Analysis of Kinetic Load of Plastic Ankle Foot Orthosis during Swing Phase

Transcription

1 Advanced Experimental Mechanics, Vol.2 (2017), Copyright C 2017 JSEM Analysis of Kinetic Load of Plastic Ankle Foot Orthosis during Swing Phase Daisuke MORIOKA 1, Ichiro KITAYAMA 2, Tsutomu NISHIGAKI 2, Masato KITANO 1, Takashi YAMANAKA 1, Yosuke IKEHARA 1, Hideyo KOYAMA 3, Takashi MORIMOTO 3, Hideki SONOBE 3 and Noriyuki MIYAZAKI 3 1 Department of Biological System Engineering, Graduate School of Biology-oriented Science and Technology, Kindai University, Wakayama , Japan 2 Department of Biomechanical and Human Factors Engineering, Faculty of Biology-oriented Science and Technology, Kindai University, Wakayama , Japan 3 Kotonoura Rehabilitation Center Hospital, Wakayama , Japan (Received 20 January 2017; received in revised form 2 May 2017; accepted 18 May 2017) Abstract: We developed a mechanical measurement system, which we then used in this PAFO gait experiment. We report that differences between hemiplegic patients and healthy subjects in the mechanical load acting on gait using a PAFO in swing phase. Moreover, there is a large difference between volunteers and patients in initial swing phase and especially in decreasing belt reaction force toward the end of terminal swing phase. We consider that this difference is relevant to control of voluntary movement. In conclusion, a large deformation can be observed between volunteers and patients by analyzing the belt reaction force and load waveform. Keywords: Force-torque sensor, Swing phase, Orthosis, Hemiplegia, Kinetic load 1. Introduction According to statistics provided by the United Nations, as of 2015 about 12% of the world's population is over sixty [1], and is expected to increase to 21.5% by Population aging is a serious issue in Japan; as of July 2016, 27.1% [2] of the total population is over 65. Illness and injury occur more readily in this group due to the reduced bodily functions associated with aging. According to the 2014 investigation by the World Health Organization, the top leading causes of death [3] are ischemic heart disease (IHD), cerebrovascular disease (stroke), and chronic obstructive pulmonary disease (COPD). In 2012, there were about 6.7 million deaths from stroke worldwide. Although mortality is comparatively lower for stroke than for IHD and COPD, there is a high risk of brain damage, and it has been reported that 21.7% [4] of stroke patients are in need of care, especially those with sequelae such as hemiplegia who are in need of advanced care. If lower limb failure has occurred as a result of stroke, everyday life becomes affected with markedly decreased quality of life. Methods for aiding such hemiplegics include the use of devices such as orthoses and canes as well as assistance from others. Among them, plastic ankle foot orthoses (PAFOs), a representative type of ankle foot orthosis (AFO), are usually used to assist ambulation in patients with stroke sequelae. An investigation by Takashima et al. found that approximately 75% of AFOs are fabricated from plastic and 23% from metal [5]. There are some studies, such as Kobayashi et al., that investigated the effectiveness of orthoses, including AFOs, using a gait analyzing system [6], and Yamamoto et al., who developed an ankle foot orthosis with dorsiflexion assist [7]. However, there are very few studies that have analyzed mechanical load acting on AFO during swing phase. The goal of this study is to develop a new form to add comfort to PAFO. In the process, we developed an analysis system with force-torque sensors that are able to analyze the mechanical load acting on the AFO during orthotic gait. In this study, we report differences found between hemiplegic patients (patients) and healthy subjects (volunteers) in mechanical load acting on PAFO gait during swing phase. 2. Experimental Environment 2.1 Experimental device Figure 1 shows the mechanical measurement system used in this PAFO gait experimental. This system has three 6-axis force-torque sensors (6-axis sensor) (United States, ATI, Mini-45-E), one 3-axis force sensor (Japan, Tec Gihan, USL06-H5-100N), and one flexible goniometer (United Kingdom, Biometrics Ltd, SG150). We installed the 6-axis sensors at three locations, two on the sole of the foot and one on the posterior calf, and the 3-axis force sensor in one location, on the device ankle belt. The two 6-axis sensors on the sole of foot were used for identification of gait cycle, as an alternative to the force plate. The 6-axis sensor on the posterior calf, and the flexible goniometer, can be used to analyze the mechanical load applied to the PAFO covering, but this was not performed in the current gait experiment. The 3-axis sensor is used to analyze ankle belt load measurement, and is the only sensor that can measure the applied mechanical load during swing phase. The data from the 6-axis sensors at the soles of the feet passed through a DAQ device (Data collecting device of LabVIEW, Japan, National Instruments Japan Corp.) and were recorded on a personal computer, while the data from the 3-axis force sensor passed through an amplifier and an AD converter before being stored on the PC. Major forces applied to orthoses during swing phase work on (1) calf, (2) ankle belt and (3) orthotic inner sole surface. While all of these data are needed to analyze perfectly during swing phase, it is difficult to measure forces about (3) which is applied between orthotic sole inner surface and human sole. Here, listed above (1) (2) and (3) are so-called three-point support system which is a basic fixation method used in clinical practice, and we decide to make research to forces applied on (2) which have a key role in orthotic fixation. 147

2 D. MORIOKA, I. KITAYAMA, T. NISHIGAKI, M. KITANO, T. YAMANAKA, Y. IKEHARA, H. KOYAMA, T. MORIMOTO, H. SONOBE and N. MIYAZAKI LabVIEW 2013 was used as the data storage software, and the recording frequency was 1000 Hz. Elements of the load vector follow the right-handed orthogonal coordinate system shown in Fig.2. We do not use data about the 6-axis sensor on the posterior calf, because 3-axis sensor on the ankle belt show the major result than any other sensors during swing phase, and we have reported about the results of that sensor on a couple of papers [8, 9] which are written by our project. Table 1 Subjects information Subject Patient A Patient B Patient C Patient D Age Sex Female Male Male Female Body Weight [kg] Subject Patient E Patient F Volunteer G Volunteer H Age Sex Male Male Male Male Body Weight [kg] Subject Volunteer I Volunteer J Volunteer K Volunteer L Age Sex Male Male Male Male Body Weight [kg] Table 2 Brunnstrom stage of each subjects Fig.1 Measurement system Rf : Toe vertical force Rh : Heel vertical force Fx : Medial direction force Fy : Under direction force Fz : Belt reaction force Fig. 2 Axial direction of each component of the six-axis force torque sensor and of the 3-axis sensor 2.2 Subjects Twelve subjects participated (average age 42.9 year, average weight 66.9 kg) in this study. Six participants have left paralysis due to cerebral hemorrhage (average age 57.0 years, average weight 68.4 kg) and six participants were healthy volunteers (average age 28.8 years, average weight 66.9 kg). Subject information is shown in Table 1 and patient Brunnstrom stage (Brs) the level of spasticity of upper extremity (U/E), finger, lower extremity (L/E) is shown in Table 2. Brunnstrom stage indicates level of paralysis in the extremities and is often used for functional patient evaluation. The Brs sets out a sequence of stages of recovery from hemiplegia after a stroke where Stage I (loss of voluntary movement) is most severe. All PAFOs were custom-made by a prosthetist. Ethical approval was obtained from the ethics committee of Kindai University Graduate School of Biology-Oriented Science and Technology (H ). Informed consent was obtained from all participants. Patient A Patient B Patient C U/E Ⅴ Ⅲ Ⅲ finger Ⅳ Ⅲ Ⅲ L/E Ⅴ Ⅳ Ⅲ Patient D Patient E Patient F U/E Ⅳ Ⅲ Ⅲ finger Ⅴ Ⅲ Ⅱ L/E Ⅳ Ⅳ Ⅱ-Ⅲ Level of spasticity of upper extremity (U/E), finger, lower extremity (L/E) I (severely) < II < III < IV < V 2.3 Experimental method Five trials of walking tests were used in healthy volunteers. Number of tests and walking speed are adjusted according to the patient's symptoms. Cadence 90 and 20 cm step length were used for volunteers, and in hemiplegic patients were set according to individual symptoms. 3. Results Ground reaction force (Fig.2, Rf + Rh) of one gait test is shown for patient E (Fig.3a) and volunteer J (Fig.3b). We define the swing phase as the point where the two 6-axis sensors on the sole of foot are at near zero. The belt reaction force of ankle was determined as an evaluation index of the swing phase. Gait cycle data in the middle gait cycle was used instead of those in the initial gait cycle so that the gait cycle data would be time-homogeneous. To remove individual differences, time was normalized to swing duration, with 0% being toe-off and 100% representing heel contact. Ground reaction force was normalized to the subject's weight. Measured results are shown in Fig.4, which is a sample graph of the belt reaction force in swing phase. Figure 4a shows belt reaction force in one swing phase for Patient E. Figure 4b shows belt reaction force in one swing phase for Volunteer J. The graph for patient E is a shallow U-shape, 148

3 Advanced Experimental Mechanics, Vol.2 (2017) whereas is it slopes downward after a peak for volunteer J. Measured results are shown in Fig.5, which is sample graph of belt reaction force in stance phase. Figure 5a shows belt reaction force in one stance phase for patient E. Figure 5b shows belt reaction force in one stance phase for volunteer J. Both subject's belt reaction force waveforms during stance phase are clearly greatly changed. Fig.5a Sample belt reaction force in stance phase for patient E Fig.3a Sample ground reaction force for patient E Fig.5b Sample belt reaction force in stance phase for volunteer J Fig.3b Sample ground reaction force for volunteer J Fig.4a Sample belt reaction force in swing phase for patient E Table 3 shows average belt reaction force, swing phase times and stance phase times for each subject in one gait cycle of one swing phase and one stance phase. Table 4 shows total average for patient group and volunteer group. The belt reaction force data as indicated in this study are given as the difference between the maximum and minimum belt reaction force, in order to cancel out the belt tension applied when the sensor was fastened. In the volunteer group, it seems average swing phase time was about s and average belt reaction force was N/kg. However, in the patient group, average swing phase time was s, which is about 1.6 times longer than that in the volunteer group. However, average belt reaction force in the patient group is N/kg, which is about 3.5 times smaller than that in the volunteer group. The difference is up to about 3.5 times smaller than volunteer group. Table 4 shows belt reaction force in the volunteer group, which is about N/kg in stance phase and N/kg in swing phase. There is a large difference of about N/kg between these values. In the patient group, belt reaction force was about N/kg in stance phase and N/kg in swing phase, which is lower by about N/kg. This difference is about 1.7 larger than the difference seen in the volunteer group. Fig.4b Sample belt reaction force in swing phase for volunteer J 149

4 D. MORIOKA, I. KITAYAMA, T. NISHIGAKI, M. KITANO, T. YAMANAKA, Y. IKEHARA, H. KOYAMA, T. MORIMOTO, H. SONOBE and N. MIYAZAKI Table 3 Average time and belt reaction force in each phase for each subject Patient A Patient B Patient C Patient D Swing Time [s] Stance Time [s] during Swing Phase [N/kg] during Stance Phase [N/kg] Patient E Patient F Volunteer G Volunteer H Swing Time [s] Stance Time [s] during Swing Phase [N/kg] during Stance Phase [N/kg] Volunteer I Volunteer J Volunteer K Volunteer L Swing Time [s] Stance Time [s] during Swing Phase [N/kg] during Stance Phase [N/kg] Table 4 Total average of phase time and belt reaction force in each group Phase Time Swing [s] Stance [s] Patients Average Volunteers Average for Each Phase Swing [N/kg] Stance [N/kg] Patients Average Volunteers Average Discussion Load waveforms are shown in Fig.4a and 4b. Results are divided into volunteers and patients. There is a large difference between volunteers and patients in initial swing phase and especially in decreasing belt reaction force toward the end of terminal swing phase. We believe that the belt reaction force unconsciously affected the ankle, allowing the volunteers to walk without the PAFO and to perform voluntary movement. Although all patients demonstrate each type of Brs, all patients' load waveforms had the same U-shaped pattern, and the patients results in Fig.4a-5b are very different from the volunteers results. This difference is likely relevant to control of voluntary movement. Normal healthy individuals are performing voluntary movement in combination with increasing (early and medial swing phase) and decreasing (late swing phase) of planter flexion such as a literature s Fig.3-16 [10] during swing phase. We consider that data indicate dynamic reaction force such as Fig.4b which comes into existence because of restriction by the ankle best on the dorsum of a foot. Looking at the swing phase and stance belt reaction force in Table 4, patients' average belt reaction force in swing phase is N/kg and N/kg in stance phase; the difference is about N/kg (about a 230% increase). However, volunteers' average belt reaction force in swing phase is N/kg and N/kg in stance phase, with a difference of about N/kg (about a 40% increase). Hemiplegia can be considered the reason for the large difference seen in the patients, as it likely reduced the belt reaction force during the swing phase due to increased difficulty in voluntary movement. Main reasons for applied to belt reaction force are conceivable that (1) fall prevention of PAFO, (2) limit of the legs voluntary movement and (3) reaction force against mounding around the ankle joint of lower leg muscles activities in stance phase. As against force data applied by (1) are supposedly flat, force data by (2) become supposedly altered. We have a think that forces by (2) are the main data about a statistics analysis aiming at obtaining between groups such as Fig.6a and Fig.6b. Based on these results, assessed the statistical significance of the belt reaction force at each phase in the hemiplegic patients and healthy volunteers (Fig.6a, 6b) using a Student s t-test. The normality of each group was determined from the absolute value of the sum of the peakedness and skewness [11]. We also used an F-test [12] was used to determine whether there was equal variance. As indicated in Table 5, normality was assured and equal variance was confirmed for all combinations. Figure 6a graphs the belt reaction force during the swing phase and Fig.6b shows that during the stance phase. Based on these graphs, we performed a t-test [13] with a significance threshold of 5% (p<0.05). As a result, the belt reaction force during the swing phase was found to be significantly lower in the hemiplegic patients compared to the healthy volunteers, as seen in Fig.6a. However, no significant difference was seen between the two groups in the belt reaction force during the stance phase as indicated in Fig.6b. These results showed that the belt reaction force exhibited a greater change (difference) in load in the healthy volunteers compared to the hemiplegic patients during both the swing phase and the stance phase. In particular, the belt reaction force acting during the swing phase in the healthy volunteers was approximately 3.5 times greater than that in the hemiplegic patients. It is conceivable that the changes in load were greater among the healthy volunteers, as they were capable of voluntary movement. As shown in Fig. 4b, the magnitude of the load in the healthy volunteers became smaller from the late swing phase until immediately before heel contact. This is believed to result from voluntary muscle power being generated during the time that the foot was flexed dorsally. However, only the load from the foot that resulted from paralysis appeared to be supported by the orthosis among the hemiplegic patients. Nevertheless, the current study focused on patients with low spasticity, and so future experiments on patients with high spasticity will also be important. 150

5 Advanced Experimental Mechanics, Vol.2 (2017) Table 5 Confirmation of normality and dispersion Patients Swing Phase Stance Phase Peakedness Skewness Normality Test * normality * * ** disnormality F Test * equally variance * * ** unequal Volunteers Swing Phase Stance Phase Peakedness Skewness Normality Test * normality * * ** disnormality F Test * equally variance * * ** unequal * p<0.05 Fig.6a Belt reaction force during swing phase * p<0.05 Fig.6b Belt reaction force during stance phase 5. Conclusions Use of the measurement system produced in this study enabled measurements of the forces acting on the orthosis during the swing phase, which were previously unknown. A large discrepancy was observed between volunteers and patients upon analyzing the belt reaction force and load waveform using this measurement system. In spite of it is easy to suppose the result about difference between healthy subjects and patient subjects; we never had seen the paper in the world which indicates these result in so * far as our searching. We submit that it is our paper s notable advantage. In future research, we will consider the development of a new form of PAFO that allows for both good comfort and high quality of life by collecting subject data, collecting attitude variation data such as initial rise motion, using sensory evaluation, and other forms of assessment. Acknowledgment This work was partially supported by a JSPS Grant-in-Aid for Scientific Research (C) (No. 16K01579) and a grant for Strategic Research Foundation Grant-aided Project for Private Universities from the Ministry of Education, Culture, Sports, Science and Technology, Japan (MEXT), (S ). References [1] United Nations New York 2015, World Population Prospects The 2015 Revision Key Findings and Advance Tables, publications/files/key_findings_wpp_2015.pdf (accessed ). [2] Statistics Bureau Japan, Population Estimates by Age (5 Year Age Group) and Sex - July 1, 2016 (Final estimates), December 1, 2016 (Provisional estimates). [3] World Health Organization, The top 10 causes of death, May, 2014, factsheets/fs310/en/ (accessed ). [4] Ministry of Health Labour and Welfare, Overview of the National Livelihood Survey (2014). [5] Takashima, T.: Situation of the plastic ankle foot orthoses, Discussion of adaptability to correction moment and spasticity (in Japanese), The Japanese Society of Prosthetics and Orthotics (JSPO), 19-2 (2003), [6] Kobayashi, S., Kinase, A., Simanuki, S. and Fukuya, T.: Effect of ankle foot orthosis on leg movements during stance phase of walking (in Japanese), Bulletin of Tsukuba International University, 1 (2010), [7] Yamamoto, S., Ebina, M., Kubo, S., Hayashi, T., Doi, T., Kawai, H., Akita, Y. and Hayakawa, Y.: Development of an ankle foot orthosis with dorsiflexion assist (DACS AFO) (in Japanese), Bulletin of the Japanese Society of Prosthetics and Orthotics, 13-2 (1977), [8] Kitano, K., Kitayama, I., Morioka, D., Nakano, K., Yamanaka, T., Ohmasa, M., Koyama, H., Sonobe, H. and Miyazaki, N.: Analysis of plastic ankle-foot orthoses: a comparison of healthy individuals and patients (in Japanese), Bulletin of Japanese Society for Clinical Biomechanics, 37 (2016), [9] Kitano, K., Kitayama, I., Morioka, D., Nakano, K., Yamanaka, T., Ikehara, Y., Koyama, H., Morimoto, T., Sonobe, H. and Miyazaki, N.: Movement and relation between force vector and plantar/dorsiflexion angle during walking with plastic ankle foot orthoses (in Japanese), J. JSEM, 16-2 (2016), [10] Rose, J. and Gamble, J. G.: Human Walking (3rd ed.), Lippincott Williams & Wilkins, Philadelphia, United States of America (2005),

6 D. MORIOKA, I. KITAYAMA, T. NISHIGAKI, M. KITANO, T. YAMANAKA, Y. IKEHARA, H. KOYAMA, T. MORIMOTO, H. SONOBE and N. MIYAZAKI [11] Katsumi, K.: Biostatistics 2011, Shizuoka Prefectural Agriculture and Forestry College, Shizuoka, Japan, (2011), (accessed ) [12] Hisae, Y.: Statcel -The Useful Addin Forms on Excel- (3rd ed.) (in Japanese), Seiun Co., Ltd., Saitama, Japan, (2011), [13] ibid. [12],