Abstract
Background:
Carbon fiber custom dynamic orthoses have been used to improve gait mechanics after lower limb trauma in military service members, with the goal of restoring function and improving outcomes. However, the effects of commercially available carbon fiber orthoses available to civilians on lower extremity joint kinetics and kinematics are poorly understood.
Research question:
The aim of this study was to examine the effect of two commercially available orthoses on lower extremity kinematics and kinetics in individuals with lower limb trauma.
Methods:
A total of 23 participants with a lower extremity traumatic injury underwent gait analysis while walking without an orthosis, and while wearing a monolithic carbon fiber orthosis or while wearing a modular carbon fiber orthosis, in a randomized order. Study participants accommodated to each orthosis for three months prior to testing. Joint kinematics and kinetics at the ankle, knee, and hip joints, and ground reaction forces were assessed.
Results:
The two study orthoses significantly reduced ankle motion compared to no orthosis, with large effect sizes observed. Peak plantarflexor moment was greater with the modular orthosis compared to the monolithic orthosis. Ankle push-off power did not differ between orthoses but was significantly reduced relative to no orthosis. Push-off power with the study orthoses was over 25% greater as compared to previous studies with military orthoses. Peak loading response power generation at the knee was greater with the monolithic orthosis as compared to the modular orthosis. The kinematics and kinetics at the hip did not differ between orthoses.
Significance:
Orthoses commonly used in civilian settings to treat limb trauma primarily alter joint kinematics and kinetics at the ankle, in a manner consistent with orthoses used in the military. Additionally, despite the apparent large differences in the designs of the two study orthoses, between-orthosis differences on gait mechanics were limited.
Keywords: Orthosis, Lower limb trauma, Joint kinematics, Joint kinetics, Gait analysis
1. Introduction
Lower limb traumatic injuries commonly result in poor functional outcomes in military and civilian populations, causing pain, psychological distress, and preventing return to work or active military duty [1], [2], [3]. The Lower Extremity Assessment Project reported that nearly half of patients with lower limb trauma had persistent disability, with only 58% of patients returning to work 7 years post-injury [2]. These results mirrored those found in service members in the Military Extremity Trauma Amputation/Limb Salvage study [1]. Carbon fiber custom dynamic orthoses (CDOs) are a type of ankle foot orthosis (AFO) comprised predominantly of carbon fiber, that include a proximal cuff just below the knee, a posterior carbon fiber strut, and a foot plate. CDOs are provided to individuals who have sustained limb trauma to improve gait, quality of life, and limb function by supporting the limb and reducing pain [4], [5], [6], [7], [8], [9], [10], [11], [12].
The majority of studies evaluating the effect of CDOs on gait mechanics after trauma focus on a single orthosis, and are from a military setting [13]. The primary orthosis, named the Intrepid Dynamic Exoskeletal Orthosis (IDEO), is a very stiff, modular (MOD) CDO, which is well suited for improving function during high impact activities commonly experienced by military service members [14]. The IDEO is far stiffer than other carbon fiber orthoses often used in clinical practice [15], [16], and has been found to improve physical performance and walking speed relative to more commonly used compliant orthoses [6]. Over the past decade, civilian variants of CDOs have become more widely used to restore lower limb function following trauma [17], [18] (e.g. Reaktiv; Fig. 1). The MOD CDOs used in civilian practice are generally less stiff than the IDEO and are often secured using an adjustable cuff. Monolithic CDOs being used in civilian practice (MONO, e.g. Posterior Spring Orthosis; Fig. 1), are more compliant than most MOD CDOs, have a more compliant foot plate, and a more compliant padded proximal cuff [15], [19]. Given the limited studies evaluating these common orthoses, and limb mechanics with the use of CDOs in a civilian population in particular [20], [21], [22], it is unclear if differences in design across CDOs results in similar lower limb mechanics. The lack of information on how these CDOs influence gait limits the ability to effectively decide between different design options.
Fig. 1.
Commercially available monolithic (MONO) (Left) and modular (MOD) (Right) carbon fiber custom dynamic orthoses (CDOs) compared in this study. orthoses.
Changes in orthosis stiffness have been shown to affect gait biomechanics [16], with stiffer orthoses primarily reducing ankle motion and having mixed effects at more proximal joints. Due to the high stiffness of CDOs, their use is associated with significantly lower ankle range of motion and push-off power relative to able-bodied individuals [23], [24]. Varying stiffness levels by as much as 40% in stiff CDOs has a limited effect on lower limb joint mechanics [25], [26], [27]. However, it is unclear if the lower stiffness or other design differences associated with CDOs primarily used with civilians will result in different lower limb joint mechanics. Changes in joint mechanics at more proximal joints of the lower extremity, such as the knee joint, have been reported with the use of CDOs [26]. A study examining the effect of posterior strut stiffness on joint mechanics in individuals with lower limb trauma found that a more compliant strut resulted in stance phase knee flexion that did not differ from healthy controls, and was over 20% lower than that observed with the clinically prescribed strut, or a stiffer strut [26]. In addition, the peak knee extensor moment was greatest with the stiffest strut and differed from healthy controls [26]. In individuals with peripheral neuropathy, increased CDO stiffness reduced ankle dorsiflexion angle and decreased peak ankle power [28]. Further, optimizing stiffness had positive outcomes on gait biomechanics, and CDOs with lower stiffness have been shown to improve walking energy cost and walking speed [28]. In contrast, stiffer CDOs may be preferred for high intensity activities after trauma to prevent painful motion and due to a lower risk of mechanical failure [26]. However, the effect of other commonly available commercial CDOs on lower limb biomechanics in individuals with lower limb trauma, which primarily use CDOs for performing lower impact activities, such as walking, has not been investigated.
The effect of CDOs used in the civilian setting on joint mechanics of the lower extremity following lower limb trauma are poorly understood. It is unclear if civilian CDOs have similar effects on ankle range of motion and power, and if compensations are seen at other proximal joints of the lower extremity. Further, prior studies on military CDOs did not compare walking with the CDO to walking without a CDO (NoCDO) in individuals with lower limb trauma. It is important to establish if kinematics and kinetics of the lower limb extremity joints with CDOs used in civilian settings resemble those of military CDOs, before determining if further design changes are required to improve patient rehabilitation. Therefore, the purpose of this study was to examine the effect of two commercially available CDOs (MONO and MOD) relative to each other and no CDO, on kinematics and kinetics of lower extremity joints, and ground reaction forces, in individuals with lower limb trauma.
2. Methods
2.1. Participants
Inclusion criteria included individuals aged 18–65 years, who had experienced a lower extremity traumatic injury at least 2 years prior to testing, with the ability to walk 15.2 m without using a crane or crutch, and experienced any of the following: pain with foot loading (>4/10 on verbal numeric pain rating scale), limited pain free ankle range of motion (dorsiflexion<10 deg and plantarflexion<20 deg), weakness of ankle plantarflexors (<4/5 on manual muscle testing), history of ankle/hindfoot fusion, or candidate for ankle/hindfoot fusion or amputation secondary to ankle/foot injury and impairment. Exclusion criteria included any of the following: pain > 8/10 while walking, ankle weakness due to spinal cord injury or central nervous system pathology, planned surgery on the involved limb in the next 6 months, use of an orthosis crossing the knee or assistive device to perform activities of daily living, other medical conditions (e.g. traumatic brain injury, stroke, heart disease, vestibular disorder), psychological conditions, neurologic, musculoskeletal or other conditions affecting the contralateral limb, visual or hearing impairment, body mass index greater than 45, and pregnancy. This study was approved by the Institutional Review Board at each site, and all study participants provided written informed consent prior to participation.
2.2. Testing protocol
Data were collected while walking in three conditions; without a CDO (NoCDO), and with modular (MOD) and monolithic (MONO) CDOs that represent two primary types of clinically provided CDOs (Fig. 1), in a randomized order. The MOD CDO (Reaktiv, FabTech Systems, Everett, WA, USA) consisted of three carbon-fiber components: a proximal cuff for circumferential support with a patellar tendon bearing design, a stiff posterior strut, and semi-rigid full-length footplate. The MOD CDO generally resembles CDOs used in previous literature and in the military, except that the rigid carbon fiber anterior shell on the proximal cuff was secured using a cable based ratcheting system (BOA Technology Inc., Denver CO). The MONO CDO (PhatBrace, Bio-Mechanical Composites Inc., Des Moines, IA, USA) consisted of a single carbon fiber strut running from a compliant proximal cuff, down the posterior aspect of the leg, and through a more compliant foot plate than the MOD CDO. The proximal cuff of the MONO CDO was padded and secured using two Velcro straps. Certified Prosthetist Orthotists (CPOs) performed participant casting and fitting with the study CDOs based on a standardized fitting manual. Participants accommodated to each CDO for at least three months prior to testing. Participants began their three-month accommodation for the next orthosis immediately after testing the first.
A minimum of 12 cameras (120 Hz; Vicon Motion Systems Ltd., Denver, CO, USA or Qualisys AB, Gothenburg, Sweden) and three force plates (1200 Hz; AMTI, Watertown, MA) collected kinematic and kinetic data using reliable methods [29]. A 57-marker set was used to track the position of the foot, shank, thigh, and pelvis as described previously [14], [30], [31]. Participants walked across a level walkway at a standardized speed based on leg length, a Froude number of 0.16 [30], while wearing athletic shoes. The walking speed was tracked using a marker placed on the C7 vertebra and was kept consistent between sessions. Verbal feedback, if required, was provided to the study participants during walking, to ensure they were within +/− 5% of the target speed. Data were collected from a minimum of three complete gait cycles and analyzed using Visual-3D software (C-motion Inc., Germantown, MD). Sagittal plane ankle, knee, and hip angles, net internal joint moments, and joint powers, and ground reaction forces (GRF) were calculated for the injured limb using Visual-3D, and peak values were extracted using max and min functions in MATLAB (The MathWorks, Inc., Natick, MA, USA) consistent with previously published methods [29], [30]. Two descriptive variables, stiffness and alignment, were assessed to compare the design characteristics between the MONO and MOD CDOs. Stiffness testing was completed with a mechanical testing system, and stiffness was calculated using the ascending slope of the torque angle curve. Alignment was determined using ankle angle when the orthosis is unloaded during the swing phase [20], [31]. Previous studies investigating similar outcome measures indicate a sample of greater than 20 participants is sufficient to detect between device differences [6], [14], [25], [26], [27].
2.3. Statistical analysis
Shapiro-Wilk test was used to assess normality of the data, and statistical analysis was performed using SPSS version 28 (SPSS, IBM Corporation, Armonk, NY). The data were normally distributed., and descriptive statistics, including the mean and standard deviation (SD), were calculated for each dependent measure. A repeated measures analysis of variance (ANOVA) was performed to assess the main effect of the condition (NoCDO, MONO, and MOD) for each dependent measure. Significance level was set at α ≤ 0.05. Post-hoc analysis using paired t-tests with Bonferroni-Holm corrections were used to test the significant main effects of the condition. Cohen’s d, was used to determine the effect size for each significant pairwise comparison [32]. Cohen’s d was interpreted as 0.20 ≤ d < 0.50 representing a small effect, 0.50 ≤ d < 0.80 representing a medium effect, and d ≥ 0.80 representing a large effect. Effect size data is available in supplemental materials.
3. Results
In total, 23 individuals (15 males and 8 females) with mean (SD) age of 42.1 (11.4) years, height of 1.8 (0.2) m, and body mass of 101.5 (18.9) kg completed testing in all study conditions. Participants were 6.1 (5.7) years post injury. The MOD was significantly stiffer than MONO (stiffness: MOD=8.7[2.7] Nm/degree, MONO=4.6[2.4] Nm/degree). The device alignment did not differ between the MONO (5.2[2.7] degrees) and MOD (5.9[2.0] degrees) CDOs. Multiple pairwise differences in kinematic and kinetic measures were found between the study conditions. Table 1 presents biomechanical data for the three conditions, along with representative data from a prior publication that evaluated gait biomechanics using similar methods in twenty-four individuals with unilateral lower limb trauma with an IDEO [33]. Fig. 2 displays joint angle, moment, and power profiles, and GRF profiles across the gait cycle.
Table 1.
Mean (SD) kinematic, kinetic, and temporal-spatial parameters of gait collected at the standardized speed for the three study conditions (NoCDO = no device, MONO = monolithic device, MOD = modular device), and in comparison to lower limb reconstruction (LLR) data from Esposito et al. [33].
| Empty Cell | NoCDO | MONO | MOD | LLR Affected (Esposito et al. 2017) |
|---|---|---|---|---|
| Kinematics (deg) | ||||
| Ankle | ||||
| Peak PF (early stance) | −2.6 (3.7) | −0.1 (3.3) | 1.5 (2.3) | −0.3 (3.0) |
| Peak DF (stance) | 14.6 (4.4) | 11.3 (3.8) | 12.1 (2.9) | 12.9 (3.6) |
| Peak PF (early swing) | −7.8 (8.7) | 2.5 (3.2) | 4.1 (2.2) | 5.7 (3.2) |
| ROM | 23.6 (6.3) | 11.6 (2.5) | 10.7 (2.5) | 13.2 (3.1) |
| Knee | ||||
| Peak extension (initial contact) | −2.1 (5.3) | −3.9 (4.7) | −1.1 (4.2) | 1.2 (6.5) |
| Peak flexion (midstance) | 11.8 (5.2) | 10.7 (5.2) | 12.9 (5.3) | 15.1 (6.7) |
| Peak extension (terminal stance) | 1.2 (5.7) | −0.7 (4.4) | 2.4 (4.9) | - |
| ROM | 59.8 (6.3) | 59.7 (5.3) | 59.8 (5.2) | 63.9 (4.9) |
| Hip | ||||
| Peak flexion (early stance) | 29.2 (8.5) | 29.6 (6.4) | 30.3 (6.8) | 35.5 (4.9) |
| Peak extension (pre-swing) | −4.9 (8.6) | −5.8 (6.4) | −5.5 (5.9) | −4.9 (5.1) |
| Peak flexion (terminal swing) | 32.1 (8.8) | 32.2 (5.9) | 32.5 (6.9) | - |
| ROM | 37.1 (3.3) | 38.1 (3.6) | 38.1 (3.4) | 41.9 (3.6) |
| Joint Moments (Nm/kg) | ||||
| Ankle | ||||
| Peak dorsiflexor (early stance) | −0.23 (0.08) | −0.31 (0.13) | −0.31 (0.09) | −0.38 (0.09) |
| Peak plantarflexor (terminal stance) | 1.21 (0.15) | 1.31 (0.28) | 1.44 (0.17) | 1.49 (0.22) |
| Knee | ||||
| Peak flexor (initial contact) | −0.38 (0.11) | −0.36 (0.15) | −0.31 (0.11) | - |
| Peak extensor (early stance) | 0.57 (0.28) | 0.66 (0.25) | 0.65 (0.19) | 0.50 (0.23) |
| Peak flexor (terminal stance) | −0.15 (0.14) | −0.21 (0.15) | −0.23 (0.14) | −0.43 (0.16) |
| Hip | ||||
| Peak extensor (initial contact) | 0.74 (0.17) | 0.71 (0.24) | 0.67 (0.17) | 0.91 (0.24) |
| Peak flexor (terminal stance) | −0.95 (0.30) | −1.05 (0.28) | −0.95 (0.32) | −0.59 (0.16) |
| Powers (W/kg) | ||||
| Ankle | ||||
| Peak absorption (loading response) | −0.26 (0.11) | −0.22 (0.10) | −0.20 (0.09) | −0.35 (0.10) |
| Peak absorption (terminal stance) | −0.81 (0.41) | −0.54 (0.21) | −0.51 (0.16) | −0.74 (0.35) |
| Peak generation (pre-swing) | 2.03 (0.74) | 1.22 (0.36) | 1.25 (0.39) | 0.90 (0.37) |
| Knee | ||||
| 1st peak generation (early stance) | 0.84 (0.31) | 0.88 (0.40) | 0.67 (0.22) | 0.72 (0.33) |
| 1st peak absorption (midstance) | −0.66 (0.42) | −0.78 (0.45) | −0.71 (0.36) | −0.50 (0.29) |
| 2nd peak generation (terminal stance) | 0.42 (0.22) | 0.43 (0.18) | 0.46 (0.22) | 0.62 (0.30) |
| 2nd peak absorption (pre-swing) | −1.39 (0.56) | −1.45 (0.47) | −1.54 (0.77) | −0.49 (0.23) |
| Hip | ||||
| 1st peak generation (midstance) | 0.36 (0.27) | 0.32 (0.16) | 0.40 (0.23) | 0.85 (0.27) |
| Peak absorption (terminal stance) | −0.70 (0.36) | −0.74 (0.29) | −0.62 (0.33) | −0.43 (0.15) |
| 2nd peak generation (Pre-swing) | 1.04 (0.29) | 1.09 (0.36) | 1.01 (0.41) | 0.78 (0.21) |
| GRF (x body weight) | ||||
| Vertical | ||||
| 1st Peak vertical force | 1.06 (0.05) | 1.12 (0.27) | 1.07 (0.06) | 1.07 (0.06) |
| 2nd Peak vertical force | 1.00 (0.06) | 1.07 (0.22) | 1.04 (0.05) | - |
| ML | ||||
| Peak medially directed force | −0.03 (0.02) | −0.03 (0.02) | −0.03 (0.01) | 0.07 (0.01) |
| Peak laterally loaded force (stance) | 0.07 (0.02) | 0.08 (0.03) | 0.07 (0.01) | - |
| AP | ||||
| Peak braking force | −0.15 (0.03) | −0.16 (0.03) | −0.14 (0.02) | −0.14 (0.03) |
| Peak propulsive force | 0.16 (0.03) | 0.17 (0.03) | 0.15 (0.02) | 0.15 (0.02) |
PF – plantarflexion, DF – dorsiflexion, ROM – range of motion, GRF – ground reaction forces, AP – anterior/posterior, ML – medial/lateral.
Significantly different from NoCDO – in bold.
MOD significantly different from MONO – underlined.
Fig. 2.
Ankle, knee and hip joint angles, moments and powers in the sagittal plane, and vertical, anterior/posterior and medial/lateral ground reaction force (GRF) across the entire gait cycle for the three study conditions (NoCDO = no device, MOD = modular device, MONO = monolithic device). Positive angle values represent ankle dorsiflexion, knee flexion, and hip flexion angles. Positive moment values represent internal ankle plantarflexor, knee extensor, and hip extensor moments. Positive power values represent power generation and negative values represent power absorption.
4. Discussion
Studies evaluating the effect of CDOs primarily used in the civilian settings on gait mechanics in individuals with lower limb trauma are limited, as are studies comparing lower extremity kinematics and kinetics with and without CDOs. The results of this study were consistent with previous publications including CDOs, showing that CDOs used in civilian settings reduce ankle motion and power compared to the injured limb with NoCDO, with a limited effect at more proximal joints. The CDOs studied here resulted in generally similar limb mechanics to those used in a military setting [27], [33].
The study CDOs restricted ankle motion and push-off power relative to NoCDO, however, the push-off power was over 25% greater than previously reported with military CDOs [33]. The MONO and MOD CDOs reduced the total sagittal plane range of motion of the ankle to just over 10 degrees, which is less than half of that with a NoCDO, with total motion within two degrees of prior studies with CDOs used in a military setting [33]. Similar to other CDO studies, the reduced motion was primarily due to the elimination of plantarflexion in pre-swing and initial-swing as the CDO returns to its unloaded position [33], [34]. This reduced motion contributed to an approximately 40% reduction in ankle power generation as compared to NoCDO. The reduced motion and power at the ankle with the study CDOs is similar to other AFO related studies in other populations [28]. In individuals with calf muscle weakness, the dorsiflexion angle at the ankle reduced by approximately 1⁰ with each additional 1 Nm/degree of AFO stiffness, reducing the ability to generate ankle power [28]. Findings from our study further suggest that the decreased ankle power generation seen with the study CDOs was primarily due to the reduced ankle joint motion and associated change in angular velocity as compared to changes in the ankle moment. Multiple factors could contribute to the greater reduction in peak ankle power generation with military CDOs reported previously, including a design that is tailored for performing high-impact activities such as load carriage or running and limits the ability to deform the CDO during gait [9], [35], or even differences between military and civilian gait [14], [31], [36].
Despite the lack of differences in stance phase knee kinematics with the study CDOs, changes were observed in both knee joint moments and powers. In this study, peak power generation at the knee was over 20% greater with the MONO CDO as compared to the stiffer MOD CDO. Additionally, the MONO and MOD CDOs resulted in greater peak flexor moments at the knee during terminal stance, which were about 50% greater as compared to the NoCDO. In contrast, the peak flexor moments at the knee with the two study CDOs were about 50% lower as compared to that reported previously with military CDOs [33]. There can be many potential reasons for this strategy, including perceived comfort, alignment, and other design characteristics of the CDOs. Previous studies have shown that an increase in ankle joint stiffness associated with the use of a CDO was countered by increased knee compliance, in an effort to maintain whole-limb stiffness [34]. The peak knee power generation during early stance was about 20% lower with the MOD CDO as compared to MONO CDO, and similar to that reported previously with military CDOs [26], [33]. Decreased power generation at the knee may be a compensatory strategy related to a greater reduction in range of motion and power generation at the ankle, especially with stiffer CDOs such as the MOD, as compared to the MONO and NoCDO conditions. Future studies aiming to evaluate changes in knee joint mechanics with other commonly used CDOs in the civilian setting are needed to establish the consistency of these findings in individuals with lower limb trauma.
The study CDOs did not alter the hip joint motion, however, the peak power generation at the hip was about 50% less than previously reported with military CDOs [33]. Joint bracing has been shown to affect gait mechanics in healthy individuals, with ankle-bracing creating asymmetries at the ankle and hip joints during terminal stance [37]. In a previous study, greater hip flexion angles were observed with the use of military CDOs in individuals with lower limb trauma [26]. Findings from this study do not provide support for a trade-off between the ankle and the hip joints while walking with CDOs commonly used in civilian settings. Future studies are needed to evaluate the effect of other commercially available CDOs on joint kinematics and kinetics of other lower limb joints, such as the hip, during walking.
Direct comparison between the military and civilian populations can be difficult. The injury characteristics of military service members may have differences from civilians, and service members tend to be physically fit prior to injury. In addition, service members generally receive CDO specific multi-disciplinary rehabilitation, such as the Return to Run (RTR) rehabilitation program, which is designed to enable a return to active duty and sports [9]. However, the effect of CDOs on gait mechanics are generally similar between CDOs included in this study and with those studied in service members, even though the study participants did not receive any specific rehabilitation training. However, it is difficult to determine if it is a single design characteristic or a combination of multiple characteristics of the study CDOs that produced a similar effect on limb mechanics as observed with the orthosis commonly used in the military, and is an important consideration for future research in the area. From a clinical perspective, the current study indicates that that practitioners may achieve similar limb mechanics with orthoses with many apparent design differences, including stiffness. This provides clinicians flexibility when selecting the preferred device for their patient, It is important to note that alignment, which has been shown to have a large effect on CDO related limb mechanics, was not different between the devices in this study. The premise that device configuration, including alignment, is more important than the specific type of orthosis warrants further study.
This study had several limitations. The study CDOs were not directly compared to the IDEO, however, methods identical to those used in previous studies were followed to facilitate comparison. Additionally, the CDOs used in this study had multiple design differences including in cuff type, foot plate geometry, and posterior strut, preventing attribution of differences to a singular design difference. However, the study CDOs are representative of commercially available options and would be expected to show differences in limb mechanics, if they were present. Further, while the MONO and MOD are commonly used CDOs in the civilian setting, the findings from this study can’t be generalized to other commercially available CDOs.
5. Conclusion
Despite apparent design differences between the two study CDOs, differences in their effect on gait mechanics in individuals with lower limb trauma are limited. Commonly available CDOs in civilian settings reduced ankle range of motion and power generation during walking, consistent with prior studies on CDOs used in the military. As compared to the MONO, the MOD resulted in a greater peak flexor moment at the knee joint and lower ground reaction braking force. Commonly available CDOs used in civilian settings affect lower limb biomechanics in a manner similar to previous reports with military CDOs following lower limb trauma.
Supplementary Material
Highlights.
Orthosis use affects lower extremity joint kinetics and kinematics.
The effect of commonly available orthoses on joint mechanics is poorly understood.
Study orthoses generated greater ankle push-off power than previously reported.
Design differences between study orthoses had a limited effect on gait mechanics.
Acknowledgements
Research reported in this publication was supported in part by a Department of Defense grant under award W81XWH-18-2-0073 and by the National Center for Advancing Translational Sciences of the National Institutes of Health under Award Number UM1TR004403. The content is solely the responsibility of the authors and does not necessarily represent the official views of the National Institutes of Health, the U.S. Departments of the Army, Navy, Air Force, Defense, the Henry M. Jackson Foundation for the Advancement of Military Medicine, Inc., Uniformed Services University of the Health Sciences, nor the U.S. Government. Identifying specific products or instrumentation is considered an integral part of the scientific endeavor and does not constitute an endorsement or implied endorsement on the part of the authors, Department of Defense, or any other entity.
We would also like to acknowledge the contributions of Jeff Palmer, Matt Husnik, Jeff Fay, Christopher Dearth, Ashley Knight, and Sara Magdziarz for their efforts in support of study completion.
Footnotes
CRediT authorship contribution statement
Molly S Pacha: Writing – review & editing, Validation, Supervision, Software, Project administration, Methodology, Investigation, Funding acquisition, Data curation. Kierra J. Falbo: Writing – review & editing, Project administration, Investigation. Kirsten M Anderson: Writing – review & editing, Validation, Supervision, Software, Project administration, Methodology, Investigation, Formal analysis, Data curation. Sapna Sharma: Writing – review & editing, Writing – original draft, Visualization, Validation, Investigation, Formal analysis, Data curation. Brad D. Hendershot: Writing – review & editing, Supervision, Resources, Project administration, Methodology, Investigation, Funding acquisition, Conceptualization. Jason M. Wilken: Writing – review & editing, Writing – original draft, Visualization, Validation, Supervision, Software, Resources, Project administration, Methodology, Funding acquisition, Data curation, Conceptualization. Clare Severe: Writing – review & editing, Validation, Supervision, Resources, Investigation, Data curation. Andrew H. Hansen: Writing – review & editing, Supervision, Project administration, Funding acquisition, Conceptualization.
Declaration of Competing Interest
The authors declare that they have no known competing financial interests or personal relationships that could have appeared to influence the work reported in this paper.
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