Alignment in TKA using a Simulated Oxford Rig. Thesis. Julie Ann Thompson. Graduate Program in Mechanical Engineering. The Ohio State University

Transcription

1 The Biomechanical Effects of Variability in Femoral and Tibial Component Rotational Alignment in TKA using a Simulated Oxford Rig Thesis Presented in Partial Fulfillment of the Requirements for the Degree Master of Science in the Graduate School of The Ohio State University By Julie Ann Thompson Graduate Program in Mechanical Engineering The Ohio State University 2009 Thesis Committee: Robert A. Siston, Advisor Ajit M.W. Chaudhari

2 Copyright by Julie Ann Thompson 2009

3 Abstract The success of total knee arthroplasty (TKA) depends on many factors, but malrotation of the prosthetic components, in particular, is a major cause of patellofemoral complications and can lead to revision surgery. Significant variability can be associated with femoral and tibial component rotational alignment, but how this variability translates into functional outcome remains unknown. The purpose of this thesis was to determine the biomechanical effects of variability in femoral and tibial component rotational alignment in TKA using a forward-dynamic computer model of an Oxford Rig, which simulates flexed-knee stance, such as occurs when riding a bicycle, rising from a chair, or climbing stairs. To become familiar with the computer modeling environment, we performed a study of hip kinematics in OpenSim, an open-source software package that was developed for the purpose of creating and analyzing musculoskeletal models and dynamic simulations of movement. The simulations commonly use motion capture data as input, but frequently parameters such as the degrees of freedom (DOF) of certain joints is chosen by the user and may not match the same DOF used by the motion-capture software. OpenSim computes kinematics using a least squares approach to minimize the difference between experimental marker location and virtual markers on the model while maintaining joint constraints. We used a simple model, looking only at the motion of the ii

4 hip, to investigate how marker weights and choice of model degrees of freedom affect kinematics and to compare the simulated kinematics with the same results from a common motion capture analysis technique. We found that high pelvis marker weights and a 6 degree-of-freedom hip model resulted in kinematics that most closely matched the results from motion capture, but large translations of the femoral head were present. We then used a forward-dynamic model of an Oxford Rig to perform a parametric study on the effects of variations in component rotational alignment in TKA. The femoral component rotational alignment was varied from 15 internal rotation to 15 external rotation in 5 increments and the tibial component rotational alignment was varied from 20 internal rotation to 20 external rotation in 5 increments. The effects of component rotational alignment on knee kinematics, quadriceps muscle force, ligament forces, and contact forces were analyzed for the cruciate-retaining and posteriorsubstituting versions of the Scorpio implant from Stryker Orthopaedics. We found that femoral component alignment, in general, had a much greater effect on our variables of interest than tibial component alignment or choice of implant design. Internal rotation of the femoral component led to a reversal of the natural screw-home motion of the knee (internal rotation of the tibia with respect to the femur during flexion) as well as high MCL force, quadriceps muscle force, and contact forces between the femoral and tibial components. Our findings suggest variability in component rotational alignment, especially internal femoral component alignment, may impact post-operative performance and further emphasize the need to accurately establish rotational alignment. iii

5 Acknowledgments I have many people to thank for helping me to get to this next stage of my education. First and foremost, I need to thank my advisor, Rob Siston. It has been almost 3 years now since I first stepped into one of his classes, and I have learned more in those 3 years than I could have imagined. I am grateful for his guidance and confidence in me and for always challenging me to do better. I know I have so much more to learn and I look forward to the journey ahead. I also want to thank the other member of my defense committee, Ajit Chaudhari. This is the second time he has been on my defense committee and I deeply appreciate all of his help and feedback. He has been an excellent resource and I look forward to continuing to work with him in the future. The work in this thesis would not have been possible without Stephen Piazza and his grad student Mike Hast at Penn State University. They not only developed the Oxford Rig computer model and allowed me to use it for my research, but they have also provided me with an endless amount of support, advice, suggestions and feedback via s, g-chats, phone conversations, and lab visits. They are both incredible engineers and it has been a joy to work with them. I am excited to see what new and exciting paths our collaboration may take us down. iv

6 I am also grateful to work with many wonderful students right here in our own lab. Specifically, I want to thank Erin Hutter for helping me to run some of the many, many computer simulations so I didn t have to do them all by myself. I also need to give a huge thanks to Becky Lathrop. She is not only one of the brightest people I have ever met, she is also one of the kindest. Becky has always been more than willing to offer help and advice with my work as well as to listen to my concerns and frustrations. I am really blessed to be able to work with her. She is a true inspiration for me. I am also thankful for my other NMBL labmates, who have all offered plenty of encouragement and advice throughout this process. And last, but most importantly, I thank my wonderful family and friends. Their unconditional love and support has pulled me through so many tough times and my life would be nothing without them. Their constant encouragement has inspired me to try harder, reach further, and accomplish things I never could have imagined for myself. Finally, I would like to acknowledge the National Science Foundation for providing me with funding through their Graduate Research Fellowship program. v

7 Vita June Grove City High School June B.S. Mechanical Engineering, The Ohio State University September 2008 to present...national Science Foundation Graduate Research Fellow Major Field: Mechanical Engineering Fields of Study vi

8 Table of Contents Abstract... ii Acknowledgments... iv Vita... vi List of Tables... viii List of Figures... ix Chapter 1: Introduction Focus of Thesis Significance of Research Overview of Thesis...9 Chapter 2: Hip Kinematics Simulation Study Introduction Methods Results Discussion Chapter 3: Oxford Rig Simulation Study Introduction Methods Results Variability in Femoral Component Alignment Variability in Tibial Component Alignment Variability in Both Femoral and Tibial Component Alignments Discussion Chapter 4: Conclusion Contributions Additional Applications Future Work Summary References vii

9 List of Tables Table 1: Summary of the effect of marker weight and model degrees-of-freedom on peak hip angles viii

10 List of Figures Figure 1: Markers used for scaling and inverse kinematics Figure 2: Hip flexion angle Figure 3: Hip adduction angle Figure 4: Hip translations Figure 5: The Oxford Knee-Testing Rig Figure 6: Oxford rig computer model Figure 7: Femoral component alignment Figure 8: Tibial component alignment Figure 9: Effect of femoral component rotational alignment on internal/external rotation angle for the CR implant design Figure 10: Effect of femoral component rotational alignment on varus/valgus angle for the CR implant design Figure 11: Effect of femoral component rotational alignment on patellar tilt angle for the CR implant design Figure 12: Effect of femoral component rotational alignment on patellar medial/lateral (M/L) translation for the CR implant design Figure 13: Effect of femoral component rotational alignment on patellar superior/inferior (S/I) location for the CR implant design Figure 14: Effect of femoral component rotational alignment on MCL force for the CR and PS implant designs Figure 15: Effect of femoral component rotational alignment on LCL force for the CR and PS implant designs Figure 16: Effect of femoral component rotational alignment on PCL force for the CR implant design Figure 17: Effect of femoral component rotational alignment on patellar ligament force for the CR implant design Figure 18: Effect of femoral component rotational alignment on quadriceps muscle force for the CR implant design Figure 19: Effect of femoral component rotational alignment on TF contact force at various knee flexion angles for the CR implant design Figure 20: Effect of femoral component rotational alignment on PF contact force at various knee flexion angles for the CR implant design Figure 21: Effect of femoral component rotational alignment on cam-post contact force at various knee flexion angles for the PS implant design Figure 22: Effect of tibial component rotational alignment on internal/external tibiofemoral rotation angle for the CR implant design ix

11 Figure 23: Effect of tibial component rotational alignment on varus/valgus angle for the CR implant design Figure 24: Effect of tibial component rotational alignment on varus/valgus angle for the PS implant design Figure 25: Effect of tibial component rotational alignment on patellar tilt for the CR implant design Figure 26: Effect of tibial component rotational alignment on patellar medial/lateral (M/L) location for the CR implant design Figure 27: Effect of tibial component rotational alignment on PCL force for the CR implant design Figure 28: Effect of tibial component rotational alignment on quadriceps muscle force for the CR implant design Figure 29: Effect of tibial component rotational alignment on TF contact force for the CR implant design Figure 30: Effect of tibial component rotational alignment on TF contact force for the PS implant design Figure 31: Effect of tibial component rotational alignment on PF contact force for the CR implant design Figure 32: Effect of tibial component rotational alignment on PF contact force for the PS implant design Figure 33: Effect of tibial component rotational alignment on CP contact force for the PS implant design Figure 34: Effect of femoral and tibial component rotational alignment on screw-home motion for the CR implant design Figure 35: Effect of femoral and tibial component rotational alignment on screw-home motion for the PS implant design Figure 36: Effect of femoral and tibial component rotational alignment on I/E rotation angle at 20 knee flexion for the CR implant design Figure 37: Effect of femoral and tibial component rotational alignment on I/E rotation angle at 70 knee flexion for the CR implant design Figure 38: Effect of femoral and tibial component rotational alignment on I/E rotation angle at 120 knee flexion for the CR implant design Figure 39: Effect of femoral and tibial component rotational alignment on I/E rotation angle at 120 knee flexion for the PS implant design Figure 40: Effect of femoral and tibial component rotational alignment on varus/valgus angle at 20 knee flexion for the CR implant design Figure 41: Effect of femoral and tibial component rotational alignment on varus/valgus angle at 120 knee flexion for the CR implant design Figure 42: Effect of femoral and tibial component rotational alignment on patellar tilt at 70 knee flexion for the CR implant design x

12 Figure 43: Effect of femoral and tibial component rotational alignment on patellar M/L location at 20 knee flexion for the CR implant design Figure 44: Effect of femoral and tibial component rotational alignment on patellar M/L location at 120 knee flexion for the CR implant design Figure 45: Effect of femoral and tibial component rotational alignment on patellar S/I location at 70 knee flexion for the CR implant design Figure 46: Effect of femoral and tibial component rotational alignment on MCL force at 90 knee flexion for the CR implant design Figure 47: Effect of femoral and tibial component rotational alignment on MCL force at 120 knee flexion for the PS implant design Figure 48: Effect of femoral and tibial component rotational alignment on LCL force at 20 knee flexion for the CR implant design Figure 49: Effect of femoral and tibial component rotational alignment on LCL force at 20 knee flexion for the PS implant design Figure 50: Effect of femoral and tibial component rotational alignment on PCL force at 120 knee flexion for the CR implant design Figure 51: Effect of femoral and tibial component rotational alignment on patellar ligament force at 30 knee flexion for the CR implant design Figure 52: Effect of femoral and tibial component rotational alignment on patellar ligament force at 60 knee flexion for the CR implant design Figure 53: Effect of femoral and tibial component rotational alignment on patellar ligament force at 90 knee flexion for the CR implant design Figure 54: Effect of femoral and tibial component rotational alignment on patellar ligament force at 120 knee flexion for the CR implant design Figure 55: Effect of femoral and tibial component rotational alignment on patellar ligament force at 30 knee flexion for the PS implant design Figure 56: Effect of femoral and tibial component rotational alignment on patellar ligament force at 60 knee flexion for the PS implant design Figure 57: Effect of femoral and tibial component rotational alignment on patellar ligament force at 90 knee flexion for the PS implant design Figure 58: Effect of femoral and tibial component rotational alignment on patellar ligament force at 120 knee flexion for the PS implant design Figure 59: Effect of femoral and tibial component rotational alignment on quadriceps force at 30 knee flexion for the CR implant design Figure 60: Effect of femoral and tibial component rotational alignment on quadriceps force at 60 knee flexion for the CR implant design Figure 61: Effect of femoral and tibial component rotational alignment on quadriceps force at 90 knee flexion for the CR implant design Figure 62: Effect of femoral and tibial component rotational alignment on quadriceps force at 120 knee flexion for the CR implant design xi

13 Figure 63: Effect of femoral and tibial component rotational alignment on quadriceps force at 30 knee flexion for the PS implant design Figure 64: Effect of femoral and tibial component rotational alignment on quadriceps force at 60 knee flexion for the PS implant design Figure 65: Effect of femoral and tibial component rotational alignment on quadriceps force at 90 knee flexion for the PS implant design Figure 66: Effect of femoral and tibial component rotational alignment on quadriceps force at 120 knee flexion for the PS implant design Figure 67: Effect of femoral and tibial component rotational alignment on TF contact force at 30 knee flexion for the CR implant design Figure 68: Effect of femoral and tibial component rotational alignment on TF contact force at 120 knee flexion for the CR implant design Figure 69: Effect of femoral and tibial component rotational alignment on TF contact force at 30 knee flexion for the PS implant design Figure 70: Effect of femoral and tibial component rotational alignment on TF contact force at 120 knee flexion for the PS implant design Figure 71: Effect of femoral and tibial component rotational alignment on PF contact force at 30 knee flexion for the CR implant design Figure 72: Effect of femoral and tibial component rotational alignment on PF contact force at 120 knee flexion for the CR implant design Figure 73: Effect of femoral and tibial component rotational alignment on PF contact force at 30 knee flexion for the PS implant design Figure 74: Effect of femoral and tibial component rotational alignment on PF contact force at 120 knee flexion for the PS implant design Figure 75: Effect of femoral and tibial component rotational alignment on CP contact force at 60 knee flexion for the PS implant design Figure 76: Effect of femoral and tibial component rotational alignment on CP contact force at 90 knee flexion for the PS implant design Figure 77: Effect of femoral and tibial component rotational alignment on CP contact force at 120 knee flexion for the PS implant design Figure 78: Mechanical Oxford Rig Figure 79: Close-up of femoral and tibal components with attached motion-capture markers on the mechanical Oxford Rig xii

14 Chapter 1: Introduction Total Knee Arthroplasty (TKA) is a surgical procedure which helps to alleviate pain and restore more normal function in patients who suffer from joint diseases such as osteoarthritis. While TKA provides excellent function for the low-demand everyday activities of older patients, there is a need to study how TKA can be improved to allow more active patients to perform high-demand activities that are important to them. It has been shown that many TKA patients value activities such as kneeling, gardening, squatting, and dancing, but most experience difficulty in performing these activities after knee arthroplasty (Weiss et al., 2002). Gill et al. (1997) investigated TKA in patients 55 years old or younger and found that 52% were able to walk unlimited distances and only 42% were able to climb stairs without difficulty. The success of total knee arthroplasty depends on many factors, including prosthesis design, preoperative condition of the joint, surgical technique, and postoperative rehabilitation. Previous research has suggested that error in surgical technique is the most common cause for revision TKA (Stulberg et al., 2002), since incorrect positioning or orientation of the implant can lead to accelerated wear and loosening of the implant as well as suboptimal functional performance. When the prosthetic components are positioned incorrectly, the bone experiences abnormal loading (Dorr and Boiardo, 1986), which can lead to failure at the bone-cement interface and 1

15 result in component loosening. An investigation of frontal plane alignment by Jeffery et al. (1991) found that alignment errors of greater than 3 are associated with component loosening. It has been shown that a posterior tilt of the tibial component between 5 and 10 is ideal for providing a more normal range of motion in cruciate-retaining knee replacement (Dorr and Boiardo, 1986; Garg and Walker, 1990), but a posterior tilt in posterior-substituting knee replacement may not produce the same beneficial effects (Piazza et al., 1998). Poor rotational alignment of the femoral component is a major cause of patellofemoral complications (Clayton and Thirupathi, 1982; Dennis et al., 1992; Huo and Sculco, 1990), the most common postoperative problem of TKA and one of the major causes for revision surgery (Clayton and Thirupathi, 1982; Dennis et al., 1992; Huo and Sculco, 1990). Rotation of the femoral component of 5 from the transepicondylar axis has been reported to alter tibiofemoral kinematics and patellar tracking and increase patellar shear force (Miller et al., 2001). A study by Anouchi et al. (1993) found that external rotation of the femoral component of 5 from the posterior femoral condyles resulted in knee stability and patellar tracking that most closely represented the stability and tracking characteristics of the intact knee. Rhoads et al. (1990) found that a 10 internal rotation of the femoral component altered patellar tracking, but a 10 external rotation produced the most normal pattern of patellar tracking. Berger et al. (1998) found a direct correlation of combined femoral and tibial component internal rotation to the severity of patellofemoral complications. Combined internal rotation of 1-4 was associated with lateral tracking and tilting of the patella while internal rotation of 7-17 was associated with patellar dislocation and failure 2

16 (Berger et al., 1998). Ranawat et al. (1986) concluded that patellar dislocation, subluxation, tilt, and excessive wear result from malrotation of the femoral and tibial components. Poor rotational alignment of the tibial component also leads to complications. Dalury et al. (2001) reported that internal or external rotation of the tibial component will lead to impingement of the polyethylene, which could lead to loss of motion and wear. In a study of patients with anterior knee pain, Barrack et al. (2001) found that the average tibial component alignment was 6.2 internal rotation, compared with 0.4 external rotation for the control group without anterior knee pain. The study also found that patients with combined component internal rotation were more than five times as likely to experience anterior knee pain after TKA compared with those patients with combined component external rotation (Barrack et al., 2001). It has been shown that there is significant variability associated with femoral component rotational alignment (Jenny and Boeri, 2004; Kinzel et al., 2005; Robinson et al., 2006; Siston et al., 2008; Siston et al., 2005) as well as tibial component rotational alignment (Akagi et al., 2005; Eckhoff et al., 1995; Ikeuchi et al., 2007; Siston et al., 2006). Siston et al. (2005) found alignment errors for femoral component rotation ranging from 13 internal rotation to 16 external rotation. A subsequent study of variability in tibial component rotational alignment found standard deviations as high as 28.1 (Siston et al., 2006). Although past studies have investigated the effect of varying femoral and tibial component rotational alignment on joint stability and patellofemoral and tibiofemoral kinematics, how this variability translates into functional performance still remains unknown. 3

17 One measure of a successful functional outcome after TKA is a patient s ability to perform certain activities. Stair climbing is a functional task that is considered a critical activity of daily living (Jevsevar et al., 1993) and has been studied by previous researchers (Andriacchi et al., 1980; Andriacchi et al., 2003; Banks and Hodge, 2004; Jevsevar et al., 1993; Walsh et al., 1998). This task has also been shown to be more challenging than other activities of daily living, such as level walking (Jevsevar et al., 1993; Walsh et al., 1998). TKA patients walk at significantly slower speeds than their healthy age-matched counterparts, but stair-climbing ability is even more compromised: 51% slower for men and 43% slower for women (Walsh et al., 1998). Other studies have shown that stair-climbing requires a greater range of motion and moment production at the knee than level walking (Andriacchi et al., 1980; Jevsevar et al., 1993). Kneeling and squatting activities have also been studied by researchers (Andriacchi et al., 2003; Argenson et al., 2004; Incavo et al., 2004). Andriacchi et al. (2003) found that substantial external femoral rotation is required to achieve deep flexion. Incavo et al. (2004) found that subjects with TKA can kneel with articulation of the tibia relative to the femur within the intended design constraints of the implants. They concluded that careful and occasional kneeling may not be detrimental to the longevity of the TKA (Incavo et al., 2004). Argenson et al. (2004) found that a knee replacement designed for deep flexion was able to replicate the kinematic patterns of a healthy knee. These activities are essential for daily living in cultures that require deep flexion for social or religious purposes. Even in western cultures there are many recreational activities, such as gardening, that require deep flexion. A study of 176 TKA patients found that many considered kneeling and squatting to be important to their lives, 4

18 but more than ¾ of these patients reported limitations doing these activities because of their knee replacement (Weiss et al., 2002). Despite the importance of these activities to many TKA patients, it remains unknown how variability in the prosthetic component alignment affects these activities and the biomechanical reasons for why many patients have difficulty performing them. The Oxford Rig is a device designed for biomechanical testing of cadaveric knee-joint specimens during simulated flexed-knee stance, such as occurs when riding a bicycle, rising from a chair, or climbing stairs (Zavatsky, 1997). However, it has never been used to investigate the effects of variations in component alignment in TKA. A powerful tool for investigating cause-effect relationships, such as how prosthetic component alignment affects functional performance, is a forward-dynamic computer simulation. Simulations can be used to help understand muscle activations (Anderson and Pandy, 2001), kinematics (Anderson and Pandy, 2001) and internal joint loading (Bei and Fregly, 2004) and thereby can shed light onto mechanisms of pathological or altered kinematics (Piazza, 2006). Dynamic simulations are well-suited for performing what if? studies (Delp et al., 2007) in which, for example, the alignment of a prosthetic component in the knee can be changed to observe the resulting motion and muscle forces. Forward-dynamic simulations are especially useful for studying functional tasks because they involve the application of forces to produce motions, offering potential insights into the roles played by individual muscles during a task. These simulations also permit monitoring of other variables of interest such as ligament forces and joint contact forces which may affect functional performance and offer insight as to the biomechanical reasons for suboptimal outcomes. 5

19 Previous researchers have developed computer models of TKA to predict knee kinematics (Garg and Walker, 1990; Godest et al., 2000; Piazza et al., 1998) as well as wear patterns in the prosthetic components (Fregly et al., 2005). Forward-dynamic computer simulations have been used to investigate the effect of varying the sagittal plane alignment (tilt) of the tibial component on range of motion (Garg and Walker, 1990; Piazza et al., 1998) as well as to investigate a specific functional task such as stairstepping (Piazza and Delp, 2001) and dynamic knee extension (Lanovaz and Ellis, 2005). However, no study to date has combined these areas of research and used a forwarddynamic computer simulation to investigate the biomechanical effects of varying prosthetic component alignment during a functional task. Establishing ranges of acceptable and unacceptable alignments may provide orthopaedic surgeons with guidelines that may help to ensure a successful postoperative outcome. 1.1 Focus of Thesis The purpose of this thesis was to determine the biomechanical effects of variations in femoral and tibial component rotational alignment in TKA using a forwarddynamic computer model of an Oxford Rig. First, to become familiar with the computer modeling environment, we performed a study of hip kinematics in OpenSim, an opensource software package that was developed for the purpose of creating and analyzing musculoskeletal models and dynamic simulations of movement. The simulations commonly use motion capture data as input, but frequently parameters such as the degrees of freedom (DOF) of certain joints is chosen by the user and may not match the same DOF used by the motion-capture software. OpenSim computes kinematics using a 6

20 least squares approach to minimize the difference between experimental marker location and virtual markers on the model while maintaining joint constraints. We used a simple model, looking only at the motion of the hip, to investigate how marker weights and choice of model degrees of freedom affect kinematics and to compare the kinematics with the same results from a common motion capture analysis technique. We then used a forward-dynamic model of an Oxford Rig to perform a parametric study on the effects of variations in femoral and tibial component rotational alignment in TKA. The femoral component rotational alignment was varied from 15 internal rotation to 15 external rotation in 5 increments and the tibial component rotational alignment was varied from 20 internal rotation to 20 external rotation in 5 increments. The effects of component rotational alignment on knee kinematics, quadriceps muscle force, ligament forces, and contact forces were analyzed for the cruciate-retaining and posterior-substituting versions of the Scorpio implant from Stryker Orthopaedics. 1.2 Significance of Research According to the NIH, 300,000 TKAs were performed in 2003 (2004) and this number is expected to reach 3.48 million in 2030 (Kurtz et al., 2007). The prevalence of arthritis is expected to continue to increase as life expectancy increases (Hootman and Helmick, 2006). Although patient satisfaction is between 90% and 95% for TKA, the large volume of TKA performed means that a small percentage of revisions translate into large costs (Sharkey et al., 2002). The cost of a TKA is around $12,000, not including non-monetary costs such as time and emotional stress (Sharkey et al., 2002). Additionally, as patients receiving TKA become younger, there is an increased need for 7

21 longer use and better functional performance to allow patients to continue to participate in the high-demand activities that are important to them. Patellofemoral complications are the most common postoperative problem and one of the major causes leading to revision surgery (Clayton and Thirupathi, 1982; Dennis et al., 1992; Huo and Sculco, 1990). These complications have been shown to be primarily caused by errors in rotational alignment of the femoral and tibial prosthetic components (Berger et al., 1998; Clayton and Thirupathi, 1982; Dennis et al., 1992; Huo and Sculco, 1990; Ranawat, 1986). The effects of poor alignment on knee stability, kinematics, and range of motion have been well characterized for the femoral component (Anouchi et al., 1993; Dorr and Boiardo, 1986; Miller et al., 2001; Rhoads et al., 1990), but how this translates into functional performance is unknown. Additionally, relatively few studies have investigated the effects of variability in tibial component rotational alignment. It has been shown that internal component rotation is associated with more severe complications than external rotation (Barrack et al., 2001; Berger et al., 1998; Rhoads et al., 1990), but the biomechanical reasons for this as well as the specific values at which deleterious effects may occur is largely unknown. A computer simulation is a useful tool for looking at relationships between malalignment and parameters such as kinematics, ligament forces, and contact forces, as it allows us to make quick changes to the alignment and run many simulations in a controlled environment. A dynamic simulation of movement also has a few important advantages over experimental approaches. One advantage is that a simulation is able to provide estimates of some variables that are not typically measurable in experiments, 8

22 such as muscle and joint forces. Another advantage is that simulations allow cause-effect relationships to be identified, which are difficult to establish from experimental data. No study has extensively investigated the biomechanical effects of variations in femoral and tibial component rotational alignment during a functional task. Understanding what alignments have undesirable effects on kinematics, ligament forces, muscle force, and contact forces can provide a rational basis for suggesting alignments that surgeons should avoid when placing the prosthetic components. In the future, simulations such as this may become invaluable tools for creating patient specific protocols, helping doctors tailor the surgery to the needs of each patient. This thesis is a first step towards achieving that goal. 1.3 Overview of Thesis This thesis has three subsequent chapters. Chapter 2 (presented as an abstract for the 33 rd annual meeting of the American Society of Biomechanics at the Pennsylvania State University) describes an initial experience in the OpenSim environment in which the effects of model degrees of freedom and marker weight on resultant hip kinematics were determined. Chapter 3 (submitted in part as an abstract to the 56 th annual meeting of the Orthopaedic Research Society in New Orleans, Louisiana) presents the results of a collaborative study with Dr. Stephen Piazza and Michael Hast at the Pennsylvania State University. This study determined the biomechanical effects of variability in femoral and tibial component rotational alignment in a simulated Oxford Rig. Chapter 4, the conclusion, summarizes the key contributions of this thesis, discusses the additional applications of this research, and proposes future directions of study. 9

23 Chapter 2: Hip Kinematics Simulation Study 2.1 Introduction Dynamic simulations of movement are useful tools for studying how the neuromuscular and musculoskeletal systems interact to produce coordinated movement. These simulations commonly use motion capture data as input, but frequently the complexity of the model, including parameters such as the degrees of freedom (DOF) of certain joints, is chosen by the user and may not match the same degrees of freedom used by the motion-capture software. OpenSim is an open-source software package that was developed for the purpose of creating and analyzing musculoskeletal models and dynamic simulations of movement (Delp et al., 2007). The software computes kinematics using a least squares approach to minimize the difference between experimental marker location and virtual markers on the model while maintaining joint constraints (Delp et al., 2007). However, this approach presents a challenge when creating musculoskeletal simulations from gait data. Since there is no set of established rules on how to weigh the markers, what is the effect on resultant kinematics of choosing different weights? The purpose of this study was to use a simple model, looking only at the motion of the hip, to investigate how marker weights and choice of model DOF affect kinematics 10

24 and to compare the kinematics with the same results from a common motion capture analysis technique. 2.2 Methods Gait data was obtained for 1 healthy subject while walking at a self-selected speed in the OSU Sports Biomechanics Laboratory using a Vicon motion capture system. The experimental kinematics of the femur segment were estimated using the Point-Cluster Technique (Andriacchi et al., 1998), which uses a redundant cluster of markers on the segment to partially correct for skin deformation (Figure 1-A and 1-B), while the experimental kinematics of the pelvis segment were estimated using 4 markers on the ASIS and PSIS landmarks. A functional hip joint center (fhjc) was obtained by having the subject trace out a star-arc pattern with a straight leg (Camomilla et al., 2006). Using custom scripts in Matlab and Vicon Bodybuilder, joint angles between the femoral coordinate system and the pelvic coordinate system were calculated using a standard Euler method (Mocap results). A generic musculoskeletal model was scaled in OpenSim using static calibration data obtained in the gait laboratory. We systematically adjusted the static pose weights for the markers on the pelvis, left greater trochanter, left femur, and a virtual marker that represented the left fhjc (Figure 1-C). 11

25 A B C Figure 1: (A) Vicon model of cluster of markers placed on subject s thigh. (B) Vicon model of coordinate system (gray spheres) derived from cluster. (C) OpenSim model showing markers used for scaling and inverse kinematics (pink = experimental markers, blue = virtual markers). The inverse kinematics problem was then solved using the same weights as the scaling procedure, including 4 markers that represent the centroid and principal axes of the cluster of 10 femoral markers. 28 solutions of the inverse kinematics problem were calculated; 16 with equal weighting on all markers of 1, 10, 100, or 1000 and 12 with differential weighting. For differential weighting, the pelvis marker weights were fixed at 1 or 1000 while the markers on the leg were varied. 8 of the equal weight trials used the fhjc while the other 8 did not. We used a 3 DOF hip model for 14 of the trials, which is commonly used in OpenSim, then released the translational constraints for the other 14 (6 DOF model). The resultant maximum and minimum hip angles in the sagittal, frontal, and transverse planes (OpenSim results) were compared amongst trials and with the Mocap results. 12

26 2.3 Results Resultant kinematics did not change in a given model when all markers were given equal weighting (Trials 1-16, Table 1). For a given series of trials (e.g. 1-4) all angles were within one-tenth of a degree. However, differential weighting led to large differences. The resultant OpenSim kinematics were closest to the Mocap results when the pelvis markers were weighted highly (Figures 2-3). Hip flexion angles matched most closely when using the fhjc than when not using it, which may have been expected. Table 1: Summary of the effect of marker weight and model degrees-of-freedom on peak hip angles. Trial Hip DOF Hip Flexion Hip Adduction Hip Rotation Marker Weights ( ) ( ) ( ) Pelvis HJC Femur Cluster Max Min Max Min Max Min ,10,100,1000 1,10,100,1000 1,10,100,1000 1,10,100, ,10,100,1000 1,10,100,1000 1,10,100,1000 1,10,100, ,10,100,1000 No 1,10,100,1000 1,10,100, ,10,100,1000 No 1,10,100,1000 1,10,100, Mocap

27 Figure 2: Hip flexion angle for differential marker weighting on 3-degree-of-freedom hip and 6-degree-offreedom hip. Figure 3: Hip adduction angle for differential marker weighting on 3-degree-of-freedom hip and 6-degreeof-freedom hip. 14

28 While using a 6 DOF hip model in OpenSim resulted in kinematics that more closely matched the Mocap results than a 3 DOF hip model, we noticed large translations of the femoral head (Figure 4). The average sagittal plane (anterior/posterior) translation was 1.4 mm, after the femoral head jumped anteriorly about 3 cm. The average frontal plane (superior/inferior) translation was 2.17 cm and the average transverse plane (medial/lateral) translation was 8.9 mm. The 6 DOF hip model with differential weighting and a high pelvis weight resulted in kinematics that most closely matched the Mocap results for hip flexion and adduction. No trial was able to closely replicate internal/external rotation. Figure 4: Hip translations in the sagittal (A/P), frontal (S/I), and transverse (M/L) planes with a 6 DOF model. Smaller negative values correspond to anterior, superior, or medial translation. 15

29 2.4 Discussion The purpose of this study was not to determine whether OpenSim or motion capture provided the right or wrong kinematic result. Instead, this work illustrates the need for researchers to carefully establish marker weights in OpenSim and to select the appropriate model specific to the questions he or she seeks to answer. Since this work only used 1 healthy subject, more work is necessary to determine if these results are generalizable across all healthy subjects or in pathological populations. 16

30 Chapter 3: Oxford Rig Simulation Study 3.1 Introduction The success of total knee arthroplasty (TKA) depends on many factors, but malrotation of the femoral and tibial components is a major cause of patellofemoral complications and can lead to revision surgery (Clayton and Thirupathi, 1982; Dennis et al., 1992; Huo and Sculco, 1990). Rotation of the femoral component of 5 from the transepicondylar axis has been reported to alter tibiofemoral kinematics and patellar tracking and increase patellar shear force (Miller et al., 2001). Anouchi et al. (1993) found that external rotation of the femoral component of 5 from the posterior femoral condyles resulted in knee stability and patellar tracking that most closely represented the stability and tracking characteristics of the intact knee. Rhoads et al. (1990) found that a 10 external rotation produced the most normal pattern of patellar tracking. Dalury (2001) reported that internal or external rotation of the tibial component will lead to impingement of the polyethylene, which could lead to loss of motion and wear. A study of patients with anterior knee pain found that the average tibial component alignment was 6.2 internal rotation compared with 0.4 external rotation for the control group without anterior knee pain (Barrack et al., 2001). Combined femoral and tibial component internal rotation has also been directly correlated to the severity of patellofemoral complications (Berger et al., 1998). 17

31 We have previously discovered that significant variability can be associated with femoral and tibial component rotational alignment, with errors in femoral component alignment ranging from 13º internal rotation to 16º external rotation (Siston et al., 2005) and standard deviations in tibial component alignment of up to 28.1 (Siston et al., 2006). The effects of malalignment on knee stability and kinematics have been well characterized (Anouchi et al., 1993; Garg and Walker, 1990; Miller et al., 2001; Piazza et al., 1998; Rhoads et al., 1990), but how this variability translates into functional performance is not known. It has been shown that internal component rotation is associated with more severe complications than external rotation (Barrack et al., 2001; Berger et al., 1998; Rhoads et al., 1990), but the biomechanical reasons for this as well as the specific values at which deleterious effects may occur is largely unknown. The Oxford Rig is a device designed for biomechanical testing of cadaveric kneejoint specimens during simulated flexed-knee stance, such as occurs when riding a bicycle, rising from a chair, or climbing stairs (Zavatsky, 1997). The movements of the ankle and hip assemblies of the rig combine to allow a knee specimen its full six degrees-of-freedom of movement (Figure 5). 18

32 Figure 5: The Oxford Knee-Testing Rig. The purpose of this study was to determine the effects of femoral and tibial component rotational alignment in TKA on knee kinematics, quadriceps force, ligament forces, and contact forces for the cruciate-retaining and posterior-substituting versions of the Scorpio implant from Stryker Orthopaedics using a forward-dynamic computer model of an Oxford Rig. 3.2 Methods We created a forward-dynamic model of an Oxford Rig, which simulated controlled knee flexion from (Figure 6). The dynamic equations of motion were 19

33 solved using SD/FAST (Symbolic Dynamics, Inc.; Mountain View, CA) and the simulations were visualized in OpenSim. Muscular architecture and limb lengths were based on the model described by Delp (1990), and a 30 kg mass was placed at the pelvis to simulate body weight. The quadriceps muscles were represented by a single muscle acting along the line of action of the vastus intermedius, and the force necessary to lower the pelvis in a controlled manner was determined using a proportional derivative controller. The PCL was modeled as 10 fibers with slack lengths and attachment points that begin to generate forces at 80 of flexion of a natural knee motion (Makino et al., 2006). Contact between the implants was modeled using a rigid body spring model (Li et al., 1997). Knee flexion angles were determined using the convention described by Grood and Suntay (1983). 20

34 A B Figure 6: Oxford rig computer model at (A) 20 of knee flexion and at (B) 120 of knee flexion. Cruciate-retaining (CR) and posterior-substituting (PS) versions of the same implant (Scorpio; Stryker Orthopaedics) were used in the simulations. Femoral rotational alignment was first varied between 15 external rotation and 15 internal rotation (Figure 7) in 5 increments while keeping the tibial component at 0 of rotation. Tibial component alignment was then varied between 20 external rotation and 20 internal rotation (Figure 8) in 5 increments while keeping the femoral component at 0 of rotation. Finally, both the femoral component alignment and tibial component alignment were varied together over all combinations of the alignment ranges mentioned above, for 21

35 a total of 63 simulations per implant design (126 total simulations). The effects of these alignments on knee joint kinematics, quadriceps muscle force, ligament forces, and contact forces were analyzed for the two implant designs. A B C Figure 7: (A) 15 external rotation of the femoral component. (B) 0 (neutral) rotation of the femoral component. (C) 15 internal rotation of the femoral component. A B C Figure 8: (A) 20 external rotation of the tibial component. (B) 0 (neutral) rotation of the tibial component. (C) 20 internal rotation of the tibial component. 22

36 3.3 Results Variability in Femoral Component Alignment Internal/External Knee Rotation Angle The internal/external rotation angle for the CR and PS designs was similarly affected by femoral rotational alignment. Externally aligned components exhibited the so-called screw-home motion (internal rotation of the tibia with respect to the femur during flexion) (Blankevoort et al., 1988) while the internally aligned components reversed this screw-home motion (Figure 9). The internal/external rotation angle value for all alignments converged at approximately 80 flexion and then diverged throughout the rest of flexion. Figure 9: Effect of femoral component rotational alignment on internal/external rotation angle for the CR implant design. 23

37 Varus/Valgus Knee Rotation Angle The varus/valgus angle for the CR and PS designs was similarly affected by femoral rotational alignment. The externally aligned components resulted in a varus alignment throughout flexion while the internally aligned components resulted in a valgus alignment (Figure 10). The magnitude of the varus/valgus angle for all alignments reached a maximum at approximately 60 flexion and converged to neutral alignment at 120 flexion. Figure 10: Effect of femoral component rotational alignment on varus/valgus angle for the CR implant design. Patellar Tilt The patellar tilt for the CR and PS designs was similarly affected by femoral rotational alignment. Internally aligned components had a more medial tilt than 24

38 externally aligned components (Figure 11). For all component alignments, the patella tilted laterally with increasing flexion. Figure 11: Effect of femoral component rotational alignment on patellar tilt angle for the CR implant design. Patellar Medial/Lateral Translation The patellar medial/lateral translation for the CR and PS designs was similarly affected by femoral rotational alignment. The patella was positioned laterally in early flexion and tracked medially with increasing flexion for the externally aligned and neutrally aligned components and was positioned medially throughout flexion for the internally aligned components (Figure 12). The medial/lateral position for all alignments converged at approximately 90 flexion and then diverged throughout the rest of flexion. 25

39 Figure 12: Effect of femoral component rotational alignment on patellar medial/lateral (M/L) translation for the CR implant design. Patellar Superior/Inferior Translation The patellar superior/inferior translation was similar for both the CR and PS designs and was very similar for all component alignments. The patella translated inferiorly until approximately 90 flexion, after which there was no translation (Figure 13). 26

40 Figure 13: Effect of femoral component rotational alignment on patellar superior/inferior (S/I) location with the respect to the femoral coordinate system for the CR implant design. Medial Collateral Ligament (MCL) Force The MCL force for the CR and PS designs was affected differently by femoral rotational alignment. The maximum force corresponded to the 15º internal alignment for both designs, with values of approximately 475 N at 90 of flexion for the CR knee and approximately 700 N at 120 of flexion for the PS knee (Figure 14). There were also smaller amounts of force in the 10 internal alignments for both designs. No force was generated for any of the other alignments. 27

41 Figure 14: Effect of femoral component rotational alignment on MCL force for the CR and PS implant designs. The MCL force for the 15 alignment in both knees is above the published yield point value (Kennedy et al., 1976). Lateral Collateral Ligament (LCL) Force The LCL force for the CR and PS designs was similarly affected by femoral rotational alignment. The maximum force corresponded to the 15º external alignment for both designs, with values of approximately 75 N for the CR knee and approximately 100 N for the PS knee (Figure 15). The maximum force values occurred at 20 flexion and decreased to zero by 80 flexion. A small amount of force was also generated in the 10 external alignments for both designs. No force was generated for any of the other alignments. 28

42 Figure 15: Effect of femoral component rotational alignment on LCL force for the CR and PS implant designs. Posterior Cruciate Ligament (PCL) Force For the CR knee, no force was generated in the PCL until high flexion for all component alignments, with a maximum force that was approximately 300 N higher in the 15º internal alignment than in the 15º external alignment (Figure 16). 29

43 Figure 16: Effect of femoral component rotational alignment on PCL force for the CR implant design. Patellar Ligament Force For both the CR and PS knees, the patellar ligament (PL) force was similar for all component alignments throughout flexion (Figure 17). However, the maximum PL force in the CR knee was approximately 400 N higher than in the PS knee. Figure 17: Effect of femoral component rotational alignment on patellar ligament force for the CR implant design. 30

44 Quadriceps Muscle Force The lumped quadriceps muscle force was similar for both the CR and PS designs and was the same for all component alignments until approximately 80 of flexion (Figure 18). At 80 of flexion, the internally aligned components began generating higher forces than the externally aligned components, with a maximum force that was approximately 500 N higher in the 15º internal alignment than in the 15º external alignment. Figure 18: Effect of femoral component rotational alignment on quadriceps muscle force for the CR implant design. Tibiofemoral Contact Force The tibiofemoral (TF) contact forces were lowest for the 0 alignment in the CR knee and for the 5 external alignment in the PS knee and were highest for the 15 31

45 internal alignment for both knees, with maximum values of 4119 N for the CR knee (Figure 19) and 4964 N for the PS knee occurring at 120 of flexion. In general, the TF contact forces were highest for the PS knee. Figure 19: Effect of femoral component rotational alignment on TF contact force at various knee flexion angles for the CR implant design. Patellofemoral Contact Force For both the CR and PS knees, the patellofemoral (PF) contact forces were lowest for the 15º external alignment and highest for the 15º internal alignment, with maximum values of 7849 N for the CR knee (Figure 20) and 7550 N for the PS knee occurring at 120 of flexion. In general, the PF contact forces were highest for the CR knee. 32

46 Figure 20: Effect of femoral component rotational alignment on PF contact force at various knee flexion angles for the CR implant design. Cam-Post Contact Force Cam-post forces in the PS knee were not generated until late flexion and increased with increasing flexion, with the maximum cam-post force of 1237 N occurring at 120 of flexion for the 0 alignment (Figure 21). 33

47 Figure 21: Effect of femoral component rotational alignment on cam-post contact force at various knee flexion angles for the PS implant design Variability in Tibial Component Alignment Internal/External Knee Rotation Angle The internal/external knee rotation angle for the CR and PS designs was similarly affected by tibial rotational alignment. Externally aligned components induced internal knee rotation angles while internally aligned components induced external knee rotation angles (Figure 22), which contrasts what was observed when the tibial component was fixed at 0 of rotation and the alignment of the femoral component was varied (femoral alignment analysis). 34

48 Figure 22: Effect of tibial component rotational alignment on internal/external tibiofemoral rotation angle for the CR implant design. For the CR knee, all alignments exhibited the screw-home motion. For the PS knee, the internal alignments exhibited the screw-home motion, but the external alignments maintained a relatively constant internal/external knee rotation angle throughout flexion. There was no reversal of screw-home motion, which contrasts what was observed in the femoral alignment analysis. Varus/Valgus Knee Rotation Angle For both the CR and PS knees, the varus/valgus angle was within the range of 1 valgus to 1 varus throughout knee flexion for all alignments. This was a much smaller range than that for the femoral alignment analysis, suggesting that variations in the tibial component had a much smaller effect on varus/valgus angle than variations in the 35

49 femoral component. In general, the internal alignments induced a small varus alignment throughout flexion, except the 5 internal alignment, which had a small valgus value until approximately 30 knee flexion. The external alignments induced a small valgus alignment throughout flexion, with the exception of the 5 and 10 external alignments, which induced a very small varus alignment during flexion and flexion, respectively, for the CR knee (Figure 23) and during flexion and flexion, respectively, for the PS knee (Figure 24). This was the opposite of the femoral alignment analysis, where internal alignment induced a valgus orientation and external alignment induced a varus orientation. Figure 23: Effect of tibial component rotational alignment on varus/valgus angle for the CR implant design. 36

50 Figure 24: Effect of tibial component rotational alignment on varus/valgus angle for the PS implant design. Patellar Tilt The patellar tilt for the CR and PS designs was similarly affected by tibial rotational alignment. For all component alignments, the patella had a medial tilt in early flexion and tilted laterally as knee flexion increased, similar to the femoral alignment analysis (Figure 25). 37

51 Figure 25: Effect of tibial component rotational alignment on patellar tilt for the CR implant design. The patella transitioned from a medial tilt to a lateral tilt around mid-flexion, depending on the alignment of the tibial component. This was different from the femoral alignment analysis in which some alignments maintained a medial tilt throughout all of flexion and some maintained a lateral tilt throughout all of flexion. The total change in patellar tilt, as well as the differences in patellar tilt between the alignments, were much smaller than those for the femoral alignment analysis, suggesting that variations in femoral component alignment had a greater effect than variations in tibial component alignment. Between approximately 20 and 40 knee flexion, the internal alignments, especially the 20 internal alignment, were causing an abnormally lateral tilt compared with the trend in the external alignments. 38

52 Patellar Medial/Lateral Translation The patellar medial/lateral translation for the CR and PS designs was similarly affected by tibial rotational alignment, except the values for the PS knee were all shifted laterally by approximately 1 mm compared with the CR knee. The internally aligned components had a more lateral patellar location than the externally aligned components. This was the opposite of the femoral alignment analysis up to approximately 90 flexion, after which the trend was the same. For all alignments, the patella tracked medially with increasing knee flexion (Figure 26). Figure 26: Effect of tibial component rotational alignment on patellar medial/lateral (M/L) location for the CR implant design. 39

53 The total change in patellar medial/lateral location, as well as the differences in location between the alignments, was smaller than those for the femoral alignment analysis, except in the range of approximately 85 to 100 knee flexion (Figure 12), suggesting that variations in femoral component alignment had a greater effect than variations in tibial component alignment. Patellar Superior/Inferior Translation The patellar superior/inferior translation was very similar to that for the femoral alignment analysis, suggesting that patellar superior/inferior translation is not affected by altering the alignment of either component. There were very few differences between the CR and PS designs and the values were similar for all alignments. MCL Force The results for MCL force suggest that tibial component alignment has very little effect on force compared with the effect of the femoral component alignment. For the CR and PS knees, small forces of 10 N and 16 N, respectively, were generated in the MCL for the 20 internal alignment in early flexion. For the PS knee, a small force of 1 N was also generated for the 15 internal alignment. No force was generated for any other alignment. This was different from the femoral alignment analysis in which the 10 and 15 internal femoral component alignments generated forces of up to 100 N and 475 N, respectively, for the CR knee and forces of up to 225 N and 700 N, respectively, for the PS knee. LCL Force The results for LCL force suggest that tibial component alignment has very little effect on force compared with the effect of the femoral component alignment. For the 40

54 CR and PS knees, a small force of 0.25 N and 9 N, respectively, was generated in the LCL for the 20 internal alignment in early flexion. For the PS knee, a small force of 1 N was also generated for the 15 internal alignment. No force was generated for any other alignment. This was different from the femoral alignment analysis in which the 15 external femoral component alignment generated forces of up to 75 N and 100 N for the CR and PS knees, respectively. PCL Force For the CR knee, no force was generated in the PCL until high flexion for all of the component alignments (Figure 27), which was similar to the results from the femoral alignment analysis. The 20 internal alignment generated no force. The other internal alignments followed a trend of generating less force for higher values of internal alignment, which was the opposite of the femoral alignment analysis. The maximum PCL force was approximately 160 N at the 5 external alignment compared with a maximum force of 350 N at the 15 internal alignment for the femoral alignment analysis. Also unlike the femoral alignment analysis, the external alignments showed no discernible relationship between alignment and force. 41

55 Figure 27: Effect of tibial component rotational alignment on PCL force for the CR implant design. Patellar Ligament Force The patellar ligament force for the CR and PS designs was similarly affected by tibial rotational alignment and was approximately the same for all alignments throughout flexion. However, the maximum patellar ligament force was approximately 3300 N for the CR knee compared with approximately 2800 N for the PS knee. These results were very similar to the femoral alignment analysis, suggesting that variations in femoral and tibial component alignments had little effect on patellar ligament force. Quadriceps Muscle Force The lumped quadriceps muscle force for the CR and PS designs was similarly affected by tibial rotational alignment and was the same for all alignments throughout flexion (Figure 28). This was different from the femoral alignment analysis in which 42

56 internal alignments produced slightly higher forces than external alignments at flexion angles higher than 80. This suggests that quadriceps force was more sensitive to variations in femoral component alignment in high flexion than to variations in tibial component alignment. Figure 28: Effect of tibial component rotational alignment on quadriceps muscle force for the CR implant design. Tibiofemoral Contact Force In general, the TF contact forces were higher for the PS knee than for the CR knee in high flexion, but were similar in early flexion. The maximum TF contact force for the CR knee was 3897 N, which occurred for the 5 internal alignment at 120 flexion (Figure 29). The maximum TF contact force for the PS knee was 4606 N, which occurred for the 20 internal alignment at 120 flexion (Figure 30). Unlike the femoral alignment 43

57 analysis, there was not a clear relationship between alignment and force. Additionally, the maximum TF force values for both knees are lower than the maximum force values for the femoral alignment analysis. Figure 29: Effect of tibial component rotational alignment on TF contact force for the CR implant design. 44

58 Figure 30: Effect of tibial component rotational alignment on TF contact force for the PS implant design. Patellofemoral Contact Force In general, the PF contact forces were higher for the CR knee than for the PS knee in high flexion (90 and 120 ), but were similar in early flexion. The maximum PF contact force for the CR knee was 7656 N, which occurred for the 5 external alignment at 120 flexion (Figure 31). The maximum PF contact force for the PS knee was 7155 N, which occurred for the 20 internal alignment at 120 flexion (Figure 32). Unlike the femoral alignment analysis, there was not a clear relationship between alignment and force. Additionally, the maximum PF force values for both knees are lower than the maximum force values for the femoral alignment analysis. 45

59 Figure 31: Effect of tibial component rotational alignment on PF contact force for the CR implant design. Figure 32: Effect of tibial component rotational alignment on PF contact force for the PS implant design. 46

60 Cam-Post Contact Force There did not appear to be any relationship between alignment and CP force, especially at higher degrees of flexion where the force values were highly variable. The maximum CP force of 1237 N occurred for the 0 (neutral) alignment at 120 flexion (Figure 33), similar to the femoral alignment analysis. Unlike the femoral alignment analysis, there was CP contact in early flexion for some of the alignments. Figure 33: Effect of tibial component rotational alignment on CP contact force for the PS implant design. 47

61 3.3.3 Variability in Both Femoral and Tibial Component Alignments Screw-home Motion The magnitude of the screw-home motion for both the CR and PS designs was largely driven by the femoral component alignment, since the iso-lines on the contour plot were relatively vertical (Figure 34 and Figure 35). The femoral component alignment is displayed on the horizontal axis of the plot and the tibial component alignment is displayed on the vertical axis, with the color of the iso-lines corresponding to the value of the variable of interest, in this case screw-home motion. The vertical isolines indicate that for any given femoral component alignment, the screw-home motion is not affected by variations in the tibial component alignment. The iso-lines were slightly less vertical for the PS knee than they were for the CR knee, suggesting that the tibial component alignment had a greater effect on the screwhome motion of the PS knee than it did for the CR knee. In general, however, externally aligned femoral components exhibited the natural screw-home motion while internally aligned femoral components reversed the screw-home motion in both designs. 48

62 Figure 34: Effect of femoral and tibial component rotational alignment on screw-home motion for the CR implant design. Figure 35: Effect of femoral and tibial component rotational alignment on screw-home motion for the PS implant design. 49

63 Internal/External Knee Rotation Angle The internal/external (I/E) rotation angle for the CR and PS designs was similarly affected by variations in tibial and femoral rotational alignment. In early flexion, both the tibial and femoral component alignment had an effect on the I/E angle (Figure 36). Internal femoral component alignment and external tibial component alignment induced an internal I/E angle while external femoral component alignment and internal tibial component alignment induced an external I/E angle. Figure 36: Effect of femoral and tibial component rotational alignment on I/E rotation angle at 20 knee flexion for the CR implant design. 50

64 At 70 flexion, internal tibial component alignments induced an external I/E angle while external tibial component alignments induced an internal I/E angle, with the femoral component alignment not playing a significant role (Figure 37). Figure 37: Effect of femoral and tibial component rotational alignment on I/E rotation angle at 70 knee flexion for the CR implant design. At 120 flexion, internal femoral and tibial component alignment induced an external I/E angle while external femoral and tibial component alignment induced an internal I/E angle (Figure 38). The results for the PS knee (Figure 39) were slightly different from those for the CR knee at this flexion angle, possibly explaining why the screw-home motion for the PS knee was slightly different from that for the CR knee. 51

65 Figure 38: Effect of femoral and tibial component rotational alignment on I/E rotation angle at 120 knee flexion for the CR implant design. Figure 39: Effect of femoral and tibial component rotational alignment on I/E rotation angle at 120 knee flexion for the PS implant design. 52

66 Varus/Valgus Knee Rotation Angle The varus/valgus angle for the CR and PS designs was similarly affected by variations in tibial and femoral rotational alignment. In early flexion, variations in tibial component alignment had little effect (Figure 40). For both designs, externally aligned femoral components induced a varus orientation while internally aligned femoral components induced a valgus orientation. Figure 40: Effect of femoral and tibial component rotational alignment on varus/valgus angle at 20 knee flexion for the CR implant design. At 120 flexion, the tibial component alignment appeared to have an effect, but the varus/valgus values dropped to less than 1 (Figure 41). Additionally, tibial 53

67 component alignment appeared to have a slightly larger effect on the varus/valgus angle for the PS knee than for the CR knee. Figure 41: Effect of femoral and tibial component rotational alignment on varus/valgus angle at 120 knee flexion for the CR implant design. Patellar Tilt For both the CR and PS designs, changes in tibial component alignment had little to no effect on patellar tilt. Internally aligned femoral components resulted in a medial patellar tilt while externally aligned femoral components resulted in lateral patellar tilt (Figure 42). In general, the patella tilted laterally with increasing knee flexion. 54

68 Figure 42: Effect of femoral and tibial component rotational alignment on patellar tilt at 70 knee flexion for the CR implant design. Patellar Medial/Lateral Translation The patellar M/L location for the CR and PS designs was similarly affected by variations in tibial and femoral rotational alignment. At 20 flexion, variations in tibial component alignment had little effect. Internally aligned femoral components induced a medial patellar location while externally aligned femoral components induced a lateral patellar location (Figure 43). 55

69 Figure 43: Effect of femoral and tibial component rotational alignment on patellar M/L location at 20 knee flexion for the CR implant design. At 70 knee flexion, both the femoral and tibial component alignment had an effect on patellar M/L location. Internal femoral component alignment and external tibial component alignment induced a more medial patellar location while external femoral component alignment and internal tibial component alignment induced a more lateral patellar location. At 120 knee flexion, the tibial component alignment still had an effect, but external femoral and tibial component alignment induced a more medial patellar location while internal femoral and tibial component alignment induced a more lateral patellar location (Figure 44). 56

70 Figure 44: Effect of femoral and tibial component rotational alignment on patellar M/L location at 120 knee flexion for the CR implant design. Patellar Superior/Inferior Translation The patellar S/I location for the CR and PS designs was similarly affected by variations in tibial and femoral rotational alignment. The values between the alignments never differed by more than 2 mm, and, in many cases, the values only changed by a fraction of 1 mm, suggesting that changing the alignment of the components did not have much effect. In general, the patella moved inferiorly with increasing knee flexion. At 20 knee flexion, the patella was located more superior at external femoral component alignments and located more inferior at internal femoral and external tibial component 57

71 alignments. At 70 and 120 knee flexion, the patella was more inferior at the internal femoral and tibial component alignments (Figure 45). Figure 45: Effect of femoral and tibial component rotational alignment on patellar S/I location with respect to the femoral coordinate system at 70 knee flexion for the CR implant design. MCL Force For both the CR and PS knees, the MCL force was zero for all external femoral component alignments and for small amounts of internal femoral component alignment, regardless of the tibial component alignment. Force was only generated for internal alignment of the femoral component higher than 10. For the CR knee, the tibial component alignment only had an effect when the femoral component was at 15 internal 58

72 alignment. The PS knee exhibited a similar effect except at higher flexion angles (especially 120 ) where the tibial component alignment had little, if any, effect. The maximum force occurred at 15 internal femoral component alignment, with values of approximately 745 N at 90 flexion for the CR knee (Figure 46) and approximately 775 N at 120 flexion for the PS knee (Figure 47). Figure 46: Effect of femoral and tibial component rotational alignment on MCL force at 90 knee flexion for the CR implant design. 59

73 Figure 47: Effect of femoral and tibial component rotational alignment on MCL force at 120 knee flexion for the PS implant design. LCL Force For both the CR and PS knees, the LCL force was zero for all internal femoral component alignments and for small amounts of external femoral component alignment, regardless of tibial component alignment. Force was only generated for external femoral component alignments of 10 or higher. High force occurred at external femoral and internal tibial component alignments, with maximum values of approximately 100 N for the CR knee (Figure 48) and approximately 220 N for the PS knee (Figure 49) at 20 knee flexion. 60

74 Figure 48: Effect of femoral and tibial component rotational alignment on LCL force at 20 knee flexion for the CR implant design. Figure 49: Effect of femoral and tibial component rotational alignment on LCL force at 20 knee flexion for the PS implant design. 61

75 PCL Force The PCL force was zero for all component alignments at low flexion angle. At higher flexion angles, the force increased with increasing amounts of internal femoral component alignment. The maximum force of approximately 375 N corresponded to the 15 internal femoral component alignment and 5 external tibial component alignment at 120 knee flexion (Figure 50). Figure 50: Effect of femoral and tibial component rotational alignment on PCL force at 120 knee flexion for the CR implant design. 62

76 Patellar Ligament Force The patellar ligament force was affected differently for the CR and PS knees by variations in tibial and femoral rotational alignment. For the CR knee at 30 knee flexion, the highest force corresponded to the neutrally aligned tibial component (Figure 51). All other alignments had relatively low force. At 60 flexion this trend was reversed, with low forces corresponding to the neutrally aligned tibial component (Figure 52). At 90 flexion there was no discernible trend in force and the maximum force occurred at the 10 internal femoral and 10 external tibial component alignment (Figure 53). At 120 flexion, the maximum force of approximately 3370 N corresponded to the 15 external femoral and 20 external tibial component alignment (Figure 54). Figure 51: Effect of femoral and tibial component rotational alignment on patellar ligament force at 30 knee flexion for the CR implant design. 63

77 Figure 52: Effect of femoral and tibial component rotational alignment on patellar ligament force at 60 knee flexion for the CR implant design. Figure 53: Effect of femoral and tibial component rotational alignment on patellar ligament force at 90 knee flexion for the CR implant design. 64

78 Figure 54: Effect of femoral and tibial component rotational alignment on patellar ligament force at 120 knee flexion for the CR implant design. For the PS knee at 30 knee flexion, the highest patellar ligament forces corresponded to external femoral component alignment and internal tibial component alignment (Figure 55). At 60 flexion, relatively high forces were seen for most alignments while the lowest forces corresponded to high internal femoral and tibial component alignments (Figure 56). At 90 flexion, high force generally corresponded to high internal and external tibial component alignments while low force corresponded to high internal femoral component alignment (Figure 57). At 120 flexion, the lowest forces corresponded to internal femoral and tibial component alignments (Figure 58). The maximum force of approximately 2980 N corresponded to the 0 femoral and 15 65

79 external tibial component alignment and was lower than the maximum force for the CR knee. Figure 55: Effect of femoral and tibial component rotational alignment on patellar ligament force at 30 knee flexion for the PS implant design. 66

80 Figure 56: Effect of femoral and tibial component rotational alignment on patellar ligament force at 60 knee flexion for the PS implant design. Figure 57: Effect of femoral and tibial component rotational alignment on patellar ligament force at 90 knee flexion for the PS implant design. 67

81 Figure 58: Effect of femoral and tibial component rotational alignment on patellar ligament force at 120 knee flexion for the PS implant design. Quadriceps Muscle Force The quadriceps muscle force was affected differently for the CR and PS knees by variations in tibial and femoral rotational alignment. For the CR knee at 30 knee flexion, the highest force corresponded to the 0 tibial component alignment (Figure 59). At 60 flexion, the force was lowest for the 0 tibial component alignment (for all femoral component alignments) and highest for the 15 internal femoral and 20 external tibial component alignment (Figure 60). At 90 flexion, the maximum force corresponded to the 15 internal femoral and 20 external tibial component alignment (Figure 61). At 120 flexion the maximum quadriceps muscle force of approximately 5500 N corresponded to the 15 internal femoral and 20 internal tibial component alignment (Figure 62). 68

82 Figure 59: Effect of femoral and tibial component rotational alignment on quadriceps force at 30 knee flexion for the CR implant design. Figure 60: Effect of femoral and tibial component rotational alignment on quadriceps force at 60 knee flexion for the CR implant design. 69

83 Figure 61: Effect of femoral and tibial component rotational alignment on quadriceps force at 90 knee flexion for the CR implant design. Figure 62: Effect of femoral and tibial component rotational alignment on quadriceps force at 120 knee flexion for the CR implant design. 70

84 For the PS knee at 30 knee flexion, the highest forces corresponded to high internal femoral and high external tibial component alignment as well as to high external femoral and high internal tibial component alignment (Figure 63). At 60 flexion, the lowest forces corresponded to external femoral and internal tibial component alignments and the highest forces corresponded to the 15 internal femoral and 20 external tibial component alignment (Figure 64). At 90 flexion, the highest forces corresponded to internal femoral and external tibial component alignments (Figure 65). Unlike the CR knee, the maximum quadriceps force at 90 flexion corresponded to the 10 internal femoral and 15 external tibial component alignment. At 120 flexion, similar to the CR knee, the maximum quadriceps muscle force of approximately 5425 N corresponded to the 15 internal femoral component alignment (Figure 66). Figure 63: Effect of femoral and tibial component rotational alignment on quadriceps force at 30 knee flexion for the PS implant design. 71

85 Figure 64: Effect of femoral and tibial component rotational alignment on quadriceps force at 60 knee flexion for the PS implant design. Figure 65: Effect of femoral and tibial component rotational alignment on quadriceps force at 90 knee flexion for the PS implant design. 72

86 Figure 66: Effect of femoral and tibial component rotational alignment on quadriceps force at 120 knee flexion for the PS implant design. Tibiofemoral Contact Force The TF contact force was affected differently for the CR and PS knees by variations in tibial and femoral rotational alignment. For the CR knee at 30 flexion, the TF contact forces were relatively low for all alignments except the 5 internal femoral and 20 external tibial component alignment (Figure 67). At higher flexion angles, the maximum force corresponded to the 15 internal femoral and 20 internal tibial component alignment. The maximum force of approximately 4425 N occurred at 120 flexion (Figure 68). 73

87 Figure 67: Effect of femoral and tibial component rotational alignment on TF contact force at 30 knee flexion for the CR implant design. Figure 68: Effect of femoral and tibial component rotational alignment on TF contact force at 120 knee flexion for the CR implant design. 74

88 For the PS knee at 30 flexion, contact forces were relatively low, except for the 15 external femoral and 15 internal tibial component alignment, where the maximum force occurred (Figure 69). At 60 and 90, similar to the CR knee, the maximum force corresponded to the 15 internal femoral and 15 internal tibial component alignment. At 120, the force values were higher than for the CR knee, with a maximum force of approximately 5250 N (Figure 70). Figure 69: Effect of femoral and tibial component rotational alignment on TF contact force at 30 knee flexion for the PS implant design. 75

89 Figure 70: Effect of femoral and tibial component rotational alignment on TF contact force at 120 knee flexion for the PS implant design. Patellofemoral Contact Force The PF contact force was affected differently for the CR and PS knees by variations in tibial and femoral rotational alignment. For the CR knee at 30 flexion, the lowest forces corresponded to externally aligned femoral and tibial components. The highest forces corresponded to internal femoral component alignment, with the maximum force occurring at 15 internal femoral and 0 tibial component alignment (Figure 71). At 60 flexion, the low forces corresponded to externally aligned femoral components and high forces corresponded to internally aligned femoral components. At 90 flexion, there was no discernible trend or pattern in force values. At 120 flexion, low forces corresponded to externally aligned femoral components and high forces corresponded to 76

90 internally aligned femoral components (Figure 72). The maximum force of approximately 8040 N corresponded to the 15 internal femoral and 20 internal tibial component alignment at 120 flexion. Figure 71: Effect of femoral and tibial component rotational alignment on PF contact force at 30 knee flexion for the CR implant design. 77

91 Figure 72: Effect of femoral and tibial component rotational alignment on PF contact force at 120 knee flexion for the CR implant design. For the PS knee at 30 flexion, the lowest forces corresponded to externally aligned femoral and tibial components (similar to the CR knee). The highest forces corresponded to internally aligned femoral components (Figure 73). At higher flexion angles, low forces corresponded to externally aligned femoral components and high forces corresponded to internally aligned femoral components (Figure 74). The maximum force for the PS knee was approximately 7560 N, compared with approximately 8040 N for the CR knee. 78

92 Figure 73: Effect of femoral and tibial component rotational alignment on PF contact force at 30 knee flexion for the PS implant design. Figure 74: Effect of femoral and tibial component rotational alignment on PF contact force at 120 knee flexion for the PS implant design. 79

93 Cam-Post Contact Force At flexion angles up to 60, the CP force was zero for all alignments except high external and high internal tibial component alignments (Figure 75). At 90 flexion, the CP force was relatively low for all alignments except internal femoral component alignments greater than 10, with the maximum force corresponding to high internal femoral and tibial component alignment (Figure 76). At 120 flexion, there was a less discernible trend in forces (Figure 77). In general, however, the highest forces corresponded to a 15 internal femoral component alignment. Figure 75: Effect of femoral and tibial component rotational alignment on CP contact force at 60 knee flexion for the PS implant design. 80

94 Figure 76: Effect of femoral and tibial component rotational alignment on CP contact force at 90 knee flexion for the PS implant design. Figure 77: Effect of femoral and tibial component rotational alignment on CP contact force at 120 knee flexion for the PS implant design. 81

95 3.4 Discussion The effects of variability in the prosthetic component alignments in TKA have not been studied in terms of their biomechanical impact on functional tasks. The purpose of this study was to determine the effects of variability in femoral and tibial component rotational alignment in TKA on knee kinematics, quadriceps muscle force, ligament forces, and contact forces using a forward-dynamic computer model of an Oxford Rig. We found that femoral component alignment, in general, had a much greater effect on our variables of interest than the tibial component alignment or choice of implant design. Internal rotation of the femoral component led to a reversal of the natural screw-home motion of the knee as well as high quadriceps muscle force, MCL force, and contact forces. Previous research on femoral component rotation has investigated the variability associated with various alignment techniques (Jenny and Boeri, 2004; Kinzel et al., 2005; Robinson et al., 2006; Siston et al., 2008; Siston et al., 2005), as well as the effect of femoral component rotational alignment on knee kinematics (Anouchi et al., 1993; Miller et al., 2001; Rhoads et al., 1990). Our results for tibiofemoral and patellofemoral kinematics generally agreed with previous researchers. Miller et al. (2001) found that 5 external rotation, 0 rotation, and 5 internal rotation of the femoral component all caused the tibia to internally rotate with knee flexion up to 90, after which the neutrally aligned component caused the tibia to cease to rotate and the internally aligned component caused the tibia to externally rotate. These results are consistent with our findings, except that our internally aligned component caused the tibia to externally rotate throughout flexion (a reversal of the screw-home motion). Previous studies found that externally 82

96 rotated femoral components induce dynamic varus in the knee (greater varus orientation with knee flexion) and internally rotated femoral components induce dynamic valgus in the knee (Anouchi et al., 1993; Miller et al., 2001), which is consistent with our results. Our results also agreed with previous researchers who found that internal rotation of the femoral component tilted the patella more medially compared with the neutral alignment while external rotation of the femoral component tilted the patella more laterally (Miller et al., 2001; Rhoads et al., 1990). Our results for patellar medial/lateral translation showed that, up to approximately 90 knee flexion, internal rotation of the femoral component displaced the patella medially compared with the neutral alignment while external rotation of the femoral component displaced the patella laterally, which is consistent with the findings of previous studies (Anouchi et al., 1993; Rhoads et al., 1990). The quadriceps force required to perform the Oxford Rig motion was higher for the internally aligned femoral components than for the externally aligned femoral components for knee flexion angles higher than 80. These higher forces may have functional consequences for activities requiring deep knee flexion such as kneeling and squatting, since many TKA patients experience quadriceps weakness after surgery (Huang et al., 1996; Mizner et al., 2005; Stevens et al., 2003). Previous studies show that quadriceps weakness is also present prior to surgery (Mizner et al., 2005; Stevens et al., 2003), and increases by as much as 60% after surgery (Stevens et al., 2003). However, the effect of component alignment on post-operative quadriceps weakness remains unknown. 83

97 Additionally, since quadriceps force was the same for all component alignments below 80 flexion, our results suggest that variation in component rotational alignment may not have an effect on functional activities that do not require a large amount of knee flexion, such as stair-climbing. The fact that rotational alignment of the femoral component would appear to have no effect on stair-climbing does not mean that alignment variations in the other planes would not have an effect. This finding may even be desirable, since accurate rotational alignment of the femoral component is highly variable (Jenny and Boeri, 2004; Kinzel et al., 2005; Robinson et al., 2006; Siston et al., 2008; Siston et al., 2005). Future work is necessary to determine the effect of variations in component alignment in the frontal and sagittal planes on functional activities. We found that large amounts of internal femoral component alignment may be detrimental to the MCL. For the 15 internal femoral component alignment in both implant designs, the MCL force is above the published yield point value of 453 N (Kennedy et al., 1976). However, even with high values of external rotation, the LCL force is below the published yield point value of 309 N (Sugita and Amis, 2001). Failure of the MCL is not a common occurrence, so these results may instead suggest that high forces in the MCL prevent TKA patients from doing certain activities. If patients experience high tension in the MCL during squatting or kneeling, we can assume that they would cease to perform that activity and would avoid any functional activities which require knee flexion angles that may put their MCL in danger of failing. Internal femoral component alignment of 15 corresponded to the highest tibiofemoral and patellofemoral contact forces for both implant designs. Singerman et al. (1997) investigated the effect of femoral component rotational alignment on 84

98 patellofemoral contact forces for a CR and PS implant design. They found that there were no significant changes in contact force for the PS design after 10 internal or external rotation, but there was a significant increase in patellofemoral contact force for the CR design after 10 internal rotation compared with 10 external rotation. Our results, in contrast, showed greater changes in patellofemoral contact force after malrotation of the femoral component for the PS design compared with the CR design. The differences between their results and ours may be due to differences in the design of the patellar prosthetic component. The CR implant design in Singerman s study used a modified dome shaped patella, characterized by a central projection and peripheral flats (Singerman et al., 1997). We did, however, find an increase in patellofemoral contact force for the CR design after 10 internal rotation compared with 10 external rotation, consistent with the findings of Singerman et al. (1997). Previous research on tibial component rotation has focused on repeatable methods of alignment using various landmarks (Akagi et al., 2005; Eckhoff et al., 1995; Ikeuchi et al., 2007). To our knowledge, no study has investigated the biomechanical effects of systematic variation in the rotational alignment of the tibial component. However, reports of patellofemoral complications such as abnormal patellar tracking and tilt, subluxation, dislocation, and excessive wear have been linked to internal rotation of the tibial component (Barrack et al., 2001; Dalury, 2001; Ranawat, 1986). We found that patellar tilt and tracking were affected by variations in the tibial component alignment, with a greater medial patellar tilt and lateral patellar location for the internally aligned tibial components compared with the externally aligned components. 85

99 Our results for variations in tibial component rotational alignment showed that internally aligned components induced an external tibiofemoral rotation angle while externally aligned components induced an internal tibiofemoral rotation angle. This was the opposite of our results for the femoral component analysis. We believe this may suggest that the femoral component alignment drives the internal/external knee rotation angle and the tibial component, as well as the tibia itself, will rotate accordingly to follow the femoral component. This phenomenon would explain why the maximum internal/external rotation angle values occur for large alignment differences between the femoral and tibial component. For example, when the femoral component is internally rotated and the tibial component is externally rotated, the tibia will rotate internally by a large amount so that the tibial component is seated with the femoral component. Unlike the femoral alignment analysis, variations in tibial component alignment had little effect on the MCL force, LCL force, quadriceps muscle force, or contact forces. For both implant designs, MCL and LCL forces were only generated for the 20 internal tibial component alignment and these forces were very small compared with the results from the femoral alignment analysis. The quadriceps muscle force was approximately the same for all tibial component alignments throughout flexion and there was no discernible trend in contact forces. To our knowledge, no study has investigated the effects of systematic variation in the rotational alignment of both the femoral and tibial component. However, reports of patellofemoral complications (Berger et al., 1998) as well as anterior knee pain (Barrack et al., 2001) have been linked with combined internal rotation of the femoral and tibial components. Berger et al. (1998) found a direct correlation between the amount of 86

100 combined internal component rotation and the severity of patellofemoral complications. Our results for variations in both femoral and tibial component alignment showed that, in general, the femoral component alignment had a much greater effect on our variables of interest than the tibial component alignment. The tibial component alignment only had an effect on the MCL force, quadriceps force, and contact forces when the femoral component was internally aligned to 15. At this femoral component alignment, internally aligned tibial components generated higher forces than externally aligned tibial components. There are some important limitations of the Oxford Rig computer model. The Oxford Rig is an idealized motion, simulating a perfect up-down movement with the hip directly above the ankle. Additionally, the pelvis is only permitted to translate vertically, with all of its other degrees of freedom fixed, while the foot is permitted to plantarflex a large amount during knee flexion. The 30 kg mass that is used in the model to simulate body weight is positioned directly above the hip and the only muscle in the model that carries any force is the lumped quadriceps muscle. Obviously, a true squat is a much more complex task than what the Oxford Rig is simulating. However, we were still able to use this model to find trends in our variables of interest and investigate the effects of altering component alignment. In the future, we can build upon the Oxford Rig model to create more realistic simulations. Femoral component rotational alignment had a much greater effect on our variables of interest than tibial component rotational alignment, suggesting that tibial component rotational alignment may not matter in terms of post-operative performance. In light of these results, the debate as to which alignment technique should be used to 87

101 establish tibial rotational alignment, as well as the difficulty in accurately aligning this component, lose significance. Our findings instead emphasize the need to accurately establish the rotational alignment of the femoral component, and failure to do so may impact functional outcome. 88

102 Chapter 4: Conclusion Although total knee arthroplasty is an effective intervention for various disabling disorders of the knee, suboptimal outcomes do occur. Patellofemoral complications are among the most common post-operative problems and are a major cause for revision surgery (Clayton and Thirupathi, 1982; Dennis et al., 1992; Huo and Sculco, 1990). Poor rotational alignment of the prosthetic components in particular is a major cause of patellofemoral complications (Berger et al., 1998; Clayton and Thirupathi, 1982; Dennis et al., 1992; Huo and Sculco, 1990; Ranawat, 1986). Additionally, significant variability can be associated with femoral and tibial component rotational alignment (Akagi et al., 2005; Eckhoff et al., 1995; Ikeuchi et al., 2007; Jenny and Boeri, 2004; Kinzel et al., 2005; Siston et al., 2008; Siston et al., 2006; Siston et al., 2005), but how this variability translates into functional outcome remains unknown. We determined the effects of variability in femoral and tibial component rotational alignment in TKA on tibiofemoral and patellofemoral kinematics, quadriceps muscle force, ligament forces, and contact forces for two implant designs using a forward-dynamic computer model of an Oxford Rig. The kinematics, MCL force, LCL force, PCL force, quadriceps muscle force, and contact forces were all largely driven by femoral component alignment, with less influence from the tibial component alignment or choice of implant. In general, the tibial component alignment only appeared to have a 89

103 significant effect on these variables when the femoral component was internally aligned to 15, with the exception of the LCL force for which the femoral component was externally aligned to 15 before the tibial component alignment had an effect. Internal femoral component alignment led to a reversal of the screw-home motion, high quadriceps muscle forces, possibly detrimental forces in the MCL, and high contact forces, highlighting the need to accurately align this component and avoid internal rotation during surgery. 4.1 Contributions No study has extensively investigated the biomechanical effects of variations in both femoral and tibial component alignment in TKA. Additionally, the significant variability associated with component alignment has not been translated into functional outcome. Since TKA s are not failing in large numbers, there must be a window of variability that is acceptable and a window of alignment variability that leads to suboptimal outcomes, but these ranges have not been established. This thesis determined the effects of variability in rotational alignment during an active knee flexion activity, similar to what is seen in functional tasks such as stair climbing, as a first step towards understanding the impact of this variability on functional outcome. We also investigated a larger range of femoral rotational alignments than what has been done in previous experiments (Anouchi et al., 1993; Miller et al., 2001; Rhoads et al., 1990), as well as all combinations of these alignments, which was made possible through the use of a forward-dynamic computer simulation. It had been shown previously that internal component rotation is associated with more severe complications than external rotation 90

104 (Barrack et al., 2001; Berger et al., 1998; Rhoads et al., 1990), but the biomechanical reasons for this were largely unknown. Our findings not only provide a biomechanical basis for why internal rotation, especially of the femoral component, is undesirable, but they also provide specific values at which complications may occur. 4.2 Additional Applications Variability in Other Alignment Planes This thesis determined the effects of variations in internal/external (transverse plane) rotational alignment of the prosthetic components in TKA, but variability in the other alignment planes is also an important consideration. We can determine the biomechanical effects of variations in varus/valgus (frontal plane) alignment as well as variations in flexion/recurvatum (sagittal plane) alignment of the femoral and tibial components in the simulated Oxford Rig. The large number of possible alignment combinations means the simulations will be very time-consuming. It may be necessary to develop algorithms to reduce the computational time of the simulations to make this alignment study feasible. Investigation of Additional Implant Designs The work in this thesis used two implant designs: a cruciate-retaining and a posterior-substituting version of the Scorpio implant from Stryker Orthopaedics. Although we found that many of our variables of interest for both implant designs were similarly affected by variations in component alignment, greater differences may exist between other types of implants. We can use our computer model to investigate the 91

105 effects of other versions of the Scorpio implant, as well as implants from other companies, such as Zimmer. Investigation of Changes in Soft Tissue Properties In addition to prosthetic implant design and alignment, soft tissue properties are also important factors affecting successful post-operative outcome following TKA (Insall et al., 1985). Ligament stiffness and soft tissue balancing are important for proper knee joint stability. We can use our computer model to investigate the effects of changes in these properties. For example, decreasing the MCL length in the model would allow us to determine the effects of an acquired varus deformity on our variables of interest. Prediction of Implant Wear We found that the highest tibiofemoral and patellofemoral contact forces were associated with internal femoral component alignment. These high contact forces may lead to accelerated wear of the prosthetic components and the need for revision surgery. We can use our contact force results to estimate the amount of wear seen for the internal component alignments using the methods of Fregly et al. (2005). 4.3 Future Work Validation with Mechanical Oxford Rig The results from our simulations compared well with published results, but we intend to provide further validation with experiments from a mechanical Oxford Rig (Figure 78). We can adjust the rotational alignment of the prosthetic components in the rig and use a motion capture system to collect kinematic data during knee flexion (Figure 79). 92

106 Figure 78: Mechanical Oxford Rig 93

107 Figure 79: Close-up of femoral and tibal components with attached motion-capture markers on the mechanical Oxford Rig. The Oxford Rig computer model used for our simulations is modeled as a realistic knee, with muscle and ligament properties as described earlier, so it cannot be used for direct comparison with the results from the mechanical Oxford Rig. Instead, the data collected from the experiments with the mechanical rig will be compared with simulation results from a second computer model which accurately represents the properties of the mechanical rig. Agreement between the experiments and computer simulations will provide validation that our computer model results are reliable. These experiments are currently in-progress. 94