One of the main concerns about the currently available simulators is that the TKA is driven in a “passive way” for assessment. For the simulators for the wear assessment, the tibio-femoral relative motion is automatically made by using the knee kinematics and loading profile of a normal gait. As for the simulators for the kinematics and kinetics assessment of TKA, also the predicted loading profiles introduced from the theoretical model are applied as the input data to drive the simulator. It should be noted that the human joints are driven by the muscles' forces and external loads, and their kinematics and kinetics are the “outcome”. This being so, the knee simulator should be driven by the muscles' forces and upon these conditions the TKA performance is to be assessed. Some other concerns about the current simulations are as follows. The effects of hip joint motion are not taken into account. The upper body weight is applied along a vertical rod in such a way as a crank-slider. Furthermore, few simulators are capable of knee flexion greater than about 110°.

Considering the above, we have developed a novel knee simulator which makes it possible to reproduce the active and natural knee motion to assess kinematics and kinetics of TKA. In the experiment, the custom-designed PS type TKA was attached and the simulator was operated so as to reproduce the sit-to-stand features, thereby introducing the tibio-femoral loading profiles during the motion.

Figure 1 illustrates the external appearances of the simulator and a close view of the knee joint compartment. Since our simulator is composed of a multiple inverted pendulum, the knee part bears the upper body weight in a physiological way. The holder bracket is set to prevent the simulator from collapsing for security. The dimension and weight of each link were set as close as those of each segment of a normal male subject. Our simulator is driven by the wire pull mechanism which substitutes the human musculo-skeletal system of lower limb. Figure 2 shows close views of tibial tray with load cells. In Fig.2a, cell FR, FC and FL are to measure the tangential components of tibio-femoral contact force, i.e., the Anterio-Posterior force (AP force). The rest five cells are to measure the normal components of tibio-femoral contact force (normal force). As shown in Fig.2c, the tibial insert of TKA is mounted on the lid of the tibial tray box.

In the experiment, a PS type TKA whose maximum flexion angle of 150° was attached to the simulator for evaluation. The simulator was operated so as to reproduce the sit-to-stand features and the data concerning about the AP force, F_t, and the normal force, F_n were recorded.

Figure 3 shows the variations of knee flexion angles and knee contact forces respectively as a function of normalized time. Our knee simulator may have a potential for substituting the in vivo measurement.

There have been a large number of studies reporting the knee joint force during level walking, however, the data of during deep knee flexion are scarce, and especially the data about patellofemoral joint force are lacking. Deep knee flexion is a important motion in Japan and some regions of Asia and Arab, because there are the lifestyle of sitting down and lying on the floor directly. Such data is necessary for designing and evaluating the new type of knee prosthesis which can flex deeply. Therefore we estimated the patellofemoral and tibiofemoral forces in deep knee flexion by using the masculoskeltal model of the lower limb.

The model for the calculation was constructed by open chain of three bar link mechanism, and each link stood for thigh, lower leg and foot. And six muscles, gluteus maximus, hamstrings, rectus, vastus, gastrocnemius and soleus were modeled as the lines connecting the both end of insertion, which apply tensile force at the insertion on the links. And the model also included the gravity forces, thigh-calf contact forces on the Inputting the data of floor reacting forces and joint angles, the model calculated the muscle forces by the moment equilibrium conditions around each joint, and some assumptions about the ratio of the biarticular muscles. And then, the joint forces were estimated from the muscle forces, using the force equilibrium conditions on patella and tibia. The position/orientation of each segments, femur, patella and tibia, were decided by referring the literature.

The motion to be analyzed was standing up from kneeling posture. The joint angles during the motion are shown in Fig.1. This motion included the motion from kneeling to squatting, rising the knee from the floor by flexing hip joint, and the motion from squatting to standing. The test subject was a healthy male, age 23[years], height 1.7[m], weight 65[kgw].

Results were shown in Fig.2. The patellofemoral force was little at standing posture, the end of the motion, however, was as large as tibiofemoral force during the knee joint angle was over 130 degrees. The reason of this was that the patellofemoral joint force was heavily dependent on the quadriceps forces, and the quadriceps tensile force was large at deep knee flexion, at kneeling or squatting posture. The maximum tibiofemoral force was 3.5[BW] at the beginning of standing up from squatting posture. And the maximum patellofemoral force was 3.8[BW] at the motion from kneeling to squatting posture. The conclusion was that the patellofemoral joint force might not be ignored in deep knee flexion and the design of the knee prosthesis should be include the strength design of patellofemoral joint.

1. Introduction

Such a Total Knee Arthroplasty (TKA) that is capable of making high knee flexion has been long awaited for the Asian and Muslim people. Our research group has developed the TKA possible to attain complete deep knee flexion such as seiza sitting. Yet as seiza is peculiar to the Japanese, other strategies will be necessary for our TKA to be on the overseas market. Still it is impractical to prepare many kinds of modifications of our TKA to meet various demands from every country/region. To this end, we contrived a way to modularize the post-cum alignment of our TKA in order to facilitate the following three activities containing high knee flexion: praying for the Muslim, gardening or golfing for the Westerner, sedentary siting on a floor for the Asian. We performed simulation and experiment, such as a mathematical model analysis, FEM analysis and a cadaveric study, thereby determining the optimal combination of moduli for the above activities respectively.

Introduction

Mobility at insert-tray articulations in mobile bearing knee implant accommodates lower cross-shear at polyethylene (PE) insert, which in turn reduces wear and delamination as well as decreasing constraint forces at implant-bone interfaces. Though, clinical studies disclosed damage due to wear has occurred at these mobile bearing articulations. The primary goal of this study is to investigate the effect of second articulations bearing mobility and surface friction at insert-tray interfaces to stress states at tibial post during deep flexion motion.

Purpose:

Differences in the sizes of femoral and tibial components between females and males, between osteoarthritis (OA) and rheumatoid arthritis (RA), and between measured bone resection and the gap control technique during TKA were assessed.

Knowledge of joint kinematics in the lower limb is important for understanding joint injuries and diseases and evaluating treatment outcomes. However, limited information is available about the joint kinematics required for high flexion activities necessary for floor sitting life style. In this study, the hip and knee joint kinematics of ten healthy male and ten healthy female subjects were investigated using an electromagnetic motion tracking system. We measured the hip and knee joints' functions moving into 1) kneeling on knees with legs parallel without using arms, 2) kneeling on knees with legs parallel with using arms, 3) kneeling on knees with one foot forward without using arms, 4) cross-legged sitting, 5) kneeling with legs to the side, 6) sitting with legs stretched out, and 7) deep squatting, and moving out of the above seven conditions. Conditions 1) through 3) were Japanese seiza style. On conditions 4) through 7), arms were not used. We further measured the functions of putting on and taking off a sock under such conditions as 8) with standing position and 9) sitting position (Fig 1). Here special attention was paid for flexion and extension motion. The data were used to produce the pattern of joint angulation against the percentage of the cycle for each individual conducting each activity. The kinematic curves were split into 3 phases: moving into the rest position, the rest position and out of the rest position. It should be noted that the moving into and the rest phases were split at the moment when the peak value was determined during the moving into phase. Thus the initiation of the rest phase on the curve was not coinciding with the moment the subject reached at the rest position. This was necessary in order not for the mean kinematic curve to become too dull in shape. Same was true when the end of rest phase was determined. The maximum hip and knee joint angles during the cycle were determined. Further a relationship between the hip and knee joint excursions were investigated. The results indicated condition 8) requires the maximum flexion angles to the hip among all conditions, 157.5 ± 20.4° and condition 3) to the knee joint, 157.1 ± 10.0° respectively (Fig 2). The results also indicated in many activities, the maximum joint angles were recorded not during the rest phase but during the moving into or out of phase. In any conditions even including donning on and off a sock, a strong relationship was found between the hip and knee joints motion (Fig 3), indicating the bi-articular muscles' co-contraction during the sit to stand activities. The data presented in this study will increase the knowledge of high-flexion needs especially in non-Western cultures and provide an initial characterization of the prosthesis kinematics in high flexion.

There are several concerns about the current simulators for TKA. First, the knee is flexed in a “passive way” under the condition of applying constant muscular tension forces. Second, the effects of hip joint motion are not taken into account. Thirdly, the external load for example, upper body weight is not applied in a natural way. Finally, few simulators are capable of knee flexion greater than about 100°.

To this end, we have developed a novel knee simulator system that reproduces the active and natural knee motion to evaluate kinematics and joint forces of TKA. Our simulator system has the following advantages and innovative features. First, it is driven directly by muscles' tension forces, and the knee is capable of active flexion. Secondly, a hip joint is incorporated into it and the lower limb motion is achieved in a synergistic way between the hip and knee joints. Thirdly, it is capable of complete deep knee flexion up to 180°.

Figure 1 shows the structure of the system. Both the hip and knee joints are moved by the tension forces of four wires that simulate the functions of the mono-articular muscles ((1), (3)) and the bi-articular muscles ((2), (4)) by means of a multiple pulley system (Fig 2). The femoral and tibial components of TKA are secured in the distal end of the upper link (thigh) and the proximal end of the lower link (shank) respectively. The ankle assembly has three sets of rotary bearings whose axes intersect at a fixed point, the center of the ankle, allowing spherical movement of the tibia about the ankle center. Springs were stretched around the ankle center to substitute the muscles around the ankle. Weights I and II are counterweights so as to duplicate the weights of the human upper body, thigh and shank respectively. The wires are pulled to produce the hip and knee motions. The linear bearings running along vertical rods also prevent the system from collapsing.

In the experiment, a custom-designed posterior stabilized type TKA was attached to the simulator system for evaluation. The system was operated so as to reproduce the sit-to-stand features in a quasi-static manner in order to study the kinematics of TKA. Beyond 130°, the knee proceeded to flex passively because of upper body weight. Conspicuous internal/external rotation or valgus/varus motion of the tibia relative to the femur was not observed as the knee flexed. When our simulator system was driven in a quasi-static manner, it was able to measure the kinematics of TKA however, when the system was driven in a dynamic manner, it oscillated because the springs around the ankle were not stiff enough to hold the inverted pendulum-like system upright and the ratios of the tension force exerted by the four wires simulating muscles could not be determined appropriately.

The objective of this study is to investigate the effect of the tensile force ratio between the two extensor muscles for the hip joint on the forces acting on the knee joint. We have created a mathematical model of lower limb and have performed some simulations to introduce the forces acting on the knee joint for various daily activities. With only one exception, our results for knee joint forces were in good or close agreement involving all range of knee flexion either with the in vivo data or other literature data. The exception was that, at high knee flexion angle (knee bend), the tangential components of knee joint force became pretty larger than those from the in vivo data, while the normal components did not differ much with each other though as shown in Fig. 1.

We considered that the above mentioned discrepancy was attributed to the fact that in order to solve an indeterminate problem, we had assumed the hamstrings and the gluteus maximus work together with the same force with each other, thereby introducing the hamstrings force too great. Then we expected that the above discrepancy could be eliminated if we change the tensile force ratio between the hamstrings and the gluteus maximus basing upon a certain biomechanical criterion, for example the biological cross-sectional areas.

Thus we modified our model so that we could introduce the knee joint forces as a function of the tensile force ratio. Simulation was performed for the various tensile ratio values and it was found that the knee joint force was sensitively affected by the tensile ratio and the above mentioned discrepancy between the simulation results and the in vivo data could be eliminated if the ratio value was appropriately chosen. Figure 2 shows the situation; Variations of F_n and F_t as a function of knee angle q for the various tensile force ratio r between the hamstrings and the gluteus maximus. Where, r=1.56 was determined from the biological cross-sectional areas of the hamstrings and the gluteus maximus and r=4.5 was determined so that the simulation results best fit to the in vivo data.

It has been criticized that there exist large variations of knee joint forces obtained from model analyses. And the reasons for this have been attributed to for example such facts that the model is 2D and the parameter values are incorrect. Yet, another important issue may be to find out the way how to determine the value of the synergetic muscles' force ratio with reflecting a biological rationality.

Introduction

Reliability of a gap control technique with the tensor/balancer during PS-TKA was assessed by means of fluoroscopic images after TKA.

We have developed a novel knee simulator that reproduces the active knee motion to evaluate kinematics and joint reaction forces of TKA.

There have been developed many kinds of knee simulators; Most of them are to predict TKA component wear and the others are to evaluate the kinematics and/or kinetics of TKA. The most simulators have been operated using the data of the loading and kinematics profile of the knee obtained from normal gait. Here a problem is that such variables as joint force and kinematics are the outcome caused by the application of muscles' and external forces. If so, a simulator should be operated by the muscles' and external forces so as to duplicate the in vivo condition. Other disadvantages for the current knee simulators are; a knee joint motion is made passively, the effects of the hip joint motion are not taken into account, and the maximum flexion angle is usually limited at about 100°.

Considering the above, we have developed a knee simulator with the following advantages and innovative features. First, the simulator is driven by the muscles' forces and an active knee motion is made with bearing the upper body weight. As a result, the knee shows a 3D kinematics and generates the tibio-femoral contact forces. Under this condition, the TKA performance is to be assessed. Secondly, a hip joint mechanism is also incorporated into the simulator. The lower limb motion is achieved by the synergistic function between the hip and knee joints. Under this condition, a natural knee motion is to be reproduced. Thirdly, the simulator can make complete deep knee flexion up to 180°. Thus not only the conventional TKA but also a new TKA for high flexion can be attached to it for the evaluation.

Figure 1 shows the structure of the simulator, in which both the hip and knee joints are moved in a synergistic fashion by the pull forces of four wires. The four wires are pulled by the four servomotors respectively and reproduce the functions of the mono-articular muscles ((1), (3)) and the bi-articular muscles ((2), (4)) through the multiple pulley system. It should be noted that weight A and B are not heavy enough for the inverted double pendulum to stand up straight. They are applied as counter weights so that each segment duplicate the each segmental weight of the human lower limb. Figure 2 shows a sequential representation of stand to sit features: (a) at standing, (b) at high flexion, and (c) at deep flexion. At a state of 130° knee flexion between (b) and (c), hamstrings wire (4) becomes shortest and then exhibits an eccentric contraction, thereby attaining deep flexion.

Our knee simulator can be a useful tool for the evaluation of TKA performance and may potentially substitute the in vivo experiments.

The objective of this study is to determine the knee joint forces when rising from a kneeling position. We have developed a new type of knee prosthesis which is capable of attaining Japanese style sitting. To run the simulations and experiments needed to assess the performance of this prosthesis, it is necessary to know what forces act on the knee during deep flexion. Because these data are lacking, we created a 2D mathematical model of the lower leg to help determine knee joint forces during deep flexion. Healthy subjects of ten males (age of 25±4years, height of 170.3±9.1cm, and weight of 67.0±22.2kg) and five females (25±3years, 161±7.1cm, 47.7±6.2kg) participated in the experiment. Ground reaction force and joints angles were measured using a force plate and a motion recording system respectively. The collected data were entered into our mathematical model, and the muscle forces and the knee joint forces were calculated. To verify our model, we first used it to run simulation of middle and high flexions of the knee joint. In vivo data for these actions are available in the literature, and the results from our simulation were in good agreement with these data. We then collected the data and run simulation when rising from a kneeling position under the conditions shown in Fig. 1. They were a) double leg rising (both legs are aligned) without using the arms, b) ditto but using the arms, c) single leg rising (legs are in the front and the rear respectively) without using the arms, and d) ditto but using the arms. We obtained the following results. The statistics of the maximum values on the single knee joint for each condition were; a) F_max=5.1±0.4 [BW: (force on the knee joint)/(body weight)] at knee flexion angle of Q=140±8°, b) F_max=3.2±0.9[BW] at Q=90±10°, c) F_max-d=5.4±0.5[BW] at Q_d=62±20° for the dominant leg and F_max-s=3.0±0.5[BW] at Q_s=138±6° for the supporting leg respectively, and d) F_max-d=3.9±1.5[BW] at Q_d=70±17° for the dominant, and F_max-s=2.1±0.5 [BW] at Q_s=130±11° for the supporting. We may conclude that the single leg rising should be recommended since the maximum knee joint force did not become large as long as the knee was at deep flexion. The values introduced in this study could be used to assess the strength of the knee prosthesis at deep flexion. To obtain more realistic values of the joint forces, it is necessary to determine the ratio of the forces exerted by the mono-articular and the bi-articular joint muscles.

The objective of this study is to introduce the forces acting on the knee joint while ascending from kneeling. Our research group has developed a new type of knee prosthesis which is capable of attaining complete deep knee flexion such as a Japanese style sitting, seiza. Yet we could not set up various kinds of simulation or experiment to assess the performance of our prosthesis because the data about joints' forces during the ascent from deep knee flexion are lacking. Considering this circumstance, we created a 2D mathematical model of lower limb and determined knee joint force during ascent from kneeling to apply them for the assessment of our prosthesis.

Ten male and five female healthy subjects participated in the measurement experiment. Although the measurement of subjects' physical parameters was non-invasive and direct, some parameters had to be determined by referring to the literature. The data of ground reaction force and each joint's angle during the motion were collected using a force plate and video recording system respectively. Then the muscle forces and the joints' forces were calculated through our mathematical model. In order to verify the validity of our model approach, we first introduced the data during the activities with small/middle knee flexion such as level walking and rising from a chair; these kinds of data are available in the literature. Then we found our results were in good agreement with the literature data. Next, we introduced the data during the activities with deep knee flexion; double leg ascent [Fig.1 (a)] and single leg ascent [Fig.1 (b)] from kneeling without using the upper limbs.

The statistics of the maximum values on the single knee joint for all the subjects were; during double leg ascent, F_max = 4.6±0.6 (4.3-5.2) [BW: (force on the knee joint)/(body weight)] at knee flexion angle of b =140±8 (134-147)°, during double leg ascent, F_max = 4.9±0.5 (4.0-5.6) [BW] at b = 62±33 (28-110)° for the dominant leg, and F_max = 3.0±0.5 (22.2-3.8) [BW] at b = 138±6 (130-150)° for the supporting leg respectively. We found that the moment arm length, i.e., the location of muscle insertion significantly affected the results, while ascending speeds did not affect the results much. We may conclude that the single leg ascent should be recommended since F_maxdid not become large while deep knee flexion. The values could be used for assessing the strength of our knee prosthesis from the risk analysis view point.

We have developed a new type of knee prosthesis which is capable to make 180° knee flexion, and have designated it as Complete flexion knee (CFK). Since the kinematics and kinetics of knee prosthesis vary depending not only on its articulating surface shapes but also on the stiffness of soft tissues, its performance should be assessed under various kinds of lower limb activities.

The objective of this study is to perform simulation analysis of various lower limb activities to evaluate the performance of CFK using the 2D and the 3D mathematical models. Kinematic analyses using X-ray picture or stress analyses using FEM are extensive however, kinematic analyses can not introduce stresses and FEM can not introduce kinematics. Mathematical model analyses can introduce vital information about kinematics and kinetics at the same time.

First, we carried out an in-vitro experiment using cadaver knee under the condition of passive knee flexion-extension. After that, we performed a simulation using the same parameter variables as the in-vitro experiment in order to assess the validity of our 2D and 3D models by comparing the results about the joint contact forces and kinematics with those from the experiment.

In the in-vitro experiment, the femoral bone of a cadaver knee was fixed on a jig. In order to secure the tibiofemoral contact, each muscle was pulled with constant force respectively. Then the tibia was carried through from 40° to 140° of knee flexion. The contact forces between the femur and the tibia were measured by a load sensor. During the process, fluoroscopic images were taken, and then 3D positions/orientations of the tibia relative to the femur were introduced from the images using the pattern matching method.

Our 2D and 3D models of total knee arthroplastic joint included the tibio-femoral and patello-femoral compartments, incorporating major muscles, patella tendon and primary ligaments. The patella tendon and primary ligaments were represented with non-linear springs, whose mechanical properties were determined from the literature. In our 2D model, “thigh and calf” contact was taken into account at deep knee flexion.

Using our 3D model, the simulation was performed up to 100° of knee flexion. After that we had to alternate the model from the 3D to the 2D because the patella stacked into the femoral intercondylar, the thigh-calf contact occurred and the 3D model did not introduce the converged solution.

Over all, both the experimental and simulation results were in good agreement with each other. The results from the simulation showed that the contact points were located unusually anteriorly. The post-cam contact occurred at 44° of knee flexion, indicating that the tibia was strongly pulled to the posterior. As for the contact resultant force, large differences between simulation and experiment were found. This may be because the soft tissues of the cadaver were not intact, while we determined their properties from the literature in the simulation.

The objective of this study is to determine the in-vivo knee joint kinematics of patients having specially designed knee prosthesis (Bi-surface) at sitting sedentary (seiza) state.

An increase in the demand for TKA has required improvement in the durability and flexibility of knee prostheses. One of the representative cases which have improved knee flexion is Bi-surface knee. Bi-surface knee has two joint surfaces; one for weight bearing and the other for flexion motion which has a unique ball-and-socket joint. This knee prosthesis, having been applied for two decades, has not yet been precisely analyzed how the femoral and tibial components are articulating at deep knee flexion.

Since there is no practical method to measure directly prosthetic kinematics in-vivo; we applied indirect techniques, pattern matching method to the Bi-surface patients. The method has been originated by Banks and Hodge (1964), and we have improved it in order to obtain higher and more reliable accuracies.

The number of subjects examined by X-ray apparatus was 18 knees of 14 patients (3 male and 11 female) who could attain the seiza. Patients were asked to sit at seiza state and their Bi-surface knees were X-ray photographed from lateral side. We focused if the internal rotation was shown at maximum flexion as commonly shown for a normal knee. We also represented the CAD models with the same position/orientation as the data from the pattern matching, thereby investigating the contact states between the ball and socket by viewing them from the desired direction.

The following results were introduced. The mean maximum flexion angle was 144.1° (SD=5.3°), and the mean internal rotation angle at maximum flexion was 15.2° (SD=6.6). The maximum flexion angle among all subjects was 153.3° and internal rotation was 19.5° at that flexion angle. The number of subjects which had (a) contact point(s) on the tibio-femoral and/or ball-socket surface(s) was 5 knees (2 knees had contact point on both the ball-and-socket and the tibio-femoral lateral surfaces, 3 knees had only on either surface) and the other 13 knees had a slight gap between two components.

Correlation was found between the value of the maximum flexion angle and the value of internal rotation angle at that flexion; the subjects of larger maximum flexion angle also demonstrated larger internal rotation angle. This suggests that at deep knee flexion, the tibial internal rotation may play an important role after TKA as a normal knee does. By checking the CAD representations, we found that the tibio-femoral and ball-socket surfaces were separate for most subjects at seiza state. Although serious impingements were not found, it was suggested the risk of subluxation when a patient rises up.

The limitation of our study is that we used simple still X-ray pictures. In order to assess kinematics for ascending from seiza state, kinematic analyses from fluoroscopic images are needed.

A new type of knee prosthesis capable of making deep knee flexion has been long awaited for Asian and Muslim people. Our research group has developed such a prosthesis and designated it as CFK (Complete Flexion Knee). In order to assess the performance of CFK, we have set up various kinds of simulation/experimental projects, such as a cadaveric study, a mathematical model analysis, a photoelastic analysis and FEM analysis.

For carrying out the above-mentioned projects, we faced the most fundamental problem; the information about the muscles’ forces and the forces acting on the joints is limited to that for ambulatory activities but not for squatting or sedentary sitting.

The objective of this study is to introduce the force acting on the knee joint and the lower limbs’ muscle forces at deep knee flexion. A 2D mathematical model was used. The model was composed with three segments: upper leg, lower leg, and foot. The muscle groups incorporated into our model were gluteal muscles, quadriceps including rectus femoris and the vasti, hamstrings, and calf muscles including gastrocnemius and soleus. And thigh-calf contact was assumed to take place at 130° of knee flexion. Three equations were introduced from the moment equilibrium condition about each joint. Since the number of unknowns was six, being surplus to the number of equations, several muscles were grouped into one basing upon the EMG data.

Double leg ascending motions from deep squatting with heel rising were studied for 10 healthy male subjects age of 24±2 years, height of 172±5.8 cm, weight of 66.5±8.7 Kg. The data of ground reaction force and angle of each joint during the motion were collected using a force plate and video recording system respectively. The length of each segment for each subject was directly measured. The mass of each segment and center of gravity was determined by referring to the literature.

The results demonstrated that both the normal force acting on the knee joint and the quadriceps force became maximum when knee flexion angle became 130°(the angle at which the thigh-calf contact diminished), then decreased according as the knees extended. Both of their maximum value were proportional to the subject’s body weight and about seven times larger than that. Therefore it was justified that the joint force and quadriceps force were normalized by dividing them by the body weight. Ascending speeds did not affect the values of joint and quadriceps forces unless the motion was jumping.

One third of the world population have a life style to sit sedentary on a floor. Thus far the patients who had undergone TKA surgery loose deep flexion of the knee and various designs of artificial knee joint capable of deep knee flexion have been proposed. Among them, Bi-surface knee prosthesis (Kyocera Inc., Japan) is of special interest because of its unique design with a ball-and-socket joint. Although some patients attained a sedentary sitting with this prosthesis, the X-ray studies revealed that the femoral condyles and tibial insert tended to separate at about 150° of knee flexion, indicating a risk of subluxation when standing up.

Thus we have developed CFK (Complete Flexion Knee, Japan Medical Material Co., Japan) by further improving Bi-surface knee to enable the patient to make knee flexion as much as 180°. Our CFK has a ball-and-socket joint and whose socket part is jutted to form a tibial post. Since the ball and the cam become into a single sphere and the ball-socket and post-cam joints form a spherical bearing, CFK can provide high stability and mobility at the same time.

Besides its kinematic performance, CFK has to be assessed with its strength and durability. Since the durability of an artificial knee joint is attributed to wear of the polyethylene insert, it is essential to focus on determining the stress on it. Although the FEM analyses have been most extensive for stress analysis, whose results greatly depend upon the way how to create the meshes. The stress values introduced from the FEM are the Von Misses stresses; while wear is mainly attributed to the shear stresses. For these reasons, we employed a photoelasticity for determining the magnitude and distributions of stresses on the insert.

The models of Bi-surface, CFK and a conventional posterior stabilizer knee, Scorpio NRG (Stryker Co., USA) were used for the experiments. Epoxy resins (Araldite AER 250, 2400, Ciba Geigy Co., Japan) were selected to fabricate the tibial insert models. Special equipment was used to apply 2 kg force on the model by setting knee flexion angle at 0°, 30°, 60°, 90° and 120° respectively. After that the stressed model was sliced along the anterio-posterior direction and photoelastic fringes in each slice were observed. The results demonstrated that while knee angle was smaller than 90°, shear stress on the lateral slice became higher in the order, NRG, CFK, and Bi-surface, indicating NRG has high conformity in the condylar surface. After knee angle bacame larger than 90°, shear stress on the mid-posterior slice became higher in the order, CFK, Bi-surface and NRG. We may conclude that CFK has optimal configuration at deep knee flexion not only for kinematic but also for load bearing viewpoints.

To obtain correct soft tissue balance during TKA is an important operative technique for successful clinical outcome. The soft tissue balancing has been assessed by the tibiofemoral joint gap in full extension, and at 90°. Since recent advancements in the design of femoral components, tibial articular surface and operative techniques have enabled a prosthetic deep knee flexion, the joint gap measurement in such a condition became necessary. Also it should be noted that the joint gap directly reflects on the clinical outcomes such as range of motion, laxity and instability.

In recent years, many in-vivo kinematic measurement methods were developed, which measure the 3D position and orientation from the 2D X-ray image. Among them a pattern-matching method is representative, which is the method by comparing the contour shape from the X-ray image with a predicted contour to seek the 3D position and orientation.

The objective of this study is to measure the range of motion of knee prostheses from their X-ray fluoroscopic images by using the pattern-matching method.

We analyzed 7 knee prostheses of 7 female patients, age of 59 to 77 years, height of 149.5 to 159 cm, weight of 43 kg to 72 kg. Their knee prostheses were all NRG-PS type (Striker Co., USA) with various sizes. During the fluoroscopy measurement, the patient was lying supine on a bed with her both legs free. First the patients were asked to make flexion-extension with their prosthetic knees by themselves and their fluoroscopic images were recorded for analysis. Next the following motions were done passively. Starting with 0°, the knee angle was gradually increased and fixed at 30°, 60°, 90° and up to 120° respectively. At each flexion angle, the knee was internally rotated as possible as the maximum limit of the patient capacity and then externally rotated in the same way. Similarly, the knee was made varusly and then valgusly at each flexion angle respectively.

The results of kinematic analyses were arranged by the tibial orientations relative to the femur. The range of flexion-extension angles were from 113.9° (SD=8.3°) to 5.2° (SD=8.2°). At maximum flexion for each patient, the orientation in terms of internal-external rotation and varus-valgus was measured and averaged; they were internally rotated by 6.0° (SD=0.6°) and varusly inclined by 1.2° (SD=1.0°). At full extension (minimum flexion), they were externally rotated by 4.3° (SD=1.9°) and varusly inclined by 0.1° (SD=0.7°) respectively. The maximum value of internal-external rotation range was recorded at 89.4° (SD=2.4°) of knee flexion and they were from 5.4° (SD=1.3°) of internal rotation to 12.9° (SD=6.0) of external rotation. The varus-valgus motion was small, from 1.7° (SD=1.6°) of varus to 0.1° (SD=2.2°) of valgus through the whole range of knee flexion.

Important findings were that the range of varus-valgus was smallest for the prosthesis with the thickest insert, and the knee whose collateral ligaments were loose tended to incline varusly.