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Claudication is the most common manifestation of peripheral arterial disease, producing significant ambulatory compromise. The purpose of our study was to evaluate patients with bilateral lower limb claudication and characterize their gait abnormality based on advanced biomechanical analysis using joint torques and powers.
Twenty patients with bilateral claudication (ten with isolated aortoiliac disease and ten with combined aortoiliac and femoropopliteal disease) and sixteen matched controls ambulated on a walkway while three dimensional biomechanical data were collected. Patients walked before and after onset of claudication pain. Joint torques and powers at early-, mid-, and late-stance for the hip, knee and ankle joints were calculated for claudicating patients before and after the onset of claudication pain, and were compared to control subjects.
Claudicating patients exhibited significantly reduced hip and knee power at early-stance (weight acceptance phase) due to decreased torques produced by the hip and knee extensors. In mid-stance (single limb support phase), patients had significantly reduced knee and hip power due to the decreased torques produced by the knee extensors and the hip flexors. In late-stance (propulsion phase), reduced propulsion was noted with significant reduction in ankle plantar flexor torques and power. These differences were present before and after the onset of pain with certain parameters worsening in association with pain.
The gait of claudication is characterized by failure of specific and identifiable muscle groups needed to perform normal walking (weight acceptance, single limb support and propulsion). Parameters of gait are abnormal with the first steps taken, in the absence of pain, and certain of these parameters worsen after the onset of claudication pain.
Intermittent claudication is the most common clinical manifestation of peripheral arterial disease (PAD) presenting as exercise induced leg muscle pain and gait dysfunction(1). Claudication and its associated ambulatory impairment produce impaired quality of life(2), physical dependence(3) and poor health outcomes(4). Previous work suggests that PAD patients walk slower with decreased cadence, increased stance time, shorter stride length and a narrower step width as compared with healthy controls(5, 6). However, these changes alone are unable to describe in sufficient detail the locomotor impairments of claudicating patients and aid in our understanding of its underlying pathophysiology.
A more detailed quantitative evaluation of gait can be obtained using advanced biomechanical analysis including joint torques and powers(7, 8). Although muscles produce linear forces, motions at joints are all rotary. The rotary torque is a measure of the tendency of a force to rotate the limb around a joint and is calculated as the product of the muscle force and the distance from the joint center that the force is being applied. The net muscle torque does not represent any one particular muscle but rather describes the net activity of all the muscles acting across a joint. Joint power can be defined as the rate of work produced by muscles contracting to move a joint and is determined as the product of the net torque (moment) of the muscles acting across a joint and the resulting angular velocity of the joint. Joint powers have been used extensively to identify the mechanisms responsible for pathological gait (in populations such as elderly, patients with knee and hip arthritis and arthroplasty and anterior cruciate ligament reconstruction) and to assess and guide successful rehabilitation strategies(8–10). Similar insights can be gained from patients with PAD utilizing this approach.
Previous studies of claudicating PAD patients from our laboratory utilizing basic biomechanical analysis(7, 11–13) suggested a potential weakness of the posterior compartment muscles of the hip and calf as key components of PAD gait impairment. The purpose of the current study is to utilize joint torques and powers to isolate and identify the individual muscle compartments responsible for the gait impairment of claudicating patients.
The gait cycle (from heel touchdown to heel touchdown) consists of a stance and a swing period. The stance phase is the most important segment of the gait cycle because the ambulating limb accepts, supports and propels forward the weight of the body. Furthermore, it is the only portion of the gait cycle that can be accurately evaluated for joint moments and powers. The stance segment can be divided in three distinct phases including the weight acceptance, the single limb support and the propulsion phases (Figure 1).
Twenty patients diagnosed with moderate arterial occlusive disease and bilateral claudication (Fontaine stage II and Rutherford Grade 1 categories 2 and 3)(14) were recruited from the vascular surgery clinics of the VA Nebraska and Western Iowa and University of Nebraska Medical Centers and signed an inform consent prior their participation to the study that was approved by the institutional review boards of each respective institution. All twenty PAD patients had aortoiliac occlusive disease with ten having isolated aortoiliac disease and the other ten having aortoiliac and femoropopliteal disease. The diagnosis of level of disease was made with the use of computerized tomographic angiography and ultrasonographic evaluation in all patients. In addition, 16 gender, age, body-mass, height matched healthy controls were recruited. The controls were matched to PAD patients by group/frequency matching. Patients and controls were screened and evaluated by two board certified vascular surgeons. Those PAD patients with ambulation limiting cardiac, pulmonary, neuromuscular or musculoskeletal disease or those who experienced pain or discomfort during walking for any reason other than claudication (i.e. arthritis, low back pain, musculoskeletal problems, and neuropathy) were excluded. Patient evaluation included resting ankle brachial index (ABI; a measurement below 0.90 was present in all subjects with claudication), detailed history, physical exam and direct assessment/observation of the patient’s walking impairment. A vascular surgeon observed the patient walking and recorded all symptoms and signs affecting ambulation to insure limitation was secondary to claudication pain. Control subjects had a resting ABI greater than 0.90 and no subjective or objective ambulatory dysfunction. Controls were screened in a similar fashion as PAD patients and were excluded for the same ambulation limiting problems or if pain was experienced during walking. Informed consent was obtained from all subjects prior to data collection according to the guidelines of the Institutional Review Boards of the two medical centers. The gait of all recruited participants was tested in the Biomechanics Laboratory of the University of Nebraska at Omaha.
Prior to data collection, reflective markers were placed at specific anatomical locations of each subject’s lower limb utilizing the modified Helen Hayes marker set(12). Each subject was directed to walk using their self-selected pace over a ten meters pathway, while three-dimensional marker trajectories (kinematics) and ground reaction forces (kinetics) were simultaneously collected. The marker trajectories were captured with an eight high-speed real-time camera system (EvaRT 5.0, Motion Analysis Corp., Santa Rosa, CA) sampling at 60Hz. The ground reaction force data were acquired with a Kistler force platform sampling at 600 Hz.
Each patient was tested before (“Pain Free” condition) and after (“Pain” condition) the onset of claudication pain. During “Pain Free” testing, mandatory rest occurred between the walking trials to insure that all trials were in a “Pain Free” condition. Once patients completed all “Pain Free” trials, “Pain” trials were performed. In order to accomplish this, each patient was asked to walk on an inclined treadmill with 10% grade at a speed of 0.67m/s(15) until claudication pain was established. The patients were then immediately removed from the treadmill and returned to the collection walk-way to acquire the data for the “Pain” condition without the mandatory resting periods. Controls completed only the “Pain Free” condition trials. A total of five walking trials were collected from each leg of the subjects for each condition.
Joint kinetics and kinematics were calculated for the sagittal plane during the stance phase of walking. An inverse dynamic solution was performed to calculate joint muscle torques and powers from the joint kinetics and kinematics(8). Joint muscle power (Pj) is calculated as the product of the net torque of force at a joint (Tj) and the relative joint angular velocity (ωj) or Pj=Tj * ωj (Joules/sec or Watts). Power combines both kinetic (forces) and kinematic (angles and velocities) information and can be expressed positively or negatively. Positive power indicates that energy is being generated and negative power that energy is being absorbed by the muscle group under study. Thus, positive joint muscle power is associated with concentric muscular contractions, while negative power is associated with eccentric muscular contractions(8). Joint torques and joint muscle powers were normalized with respect to the subject’s body mass and expressed as a percentage of the stance phase. The peak values for extensor and flexor torques were identified for the ankle, knee and hip joints(8). The variables identified for the ankle were the ankle dorsiflexor torque (ADT) in early stance and the ankle plantar flexor torque (APT) in late stance; for the knee were the knee extensor torque (KET) in early stance and the knee flexor torque (KFT) in late stance; for the hip were the hip extensor torque (HET) in early stance and the hip flexor torque (HFT) in late stance. In addition, the peak values for power absorption (eccentric contraction) and generation (concentric contraction) were identified for the ankle, knee and hip joints(8). The power variables indentified for the ankle were the power absorption in early (A1) and mid (A2) stance and the power generation in late stance (A3); for the knee were the power absorption in early stance (K1), the power generation in the early part of mid-stance (K2), and the knee power absorption in late stance (K3); for the hip joint were the power generation in early stance (H1), the power absorption in mid-stance (H2) and the power generation in late stance (H3)(8). Custom made Matlab (Matlab 2008b, Mathworks, Inc., Concord, MA) software was used to calculate the joint torques and powers.
Data were summarized by group means and standard deviations of the peak joint torques and powers. These were calculated for each testing condition (Controls, and “Pain Free” and “Pain” conditions for claudicating patients) by first averaging the limb measurements for each subject, and then averaging the obtained averages over all subjects in each group (i.e. for 20 subjects in the claudicating group and, respectively, for 16 subjects in the Control group). When modeling the effects of predictors on response variables, repeated measures models were developed, where the unit of analysis was limb (not person/subject), and the two limbs of a subject were “nested” in him/her, to account for possible correlation between the two limbs in a subject. Three separate comparisons were performed: (i) Controls vs. claudicating patients in “Pain Free” condition; (ii) Controls vs. claudicating patients in “Pain” condition; and (iii) claudicating patients in “Pain Free” condition vs. claudicating patients in “Pain” condition. For comparisons (i) and (ii), linear models with repeated measures on variable Leg (left vs. right) were used, whereas for comparison (iii), linear models with doubly-repeated measures on variables Leg (left vs. right) and Pain status (before pain vs. after pain) were used. In addition to the main effects of Group and Leg for (i) and (ii), or Pain status and Leg for (iii), first order interaction terms were also included in the models. Additionally, in all models we have controlled for variable ABI and, particularly for (iii), we also controlled for variable Level-of-disease (aortoilliac only vs. aortoilliac plus femoropopliteal). The concrete P-values were always reported. The level of significance was set to 0.05. The procedure MIXED in the statistical package SAS 9.2 (SAS Inst, Cary, NC, USA) was used to develop all linear models.
The baseline clinical characteristics of patients and healthy controls are presented in Table 1. No significant differences were found between groups regarding age, body mass and body height. When compared to controls, PAD patients had significantly decreased gait velocity, stride length and step length and increased step width (Table 2). The differences for these parameters remained significant when the patients walked experiencing muscle pain in the “Pain” condition (Table 2). Comparing temporal and spatial gait measurements before and after onset of claudication, there was a significant decrease in gait velocity (Table 2).
In comparison to the controls (Tables 3, ,4),4), patients in the “Pain Free” condition generated significantly decreased torque by the hip extensors (HET) and by the knee extensors (KET) which translated to significantly decreased power at the hip (H1, reduced concentric contraction of the hip extensors) and at the knee (K1, reduced eccentric contraction of the knee extensors). Decreased (although not significant) power absorption was produced at the ankle (A1, reduced eccentric contraction of the ankle dorsiflexors). In the “Pain” condition, the results for the the knee extensor torque (KET), the concentric contraction of the hip extensors (H1) and the eccentric contraction of the ankle dorsiflexors (A1) remained significantly different compared to healthy controls (Table 3).
In comparison to the controls (Table 3,,4),4), patients in both “Pain Free” and “Pain” conditions generated significantly decreased torque by the knee extensors (KET) and the hip flexors (HFT) which then translated to significantly decreased knee joint power generation in the early part of mid-stance phase (K2, reduced concentric contraction of the knee extensors) and decreased hip power absorption in the late part of mid-stance (H2, reduced eccentric contraction of the hip flexors).
In comparison to the controls (Table 3,,4),4), patients in both “Pain Free” and “Pain” conditions generated significantly decreased torque by the ankle plantarflexors (APT) and the hip flexors (HFT) which translated to significantly less power at the ankle (A3, reduced concentric contraction of the ankle plantarflexors) and the hip (H3, reduced concentric contraction of the hip flexors in the pain condition). Additionally, patients in both conditions absorbed significantly less power at the level of the knee flexors (K3, reduced eccentric contraction of the knee flexors). Once in the “Pain” condition, patients generated significantly decreased ankle plantarflexor torque (APT) compared to the “Pain Free” condition. This significant decrease in torque (reduced concentric contraction of the ankle plantar flexors) translated into further decrease of power generation at the ankle (A3).
The purpose of this study was to utilize joint torques and powers in order to characterize and provide an in depth understanding of the gait impairment of claudicating patients. Joint torques and powers were measured while patients walked both with and without claudication pain and were compared to those of gender-, height-, mass-, and age-matched healthy controls. Our results from the temporal and spatial gait parameters demonstrate that the character of the PAD gait is overall “sluggish and tired”. Patients with claudication have decreased gait velocity, decreased stride and step length, and increased step width. These findings are in agreement with previous studies and unequivocally document the abnormal temporal and spatial gait parameters of claudicating PAD patients(5,6). Utilizing advanced biomechanical analysis in the form of joint torques and powers, we were able to isolate and describe the specific muscle group impairments that operate to produce the gait deficit in claudicating patients. Our data demonstrate a decreased ability of the knee and hip extensors to control weight acceptance and ankle dorsiflexors to eccentrically control the lowering of the foot to the ground after heel strike in early stance. In mid stance we found a decreased ability of the knee extensors to concentrically extend the knee and of the hip flexors to eccentrically control the movement of the pelvis. Finally, in late stance we demonstrated a decreased ability of the ankle plantarflexors and of the hip flexors to concentrically propel the body forward and of the knee flexors to eccentrically control knee hyperextension as the trunk is accelerated forward.
Trunk support in early stance is provided by the hip extensors concentrically contracting to extend the hip (H1), the knee extensors eccentrically contracting to allow the knee to flex (K1) and the ankle dorsiflexors which eccentrically control the movement of the foot towards full contact with the ground (A1)(16). Our findings in PAD patients demonstrate that all three muscle groups involved in the weight acceptance phase produce less power than in controls. Specifically PAD patients have decreased power generation by the hip extensors (Gluteus muscles, H1) and the knee extensors (quadriceps) in early stance (K1) indicating decreased ability to support the body weight. These power results are supported by our joint torque findings which showed significantly decreased torque development by the hip (HET) and knee (KET) extensors(12). The demonstrated weakness of the hip and knee extensors in early stance is results in diminished ability for weight acceptance and control of forward momentum when a claudicating patient walks. The demonstrated weakness of the ankle dorsiflexors in early stance is in agreement with findings our group has previously published showing that PAD patients have a “foot drop” upon heel strike(11).
In the early part of the mid-stance phase the knee extensors concentrically contract to extend the knee joint and in the late part the hip flexors contract eccentrically to control the movement of the pelvis during single leg support. To maintain the energy required for walking(16, 17), it is necessary to “straighten” the leg and stabilize the pelvis to maximize the ability to generate potential energy at the highest point of the body’s center of mass during the gait cycle. Our work shows that in PAD patients both muscle groups (knee extensors, K2 and hip flexors, H2) involved in single limb support produce less power than in controls.
In late stance the body is propelled forward mainly by the action of the ankle plantarflexors. Functionally, these muscles contract concentrically and accelerate the leg and the trunk forward to initiate swing, while decelerating the downward motion of the trunk (i.e., providing forward progression and support)(18). Our results in the PAD patients, demonstrate that power generation via concentric contractions of the ankle plantarflexor muscles in late stance (A3) is decreased in the “Pain Free” condition and worsens in the “Pain” condition. This hypothesis is supported by previous findings from our and other laboratories demonstrating that PAD patients have significantly decreased ankle plantarflexor torques(7, 12) and strength(13). Our additional findings of decreased knee and hip flexor powers further demonstrate the failure of the PAD limb to appropriately utilize these muscle groups to assist the ankle plantar flexors to accelerate the trunk forward.
Our findings for the temporal and spatial gait parameters provide definitive evidence of abnormal temporal and spatial parameters in patients with PAD and confirm the generally accepted thought that claudicating patients have an abnormal gait(5), which leads to increased energy cost and earlier fatigue(19). More importantly, our advanced analysis with joint torques and powers provides a more detailed delineation of the gait disturbance than that previously documented by our group and others using spatial (joint angles)(11, 12, 20, 21) and temporal (velocity, cadence and step/stride characteristics) parameters(5, 6, 20).
When comparing the abnormalities of joint torques and powers in PAD to other conditions, our values are in line with those of healthy elderly and elderly patients with osteoarthritis(22–25). In contrast to these two groups however, the gait biomechanics of PAD patients appear to be significantly worse than healthy elderly subjects and those patients with severe arthritis. Specifically our data demonstrate that from the first few steps they take and before they experience any muscle pain claudicating patients walk with 26% decrease (versus controls) of their ankle plantarflexor power compared to a 13% for elderly osteoarthritis patients(26). Arthritis patients compensate for the 13% decrease in the power of their ankle plantarflexors by increasing the power of their knee and hip extensors (by 13% and 28% respectively)(23). In marked contrast to the arthritic patients, claudicating patients demonstrate a drop of compensatory power in these muscle groups by 39 and 14% respectively(23). It is clear from a biomechanical standpoint that claudication produces a considerably worse functional limitation than osteoarthritis. Our data support the findings that claudicating patients typically gather around the extreme low end of the physical activity spectrum(27) and experience a severe decline in all domains of physical function(6, 28).
The present data demonstrate significant abnormalities in the gait of claudicating patients that are present at the initiation of ambulation and prior to onset of pain. Work of other investigators evaluating PAD patients has shown that tissue oxygen levels and blood flow, in the leg, during the first few steps of walking, are very similar to (and sometimes even higher than) those of controls(29–31). These results together with our finding of gait impairments present within the first few steps taken by the patient suggest that a mechanism other than blood flow is responsible for a significant portion of the limb dysfunction of claudication. This conclusion is supported by the work of McDermott et al. who measured muscle strength in PAD patients, using a 5-second maximal isometric strength assessment(32). This test, which is clearly blood flow–independent, demonstrated that muscle strength in PAD patients is significantly reduced compared to controls, again, indicating that blood flow is not the only determinant of the dysfunction of claudicating muscle. These baseline biomechanical impairments likely reflect a muscle metabolic myopathy and an axonal polyneuropathy in the lower extremities of patients with PAD(33–35). Specifically, a number of reports have documented a metabolic myopathy in the PAD muscle that is related to defective mitochondrial bioenergetics, oxidative damage and inflammation in the skeletal muscle(34, 36–39). Furthermore, there is accumulating evidence suggesting that chronic ischemia in patients with PAD results in a consistent pattern of electrodiagnostic abnormalities indicating axonal nerve loss(35, 37). Our data further demonstrate that certain baseline impairments worsen after the onset of claudication pain. This reflects exercise induced ischemia and increased workload, restricting lower extremity bioenergetics and producing muscle pain. Because of this ischemia-induced neuromyopathy patients with claudication become severely debilitated and adopt a sedentary lifestyle characterized by limited use of the myopathic limb, which may exacerbate the neuromyopathy. Figure 2 illustrates a proposed pathway linking these basic pathophysiologic mechanisms (exercise associated ischemia-reperfusion, myopathy, neuropathy) with the specific biomechanic deficits identified in this work. The role played by each one of these mechanisms and the way they are related to the clinical biomechanical findings of leg dysfunction should be the focus of intense future investigation and may hold the key to understanding PAD pathophysiology.
On a clinical level, only recently have studies with large sample sizes (N= 700–2000) been able to demonstrate the long held assumption that claudicating patients have significantly reduced muscle strength(32, 40). Our study utilizing advanced biomechanical analysis has allowed us to confirm these large scale studies with a limited number of patients and to implicate the involved muscle groups. Advanced biomechanical techniques thus provide a new avenue for evaluation, treatment, and rehabilitation of the PAD patient.
Biomechanical analysis using joint torques and powers demonstrates significant abnormalities in the gait of claudicating patients with bilateral PAD. These abnormalities are present at the onset of ambulation and worsen with the pain of claudication. Our work points to a failure of major muscle groups to optimally perform the sequence of functions (weight acceptance, transfer and propulsion) that characterize normal gait. In patients with occlusive disease affecting the proximal arterial tree the muscle groups most affected by the chronic ischemia are the hip flexors and extensors, the knee extensors and the ankle plantar flexors. These findings introduce new insights into the pathophysiology of claudicating gait. In the future these advanced biomechanical techniques will provide for detailed objective and quantitative evaluation of the gait deficit of the claudicating patient, allowing for evaluation of new treatment and rehabilitation strategies.
Support for this work was provided by funds from the Alexander S. Onassis Public Benefit Foundation to PK, the American Geriatrics Society’s Hartford Foundation Dennis W. Jahnigen Award to JMJ, the Nebraska Research Initiative to NS, the Lifeline Programs of the American Vascular Association to IIP and the NIH to NS (K25HD047194) and IIP (K08HL079967).
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