Annually, 795,000 individuals have a stroke, which is the leading cause of disability in the United States.1 Even after rehabilitation, poststroke gait dysfunction is common and has been linked to activity and participation limitations.2,3 This disability is both due to and a risk factor for physical inactivity,4,5 which has serious health consequences for chronic stroke survivors.6,7
Muscle weakness is a primary impairment and contributor to gait dysfunctions in individuals poststroke.8,9 Plantarflexor weakness is a particularly impactful impairment for stroke survivors.9–13 The plantarflexors play an important role in gait.14–16 During mid stance when the ankle is dorsiflexing, the plantarflexors eccentrically control the shank's rotation14,15,17,18 (supported dorsiflexion). Supported dorsiflexion can be measured via net ankle pseudostiffness (NAS), defined as the ratio of the change in plantarflexion moment to ankle angle during the period of dorsiflexion in stance.19,20 This supported dorsiflexion provides the majority of the stability for the entire body during this phase of gait15,16,21–23 and sets up the proper positioning for the limb's kinematics.14,18,24 Through these functions, supported dorsiflexion establishes proper kinematics and energetic states for the plantarflexors to effectively aid forward propulsion10,16,25–27 and swing initiation.10,16,28 Thus, plantarflexor weakness greatly impairs gait function. During stance, plantarflexor weakness commonly presents as excessive dorsiflexion17,18 or knee hyperextension.14,29 Both gait impairments hinder propulsion and forward progression and result in shorter and often asymmetrical step lengths as well as decreased walking speeds.9,14,18,30,31
The spring-like bending stiffness of passive-dynamic ankle-foot orthoses (PD-AFOs) has the capacity to replicate many of the functions of supported dorsiflexion.32 Similar to NAS, PD-AFO bending stiffness is defined as the change in moment divided by angular deformation.33 It is believed that PD-AFO characteristics, including bending stiffness, must be personalized.34–38 However, a method to readily and effectively personalize PD-AFO bending stiffness has not yet been established. Currently, AFOs are fabricated by hand. This craft-based fabrication process results in orthoses with varying characteristics, which are typically never quantified. This craft-based process has made it difficult to understand the interaction between orthoses and gait function and has limited establishment of objective prescription guidelines for PD-AFOs. Furthermore, this lack of standardization results in many orthoses having variable effectiveness and suboptimal enhancements in patient function.37
A technology that harnesses additive manufacturing to rapidly fabricate orthoses with objectively personalized characteristics has been developed.39–41 While this technology overcomes the craft-based limitations of the current process, an objective method to effectively personalize PD-AFO characteristics must be established to fully leverage this technology. Other studies have begun to explore the effects of experimentally measured42–45 or simulated46,47 orthosis bending stiffness on gait, but no study to date has presented and evaluated an objective method to personalize PD-AFO bending stiffness.
It was recently demonstrated that PD-AFO bending stiffness can substitute for healthy plantarflexor function.32 Accordingly, for individuals poststroke with plantarflexor weakness, we hypothesized that PD-AFO bending stiffness can add to the individual's remaining plantarflexor function, which can be quantified via the peak plantarflexion moment during stance in gait. Thus, we predicted that PD-AFO bending stiffness can add to the individual's existing plantarflexion moment to make up for the lost moment and restore net (individual + PD-AFO)-supported dorsiflexion to a typical level. The purpose of this study was to investigate if PD-AFO bending stiffness personalized to account for an individual's level of plantarflexor deficit can enhance the net-supported dorsiflexion. By understanding if PD-AFO bending stiffness can enhance an individual's plantarflexor function, this is a first step toward developing a prescription model to objectively personalize PD-AFO bending stiffness for patients poststroke with plantarflexor muscle weakness.
Two individuals poststroke (subject A: male, 43 years, 1.82 m, 66 kg; subject B: female, 49 years, 1.70 m, 84.2 kg) participated in the study, which was approved by the University of Delaware Institutional Review Board. To be included in this study, subjects had to be at least 6 months poststroke, prescribed an AFO by a clinician, have adequate paretic ankle dorsiflexion range of motion (ROM), defined as 15° or greater measured via instrumented ROM test (see Baseline Visit), and have plantarflexor strength deficits, defined as decreased paretic peak plantarflexion moment during stance when compared with scaled speed-matched normative values. Medical clearance was obtained from subjects' physicians before their participation.
Each subject underwent 2 visits to the laboratory. First, a baseline visit was done where measurements to personalize the PD-AFO were obtained. A PD-AFO was then personalized based on each subject's level of plantarflexor function deficit and manufactured for each subject using the patented technology. Next, subjects returned to the laboratory for an evaluation visit where motion analysis data were collected as subjects walked while wearing the personalized PD-AFO.
Each subject's self-selected walking speed was determined using a 10-m walk test. The subject walked along a 10-m walkway with time recorded during the middle 6 m to allow for acceleration and deceleration. The subject performed two trials; his/her self-selected speed and results were averaged.48
Second, an instrumented ankle ROM evaluation was performed to verify that the subject had sufficient dorsiflexion ROM. The subject's paretic side ankle ROM was measured as a physical therapist maximally dorsiflexed the ankle with the subject seated and relaxed 1) with knee at 90° and 2) with knee straight. Foot and shank kinematics were recorded via a motion analysis system (Motion Analysis Corporation, Santa Rosa, CA, USA) using a six degree-of-freedom marker set49 on the subject's foot and shank. From these data, dorsiflexion ROM was calculated.
Next, the subject's plantarflexor function deficit was quantified. The subject walked without any orthosis, if possible, at his/her previously determined self-selected speed on a split-belt, instrumented treadmill (Bertec Corp, Columbus, OH, USA). Subjects were allowed a light touch on the handrails if necessary and wore a safety harness that provided no body weight support. The subjects were given time to acclimate to treadmill walking. Then, 3-dimensional (3D) kinematic and kinetic data to describe five gait cycles were captured via the treadmill force plates and a six-camera motion analysis system. Retroreflective markers were placed bilaterally on the subject's lower limbs using a six degree-of-freedom marker set.49 Kinematic and kinetic data were collected at 240 Hz and 1200 Hz and filtered at 6 Hz and 25 Hz, respectively, using a zero-lag low-pass Butterworth filter.50
Finally, a set of discrete 3D landmarks on each subject's paretic lower leg were digitized using a 3D Fusion FaroArm (FARO Technologies Inc, Lake Mary, FL, USA).39,41 These landmarks captured the size and shape of each subject's shank and foot and were used to customize the PD-AFO fit.39,41
PD-AFO PERSONALIZATION AND MANUFACTURING
From the baseline data, the subject's paretic net ankle joint moment was calculated via inverse dynamics in Visual 3D (C-Motion Inc, Germantown, MD, USA), averaged across the five gait cycles and scaled by body mass.50,51 In addition, self-selected walking speed was scaled by height, resulting in a speed in statures/second. The subject's scaled paretic net plantarflexion moment was compared with a scaled speed-matched net plantarflexion moment from our laboratory's normative database. This database included lower-limb motion analysis data for 10 individuals (average age, 24.2 ± 3.0 years; average height, 1.71 ± 0.07 m; average weight, 69.6 ± 11.3 kg) with no prior history of musculoskeletal injury or disease each walking at a range of scaled speeds. The subject's plantarflexion function deficit was quantified as the difference between the speed-matched typical peak plantarflexion moment and the subject's peak plantarflexion moment. PD-AFO bending stiffness was therefore personalized as the difference in peak moments (typical − paretic) divided by the average typical change in ankle angle during the period of dorsiflexion (15°). Using the digitized fit characterization data and this bending stiffness prescription, a PD-AFO was personalized then manufactured for each subject.39–41 PD-AFOs were manufactured out of medical-grade polycarbonate (Fortus 3D Production Systems, Stratasys Inc; Eden Prairie, MN, USA) via fused deposition modeling.40 Once manufactured, PD-AFOs were verified for proper bending stiffness40 and padded (Fig. 1).
To note, these PD-AFOs were research orthoses. In particular, this PD-AFO was designed to be worn without a shoe. Instead of a shoe, the PD-AFO footplate interfaced with the ground (Fig. 1). This PD-AFO design was ideally suited for this study as all PD-AFOs' characteristics were precisely known and could be controlled.
The subject donned the personalized PD-AFO, which was visually assessed for fit, and then walked over ground for several minutes to initially acclimate to wearing the PD-AFO. Then, with the subject walking on a treadmill, data were collected under three randomly presented conditions: without any AFO if possible (No AFO), while wearing his/her originally prescribed AFO (Original AFO), and while wearing his/her bending stiffness-personalized PD-AFO (Personalized PD-AFO). Subject B was unable to walk without an orthosis. For the Personalized PD-AFO condition, kinematics of the PD-AFO cuff were recorded separately from subject's shank with three additional tracking markers placed directly on the PD-AFO cuff so angular deformation of the PD-AFO could be determined independently from the ankle angle. Subjects walked at their previously determined self-selected walking speed for all conditions and were given time to acclimate to each condition while walking on the treadmill prior to the data being collected. Three-dimensional kinematic and kinetic data were collected using the same procedures as the baseline visit.
Lower-limb joint kinematics and kinetics during stance were calculated via inverse dynamics in Visual 3D and averaged across five gait cycles for each condition for both subjects. To evaluate this study's primary hypothesis that stiffness-personalized PD-AFOs could enhance net plantarflexor function during gait, peak net plantarflexion moments, NAS, and the PD-AFO's and subject's contribution to the net plantarflexion moment during the Personalized PD-AFO condition were calculated and compared among conditions. The ankle moment provided by the PD-AFO and subject were determined by calculating the PD-AFO moment (bending stiffness times orthosis angular deformation), then subtracting the PD-AFO moment from the net ankle joint moment to determine the subject's moment. The subject's peak plantarflexion moment was compared to the peak net plantarflexion moment in the No AFO condition to determine if the PD-AFO bending stiffness enhanced the subject's plantarflexor function (if the moment generated by the subject remained the same with and without PD-AFO use) or if the PD-AFO bending stiffness substituted for some of the subject's available function (if the moment generated by the subject decreased with PD-AFO use). Finally, paretic side ankle, knee, and hip angles during mid and late stance, which is the period of interest for these bending stiffness-personalized PD-AFOs, were compared across conditions to evaluate the influence of the personalized PD-AFOs on joint kinematics.
Subjects A and B walked at 0.6 statures/second and 0.36 statures/second, respectively, with ∼0.80 statures/second an average, typical comfortable walking speed. From the instrumented ankle ROM evaluation, both subjects had sufficient dorsiflexion range (subject A, 18.9°; subject B, 15.2°). Subjects A and B had 25% and 31% plantarflexor function weakness, respectively. PD-AFOs tuned to 2.0 Nm/deg and 2.5 Nm/deg were personalized and fabricated for subjects A and B, respectively.
In support of our hypothesis, results suggest that stiffness-personalized PD-AFOs can increase the net peak plantarflexion moment (Fig. 2) and NAS (Table 1) compared with the baseline conditions. The Personalized PD-AFOs returned NAS to slightly higher than typical levels for both subjects (Table 1). Furthermore, importantly, the PD-AFO bending stiffness added to, and did not substitute for, the subject's plantarflexion function as the peak plantarflexion moment generated by the subject remained the same with and without PD-AFO use (Fig. 2).
Dorsiflexion excursions decreased with Personalized PD-AFO use compared with the baseline conditions and were lower than average excursions for speed-matched typical individuals (Table 1). Subject A walked with excessive knee extension during mid stance without any orthosis and with his Original AFO (Fig. 3C). The PD-AFO brought peak knee extension during mid stance back to typical levels (Fig. 3C). In contrast, subject B walked with an excessively flexed knee throughout stance with her Original AFO (Fig. 3D). Like subject A, use of the personalized PD-AFO brought subject B's knee kinematics to a more typical pattern during mid and late stance (Fig. 3D). Subject A only had slight hip deviations during stance without any orthosis while subject B walked with considerable and sustained hip flexion in her baseline condition (Fig. 3E–F). Still, personalized PD-AFO use resulted in improved hip angle during mid and late stance for both subjects (Fig. 3E–F).
This study evaluated whether or not bending stiffness-personalized PD-AFOs could enhance the net plantarflexor function of individuals poststroke with plantarflexor weakness. Results suggest that PD-AFOs personalized to account for an individual's level of plantarflexor function deficit can add to the individual's existing plantarflexor function and improve net-supported dorsiflexion, quantified via NAS. Furthermore, while the subjects in this study walked with very different kinematic deviations, the stiffness personalized PD-AFOs improved both gait patterns during mid and late stance, which is when the PD-AFO bending stiffness was utilized.
Both subjects had substantially reduced peak plantarflexion moments during stance when compared with scaled speed-matched normative values. Subject A's originally prescribed orthosis made little to no improvement in this moment, and subject B's peak plantarflexion moment showed substantial deficits with the originally prescribed orthosis. However, in agreement with our hypothesis, the personalized PD-AFOs enhanced both subjects' peak moment, restoring subject A's peak moment to typical levels and considerably improving subject B's peak moment. Moreover, the assistance provided by these PD-AFOs added to, and did not substitute for, the subjects' existing plantarflexor function as the subjects' moments remained consistent with and without orthosis use. Personalized PD-AFO use also resulted in increases in NAS. These results support the potential efficacy of this objective PD-AFO bending stiffness personalization method for individuals with plantarflexor weakness.
In contrast to results from our study, Takahashi and colleagues52 found paretic soleus activity decreased when individuals poststroke used a powered exoskeleton. In other words, the powered exoskeleton appeared to substitute for some of the subjects' existing soleus function. Decreasing paretic muscle activity may be counterproductive to the rehabilitation and long-term function of individuals poststroke. Similarly, results reported by Bregman and colleagues42 also appear to show that the AFOs tested substituted for some of the subjects' existing plantarflexor function. The orthoses used in the study by Bregman et al were passive devices and prescribed by a clinician. A key difference between these studies and the present study is that in this study the orthosis bending stiffness was personalized based on the subject's level of plantarflexor weakness. In contrast, the orthosis personalization method in the study by Takahashi et al was unclear, and the orthoses used by Bregman et al were likely not objectively personalized as they were prescribed clinically. Therefore, these results support the importance of objectively personalized orthosis bending stiffness based on an individual's unique impairment profile, such as plantarflexor function deficits.
Interestingly, all orthoses also induced a premature increase in the plantarflexion moment earlier in stance before the peak moment. For subject A, this premature increase was exacerbated with personalized PD-AFO use. Similar premature increases were found in other studies with individuals poststroke42 as well as when typical individuals walked while wearing bending stiffness-tuned PD-AFOs.32 In typical individuals, this premature moment may be a result of the PD-AFOs replicating soleus function but not gastrocnemius function.53 However, the cause of this premature plantarflexion moment increase and it implications on gait function needs to be investigated further for individuals poststroke.
While the PD-AFOs provided benefits to the plantarflexion joint moment, their use excessively reduced dorsiflexion excursions for both subjects, which resulted in NAS values above typical levels. The reduced ankle dorsiflexion may have been because the subjects were not trained to use the PD-AFOs. If the subjects are trained to bend into the orthoses and use the PD-AFOs more, dorsiflexion excursions may improve. Alternatively, even though these PD-AFOs provided near optimal improvements in plantarflexion moments, these results may indicate that the PD-AFOs were too stiff for the subjects to fully deform. In concurrence with this finding, a study by Kobayashi and colleagues44 found a decrease in dorsiflexion excursions as orthosis bending stiffness increased for individuals poststroke. Future research should study the energetic implications of bending stiffness-tuned PD-AFOs to explore this concept further.
Both subjects had impaired, yet different, paretic knee and hip kinematics in the baseline conditions. Use of the Personalized PD-AFOs brought both subject's knee and hip kinematics toward a typical pattern, particularly during mid stance when the PD-AFO bending stiffness was being used. The ability of the Personalized PD-AFO to address opposing kinematic deviations demonstrates the importance of enhancing supported dorsiflexion. Given the immediate improvements in NAS and knee and hip kinematics during mid and late stance with PD-AFO use, training the subjects how to use the PD-AFO might result in even greater improvements.
While this study demonstrated the efficacy of the personalization method in improving supported dorsiflexion for two individuals poststroke, some limitations should be noted. This study evaluated the feasibility of this novel personalization method in improving supported dorsiflexion via a dual subject study design. With feasibility now demonstrated, this method needs to be assessed on more subjects as the poststroke population exhibits great variability. With additional subjects, the statistical significance of the findings can be evaluated to determine if this method can be generalized across the population of stroke survivors. In addition, while the subjects in this study were required to have adequate dorsiflexion ROM to be able to use the PD-AFO bending stiffness, other characteristics such as knee motion, tone, clonus, and proprioception may also influence to what extent individuals use the PD-AFO. Future studies should assess these additional factors, which may help identify for which patients the stiffness-personalized PD-AFOs are ideal. Moreover, while this study focused on the PD-AFOs' effect during mid and late stance when the orthosis bending stiffness was used, it is noted that for subject A, PD-AFO use increased knee flexion in early stance. These PD-AFOs were set into a few degrees of dorsiflexion, as is often done clinically, to facilitate tibial progression and utilization of the bending stiffness during mid and late stance. While this customization appeared to be useful for mid and late stance function, future research should investigate its effects on early stance and possibly optimize PD-AFO design to assist mid and late stance function without impeding early stance kinematics. Furthermore, this study focused on the personalized PD-AFOs' influence on ankle-foot function and paretic joint kinematics. Future studies should investigate the influence of these PD-AFOs on other aspects of gait function. Finally, results presented in this study are immediate effects of the PD-AFOs. With more training and use, the subjects may use the personalized PD-AFOs more (reach typical dorsiflexion excursions) and reap even greater benefits; however, the long-term influences of these personalized PD-AFOs are not yet known.
This study can be used to inform clinical practice as results show that personalized PD-AFO bending stiffness has the capacity to improve paretic ankle-foot function during gait, specifically during mid and late stance. Thus, AFO efficacy and poststroke gait function may be improved with objectively personalized PD-AFOs. To apply this knowledge in a clinical setting, practitioners need to be able to measure each individual's plantarflexor function deficit and have PD-AFOs with objectively personalized bending stiffness levels manufactured. While motion analysis is used in this study to quantify plantarflexor function deficits, less expensive clinically viable tools to quantify plantarflexor function are in development so practitioners will be able to obtain this measure in a clinical setting. In addition, several methods to customize and manufacture stiffness-personalized PD-AFOs are in various stages of development,39–41,54,55 which will ultimately facilitate the clinical implementation of knowledge gained from this study.
This study demonstrated that objectively personalized PD-AFO bending stiffness has the potential to enhance net plantarflexor function for individuals poststroke. Orthotists today certainly customize orthoses for patients by selecting the orthosis type and shaping the orthosis to qualitatively adjust its characteristics. However, the craft-based nature of the current process results in great variability in orthosis efficacy. The personalization and manufacturing technology applied in this study provides a unique tool to develop and eventually implement quantitative prescription models. With the PD-AFOs showing positive outcomes on paretic ankle-foot function during mid and late stance for the 2 subjects studied, this work lays the foundation for continued advancement, and possible implementation, of objectively personalized PD-AFOs for individuals with plantarflexor weakness.
1. Go A, Mozaffarian D, Roger V, et al. Heart disease and stroke
statistics—2014 update: a report from the American Heart Association. Circulation
2014; 129: e28–e292.
2. Vestling M, Tufvesson B, Iwarsson S. Indicators for return to work after stroke
and the importance of work for subjective well-being and life satisfaction. J Rehabil Med
2003; 35(3): 127–131.
3. Mayo NE, Wood-Dauphinee S, Ahmed S, et al. Disablement following stroke
. Disabil Rehabil
1999; 21(5–6): 258–268.
4. Mayo NE, Wood-Dauphinee S, Côté R, et al. Activity, participation, and quality of life 6 months poststroke. Arch Phys Med Rehabil
2002; 83(8): 1035–1042.
5. Rimmer JH, Wang E. Aerobic exercise training in stroke
survivors. Top Stroke Rehabil
2005; 12(1): 17–30.
6. Hornnes N, Larsen K, Boysen G. Little change of modifiable risk factors 1 year after stroke
: a pilot study. Int J Stroke
2010; 5(3): 157–162.
7. Lloyd-Jones D, Adams RJ, Brown TM, et al. Heart disease and stroke
statistics—2010 update: a report from the American Heart Association. Circulation
2010; 121(7): e46–e215.
8. Pak S, Patten C. Strengthening to promote functional recovery poststroke: an evidence-based review. Top Stroke Rehabil
2008; 15: 177–199.
9. Nadeau S, Gravel D, Arsenault B, Bourbonnais D. Plantarflexor weakness as a limiting factor of gait
speed in stroke
subjects and the compensating role of hip flexors. Clin Biomech
1999; 14(2): 125–135.
10. Peterson CL, Hall AL, Kautz SA, Neptune RR. Pre-swing deficits in forward propulsion, swing initiation and power generation by individual muscles during hemiparetic walking. J Biomech
2010; 43(12): 2348–2355.
11. Jonkers I, Delp S, Patten C. Capacity to increase walking speed is limited by impaired hip and ankle power generation in lower functioning persons post-stroke
. Gait Posture
2009; 29(1): 129–137.
12. Kim CM, Janice JE. The relationship of lower extremity muscle torque to locomotor performance in people with stroke
. Phys Ther
2003; 83: 49–57.
13. Parvataneni K, Olney SJ, Brouwer B. Changes in muscle group work associated with changes in gait
speed of persons with stroke
. Clin Biomech (Bristol, Avon)
2007; 22(7): 813–820.
14. Perry J, Burnfield JM. Gait
Analysis: Normal and Pathological Function. 2nd ed. SLACK Inc: Thorofare; 2010.
15. Kepple T, Siegel K, Stanhope SJ. Relative contributions of the lower extremity joint moments to forward progression and support during gait
. Gait Posture
1997; 6(1): 1–8.
16. 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(11): 1387–1398.
17. Jonkers I, Stewart C, Spaepen A. The complementary role of the plantarflexors
, hamstrings and gluteus maximus in the control of stance limb stability during gait
. Gait Posture
2003; 17(3): 264–272.
18. Sutherland DH, Cooper L, Daniel D. The role of the ankle plantar flexors in normal walking. J Bone Joint Surg Am
1980; 62-A(3): 354–363.
19. Davis RAB, Deluca PA. Gait
characterization via dynamic joint stiffness. Gait Posture
1996; 4: 224–231.
20. Latash ML, Zatsiorsky VM. Joint stiffness: myth or reality? Hum Mov Sci
1993; 12(6): 653–692.
21. McGowan CP, Kram R, Neptune RR. Modulation of leg muscle function in response to altered demand for body support and forward propulsion during walking. J Biomech
2009; 42(7): 850–856.
22. Francis CA, Lenz AL, Lenhart RL, Thelen DG. The modulation of forward propulsion, vertical support, and center of pressure by the plantarflexors
during human walking. Gait Posture
2013; 38(4): 993–997.
23. McGowan CP, Neptune RR, Kram R. Independent effects of weight and mass on plantar flexor activity during walking: implications for their contributions to body support and forward propulsion. J Appl Physiol
2008; 105(2): 486–494.
24. Owen E. The importance of being earnest about shank and thigh kinematics especially when using ankle-foot orthoses. Prosthet Orthot Int
2010; 34(3): 254–269.
25. Allen JL, Kautz SA, Neptune RR. Modular control of walking: a 3D simulation study. In: Proceedings of American Society of Biomechanics Annual Meeting. Long Beach: 2011: 3–4.
26. Liu MQ, Anderson FC, Pandy MG, Delp SL. Muscles that support the body also modulate forward progression during walking. J Biomech
2006; 39(14): 2623–2630.
27. Turns LJ, Neptune RR, Kautz SA. Relationships between muscle activity and anteroposterior ground reaction forces in hemiparetic walking. Arch Phys Med Rehabil
2007; 88(9): 1127–i).
28. Chen G, Patten C. Joint moment work during the stance-to-swing transition in hemiparetic subjects. J Biomech
2008; 41(4): 877–883.
29. Mulroy S, Gronley J, Weiss W, et al. Use of cluster analysis for gait
pattern classification of patients in the early and late recovery phases following stroke
. Gait Posture
2003; 18(1): 114–125.
30. Olney SJ, Richards CL. Hemiparetic gait
. Part I: characteristics. Gait Posture
1996; 4: 136–148.
31. Allen JL, Kautz SA, Neptune RR. Step length asymmetry is representative of compensatory mechanisms used in post-stroke
hemiparetic walking. Gait Posture
2011; 33(4): 538–543.
32. Arch ES, Stanhope SJ. Passive-dynamic ankle-foot orthoses substitute for ankle strength while causing adaptive gait
strategies: a feasibility study. Ann Biomed Eng
2015; 43(2): 442–450.
33. 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(2 Pt 1): 784–790.
34. Condie DN. The modern era of orthotics. Prosthet Orthot Int
2008; 32(3): 313–323.
35. Klapsing GM, Marin TJ, Parreño EM. Applied biomechanics: footwear industry. J Biomech
2010; 43: S26–S27.
36. Bregman DJJ, Rozumalski A, Koops D, et al. A new method for evaluating ankle foot orthosis characteristics: BRUCE. Gait Posture
2009; 30(2): 144–149.
37. Harlaar J, Brehm M, Becher JG, et al. Studies examining the efficacy of ankle foot orthoses should report activity level and mechanical evidence. Prosthet Orthot Int
2010; 34(3): 327–335.
38. Dufek JS, Neumann ES, Hawkins MC, O'Toole B. Functional and dynamic response characteristics of a custom composite ankle foot orthosis for Charcot-Marie-Tooth patients. Gait Posture
2013; 39(1): 308–313.
39. Schrank ES, Stanhope SJ. Dimensional accuracy of ankle-foot orthoses constructed by rapid customization and manufacturing framework. J Rehabil Res Dev
2011; 48(1): 31–42.
40. Schrank ES, Hitch L, Wallace K, Moore R, et al. Assessment of a virtual functional prototyping process for the rapid manufacture of passive-dynamic ankle-foot orthoses. J Biomech Eng
2013; 135(10): 101011–101017.
41. Stanhope SJ, Schrank ES. Process and System for Manufacturing a Customized Orthosis. Patent: United States Patent No. 8,538,570, 2013.
42. 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(1): 148–153.
43. Russell Esposito E, Blanck RV, Harper NG, et al. How does ankle-foot orthosis
stiffness affect gait
in patients with lower limb salvage? Clin Orthop Relat Res
2014; 472(10): 3026–3035.
44. Kobayashi T, Leung AKL, Akazawa Y, et al. Design of a stiffness-adjustable ankle-foot orthosis
and its effect on ankle joint kinematics in patients with stroke
. Gait Posture
2011; 33(4): 721–723.
45. Harper NG, Esposito ER, Wilken JM, Neptune RR. The influence of ankle-foot orthosis
stiffness on walking performance in individuals with lower limb impairments. Clin Biomech (Bristol, Avon)
2014; 29(8): 877–884.
46. Bregman DJJ, van der Krogt MM, de Groot V, et al. The effect of ankle foot orthosis stiffness on the energy cost of walking: a simulation study. Clin Biomech (Bristol, Avon)
2011; 26(9): 955–961.
47. Crabtree CA, Higginson JS. Modeling neuromuscular effects of ankle foot orthoses (AFOs) in computer simulations of gait
. Gait Posture
2009; 29(1): 65–70.
48. Plummer P, Behrman AL, Duncan PW, et al. Effects of stroke
severity and training duration on locomotor recovery after stroke
: a pilot study. Neurorehabil Neural Repair
2007; 21(2): 137–151.
49. Holden J, Stanhope SJ. The effect of variation in knee center location estimates on net knee joint moments. Gait Posture
1998; 7(1): 1–6.
50. Winter DA. Biomechanics and Motor Control of Human Movement. 4th ed. Hoboken: Wiley; 2009.
51. Pierrynowski MR, Galea V. Enhancing the ability of gait
analyses to differentiate between groups: scaling gait
data to body size. Gait Posture
2001; 13(3): 193–201.
52. Takahashi KZ, Lewek MD, Sawicki GS. A neuromechanics-based powered ankle exoskeleton to assist walking post-stroke
: a feasibility study. J Neuroeng Rehabil
2015; 12(1): 1–13.
53. Arch ES, Stanhope SJ, Higginson JS. Passive-dynamic ankle-foot orthosis
replicates soleus but not gastrocnemius muscle function during stance in gait
: insights for orthosis prescription. Prosthetics and Orthotics International
. [Epub ahead of print].
54. Telfer S, Pallari J, Munguia J, et al. Embracing additive manufacture: implications for foot and ankle orthosis design. BMC Musculoskelet Disord
2012; 13: 84.
55. Patzkowski JC, Blanck RV, Owens JG, et al. Can an ankle-foot orthosis
change hearts and minds? J Surg Orthop Adv
2011; 20(1): 8–18.