Impact on metabolic rate

7.1.1 Testing protocols

To validate the walking assistant function of our device, we set up a treadmill which enables walking test at constant speed of 4.3 km/h. The test has four scenarios:

(a) wearing the knee braces only; (b) wearing the device while the cable is loosened;

(c) wearing the device while the power of device is randomly turned on or off; (d) wearing the device, fully functioned. Each test lasts 7 minutes. During the test, the metabolic rate of the tester is measured and recorded with Cosmed K4b2, a portable pulmonary gas exchange measurement system. The data in scenario (a) is the reference, and that in scenario (b) shows the effect of the load. Scenarios (c) and (d) show the impact from the device. Ｔhe device senses the force that the limbs pull the slider when swinging backward and how fast the motor is driven. The obtained reading is important to start actuation. The straddle cables are pre-tensioned with 2 kgw force while wearing, to ensure that the motor can be pulled by tester’s limbs while walking. Also, the attachment on the waist should be adjusted to make sure that the cable is parallel to the sagittal plane.

The metabolic rate (MBR) during walking is evaluated by measuring the O2 and CO2 gas exchanges between the subject and the environment. We filter the data through a half-minute moving average.

7.1.2 Metabolic rate reduction

Figures 7-1(a)~(f) and Table 7-1 show the 7-min results of the four scenario of 6 testers. The solid lines show the results of scenario (a), the dash lines represent scenario (b), the dotted lines represent scenario (c), and the lines with square dot represent scenario (d). Commonly, the average metabolic rates of the testers increase from 2% to 7% in the scenario (b), showing that the load effect. In the scenario (c), the metabolic rates are either high or lower than that in scenario (b) from tester by tester since the nature gait pattern differs from person to person and is either assisted or constrained by the device passively. In the scenario (d), the increase of metabolic rates of testers are lower than the result in scenario (b) except for the tester #6, generally showing a proof-of-concept result that the assistance from our device is able to help wearer walking easier.

Figure 7-1 Measured metabolic rates (MBR) of six testers

Table 7-1 Change of MBR in scenarios (b)-(d) with respect to the result of scenario (a)

-30.00%

-20.00%

-10.00%

0.00%

10.00%

20.00%

30.00%

40.00%

#1 #2 #3 #4 #5 #6

Loaded Power-off Assisted

7.2 Effect on gait

7.2.1 Testing protocols

To observe how our device affects human gait, we performed another experiment.

As shown in Figure 7-2, testers in this experiment are asked to walk along a straight line freely with three markers pasted on each outward side of his/her shoulder, hip, and knee while the whole process are recorded via a camera 5.3 meters away, pointing to the mid-point of the straight line. There are three scenarios in the experiment: (a) Tester wears the device correctly with straddle cable pre-tensioned with 2 kgw force while the power of the device is on. (b) Tester wears the device in the way as the same as that in (a) while the power of device is off. (c) Tester wears the device while the straddle cable are loosened.

The gait patterns are recorded at the frame rate of 30 fps. Each frame is analyzed via image processing to track the positions of three markers. The angle formed by the markers are measured, transformed, de-trended, and at last separated through every two heel strikes (one gait cycle).

Figure 7-2 Diagram of set up of experiment to observe the effect on gait

7.2.2 Gait variation

Figures 7-3 (a) to (h) show the experimental result of 7 testers. The solid lines show the results of scenario (a) (assisted), the dash lines are for scenario (b) (power off); the dotted lines are for scenario (c) (free). Comparing to the result of free-walking in scenario (c), the power-off device added a passive damper to the testers in swinging thigh backward in scenario (b). To compromise this effect, testers tend to either lessen the hip extension or to extend the duration of hip extension. As a result, there is a phase lag for the tester to reach the maximum hip extension in scenario (b). In scenario (a), the damping effect of the device is minimized since the slider will move and return to near center after being pulled by wearer rather than remaining its last position.

Therefore, the phase lag of the maximum hip extension is eliminated. In addition, the hip flexion motion from θhip = -10° to θhip =0°, which is assisted by the device, is accelerated as expected. Except the changes mentioned, the device shows no obvious, unexpected effect on gait.

Figure 7-3 Gait data of tester #1 ….

Figure 7-4 Gait data of tester #2 ….

Figure 7-5 Gait data of tester #3 ….

Figure 7-6 Gait data of tester #4 ….

Figure 7-7 Gait data of tester #5 ….

Figure 7-8 Gait data of tester #6 ….

Figure 7-9 Gait data of tester #7 ….

Figure 7-10 Summarized gait data of seven testers ….

Chapter 8 Conclusion and future works

8.1 Conclusion

In this thesis, we have presented a design of walking assistance device. The device is capable of generating reciprocating linear motion within stroke length of 9 cm with the maximum force of 192N and the maximum speed of 34.22 cm/s to withdraw the cable connecting waist and knee brace, providing assistive torque on the hip joint for hip flexion. The linear motion is achieved by a timing-belt transmission train with gear ratio of 3 and a pulley system. With the low-gear-ratio transmission and limited stroke length, the device is mechanically constrained and is able to be overridden by wearer when emergency. After analyzing known gait patterns, we have designed a logical control strategy. Sensors like arranged photo-interrupters, a minimized load cell, and the speed sensor built-in in the motor are used to provide necessary position, force, and speed signals feedback for control. The load cell is compatible with straddle cable and capable of withstanding more than 200 N tensile force, with the measurement range of 100 N and at 1.755 N resolution.

To validate the effect of walking assistance, we performed 28-minute walking tests under constant walking speed of 1.2m/s, with 6 testers involved. By comparing the four 7-minute test scenarios with different walking conditions, we found that the assistive force is capable of reducing 2%~20% of tester’s metabolic rate, while carrying the 4.3kg device increases 2% to 6.96% of metabolic rate, referencing to test results of walking freely with wearing only knee brace.

To observe how the device affects the gait, we performed walking tests with another 7 testers, with gait patterns recorded via camera. As a result, the swing phase of hip flexion motion from θhip = -10° to θhip =0° is accelerated with the assistance of the device, and the maximum hip extension is slightly delayed when the power of the device is off. Both changes observed meet the expectation.

8.2 Future works

Despite the validation result, since the control strategy is simple and a nearly constant assistive force is generated during actuation, the assistance of our device causes discomfort to tester at the beginning of actuation. As the end of actuation, there are overshoot of position due to the angular momentum of the motor, which also causes instability of the device. Theoretically, with the incremental encoder built-in in the motor and the photo-interrupters, a reliable and accurate positioning is achievable, enabling a smoother actuation with modified position trail of slider.

Though the load cell is available, the effect of drift is critical, causing an unstable reference of force reading for feedback control. Since the quality of circuitry for load cell is also related to the stability, a finely-welded circuity might help giving a more stable reference. Then, with a reliable force sensor, force feedback control is possible.

When validating the powertrain, we notice that there is a shortage in power of motor to meet the expected actuation speed when walking at a faster speed. However, the maximum force requirement is satisfied. Hence, either reducing the gear ratio of the powertrain or using a more powerful motor can help.

According to current design, there is an under-design in the powertrain, in which the fixation between the driving pulley and the timing belt pulley is not strong enough to keep the two pulleys synchronize when shock. That is, slip often occurs between the two pulleys, causing failure of force transmission without frequent maintenance.

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