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Journal of Biomechanical Engineering logoLink to Journal of Biomechanical Engineering
. 2021 Sep 1;144(1):011004. doi: 10.1115/1.4051845

Multiplanar Stiffness of Commercial Carbon Composite Ankle-Foot Orthoses

Benjamin R Shuman 1,✉,c, Elizabeth Russell Esposito 2,
PMCID: PMC8420787  PMID: 34286822

Abstract

The mechanical properties of an ankle-foot orthosis (AFO) can impact how a user's movement is either restricted or augmented by the device. However, standardized methods for assessing stiffness properties of AFOs are lacking, posing a challenge for comparing between devices and across vendors. Therefore, the purpose of this study was to quantify the rotational stiffness of thirteen commercial, nonarticulated, carbon composite ankle-foot orthoses. A custom, instrumented test fixture, for evaluating mechanical properties in rotating exoskeletons (EMPIRE), deflected an AFO through 20 deg of plantar/dorsiflexion motion about a specified, but adjustable, ankle axis. Sagittal, frontal, and transverse plane rotational stiffness were calculated, and reliability was assessed between cycles, sessions, and testers. The EMPIRE demonstrated good-to-excellent reliability between testers, sessions, and cycles (intraclass correlation coefficients all ≥0.95 for sagittal plane stiffness measures). Sagittal plane AFO stiffness ranged from 0.58 N·m/deg to 3.66 N·m/deg. AFOs with a lateral strut demonstrated frontal plane stiffnesses up to 0.71 N·m/deg of eversion while those with a medial strut demonstrated frontal plane stiffnesses up to 0.53 N·m/deg of inversion. Transverse plane stiffnesses were less than 0.30 N·m/deg of internal or external rotation. These results directly compare AFOs of different models and from different manufacturers using consistent methodology and are intended as a resource for clinicians in identifying a device with stiffness properties for individual patients.

Keywords: ankle-foot orthosis, rotational stiffness, reliability testing, ankle axis

Introduction

Ankle foot orthoses (AFOs) are often necessary to overcome a variety of mobility limitations during gait. The mechanical properties of the AFO largely determine the degree to which an individual's movement is either restricted or augmented. [15], and are related to the AFO material, proprietary geometry, and fabrication practices [6]. Broadly, AFOs have been shown to restrict ankle kinematics and alter gait timings, with secondary changes at the knee [7]. It has been suggested that the mechanical properties of an AFO can be matched to an individual's specific impairments and thus should be tailored to the individual [1,8,9]. As such, the reporting of mechanical properties has been suggested as a best practice in studies involving AFO interventions [10] and the importance of understanding the mechanical characteristics of AFOs has received increasing attention in the literature [7,11,12].

Given the interest in AFO design properties, a number of different methods for evaluating the mechanical properties of an AFO have been developed. Some investigators examine linear stiffness [1317] while others examine the rotational stiffness about the ankle measured as the change in resistive torque versus ankle angle [1,9,1822]. However, nonstandardized test methodologies [12,23] and differences in outcome variables [7] pose a challenge for comparing outcomes across the literature. In addition, limited, if any, information is provided on the reliability of many of these testing methods. Those studies which do provide reliability analyses focused on the rotational stiffness of an AFO about a defined axis (as in an ankle joint) and evaluated intracycle variability [1820,24], intersession variability [3,18,19,24,25], and/or intertester variability [18,19,24]. Intracycle, intersession and intertester variability are all important to report in characterizing AFO properties.

The sagittal plane stiffness of an AFO is widely recognized as an important component in the design and prescription of a device. However, many AFOs, including those with a single medial or lateral strut, may, by the nature of their design, also impart coupled frontal and transverse plane resistive moments. Examinations of AFO properties out of the sagittal plane are less common in the literature, vary in how out of plane properties are loaded and measured, and have not been evaluated with a reliability analysis [2633]. These frontal and transverse plane contributions may be important for certain patients [7], including those who experience pain or functional deficits during out of plane rotations such as weakness or ankle instability.

Ideally, the information on AFO mechanical properties would come from manufacturers using uniform methodology. Many manufacturers do provide qualitative information related to the stiffness characteristics of their products rating rigidity, function scales [34], or identifying a list of indications such as drop foot or mild versus moderate muscle plantarflexor weaknesses [35,36]. However, stiffness, specifically, is rarely provided in units that enable comparisons across manufacturers. For example, one manufacturer may offer stiffnesses ranging from 1 to 7 in their own units of measurement [37] while another offers a choice of different colors or names intended to represent different stiffnesses [34]. Quantitative stiffness data is generally lacking for these off-the shelf models.

Many AFOs evaluated in the literature are not off-the-shelf models, but customized to an individual, often made of thermoplastic [2,24,30,32,38], or three-dimensional printed [23,39]. Quantitative evaluations in the literature of commercial devices include how stiffness varies in the Neuroswing (Fior & Gentz, Lüneburg, Germany) [4042] and Tammerak (Tamarack Habilitation Technologies Inc., Blaine, MN) [20,21,28,43] joints, the Carbon Ankle Seven struts (Ottobock, Duderstadt, Germany) [17,44] and varying layups of the Orthotics Composites Helix (Thusane, Levallois-Perret, France) [45]. Dynamic-response custom carbon AFOs, such as the Intrepid Dynamic Exoskeletal Orthosis, have also been evaluated in how posterior strut bending stiffness affects walking, running, and inclined walking mechanics. [1317]. While AFOs with defined hinges or joints have designed axes of rotation (e.g., Neuroswing and Tammerak), many composite AFOs do not, instead deflecting in a strut during motion. The lack of a defined AFO axis of rotation presents a challenge in defining the experimental axis and aligning/testing the devices, highlighting the need for test setups evaluated for reliability.

As the mechanical properties of most common commercially available AFOs have not been evaluated in the literature, the aim of this study was to provide clinicians with a direct comparison of multiplanar stiffness data across different manufacturers. In this study, a mechanical testing device for evaluating mechanical properties in rotating exoskeletons (EMPIRE) is described, and the reliability quantified to measure AFO stiffness about an ankle axis.

Methods

Ankle-Foot Orthoses Tested.

Thirteen commercially available AFOs from four manufacturers were selected for analysis. Criteria for selected AFOs included nonarticulated, nonmodular, and constructed of carbon composite. Their use case was intended to achieve or improve symmetrical gait patterns for activities that include walking and are not intended solely for foot drop. An internal survey of VA Puget Sound orthotists identified Allard's (Helsingborg, Sweden) Blue Rocker, Blue Rocker 2½, ToeOFF, and ToeOFF 2½, Ottobock's (Duderstadt, Germany) WalkOn Reaction and WalkOn Reaction Plus, Thuasne's (Levallois-Perret, France) SpryStep, SpryStep Max and SpryStep Plus, and Trulife's (Dublin, Ireland) Matrix, Matrix Max, Matrix Max2 and Matrix Supermax as the most commonly prescribed off-the-shelf AFOs. All AFOs utilized a single strut between the tibial cuff and footplate, on either the medial (Ottobock models) or lateral (all others) side of the foot. The strut connection to the footplate was anterior to the ankle joint for all but the SpryStep and SpryStep Plus models. The model of each tested AFO was selected to fit the right foot of a hypothetical male with a size U.S. Men's 10½ shoe (275 mm), which corresponded to a size large AFO in all cases.

Fixture Design.

A test fixture, the EMPIRE, was designed to evaluate rotational stiffness about a defined ankle axis. The EMPIRE was conceptually similar to the fixture previously described by Ielapi et al. [19] (Fig. 1). Engineering drawings can be found in the Supplemental Materials on the ASME Digital Collection. The footplate was securely clamped to the base of the fixture using two screws. A contoured aluminum plate with small rubber pads was placed between the screws and the AFO to apply uniform loading and constrain deflection to the strut of the AFO. A thin rubber pad was placed underneath the footplate to minimize slippage during clamping. The tibial cuff of the AFO was strapped around a surrogate shank consisting of a 115 mm diameter, 25 mm thick disk with a rubber interface that is height-adjustable and designed to minimize the relative movement at the AFO interface (similar to the movement constraints in Ref. [21]). This differs from some prior test fixtures [18,19,24], which allow the AFO to rotate and translate along the surrogate shank axis.

Fig. 1.

Overview of the EMPIRE experimental testing fixture (1). An AFO (2) is mounted by securing the footplate with a clamping block (3) consisting of screws and a plate and strapping the tibial cuff to a surrogate shank (4). A linear actuator (5) deflects the AFO by rotating a frame (6) about a specified ankle axis (7). The position of the ankle axis can be adjusted using alignment brackets (8). Angular position is measured by an encoder at the ankle axis and loads through a load cell (9) in between the surrogate shank and the rotating frame. The base plate (10) has marked lines (representative line highlighted, (11) to aid in aligning the two medial points of the AFO.

Overview of the EMPIRE experimental testing fixture (1). An AFO (2) is mounted by securing the footplate with a clamping block (3) consisting of screws and a plate and strapping the tibial cuff to a surrogate shank (4). A linear actuator (5) deflects the AFO by rotating a frame (6) about a specified ankle axis (7). The position of the ankle axis can be adjusted using alignment brackets (8). Angular position is measured by an encoder at the ankle axis and loads through a load cell (9) in between the surrogate shank and the rotating frame. The base plate (10) has marked lines (representative line highlighted, (11) to aid in aligning the two medial points of the AFO.

The EMPIRE was designed to rotate an AFO about a simulated ankle joint, from 0 up to 25 degrees of dorsiflexion. An analog magnetic angle encoder (model RM08A, Renishaw, West Dundee, IL) measured angular deflection at the point of rotation. Loads were applied to a rotating frame using a linear actuator with a 406 mm stroke (model PA04, Progressive Automations, Arlington, WA). The actuator operated at an unloaded speed of 10 mm/sec which imparted a rotational speed of roughly 0.75 deg/sec. A six-axis load cell (Mini 58 SI-2800-120, ATI, Apex, NC) measured the resistive moments of the AFO 616 mm above the rotation axis. Ankle axis position could be adjusted vertically by adjusting the location of the rotational frame. For the tests presented here, the surrogate shank was adjusted such that loads were applied to the AFO 273 mm above the rotation axis.

Angular position and load cell data were captured simultaneously at 100 Hz using a custom interface in LabVIEW (National Instruments, Austin, TX). Angular position was controlled using preset positions in the linear actuator. To avoid overloading the load cell, positions and loads were monitored in real-time and the actuator could be paused at any time with input from the operator.

Ankle-Axis Definition.

We defined a set of generic ankle axis locations (Appendix A) as being located halfway between the positions of the medial and lateral malleoli measured from the heel [18,46]. Axis locations were computed for foot lengths corresponding to U.S. standard shoe sizes [47]. Linear regressions relating overall foot length with malleolar positions were computed from anthropometry data [48,49] and computerized tomography (CT) data previously collected from 21 healthy individuals (11 male, age: 61.5 ± 7.2 years, height: 1740 ± 96 mm, weight: 80.5 ± 13.2 kg) under 25% bodyweight loading. Sagittal plane horizontal (relative to heel) and vertical (relative to floor) positioning of the malleoli was taken as the central point in the first sagittal slice where the medial/lateral malleoli appeared and foot length was taken as the first frontal slice containing the hallux. For a size 10 ½ shoe the average vertical position of the ankle axis was 81 mm.

Test Procedures and Alignment.

The EMPIRE was first calibrated by moving through the range of motion without an AFO to account for the gravitational loading of the fixturing hardware for a set distance between the load cell and the surrogate shank mounting disk. Prior to mounting an AFO, the surrogate shank was oriented vertically. Padding on the AFO was removed to minimize movement between the AFO and the fixture. The AFO was aligned such that the two most medial prominences of the AFO footplate was parallel with the path of deflection. As the AFOs tested did not have a defined heel cup and many footplates are designed to be trimmed to fit an individual's shoe, the fore aft position was determined as the point at which the tibial cuff of the AFO contacted the mounting disk on the surrogate shank. The AFO's footplate was securely clamped using the aluminum plate and clamping screws described above. The mounting disk/surrogate shank was secured within the AFO's tibial cuff using Velcro.

To determine the repeatability of the measurements in the EMPIRE, each AFO was evaluated by two testers on two separate sessions. Prior to any testing, both testers were oriented to the fixture and instructed in the alignment and testing procedure. Each session first deflected the AFO to 3 deg of plantarflexion and then through three complete cycles moving between 3 deg of plantarflexion and 20 deg of dorsiflexion without removing the AFO from the fixture. The order in which AFOs were tested was randomized for each tester and session.

Data Analysis.

Data analyses were performed using custom scripts in matlab (MathWorks, Natick, MA). Load cell and angular position data were filtered using a 1 Hz low-pass filter (fourth order, zero-lag, Butterworth). To account for the gravitational loads imposed by the fixture, a second-order polynomial fit was computed for the calibration (no AFO) trials between the filtered angular position and each load cell channel. These calibration curves were used to adjust the filtered load cell data from the AFO testing sessions at each angular position. Computation of the AFO resistive moment in the sagittal, frontal, and transverse planes were computed per Appendix B. The first loading cycle was treated as a preconditioning cycle and removed from further analysis. AFO sagittal, frontal, and transverse stiffness, measured in N·m/deg, were computed for the second and third cycles from a linear fit of the AFO moments relative to dorsiflexion angle during the loading phase from 0 to 18 deg of dorsiflexion. The range of motion from 3 to 0 deg of plantarflexion and from 18 to 20 deg of dorsiflexion was not fit to avoid periods when the actuator was accelerating/decelerating. To account for nonlinearities in the AFO moment versus dorsiflexion angle curve, we also report the sagittal, frontal, and transverse plane moments at 10 and 15 deg.

Statistical Analyses.

Each AFO was tested twice by two separate operators (four sessions total). Descriptive statistics include the calculation of the mean and range of AFO stiffnesses. The intrasession, intersession, and interrater reliability of the AFO stiffnesses during loading were assessed with interclass correlation coefficients (ICC's) using a one and a two-way mixed model for absolute agreement (ICC (A,1) and (A,K)) for single and mean scores [50,51]. Within session reliability was calculated using the second and third cycle from all four sessions for each AFO. Between session reliability was calculated as the average stiffness from the second and third cycle of each session for both operators. Between operator reliability was calculated as the average stiffness from the second and third cycle of each session including both days. The standard error of measurement (SEM) was estimated using the square root of the mean square error [51]. The minimum detectable difference (MDD) was calculated from the SEM as MDD=SEM×1.96×2 [51].

Results

Intrasession, intersession, and interrater reliability of AFO stiffness were good to excellent [50] with all ICC's ≥ 0.95, ≥0.86, and ≥ 0.95 for sagittal, frontal and transverse planes, respectively (Table 1). The MDD was 0.53, 0.36, and 0.06 N·m/deg, in the sagittal, frontal, and transverse planes, respectively.

Table 1.

Reliability for AFO stiffness in the sagittal, frontal, and transverse planes

Sagittal stiffness Frontal stiffness Transverse stiffness
Within session ICC (A,1) [95% CI] 1.00 [1.00–1.00] 1.00 [1.00–1.00] 1.00 [1.00–1.00]
ICC (A,k) [95% CI] 1.00 [1.00–1.00] 1.00 [1.00–1.00] 1.00 [1.00–1.00]
SEM (N·m/deg) 0.01 0.01 0.00
MDD (N·m/deg) 0.03 0.03 0.01
Between session ICC (A,1) [95% CI] 0.95 [0.89–0.98] 0.88 [0.75–0.95] 0.95 [0.87–0.98]
ICC (A,k) [95% CI] 0.98 [0.94–0.99] 0.94 [0.86–0.97] 0.97 [0.93–0.99]
SEM (N·m/deg) 0.18 0.11 0.02
MDD (N·m/deg) 0.50 0.31 0.06
Between operator ICC (A,1) [95% CI] 0.95 [0.90–0.98] 0.86 [0.71–0.93] 0.96 [0.90–0.98]
ICC (A,k) [95% CI] 0.98 [0.95–0.99] 0.92 [0.83–0.97] 0.98 [0.95–0.99]
SEM (N·m/deg) 0.19 0.13 0.02
MDD (N·m/deg) 0.53 0.36 0.06

Linear fits of AFO stiffness accounted for greater than 99% of the variance in the sagittal plane moment across all AFOs and sessions. In the frontal and transverse planes, linear fits of the AFO stiffness accounted for an average of 99% and 95% of the variance in moments (minimum (92% and 53%, respectively). A representative AFO moment versus dorsiflexion angle is shown in Fig. 2. Measured AFO stiffnesses in the sagittal plane ranged from 0.58 (Matrix) to 3.66 N·m/deg (Blue Rocker) (Fig. 3, Table 2) with an average measured plantarflexion moment of 7.4–52.2 N·m at 15 degrees of dorsiflexion (Table 3). While all deflection occurred in dorsiflexion, the AFOs generated frontal plane moments (at 15 degrees of dorsiflexion), ranging from 12.3 N·m of eversion (Sprystep Max) to 9.9 N·m of inversion (WalkOn Reaction) and transverse plane moments of 4.1 N·m of internal rotation (Sprystep Plus) to 1.4 N·m of external rotation (WalkOn Reaction).

Fig. 2.

Representative moments versus dorsiflexion angle curves for one session (two cycles shown). Defection in the sagittal plane elicited reaction moments in all three planes. AFO stiffness was taken as the linear fit in each plane between 0 and 18 degrees of dorsiflexion during loading. AFO shown in figure: Allard ToeOff.

Representative moments versus dorsiflexion angle curves for one session (two cycles shown). Defection in the sagittal plane elicited reaction moments in all three planes. AFO stiffness was taken as the linear fit in each plane between 0 and 18 degrees of dorsiflexion during loading. AFO shown in figure: Allard ToeOff.

Fig. 3.

Measured AFO stiffness (mean [range]) across a total of four sessions (two per operator). Sagittal plane stiffness across AFOs varied from 0.58 to 3.66 N·m/deg of dorsiflexion. Most AFOs produced an eversion moment in the frontal plane, while the two AFOs with medial struts produced an inversion moment.

Measured AFO stiffness (mean [range]) across a total of four sessions (two per operator). Sagittal plane stiffness across AFOs varied from 0.58 to 3.66 N·m/deg of dorsiflexion. Most AFOs produced an eversion moment in the frontal plane, while the two AFOs with medial struts produced an inversion moment.

Table 2.

Calculated AFO stiffness across four sessions

Sagittal stiffness Frontal stiffnessa Transverse stiffnessa
(N·m/deg) (N·m/deg) (N·m/deg)
AFO Mean (SD) [Range] Mean (SD) [Range] Mean (SD) [Range]
Blue Rocker 3.66 (0.15) [3.42 to 3.80] −0.27 (0.09) [−0.42 to −0.18] 0.03 (0.02) [0.01 to 0.06]
Blue Rocker 2.5 3.08 (0.10) [2.96 to 3.19] −0.29 (0.07) [−0.41 to −0.21] −0.01 (0.02) [−0.03 to 0.02]
WalkOn Reaction 2.88 (0.14) [2.67 to 3.03] 0.53 (0.05) [0.47 to 0.60] −0.11 (0.01) [−0.12 to −0.10]
WalkOn Reaction Plus 2.75 (0.19) [2.48 to 2.99] 0.52 (0.08) [0.42 to 0.61] −0.10 (0.02) [−0.13 to −0.09]
SpryStep Max 2.70 (0.05) [2.66 to 2.78] −0.24 (0.14) [−0.46 to −0.13] 0.14 (0.02) [0.11 to 0.16]
SpryStep Plus 2.68 (0.30) [2.35 to 3.06] −0.71 (0.27) [−1.07 to −0.35] 0.26 (0.05) [0.21 to 0.34]
ToeOff 2.5 2.02 (0.06) [1.97 to 2.14] −0.23 (0.12) [−0.40 to −0.11] 0.04 (0.01) [0.03 to 0.05]
ToeOff 2.5 1.99 (0.06) [1.93 to 2.09] −0.01 (0.01) [−0.03 to 0.00] 0.03 (0.01) [0.02 to 0.04]
Matrix SuperMax 1.69 (0.14) [1.50 to 1.85] −0.22 (0.15) [−0.46 to −0.06] −0.05 (0.02) [−0.07 to −0.03]
SpryStep 1.51 (0.12) [1.35 to 1.65] −0.37 (0.11) [−0.50 to −0.21] 0.02 (0.02) [−0.02 to 0.04]
Matrix Max 2 1.35 (0.18) [1.24 to 1.64] −0.17 (0.06) [−0.22 to −0.11] −0.06 (0.02) [−0.09 to −0.04]
Matrix Max 1.34 (0.33) [1.14 to 1.88] −0.13 (0.04) [−0.18 to −0.09] −0.04 (0.01) [−0.06 to −0.02]
Matrix 0.58 (0.08) [0.53 to 0.71] −0.19 (0.07) [−0.30 to −0.12] −0.01 (0.01) [−0.02 to −0.01]

aNegative stiffness values in the frontal and transverse plane represent eversion and external rotation.

Table 3.

Calculated AFO moments at 10 and 15 deg of dorsiflexion

Sagittal moment (N·m) Frontal moment (N·m)a Transverse moment (N·m)b
10 deg 15 deg 10 deg 15 deg 10 deg 15 deg
Blue Rocker 32.1 52.2 −9.4 −11.2 0.4 0.4
Blue Rocker 2.5 25.8 42.6 −3.7 −5.7 0.1 0.2
WalkOn Reaction 25.6 40.9 6.6 9.9 −0.9 −1.4
WalkOn Reaction Plus 25.5 39.7 7.0 9.8 −0.8 −1.4
SpryStep Max 21.6 35.5 −9.1 −12.3 0.3 0.8
SpryStep Plus 23.5 36.6 −3.9 −7.9 2.4 4.1
ToeOff 16.8 27.5 −3.4 −5.4 0.2 0.3
ToeOff 2.5 18.4 29.6 −1.4 −0.1 0.1 0.5
Matrix SuperMax 10.5 19.8 −6.5 −8.1 −0.3 −0.8
SpryStep 12.4 19.4 −2.3 −4.2 0.2 0.3
Matrix Max 2 9.1 16.4 −4.1 −5.4 −0.2 −0.6
Matrix Max 9.7 16.8 −3.3 −4.4 −0.4 −0.7
Matrix 4.2 7.4 −2.5 −3.9 −0.3 −0.4
a

Negative stiffness values in the frontal and transverse plane represents eversion

b

Negative stiffness values in the frontal and transverse plane represents external rotation.

Discussion

This study evaluated the rotational stiffness about an ankle axis of commercially available AFOs across models and manufacturers. With only a few exceptions, the stiffnesses measured in this work are largely in line with the progressive order expected from manufacturer provided information and a prior comparison of AFOs [22]. There were, however, two instances where our findings did not align with the manufacturer-provided information. We found that the Thuasne SpryStep Max and SpryStep Plus had similar stiffness values (within 1% of each other), which was unexpected as the SpryStep Max is indicated to be stiffer than the Plus [35]. Similarly, we found comparable stiffness values (within 5% of each other) between the Ottobock WalkOn Reaction and WalkOn Reaction Plus, despite the manufacturer indicating that the WalkOn Reaction Plus is stiffer. We also note that the newer Blue Rocker 2½ is 16% less stiff than the original Blue Rocker.

There are several likely explanations for any inconsistencies with prior reports. Perhaps the foremost factor is that the methods by which manufacturers evaluate AFOs are not broadly known. As noted in the introduction, testing methods may characterize either linear or rotational stiffness. Our testing occurred on new devices. To what extent the stiffness of the tested AFOs may change with repeated cycling is unclear but prior examinations of residual stiffness in carbon composites suggest that the stiffness may decrease by as much as 5% to 30% [52]. The size of an AFO has previously been shown to impact the rotational stiffness [53] and some models are indicated to have different mechanical properties targeted to the size of a patient [54]. Another factor was that we confined our investigation to a single speed of rotation (0.75 deg/sec). Previous investigations have found either no impact of loading speed on AFO stiffness [25,30,53] or small, but significant differences that are generally less than the minimal detectible difference of the fixture [55]. It is also currently unclear how much individual AFOs may vary between batches and specimens, as the mechanical properties of carbon composites are dependent not only on the geometry and material but also on the process control in many manufacturing steps including layup of the individual carbon fiber sheets [6].

In addition, this study constrained deflection of the AFO purely to the sagittal plane. Frontal and transverse plane stiffness were calculated from the out-of-plane moments developed while moving through the range of sagittal motion similar to the methods described by Klassen et al. [33] and Singerman et al. [28]. Other methods described in the literature involve deflecting the AFO in multiple planes simultaneously [26,27], or deflecting an AFO through a range of motion in the frontal plane independently [3032]. These different approaches are clearly related but not readily comparable, as independently testing the frontal plane measures a resistance to frontal plane motion, while our methods indicate a preferred direction of coupled loading by the AFO.

Across the AFOs tested, we found that AFOs with a lateral strut tended to exert an eversion moment during dorsiflexion, while the two AFOs with a medial strut tended to exert an inversion moment. This was expected and broadly agrees with the work by Cappa et al. [27], who found that internally wound spiral AFOs (most similar to the lateral strut AFOs tested here) were stiffer in inversion than eversion. Motion of the ankle joint is not purely sagittal, coupling dorsiflexion with eversion [46]. Our study suggests that AFOs with lateral struts may deflect with this coupled motion while medial struts may resist it, but further work is needed to evaluate the impact of strut orientation on patient frontal plane motion during gait.

The design of the commercial AFOs tested in this study required some alignment assumptions for standardization of testing. Features common in AFOs tested in the literature including hinges and heel cups defining the position of the foot were absent. Instead, all included AFOs had a single strut and trimmable foot plates, which allow an off-the-shelf AFO to be easily customized to an individual. Considering these features, we defined a sagittal orientation by aligning the two most medial prominences of the AFOs footplate parallel with the path of deflection and the fore/aft position by lightly contacting the tibial cuff to the surrogate shank. However, foot plate geometry varied by manufacturer, which may impact the alignment between models. Moreover, the bottom of many of the foot plates were contoured causing the AFO to rotate during clamping. We were able to mitigate the challenges imparted by AFO designs with our alignment protocol, resulting in robust intersession and intertester reliability comparable to previous reports in the sagittal plane (ICC's ≥ 0.95 here versus 1.00 [18], 0.99 [19,24], and 0.97 [3]). To our knowledge, this is the first study to examine the repeatability of out of plane stiffnesses, finding good reliability [50], in the frontal and transverse planes (ICC's ≥ 0.83). This is a critical metric to report when evaluating a test fixture for future use in both clinical and research areas.

There are several limitations to this study. While we were able to test thirteen commercial AFOs, our investigation was by no means comprehensive. Our tests were limited to a single size, specimen, testing speed, and ankle axis location for each model, as described above, and only performed on new AFOs. An important feature to note about the design of the EMPIRE is that we intentionally defined a generic ankle axis location based upon the position of the malleoli in the anthropometric literature. However, there exists considerable variability in ankle axis position between individuals. While a significant impact of ankle axis position on AFO stiffness has been demonstrated in hinged AFOs, [20,21] it remains unclear how the choice of ankle axis location may impact AFO stiffness in the AFOs tested here. Similarly, it is currently unclear how the stiffness of each AFO may change over time with use. We did not evaluate how the footplate may contribute to the overall stiffness of the device [18,45,56], although this contribution would be largely impacted by the choice of shoe and how much the footplate was trimmed or altered during fitting (e.g., the application of wedges or supplemental foot orthotics). The contribution of the footplate may partially explain the inconsistencies in stiffness between our results and prior reports. Finally, while we evaluated repeatability of our fixture, we did not directly compare our findings with any other test fixtures. As such, this study did not seek to establish the validity of the test fixture, but to provide a comparison across commercially available AFOs from a range of manufacturers. A previous cross fixture comparison [24] found some systemic differences in stiffness, and the numerous methodological decisions presented in this study highlight how important it is to rigorously evaluate test methods across the literature in the absence of standardized procedures. Future work will compare outcomes from a range of evaluation methods previously presented in the literature. Finally, this work does not identify clinically meaningful differences in stiffness, as individual responses to AFO's are highly heterogeneous. Thus, our results only provide a tool to augment clinical decision-making, realizing that the clinician must interpret these results within the context of the individual.

Conclusions

We directly compared AFO stiffnesses across 13 commercially available carbon composite AFOs using the EMPIRE, a reliable, custom mechanical testing device. In this study, we evaluated the multiplanar stiffness of thirteen nonarticulated, nonmodular, carbon composite AFOs ranging from 0.58 to 3.66 N·m/deg. This research provides a tool through which clinicians can apply their expertise in identifying appropriate AFOs for their patients.

Supplementary Material

Supplementary Material

Supplementary File

Acknowledgment

We would like to thank G. Eli Kaufman, CPO, for his clinical recommendations and Matt Kindig for his support in the design and engineering of the test fixture. The WalkOn Reaction and WalkOn Reaction Plus AFOs were graciously lent to us by Ottobock. The view(s) expressed herein are those of the author(s) and do not reflect the official policy or position of the U.S. Army Medical Department, the U.S. Army Office of the Surgeon General, the Department of the Army, Department of Defense, Department of Veterans Affairs, or the U.S. Government.

Appendix A

Appendix B

The measured torques at the load cell can be written as a combination of the measured forces multiplied by the load cell moment arm offsets

TX=FZ*DY,LCFY*DZ,LCTY=FX*DZ,LCFZ*DX,LCTZ=FY*DX,LCFX*D,LC

Similarly, the torques at the AFO can be written as a combination of the measured forces multiplied by the AFO moment arm offset

TSagittal=FX*DZ,AFOFZ*DX,AFOTFrontal=FZ*DY,AFOFY*DZ,AFOTTransverse=FY*DX,AFOFX*DY,AFO

We defined the AFO origin as being located on the axis of rotation in line with the shank axis (load cell Z). Thus, the moment arms between the load cell and the point of force application relate to the moment arms between the AFO origin through the following, where the height of the load cell is defined by the fixture design and the vertical (z) distance between the load cell and the point of force application is adjustable and measured:

DX,LC=DX,AFODY,LC=DY,AFODZ,LC=HeightLCDZ,AFO

The following equations were then used to convert between the loads measured by the load cell and the moments generated by the ankle:

TSagittal=FX*(HeightLC2*DZ,LC)+TYTFrontal=FY*(HeightLC2*DZ,LC)+TXTTransverse=TZ

Funding Data

This study was funded by the Extremity Trauma and Amputation Center of Excellence (EACE) Rehabilitation Technology Development Fellowship by an appointment to the Department of Defense (DOD) Research Participation Program administered by the Oak Ridge Institute for Science and Education (ORISE) through an interagency agreement between the U.S. Department of Energy (DOE) and the DOD. ORISE is managed by Oak Ridge Associated Universities (ORAU) under DOE Contract No. DE-SC0014664; Funder ID: 10.13039/100006229. All opinions expressed in this paper are the author's and do not necessarily reflect the policies and views of DOD, DOE, or ORAU/ORISE. CT data used to define the ankle axis location was originally funded under National Institute of Health (NIH) Grant No. R21 AR069283; Funder ID: 10.13039/100000002.

Nomenclature

AFO =

ankle-foot orthosis

CT =

computerized tomography

DX,Y,ZLC =

distance between the load cell origin and point of force application

DX,Y,ZAFO =

distance between the center of the AFO axis and point of force application

EMPIRE =

test fixture for evaluating mechanical properties in rotating exoskeletons

FX,FY,FZ =

forces measured in the load cell

HLCAFO =

vertical distance between the load cell origin and the ankle axis

ICC =

interclass correlation coefficient

MDD =

minimum detectable difference

SEM =

standard error of measurement

TSagittal,TFrontal,TTransverse =

torques at the AFO rotation axis

TX,TY,TZ =

torques measured in the load cell

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