000N compressive load [79]. Stresses as great as 2.91 ± 1.33 MPa was found in the posterior annulus in extension with compression [79]. In the Zander et al FE study previously described in chapter two, disc stresses were reported to change as a result of different grades of facetectomies and laminectomies [30]. Maximum von Mises stresses in the annulus occurred in models of bilateral and two-level laminectomy, as compared to intact, and increased more when the disc was degenerated such that maximum von Mises stress was greater than 3.5 MPa in flexion (intact stress was approximately 1.5 MPa) [30]. These larger annular stresses may result in disc degeneration at the affected level [30]. Disc degeneration at one level may also result in changes in adjacent levels shown in an FE study performed by Kim et al, which is also a clinically observed event [80]. In the Kim et al study described in more detail in chapter two, the L4-L5 disc was degenerated and von Mises stresses increased by up to 6% in the L3-L4 annulus [19]. Stresses in the annulus fibrosus and nucleus pulposus can be predicted by the FE method used in the current study for intact, destabilized, and artificial facet models. It is important to understand the effect of artificial facet implants on the annular and nuclear disc stresses. If stress increases in the discs, it may be likely that disc degeneration will be caused by the implant and may lead to future pain for the patient. Stresses similar to those predicted in the intact model are desirable and predicted stresses in all models are shown below.

196 174 Disc Stress Results Intervertebral disc stresses in the nucleus and annulus predicted by the artificial facet finite element models for L3-L4, L4-L5, and L5-S1 discs are presented compared to intact and destabilized condition with wide laminectomy. Two loading conditions were simulated for the FE models resulting in peak von Mises disc stress values and stress contours: 6 N-m bending moment and 400N follower load with 6 N-m bending moment. The stress contours are found in Figures F.1 to 6 and tables giving maximum von Mises values and percent changes compared to intact are in Tables F.1 to 4. Stresses were also found to help determine the effectiveness of an artificial facet in stabilizing the spine with a wide laminectomy. The wide laminectomy model stress contour results for 6 N-m and 400N follower load with 6 N-m applied moment are shown in Figures F.7 to 12 and peak von Mises stress values are given in Tables F.5 to 8. All percent changes reported are compared to the intact model stress predictions. Artificial Facet Caps Stresses in general increased in the nucleus due to capping the L4-L5 facet joint. The L3- L4 nucleus experienced increased stress in all loading modes regardless of follower load application except in rotation where stress decreased as compared to intact. Greater increases in stress were predicted for the L4-L5 nucleus than at other levels. Stress increased by 10.3, 3.2, 1.0, and 18.5% in extension, flexion, bending, and rotation NFL, respectively and 17.5 and 1.1% in extension and flexion FL. The L5-S1 nucleus also experienced greater stress than in intact in extension and flexion NFL and FL, but stress decreased in bending and rotation.

197 175 Capping the facet joint resulted in changed annular disc stresses. Annular stresses increased in all loading modes but bending in the L3-L4 disc. The L4-L5 disc annulus experienced increased stresses in all loading modes regardless of follower load application, greater than the other levels. Stress increased by 11.1, 4.6, 0.5, and 16.6% for extension, flexion, bending, and rotation NFL, and 13.8 and 0.7% in extension and flexion FL. In extension and flexion NFL and FL stress increased at the L5-S1 annulus. However, stresses decreased in bending and rotation at L5-S1. Artificial Facet Caps with Screws Securing the caps with screws did not change the disc stresses greatly from the cap design. The L3-L4 disc nucleus was stressed the same as the cap model without screws. Slight changes were found in the L4-L5 nucleus. In flexion and rotation, stress increased by 3.3 and 18.6% respectively with no follower load, and 17.7% in extension with follower load in the L4-L5 nucleus. In extension NFL and bending, stresses were the same as the cap model. Stress in the L5-S1 nucleus only changed from the cap model in extension FL where an increase was predicted. The annular stresses were also very similar to stresses found in the cap without screws model. Differences occurred in L4-L5 annulus in rotation and extension FL (increase 16.7 and 13.9%, respectively) and in the L5-S1 annulus in rotation (decreased) and extension FL (increased).

198 176 3mm Pedicle Screw Based Artificial Facet The 3mm stem pedicle screw based design changed disc stresses from intact in all modes. In the nucleus at L3-L4, stresses increased in all loading modes but bending. The L4-L5 nucleus experienced increased stress by 11.9, 0.1, and 18.5% in extension, bending, and rotation NFL, respectively, and 27.3% in extension FL. While in flexion, stress decreased by 0.6 and 0.4% with no follower load and with follower load, respectively. In the L5-S1 nucleus, stress decreased in all loading modes but rotation. The annulus also experienced changed von Mises stresses as compared to intact. The L3- L4 disc experienced an increase in stress in all loading modes regardless of follower load application. Stress increased in the L4-L5 annulus by 12.3 and 23.8% in extension and rotation NFL, respectively and 17.4 and 0.1% in extension and flexion with follower load. A decrease in stress was experienced in flexion NFL by 0.3% and bending by 0.4%. Most loading modes resulted in a decrease in annular stress in the L5-S1 disc, but stress increased in rotation and flexion FL. 5mm Pedicle Screw Based Artificial Facet Thickening the pedicle screw based design stem to 5mm did not change the stress distribution as compared to the 3mm design greatly. The L3-L4 and L5-S1 nucleuses experienced the same stress changes as compared to intact as the 3mm design did. The L4-L5 nucleus was different from the 3mm design stress distribution in extension NFL (11.6% increase from intact), rotation (18.3% increase), and extension FL (26.4% increase).

199 177 The annular stresses due to the 5mm design were also very similar to the 3mm design. The L3-L4 annulus experienced the same stress changes from intact as the 3mm design. In the L4-L5 annulus, all loading modes, but extension, regardless of follower load application, were the same as the 3mm design. In extension NFL stress increased by 12.0% and FL by 16.9% in the L4-L5 annulus. L5-S1 experienced the same stress results as the 3mm design but a slight decrease in extension NFL and rotation. Pedicle Screw Based Artificial Facet with Pedicle Screw Support When the support around the pedicle screw was added, disc stresses were more similar to intact than with the other designs. The L3-L4 nucleus experienced stresses increased slightly in all loading modes with and without follower load except in bending where stress decreased slightly from intact. The L4-L5 nucleus increased in stress in all loading modes but flexion NFL and FL. In extension NFL stress increased by 5.3%, 0.2% in bending, 21.1% in rotation, and 6.9% in extension FL in the L4-L5 nucleus. In flexion, stress decreased by 0.6% NFL and 0.4% FL. The L5-S1 experienced a decrease in nuclear stresses in all modes but rotation (1.0% increase). Annular stresses also increased over intact in L3-L4 in all loading modes regardless of follower load application. The L4-L5 annulus experienced an increase in stress from intact for all loading modes but flexion NFL. Stress increased by 5.5, 0.4, 17.1, 7.2, and 0.1% in extension, bending, and rotation NFL, and extension and flexion FL. Stress decreased in flexion NFL by 0.2% for the L4-L5 annulus as compared to intact. L5-S1 annular stresses decreased from intact in all modes by rotation and flexion FL.

200 178 Destabilized Destabilization of the lumbar spine by performing a wide laminectomy at L4-L5 resulted in large stress changes in the L4-L5 disc as compared to intact, but little change in the stresses at L3-L4 and L5-S1 except in flexion. The von Mises stress experienced by the L3-L4 nucleus increased in all loading modes, but bending did not result in a change in stress. The L4-L5 nuclear stresses increased greatly in flexion NFL and FL. Stress increased by 12.8 and 106.9% for extension and flexion NFL and 20.8 and 53.7% in extension and flexion FL. Stress decreased from intact in L4-L5 nucleus by 0.5 and 6.9% in bending and rotation. L5-S1 nuclear stress decreased in extension NFL and rotation, but increased in all other loading modes. Annular stresses also increased dramatically in flexion at all levels, especially L4-L5. Stress increased from intact in the L3-L4 annulus in all bending modes with and without follower load, but lateral bending did not change. In extension NFL stress increased by 11.0%, 114.4% in flexion NFL, 11.9% in rotation, in extension FL by 7.3%, and 40.2% in flexion FL, in the L4-L5 annulus. In bending, stress decreased by 1.4% at L4-L5. Annular stress increased at L5-S1 in all modes but extension FL where a slight decrease occurred. 3mm Pedicle Screw Based Artificial Facet with Wide Laminectomy Stresses in the disc changed due to implantation of the 3mm stem pedicle screw based artificial facet in the destabilized model. Stresses increased in all loading modes but bending where a slight decrease occurred in the L3-L4 nucleus. All loading modes resulted in increased stress as compared to intact in the L4-L5 nucleus. Stress increased

201 179 by 12.3, 130.9, 0.1, 19.7, 28.0, and 58.6%, in extension, flexion, bending, and rotation with no follower load, and extension and flexion with follower load, respectively. L5-S1 nuclear stresses increased in flexion NFL and FL and rotation and decreased in extension NFL and FL and bending. Annular stresses increased in all loading modes in the L3-L4 annulus regardless of follower load application. An increase in stress was also experienced by the L4-L5 annulus by 12.6, 138.6, 24.8, 17.8, and 46.8% in extension, flexion, and rotation NFL and extension and flexion FL. Bending resulted in a slight decrease of 0.4% at L4-L5. Stress increased in the L5-S1 annulus in flexion NFL and FL and rotation, but decreased in extension NFL and FL and bending. 5mm Pedicle Screw Based Artificial Facet with Wide Laminectomy Changing the stem thickness to 5mm in the wide laminectomy model resulted in similar maximum stresses in the nucleus and annulus as the 3mm design. The stresses experienced at the L3-L4 and L5-S1 discs were the same as in the 3mm design. All loading modes but extension also resulted in the same stress values as the 3mm design in the L4-L5 disc. The nuclear stresses increased by 12.0 and 27.2% in extension NFL and FL and annular stresses increased by 12.4 and 17.3% in extension NFL and FL in L4-L5. Pedicle Screw Based Artificial Facet with Pedicle Screw Support with Wide Laminectomy Increasing the support on the pedicle screw based artificial facet design in the wide laminectomy resulted in stress changes similar to those of the 3mm and 5mm designs. Stress in the nucleus increased in all loading modes but bending where a slight decrease

202 180 occurred in the L3-L4 nucleus. The L4-L5 nuclear stresses increased in all loading modes by 5.8, 130.9, 0.2, 23.2, 8.1, 58.6% in extension, flexion, bending, and rotation NFL, and extension and flexion FL. Nuclear stresses increased in flexion NFL and FL and rotation, but decreased in extension NFL and FL and bending in the L5-S1 nucleus. Annular stresses increased in all loading modes in the L3-L4 annulus. Stress increased by 6.2, 138.7, 0.4, 18.8, 7.8, and 46.8% in extension, flexion, bending, and rotation with no follower load and in extension and flexion with follower load, respectively. Annular stresses increased in flexion NFL and FL and rotation in the L5-S1 annulus and decreased in extension NFL and FL and bending. P A Intact 0N + 6 N-m Cap 0N + 6 N-m Cap Screw 0N + 6 N-m 3mm 0N + 6 N-m 5mm 0N + 6 N-m Support 0N + 6 N-m Figure F.1: L3-L4, L4-L5, and L5-S1 intervertebral disc stress contours (MPa) of the nucleus pulposus and annulus fibrosus for each artificial facet and intact FE model with 6 N-m applied moment in extension.

203 181 P A Intact 0N + 6 N-m Cap 0N + 6 N-m Cap Screw 0N + 6 N-m 3mm 0N + 6 N-m 5mm 0N + 6 N-m Support 0N + 6 N-m Figure F.2: L3-L4, L4-L5, and L5-S1 intervertebral disc stress contours (MPa) of the nucleus pulposus and annulus fibrosus for each artificial facet and intact FE model with 6 N-m applied moment in flexion. P A Intact 0N + 6 N-m Cap 0N + 6 N-m Cap Screw 0N + 6 N-m 3mm 0N + 6 N-m 5mm 0N + 6 N-m Support 0N + 6 N-m Figure F.3: L3-L4, L4-L5, and L5-S1 intervertebral disc stress contours (MPa) of the nucleus pulposus and annulus fibrosus for each artificial facet and intact FE model with 6 N-m applied moment in lateral bending.

204 182 P A Intact 0N + 6 N-m Cap 0N + 6 N-m Cap Screw 0N + 6 N-m 3mm 0N + 6 N-m 5mm 0N + 6 N-m Support 0N + 6 N-m Figure F.4: L3-L4, L4-L5, and L5-S1 intervertebral disc stress contours (MPa) of the nucleus pulposus and annulus fibrosus for each artificial facet and intact FE model with 6 N-m applied moment in axial rotation. P A Intact 400N + 6 N-m Cap 400N + 6 N-m Cap Screw 400N + 6 N-m 3mm 400N + 6 N-m 5mm 400N + 6 N-m Support 400N + 6 N-m Figure F.5: L3-L4, L4-L5, and L5-S1 intervertebral disc stress contours (MPa) of the nucleus pulposus and annulus fibrosus for each artificial facet and intact FE model with 400N follower load and 6 N-m applied moment in extension.

205 183 P A Intact 400N + 6 N-m Cap 400N + 6 N-m Cap Screw 400N + 6 N-m 3mm 400N + 6 N-m 5mm 400N + 6 N-m Support 400N + 6 N-m Figure F.6: L3-L4, L4-L5, and L5-S1 intervertebral disc stress contours (MPa) of the nucleus pulposus and annulus fibrosus for each artificial facet and intact FE model with 400N follower load and 6 N-m applied moment in flexion.

206 184 Nucleus Pulposus Peak von Mises Stresses (MPa) and Percent Changes as Compared to Intact for 6 N-m Moment L3/L4 % Change L4/L5 % Change L5/S1 % Change Intact Cap Cap Screw Extension 3mm mm Support Intact Cap Flexion Cap Screw mm mm Support Intact Cap Bending Cap Screw mm mm Support Intact Cap Rotation Cap Screw mm mm Support Table F.1: L3-L4, L4-L5, and L5-S1 intervertebral disc peak von Mises stress values (MPa) and percent changes as compared to intact of the nucleus pulposus for each artificial facet and intact FE model with 6 N-m applied moment in extension, flexion, bending, and rotation. A positive percent change indicates an increase in stresses where as a negative percent change indicates a decrease in stresses as compared to intact.

207 185 Nucleus Pulposus Peak von Mises Stresses (MPa) and Percent Changes as Compared to Intact for 400N FL and 6 N-m Moment L3/L4 % Change L4/L5 % Change L5/S1 % Change Intact Cap Cap Screw Extension 3mm mm Support Intact Cap Flexion Cap Screw mm mm Support Table F.2: L3-L4, L4-L5, and L5-S1 intervertebral disc peak von Mises stress values (MPa) and percent changes as compared to intact of the nucleus pulposus for each artificial facet and intact FE model with 400N follower load and 6 N-m applied moment in extension and flexion. A positive percent change indicates an increase in stresses where as a negative percent change indicates a decrease in stresses as compared to intact.

208 186 Annulus Fibrosus Peak von Mises Stresses (MPa) and Percent Changes as Compared to Intact for 6 N-m Moment L3/L4 % Change L4/L5 % Change L5/S1 % Change Intact Cap Cap Screw Extension 3mm mm Support Intact Cap Flexion Cap Screw mm mm Support Intact Cap Bending Cap Screw mm mm Support Intact Cap Rotation Cap Screw mm mm Support Table F.3: L3-L4, L4-L5, and L5-S1 intervertebral disc peak von Mises stress values (MPa) and percent changes as compared to intact of the annulus fibrosus for each artificial facet and intact FE model with 6 N-m applied moment in extension, flexion, bending, and rotation. A positive percent change indicates an increase in stresses where as a negative percent change indicates a decrease in stresses as compared to intact.

209 187 Annulus Fibrosus Peak von Mises Stresses (MPa) and Percent Changes as Compared to Intact for 400N FL and 6 N-m Moment L3/L4 % Change L4/L5 % Change L5/S1 % Change Intact Cap Cap Screw Extension 3mm mm Support Intact Cap Flexion Cap Screw mm mm Support Table F.4: L3-L4, L4-L5, and L5-S1 intervertebral disc peak von Mises stress values (MPa) and percent changes as compared to intact of the annulus fibrosus for each artificial facet and intact FE model with 400N follower load and 6 N-m applied moment in extension and flexion. A positive percent change indicates an increase in stresses where as a negative percent change indicates a decrease in stresses as compared to intact. P A Intact 0N + 6 N-m DestabL 0N + 6 N-m 3mmL 0N + 6 N-m 5mmL 0N + 6 N-m SupportL 0N + 6 N-m Figure F.7: L3-L4, L4-L5, and L5-S1 intervertebral disc stress contours (MPa) of the nucleus pulposus and annulus fibrosus for each artificial facet with wide laminectomy and intact FE model with 6 N-m applied moment in extension.

210 188 P A Intact 0N + 6 N-m DestabL 0N + 6 N-m 3mmL 0N + 6 N-m 5mmL 0N + 6 N-m SupportL 0N + 6 N-m Figure F.8: L3-L4, L4-L5, and L5-S1 intervertebral disc stress contours (MPa) of the nucleus pulposus and annulus fibrosus for each artificial facet with wide laminectomy and intact FE model with 6 N-m applied moment in flexion. P A Intact 0N + 6 N-m DestabL 0N + 6 N-m 3mmL 0N + 6 N-m 5mmL 0N + 6 N-m SupportL 0N + 6 N-m Figure F.9: L3-L4, L4-L5, and L5-S1 intervertebral disc stress contours (MPa) of the nucleus pulposus and annulus fibrosus for each artificial facet with wide laminectomy and intact FE model with 6 N-m applied moment in lateral bending.

211 189 P A Intact 0N + 6 N-m DestabL 0N + 6 N-m 3mmL 0N + 6 N-m 5mmL 0N + 6 N-m SupportL 0N + 6 N-m Figure F.10: L3-L4, L4-L5, and L5-S1 intervertebral disc stress contours (MPa) of the nucleus pulposus and annulus fibrosus for each artificial facet with wide laminectomy and intact FE model with 6 N-m applied moment in axial rotation. P A Intact 400N + 6 N-m DestabL 400N + 6 N-m 3mmL 400N + 6 N-m 5mmL 400N + 6 N-m SupportL 400N + 6 N-m Figure F.11: L3-L4, L4-L5, and L5-S1 intervertebral disc stress contours (MPa) of the nucleus pulposus and annulus fibrosus for each artificial facet with wide laminectomy and intact FE model with 400N follower load and 6 N-m applied moment in extension.

212 190 P A Intact 400N + 6 N-m DestabL 400N + 6 N-m 3mmL 400N + 6 N-m 5mmL 400N + 6 N-m SupportL 400N + 6 N-m Figure F.12: L3-L4, L4-L5, and L5-S1 intervertebral disc stress contours (MPa) of the nucleus pulposus and annulus fibrosus for each artificial facet with wide laminectomy and intact FE model with 400N follower load and 6 N-m applied moment in flexion.