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Biomechanics of uphill walking using custom ankle-foot orthoses of three different stiffnesses§ Derek J. Haight, Elizabeth Russell Esposito *, Jason M. Wilken Center for the Intrepid, Department of Orthopaedics and Rehabilitation, Brooke Army Medical Center, Ft. Sam Houston, TX 78234, United States
A R T I C L E I N F O
A B S T R A C T
Article history: Received 14 August 2014 Received in revised form 12 December 2014 Accepted 3 January 2015
Ankle-foot orthoses (AFOs) can provide support and improve walking ability in individuals with plantarflexor weakness. Passive-dynamic AFO stiffness can be optimized for over-ground walking, however little research exists for uphill walking, when plantarflexor contributions are key. Purpose: Compare uphill walking biomechanics (1) between dynamic AFO users and able-bodied control subjects. (2) between injured and sound limbs (3) across different AFO stiffnesses. Methods: Twelve patients with unilateral limb-salvage and twelve matched, able-bodied controls underwent biomechanical gait analysis when walking up a 108 incline. Three AFO stiffnesses were tested in the patient group: Nominal (clinically prescribed), Compliant (20% less stiff), and Stiff (20% more stiff). Results and discussion: AFO users experienced less ankle motion and power generation, lower knee extensor moments, and greater hip flexion and power generation than controls during uphill walking. Despite these deviations, they walked at equivalent self-selected velocities and stride lengths. Asymmetries were present at the ankle and knee with decreased ankle motion and power, and lower knee extensor moments on the AFO limb. Stiffer AFOs increased knee joint flexion but a 40% range in AFO stiffness had few other effects on gait. Therefore, a wide range of clinically prescribed AFO stiffnesses may adequately assist uphill walking. Published by Elsevier B.V.
Keywords: Incline IDEO Ankle brace Gait Limb salvage Military
1. Introduction Ankle-foot orthoses (AFOs) have demonstrated ability to positively affect gait biomechanics [1], improve confidence in walking ability [2], and reduce energy cost of walking [3] across a range of conditions. The plantarflexors play an important role in providing support and forward progression during gait [4], and impaired plantarflexor power generation has been shown to hinder gait and lead to compensatory strategies (e.g. greater hip involvement) [5]. An AFO can help compensate for plantarflexor weakness and improve walking economy by both stabilizing the ankle joint [6] and utilizing elastic storage and return from AFO deformation during stance [7,8]. Recently developed passive-dynamic AFOs have been designed to provide energy-storage-and-return during a range of physical activities. One example is the intrepid dynamic exoskeletal orthosis (IDEO). The IDEO (Fig. 1) is a custom passive-dynamic
§ The view(s) expressed herein are those of the author(s) and do not reflect the official policy or position of Brooke Army Medical Center, the U.S. Army Medical Department, the U.S. Army Office of the Surgeon General, the Department of the Army, Department of Defense or the U.S. Government. * Corresponding author. Tel.: +1 2105395824. E-mail address: [email protected] (E. Russell Esposito).
carbon fiber AFO available to lower extremity limb salvage patients in the military [9]. The IDEO has been shown to result in greater functional and performance improvements than other commercially available AFOs [10,11] and has enabled limb salvage patients to walk at speeds equivalent to uninjured individuals [12]. However, the full benefits and/or limitations of the IDEO still remain unknown across a variety of other tasks, such as uphill walking. Uphill terrain is frequently encountered in activities of daily living (e.g. handicap accessible ramps) and has different musculoskeletal requirements than level ground walking. While the plantarflexor muscles are an important component of forward propulsion during level ground walking, studies have shown even greater plantarflexor activations during uphill walking for both forward and upward propulsion [13,14]. Greater flexion and range of motion (ROM) at the ankle, knee, and hip [15], as well as greater power generation, particularly at the ankle and hip, have also been reported during uphill walking compared to level walking [16]. The combination of greater plantarflexor muscle activity and ankle power generation requirements suggests that uphill walking may be particularly challenging for patients with traumatic limb salvage when range of motion is relatively constrained by an AFO. To date, research involving uphill walking with the use of AFOs is limited. Spaulding and colleagues [17] found that participants with ankle
http://dx.doi.org/10.1016/j.gaitpost.2015.01.001 0966-6362/Published by Elsevier B.V.
Please cite this article in press as: Haight DJ, et al. Biomechanics of uphill walking using custom ankle-foot orthoses of three different stiffnesses. Gait Posture (2015), http://dx.doi.org/10.1016/j.gaitpost.2015.01.001
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2. Materials and methods 2.1. Subjects Twelve males with unilateral lower-limb reconstruction volunteered to participate. Mechanisms of injuries included motor vehicle accidents (N = 3), gunshot wounds (N = 3), blast injuries (N = 5) and a high-impact landing in foot eversion (N = 1). Primary sustained injuries included multiple fractures combined with soft tissue injuries to the lower leg, ankle and foot (N = 2), massive soft tissue injuries (N = 2), combined soft tissue injury with neuropathy (N = 1), tibia/fibula fractures (N = 5), talar fracture (N = 1) and ankle fracture (N = 1). A common product of all of these injuries was impaired plantarflexion strength and/or ROM relative to the sound limb. Subjects were on average (SD) 29.3(6.0) years old, with a mean height and body mass of 1.80(0.08) m and 87.4(11.0) kg, respectively. All subjects were ambulatory without assistance other than their custom IDEO and had been using the IDEO for an average of 8(5) months. Data from 12 gender, height, and mass (5%) matched able-bodied control subjects were also used. Control subjects were on average 23.3(4.6) years old, with a mean height of 1.82(0.07) m and mean mass of 85.9(11.5) kg. This study was approved by the Institutional Review Board at Brooke Army Medical Center, Ft. Sam Houston, TX. 2.2. IDEO and stiffness modification
Fig. 1. Intrepid dynamic exoskeletal orthosis (IDEO). The IDEO brace consists of a (A) proximal tibial ‘‘ground reaction’’ cuff (B) removable, posteriorly mounted strut and (C) supramalleolar foot/ankle plate. The IDEO is augmented by a foam wedge placed under the heel of the AFO, within the shoe. All IDEO braces were custom made and fit for each patient by the same prosthetist for all patients. Prescribed height of the foam heel wedge varied between patients, but was constant across testing sessions. Photograph Courtesy of Robert Shields, Brooke Army Medical Center.
instability readily adapted their biomechanics to AFO use on a 58 incline and were only limited by the AFO in their plantarflexion motion at toe-off. Uphill walking biomechanics in a population with plantar flexion weakness from limb salvage or with IDEO use has not been studied, despite the prevalence of inclines in daily living environments. The energy-storage-and-return properties of an AFO, such as the IDEO, can be optimized to assist uphill walking, as the mechanical deformation of an AFO can be modulated to accommodate the needs of the patient. For example, simulations by Bregman et al. [18] demonstrated that AFO stiffness could be adjusted to optimize elastic storage and return and improve gait. Inadequate energy return could warrant significant strength requirements on the part of the user, such as active plantarflexion, to sufficiently deform the AFO [19]. The results of Harper et al. [20] and Russell Esposito et al. [12] showed that IDEO stiffness may influence ankle plantarflexor muscle activation and knee joint compliance during level walking, but optimal AFO mechanical properties for uphill walking have not been explored. Therefore, the purpose of the present study was to examine lower extremity kinematics and kinetics during incline walking in patients who had undergone unilateral limb salvage, and were currently using the custom IDEO. Specifically, we compared uphill walking biomechanics (1) between IDEO users and able-bodied controls subjects, (2) between the patient’s salvaged and sound limbs, and (3) across three IDEO strut stiffnesses.
The IDEO is a custom AFO consisting of a proximal carbon fiber tibial cuff and a custom molded carbon fiber foot plate connected on the posterior aspects of each via a deformable carbon fiber strut. This ankle-foot unit is augmented by a foam wedge underneath the heel, which acts to provide shock absorption during loading response (Fig. 1) [9]. In order to test the effect of AFO strut stiffness, three experimental struts were constructed using selective laser sintering [21]. A nominal strut matched the clinically prescribed stiffness (based on the patient’s available range of motion, activity level, types of activities performed, body mass, and load carriage requirements) and two other struts were constructed to be 20% less stiff (Compliant) and 20% more stiff (Stiff). The experimental struts were mechanically tested (three-point bend test) to be within 5% of the desired stiffness and weight-matched using lead tape [20]. Clinically prescribed IDEO strut stiffness averaged 798(189) Nm m 1 and ranged from 493 Nm m 1 to 1030 Nm m 1. 2.3. Experimental setup The experimental setup consisted of a 26 camera optoelectronic motion capture system (120 Hz, Motion Analysis Corp., Santa Rosa, CA) and a six degree-of-freedom marker set which allows independent tracking of each body segment of interest [22]. Fifty-seven reflective markers were affixed to anatomical landmarks and segments [23] and a digitizing pointer (C-motion, Inc., Germantown, MD) was used to identify anatomical landmarks for joint center calculations [23]. An inclinable surface, raised to 108 (Fig. 2), was instrumented with two force platforms embedded in series in the walkway (1200 Hz, AMTI, Watertown, MA). Four non co-linear markers on the walkway were used to determine the location and orientation of the force plates in the lab coordinate system. The order of the strut conditions was randomized and collected on separate days. In each session subjects walked up the incline at a controlled velocity (5%) based on leg length and a dimensionless Froude number of 0.16 [24]. Five successful trials on each side were analyzed. Self-selected walking velocity was also calculated as a descriptive characteristic. Patient preference for each strut was ranked after the final testing session. Maximal ankle plantar flexion and knee extension strength were tested on an
Please cite this article in press as: Haight DJ, et al. Biomechanics of uphill walking using custom ankle-foot orthoses of three different stiffnesses. Gait Posture (2015), http://dx.doi.org/10.1016/j.gaitpost.2015.01.001
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series data were normalized to 101 data points across the full gait cycle. Peak values and ROMs were extracted using a custom MATLAB program (Mathworks Inc., Rockville, MD). The mean peak value of five trials (for both the IDEO and sound sides) was calculated and used in statistical analysis of each dependent measure (SPSS Inc., version 19, Chicago, IL). One-way ANOVAs with Dunnet’s post hoc comparisons were used to determine differences between the control group and the salvaged and sound limb in each strut condition. A two-way (2 limbs 3 strut stiffnesses) repeated measures ANOVA was used to determine statistical significance within the patient group. Post hoc paired t-tests with Bonferroni–Holm adjustments separated significant main effects. The unadjusted criterion for statistical significance was set at P < 0.05. 3. Results Groups were not significantly different in terms of height, mass, or self-selected walking velocity but age was greater in the patient group than the control group (P = 0.01). Maximal isometric testing was only completed on subjects who could do so without pain above 4/10 on a visual analog scale. Eight of the 12 subjects completed maximal isometric plantarflexion contractions, nine completed knee extension contractions, and deficits in the affected side were 50.0 and 25.1%, respectively, relative to the sound side. Of the 12 patients, three preferred the compliant strut, four preferred the nominal strut, and three preferred the stiff strut. One subject preferred the compliant and nominal struts equally, while another was not able to discern any differences among strut stiffnesses.
Fig. 2. Instrumented, inclinable walkway used to collect inclined gait biomechanics. As shown, the walkway is inclined to ten degrees. The walkway contains two 600 mm 600 mm force platforms embedded in series. Force platforms have been outlined in black to show location. Photograph courtesy of Robert Shields, Brooke Army Medical Center.
3.1. Temporal-spatial parameters (Table 1)
isometric dynamometer (Biodex System 3, Biodex Medical Systems, Shirley, NY).
Self-selected walking velocity, stride length, and cycle times were not different between groups. Step width was significantly greater in the patient group for all strut conditions (P < 0.020), while double support time was only less in the patient group while using the nominal and stiff struts (P < 0.034). Stance time was shorter (P < 0.001) and swing time was longer (P < 0.033) for the IDEO limb, compared to the sound limb but there were no other significant inter-limb differences or differences among strut stiffness conditions.
2.4. Analysis The platform was tracked as a rigid body in order to determine the precise orientation and location of the force plates. The use of force structures in Visual 3D (C-motion, Inc., Germantown, MD) makes it possible to correctly associate forces from an inclined force plate to the appropriate limb. Marker trajectory and analog data were digitally low-pass filtered at 6 and 50 Hz, respectively, using fourth-order Butterworth filters. Coordinate systems of the foot, leg, thigh, pelvis, and trunk segments were defined in a manner consistent with International Society of Biomechanics standards [25]. Lower extremity kinematics were computed using a Newton–Euler angle approach within Visual3D and kinematics were combined with GRF data to calculate net internal joint moments and powers using an inverse dynamics approach. Time
3.2. Kinematics (Fig. 3, Table 2) IDEO use significantly reduced ankle ROM compared to control subjects (P < 0.001) and the sound limb (P < 0.001). This finding resulted from a combination of less dorsiflexion during stance (P < 0.002) and less peak plantarflexion (P < 0.001). The patient’s sound limb responded to the stiffest strut by increasing ankle ROM to a greater extent than controls (P = 0.029). In all stiffness conditions both the IDEO (P < 0.030) and sound (P < 0.013) limbs underwent greater hip flexion than controls. However, the sound limb underwent greater hip extension compared to the sound limb in the nominal (P = 0.039) and stiff (P = 0.048) struts. Kinematic differences among strut conditions were limited, however, the knee was more flexed at initial contact in the complaint (P = 0.009) and nominal (P = 0.015) conditions compared to the stiff.
Table 1 Mean (SD) temporal-spatial gait parameters. Patient
Controls
Between groups Self-selected vel. Stride length Stride width Stride time Double support time
1.22 1.47 0.12 1.19 0.31
(0.13) (0.11) (0.03) (0.09) (0.04)
Compliant
Nominal
Stiff
1.21 1.41 0.15 1.16 0.29
1.18 1.42 0.15 1.13 0.27
1.19 1.40 0.15 1.13 0.27
(0.15) (0.14) (0.03)* (0.08) (0.05)
IDEO limb
Between limbs Step length Step time Stance time Swing time
Nominal
Stiff
0.72 0.59 0.70 0.47
0.71 0.57 0.67 0.46
0.69 0.57 0.68 0.46
(0.08) (0.04) (0.03)L (0.04)L
(0.12) (0.13) (0.02)* (0.7) (0.03)*
Sound limb
Compliant (0.07) (0.05) (0.06)L (0.04)L
(0.10) (0.13) (0.03)* (0.06) (0.03)*
(0.06) (0.05) (0.04)L (0.04)L
Compliant
Nominal
Stiff
0.70 0.57 0.75 0.40
0.72 0.56 0.73 0.40
0.71 0.56 0.73 0.40
(0.08) (0.03) (0.06) (0.03)
(0.07) (0.03) (0.05) (0.02)
(0.08) (0.03) (0.05) (0.03)
No significant differences were found among strut conditions. * Indicates significantly different from control subjects. L Indicates significantly different from sound limb.
Please cite this article in press as: Haight DJ, et al. Biomechanics of uphill walking using custom ankle-foot orthoses of three different stiffnesses. Gait Posture (2015), http://dx.doi.org/10.1016/j.gaitpost.2015.01.001
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Fig. 3. Mean joint angles (ankle dorsi-flexion(+)/plantar-flexion, knee flexion(+)/extension, hip flexion(+)/extension), internal joint moments (plantarflexor(+)/dorsiflexor, knee extensor(+)/flexor, and hip extension(+)/flexion), and joint powers for the (A) IDEO limb and (B) sound limb. Black vertical lines indicate toe-off for controls subjects and gray vertical lines indicate toe-off for IDEO users.
Please cite this article in press as: Haight DJ, et al. Biomechanics of uphill walking using custom ankle-foot orthoses of three different stiffnesses. Gait Posture (2015), http://dx.doi.org/10.1016/j.gaitpost.2015.01.001
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Table 2 Kinetics and kinematics. IDEO limb
Controls Compliant Joint angles Peak DF stance Peak PF – initial swing Ankle ROM Knee flex. – initial contact Peak knee flex. – loading Peak knee ext. – stance Peak knee flex. – swing Knee ROM Peak hip flex. – stance Peak hip ext. – stance Peak hip flex – swing Hip ROM – sagittal
21.5 14.5 36.0 28.4 34.8 1.4 61.7 60.3 53.4 6.7 53.6 60.5
(3.3) (4.7) (5.0) (7.0) (6.3) (3.3) (3.6) (3.8) (6.6) (7.0) (6.3) (3.2)
15.3 6.4 10.0 30.6 34.4 1.1 62.9 61.8 59.7 0.2 60.7 61.0
(5.4)*L (4.7)*L (2.9)*L (12.1)S (13.1) (10.0) (7.1) (9.6) (6.4)* (7.7) (5.9)* (6.1)
Nominal 15.3 7.3 8.6 36.3 41.1 6.2 63.0 56.9 62.5 1.5 63.1 62.1
Sound limb Stiff
(4.4)*L (3.3)*L (2.1)*L (11.1)S (12.9) (10.2) (6.6) (12.1) (5.3)* (7.4)* (4.7)* (5.9)
15.0 7.6 8.2 39.6 43.3 7.1 65.2 58.0 61.8 1.2 62.4 61.9
Compliant (2.8)*L (2.2)*L (2.2)*L (12.2)*CN (13.6) (10.4) (4.0) (9.7) (5.0)* (7.3)* (4.4)* (4.9)
22.8 (2.9) 19.2 (5.8) 41.9 (6.7) 31.0 (8.7) 36.8 (8.0) 0.32 (5.4) 60.9 (4.7) 61.2 (6.5) 59.3 (4.9)* 0.7 (7.8) 60.3 (5.1)* 61.8 (7.5)
Nominal 24.7 17.1 41.7 34.4 40.2 1.8 62.9 61.1 61.6 0.1 61.0 62.1
(4.8) (7.2) (6.7)* (8.0) (8.0) (4.0) (5.3) (3.2) (3.2)* (7.6) (4.3)* (6.1)
Stiff 23.8 18.8 42.6 34.1 39.6 0.9 61.6 60.7 59.9 1.9 60.0 62.8
(3.6) (6.4) (5.8)* (9.5) (8.5) (5.6) (6.0) (7.0) (4.1)* (8.0) (4.7)* (7.4)
GRF (BW) Peak A/P – braking Peak A/P – propulsive Peak vertical – early stance Peak vertical – late stance Peak M/L
0.15 0.17 1.05 1.15 0.05
(0.03) (0.03) (0.06) (0.09) (0.01)
0.17 0.14 1.03 1.08 0.07
(0.05)L (0.07)L (0.08)L (0.13)L (0.01)*
0.18 0.16 1.05 1.10 0.08
(0.05) (0.06)L (0.06)L (0.13)L (0.02)*
0.18 0.15 1.06 1.11 0.07
(0.04)L (0.06)L (0.05)L (0.14)L (0.01)*
0.14 0.19 1.22 1.23 0.07
(0.04) (0.05) (0.10)*NS (0.12) (0.02)*NS
0.16 0.20 1.30 1.25 0.08
(0.04) (0.05) (0.11)*C (0.11) (0.02)*C
0.15 0.20 1.28 1.24 0.08
(0.04) (0.05) (0.09)*C (0.11) (0.02)*C
Moments (Nm kg 1) Peak DF – loading response Peak PF – terminal stance Peak knee ext. – loading response Peak knee flex. – terminal stance Peak hip ext. – stance Peak hip flex. – terminal stance
0.09 1.73 0.87 0.54 1.26 0.61
(0.05) (0.15) (0.32) (0.15) (0.26) (0.23)
0.03 1.76 0.19 0.72 1.49 0.55
(0.14) (0.41) (0.31)*L (0.30) (0.29) (0.18)
0.02 1.82 0.29 0.63 1.58 0.55
(0.11) (0.34) (0.38)*L (0.38) (0.34)* (0.14)L
0.01 1.81 0.38 0.59 1.52 0.56
(0.12) (0.40) (0.38)*L (0.35) (0.33) (0.22)
0.08 1.78 0.86 0.71 1.63 0.49
(0.06) (0.25) (0.39) (0.24) (0.38)* (0.20)
0.06 1.85 1.02 0.72 1.86 0.45
(0.08) (0.28) 0.46) (0.20) (0.47)* (0.20)
0.06 1.83 0.92 0.74 1.71 0.48
(0.08) (0.22) (0.33) (0.23) (0.35)* (0.15)
Powers (W kg 1) Ankle abs. – loading response Ankle abs. – terminal stance Peak ankle gen. – pre-swing Peak knee gen. – early mid stance Peak hip gen. – loading response Peak hip abs. – terminal stance
0.21 0.31 3.67 1.19 1.74 0.58
(0.21) (0.42) (0.83) (0.57) (0.24) (0.30)
0.86 0.80 1.23 0.47 2.51 0.50
(0.78)* (0.35)* (0.50)*L (0.56)*L (0.47)* (0.33)
0.87 0.71 1.19 0.56 2.82 0.58
(0.38)* (0.36)* (0.28)*L (0.58)*L (0.73)* (0.42)
0.72 0.69 1.20 0.69 2.68 0.53
(0.42)* (0.38) (0.43)*L (0.51)L (0.63)* (0.27)
0.66 0.74 4.02 1.29 2.41 0.40
(1.22) (0.49) (1.03) (0.62) (0.83)* (0.35)
0.69 0.67 4.05 1.51 2.70 0.36
(0.94) (0.48) (0.98) (0.77) (0.65)* (0.35)
0.87 0.56 4.04 1.35 2.39 0.38
(1.03) (0.33) (0.94) (0.53) (0.74)* (0.24)
* C N S L
Indicates significantly different from controls. Indicates significantly different from compliant strut condition (within limb). Indicates significantly different from nominal strut condition (within limb). Indicates significantly different from stiff strut condition (within limb). Indicates significantly different from sound limb (within strut condition).
3.3. GRF (Table 2) A significant interaction was observed for the first peak in the vertical GRF (P = 0.008) such that the nominal and stiff struts resulted in greater forces than the compliant in the sound limb, but not the IDEO limb. Overall, the sound limb experienced greater peak forces than the IDEO limb (P < 0.001) and the control group (P < 0.002). IDEO users experienced other inter-limb differences, as well. Both the first (P < 0.001) and second (P = 0.022) peaks in the vertical force and the peak propulsive force (P = 0.001) were greater on the sound side compared to the IDEO side. On the sound side, the first peak vertical GRF and the peak medial GRF were lowest in the compliant strut (P < 0.030). 3.4. Moments (Fig. 3, Table 2) Peak net internal joint moments differed between patients and controls, and between IDEO and sound limbs, but were not significantly different among strut conditions. Internal ankle plantarflexor moments exhibited a distinct local minima in mid stance that was not found in the control group, thus no statistical comparisons were made, and were not significantly different between groups in late stance. There was a significant interaction between stiffness condition and limb in the peak knee extensor moment (P = 0.043) such that it was lower in patients’ IDEO limb compared to sound limb (P < 0.001), but post hoc tests revealed no statistically significant differences among struts. Comparisons to controls indicated significantly lower peak knee extensor moments in the IDEO limb (P < 0.004) and greater hip extensor moments in the sound limb (P < 0.047) of the patient group.
3.5. Powers (Fig. 3, Table 2) The patients’ IDEO limb exhibited significantly less peak ankle power generation compared to both the control group (P < 0.001) and the sound limb (P < 0.001). Peak knee power generation of the IDEO limb during mid-stance was less than that
of the control group while using the compliant (P = 0.008) and nominal struts (P = 0.020), but not the stiff strut. Between limb comparisons indicated that the peak knee power generation during mid-stance was less in the IDEO limb than sound limb (P < 0.003). There was also a bi-lateral increase in peak hip power generation in the patient group, compared to controls for all strut conditions (P < 0.050).
4. Discussion 4.1. Comparisons to controls Use of the custom IDEO enabled limb salvage patients to successfully complete an uphill walking protocol at equivalent self-selected speeds and stride lengths to able-bodied individuals. The IDEO users accomplished the uphill walking task despite deviations in their lower extremity mechanics. Consistent with existing level ground walking literature [7,12,20,26], AFO use limited ankle ROM, which resulted in 77% less ankle power generation at push-off. Previous investigators have found between 42% [26] and 80% [8] less power generation than controls while using carbon fiber AFOs. Limited ankle dorsiflexion during stance due to the semi-rigid nature of the IDEO led to decreased ipsilateral knee extensor moment (all strut conditions) and peak power generation (compliant and nominal struts). The bilateral increase in peak hip power generation during early stance in the patient group compared to control subjects was likely a compensatory strategy
Please cite this article in press as: Haight DJ, et al. Biomechanics of uphill walking using custom ankle-foot orthoses of three different stiffnesses. Gait Posture (2015), http://dx.doi.org/10.1016/j.gaitpost.2015.01.001
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[5] for less ankle and knee power generation while using the IDEO. Although research on uphill walking with an AFO is very limited, the observed hip strategy, characterized by greater hip flexion and power generation, has been described during over-ground walking with a conceptually similar AFO [26]. The adaptations at proximal joints may be influenced by changes in AFO design. Changes such as increasing motion by using an articulated ankle joint or other approaches to increase compliance or ankle power may preclude the need for greater hip flexion and power generation in the AFO users. Further research is needed to draw firm conclusions. 4.2. Comparisons between limbs While it is recognized that the IDEO improves functional capacity to a greater extent than other AFOs in this population [11], it is not realistic to expect symmetrical gait, given the restrictions in ankle ROM. Ankle plantarflexion plays an important role at push-off for ankle power generation and ankle dorsiflexion is key for aligning the limb to utilize ipsilateral knee power. It is possible that new powered AFOs could address power generation asymmetries at the ankle, although, compensatory hip strategies and asymmetries are likely to persist unless a full ankle ROM is achieved. However, the IDEO was created for patients with a variety of injuries and surgical procedures, including ankle fracture or fusion, and as a result large excursions can be contraindicated. Therefore, sacrifices in normative biomechanics may be necessary in order to improve overall functional capacity. 4.3. Comparisons among struts Optimal AFO stiffness has been previously studied and is thought to play an important role in the optimization of walking [18,27]. However, modulating IDEO strut stiffness by up to 40% in the present study yielded few significant biomechanical effects during uphill walking. These results are consistent with our previously published results from overground walking [12,20] but it was expected that due to the greater contribution of the plantarflexors and the greater dorsiflexion range of motion during uphill walking lower extremity mechanics would be affected to a greater extent by strut stiffness. The primary kinematic difference among strut stiffnesses was that the knee of the IDEO limb contacted the ground in a more extended position in the two more compliant struts. This finding is consistent with previous overground research [12,28], where the authors showed that early stance knee angles increased with greater resistance to ankle plantarflexion, as would be found with stiffer struts. 4.4. Applications Patients with lower limb salvage who used the IDEO were able to accomplish uphill walking despite substantial plantar flexor weakness. AFOs of a similar design may assist other populations with lower-limb weakness caused by stroke, brain injury, or other neuropathies to achieve functional improvements. The limited effects of stiffness suggest that a range of clinically prescribed IDEO stiffnesses could adequately address the biomechanical requirements of uphill walking. It is possible that the effects of modulating stiffness may have been greater if the initial prescribed strut was more compliant. However, a stiffer strut may better transition to high impact activities without negatively affecting the mechanics of lower impact walking activities. Additional research on high intensity activities is needed to determine if clinically prescribed IDEO stiffness is robust across a variety of activities or if an optimal stiffness becomes more important at higher intensities.
4.5. Limitations Not all patients were able to successfully complete a gait analysis without their IDEO; hence, there is no comparison to walking without an AFO. The authors expect that IDEO use improved gait, but are not able to quantify the extent of its benefit or delineate explicitly if certain gait deviations were from injuries/ plantarflexor weakness or a result of using the IDEO. Furthermore, the combat-related injuries sustained by the limb salvage patients were diverse and subjects were at varying stages of rehabilitation at the time of their participation. However, self-selected walking velocity has been used as an indicator of recovery following lower limb trauma [29] and the equivalent velocities between the two groups may indicate that patients reached a level of recovery where relevant comparisons could be made. 5. Conclusions Use of the passive-dynamic IDEO enabled subjects with substantial plantarflexor weakness to walk uphill at speeds and stride lengths equivalent to able-bodies controls. However, due to limitations in ankle ROM, compensations at proximal joints were observed and asymmetries persisted. Modulating IDEO stiffness had few effects on the biomechanics of uphill walking; however future study is needed to determine if this is robust across a wider range of activities. Acknowledgements The authors acknowledge Nicole Harper and Dr. Richard Neptune at the University of Texas for strut manufacturing and testing. Support was provided by the Center for Rehabilitation Sciences Research Department of Physical Medicine and Rehabilitation, Uniformed Services University of Health Sciences, Bethesda, MD. Conflict of interest The authors, Derek J. Haight, MS, Elizabeth Russell Esposito, PhD, and Jason M Wilken, PhD, MPT, declare that they have no competing interests that would affect interpretation of the data. References [1] Franceschini M, Massucci M, Ferrari L, Agosti M, Paroli C. Effects of an anklefoot orthosis on spatiotemporal parameters and energy cost of hemiparetic gait. Clin Rehabil 2003;17:368–72. [2] de Wit DCM, Buurke JH, Nijlant JMM, Ijzerman MJ, Hermens HJ. The effect of an ankle-foot orthosis on walking ability in chronic stroke patients: a randomized controlled trial. Clin Rehabil 2004;18:550–7. [3] Danielsson A, Sunnerhagen KS. Energy expenditure in stroke subjects walking with a carbon composite ankle foot orthosis. J Rehabil Med 2004;36:165–8. [4] Neptune RR, Kautz SA, Zajac FE. Contributions of the individual ankle plantar flexors to support, forward progression and swing initiation during walking. J Biomech 2001;34:1387–98. [5] Nadeau S, Gravel D, Arsenault AB, Bourbonnais D. Plantarflexor weakness as a limiting factor of gait speed in stroke subjects and the compensating role of hip flexors. Clin Biochem 1999;14:125–35. [6] Esquenazi A, Ofluoglu D, Hirai B, Kim S. The effect of an ankle-foot orthosis on temporal spatial parameters and asymmetry of gait in hemiparetic patients. PM R 2009;1:1014–8. [7] Bregman DJJ, Harlaar J, Meskers CGM, de Groot V. Spring-like ankle foot orthoses reduce the energy cost of walking by taking over ankle work. Gait Posture 2012;35:148–53. [8] Wolf SI, Alimusaj M, Rettig O, Do¨derlein L. Dynamic assist by carbon fiber spring AFOs for patients with myelomeningocele. Gait Posture 2008;28:175–7. [9] Patzkowski JC, Blanck RV, Owens JG, Wilken JM, Blair JA, Hsu JR. Can an anklefoot orthosis change hearts and minds. J Surg Orthop Adv 2011;20:8–18. [10] Bedigrew KM, Patzkowski JC, Wilken JM, Owens JG, Blanck RV, Stinner DJ, et al. Can an integrated orthotic and rehabilitation program decrease pain and improve function after lower extremity trauma? Clin Orthop Relat Res 2014;472:3017–25.
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[21] Faustini MC, Neptune RR, Crawford RH, Stanhope SJ. Manufacture of passive dynamic ankle – foot orthoses using selective laser sintering. IEEE Trans Biomed Eng 2008;55:784–90. [22] Collins TD, Ghoussayni SN, Ewins DJ, Kent JA. A six degrees-of-freedom marker set for gait analysis: repeatability and comparison with a modified Helen Hayes set. Gait Posture 2009;30:173–80. [23] Wilken JM, Rodriguez KM, Brawner M, Darter BJ. Reliability and minimal detectible change values for gait kinematics and kinetics in healthy adults. Gait Posture 2012;35:301–7. [24] Vaughan CL, O’Malley MJ. Froude and the contribution of naval architecture to our understanding of bipedal locomotion. Gait Posture 2005;21:350–62. [25] Wu G, Siegler S, Allard P, Kirtley C, Leardini A, Rosenbaum D, et al. ISB recommendation on definitions of joint coordinate system of various joints for the reporting of human joint motion – Part I: Ankle, hip, and spine. J Biomech 2002;35:543–8. [26] Desloovere K, Molenaers G, Van Gestel L, Huenaerts C, Van Campenhout A, Callewaert B, et al. How can push-off be preserved during use of an ankle foot orthosis in children with hemiplegia? A prospective controlled study. Gait Posture 2006;24:142–51. [27] Sumiya T, Suzuki Y, Kasahara T. Stiffness control in posterior-type plastic ankle-foot orthoses: effect of ankle trimline. Part 2: Orthosis characteristics and orthosis/patient matching. Prosthet Orthot Int 1996;20:132–7. [28] Kobayashi T, Leung AKL, Akazawa Y, Hutchins SW. The effect of varying the plantarflexion resistance of an ankle-foot orthosis on knee joint kinematics in patients with stroke. Gait Posture 2013;37:457–9. [29] Baker P, Hewison S. Gait recovery pattern of unilateral lower limb amputees during rehabilitation. Prosthet Orthot Int 1990;14:80–4.
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Biomechanics of uphill walking using custom ankle-foot orthoses of three different stiffnesses§ Derek J. Haight, Elizabeth Russell Esposito *, Jason M. Wilken Center for the Intrepid, Department of Orthopaedics and Rehabilitation, Brooke Army Medical Center, Ft. Sam Houston, TX 78234, United States
A R T I C L E I N F O
A B S T R A C T
Article history: Received 14 August 2014 Received in revised form 12 December 2014 Accepted 3 January 2015
Ankle-foot orthoses (AFOs) can provide support and improve walking ability in individuals with plantarflexor weakness. Passive-dynamic AFO stiffness can be optimized for over-ground walking, however little research exists for uphill walking, when plantarflexor contributions are key. Purpose: Compare uphill walking biomechanics (1) between dynamic AFO users and able-bodied control subjects. (2) between injured and sound limbs (3) across different AFO stiffnesses. Methods: Twelve patients with unilateral limb-salvage and twelve matched, able-bodied controls underwent biomechanical gait analysis when walking up a 108 incline. Three AFO stiffnesses were tested in the patient group: Nominal (clinically prescribed), Compliant (20% less stiff), and Stiff (20% more stiff). Results and discussion: AFO users experienced less ankle motion and power generation, lower knee extensor moments, and greater hip flexion and power generation than controls during uphill walking. Despite these deviations, they walked at equivalent self-selected velocities and stride lengths. Asymmetries were present at the ankle and knee with decreased ankle motion and power, and lower knee extensor moments on the AFO limb. Stiffer AFOs increased knee joint flexion but a 40% range in AFO stiffness had few other effects on gait. Therefore, a wide range of clinically prescribed AFO stiffnesses may adequately assist uphill walking. Published by Elsevier B.V.
Keywords: Incline IDEO Ankle brace Gait Limb salvage Military
1. Introduction Ankle-foot orthoses (AFOs) have demonstrated ability to positively affect gait biomechanics [1], improve confidence in walking ability [2], and reduce energy cost of walking [3] across a range of conditions. The plantarflexors play an important role in providing support and forward progression during gait [4], and impaired plantarflexor power generation has been shown to hinder gait and lead to compensatory strategies (e.g. greater hip involvement) [5]. An AFO can help compensate for plantarflexor weakness and improve walking economy by both stabilizing the ankle joint [6] and utilizing elastic storage and return from AFO deformation during stance [7,8]. Recently developed passive-dynamic AFOs have been designed to provide energy-storage-and-return during a range of physical activities. One example is the intrepid dynamic exoskeletal orthosis (IDEO). The IDEO (Fig. 1) is a custom passive-dynamic
§ The view(s) expressed herein are those of the author(s) and do not reflect the official policy or position of Brooke Army Medical Center, the U.S. Army Medical Department, the U.S. Army Office of the Surgeon General, the Department of the Army, Department of Defense or the U.S. Government. * Corresponding author. Tel.: +1 2105395824. E-mail address: [email protected] (E. Russell Esposito).
carbon fiber AFO available to lower extremity limb salvage patients in the military [9]. The IDEO has been shown to result in greater functional and performance improvements than other commercially available AFOs [10,11] and has enabled limb salvage patients to walk at speeds equivalent to uninjured individuals [12]. However, the full benefits and/or limitations of the IDEO still remain unknown across a variety of other tasks, such as uphill walking. Uphill terrain is frequently encountered in activities of daily living (e.g. handicap accessible ramps) and has different musculoskeletal requirements than level ground walking. While the plantarflexor muscles are an important component of forward propulsion during level ground walking, studies have shown even greater plantarflexor activations during uphill walking for both forward and upward propulsion [13,14]. Greater flexion and range of motion (ROM) at the ankle, knee, and hip [15], as well as greater power generation, particularly at the ankle and hip, have also been reported during uphill walking compared to level walking [16]. The combination of greater plantarflexor muscle activity and ankle power generation requirements suggests that uphill walking may be particularly challenging for patients with traumatic limb salvage when range of motion is relatively constrained by an AFO. To date, research involving uphill walking with the use of AFOs is limited. Spaulding and colleagues [17] found that participants with ankle
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2. Materials and methods 2.1. Subjects Twelve males with unilateral lower-limb reconstruction volunteered to participate. Mechanisms of injuries included motor vehicle accidents (N = 3), gunshot wounds (N = 3), blast injuries (N = 5) and a high-impact landing in foot eversion (N = 1). Primary sustained injuries included multiple fractures combined with soft tissue injuries to the lower leg, ankle and foot (N = 2), massive soft tissue injuries (N = 2), combined soft tissue injury with neuropathy (N = 1), tibia/fibula fractures (N = 5), talar fracture (N = 1) and ankle fracture (N = 1). A common product of all of these injuries was impaired plantarflexion strength and/or ROM relative to the sound limb. Subjects were on average (SD) 29.3(6.0) years old, with a mean height and body mass of 1.80(0.08) m and 87.4(11.0) kg, respectively. All subjects were ambulatory without assistance other than their custom IDEO and had been using the IDEO for an average of 8(5) months. Data from 12 gender, height, and mass (5%) matched able-bodied control subjects were also used. Control subjects were on average 23.3(4.6) years old, with a mean height of 1.82(0.07) m and mean mass of 85.9(11.5) kg. This study was approved by the Institutional Review Board at Brooke Army Medical Center, Ft. Sam Houston, TX. 2.2. IDEO and stiffness modification
Fig. 1. Intrepid dynamic exoskeletal orthosis (IDEO). The IDEO brace consists of a (A) proximal tibial ‘‘ground reaction’’ cuff (B) removable, posteriorly mounted strut and (C) supramalleolar foot/ankle plate. The IDEO is augmented by a foam wedge placed under the heel of the AFO, within the shoe. All IDEO braces were custom made and fit for each patient by the same prosthetist for all patients. Prescribed height of the foam heel wedge varied between patients, but was constant across testing sessions. Photograph Courtesy of Robert Shields, Brooke Army Medical Center.
instability readily adapted their biomechanics to AFO use on a 58 incline and were only limited by the AFO in their plantarflexion motion at toe-off. Uphill walking biomechanics in a population with plantar flexion weakness from limb salvage or with IDEO use has not been studied, despite the prevalence of inclines in daily living environments. The energy-storage-and-return properties of an AFO, such as the IDEO, can be optimized to assist uphill walking, as the mechanical deformation of an AFO can be modulated to accommodate the needs of the patient. For example, simulations by Bregman et al. [18] demonstrated that AFO stiffness could be adjusted to optimize elastic storage and return and improve gait. Inadequate energy return could warrant significant strength requirements on the part of the user, such as active plantarflexion, to sufficiently deform the AFO [19]. The results of Harper et al. [20] and Russell Esposito et al. [12] showed that IDEO stiffness may influence ankle plantarflexor muscle activation and knee joint compliance during level walking, but optimal AFO mechanical properties for uphill walking have not been explored. Therefore, the purpose of the present study was to examine lower extremity kinematics and kinetics during incline walking in patients who had undergone unilateral limb salvage, and were currently using the custom IDEO. Specifically, we compared uphill walking biomechanics (1) between IDEO users and able-bodied controls subjects, (2) between the patient’s salvaged and sound limbs, and (3) across three IDEO strut stiffnesses.
The IDEO is a custom AFO consisting of a proximal carbon fiber tibial cuff and a custom molded carbon fiber foot plate connected on the posterior aspects of each via a deformable carbon fiber strut. This ankle-foot unit is augmented by a foam wedge underneath the heel, which acts to provide shock absorption during loading response (Fig. 1) [9]. In order to test the effect of AFO strut stiffness, three experimental struts were constructed using selective laser sintering [21]. A nominal strut matched the clinically prescribed stiffness (based on the patient’s available range of motion, activity level, types of activities performed, body mass, and load carriage requirements) and two other struts were constructed to be 20% less stiff (Compliant) and 20% more stiff (Stiff). The experimental struts were mechanically tested (three-point bend test) to be within 5% of the desired stiffness and weight-matched using lead tape [20]. Clinically prescribed IDEO strut stiffness averaged 798(189) Nm m 1 and ranged from 493 Nm m 1 to 1030 Nm m 1. 2.3. Experimental setup The experimental setup consisted of a 26 camera optoelectronic motion capture system (120 Hz, Motion Analysis Corp., Santa Rosa, CA) and a six degree-of-freedom marker set which allows independent tracking of each body segment of interest [22]. Fifty-seven reflective markers were affixed to anatomical landmarks and segments [23] and a digitizing pointer (C-motion, Inc., Germantown, MD) was used to identify anatomical landmarks for joint center calculations [23]. An inclinable surface, raised to 108 (Fig. 2), was instrumented with two force platforms embedded in series in the walkway (1200 Hz, AMTI, Watertown, MA). Four non co-linear markers on the walkway were used to determine the location and orientation of the force plates in the lab coordinate system. The order of the strut conditions was randomized and collected on separate days. In each session subjects walked up the incline at a controlled velocity (5%) based on leg length and a dimensionless Froude number of 0.16 [24]. Five successful trials on each side were analyzed. Self-selected walking velocity was also calculated as a descriptive characteristic. Patient preference for each strut was ranked after the final testing session. Maximal ankle plantar flexion and knee extension strength were tested on an
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series data were normalized to 101 data points across the full gait cycle. Peak values and ROMs were extracted using a custom MATLAB program (Mathworks Inc., Rockville, MD). The mean peak value of five trials (for both the IDEO and sound sides) was calculated and used in statistical analysis of each dependent measure (SPSS Inc., version 19, Chicago, IL). One-way ANOVAs with Dunnet’s post hoc comparisons were used to determine differences between the control group and the salvaged and sound limb in each strut condition. A two-way (2 limbs 3 strut stiffnesses) repeated measures ANOVA was used to determine statistical significance within the patient group. Post hoc paired t-tests with Bonferroni–Holm adjustments separated significant main effects. The unadjusted criterion for statistical significance was set at P < 0.05. 3. Results Groups were not significantly different in terms of height, mass, or self-selected walking velocity but age was greater in the patient group than the control group (P = 0.01). Maximal isometric testing was only completed on subjects who could do so without pain above 4/10 on a visual analog scale. Eight of the 12 subjects completed maximal isometric plantarflexion contractions, nine completed knee extension contractions, and deficits in the affected side were 50.0 and 25.1%, respectively, relative to the sound side. Of the 12 patients, three preferred the compliant strut, four preferred the nominal strut, and three preferred the stiff strut. One subject preferred the compliant and nominal struts equally, while another was not able to discern any differences among strut stiffnesses.
Fig. 2. Instrumented, inclinable walkway used to collect inclined gait biomechanics. As shown, the walkway is inclined to ten degrees. The walkway contains two 600 mm 600 mm force platforms embedded in series. Force platforms have been outlined in black to show location. Photograph courtesy of Robert Shields, Brooke Army Medical Center.
3.1. Temporal-spatial parameters (Table 1)
isometric dynamometer (Biodex System 3, Biodex Medical Systems, Shirley, NY).
Self-selected walking velocity, stride length, and cycle times were not different between groups. Step width was significantly greater in the patient group for all strut conditions (P < 0.020), while double support time was only less in the patient group while using the nominal and stiff struts (P < 0.034). Stance time was shorter (P < 0.001) and swing time was longer (P < 0.033) for the IDEO limb, compared to the sound limb but there were no other significant inter-limb differences or differences among strut stiffness conditions.
2.4. Analysis The platform was tracked as a rigid body in order to determine the precise orientation and location of the force plates. The use of force structures in Visual 3D (C-motion, Inc., Germantown, MD) makes it possible to correctly associate forces from an inclined force plate to the appropriate limb. Marker trajectory and analog data were digitally low-pass filtered at 6 and 50 Hz, respectively, using fourth-order Butterworth filters. Coordinate systems of the foot, leg, thigh, pelvis, and trunk segments were defined in a manner consistent with International Society of Biomechanics standards [25]. Lower extremity kinematics were computed using a Newton–Euler angle approach within Visual3D and kinematics were combined with GRF data to calculate net internal joint moments and powers using an inverse dynamics approach. Time
3.2. Kinematics (Fig. 3, Table 2) IDEO use significantly reduced ankle ROM compared to control subjects (P < 0.001) and the sound limb (P < 0.001). This finding resulted from a combination of less dorsiflexion during stance (P < 0.002) and less peak plantarflexion (P < 0.001). The patient’s sound limb responded to the stiffest strut by increasing ankle ROM to a greater extent than controls (P = 0.029). In all stiffness conditions both the IDEO (P < 0.030) and sound (P < 0.013) limbs underwent greater hip flexion than controls. However, the sound limb underwent greater hip extension compared to the sound limb in the nominal (P = 0.039) and stiff (P = 0.048) struts. Kinematic differences among strut conditions were limited, however, the knee was more flexed at initial contact in the complaint (P = 0.009) and nominal (P = 0.015) conditions compared to the stiff.
Table 1 Mean (SD) temporal-spatial gait parameters. Patient
Controls
Between groups Self-selected vel. Stride length Stride width Stride time Double support time
1.22 1.47 0.12 1.19 0.31
(0.13) (0.11) (0.03) (0.09) (0.04)
Compliant
Nominal
Stiff
1.21 1.41 0.15 1.16 0.29
1.18 1.42 0.15 1.13 0.27
1.19 1.40 0.15 1.13 0.27
(0.15) (0.14) (0.03)* (0.08) (0.05)
IDEO limb
Between limbs Step length Step time Stance time Swing time
Nominal
Stiff
0.72 0.59 0.70 0.47
0.71 0.57 0.67 0.46
0.69 0.57 0.68 0.46
(0.08) (0.04) (0.03)L (0.04)L
(0.12) (0.13) (0.02)* (0.7) (0.03)*
Sound limb
Compliant (0.07) (0.05) (0.06)L (0.04)L
(0.10) (0.13) (0.03)* (0.06) (0.03)*
(0.06) (0.05) (0.04)L (0.04)L
Compliant
Nominal
Stiff
0.70 0.57 0.75 0.40
0.72 0.56 0.73 0.40
0.71 0.56 0.73 0.40
(0.08) (0.03) (0.06) (0.03)
(0.07) (0.03) (0.05) (0.02)
(0.08) (0.03) (0.05) (0.03)
No significant differences were found among strut conditions. * Indicates significantly different from control subjects. L Indicates significantly different from sound limb.
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Fig. 3. Mean joint angles (ankle dorsi-flexion(+)/plantar-flexion, knee flexion(+)/extension, hip flexion(+)/extension), internal joint moments (plantarflexor(+)/dorsiflexor, knee extensor(+)/flexor, and hip extension(+)/flexion), and joint powers for the (A) IDEO limb and (B) sound limb. Black vertical lines indicate toe-off for controls subjects and gray vertical lines indicate toe-off for IDEO users.
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Table 2 Kinetics and kinematics. IDEO limb
Controls Compliant Joint angles Peak DF stance Peak PF – initial swing Ankle ROM Knee flex. – initial contact Peak knee flex. – loading Peak knee ext. – stance Peak knee flex. – swing Knee ROM Peak hip flex. – stance Peak hip ext. – stance Peak hip flex – swing Hip ROM – sagittal
21.5 14.5 36.0 28.4 34.8 1.4 61.7 60.3 53.4 6.7 53.6 60.5
(3.3) (4.7) (5.0) (7.0) (6.3) (3.3) (3.6) (3.8) (6.6) (7.0) (6.3) (3.2)
15.3 6.4 10.0 30.6 34.4 1.1 62.9 61.8 59.7 0.2 60.7 61.0
(5.4)*L (4.7)*L (2.9)*L (12.1)S (13.1) (10.0) (7.1) (9.6) (6.4)* (7.7) (5.9)* (6.1)
Nominal 15.3 7.3 8.6 36.3 41.1 6.2 63.0 56.9 62.5 1.5 63.1 62.1
Sound limb Stiff
(4.4)*L (3.3)*L (2.1)*L (11.1)S (12.9) (10.2) (6.6) (12.1) (5.3)* (7.4)* (4.7)* (5.9)
15.0 7.6 8.2 39.6 43.3 7.1 65.2 58.0 61.8 1.2 62.4 61.9
Compliant (2.8)*L (2.2)*L (2.2)*L (12.2)*CN (13.6) (10.4) (4.0) (9.7) (5.0)* (7.3)* (4.4)* (4.9)
22.8 (2.9) 19.2 (5.8) 41.9 (6.7) 31.0 (8.7) 36.8 (8.0) 0.32 (5.4) 60.9 (4.7) 61.2 (6.5) 59.3 (4.9)* 0.7 (7.8) 60.3 (5.1)* 61.8 (7.5)
Nominal 24.7 17.1 41.7 34.4 40.2 1.8 62.9 61.1 61.6 0.1 61.0 62.1
(4.8) (7.2) (6.7)* (8.0) (8.0) (4.0) (5.3) (3.2) (3.2)* (7.6) (4.3)* (6.1)
Stiff 23.8 18.8 42.6 34.1 39.6 0.9 61.6 60.7 59.9 1.9 60.0 62.8
(3.6) (6.4) (5.8)* (9.5) (8.5) (5.6) (6.0) (7.0) (4.1)* (8.0) (4.7)* (7.4)
GRF (BW) Peak A/P – braking Peak A/P – propulsive Peak vertical – early stance Peak vertical – late stance Peak M/L
0.15 0.17 1.05 1.15 0.05
(0.03) (0.03) (0.06) (0.09) (0.01)
0.17 0.14 1.03 1.08 0.07
(0.05)L (0.07)L (0.08)L (0.13)L (0.01)*
0.18 0.16 1.05 1.10 0.08
(0.05) (0.06)L (0.06)L (0.13)L (0.02)*
0.18 0.15 1.06 1.11 0.07
(0.04)L (0.06)L (0.05)L (0.14)L (0.01)*
0.14 0.19 1.22 1.23 0.07
(0.04) (0.05) (0.10)*NS (0.12) (0.02)*NS
0.16 0.20 1.30 1.25 0.08
(0.04) (0.05) (0.11)*C (0.11) (0.02)*C
0.15 0.20 1.28 1.24 0.08
(0.04) (0.05) (0.09)*C (0.11) (0.02)*C
Moments (Nm kg 1) Peak DF – loading response Peak PF – terminal stance Peak knee ext. – loading response Peak knee flex. – terminal stance Peak hip ext. – stance Peak hip flex. – terminal stance
0.09 1.73 0.87 0.54 1.26 0.61
(0.05) (0.15) (0.32) (0.15) (0.26) (0.23)
0.03 1.76 0.19 0.72 1.49 0.55
(0.14) (0.41) (0.31)*L (0.30) (0.29) (0.18)
0.02 1.82 0.29 0.63 1.58 0.55
(0.11) (0.34) (0.38)*L (0.38) (0.34)* (0.14)L
0.01 1.81 0.38 0.59 1.52 0.56
(0.12) (0.40) (0.38)*L (0.35) (0.33) (0.22)
0.08 1.78 0.86 0.71 1.63 0.49
(0.06) (0.25) (0.39) (0.24) (0.38)* (0.20)
0.06 1.85 1.02 0.72 1.86 0.45
(0.08) (0.28) 0.46) (0.20) (0.47)* (0.20)
0.06 1.83 0.92 0.74 1.71 0.48
(0.08) (0.22) (0.33) (0.23) (0.35)* (0.15)
Powers (W kg 1) Ankle abs. – loading response Ankle abs. – terminal stance Peak ankle gen. – pre-swing Peak knee gen. – early mid stance Peak hip gen. – loading response Peak hip abs. – terminal stance
0.21 0.31 3.67 1.19 1.74 0.58
(0.21) (0.42) (0.83) (0.57) (0.24) (0.30)
0.86 0.80 1.23 0.47 2.51 0.50
(0.78)* (0.35)* (0.50)*L (0.56)*L (0.47)* (0.33)
0.87 0.71 1.19 0.56 2.82 0.58
(0.38)* (0.36)* (0.28)*L (0.58)*L (0.73)* (0.42)
0.72 0.69 1.20 0.69 2.68 0.53
(0.42)* (0.38) (0.43)*L (0.51)L (0.63)* (0.27)
0.66 0.74 4.02 1.29 2.41 0.40
(1.22) (0.49) (1.03) (0.62) (0.83)* (0.35)
0.69 0.67 4.05 1.51 2.70 0.36
(0.94) (0.48) (0.98) (0.77) (0.65)* (0.35)
0.87 0.56 4.04 1.35 2.39 0.38
(1.03) (0.33) (0.94) (0.53) (0.74)* (0.24)
* C N S L
Indicates significantly different from controls. Indicates significantly different from compliant strut condition (within limb). Indicates significantly different from nominal strut condition (within limb). Indicates significantly different from stiff strut condition (within limb). Indicates significantly different from sound limb (within strut condition).
3.3. GRF (Table 2) A significant interaction was observed for the first peak in the vertical GRF (P = 0.008) such that the nominal and stiff struts resulted in greater forces than the compliant in the sound limb, but not the IDEO limb. Overall, the sound limb experienced greater peak forces than the IDEO limb (P < 0.001) and the control group (P < 0.002). IDEO users experienced other inter-limb differences, as well. Both the first (P < 0.001) and second (P = 0.022) peaks in the vertical force and the peak propulsive force (P = 0.001) were greater on the sound side compared to the IDEO side. On the sound side, the first peak vertical GRF and the peak medial GRF were lowest in the compliant strut (P < 0.030). 3.4. Moments (Fig. 3, Table 2) Peak net internal joint moments differed between patients and controls, and between IDEO and sound limbs, but were not significantly different among strut conditions. Internal ankle plantarflexor moments exhibited a distinct local minima in mid stance that was not found in the control group, thus no statistical comparisons were made, and were not significantly different between groups in late stance. There was a significant interaction between stiffness condition and limb in the peak knee extensor moment (P = 0.043) such that it was lower in patients’ IDEO limb compared to sound limb (P < 0.001), but post hoc tests revealed no statistically significant differences among struts. Comparisons to controls indicated significantly lower peak knee extensor moments in the IDEO limb (P < 0.004) and greater hip extensor moments in the sound limb (P < 0.047) of the patient group.
3.5. Powers (Fig. 3, Table 2) The patients’ IDEO limb exhibited significantly less peak ankle power generation compared to both the control group (P < 0.001) and the sound limb (P < 0.001). Peak knee power generation of the IDEO limb during mid-stance was less than that
of the control group while using the compliant (P = 0.008) and nominal struts (P = 0.020), but not the stiff strut. Between limb comparisons indicated that the peak knee power generation during mid-stance was less in the IDEO limb than sound limb (P < 0.003). There was also a bi-lateral increase in peak hip power generation in the patient group, compared to controls for all strut conditions (P < 0.050).
4. Discussion 4.1. Comparisons to controls Use of the custom IDEO enabled limb salvage patients to successfully complete an uphill walking protocol at equivalent self-selected speeds and stride lengths to able-bodied individuals. The IDEO users accomplished the uphill walking task despite deviations in their lower extremity mechanics. Consistent with existing level ground walking literature [7,12,20,26], AFO use limited ankle ROM, which resulted in 77% less ankle power generation at push-off. Previous investigators have found between 42% [26] and 80% [8] less power generation than controls while using carbon fiber AFOs. Limited ankle dorsiflexion during stance due to the semi-rigid nature of the IDEO led to decreased ipsilateral knee extensor moment (all strut conditions) and peak power generation (compliant and nominal struts). The bilateral increase in peak hip power generation during early stance in the patient group compared to control subjects was likely a compensatory strategy
Please cite this article in press as: Haight DJ, et al. Biomechanics of uphill walking using custom ankle-foot orthoses of three different stiffnesses. Gait Posture (2015), http://dx.doi.org/10.1016/j.gaitpost.2015.01.001
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GAIPOS-4385; No. of Pages 7 D.J. Haight et al. / Gait & Posture xxx (2015) xxx–xxx
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[5] for less ankle and knee power generation while using the IDEO. Although research on uphill walking with an AFO is very limited, the observed hip strategy, characterized by greater hip flexion and power generation, has been described during over-ground walking with a conceptually similar AFO [26]. The adaptations at proximal joints may be influenced by changes in AFO design. Changes such as increasing motion by using an articulated ankle joint or other approaches to increase compliance or ankle power may preclude the need for greater hip flexion and power generation in the AFO users. Further research is needed to draw firm conclusions. 4.2. Comparisons between limbs While it is recognized that the IDEO improves functional capacity to a greater extent than other AFOs in this population [11], it is not realistic to expect symmetrical gait, given the restrictions in ankle ROM. Ankle plantarflexion plays an important role at push-off for ankle power generation and ankle dorsiflexion is key for aligning the limb to utilize ipsilateral knee power. It is possible that new powered AFOs could address power generation asymmetries at the ankle, although, compensatory hip strategies and asymmetries are likely to persist unless a full ankle ROM is achieved. However, the IDEO was created for patients with a variety of injuries and surgical procedures, including ankle fracture or fusion, and as a result large excursions can be contraindicated. Therefore, sacrifices in normative biomechanics may be necessary in order to improve overall functional capacity. 4.3. Comparisons among struts Optimal AFO stiffness has been previously studied and is thought to play an important role in the optimization of walking [18,27]. However, modulating IDEO strut stiffness by up to 40% in the present study yielded few significant biomechanical effects during uphill walking. These results are consistent with our previously published results from overground walking [12,20] but it was expected that due to the greater contribution of the plantarflexors and the greater dorsiflexion range of motion during uphill walking lower extremity mechanics would be affected to a greater extent by strut stiffness. The primary kinematic difference among strut stiffnesses was that the knee of the IDEO limb contacted the ground in a more extended position in the two more compliant struts. This finding is consistent with previous overground research [12,28], where the authors showed that early stance knee angles increased with greater resistance to ankle plantarflexion, as would be found with stiffer struts. 4.4. Applications Patients with lower limb salvage who used the IDEO were able to accomplish uphill walking despite substantial plantar flexor weakness. AFOs of a similar design may assist other populations with lower-limb weakness caused by stroke, brain injury, or other neuropathies to achieve functional improvements. The limited effects of stiffness suggest that a range of clinically prescribed IDEO stiffnesses could adequately address the biomechanical requirements of uphill walking. It is possible that the effects of modulating stiffness may have been greater if the initial prescribed strut was more compliant. However, a stiffer strut may better transition to high impact activities without negatively affecting the mechanics of lower impact walking activities. Additional research on high intensity activities is needed to determine if clinically prescribed IDEO stiffness is robust across a variety of activities or if an optimal stiffness becomes more important at higher intensities.
4.5. Limitations Not all patients were able to successfully complete a gait analysis without their IDEO; hence, there is no comparison to walking without an AFO. The authors expect that IDEO use improved gait, but are not able to quantify the extent of its benefit or delineate explicitly if certain gait deviations were from injuries/ plantarflexor weakness or a result of using the IDEO. Furthermore, the combat-related injuries sustained by the limb salvage patients were diverse and subjects were at varying stages of rehabilitation at the time of their participation. However, self-selected walking velocity has been used as an indicator of recovery following lower limb trauma [29] and the equivalent velocities between the two groups may indicate that patients reached a level of recovery where relevant comparisons could be made. 5. Conclusions Use of the passive-dynamic IDEO enabled subjects with substantial plantarflexor weakness to walk uphill at speeds and stride lengths equivalent to able-bodies controls. However, due to limitations in ankle ROM, compensations at proximal joints were observed and asymmetries persisted. Modulating IDEO stiffness had few effects on the biomechanics of uphill walking; however future study is needed to determine if this is robust across a wider range of activities. Acknowledgements The authors acknowledge Nicole Harper and Dr. Richard Neptune at the University of Texas for strut manufacturing and testing. Support was provided by the Center for Rehabilitation Sciences Research Department of Physical Medicine and Rehabilitation, Uniformed Services University of Health Sciences, Bethesda, MD. Conflict of interest The authors, Derek J. Haight, MS, Elizabeth Russell Esposito, PhD, and Jason M Wilken, PhD, MPT, declare that they have no competing interests that would affect interpretation of the data. References [1] Franceschini M, Massucci M, Ferrari L, Agosti M, Paroli C. Effects of an anklefoot orthosis on spatiotemporal parameters and energy cost of hemiparetic gait. Clin Rehabil 2003;17:368–72. [2] de Wit DCM, Buurke JH, Nijlant JMM, Ijzerman MJ, Hermens HJ. The effect of an ankle-foot orthosis on walking ability in chronic stroke patients: a randomized controlled trial. Clin Rehabil 2004;18:550–7. [3] Danielsson A, Sunnerhagen KS. Energy expenditure in stroke subjects walking with a carbon composite ankle foot orthosis. J Rehabil Med 2004;36:165–8. [4] Neptune RR, Kautz SA, Zajac FE. Contributions of the individual ankle plantar flexors to support, forward progression and swing initiation during walking. J Biomech 2001;34:1387–98. [5] Nadeau S, Gravel D, Arsenault AB, Bourbonnais D. Plantarflexor weakness as a limiting factor of gait speed in stroke subjects and the compensating role of hip flexors. Clin Biochem 1999;14:125–35. [6] Esquenazi A, Ofluoglu D, Hirai B, Kim S. The effect of an ankle-foot orthosis on temporal spatial parameters and asymmetry of gait in hemiparetic patients. PM R 2009;1:1014–8. [7] Bregman DJJ, Harlaar J, Meskers CGM, de Groot V. Spring-like ankle foot orthoses reduce the energy cost of walking by taking over ankle work. Gait Posture 2012;35:148–53. [8] Wolf SI, Alimusaj M, Rettig O, Do¨derlein L. Dynamic assist by carbon fiber spring AFOs for patients with myelomeningocele. Gait Posture 2008;28:175–7. [9] Patzkowski JC, Blanck RV, Owens JG, Wilken JM, Blair JA, Hsu JR. Can an anklefoot orthosis change hearts and minds. J Surg Orthop Adv 2011;20:8–18. [10] Bedigrew KM, Patzkowski JC, Wilken JM, Owens JG, Blanck RV, Stinner DJ, et al. Can an integrated orthotic and rehabilitation program decrease pain and improve function after lower extremity trauma? Clin Orthop Relat Res 2014;472:3017–25.
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