Journal of Biomechanics 47 (2014) 1572–1576
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In-vivo 6 degrees-of-freedom kinematics of metal-on-polyethylene total hip arthroplasty during gait Tsung-Yuan Tsai, Jing-Sheng Li, Shaobai Wang, Donna Scarborough, Young-Min Kwon n Department of Orthopaedic Surgery, Massachusetts General Hospital, Harvard Medical School, 55 Fruit Street, Boston, MA 02114, USA
art ic l e i nf o
a b s t r a c t
Article history: Accepted 7 March 2014
Knowledge of accurate in-vivo 6 degree-of-freedom (6-DOF) kinematics of total hip arthroplasty (THA) during daily activities is critical for improvement of longevity of the components. Previous studies assessed in-vivo THA kinematics using skin marker-based motion analysis. However, skin markers are prone to move with respect to the underlying bones. A non-invasive dual fluoroscopic imaging system (DFIS) based tracking technique has been used to avoid skin artifacts and provide accurate 6-DOF kinematic measurement. This study aimed to quantify in-vivo 6-DOF THA kinematics during gait using DFIS. Twenty eight well-functioning THAs were evaluated during treadmill gait under DFIS surveillance. The maximum translations of the femoral head were 0.46 70.10 mm and 0.45 70.10 mm during the stance and swing phases (p ¼0.57), respectively. The range of hip flexion was from 8.71 to 47.61, adduction from 3.01 to 12.51 and external rotation from 19.21 to 29.71. The THA was flexed, externally rotated and adducted throughout the gait. The magnitudes of the femoral head translations were found to be within the manufacture tolerance of the components, suggesting that in-vivo hip “pistoning” during gait cycle may be minimal in well-functioning THAs. The 6-DOF kinematics could be used as the baseline knowledge for further improvement of wear-testing of hip implant, implants manufacturing and implant positioning during surgery. & 2014 Elsevier Ltd. All rights reserved.
Keywords: Total hip arthroplasty in vivo Gait Kinematics Fluoroscope
1. Introduction More than 330,000 people in the United States of America received total hip arthroplasty (THA) in 2010 (Centers for Disease Control and Prevention, 2010) and the increasing demand of THA is expected to grow to 572,000 annually by 2030 (Kurtz et al., 2007). Although significant pain relief and an improvement of the functional capacity are observed in post-operative patients (Fortina et al., 2005; Herberts and Malchau, 2000), the past and current innovation of THA has largely focused on development of improved implant bearing materials and implant fixation methods. However, the challenges in the performance of contemporary THA are related to adverse in vivo dynamic phenomenon including edge loading, impingement and dislocation. These occur as a function of both static implant factors (such as component positioning) as well as dynamic biomechanics of individual patient. Therefore, it is important to quantify the in-vivo biomechanics of the hip joint during functional activities.
n Correspondence to: Bioengineering Laboratory, Department of Orthopaedic Surgery, Massachusetts General Hospital, Harvard Medical School, 55 Fruit Street, GRJ-1215, Boston, MA 02114, USA. Tel.: þ 1 617 726 6472; fax: þ1 617 724 4392. E-mail address: [email protected] (Y.-M. Kwon).
http://dx.doi.org/10.1016/j.jbiomech.2014.03.012 0021-9290/& 2014 Elsevier Ltd. All rights reserved.
Previous studies assessed in-vivo three-dimensional (3D) hip kinematics in patients with THA during gait using motion capture systems (Bennett et al., 2008; Ewen et al., 2012; Perron et al., 2000; Sinha et al., 2011). Significant decreases in gait speed and the hip extension were observed in THA patients (Bennett et al., 2008; Perron et al., 2000). However, skin markers are prone to move in relation to the underlying bones. The soft tissue artifacts affects the measured joint kinematics (Kuo et al., 2011; Tsai et al., 2013, 2011). Other researchers adopted medical imaging techniques to measure the positions of the hip implants without the interference of soft tissues artifacts (Dennis et al., 2001; Komistek et al., 2002; Lombardi et al., 2000). In-vivo femoral translation of THA has been reported using two-dimensional (2D) to 3D registration techniques with single video fluoroscopic system and implant models (Lombardi et al., 2000). However, only few THA patients were studied in the literature. No hip rotation angles of patients with THA during a complete gait cycle have been reported using noninvasive imaging technique. Recently, a validated application of the dual fluoroscope imaging system (DFIS) was developed for measurement of in-vivo 6-degrees-of-freedom (6-DOF) kinematics of the hips in patients with THA during dynamic function (Tsai et al., 2013). The aims of this study were to quantify 6-DOF kinematics of in-vivo hip in subjects with THA during gait using the validated DFIS technique (Tsai et al., 2013), and to
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compare the hip translations between stance (weight bearing) and swing phases (non-weight bearing).
2. Materials and methods Twenty four subjects with a total of 28 cementless metal-on-polyethylene (MoP) non-constrained THAs (DJO Surgical, Encore Medical, Austin, Texas) were evaluated in this study (5 men and 19 women, with 17 right and 11 left THAs). The study was approved by the institution's Internal Review Board and each subject provided written informed consent prior to participation. The average age was 64.0 yrs ( 77.3, range 47–73). The average body height and weight were 166.8 cm ( 7 7.5, range 154.9–185.4) and 75.8 kg ( 7 18.5, range 49.0–113.4) with average BMI of 27.1 ( 75.3, range 18.5–35.8). The average follow-up time was 13.5 months ( 7 7.0, range 6.6–33.8) from surgical date. The acetabular cup sizes were from 50 to 54 mm and the femoral head sizes from 32 to 36 mm. The average cup planar anteversion and inclination were 27.61 ( 7 9.0, range 16.4–51.0) and 39.71 ( 77.9, range 24.9–59.4), respectively. All subjects included in this cohort had no history of dislocation or subluxation or report of any surgical complication. In accordance with study protocol reported in previously published validation study (Lin et al., 2013; Tsai et al., 2013), each subject received a CT scan (Sensation 64, Siemens, Germany) from the L-5 vertebra to the mid-femur using 140 kVp with image resolution of 512 512 pixels and voxel size of 0.97 0.97 0.60 mm3 for creation of surface models of the acetabular cup, femoral stem and the hip bones. Segmentation procedure was a semi-automatic process and supervised by an experienced physician. Two pelvic local coordinate systems were defined for describing hip translations and joint rotations. The origins of the two local coordinate systems were at the center of the acetabular cup (Fig. 1a and b). The orientation of the first pelvic coordinate system of the pelvis was constructed for the hip rotation using bony landmarks, including anterior and posterior superior iliac spines, following the International Society of Biomechanics (ISB) recommendation (Fig. 1a) (Wu et al., 2002). In order to describe the femoral head position with respect to the cup, the second pelvic coordinate system for the hip translation was set corresponding to the cup geometry (Fig. 1b). The z-axis was parallel to the interception line of cup opening plane and sagittal plane (Fig. 1b). The y-axis was the normal vector of the cup opening plane (Fig. 1b). The femoral coordinate system was defined according to the proximal femur and the stem. The origin of the femoral coordinate system was at the center of the femoral head (Fig. 1a). The femoral y-axis was parallel to the long axis of the proximal femoral shaft. The x-axis was parallel to the normal vector of the plane formed by the y-axis and the center of the femoral head. Each subject was asked to perform level walking on a treadmill at self-selected speed under DFIS (BV Pulsera, Phillips Medical, USA) surveillance using snapshots with a frame rate of 30 Hz (Fig. 2). The walking speed of the treadmill was recorded. The fluoroscopic images were collected with 98 cm source to image distance, 8 ms pulse width, 60–80 kV and 0.042–0.066 mAs. Two full gait cycles on the treadmill were recorded for each subject. Two thin pressure sensors (force sensor resistor, Interlink Electronics, Camarillo, CA) fixed to the bottom shoes for
a
b Yp Zp Xp
yf
zf
z x
y
xf
Fig. 1. The origin of pelvic local coordinate system located at the center of the acetabular sphere, (a and b). The origin of femoral coordinate system was at the center of the femoral head, (a and b). Two local coordinates of the acetabular cup were defined for describing the hip joint rotations (a), and the hip translations (b). The hip rotation angles were calculated using a Cardan sequence of Z–X–Y between the local coordinates of the femur relative to the pelvis (a). The 3D vector from the origin of the acetabulum to the center of the femoral head in the acetabular coordinate system was measured as the hip translation (b).
Fig. 2. A simulated virtual DFIS environment overlapping on a photo of DFIS experimental setup demonstrated an example of registration of in-vivo hip kinematics during treadmill gait. The registered hip bone and implants were superimposed on the subject. defining gait cycles during the treadmill gait (Chen et al., 2012). One pair of fluoroscopic images of the hip was captured while the subject stood at an upright position as the reference. The averaged effective dose of the CT and dual fluoroscopy procedures was 9.1 mSv (range 8.1–10.3) per subject. Surface models of the hip and fluoroscopic images were introduced into a virtual DFIS to determine the actual hip poses (Fig. 2). Models were projected onto the image intensifiers for comparison of the difference between the fluoroscopic outlines and the features of the model projection (Tsai et al., 2013). An optimization was used to minimize the objective function and to find the position of the model where its projection best matches the fluoroscopic outlines (Tsai et al., 2013). The registered poses of the models in a series of image frames can then be used to obtain the 3D kinematics of the hip joint. The position of the femoral head relative to the cup at the upright standing position was measured and considered as the reference. The hip translations and rotations of each THA during gait were then analyzed throughout gait cycles including stance (weight bearing) and swing phases (non-weight bearing). The hip rotation angles were obtained using a Cardan sequence of Z–X–Y between the local coordinates of the femur relative to the pelvis following ISB recommendation (Wu et al., 2002) (Fig. 1a). The 3D vector from the origin of the acetabulum to the center of the femoral head in the acetabular coordinate system was measured as the hip translation (Fig. 1b). Maximum, average and standard deviation of the hip translations and rotations in swing and stance phases during gait across all THAs were reported. Two-tailed paired t test was performed to compare the hip translations between the stance and swing phases (α ¼ 0.05).
3. Results The median and standard deviation of walking speeds of all the THA subjects was 2.870.5 km/h. The average maximum hip translations were 0.46 (70.10) mm for stance phase and 0.45 (70.10) mm
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Table 1 Average7 standard deviation (maximum across subjects) of maximum hip translation components along cup local axes of 28 THAs in stance and swing phases during treadmill gait were listed. The same calculation for maximum and average of the magnitudes of hip translations during stance and swing phases were also reported. P-values were calculated using two-tailed paired t test. Axis (mm)
Hip separations of 28 THA during treadmill gait Stance phase
Swing phase
P-value
x y z
0.25 7 0.13 (0.55) 0.22 7 0.13 (0.57) 0.23 7 0.12 (0.44)
0.217 0.14 (0.54) 0.25 7 0.13 (0.52) 0.22 7 0.15 (0.58)
0.24 0.22 0.80
Max translation Average translation
0.46 7 0.10 (0.70) 0.317 0.08 (0.60)
0.45 7 0.10(0.72) 0.32 7 0.09(0.57)
0.57 0.52
for swing phases during the treadmill gait in this patient group. The maximum hip translations across all the THAs were found to be 0.70 mm and 0.72 mm in stance and swing phases, respectively (Table 1). The average hip translations were 0.3170.08 mm in stance phase and 0.3270.09 mm in swing phase (Table 1). No significant differences were found in the maximum or average (P-value¼0.57, 0.52) hip translations between stance and swing phases (Table 1). No specific trend was found in the femoral head translations during the gait (Fig. 3). The maximum hip translations in medial/lateral, in/out of the cup and anterior/posterior directions were 0.2570.13 mm, 0.2270.13 mm and 0.2370.12 mm in stance phase, and 0.2170.14, 0.2570.13 and 0.2270.15 mm in swing phase (Table 1). No significant differences were found in all the components of the hip translation between stance and swing phases (Table 1). At heel strike, average hip flexion was 36.81 ( 75.21). At late stance phase the hip did not reach full extension. The average hip flexion at opposite heel contact ( 50% gait cycle) was 9.81 (7 8.71) (Fig. 3). The maximum hip flexion was 47.617 5.11 at terminal midswing ( 85% gait cycle) (Fig. 3). In the transverse plane at initial contact the average hip position was in external rotation at 23.21 (79.61), transitioning during stance to external rotation at 23.21 (77.61) at opposite heel contact. At toe off the hip was externally rotated at an average of 27.61 (78.41). Afterward, the hip kept increasing internal rotation slowly in swing phase and continued the trend till the end of gait cycle (Fig. 3). Frontal plane motion demonstrated an average hip adduction of 8.61 ( 72.31) at heel strike during early to mid-stance phase. The average hip adduction at late stance to toe off was 4.31 ( 73.51). This cohort of THA subjects demonstrated movement into adduction during swing phase (Fig. 3). Overall, the THA joint was found to be flexed, adducted and external rotated throughout the treadmill gait not attaining hip extension at terminal stance. The average of the range of motion of the THA rotations during the gait were 38.91, 9.01 and 10.51 for flexion/extension (FLEX/EXT), adduction/abduction (ADD/ABD) and internal rotation/external rotation (IR/ER), respectively (Table 2). The minimum to maximum THA rotation angles during the gait were 8.7–47.61, 3.5–12.51 and 19.2–29.71 for FLEX, ADD and ER, respectively (Table 3).
4. Discussion The challenges in the performance of contemporary THA are related to adverse in vivo dynamic phenomenon including edge loading, impingement and dislocation. Thus, accurate understanding of in vivo biomechanics of hip joint during functional activity is required to optimize outcome of patients with THA. The current study evaluated in-vivo hip translations and rotations of 28 MoP THAs during gait using a validated non-invasive DFIS technique without the effects of soft tissue artifacts. The average maximum
hip translations were less than 0.5 mm in stance and swing phases, and no significant differences were found between stance and swing phases. The average hip flexion, adduction and external rotation of all THAs ranged from 8.71 to 47.61, 3.51 to 12.51 and 19.21 to 29.71 respectively during the gait, indicating that the THA joint remained flexed, adducted and externally rotated throughout the gait. Previous studies have reported in-vivo kinematics of THA during gait using skin marker based motion analysis (Bennett et al., 2008; Kyriazis and Rigas, 2002; Perron et al., 2000). Reduced gait speed, range of hip flexion/extension, maximum hip extension and range of hip abduction/adduction were found with respect to normal elderly group (Bennett et al., 2008; Perron et al., 2000). Similar slow walking speed and no hip extension during the gait was also shown in this study. The lack of hip extension might be due to the anterior tilt of the pelvis and femoral component orientation in these THA patients. However, greater range of hip flexion/extension (Table 2), greater maximum hip flexion (Table 3), greater average hip rotation and greater range of hip ADD/ABD (Table 2) were found in this study comparing to the results of a previous study (Bennett et al., 2008). The trend of hip rotation angles in ADD/ABD and IR/ER (Fig. 3) during the gait was also different from previous observations (Bennett et al., 2008; Perron et al., 2000). No hip abduction and internal rotation during gait were found in this study (Table 3). The differences in the THA rotations among these studies might be resulted from different measurement techniques and the effects of soft tissue artifacts during gait (Stagni et al., 2005). In addition, when using skin marker based measurements, the precision in determining the positions of anatomic markers would affect the error propagation from the axes to joint kinematics (Della Croce et al., 2005). The thickness of the soft tissues surrounding the hip and individual body weight and compositions could vary case by case, which would also affect the measures of THA kinematics. In this study, the 6-DOF THA kinematics was measured according to the local coordinate systems defined on the bones and implants, which would be useful for establishing boundary conditions for finite element analysis of hip biomechanics and realistic in-vitro settings of wear-testing of hip implants. The current study indicated no specific trend of hip translation in all directions along cup local axes (Fig. 3) and the average of maximum hip translations across all the subjects during the gait was less than 0.50 mm (Table 1), which implied that the femoral head was located within the cup throughout gait cycle without significant translocation by the weight of the trailing leg in swing phase. The precision of the DFIS technique for determination of the THA translation has been reported to be 70.81 mm (Tsai et al., 2013). However, no hip translation larger than 0.72 mm was observed in any of the 28 THAs during the gait. The average magnitudes of the maximum hip translation in each cup local axes were found to be less than 0.3 mm with small variations (o0.2 mm) in both stance and swing phase during gait (Table 1), that were similar to the manufacturing tolerance of ball-in-cup clearance of 0.2 mm in radius. These findings are in inconsistent to the large hip separation in MoP THAs reported in literature during gait (10 THAs, 2.8 mm), abduction/adduction leg lift (10 THAs, 3.0 mm) (Lombardi et al., 2000) and pivoting activity (10 THAs, 3.3 mm) (Blumenfeld et al., 2011) using single fluoroscopy measurement. Small hip separations were also reported in MoP THAs during gait (15 THAs, 0.9 mm) (Glaser et al., 2008). In current study, no hip translation larger than 0.72 mm was observed in the 28 THAs during gait. The differences in hip translations among these studies might be due to various factors such as measurement techniques, different patient population, implant models, and subject's physique and movement patterns. Therefore, a direct comparison of these studies has to take these factors into consideration.
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Fig. 3. Averages and standard deviations of in-vivo hip translations and rotations of 28 THA during treadmill gait were quantified using the DFIS tracking technique. Dashed lines indicate the time of toe-off. The hip models and the men's sketches showed the corresponding THA positions and lower limb postures during gait. Table 2 Ranges of average hip rotation angles of 28 THAs in stance and swing phases during treadmill gait. Extreme angles are reported. Rotation
Stance phase
Hip angles of 28 THA during treadmill gait FLEX/EXT 9.617 8.51–36.817 5.21 ADD/ABD 4.417 3.61–11.517 4.11 IR/ER 27.51 78.31 to 21.51 77.51
Instrumented hip prosthesis for measurement of in-vivo hip contact force demonstrated that there were constant compressive forces at the THA throughout gait including stance and swing phases (Bergmann et al., 2001; Damm et al., 2013). The compressive force in swing phase have been described by others as the result of muscle co-contraction for facilitating moving the trailing
Swing phase
Range of motion
22.717 8.51–46.417 5.21 4.117 3.61–10.217 4.11 27.61 78.21 to 23.317 7.51
38.91 76.41 9.01 73.31 10.51 74.91
leg forward (Park et al., 1999). The contact force acting on the concave surface of the acetabular cup of the THA and forces from surrounding hip muscles could create cumulative compressive forces to maintain the femoral head within the acetabular cup. This implies that there would be minimal hip translations in THA during the swing phase of gait. This is consistent with the finding
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Table 3 Averages and standard deviations of maximum hip rotation angles of THA subjects during treadmill gait. Hip rotations
Maximum angle
Flexion/extension Adduction/abduction Internal/external rotation
8.717 8.31–47.617 5.11 3.517 3.41–12.517 3.81 29.717 8.91 to 19.217 8.11
of this study demonstrating no large hip translations in swing phase. Comparison of hip translations during the swing phase to the loaded weight bearing condition of the hip joint during stance phase revealed no significant differences. Therefore, in this cohort, the weight-bearing condition did not affect hip translations in this group of subjects with well-functioning THA. The results of the current study need to be interpreted in light of several limitations. Firstly, treadmill gait is not a true replicate of normal overground gait, however, previous study showed that treadmill gait is qualitatively and quantitatively similar to overground gait (Riley et al., 2007). Secondly, the subjects were asked to perform treadmill gait under DFIS surveillance at self-selected speed for safety consideration. The THA patients walk slower than normal health people on average. The hip kinematics of the treadmill gait in the environment may differ from level walking on the ground. Thirdly, the semi-automatic segmentation of CT images for bone model construction requires supervision of an experienced physician. Finally, all the subjects in the current study had MoP THAs with an average follow-up time of 13.5 months. Thus, the findings of this study cannot be generalized to different types of THA bearing surfaces with different follow-up time during different loading conditions. Current study quantified and described accurate in-vivo 6-DOF kinematics of THA during treadmill gait using a validated imaging technique. The THA joint was found to remain flexed, adducted and externally rotated throughout the gait. The magnitudes of hip translations were found to be within the manufacture tolerance of the implant components, suggesting that in-vivo hip “pistoning” during gait cycle may be minimal among these THAs. No significant difference was found in hip translations between stance and swing phases, indicating that weight-bearing does not significantly affect the hip translation during gait. The results of current study provide a framework in gaining insights into the understanding of the in-vivo function of THAs during functional activities. Further studies are necessary to investigate different types of THA bearing surfaces during different loading conditions.
Funding MAKO Surgical Corp.
Ethical approval Research protocol has been reviewed and approved by Massachusetts General Hospital IRB.
Conflict of interest The authors of this manuscript have nothing to disclose that would bias our work.
Acknowledgments Technical guidance of Dr. Harry E. Rubash and Dr. Guoan Li is greatly appreciated. References Bennett, D., Humphreys, L., O'Brien, S., Kelly, C., Orr, J.F., Beverland, D.E., 2008. Gait kinematics of age-stratified hip replacement patients—a large scale, long-term follow-up study. Gait Posture 28, 194–200. Bergmann, G., Deuretzbacher, G., Heller, M., Graichen, F., Rohlmann, A., Strauss, J., Duda, G.N., 2001. Hip contact forces and gait patterns from routine activities. J. Biomech. 34, 859–871. Blumenfeld, T.J., Glaser, D.A., Bargar, W.L., Langston, G.D., Mahfouz, M.R., Komistek, R.D., 2011. In vivo assessment of total hip femoral head separation from the acetabular cup during 4 common daily activities. Orthopedics 34, 127. Chen, C.H., Li, J.S., Hosseini, A., Gadikota, H.R., Gill, T.J., Li, G., 2012. Anteroposterior stability of the knee during the stance phase of gait after anterior cruciate ligament deficiency. Gait Posture 35, 467–471. 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In-vivo 6 degrees-of-freedom kinematics of metal-on-polyethylene total hip arthroplasty during gait Tsung-Yuan Tsai, Jing-Sheng Li, Shaobai Wang, Donna Scarborough, Young-Min Kwon n Department of Orthopaedic Surgery, Massachusetts General Hospital, Harvard Medical School, 55 Fruit Street, Boston, MA 02114, USA
art ic l e i nf o
a b s t r a c t
Article history: Accepted 7 March 2014
Knowledge of accurate in-vivo 6 degree-of-freedom (6-DOF) kinematics of total hip arthroplasty (THA) during daily activities is critical for improvement of longevity of the components. Previous studies assessed in-vivo THA kinematics using skin marker-based motion analysis. However, skin markers are prone to move with respect to the underlying bones. A non-invasive dual fluoroscopic imaging system (DFIS) based tracking technique has been used to avoid skin artifacts and provide accurate 6-DOF kinematic measurement. This study aimed to quantify in-vivo 6-DOF THA kinematics during gait using DFIS. Twenty eight well-functioning THAs were evaluated during treadmill gait under DFIS surveillance. The maximum translations of the femoral head were 0.46 70.10 mm and 0.45 70.10 mm during the stance and swing phases (p ¼0.57), respectively. The range of hip flexion was from 8.71 to 47.61, adduction from 3.01 to 12.51 and external rotation from 19.21 to 29.71. The THA was flexed, externally rotated and adducted throughout the gait. The magnitudes of the femoral head translations were found to be within the manufacture tolerance of the components, suggesting that in-vivo hip “pistoning” during gait cycle may be minimal in well-functioning THAs. The 6-DOF kinematics could be used as the baseline knowledge for further improvement of wear-testing of hip implant, implants manufacturing and implant positioning during surgery. & 2014 Elsevier Ltd. All rights reserved.
Keywords: Total hip arthroplasty in vivo Gait Kinematics Fluoroscope
1. Introduction More than 330,000 people in the United States of America received total hip arthroplasty (THA) in 2010 (Centers for Disease Control and Prevention, 2010) and the increasing demand of THA is expected to grow to 572,000 annually by 2030 (Kurtz et al., 2007). Although significant pain relief and an improvement of the functional capacity are observed in post-operative patients (Fortina et al., 2005; Herberts and Malchau, 2000), the past and current innovation of THA has largely focused on development of improved implant bearing materials and implant fixation methods. However, the challenges in the performance of contemporary THA are related to adverse in vivo dynamic phenomenon including edge loading, impingement and dislocation. These occur as a function of both static implant factors (such as component positioning) as well as dynamic biomechanics of individual patient. Therefore, it is important to quantify the in-vivo biomechanics of the hip joint during functional activities.
n Correspondence to: Bioengineering Laboratory, Department of Orthopaedic Surgery, Massachusetts General Hospital, Harvard Medical School, 55 Fruit Street, GRJ-1215, Boston, MA 02114, USA. Tel.: þ 1 617 726 6472; fax: þ1 617 724 4392. E-mail address: [email protected] (Y.-M. Kwon).
http://dx.doi.org/10.1016/j.jbiomech.2014.03.012 0021-9290/& 2014 Elsevier Ltd. All rights reserved.
Previous studies assessed in-vivo three-dimensional (3D) hip kinematics in patients with THA during gait using motion capture systems (Bennett et al., 2008; Ewen et al., 2012; Perron et al., 2000; Sinha et al., 2011). Significant decreases in gait speed and the hip extension were observed in THA patients (Bennett et al., 2008; Perron et al., 2000). However, skin markers are prone to move in relation to the underlying bones. The soft tissue artifacts affects the measured joint kinematics (Kuo et al., 2011; Tsai et al., 2013, 2011). Other researchers adopted medical imaging techniques to measure the positions of the hip implants without the interference of soft tissues artifacts (Dennis et al., 2001; Komistek et al., 2002; Lombardi et al., 2000). In-vivo femoral translation of THA has been reported using two-dimensional (2D) to 3D registration techniques with single video fluoroscopic system and implant models (Lombardi et al., 2000). However, only few THA patients were studied in the literature. No hip rotation angles of patients with THA during a complete gait cycle have been reported using noninvasive imaging technique. Recently, a validated application of the dual fluoroscope imaging system (DFIS) was developed for measurement of in-vivo 6-degrees-of-freedom (6-DOF) kinematics of the hips in patients with THA during dynamic function (Tsai et al., 2013). The aims of this study were to quantify 6-DOF kinematics of in-vivo hip in subjects with THA during gait using the validated DFIS technique (Tsai et al., 2013), and to
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compare the hip translations between stance (weight bearing) and swing phases (non-weight bearing).
2. Materials and methods Twenty four subjects with a total of 28 cementless metal-on-polyethylene (MoP) non-constrained THAs (DJO Surgical, Encore Medical, Austin, Texas) were evaluated in this study (5 men and 19 women, with 17 right and 11 left THAs). The study was approved by the institution's Internal Review Board and each subject provided written informed consent prior to participation. The average age was 64.0 yrs ( 77.3, range 47–73). The average body height and weight were 166.8 cm ( 7 7.5, range 154.9–185.4) and 75.8 kg ( 7 18.5, range 49.0–113.4) with average BMI of 27.1 ( 75.3, range 18.5–35.8). The average follow-up time was 13.5 months ( 7 7.0, range 6.6–33.8) from surgical date. The acetabular cup sizes were from 50 to 54 mm and the femoral head sizes from 32 to 36 mm. The average cup planar anteversion and inclination were 27.61 ( 7 9.0, range 16.4–51.0) and 39.71 ( 77.9, range 24.9–59.4), respectively. All subjects included in this cohort had no history of dislocation or subluxation or report of any surgical complication. In accordance with study protocol reported in previously published validation study (Lin et al., 2013; Tsai et al., 2013), each subject received a CT scan (Sensation 64, Siemens, Germany) from the L-5 vertebra to the mid-femur using 140 kVp with image resolution of 512 512 pixels and voxel size of 0.97 0.97 0.60 mm3 for creation of surface models of the acetabular cup, femoral stem and the hip bones. Segmentation procedure was a semi-automatic process and supervised by an experienced physician. Two pelvic local coordinate systems were defined for describing hip translations and joint rotations. The origins of the two local coordinate systems were at the center of the acetabular cup (Fig. 1a and b). The orientation of the first pelvic coordinate system of the pelvis was constructed for the hip rotation using bony landmarks, including anterior and posterior superior iliac spines, following the International Society of Biomechanics (ISB) recommendation (Fig. 1a) (Wu et al., 2002). In order to describe the femoral head position with respect to the cup, the second pelvic coordinate system for the hip translation was set corresponding to the cup geometry (Fig. 1b). The z-axis was parallel to the interception line of cup opening plane and sagittal plane (Fig. 1b). The y-axis was the normal vector of the cup opening plane (Fig. 1b). The femoral coordinate system was defined according to the proximal femur and the stem. The origin of the femoral coordinate system was at the center of the femoral head (Fig. 1a). The femoral y-axis was parallel to the long axis of the proximal femoral shaft. The x-axis was parallel to the normal vector of the plane formed by the y-axis and the center of the femoral head. Each subject was asked to perform level walking on a treadmill at self-selected speed under DFIS (BV Pulsera, Phillips Medical, USA) surveillance using snapshots with a frame rate of 30 Hz (Fig. 2). The walking speed of the treadmill was recorded. The fluoroscopic images were collected with 98 cm source to image distance, 8 ms pulse width, 60–80 kV and 0.042–0.066 mAs. Two full gait cycles on the treadmill were recorded for each subject. Two thin pressure sensors (force sensor resistor, Interlink Electronics, Camarillo, CA) fixed to the bottom shoes for
a
b Yp Zp Xp
yf
zf
z x
y
xf
Fig. 1. The origin of pelvic local coordinate system located at the center of the acetabular sphere, (a and b). The origin of femoral coordinate system was at the center of the femoral head, (a and b). Two local coordinates of the acetabular cup were defined for describing the hip joint rotations (a), and the hip translations (b). The hip rotation angles were calculated using a Cardan sequence of Z–X–Y between the local coordinates of the femur relative to the pelvis (a). The 3D vector from the origin of the acetabulum to the center of the femoral head in the acetabular coordinate system was measured as the hip translation (b).
Fig. 2. A simulated virtual DFIS environment overlapping on a photo of DFIS experimental setup demonstrated an example of registration of in-vivo hip kinematics during treadmill gait. The registered hip bone and implants were superimposed on the subject. defining gait cycles during the treadmill gait (Chen et al., 2012). One pair of fluoroscopic images of the hip was captured while the subject stood at an upright position as the reference. The averaged effective dose of the CT and dual fluoroscopy procedures was 9.1 mSv (range 8.1–10.3) per subject. Surface models of the hip and fluoroscopic images were introduced into a virtual DFIS to determine the actual hip poses (Fig. 2). Models were projected onto the image intensifiers for comparison of the difference between the fluoroscopic outlines and the features of the model projection (Tsai et al., 2013). An optimization was used to minimize the objective function and to find the position of the model where its projection best matches the fluoroscopic outlines (Tsai et al., 2013). The registered poses of the models in a series of image frames can then be used to obtain the 3D kinematics of the hip joint. The position of the femoral head relative to the cup at the upright standing position was measured and considered as the reference. The hip translations and rotations of each THA during gait were then analyzed throughout gait cycles including stance (weight bearing) and swing phases (non-weight bearing). The hip rotation angles were obtained using a Cardan sequence of Z–X–Y between the local coordinates of the femur relative to the pelvis following ISB recommendation (Wu et al., 2002) (Fig. 1a). The 3D vector from the origin of the acetabulum to the center of the femoral head in the acetabular coordinate system was measured as the hip translation (Fig. 1b). Maximum, average and standard deviation of the hip translations and rotations in swing and stance phases during gait across all THAs were reported. Two-tailed paired t test was performed to compare the hip translations between the stance and swing phases (α ¼ 0.05).
3. Results The median and standard deviation of walking speeds of all the THA subjects was 2.870.5 km/h. The average maximum hip translations were 0.46 (70.10) mm for stance phase and 0.45 (70.10) mm
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Table 1 Average7 standard deviation (maximum across subjects) of maximum hip translation components along cup local axes of 28 THAs in stance and swing phases during treadmill gait were listed. The same calculation for maximum and average of the magnitudes of hip translations during stance and swing phases were also reported. P-values were calculated using two-tailed paired t test. Axis (mm)
Hip separations of 28 THA during treadmill gait Stance phase
Swing phase
P-value
x y z
0.25 7 0.13 (0.55) 0.22 7 0.13 (0.57) 0.23 7 0.12 (0.44)
0.217 0.14 (0.54) 0.25 7 0.13 (0.52) 0.22 7 0.15 (0.58)
0.24 0.22 0.80
Max translation Average translation
0.46 7 0.10 (0.70) 0.317 0.08 (0.60)
0.45 7 0.10(0.72) 0.32 7 0.09(0.57)
0.57 0.52
for swing phases during the treadmill gait in this patient group. The maximum hip translations across all the THAs were found to be 0.70 mm and 0.72 mm in stance and swing phases, respectively (Table 1). The average hip translations were 0.3170.08 mm in stance phase and 0.3270.09 mm in swing phase (Table 1). No significant differences were found in the maximum or average (P-value¼0.57, 0.52) hip translations between stance and swing phases (Table 1). No specific trend was found in the femoral head translations during the gait (Fig. 3). The maximum hip translations in medial/lateral, in/out of the cup and anterior/posterior directions were 0.2570.13 mm, 0.2270.13 mm and 0.2370.12 mm in stance phase, and 0.2170.14, 0.2570.13 and 0.2270.15 mm in swing phase (Table 1). No significant differences were found in all the components of the hip translation between stance and swing phases (Table 1). At heel strike, average hip flexion was 36.81 ( 75.21). At late stance phase the hip did not reach full extension. The average hip flexion at opposite heel contact ( 50% gait cycle) was 9.81 (7 8.71) (Fig. 3). The maximum hip flexion was 47.617 5.11 at terminal midswing ( 85% gait cycle) (Fig. 3). In the transverse plane at initial contact the average hip position was in external rotation at 23.21 (79.61), transitioning during stance to external rotation at 23.21 (77.61) at opposite heel contact. At toe off the hip was externally rotated at an average of 27.61 (78.41). Afterward, the hip kept increasing internal rotation slowly in swing phase and continued the trend till the end of gait cycle (Fig. 3). Frontal plane motion demonstrated an average hip adduction of 8.61 ( 72.31) at heel strike during early to mid-stance phase. The average hip adduction at late stance to toe off was 4.31 ( 73.51). This cohort of THA subjects demonstrated movement into adduction during swing phase (Fig. 3). Overall, the THA joint was found to be flexed, adducted and external rotated throughout the treadmill gait not attaining hip extension at terminal stance. The average of the range of motion of the THA rotations during the gait were 38.91, 9.01 and 10.51 for flexion/extension (FLEX/EXT), adduction/abduction (ADD/ABD) and internal rotation/external rotation (IR/ER), respectively (Table 2). The minimum to maximum THA rotation angles during the gait were 8.7–47.61, 3.5–12.51 and 19.2–29.71 for FLEX, ADD and ER, respectively (Table 3).
4. Discussion The challenges in the performance of contemporary THA are related to adverse in vivo dynamic phenomenon including edge loading, impingement and dislocation. Thus, accurate understanding of in vivo biomechanics of hip joint during functional activity is required to optimize outcome of patients with THA. The current study evaluated in-vivo hip translations and rotations of 28 MoP THAs during gait using a validated non-invasive DFIS technique without the effects of soft tissue artifacts. The average maximum
hip translations were less than 0.5 mm in stance and swing phases, and no significant differences were found between stance and swing phases. The average hip flexion, adduction and external rotation of all THAs ranged from 8.71 to 47.61, 3.51 to 12.51 and 19.21 to 29.71 respectively during the gait, indicating that the THA joint remained flexed, adducted and externally rotated throughout the gait. Previous studies have reported in-vivo kinematics of THA during gait using skin marker based motion analysis (Bennett et al., 2008; Kyriazis and Rigas, 2002; Perron et al., 2000). Reduced gait speed, range of hip flexion/extension, maximum hip extension and range of hip abduction/adduction were found with respect to normal elderly group (Bennett et al., 2008; Perron et al., 2000). Similar slow walking speed and no hip extension during the gait was also shown in this study. The lack of hip extension might be due to the anterior tilt of the pelvis and femoral component orientation in these THA patients. However, greater range of hip flexion/extension (Table 2), greater maximum hip flexion (Table 3), greater average hip rotation and greater range of hip ADD/ABD (Table 2) were found in this study comparing to the results of a previous study (Bennett et al., 2008). The trend of hip rotation angles in ADD/ABD and IR/ER (Fig. 3) during the gait was also different from previous observations (Bennett et al., 2008; Perron et al., 2000). No hip abduction and internal rotation during gait were found in this study (Table 3). The differences in the THA rotations among these studies might be resulted from different measurement techniques and the effects of soft tissue artifacts during gait (Stagni et al., 2005). In addition, when using skin marker based measurements, the precision in determining the positions of anatomic markers would affect the error propagation from the axes to joint kinematics (Della Croce et al., 2005). The thickness of the soft tissues surrounding the hip and individual body weight and compositions could vary case by case, which would also affect the measures of THA kinematics. In this study, the 6-DOF THA kinematics was measured according to the local coordinate systems defined on the bones and implants, which would be useful for establishing boundary conditions for finite element analysis of hip biomechanics and realistic in-vitro settings of wear-testing of hip implants. The current study indicated no specific trend of hip translation in all directions along cup local axes (Fig. 3) and the average of maximum hip translations across all the subjects during the gait was less than 0.50 mm (Table 1), which implied that the femoral head was located within the cup throughout gait cycle without significant translocation by the weight of the trailing leg in swing phase. The precision of the DFIS technique for determination of the THA translation has been reported to be 70.81 mm (Tsai et al., 2013). However, no hip translation larger than 0.72 mm was observed in any of the 28 THAs during the gait. The average magnitudes of the maximum hip translation in each cup local axes were found to be less than 0.3 mm with small variations (o0.2 mm) in both stance and swing phase during gait (Table 1), that were similar to the manufacturing tolerance of ball-in-cup clearance of 0.2 mm in radius. These findings are in inconsistent to the large hip separation in MoP THAs reported in literature during gait (10 THAs, 2.8 mm), abduction/adduction leg lift (10 THAs, 3.0 mm) (Lombardi et al., 2000) and pivoting activity (10 THAs, 3.3 mm) (Blumenfeld et al., 2011) using single fluoroscopy measurement. Small hip separations were also reported in MoP THAs during gait (15 THAs, 0.9 mm) (Glaser et al., 2008). In current study, no hip translation larger than 0.72 mm was observed in the 28 THAs during gait. The differences in hip translations among these studies might be due to various factors such as measurement techniques, different patient population, implant models, and subject's physique and movement patterns. Therefore, a direct comparison of these studies has to take these factors into consideration.
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Fig. 3. Averages and standard deviations of in-vivo hip translations and rotations of 28 THA during treadmill gait were quantified using the DFIS tracking technique. Dashed lines indicate the time of toe-off. The hip models and the men's sketches showed the corresponding THA positions and lower limb postures during gait. Table 2 Ranges of average hip rotation angles of 28 THAs in stance and swing phases during treadmill gait. Extreme angles are reported. Rotation
Stance phase
Hip angles of 28 THA during treadmill gait FLEX/EXT 9.617 8.51–36.817 5.21 ADD/ABD 4.417 3.61–11.517 4.11 IR/ER 27.51 78.31 to 21.51 77.51
Instrumented hip prosthesis for measurement of in-vivo hip contact force demonstrated that there were constant compressive forces at the THA throughout gait including stance and swing phases (Bergmann et al., 2001; Damm et al., 2013). The compressive force in swing phase have been described by others as the result of muscle co-contraction for facilitating moving the trailing
Swing phase
Range of motion
22.717 8.51–46.417 5.21 4.117 3.61–10.217 4.11 27.61 78.21 to 23.317 7.51
38.91 76.41 9.01 73.31 10.51 74.91
leg forward (Park et al., 1999). The contact force acting on the concave surface of the acetabular cup of the THA and forces from surrounding hip muscles could create cumulative compressive forces to maintain the femoral head within the acetabular cup. This implies that there would be minimal hip translations in THA during the swing phase of gait. This is consistent with the finding
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Table 3 Averages and standard deviations of maximum hip rotation angles of THA subjects during treadmill gait. Hip rotations
Maximum angle
Flexion/extension Adduction/abduction Internal/external rotation
8.717 8.31–47.617 5.11 3.517 3.41–12.517 3.81 29.717 8.91 to 19.217 8.11
of this study demonstrating no large hip translations in swing phase. Comparison of hip translations during the swing phase to the loaded weight bearing condition of the hip joint during stance phase revealed no significant differences. Therefore, in this cohort, the weight-bearing condition did not affect hip translations in this group of subjects with well-functioning THA. The results of the current study need to be interpreted in light of several limitations. Firstly, treadmill gait is not a true replicate of normal overground gait, however, previous study showed that treadmill gait is qualitatively and quantitatively similar to overground gait (Riley et al., 2007). Secondly, the subjects were asked to perform treadmill gait under DFIS surveillance at self-selected speed for safety consideration. The THA patients walk slower than normal health people on average. The hip kinematics of the treadmill gait in the environment may differ from level walking on the ground. Thirdly, the semi-automatic segmentation of CT images for bone model construction requires supervision of an experienced physician. Finally, all the subjects in the current study had MoP THAs with an average follow-up time of 13.5 months. Thus, the findings of this study cannot be generalized to different types of THA bearing surfaces with different follow-up time during different loading conditions. Current study quantified and described accurate in-vivo 6-DOF kinematics of THA during treadmill gait using a validated imaging technique. The THA joint was found to remain flexed, adducted and externally rotated throughout the gait. The magnitudes of hip translations were found to be within the manufacture tolerance of the implant components, suggesting that in-vivo hip “pistoning” during gait cycle may be minimal among these THAs. No significant difference was found in hip translations between stance and swing phases, indicating that weight-bearing does not significantly affect the hip translation during gait. The results of current study provide a framework in gaining insights into the understanding of the in-vivo function of THAs during functional activities. Further studies are necessary to investigate different types of THA bearing surfaces during different loading conditions.
Funding MAKO Surgical Corp.
Ethical approval Research protocol has been reviewed and approved by Massachusetts General Hospital IRB.
Conflict of interest The authors of this manuscript have nothing to disclose that would bias our work.
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