The present research aims to develop a high-dorsiflexion assistive robotic system and to verify the potential of the system to be applied in assisting gait rehabilitation with intervention experiments on healthy participants with dorsiflexion movement restriction.
Requirements of the High-dorsiflexion Assistive Robotic System
The main requirements for the system were set as follows:
- The weight of the lower-limb part should not be higher than 0.45 kg, and the weight of the waist part should not be higher than 1 kg.
- The maximum output torque for ankle dorsiflexion assistance should be at least 25 Nm.
- The setup time on a user should be shorter than 3 minutes.
For our previous prototype, the weight of the lower-limb part was 0.56 kg and that of the waist part was close to 2 kg [20]. A pilot study on 6 stroke patients was conducted, and the potential for improving dorsiflexion movement was indicated. However, identical to other studies, the participants were chronic phase patients. Although no negative effects resulting from the weight of lower-limb part were reported from our previous study, the weight requirement was set even lower as 1% of an adult’s weight. Assuming an adult could be as light as 45 kg, the maximum weight of the lower-limb part was set as 0.45 kg. Meanwhile, most related RAFO devices and our previous prototype deployed the control unit at the waist band or pocket to avoid excess load on the lower-limbs. Conversely, the weights at waist parts are all at least 1.0 kg, which might also be a burden for subacute patients. Therefore, a requirement of a maximum 1.0-kg waist part is crucial.
A previous study reported that an external output torque of approximately 13.5 Nm was required to assist dorsiflexion movement from a drop foot posture [22]. Considering factors of individual differences and our application for subacute patients for full dorsiflexion assistance during gait rehabilitation, the output torque of the system was set close to double, at 25 Nm.
Last, the setup time on the user was shorter, and the physical and mental burden on patients was lower, especially for those in the subacute phase. Thus, an extremely short setup time of 3 minutes was set, including putting the system on the users, calibration of the sensor unit, and assistance dorsiflexion angle of the system.
System Design
The system assists dorsiflexion movement upon the swing phase with contraction of a McKibben-type artificial muscle aligned between the knee and forefoot. The whole system is shown in Fig. 1. The weight of the lower-limb part is 0.35 kg, and that of the waist part is approximately 1.0 kg, which barely fulfills the weight requirement set in the previous session.
Insert Figure 1
A. Lower-limb Part
The lower-limb part consists of a pressure sensor (FSR-402, Interlink Electronics, USA), an artificial muscle (DMSP-10, Festo Inc., Germany), and a tension spring. The pressure sensor was attached at the inner part of the forefoot. The pressure sensor was deployed at the inner side of the forefoot, which is applied to measure the pressure of the first metatarsal point. Details of the control method with pressure data were explained in Session C. The artificial muscle was fixed on a knee pad, and the tension spring is aligned with the artificial muscle between the knee and forefoot. The spring was applied to support heel rocker function during the loading response phase of gait, and the details were reported in research on the previous prototype [21].
A simplified shank model, depicted in Fig. 2, is applied to check whether the assisted dorsiflexion torque matches the main requirement set in the previous section. Displacement of the tension spring during the swing phase is small enough to be neglected. The equation of assisted dorsiflexion torque is
![](https://myfiles.space/user_files/58890_add8f4303ffe25fa/58890_custom_files/img1600076791.png)
According to research on the physical parameter measurements of stroke patients aged 50 to 99 years old in Japan [23], was set as 0.4 m. was set as 0.15 m from the measurement of the length between the ankle joint to the forefoot attachment point. was set as due to a plantarflexion angle at the pre-swing phase of gait, assuming is close to at neutral posture [19]. The maximum pulling force of the selected artificial muscle is 630 N. Therefore, the maximum assisted torque can be approximately 75.12 Nm, which is much larger than the requirement set in the previous section.
Insert Figure 2
A small piece of 2-mm-thick cow leather was glued on the shoe, and three right angle brackets were attached on the leather from the middle to the outer side of the forefoot. The cow leather prevents deformation caused from pulling of artificial muscle, and the three placements of brackets provide choices for simultaneous extroversion assistance caused from different extents of introversion posture [2, 3]. It is extremely simple and quick to put the lower-limb part on the user’s body. First, the artificial muscle is attached to the knee pad with magic tape. Second, the spring is hooked on the bracket. Last, the cable between the artificial muscle and the spring is adjusted for a sufficient dorsiflexion assistance angle. The angle was set between and the range of motion for the user of the dorsiflexion posture referenced from a healthy people’s dorsiflexion angle [19]. The total time for setting the lower-limb part on a healthy person would not be longer than 2 minutes. It is worth noting that dorsiflexion movement is assisted with full contraction of the artificial muscle, and the assisted dorsiflexion angle was set under artificial muscle’s full contraction condition. Therefore, this physical constraint of contraction length makes our system much safer than motor-actuated robotic devices. Meanwhile, excluding the right angle brackets, none the components of lower-limb parts are rigid, and the artificial muscle assists dorsiflexion movement by directly pulling the forefoot. This design not only realizes an extremely low-weighted prototype but also prevents possible joint mismatch and discomfort issues for exoskeleton-type devices [24].
B. Waist Part
The waist part consists of a microcontroller with an embedded wireless module (nRF51822, Nordic Semiconductor, Norway), two solenoid valves for carbon dioxide gas injection and release to the artificial muscle (BV214A, Mac Valves Inc., USA), a rechargeable battery compatible with medical applications (RRC1120, RRC power solutions GmbH, Germany), and a portable gas cylinder containing 74 g of carbon dioxide gas. The microcontroller unit with the battery was attached on the waist band with magic tape, and the valves and gas cylinder were placed in separate back pockets. Gas is transmitted through 6-mm-diameter tubes.
The battery is able to provide power to the system consecutively for approximately 3.5 hours, which is apparently sufficient even for a chronic stage patient. The carbon dioxide gas cylinder is enough for at least 50 full contractions of the artificial muscle through practical tests; that is, at least 50 dorsiflexion movements could be assisted through the air source of one portable cylinder. The step length of a subacute stroke patient is approximately 0.3 m [25]. With our previous experience, even some chronic patients require rest after one 10-m level ground walking. It contains approximately 13 swing phase dorsiflexion movements, which means the air source from one gas cylinder should sufficient. The cylinder can be replaced with a new one when the patient rests.
The setting of the waist part on users is also simple. Securing the part to the waist with magic tape, activating control unit, and connecting the tube to the artificial muscle should take less than 30 seconds. Therefore, the total setup time on a user for the whole system was less than 2.5 minutes by adding the setup times of the lower-limb part and the waist part. This meets the requirement for setup time. However, the current result was acquired from healthy people. The actual setup time on subacute stroke patients will be monitored in a future pilot study.
C. System Configuration and Control Method
Fig. 3 shows the configuration of our high-dorsiflexion assistive system. The microcontroller received pressure sensor data through wireless communication and determined the opening and closing of the solenoid valves for air injection and release. The air source was provided by the carbon dioxide gas cylinder to the artificial muscle through solenoid valves.
Insert Figure 3
Fig. 4(a) depicts the control flow of our system. The artificial muscle contracts when no foot pressure is detected, which means dorsiflexion support is activated when the gait phase transits from the stance to the swing phase. It has to be noted that only one pressure sensor was placed on the inner forefoot side of the insole, so the artificial muscle does not extend until the mid-stance phase when foot flat occurs. Heel rocker function, the resistive dorsiflexion mechanism in the loading response phase of gait, is assisted by the tension spring, as shown in Fig. 4(b) [21]. This mechanism prevents the requirements for suitable assistance timing and speed for heel rocker function and enables a simple control method by merely monitoring the pressure sensor. Although numerous complicated control strategies have been developed for assisting gait rehabilitation, it is agreed that a simple control method is suitable for practical clinical application [13].
Insert Figure 4
Experimental Design
The purpose of this experiment is to verify the actual intervention effects on users’ physiological and kinematic characteristics with the developed high-dorsiflexion assistive system. Ethical approval of this experiment was granted by the Ethics Committee of Waseda University.
A. Participants with Dorsiflexion Restriction
Six healthy and young people without gait disabilities participated in the experiment. Simple information of the participants is shown in Table 1. It is important to conduct trials before clinical applications, especially for subacute patients who require more safety concerns. To make participants gait characteristics close to those of stroke hemiplegic patients, they were requested to wear a dorsiflexion movement-restricted ankle-foot orthosis, as illustrated in Fig. 5. This orthosis applies a tension spring above the back of the calf with a coefficient of 10.39 N/mm. The spring was set at nominal length when the orthosis was at an angle of least of plantarflexion posture. Thus, a person who wears this orthosis should exert more dorsiflexion torque than usual and the decreases in MTC and compensatory movement should reflect those of stroke patients.
Table 1. Characteristics of the experiment participants
|
Sex
|
Age
|
Height (cm)
|
Weight (kg)
|
A
|
Male
|
30
|
172
|
73
|
B
|
Male
|
26
|
175
|
75
|
C
|
Male
|
24
|
174
|
62
|
D
|
Male
|
25
|
172
|
67
|
E
|
Male
|
23
|
180
|
66
|
F
|
Male
|
25
|
170
|
80
|
Insert Figure 5
B. Experimental Settings
Interventions were conducted on the right legs for all participants. They were requested to walk 5 m six times under all the following three conditions with their own preferred speed:
a) Normal walk (NOR): The participants wore the orthosis, but without restriction of dorsiflexion movement.
b) Dorsiflexion restriction (DFR): The participants wore the orthosis, and the spring was applied for dorsiflexion movement restriction.
c) Assistance (AST): The participants wore the orthosis with dorsiflexion movement restriction, and the high-dorsiflexion assistive system was applied to support dorsiflexion in the swing phase.
The sEMG data of the tibialis anterior muscle were collected with the TrignoTM Wireless System (Delsys Inc., USA) with a sampling rate of 2000 Hz. Meanwhile, the gait kinematic data were recorded by the motion capture system (MAC3D, NAC Inc., Japan) with a sampling rate of 100 Hz. The marker set of this experiment is shown and explained in Fig. 6. Foot pressure data were also collected by force plates (Advanced Manufacturing Technology Inc., USA) to determine the stance and swing phases of gait.
Insert Figure 6
C. Evaluations
The physiological characteristics were evaluated with the collected sEMG data and were analyzed by the software EMGworks (Delsys Inc., USA). The data were filtered through an infinite impulse response bandpass filter with cut-off frequencies equal to 100 Hz to 400 Hz. The filtered data were then processed with mean absolute value calculation, and the index was accessed by integrating the processed data during ankle dorsiflexion movement in the swing phase. This index evaluates the extent of self-activating dorsiflexion movement under each condition. For realization of full dorsiflexion assistance for passive gait training, a small sEMG signal during our system’s assistance was expected.
For kinematic characteristics, three aspects of evaluation indices were determined. The first aspect was the direct effect on ankle kinematics. As Fig. 7(a) shows, the ankle angle was calculated with the markers RVMH, RANK, and RKNE. Dorsiflexion angle was defined as the minimum ankle angle in the swing phase, and improvement of dorsiflexion angle in AST condition compared with DRF condition was expected. The other aspects were stumbling risk and compensatory gait patterns. The stumbling risk was evaluated with MTC and registered from the minimum height of the RTOE marker in the swing phase of gait, which is also shown in Fig. 7(a). The compensatory gait patterns were divided into two types: coronal plane and sagittal plane. For compensatory gait patterns in the coronal plane, as Fig. 7(b) indicates, circumduction gait was frequent for stroke patients. The lateral pelvis tilt angle , calculated with the angle between the floor and vector from the markers LASIS to RASIS, and swing width , calculated with the horizontal length of marker RTOE in the swing phase, were evaluated. For compensatory gait patterns in sagittal plane, as shown in Fig. 7(a), excess hip flexion was frequent for stroke patients. Since an increase in knee height is an obvious sign of excess hip flexion, RKNE marker’s height in the swing phase was evaluated. It was expected that either a small MTC or inclination of the compensatory gait pattern would be revealed while ankle dorsiflexion was restricted, and improvement could be observed during assistance using our high-dorsiflexion assistive system.
Insert Figure 7