Amanda Wach Department of Biomedical Engineering, Marquette University, Olin Engineering Center, Room 206, 1515 W. Wisconsin Avenue, Milwaukee, WI 53233 e-mail:
[email protected]
Linda McGrady Department of Orthopaedic Surgery, Medical College of Wisconsin, Milwaukee, WI 53226
Mei Wang Department of Orthopaedic Surgery, Medical College of Wisconsin, Milwaukee, WI 53226
Barbara Silver-Thorn Department of Biomedical Engineering, Marquette University, Milwaukee, WI 53233
Assessment of Mechanical Characteristics of Ankle-Foot Orthoses Recent designs of ankle-foot orthoses (AFOs) have been influenced by the increasing demand for higher function from active individuals. The biomechanical function of the individual and device is dependent upon the underlying mechanical characteristics of the AFO. Prior mechanical testing of AFOs has primarily focused on rotational stiffness to provide insight into expected functional outcomes; mechanical characteristics pertaining to energy storage and release have not yet been investigated. A pseudostatic bench testing method is introduced to characterize compressive stiffness, device deflection, and motion of solid-ankle, anterior floor reaction, posterior leaf spring, and the intrepid dynamic exoskeletal orthosis (IDEO) AFOs. Each of these four AFOs, donned over a surrogate limb, were compressively loaded at different joint angles to simulate the footshank orientation during various subphases of stance. In addition to force–displacement measurements, deflection of each AFO strut and rotation of proximal and supramalleolar segments were analyzed. Although similar compressive stiffness values were observed for AFOs designed to reduce ankle motion, the corresponding strut deflection profile differed based on the respective fabrication material. For example, strut deflection of carbonfiber AFOs resembled column buckling. Expanded clinical test protocols to include quantification of AFO deflection and rotation during subject use may provide additional insight into design and material effects on performance and functional outcomes, such as energy storage and release. [DOI: 10.1115/1.4039816] Keywords: ankle-foot orthoses, kinematics, stiffness, mechanical testing
Introduction With recent advancements in surgical techniques, individuals with traumatic injuries to the lower leg and ankle-foot complex may consider limb salvation as an alternative to amputation [1,2]. Consequently, the patient population requiring ankle-foot orthoses (AFOs), traditionally geriatric individuals and those with motorcontrol impairments, is broadening to include younger, more active patients eager to return to advanced functional performance [3]. The needs of this changing patient population have led to advancements in AFOs and alternative fabrication materials to facilitate advanced functional performance [3]. Carbon fiber composite materials have been integrated into thermoplastic AFOs and used as the primary fabrication material in both conventional and novel AFOs [4,5]. While the use of traditional AFOs has demonstrated improved temporal–spatial parameters (e.g., walking speed, cadence, step and stride length, energy cost, and oxygen consumption) [3,6–10], these innovative carbon fiber AFOs have facilitated return to advanced functional performance (e.g., running, jumping, and lifting) [9,11–14]. The impact of AFO modifications on functional performance can be informed by exploration of the biomechanical behavior of the AFO-limb complex in response to loading conditions that simulate activities of daily living [15]. The previous studies have focused mechanical testing of AFOs to the characterization of rotational stiffness, the resistance to ankle rotation in the sagittal plane [16]. However, as patients’ expectations of AFOs expand to include advanced functional performance, additional mechanical characteristics should be considered. Compressive stiffness, a mechanical characteristic representing an AFO’s ability to resist compressive loads or displacements, has also been investigated [17]. This measure supports classification and comparison of
Manuscript received July 7, 2017; final manuscript received March 5, 2018; published online April 30, 2018. Assoc. Editor: Brian D. Stemper.
Journal of Biomechanical Engineering
AFOs and may have relevance with regard to energy storage in various AFOs. Additional mechanical characteristics studied in conjunction with energy storage and release in leaf spring AFOs and prosthetic feet, such as device deflection [18], may provide insight into the behavior of AFOs designed for advanced functional performance. Device deflection describes changes in AFO geometry as a result of application of load. The aim of this study was to characterize the compressive stiffness and device deflection of loaded AFO-limb complexes. We hypothesize that AFOs prescribed for return to advanced functional performance will demonstrate unique mechanical characteristics compared to traditional AFOs, specifically increased compressive stiffness and device deflection. Exploration of AFO mechanical characteristics will provide further insight into the biomechanical behavior of AFO-limb complexes and the effects of AFO design structure and material composition. Improved understanding of AFOs can inform device prescription, design refinement, and alignment to better address patient-specific needs.
Methods Mechanical testing was performed on four AFOs that differ in structure and/or materials. The tested AFOs included two thermoplastic designs (solid-ankle AFO and anterior floor reaction AFO) intended to facilitate walking and community ambulation, and two carbon fiber designs (PhatBrace dynamic response AFO and the intrepid dynamic exoskeletal orthosis (IDEO)) prescribed for advanced functional performance, specifically jogging and return to sport, respectively (Fig. 1). These AFOs were included to provide a variety in design structure and materials, particularly with respect to the strut, and patient functional performance. The bench testing model incorporated a surrogate limb model to act as a consistent internal support structure and load distribution to the AFOs during testing. This surrogate limb model
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Fig. 1 Study AFOs: (a) solid-ankle AFO, (b) anterior floor reaction AFO, (c) PhatBrace AFO, and (d) IDEO
simulated clinical load application for distribution across the AFO-limb interfaces, minimizing deflection artifacts that might be observed if the AFO alone was loaded. The internal structure of the surrogate limb model was composed of an aluminum pylon (TruLife; Hannover, Germany), covered by a stiff foam (SPS National Labs; Tempe, AZ); this foam (formed from the subject cast) was fixed around the pylon to mimic the soft tissue bulk of the lower leg. A gel liner (Alps South LLC; St. Petersburg, FL) was placed over the foam to simulate the viscoelastic skin structures. A single-axis prosthetic foot (Ohio Willow Wood Co.; Sterling, OH) was included in the surrogate limb model, permitting ankle flexion/extension with minimal hindfoot inversion/eversion and modest forefoot flexion/extension. Custom AFOs were fabricated to fit a female subject (27 yr, 173 cm, 80 kg) who exhibited plantar-/dorsiflexor weakness for which an AFO was prescribed. The affected limb of a this subject was casted to fabricate the custom solid-ankle and anterior floor reaction AFOs, as well as a surrogate limb model used during bench testing. The PhatBrace dynamic response AFO was fabricated by Bio-Mechanical Composites (Des Moines, IA) based on the anthropometry of the same subject. An independent certified orthotist fabricated the thermoplastic AFOs; the IDEO was fabricated by a second independent certified orthotist, the original designer of the IDEO, using an alternative cast of the same subject, as per IDEO fabrication guidelines [3]. Mechanical testing was conducted on two separate material testing systems. A MTS 809 axial/torsional load frame was used to quantify AFO posterior strut deflection; the MTS criterion was used to assess the compressive stiffness and motion of the proximal and supramalleolar regions of the AFOs. Mechanical tests
Fig. 2 Mechanical testing setup for (a) phase 1: strut deflection with posterior strut markers and (b) phase 2: AFO proximal and supramalleolar regional motion assessment with active markers (white). These two test configurations also illustrate the adjustable loading plate to simulate the different limb orientations corresponding to the various stance subphases.
were conducted using materials testing systems, MTS 809 axial/ torsional load frame and MTS criterion (MTS Systems Co.; Eden Prairie, MN) for the first and second phases of testing, respectively, with a 5 kN load cell. To approximate AFO-limb complex loading during gait in both phases of testing, vertical load magnitudes and AFO-limb complex orientation conditions corresponding to discrete instances during midstance (MSt), terminal stance (TSt), and preswing (PSw) for an able-bodied subject (age- and weight-matched to the AFO model subject) during level over ground walking were applied (Table 1). An adjustable loading plate fixture (Fig. 2) was used to vary the shank orientation and approximate the center of pressure during these respective stance subphases. A 66.7 N (15 lbf) preload was applied to ensure proper seating of the surrogate limb model within the AFO. Two phases of testing were conducted for each AFO-limb complex. In the first test phase, the contact point between the AFO footplate and the loading plate fixture was monitored to minimize potential variation in the ankle moment lever arm and introduction of procedural error in ankle moment during strut deflection testing. If the application of the full target load determined based on gait analysis of an age- and weight-matched able-bodied subject (Table 1) resulting in excessive deflection of the AFO plantar surface or full, rather than localized, contact between the AFO plantar surface and the loading plate, the target load was reduced. The reduced load ensured that the AFO/loading plate contact remained localized such that centers of pressure typical of over ground walking were approximated; these reduced applied loads are noted in Table 2. Ten loading/unloading cycles, from the preload to the full/reduced target load, were applied using displacement control at rate of 5 mm/s [17,21]. Three-dimensional motion capture
Table 1 Discrete AFO-loading plate orientations and target loads corresponding to different subphases of stance: MSt, TSt, and PSw during level over ground walking of an age- and weight-matched able-bodied subject [19,20] Stance subphase
Shank to vertical angle (deg)
Vertical force (% body weight)
Target force (N)
Midstance Terminal stance Preswing
5 10 20
110 80 110
880.7 640.5 880.7
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Table 2 Mean peak loads (and standard deviations) applied during partial (gray) and full (black) target loading trials at each tested subphase of stance Maximum load (N)
Solid-ankle Anterior floor reaction IDEO PhatBrace Target
Midstance
Terminal stance
Preswing
328 (62) 387 (63) 923 (62) 154 (61) 880.7
316 (63) 294 (61) 721 (61) 163 (62) 640.5
299 (64) 293 (63) 871 (62) 154 (61) 880.7
(Optotrak Certus, NDI; ON, Canada, accuracy 0.1 mm) was performed during testing. Ankle-foot orthoses strut deflection was characterized using 15 7-mm active markers positioned along the AFO strut at 2 cm increments (Fig. 2(a)). To characterize AFO strut deflection and movement pattern during compressive loading, the magnitude of maximum displacement of the strut markers in the sagittal plane as well as the overall sagittal plane motion of all strut markers with respect to the preloaded position was assessed. The strut active markers, evenly spaced along the AFO length, were divided into three regions such that the location of maximum displacement could be differentiated as occurring in the proximal, mid-, or distal third of the strut. For the solid-ankle and anterior floor reaction AFOs, the shell along the posterior and posterior/anterior shank, respectively, was analyzed as the strut of the design. To evaluate compressive stiffness and regional rotation, the same loading protocol was applied during the second test phase; however, the peak applied load was not reduced from the full target load (Table 1). Under these full loading conditions, the
deformation of the thermoplastic AFOs resulted in full contact of the AFO plantar surface and loading plate. Using motion capture, AFO regional rotation was evaluated using active marker triads rigidly fixed to the AFO at the proximal cuff and supramalleolar regions (Fig. 2(b)). For both test phases, only the latter five data cycles were used for analyses to reduce potential preconditioning effects. All data were sampled at 115 Hz. Data were analyzed using MATLAB (MATLAB R2013b, The MathWorks Inc., Natick, MA). The force–displacement data from the MTS system were synchronized postacquisition with motion data using an active marker secured to the MTS crosshead. The rotations of the various triads were determined using Euler angle analysis. The final 25% of the mean force–displacement curve over the latter five loading cycles were fit via linear regression to determine the slope or compressive stiffness of each AFO.
Results During compressive loading, the resultant force increased nonlinearly with the applied displacement for each AFO and stance subphase. The corresponding compressive stiffness values during the latter portion of the loading protocol are shown in Fig. 3. The solid-ankle and IDEO AFOs demonstrated compressive stiffness 1.6–2.8 times greater than the anterior floor reaction AFO and 7.1–12.5 times greater than PhatBrace design throughout stance, respectively. As shown in Table 3, the IDEO AFO demonstrated greatest peak posterior strut displacement for all subphases of stance tested. Peak strut displacement of the solid-ankle, anterior floor reaction AFO, and PhatBrace displayed the greatest differences from the IDEO during simulated PSw, with percent differences of 84.3, 80.4, and 73.0, respectively. The location of this peak displacement along the strut differed between AFOs, with peak displacement locations observed midstrut for the carbon fiber
Fig. 3 Mean (and SD) compressive stiffness of tested AFOs during latter subphases of stance (top) across the latter five loading cycles. The corresponding regressed force–displacement curves used to determine stiffness for the IDEO are also shown (bottom).
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Table 3 Peak strut displacement relative to the preloaded position in the sagittal plane at various subphases of stance. The location of peak displacement is noted as proximal (P), mid- (M), or distal (D) third of the strut region. Data corresponding to the reduced target loading trials are shown in gray. Maximum strut displacement (mm) Midstance Solid-ankle Anterior floor reaction IDEO PhatBrace
2.10 (60.01) 2.70 (60.01) 9.07 (60.02) 3.43 (60.01)
Terminal stance P M M M
2.80 (60.01) 3.92 (60.01) 8.21 (60.02) 4.73 (60.01)
Preswing P M M M
2.02 (60.01) 2.53 (60.02) 12.88 (60.01) 3.48 (60.01)
P M M M
Fig. 4 Displacement of posterior strut markers in the sagittal plane during strut deflection testing. The strut marker positions during preload (midstance: circles; terminal stance: diamonds; preswing: squares) are shown in the inset figures; the strut displacement at the (reduced—see Table 2) target load for various subphases of stance is shown for each AFO.
AFOs. The posterior strut marker displacements from preload to the (reduced) target load demonstrated different patterns of deflection (Fig. 4). The solid-ankle and anterior floor reaction AFOs demonstrated increased anterior–posterior displacement of distal posterior strut markers relative to the proximal markers. The PhatBrace and IDEO AFOs showed nonlinear displacement along the posterior strut. The mean sagittal rotation of the proximal region of the AFO was greater for the IDEO than the solid-ankle and PhatBrace at all
subphases of stance by 3.5–75 times and 2.6–13 times, respectively (Table 4). The IDEO demonstrated a larger proximal rotation than the anterior floor reaction AFO at MSt only; the percent differences between rotation of the proximal IDEO and anterior floor reaction AFO were 2.7 at TSt and 22.6 at PSw. The sagittal plane rotation of the supramalleolar region was greater for the PhatBrace dynamic response AFO than the IDEO (Table 5), with the largest percent difference, 75.3%, occurring at PSw. Due to out of plane deformation observed in the supramalleolar region at
Table 4 Mean sagittal plane rotation of the proximal region of tested AFOs Proximal regional rotation (deg) of AFO
Solid-ankle Anterior floor reaction IDEO PhatBrace
Midstance
Terminal stance
Preswing
0.03 (60.00) 1.09 (60.01) 2.27 (60.03) 0.85 (60.05)
0.18 (60.01) 1.87 (60.01) 1.82 (60.03) 0.14 (60.07)
0.81 (60.01) 3.65 (60.02) 2.91 (60.01) 0.43 (60.05)
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Table 5 Mean sagittal plane rotation of the supramalleolar region of two AFOs Supramalleolar rotation (deg)
PhatBrace IDEO
Midstance
Terminal stance
Preswing
3.20 (60.04) 3.22 (60.05)
6.83 (60.06) 2.39 (60.04)
18.41 (61.33) 4.55 (60.01)
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the target loads, the sagittal rotation of the supramalleolar region was not considered for the thermoplastic AFOs.
Discussion The aim of this study was to characterize and compare mechanical characteristics of AFOs in clinically relevant loading conditions. The AFOs selected for testing were chosen to represent different fabrication materials, geometries, and clinically observed functional outcomes in an effort to improve understanding of the biomechanical function of AFOs. Both the solid-ankle and the anterior floor reaction AFOs functionally support the foot and ankle in a constant neutral position, resisting ankle motion throughout stance and swing phases of gait. These thermoplastic AFOs are commonly prescribed for patients with motor control and ankle stability concerns to facilitate community ambulation. The PhatBrace dynamic response AFO is a carbon fiber posterior leaf spring design that controls progression during stance and assists with toe off and foot clearance during swing. This AFO can improve balance and reduce energy cost in patients with weak lower leg musculature, helping them return to moderate activity such as prolonged walking. The IDEO (developed by the Center for the Intrepid; San Antonio, TX) is a unique carbon fiber AFO, designed to limit the ankle range of motion with the ability to provide energy storage and release for high-intensity activities. Clinically, use of the IDEO has demonstrated improved functional performance [3,8,9] compared to that observed with thermoplastic and carbon-fiber AFOs. Despite prior investigation of the mechanical behavior of AFOs, a standard test setup has not been established, nor have characteristic mechanical loading parameters been definitively identified. Mechanical testing studies have used various methods to investigate rotational stiffness of AFOs, including functional testing with human subjects as well as bench testing, with and without, various surrogate limb models [17,20–25]. In the current study, bench testing with consistent application of displacement was conducted to allow for a controlled, repeatable rate of displacement and measurement of load. Each AFO-limb complex was mechanically loaded under the same controlled testing conditions, despite potential effects of each design on gait kinematics, to characterize mechanical differences between AFOs. Additional testing and analysis revealed force–displacement measures to be independent of the applied displacement rate; preconditioning was not needed to reduce intercycle variability [19]. The study incorporated motion capture to characterize device deflection through measured strut displacement and AFO regional rotation, which has not been previously considered for AFOs. The compressive stiffness values demonstrated no consistent trend in stiffness with stance subphase. Higher compressive stiffness was observed with the IDEO and solid ankle AFO, indicating similar mechanical characteristics despite fabrication with different materials. While these two AFOs share a common clinical objective, prevention of ankle rotation, the corresponding functional outcomes differ greatly [8]. As such, compressive stiffness does not fully describe the mechanical behavior nor predict functional performance of an AFO. Of the previous studies investigating AFO mechanical properties, only Hawkins quantified the compressive stiffness of a carbon-fiber posterior leaf spring AFO [17]. The compressive stiffness (8–41 N/mm) measured by Hawkins was more compliant than those observed in the current study (64–87 N/mm). These disparities may be attributed to differences in test setup, namely the lack of a surrogate limb, reduced target loads (100–150 N), and slower compressive displacement rate (1.69 mm/min) [17] relative to the current study. These differences in test setup are due to different aims of the mechanical tests: the goal of this study to test AFOs in a controlled setup that closely mimics subphases of stance, requiring clinically based loads, load distribution, and load application rate, as opposed to Hawkin’s aim to test AFO deflection under a large pure compression or bending load. Journal of Biomechanical Engineering
In addition to quantifying AFO compressive stiffness, the present study measured AFO strut deflection and regional AFO rotational motion. To our knowledge, these mechanical characteristics of AFOs have rarely been investigated, though a study by Dyer et al. evaluated similar behaviors of carbon fiber prostheses [26]. The maximum displacement of the posterior strut in the sagittal plane of the IDEO exceeded that for the other AFOs for all latter stance subphases and occurred midstrut. Direct comparison of the magnitude of marker displacements between AFOs cannot, however, be made due to variations in the applied load (e.g., target load was reduced for some AFOs to prevent atypical full contact between the AFO plantar surface and the loading plate). However, qualitative analysis of the peak displacement location, and the resulting deflection mechanism, may provide some insight. The maximum strut displacement of the PhatBrace AFO, like the IDEO, occurred midstrut; both thermoplastic AFOs resulted in peak displacement more proximally. The deflection mechanism of the AFOs, e.g., the displacements along the entire length of the strut relative to the preload position, was also considered. The carbon-fiber AFOs demonstrated a deflection pattern similar to a beam buckling under compressive loading: the displacement of the strut peaked midstrut and decreased proximally and distally in a nonlinear fashion. This deflection mechanism may contribute to energy storage and release during stance, similar to that observed with the carbon fiber shafts of prosthetic feet such as the Flex foot or Cheetah running prosthesis. This beam bucking-like strut deflection pattern suggests sagittal plane rotation, with the center of rotation occurring at the location of minimal motion. Further testing to the full target load magnitude is necessary to confirm these deflection mechanisms and fully characterize the relationship between deflection and potential energy storage. The Euler angular rotations of the proximal and supramalleolar regions were also characterized for the AFOs. The IDEO demonstrated the greatest proximal regional rotation at the full target load for all tested subphases of stance. This finding supports the strut deflection mechanism results, as midstrut buckling could induce proximal rotation of the AFO. The rotation of the more distal supramalleolar region of the PhatBrace was greater than that of the IDEO. This rotation may have contributed to the increased contact of the AFO plantar surface with the loading plate. The lack of rotation and deformation of the IDEO at the supramalleolar region suggests distal rigidity of the device structure, protecting the ankle and subtalar joints and forcing deformation due to compressive loads to occur at the weaker region of the design, the posterior strut. Characterization of compressive stiffness and device deflection in loaded AFOs describes how the entire AFO-limb complex behaves under load, as well as how the strut of the AFO responds to loading. This study was limited in testing AFOs at discrete instances in the gait cycles using pseudostatic approximations of various subphases of stance. These conditions use normal gait as a control and do not consider the effect each AFO design has on patient gait. The AFOs were tested with a surrogate limb model. However, external orthotic modifications such as shoe or heel wedges were excluded; individual-specific AFO loading was, therefore, not replicated. In addition, the AFOs included in this study were customized for a single patient, with the exception of the commercially available PhatBrace. Device customization, such as trimlines of the thermoplastic designs and strut tuning of the IDEO, to accommodate individual needs will affect the mechanical measures. Future mechanical testing protocols incorporating actuation of the loading plate to permit dynamic and continuous testing of patient gait, as well as more rigorous activities such as running, would facilitate investigation of energy storage and release mechanisms in AFOs. The presented study details mechanical testing procedures and comparative results to introduce a mechanical characteristic measure, strut deflection, that may be used in future, related studies to investigate differences in advanced functional performance possible with AFO use. JULY 2018, Vol. 140 / 071007-5
Acknowledgment The support of the Biomedical Engineering Department at Marquette University (MU), the Orthopaedic Rehabilitation Engineering Center of MU and the Medical College of Wisconsin, and Hanger P&O (Milwaukee, WI and Draper, UT) is gratefully acknowledged. The authors would like to thank Philip Voglewede, Ph.D., Thomas Current, C.P.O., Ryan Blanck, C.O., and Dean Jeutter, Ph.D. for their assistance throughout the project. Additionally, the authors thank the Center for Motion Analysis (Greenfield, WI) for the contribution of gait kinetic and kinematic data for a matched able-bodied subject.
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