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The impact of adding trunk motion to the interpretation of the role of joint moments during normal walking

Journal of Biomechanics, 2007
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Journal of Biomechanics 40 (2007) 3563–3569 The impact of adding trunk motion to the interpretation of the role of joint moments during normal walking Mausam Patel a,Ã , Mukul Talaty b,1 , Sylvia O ˜ unpuu a,2 a Center for Motion Analysis, Connecticut Children’s Medical Center, 282 Washington Street, Hartford, CT 06106, USA b Gait and Motion Analysis, MossRehab, 60 East Township Line Road, Elkins Park, PA 19027, USA Accepted 2 June 2007 Abstract Biomechanical model assumptions affect the interpretation of the role of the muscle or joint moments to the segmental power estimated by induced acceleration analysis (IAA). We evaluated the effect of modeling the pelvis and trunk segments as two separate segments (8 SM) versus as a single segment (7 SM) on the segmental power, support of the body, knee and hip extension acceleration produced by the joint moments during the stance phase of normal walking. Significant differences were observed in the contribution of the stance hip abductor and extensor moments to support, ipsilateral knee and hip acceleration, and ipsilateral thigh and upper body power. The primary finding was that the role of the stance hip moment in generating ipsilateral thigh and upper body power differed based on degrees of freedom in the model. Secondarily, the magnitude of contributions also differed. For example, the hip abductor and extensor moments showed greater contribution to support, hip and knee acceleration in the 8 SM. IAA and segment power analysis are sensitive to the degrees of freedom between the pelvis and trunk. There is currently no gold standard by which to evaluate the accuracy of IAA predictions. However, modeling the pelvis and trunk as separate segments is closer to the anatomical architecture of the body. An 8 SM appears to be more appropriate for estimating the role of joint moments, particularly to motion of more proximal segments during normal walking. r 2007 Elsevier Ltd. All rights reserved. Keywords: Induced acceleration; Segmental power; Trunk motion; Normal gait; Joint moments 1. Introduction Recent modeling and simulation studies have used basic principles of linked segment dynamics to understand how a muscle force or joint moment contributes to the motion of segments and joints of the entire body (Anderson and Pandy, 2003; Arnold et al., 2005; Kepple et al., 1997; Liu et al., 2006; Neptune et al., 2001, 2004; Riley et al., 2001a, b). This analysis is called induced acceleration analysis (IAA). This approach has also been expanded to segmental power analysis to study mechanical energy flow (Fregly and Zajac, 1996; Neptune et al., 2001, 2004; Seigel et al., 2004). Varying degrees of complexity exist in the models used in the above studies, depending on the objective of the study and feasibility of incorporating the complexity in the model. The extent to which biomechanical model assumptions affect the interpretations of muscle function is not fully clear. Some researchers have used three-dimensional models in their studies (Anderson and Pandy, 2003; Arnold et al., 2005; Kepple et al., 1997; Liu et al., 2006; Riley et al., 2001b; Siegel et al., 2001), whereas others have used planar models (Fregly and Zajac, 1996; Higginson et al., 2006; Neptune et al., 2001, 2004). The types of foot–floor interactions used in the above studies also vary. All the studies except the one using Anderson’s model (Anderson and Pandy, 2003) have modeled everything above the hip joints as a single segment. Some of the results from the above studies have differed depending on the model used (Anderson and Pandy, 2003; Kepple et al., 1997; Neptune ARTICLE IN PRESS www.elsevier.com/locate/jbiomech www.JBiomech.com 0021-9290/$ - see front matter r 2007 Elsevier Ltd. All rights reserved. doi:10.1016/j.jbiomech.2007.06.031 Ã Corresponding author. Tel.: +1 860 380 0164; fax: +1 860 545 8842. E-mail addresses: Mvpatel@ccmckids.org (M. Patel), mctalaty@einstein.edu (M. Talaty), Sounpuu@ccmckids.org (S. O ˜ unpuu). 1 Tel.: +1 215 663 6682. 2 Tel.: +1 860 545 8710.
et al., 2004; Riley et al., 2001b). Though some of the recent studies have started using Anderson’s 10-segment model, where the pelvis and trunk are modeled as separate segments, to our knowledge there exists no literature that documents the impact on the IAA results of modeling the pelvis and trunk segments as two separate segments versus as a single segment. The effect of assuming the pelvis and trunk as a single segment on the IAA results has not been reported, despite clear relative motion between the trunk and pelvis during walking. (Riley et al., 2001c). The combined pelvis–trunk segment assumption may especially not hold for pathologic gait, as observed in patients with myelomeningocele who exhibit excessive trunk motion relative to the pelvis (Ounpuu et al., 2000). Before performing IAA on the gait of the patients with myelomeningocele, we aimed to clarify how upper body modeling assumptions affected the interpretation of the role of muscles during walking. We studied segmental power, knee and hip extension accelera- tion and support of the body (total body center of mass vertical acceleration) produced by all the joint moments to infer muscle function. 2. Methods Motion data from five healthy subjects (65710 kg; 1.6570.11 m; 2371 year; two males and three females) with no known pathology were collected at Connecticut Children’s Medical Center during normal walking. Informed consent was obtained from the subjects participating in the data collection. The subjects were instrumented with 16 reflective markers on the lower extremities, pelvis and trunk. The subjects were asked to walk barefoot at a self-selected pace along a pathway on which three force plates (AMTI, Model # OR6, Watertown, MA) were mounted in series. Data were collected from each subject using the 512 Vicon Motion Systems (Vicon Motion Systems, Lake Forest, CA). The testing protocol for these subjects has been described elsewhere (Davis et al., 1991). A single trial for each subject was chosen, in which three consecutive steps contacted the three force plates. Joint moments and powers were computed using standard inverse dynamics approach (Davis et al., 1991) using anthropometric data from Dumas et al. (2007). The load summation for the most proximal segment computed using inverse dynamics might not equal the product of mass/inertia and measured acceleration. Skin movement artifact, error in the estimated segment anthropometry, error in locating the joint center location, etc. are possible reasons (Plamondon et al., 1996). Loads that make the equations of motion of the most proximal segment dynamically consistent were calculated for the upper body and trunk in the 7- and 8-segment models, respectively, and considered when calculating the model output sum. The three-dimensional biomechanical model was created in ADAMS (MSC Software, Ann Arbor, MI) for each subject. The basis of the biomechanical model used in this work was developed based on the work done by M. Patel during the research component of her masters work that was carried out at the Gait and Motion Analysis Laboratory at Moss Rehab. Two different three-dimensional (3D) models were created for each subject, a 7-segment (7 SM), 22 degrees-of-freedom (DOF) [Upper body segment (combined head, arms, torso and pelvis) and bilateral thigh, shank, feet] model and an 8-segment (8 SM), 25 DOF [trunk (combined head, arms and torso), pelvis, and bilateral thigh, shank, feet] model. For both the models, the ankle was modeled as a universal joint, and the hip and knee as spherical joints. The knee was modeled as a spherical joint for upcoming analysis of patients with myelomeningocele who may exhibit some motion in all three planes at the knee (Ounpuu et al., 2000). For the 8 SM the lumbar joint (joint between the pelvis and trunk) was modeled as a spherical joint. The motion and orientation of the upper body segment in the 7 SM was tracked using the pelvis markers but the anthropometry of the combined pelvis and trunk was used, similar to Kepple et al. (1997). Foot–floor interaction was modeled as a revolute joint at the measured center of pressure (CoP) obtained from the forceplates during the entire stance phase. The segmental power, knee and hip extension acceleration and support of the whole body produced by the joint moments and gravity during the stance phase were analyzed for both the 7 and 8 SMs. The total body center of mass (COG) vertical acceleration was assumed to be representative of the support of the body. Input to the model was body configuration (position and orientation), location of the center-of- pressure, computed joint moments and the offset loads. The model was reconfigured for each analysis frame using motion capture data. Each joint moment’s contribution to segmental power, COG vertical acceleration and joint angular accelerations was calculated for each frame by the solver module in ADAMS, using methods similar to ones previously described (Kepple et al., 1997; Seigel et al., 2004). For the 8 SM, the lumbar joint moments were also applied. For ease of comparison of the upper body power for the two models the power of the pelvis and trunk segments were summed to estimate the contribution of each joint moment to upper body power for the 8 SM. This was then compared to the upper body power of the 7 SM. The segmental power results were normalized to body mass to compare the data across subjects. The model output sum was compared with the results computed from the motion capture data to validate the model (Seigel et al., 2004). The term ‘reconstruction’ was used to describe the computed model output sum. The contribution of the coriolis and centripetal forces during normal walking was assumed to be negligible. 3. Results Induced acceleration and segmental power using the two different models were computed for the five subjects. The results for a representative subject are graphed in Figs. 1–3, 5 and 7; and the averages of the five subjects are graphed in Figs. 4, 6 and 8. The inverse dynamics solution was used to verify the model (Kepple et al., 1997). The data computed from the model output were reasonable compared to the motion data (Fig. 1). Similar reconstructions were obtained for the two models (Fig. 1). 3.1. Segmental power Key differences in hip contribution to upper body power between the 7 and 8 SMs were noted (Fig. 2). The stance hip moment added energy to the upper body in early stance and removed energy from the upper body during late stance for the 7 SM, whereas the opposite was observed for the 8 SM. The magnitude of the power delivered or removed from the thigh, upper body and contra lateral leg by the hip moment was considerably greater for the 8 SM. The ankle moment removed less power from the upper body segment and for a shorter duration during early stance for the 8 SM. 3.2. Support of the body The contribution of the joint moments to the COG vertical acceleration for the 7 and 8 SMs are shown in Figs. 3 ARTICLE IN PRESS M. Patel et al. / Journal of Biomechanics 40 (2007) 3563–3569 3564
ARTICLE IN PRESS Journal of Biomechanics 40 (2007) 3563–3569 www.elsevier.com/locate/jbiomech www.JBiomech.com The impact of adding trunk motion to the interpretation of the role of joint moments during normal walking Mausam Patela,, Mukul Talatyb,1, Sylvia Õunpuua,2 a Center for Motion Analysis, Connecticut Children’s Medical Center, 282 Washington Street, Hartford, CT 06106, USA b Gait and Motion Analysis, MossRehab, 60 East Township Line Road, Elkins Park, PA 19027, USA Accepted 2 June 2007 Abstract Biomechanical model assumptions affect the interpretation of the role of the muscle or joint moments to the segmental power estimated by induced acceleration analysis (IAA). We evaluated the effect of modeling the pelvis and trunk segments as two separate segments (8 SM) versus as a single segment (7 SM) on the segmental power, support of the body, knee and hip extension acceleration produced by the joint moments during the stance phase of normal walking. Significant differences were observed in the contribution of the stance hip abductor and extensor moments to support, ipsilateral knee and hip acceleration, and ipsilateral thigh and upper body power. The primary finding was that the role of the stance hip moment in generating ipsilateral thigh and upper body power differed based on degrees of freedom in the model. Secondarily, the magnitude of contributions also differed. For example, the hip abductor and extensor moments showed greater contribution to support, hip and knee acceleration in the 8 SM. IAA and segment power analysis are sensitive to the degrees of freedom between the pelvis and trunk. There is currently no gold standard by which to evaluate the accuracy of IAA predictions. However, modeling the pelvis and trunk as separate segments is closer to the anatomical architecture of the body. An 8 SM appears to be more appropriate for estimating the role of joint moments, particularly to motion of more proximal segments during normal walking. r 2007 Elsevier Ltd. All rights reserved. Keywords: Induced acceleration; Segmental power; Trunk motion; Normal gait; Joint moments 1. Introduction Recent modeling and simulation studies have used basic principles of linked segment dynamics to understand how a muscle force or joint moment contributes to the motion of segments and joints of the entire body (Anderson and Pandy, 2003; Arnold et al., 2005; Kepple et al., 1997; Liu et al., 2006; Neptune et al., 2001, 2004; Riley et al., 2001a, b). This analysis is called induced acceleration analysis (IAA). This approach has also been expanded to segmental power analysis to study mechanical energy flow (Fregly and Zajac, 1996; Neptune et al., 2001, 2004; Seigel Corresponding author. Tel.: +1 860 380 0164; fax: +1 860 545 8842. E-mail addresses: Mvpatel@ccmckids.org (M. Patel), mctalaty@einstein.edu (M. Talaty), Sounpuu@ccmckids.org (S. Õunpuu). 1 Tel.: +1 215 663 6682. 2 Tel.: +1 860 545 8710. 0021-9290/$ - see front matter r 2007 Elsevier Ltd. All rights reserved. doi:10.1016/j.jbiomech.2007.06.031 et al., 2004). Varying degrees of complexity exist in the models used in the above studies, depending on the objective of the study and feasibility of incorporating the complexity in the model. The extent to which biomechanical model assumptions affect the interpretations of muscle function is not fully clear. Some researchers have used three-dimensional models in their studies (Anderson and Pandy, 2003; Arnold et al., 2005; Kepple et al., 1997; Liu et al., 2006; Riley et al., 2001b; Siegel et al., 2001), whereas others have used planar models (Fregly and Zajac, 1996; Higginson et al., 2006; Neptune et al., 2001, 2004). The types of foot–floor interactions used in the above studies also vary. All the studies except the one using Anderson’s model (Anderson and Pandy, 2003) have modeled everything above the hip joints as a single segment. Some of the results from the above studies have differed depending on the model used (Anderson and Pandy, 2003; Kepple et al., 1997; Neptune ARTICLE IN PRESS 3564 M. Patel et al. / Journal of Biomechanics 40 (2007) 3563–3569 et al., 2004; Riley et al., 2001b). Though some of the recent studies have started using Anderson’s 10-segment model, where the pelvis and trunk are modeled as separate segments, to our knowledge there exists no literature that documents the impact on the IAA results of modeling the pelvis and trunk segments as two separate segments versus as a single segment. The effect of assuming the pelvis and trunk as a single segment on the IAA results has not been reported, despite clear relative motion between the trunk and pelvis during walking. (Riley et al., 2001c). The combined pelvis–trunk segment assumption may especially not hold for pathologic gait, as observed in patients with myelomeningocele who exhibit excessive trunk motion relative to the pelvis (Ounpuu et al., 2000). Before performing IAA on the gait of the patients with myelomeningocele, we aimed to clarify how upper body modeling assumptions affected the interpretation of the role of muscles during walking. We studied segmental power, knee and hip extension acceleration and support of the body (total body center of mass vertical acceleration) produced by all the joint moments to infer muscle function. 2. Methods Motion data from five healthy subjects (65710 kg; 1.6570.11 m; 2371 year; two males and three females) with no known pathology were collected at Connecticut Children’s Medical Center during normal walking. Informed consent was obtained from the subjects participating in the data collection. The subjects were instrumented with 16 reflective markers on the lower extremities, pelvis and trunk. The subjects were asked to walk barefoot at a self-selected pace along a pathway on which three force plates (AMTI, Model # OR6, Watertown, MA) were mounted in series. Data were collected from each subject using the 512 Vicon Motion Systems (Vicon Motion Systems, Lake Forest, CA). The testing protocol for these subjects has been described elsewhere (Davis et al., 1991). A single trial for each subject was chosen, in which three consecutive steps contacted the three force plates. Joint moments and powers were computed using standard inverse dynamics approach (Davis et al., 1991) using anthropometric data from Dumas et al. (2007). The load summation for the most proximal segment computed using inverse dynamics might not equal the product of mass/inertia and measured acceleration. Skin movement artifact, error in the estimated segment anthropometry, error in locating the joint center location, etc. are possible reasons (Plamondon et al., 1996). Loads that make the equations of motion of the most proximal segment dynamically consistent were calculated for the upper body and trunk in the 7- and 8-segment models, respectively, and considered when calculating the model output sum. The three-dimensional biomechanical model was created in ADAMS (MSC Software, Ann Arbor, MI) for each subject. The basis of the biomechanical model used in this work was developed based on the work done by M. Patel during the research component of her masters work that was carried out at the Gait and Motion Analysis Laboratory at Moss Rehab. Two different three-dimensional (3D) models were created for each subject, a 7-segment (7 SM), 22 degrees-of-freedom (DOF) [Upper body segment (combined head, arms, torso and pelvis) and bilateral thigh, shank, feet] model and an 8-segment (8 SM), 25 DOF [trunk (combined head, arms and torso), pelvis, and bilateral thigh, shank, feet] model. For both the models, the ankle was modeled as a universal joint, and the hip and knee as spherical joints. The knee was modeled as a spherical joint for upcoming analysis of patients with myelomeningocele who may exhibit some motion in all three planes at the knee (Ounpuu et al., 2000). For the 8 SM the lumbar joint (joint between the pelvis and trunk) was modeled as a spherical joint. The motion and orientation of the upper body segment in the 7 SM was tracked using the pelvis markers but the anthropometry of the combined pelvis and trunk was used, similar to Kepple et al. (1997). Foot–floor interaction was modeled as a revolute joint at the measured center of pressure (CoP) obtained from the forceplates during the entire stance phase. The segmental power, knee and hip extension acceleration and support of the whole body produced by the joint moments and gravity during the stance phase were analyzed for both the 7 and 8 SMs. The total body center of mass (COG) vertical acceleration was assumed to be representative of the support of the body. Input to the model was body configuration (position and orientation), location of the center-ofpressure, computed joint moments and the offset loads. The model was reconfigured for each analysis frame using motion capture data. Each joint moment’s contribution to segmental power, COG vertical acceleration and joint angular accelerations was calculated for each frame by the solver module in ADAMS, using methods similar to ones previously described (Kepple et al., 1997; Seigel et al., 2004). For the 8 SM, the lumbar joint moments were also applied. For ease of comparison of the upper body power for the two models the power of the pelvis and trunk segments were summed to estimate the contribution of each joint moment to upper body power for the 8 SM. This was then compared to the upper body power of the 7 SM. The segmental power results were normalized to body mass to compare the data across subjects. The model output sum was compared with the results computed from the motion capture data to validate the model (Seigel et al., 2004). The term ‘reconstruction’ was used to describe the computed model output sum. The contribution of the coriolis and centripetal forces during normal walking was assumed to be negligible. 3. Results Induced acceleration and segmental power using the two different models were computed for the five subjects. The results for a representative subject are graphed in Figs. 1–3, 5 and 7; and the averages of the five subjects are graphed in Figs. 4, 6 and 8. The inverse dynamics solution was used to verify the model (Kepple et al., 1997). The data computed from the model output were reasonable compared to the motion data (Fig. 1). Similar reconstructions were obtained for the two models (Fig. 1). 3.1. Segmental power Key differences in hip contribution to upper body power between the 7 and 8 SMs were noted (Fig. 2). The stance hip moment added energy to the upper body in early stance and removed energy from the upper body during late stance for the 7 SM, whereas the opposite was observed for the 8 SM. The magnitude of the power delivered or removed from the thigh, upper body and contra lateral leg by the hip moment was considerably greater for the 8 SM. The ankle moment removed less power from the upper body segment and for a shorter duration during early stance for the 8 SM. 3.2. Support of the body The contribution of the joint moments to the COG vertical acceleration for the 7 and 8 SMs are shown in Figs. 3 ARTICLE IN PRESS M. Patel et al. / Journal of Biomechanics 40 (2007) 3563–3569 3565 Fig. 1. Net stance leg joint power, body center of gravity vertical acceleration, hip and knee extension acceleration as computed from the model data and the motion data for a single representative subject during the stance phase of gait. The solid line is as calculated/measured from the motion data, the dotted line is the 7-segment model reconstruction and the dash–dot line is the 8-segment model reconstruction. Fig. 2. Segmental power due to the stance leg hip, knee and ankle moment during the stance phase of gait for a representative subject. The abbreviations used in the legend are explained here. ITh: ipsilateral thigh (dash-dot), ISk: ipsilateral shank (dash-dash); IFt: ipsilateral foot (solid); Contra: contra lateral leg (dot); and UB: upper body (dotted line). Note: the upper body power for the 8-segment model was the sum of contributions to the pelvis and trunk segments. ARTICLE IN PRESS 3566 M. Patel et al. / Journal of Biomechanics 40 (2007) 3563–3569 Fig. 3. Contribution of the joint moments and gravity to body center of gravity vertical acceleration (support) during the stance phase of gait for a representative subject. The abbreviations used in the legends are explained here. SIA: sagittal ipsilateral ankle moment, CIA: coronal ipsilateral ankle moment, SIK: sagittal ipsilateral knee moment; CIK: coronal ipsilateral knee moment; SIH: sagittal ipsilateral hip moment; CIH: coronal ipsilateral hip moment; Contr: contra lateral leg moments; Grav: passive resistance of bones and joints to gravity; SL: sagittal lumbar moment; CL: coronal lumbar moment. and 4. For the 8 SM, the hip extensors during early stance and hip abductors during mid-stance provided greater support to the body than that observed in the 7 SM (Fig. 4). Also lumbar sagittal and coronal moments provided support to the body during the second half of stance for the 8 SM (Fig. 3). 3.3. Knee and hip extension acceleration The differences observed in the distribution of the contributions to the knee angular acceleration between the 7 and 8 SMs were small (Figs. 5 and 6), whereas considerable differences were observed in the magnitude of the contributions for the hip angular acceleration (Figs. 7 and 8). The 7 SM showed prolonged contribution of ankle sagittal moment to flexion acceleration of the knee in early stance. The contributions of the ipsilateral hip extensors and abductors to hip and knee acceleration during stance were greater for the 8 SM than for the 7 SM (Figs. 5–8). For the 8 SM, hip abductors caused knee flexion acceleration during first half of stance and hip extension acceleration for the second half of stance. However, for the 7 SM they had minimal contribution for those phases (Fig. 6). 4. Discussion We aimed to evaluate the effect of modeling the pelvis and trunk as a single versus separate segments on acceleration Fig. 4. Mean contribution across all five subjects of the major contributors to the body center of gravity vertical acceleration for the 7- and 8-segment models during early stance (up to 20% of the gait cycle) unless specified. analysis calculations during the stance phase of normal walking. Specifically, we evaluated segmental power, support of the body, and knee and hip extension acceleration produced by the joint moments for the two models. Differences were observed predominantly in the contribution of the ipsilateral hip sagittal and coronal moments to support, upper body and ipsilateral thigh power, and hip and knee acceleration. The polarities of the contributions, which are indicative of the role of the muscle or joint ARTICLE IN PRESS M. Patel et al. / Journal of Biomechanics 40 (2007) 3563–3569 3567 Fig. 5. Contribution of the joint moments and gravity to knee extension acceleration during the stance phase of gait for a representative subject. The abbreviations used in the legends are explained here. SIA: sagittal ipsilateral ankle moment, CIA: coronal ipsilateral ankle moment, SIK: sagittal ipsilateral knee moment; CIK: coronal ipsilateral knee moment; SIH: sagittal ipsilateral hip moment; CIH: coronal ipsilateral hip moment; Contr: contra lateral leg moments; Grav: passive resistance of bones and joints to gravity. Fig. 6. Mean contribution across all five subjects of the joint moments to knee extension acceleration during early stance (5–25% of the gait cycle). moment, were the main focus. Differences in the magnitudes of contributions were also noted, but their importance at this time is less clear. It is important to consider some modeling assumptions used in the present study before discussing the results. First, the same CoP location that calculated from actual force plate data was used for both the 7 and 8 SMs. Changing the upper body orientation, as done in the 7 SM, would likely change the CoP, which would affect the computed lower extremity kinetics (Westwell et al., 2005). We did not recompute the CoP, as the focus of the present work was to illustrate the impact of modeling the trunk and pelvis as separate versus a single segment on the IAA results using biomechanical models similar to those used in the literature (Seigel et al., 2004; Anderson and Pandy, 2003). Secondly, the foot–floor interaction was assumed as a rigid contact of a revolute joint at the CoP during the entire stance phase. This assumption may not hold for the entire stance phase, which may be responsible for some of the differences between the reconstruction and motion data as seen in Fig. 1. Lastly, a spherical knee joint was used in the current study instead of the most commonly used hinge joint (Seigel et al., 2004). The impact of using a spherical joint instead of a hinge joint on the differences observed between the 7- and 8-segment models was tested. It was found that the IAA results were not sensitive to the two extra knee degrees of freedom. The differences between the 7- and 8-segment models remained about the same, with small changes in the magnitude of the contributions observed for some of the variables for both the 7- and 8-segment models. Despite the similarity of model output reconstructions (Fig. 1), the distributions of the joint moments, particularly those of the hip to the segmental power or accelerations were appreciably different (see Figs. 2–8). Chen (2006) had found that increasing the degrees of freedom of the model he used resulted in greater power redistribution attributed to the joint moments. For our study, this was not true for all the joint moments. Extra lumbar degrees of freedom increased the contribution of the ipsilateral hip moments to ARTICLE IN PRESS 3568 M. Patel et al. / Journal of Biomechanics 40 (2007) 3563–3569 Fig. 7. Contribution of the joint moments and gravity to hip extension acceleration during the stance phase of gait for a representative subject. The abbreviations used in the legends are explained here. SIA: sagittal ipsilateral ankle moment, CIA: coronal ipsilateral ankle moment, SIK: sagittal ipsilateral knee moment; CIK: coronal ipsilateral knee moment; SIH: sagittal ipsilateral hip moment; CIH: coronal ipsilateral hip moment; Contr: contra lateral leg moments; Grav: passive resistance of bones and joints to gravity. Fig. 8. Mean contribution across all five subjects of the joint moments to hip extension acceleration during early stance (5–25% of the gait cycle) unless specified. the segmental power, but the contribution of the ipsilateral knee moment to the segmental power did not change much and the contribution of the ipsilateral ankle moment to the segmental power decreased (Fig. 2). The joint reaction forces more proximal to the body and on the contra-lateral side were the most affected when modeling a separate trunk and pelvis segment. The main reason for the observed differences between the two models was the extra degrees of freedom allowed for the trunk in the 8 SM and not differences in trunk orientation or its center of mass location. Because the 7 SM used pelvis markers to position and orient (POSE) the trunk, the trunk center of mass location and inertia orientation were different between the 7 and 8 SMs—in addition to the extra degrees of freedom allowed in the latter. Either could have caused the significant differences we observed. To clarify further, a test was performed where the trunk POSE was forced to be equal to the pelvis POSE for the 8 SM, while keeping the joint between the pelvis and trunk as a spherical joint. This was named the 8-segment dummy. The 8-segment dummy energy flow results were similar to the 8 SM with subtle differences in the magnitude of the contributions, indicating that most of the differences between the 7 and 8 SMs were due to the extra degrees of freedom allowed for the trunk in the 8 SM. The segmental power attributed to the joint moments for the 7 SM in our study was generally consistent with that reported in the literature (Neptune et al., 2004; Seigel et al., 2004), using 7 SMs. Differences in the segmental power between the 7 and 8 SMs were predominantly observed in the contribution of the hip moments. The hip flexor contribution to the upper body power changed on modeling the pelvis and trunk as separate segments. Neptune et al. (2004) and Seigel et al. (2004) reported that the hip flexors redistributed the power from the upper body to the ipsilateral leg during pre-swing, as we did in the 7 SM. But for the 8 SM, the opposite was observed. Transfer of power into the upper body, during pre-swing, is in ARTICLE IN PRESS M. Patel et al. / Journal of Biomechanics 40 (2007) 3563–3569 agreement with suggestions of others that the hip flexors assist forward acceleration (Sadeghi et al., 2001). The role of the ipsilateral hip extensors during early stance and ipsilateral hip abductors during mid-stance to support appears to be underestimated if the pelvis and trunk segments are fused as a single segment (Figs. 3 and 4). Unlike the 7 SM, for the 8 SM the hip abductors had significant contribution to support during mid-stance (Fig. 3) for most of the subjects; similar to that observed by others (Anderson and Pandy, 2003; Liu et al., 2006), who modeled the pelvis and trunk as separate segments. Substantially greater hip abductor contribution was observed for two of the five subjects using the 8 SM, who had their trunk leaning more towards their contra lateral side. For the 7 SM, the support provided by the hip abductors throughout stance remained small for all the five subjects; similar to that observed by Kepple et al. (1997), who used a 7 SM. It appears that to quantify the role of the hip abductors to support, it is essential to model the pelvis and trunk as separate segments and allow the three rotational degrees of freedom between the two segments. The overall trend of the major contributions other than by the ipsilateral hip abductors to hip and knee accelerations did not differ much for the two models. However, some discrepancies were observed in the magnitude of the contributions. Hip and knee extensors accelerated the hip and knee joints into extension during the first half of stance for both the models, similar to that observed by Neptune et al. (2004) and Arnold et al. (2005), but greater hip extensor contribution was observed for the 8 SM. Unlike the 7 SM, the hip abductors in the 8 SM were found to cause knee flexion acceleration during first half of single support phase (Fig. 6), and hip extension acceleration during the second half of stance (Fig. 8), similar to Arnold et al. (2005), who modeled the pelvis and trunk as separate segments. This study documents the segmental power flow due to joint moments using a model which models the pelvis and trunk as separate segments, which to date does not exist in the literature. From the current study it was found that modeling the pelvis and trunk as separate segments impacts the interpretation of the role of the joint moment during normal walking. Differences in the trends of the contribution of the stance hip moment to the ipsilateral thigh and upper body power, support, hip and knee sagittal acceleration were observed between the 7 and 8 SMs. This changes our understanding of the possible role of the hip moment during stance phase. In the future, we plan to use the 8 SM to analyze the mechanisms used to walk by patients with myelomeningocele and study the role played by the lumbar moments to substitute for the lower extremity muscle weaknesses. Conflict of interest None of the authors of this manuscript (Mausam Patel, Dr. Mukul Talaty and Sylvia Õunpuu) have any financial or personal relationships with other people or 3569 organizations that could inappropriately influence (bias) this work. Also the material in the manuscript has not been and will not be submitted for publication elsewhere except as an abstract. Acknowledgments Some of the computational aspects of this work were performed at MossRehab. The authors gratefully acknowledge the comments on an earlier draft of this manuscript by Alberto Esquenazi, MD. References Anderson, F., Pandy, M., 2003. Individual muscle contributions to support in normal walking. Gait & Posture 17, 159–169. Arnold, A., Anderson, F., Pandy, M., Delp, S., 2005. Muscular contributions to hip and knee extension during single limb stance phase of normal gait: a framework for investigating the causes of crouch gait. Journal of Biomechanics 38, 2181–2189. Chen, G., 2006. Induced acceleration contributions to locomotor dynamics are not physically well defined. Gait & Posture 23, 37–44. Davis, R., Ounpuu, S., Tyburski, D., Gage, J., 1991. A gait analysis data collection and reduction technique. Human Movement Science 10, 575–587. 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