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Technical Brief

An Experimental Powered Lower Limb Prosthesis Using Proportional Myoelectric Control

[+] Author and Article Information
Stephanie Huang

Human Neuromechanics Laboratory,
University of Michigan,
401 Washtenaw Avenue,
Ann Arbor, MI 48109;
Department of Biomedical Engineering,
University of Michigan,
Ann Arbor, MI 48109
e-mail: shuangz@umich.edu

Jeffrey P. Wensman

University of Michigan Orthotics and Prosthetics Center,
2850 South Industrial Highway,
Suite 400,
Ann Arbor, MI 48104;
Department of Physical Medicine and Rehabilitation, University of Michigan,
Ann Arbor, MI 48109

Daniel P. Ferris

Human Neuromechanics Laboratory,
University of Michigan,
401 Washtenaw Avenue,
Ann Arbor, MI 48109;
Department of Biomedical Engineering,
University of Michigan,
Ann Arbor, MI 48109;
School of Kinesiology,
University of Michigan,
Ann Arbor, MI 48109

1Corresponding author.

Manuscript received August 31, 2013; final manuscript received January 28, 2014; published online March 7, 2014. Assoc. Editor: Venketesh N. Dubey.

J. Med. Devices 8(2), 024501 (Mar 07, 2014) (5 pages) Paper No: MED-13-1205; doi: 10.1115/1.4026633 History: Received August 31, 2013; Revised January 28, 2014

One way to provide powered lower limb prostheses with greater adaptability to a wearer's intent is to use a neural signal to provide feedforward control of prosthesis mechanics. We designed and tested the feasibility of an experimental powered ankle-foot prosthesis that uses pneumatic artificial muscles and proportional myoelectric control to vary ankle mechanics during walking. The force output of the artificial plantar flexor muscles was directly proportional to the subject's residual gastrocnemius muscle activity. The maximum force generated by a pair of artificial muscles fixed at nominal length was 3513 N. The maximum planter flexion torque that could be generated during walking was 176 Nm. The force bandwidth of the pneumatic artificial muscles was 2 Hz. The electromechanical delay was 33 ms, the time to peak tension was 48 ms, and the half relaxation time was 50 ms. We used two artificial muscles as dorsiflexors and two artificial muscles as plantar flexors. The prosthetic ankle had 25 deg of dorsiflexion and 35 deg of plantar flexion with the artificial muscles uninflated. The intent of the device was not to create a commercially viable prosthesis but to have a laboratory prototype to test principles of locomotor adaptation and biomechanics. We recruited one unilateral transtibial amputee to walk on a treadmill at 1.0 m/s while wearing the powered prosthesis. We recorded muscle activity within the subject's prescribed prosthetic socket using surface electrodes. The controller was active throughout the entire gait cycle and did not rely on detection of gait phases. The amputee subject quickly adapted to the powered prosthesis and walked with a functional gait. The subject generated peak ankle power at push off that was similar between amputated and prosthetic sides. Our results suggest that amputees can use their residual muscles for proportional myoelectric control to alter prosthetic mechanics during walking.

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References

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Figures

Grahic Jump Location
Fig. 1

The powered prosthesis in three different ankle positions. Two artificial muscles acted as dorsiflexors, and two artificial muscles acted as plantar flexors. The range of motion of the ankle with uninflated actuators was 25 deg of dorsiflexion and 35 deg of plantar flexion. The prosthetic socket was interchangeable so that the amputee's prescribed socket could be used. Standard stainless steel prosthetic components above the ankle and below the socket interface allowed for proper alignment. The ankle was a modified Rampro Swim Ankle (Rampro, Oceanside, CA). The foot was a modified Seattle LiteFoot (Trulife, Dublin, Ireland).

Grahic Jump Location
Fig. 2

Tension and length data from a pair of muscles during isometric benchtop testing at 20, 22, 24, 26.5, and 27.5 cm lengths and 25%, 50%, 75%, and 100% of maximum pressure (approximately 1.6, 3.1, 4.7, and 6.2 bar). The functional length of the artificial plantar flexor muscles during walking was 23.5–27.5 cm, which is approximately 85–100% of nominal length. We calculated he functional length as the range of muscle length measured during walking using three-dimensional kinematics.

Grahic Jump Location
Fig. 3

Force bandwidth and phase lag of a pair of pneumatic artificial muscles in an isometric benchtop configuration at nominal muscle length (27.5 cm). The input was a sinusoid signal with peak-to-peak amplitude of 10 V. Point characters on the plots show the frequencies that data were recorded. The bandwidth of the artificial muscles was 2.0 Hz. At 2.0 Hz, the output lags the input by 39 deg.

Grahic Jump Location
Fig. 4

Electromechanical response times of a pair of pneumatic artificial muscles in an isometric benchtop configuration at nominal muscle length (27.5 cm). The input signal was a 5 ms square pulse whose amplitude produced a control signal with a 10 V peak. Vertical lines indicate time points used to calculate electromechanical delay, time to peak tension, and half relaxation time. EMD was the time from the start of the square pulse to the onset of artificial muscle force development (three standard deviations above baseline force). TPT was the time from the onset of force development to when peak force was achieved. HRT was the time from peak force to when the force dropped 50% from peak force.

Grahic Jump Location
Fig. 5

Residual gastrocnemius EMG control signal during walking. Each plot shows ten consecutive strides at 1 min, 5 min, and 30 min of walking. Black lines show mean of ten strides. Shaded regions show ± two standard deviations. Horizontal lines show the 1.2 V offset to achieve the ankle set-point stiffness. Vertical lines show the average toe-off timing of ten strides.

Grahic Jump Location
Fig. 6

Comparison of ankle angle, ankle moment, and ankle power for prescribed and powered prosthesis. Ankle angles, moments, and powers were calculated from ten consecutive cycles. For the powered prosthesis, ten consecutive cycles starting at the 30 min time point were used. The proportional EMG control signal for these ten cycles is shown in Fig. 6. Outlined regions (intact side) and shaded regions (prosthetic side) show ± two standard deviations about the mean. Vertical lines show the average toe-off timing for intact and prosthetic sides.

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