ft* =fz/(*Do-Tm-l) (11) where, fz = axial force acting on a stenosis, / = stenosis length ( = D0). This figure shows that the non-Newtonian effect of blood decreases the value of fz* and suggests that the nonNewtonian effect of blood has an effect to prevent the vascular lesions. Present calculation results show that the characteristics of the non-Newtonian fluid are preferred for the prevention of the vascular lesions. Of course we cannot definitely conclude this and we must wait for future detailed experimental and theoretical works. References 1 Daly, B. J., " A Numerical Study of Pulsatile Flow through Stenosed Canine Femoral Arteries," Journal of Biomechanics, Vol. 9, 1976, pp. 465-475. 2 Mann, K. A., Deutsch, S., Tarbell, J. M., Geselowitz, D. B., Rosenberg, G., and Pierce, W. S., "An Experimental Study of Newtonian and NonNewtonian Flow Dynamics in a Venticular Assist Device," ASME JOURNAL OF BlOMECHANICAl ENGINEERING, V o l . 109, 1987, p p . 139-147.
3 Nakamura, M., and Sawada, T., "Numerical Study on the Flow of NonNewtonian Fluid through an Axisymmetric Stenosis," ASME JOURNAL OF
The construction of the load cell was undertaken as part of a project to develop improved means of measurement for medical applications. In this instance, the effort was specifically directed towards improving the gait training of above-knee (A/K) amputees. The system developed includes a hip-angle transducer and a portable instrument to monitor transducers, produce audio feedback, and record data for patient evaluation [1, 2, 3]. Several load cells have been designed to accommodate nonamputees. The sensors are usually made in the form of a foot-pad shoe insert, such as the Krusen limb load monitor [4, 5, 6, 7]. Their performance as load cells is quite poor, however, since the sensor does not truly integrate the pressure over the entire footpad. Measurement error can be as much as 50 percent [8]. A force plate, while more accurate, is not useful for gait training enhancement, since the measurement is not available continuously and considerable concentration may be required for the amputee to place his leg on the force plate. Our goal was to design a load cell specifically for use in A/K prostheses which had improved accuracy and was rugged enough for daily clinical use.
BIOMECHANICAL ENGINEERING, Vol. 110, 1988, pp. 137-143.
4 Einav, S., Bermann, H. J., Fuhro, R. L., DiGiovanni, P. R., Fine, S., and Friedman, J. D., "Measurement of Velocity Profiles of Red Blood Cells in the Microcirculation by Laser Doppler Anemometry (LDA)," Biorheology, Vol. 12, 1975, pp. 207-210. 5 Keentok, M., Milthorpe, J. F., and O'donovan, E., "On the Shearing Zone around Rotating Vanes in Plastic Fluids: Theory and Experiment," Journal of Non-Newtonian Fluid Mechanics, Vol. 17, 1985, pp. 23-35. 6 McDonald, D. A., Blood Flow in Arteries, Arnold, London, 1960. 7 Nakamura, M., and Sawada, T., "Numerical Study on the Laminar Pulsatile Flow of Slurries," Journal of Non-Newtonian Fluid Mechanics, Vol. 22, 1987, pp. 191-206. 8 Crochet, M. J., Davies, A. R., and Walters, K., Numerical Simulation of Non-Newtonian Flow, Elsevier, Amsterdam, 1984, pp. 234-240. 9 Lee, J. S., and Fung, Y. C , "Flow in Locally Constricted Tubes at Low Reynolds Numbers," ASME Journal of Applied Mechanics, Vol. 137, 1970, pp. 9-16. 10 Young, D. F., "Fluid Mechanics of Arterial Stenosis," ASME JOURNAL OF BIOMECHANICAL ENGINEERING, Vol. 101, 1979, pp. 157-175.
11 Wille, S. 0., "Pressure and Flow Patterns in the vicinity of an arterial stenosis," Finite Elements in Biomechanics, (ed. B. R. Simon), The University of Arizona, Tucson, 1980, pp. 295-314. 12 Fukushima, T., and Azuma, T., "Patterns of Pulsatile Flow in Arterial Models with Stenosis," ClinicalHemorheology, Vol. 2, 1982, pp. 31-41. 13 Weinbaum, S., and Caro, C. G., " A Macromolecule Transport Model for the Arterial Wall and Endothelium based on the Ultrastructual Specialization Observed in Electron Microscopic Studies," Journal of Fluid Mechanics, Vol. 74, 1976, pp. 611-640. 14 Oka, S., Biorheology, Syokabo, Tokyo, 1984 (in Japanese). 15 Nakamura, M., and Sawada, T., "Model Study on the Strain and Stress Distributions in the Vicinity of an Arterial Stenosis," Biorheology, Vol. 25, 1988, pp. 685-695.
A Robust, Accurate Load Cell for Use in Training Above-Knee Amputees Christopher Cullen3 and Woodie Flowers4 Introduction In many physical therapy situations, it is useful to have a quantitative, real-time measurement of limb load. Such measurements can be used to generate biofeedback to help improve a patient's gait. This paper describes a load cell for use with an above-knee prosthesis. It has features not common to most load cells. Consulting Engineer, Cullen Associates, Ipswich, MA 01938 Professor, Department of Mechanical Engineering, Massachusetts Institute of Technology, Cambridge, MA 02139 Contributed by the Bioengineering Division of THE AMERICAN SOCIETY OF
MECHANICAL ENGINEERS. Manuscript received by the Bioengineering Division April 10, 1987; revised manuscript received October 23, 1989.
Design A dominant issue in this design problem is the presence of large bending moments in the leg. These moments produce strains in the prosthetic leg tube which are more than an order of magnitude greater than those caused by the compression to be measured. It is essential to measure only axial loading of the leg for the purpose of useful biofeedback. These moments arise because the amputee must maintain his or her center of gravity well forward of the knee joint to prevent buckling. The chosen design uses a dual-structure approach. There are separate structures used to measure the axial compressive force and to isolate bending moments. Figure 1 shows external and internal views of the load cell. It is comprised of two concentric aluminum cylinders with a thin gap between them. The inner cylinder is clamped around the lower leg tube, the outer is fixed to the knee unit. Silicone elastomer is cast into the gap between the two cylinders. Attached to the inner cylinder is a cap with a central screw which bears against an aluminum beam fixed to the outer cylinder. Strain gages on the beam provide a measure of the beam displacement. Amplifying electronics are mounted inside the knee unit cavity. The concentric cylinders serve to isolate bending moments from the measuring structure. The elastomer has the useful property of being relatively compliant under shear loading but much stiffer in compression. Imposing a moment about an axis perpendicular to the leg tube subjects the silicone to a combined shear/compressive stress loading. The cylinders are made sufficiently long so that the compressive stiffness dominates. Applied torque deflects this stiff structure only slightly, resulting in a small sliding motion of the screw across the beam and minimal strain. A solid film lubricant was applied to the loading surface to provide wear resistance for this sliding motion. Applied axial load directly deflects the compliant axial structure, producing a much larger measured beam deflection. The dual-structure design allows selection of the load cell sensitivity independent of the strength of the knee unit itself. This is an important safety consideration. This configuration has the additional advantage that failure due to overloading, i.e. a sensing beam failure, is quite benign. The beam itself can support a 135 percent overload (the design load is 1100 newtons (250 pounds)). Should it fail, the elastomer is strong enough to support an 1800 percent overload. If the bond between the elastomer and the aluminum cylinder fails, the inner cylinder will come to rest on the knee unit after moving less than .5 cm (.2 inches). FEBRUARY 1990, Vol. 112 /103
Journal of Biomechanical Engineering
Copyright © 1990 by ASME Downloaded From: http://biomechanical.asmedigitalcollection.asme.org/ on 05/26/2018 Terms of Use: http://www.asme.org/about-asme/terms-of-use
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Load sensing element and electronics package
Output of load cell and force plate versus time, low loading rate 7.0 ....
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This particular load cell is built into a modified brand-name prosthetic knee hinge unit, commonly used in training prostheses. It is easily transportable from one prosthesis to another by knocking out the hinge pin and unclamping the leg tube. One device can therefore serve several patients undergoing gait training concurrently. The load cell was built to fit within the envelope of a normal human lower leg. It only slightly increases the shank diameter of an immediate-post-operative prosthesis. To minimize discomfort to the patient, the load cell was made to be as lightweight as possible. It adds 125 grams (4.5 ounces) just below the knee. The load cell is supplied with + 5 volts D.C. and provides a o to + 5 volt output, proportional to the applied load. The onboard electronics include a preamplifier and a passive 160 Hz low-pass filter. Circuit board potentiometers are used for offset and gain adjustments. Nominal output sensitivity is 4.2 mV IN (18.7 mV Ipound) to give 4.5 Nlbit (I pound/bit) for the 8-bit AID converter in the biofeedback controller. The load cell has a modular construction. It may be 104/VoI.112, FEBRUARY 1990
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disassembled to replace a defective element. The sensing beam is machined from an aluminum disk. The active strain gages are mounted in the center of the beam; the temperature compensation gages are mounted on the surface of the disk, beside the beam (Fig. 2). A small circuit board containing the elecTransactions of the ASME
Downloaded From: http://biomechanical.asmedigitalcollection.asme.org/ on 05/26/2018 Terms of Use: http://www.asme.org/about-asme/terms-of-use
tronics is mounted on top of the disk. Thus the leads from the gages to the preamplifier are kept extremely short, reducing electrical noise. Testing and Performance After initial calibration, a log was kept for a one-month period to test for null drift. The output of the load cell in the unloaded condition was measured over the course of the month. Maximum drift was .025 volts, representing 5.8 N (1.3 pounds). Null drift includes the effects of both elastomer creep and drift in the electronics. A short-term loaded creep test was performed. A 220 N (fifty pound) load was applied for twenty-four hours. Change in the null point was 7 mV, or 1.8 N (.4 pounds). Again, the effect of electronics drift are included in this measurement. Bending moment isolation was evaluated by applying a varying torque of up to 35 N-m (300 inch-pounds), similar to that seen in amputee gait. Overall change in output was .047 volts, representing an average of 5.8 N (1.3 pounds) or 2.5 percent of the applied load. Linearity and hysteresis were measured by comparing the output to that of a Kistler force platform. This platform has a rated linearity of +0.5 percent of full scale output, and hysteresis of less than 0.5 percent full scale output. Different loading rates were used to determine what, if any, frequency effects were present. Figures 3 and 4 compare output of the force plate and load cell for two different loading rates. The maximum error was 2.5 percent of the largest value measured. (This test was not performed to the full design load because of the difficulty of manipulating such large weights.) One might expect significant error due to hysteresis effects of the elastomer. This error was measured under typical loading rates and found to be at most 2.5 percent of full scale, at the highest loading frequency. Overall worst-case accuracy of the load cell is estimated to be 6.5 percent of full scale.
Journal of Biomechanical Engineering
The load cell performed extremely well during six months of regular clinical use at the White 9 Physical Therapy Unit of the Massachusetts General Hospital. No failures of any kind occurred, and installation into a prosthesis was quickly and easily performed. Acknowledgments This work was supported by the National Institute of Handicapped Research, grant numbers 6008003004, 6008200048, "and 6008300074. Equipment support was also provided by Bell Laboratories of New Jersey.
References 1 Cullen, C , "Design and Evaluation of Weight and Angle Sensors for Gait Training of Above-Knee Amputees," S.M. thesis, Department of Mechanical Engineering, M.I.T., May 1984. 2 Cullen, C , and Flowers, W., " A Radio Goniometer for Use in Training Above-Knee Amputees," Proceedings of the Second International Conference on Rehabilitation Engineering, Ottawa, Canada, 1984. 3 Flowers, W., Cullen, C , and Tyra, K. P., "A Preliminary Report on the Use of a Practical Biofeedback Device for Gait Training of Above-Knee Amputees," Journal of Rehabilitation Research and Development, Vol. 23, No. 4, October 1986, pp. 7-18. 4 Craik, R., and Wannstedt, G., "The Limb Load Monitor: An Augmented Sensory Feedback Device," Proceedings: Devices and Systems for the Disabled, Philadelphia, 1975. 5 Wannstedt, G. T., and Herman, R. M., "Use of Augmented Sensory Feedback to Achieve Symmetrical Standing," Physical Therapy, 58:553-559, May 1978. 6 Wolf, S. L., and Hudson, J. E., "Feedback Signal Based on Force and Time Delay," Physical Therapy, 60:1289-1290, October 1980. 7 Ark, J. W., "Feedback Device for Lower Extremity Amputees," Proceedings of the 5th Annual Conference on Rehabilitation Engineering, Houston, Texas, 1982. 8 Wolf, S. L., and Binder-Macleod, S. A., "Use of the Krusen Limb Load Monitor to Quantify Temporal and Loading Measurements of Gait," Physical Therapy, 62:976-984, July 1982.
FEBRUARY 1990, Vol. 112/105
Downloaded From: http://biomechanical.asmedigitalcollection.asme.org/ on 05/26/2018 Terms of Use: http://www.asme.org/about-asme/terms-of-use
3 Nakamura, M., and Sawada, T., "Numerical Study on the Flow of NonNewtonian Fluid through an Axisymmetric Stenosis," ASME JOURNAL OF
The construction of the load cell was undertaken as part of a project to develop improved means of measurement for medical applications. In this instance, the effort was specifically directed towards improving the gait training of above-knee (A/K) amputees. The system developed includes a hip-angle transducer and a portable instrument to monitor transducers, produce audio feedback, and record data for patient evaluation [1, 2, 3]. Several load cells have been designed to accommodate nonamputees. The sensors are usually made in the form of a foot-pad shoe insert, such as the Krusen limb load monitor [4, 5, 6, 7]. Their performance as load cells is quite poor, however, since the sensor does not truly integrate the pressure over the entire footpad. Measurement error can be as much as 50 percent [8]. A force plate, while more accurate, is not useful for gait training enhancement, since the measurement is not available continuously and considerable concentration may be required for the amputee to place his leg on the force plate. Our goal was to design a load cell specifically for use in A/K prostheses which had improved accuracy and was rugged enough for daily clinical use.
BIOMECHANICAL ENGINEERING, Vol. 110, 1988, pp. 137-143.
4 Einav, S., Bermann, H. J., Fuhro, R. L., DiGiovanni, P. R., Fine, S., and Friedman, J. D., "Measurement of Velocity Profiles of Red Blood Cells in the Microcirculation by Laser Doppler Anemometry (LDA)," Biorheology, Vol. 12, 1975, pp. 207-210. 5 Keentok, M., Milthorpe, J. F., and O'donovan, E., "On the Shearing Zone around Rotating Vanes in Plastic Fluids: Theory and Experiment," Journal of Non-Newtonian Fluid Mechanics, Vol. 17, 1985, pp. 23-35. 6 McDonald, D. A., Blood Flow in Arteries, Arnold, London, 1960. 7 Nakamura, M., and Sawada, T., "Numerical Study on the Laminar Pulsatile Flow of Slurries," Journal of Non-Newtonian Fluid Mechanics, Vol. 22, 1987, pp. 191-206. 8 Crochet, M. J., Davies, A. R., and Walters, K., Numerical Simulation of Non-Newtonian Flow, Elsevier, Amsterdam, 1984, pp. 234-240. 9 Lee, J. S., and Fung, Y. C , "Flow in Locally Constricted Tubes at Low Reynolds Numbers," ASME Journal of Applied Mechanics, Vol. 137, 1970, pp. 9-16. 10 Young, D. F., "Fluid Mechanics of Arterial Stenosis," ASME JOURNAL OF BIOMECHANICAL ENGINEERING, Vol. 101, 1979, pp. 157-175.
11 Wille, S. 0., "Pressure and Flow Patterns in the vicinity of an arterial stenosis," Finite Elements in Biomechanics, (ed. B. R. Simon), The University of Arizona, Tucson, 1980, pp. 295-314. 12 Fukushima, T., and Azuma, T., "Patterns of Pulsatile Flow in Arterial Models with Stenosis," ClinicalHemorheology, Vol. 2, 1982, pp. 31-41. 13 Weinbaum, S., and Caro, C. G., " A Macromolecule Transport Model for the Arterial Wall and Endothelium based on the Ultrastructual Specialization Observed in Electron Microscopic Studies," Journal of Fluid Mechanics, Vol. 74, 1976, pp. 611-640. 14 Oka, S., Biorheology, Syokabo, Tokyo, 1984 (in Japanese). 15 Nakamura, M., and Sawada, T., "Model Study on the Strain and Stress Distributions in the Vicinity of an Arterial Stenosis," Biorheology, Vol. 25, 1988, pp. 685-695.
A Robust, Accurate Load Cell for Use in Training Above-Knee Amputees Christopher Cullen3 and Woodie Flowers4 Introduction In many physical therapy situations, it is useful to have a quantitative, real-time measurement of limb load. Such measurements can be used to generate biofeedback to help improve a patient's gait. This paper describes a load cell for use with an above-knee prosthesis. It has features not common to most load cells. Consulting Engineer, Cullen Associates, Ipswich, MA 01938 Professor, Department of Mechanical Engineering, Massachusetts Institute of Technology, Cambridge, MA 02139 Contributed by the Bioengineering Division of THE AMERICAN SOCIETY OF
MECHANICAL ENGINEERS. Manuscript received by the Bioengineering Division April 10, 1987; revised manuscript received October 23, 1989.
Design A dominant issue in this design problem is the presence of large bending moments in the leg. These moments produce strains in the prosthetic leg tube which are more than an order of magnitude greater than those caused by the compression to be measured. It is essential to measure only axial loading of the leg for the purpose of useful biofeedback. These moments arise because the amputee must maintain his or her center of gravity well forward of the knee joint to prevent buckling. The chosen design uses a dual-structure approach. There are separate structures used to measure the axial compressive force and to isolate bending moments. Figure 1 shows external and internal views of the load cell. It is comprised of two concentric aluminum cylinders with a thin gap between them. The inner cylinder is clamped around the lower leg tube, the outer is fixed to the knee unit. Silicone elastomer is cast into the gap between the two cylinders. Attached to the inner cylinder is a cap with a central screw which bears against an aluminum beam fixed to the outer cylinder. Strain gages on the beam provide a measure of the beam displacement. Amplifying electronics are mounted inside the knee unit cavity. The concentric cylinders serve to isolate bending moments from the measuring structure. The elastomer has the useful property of being relatively compliant under shear loading but much stiffer in compression. Imposing a moment about an axis perpendicular to the leg tube subjects the silicone to a combined shear/compressive stress loading. The cylinders are made sufficiently long so that the compressive stiffness dominates. Applied torque deflects this stiff structure only slightly, resulting in a small sliding motion of the screw across the beam and minimal strain. A solid film lubricant was applied to the loading surface to provide wear resistance for this sliding motion. Applied axial load directly deflects the compliant axial structure, producing a much larger measured beam deflection. The dual-structure design allows selection of the load cell sensitivity independent of the strength of the knee unit itself. This is an important safety consideration. This configuration has the additional advantage that failure due to overloading, i.e. a sensing beam failure, is quite benign. The beam itself can support a 135 percent overload (the design load is 1100 newtons (250 pounds)). Should it fail, the elastomer is strong enough to support an 1800 percent overload. If the bond between the elastomer and the aluminum cylinder fails, the inner cylinder will come to rest on the knee unit after moving less than .5 cm (.2 inches). FEBRUARY 1990, Vol. 112 /103
Journal of Biomechanical Engineering
Copyright © 1990 by ASME Downloaded From: http://biomechanical.asmedigitalcollection.asme.org/ on 05/26/2018 Terms of Use: http://www.asme.org/about-asme/terms-of-use
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TIME, Seconds
Flg.3 Fig. 2
Load sensing element and electronics package
Output of load cell and force plate versus time, low loading rate 7.0 ....
--,
6.0
This particular load cell is built into a modified brand-name prosthetic knee hinge unit, commonly used in training prostheses. It is easily transportable from one prosthesis to another by knocking out the hinge pin and unclamping the leg tube. One device can therefore serve several patients undergoing gait training concurrently. The load cell was built to fit within the envelope of a normal human lower leg. It only slightly increases the shank diameter of an immediate-post-operative prosthesis. To minimize discomfort to the patient, the load cell was made to be as lightweight as possible. It adds 125 grams (4.5 ounces) just below the knee. The load cell is supplied with + 5 volts D.C. and provides a o to + 5 volt output, proportional to the applied load. The onboard electronics include a preamplifier and a passive 160 Hz low-pass filter. Circuit board potentiometers are used for offset and gain adjustments. Nominal output sensitivity is 4.2 mV IN (18.7 mV Ipound) to give 4.5 Nlbit (I pound/bit) for the 8-bit AID converter in the biofeedback controller. The load cell has a modular construction. It may be 104/VoI.112, FEBRUARY 1990
¥ ~ ~
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~
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Fig. 4
Output of load cell and force plate versus time, high loading rate
disassembled to replace a defective element. The sensing beam is machined from an aluminum disk. The active strain gages are mounted in the center of the beam; the temperature compensation gages are mounted on the surface of the disk, beside the beam (Fig. 2). A small circuit board containing the elecTransactions of the ASME
Downloaded From: http://biomechanical.asmedigitalcollection.asme.org/ on 05/26/2018 Terms of Use: http://www.asme.org/about-asme/terms-of-use
tronics is mounted on top of the disk. Thus the leads from the gages to the preamplifier are kept extremely short, reducing electrical noise. Testing and Performance After initial calibration, a log was kept for a one-month period to test for null drift. The output of the load cell in the unloaded condition was measured over the course of the month. Maximum drift was .025 volts, representing 5.8 N (1.3 pounds). Null drift includes the effects of both elastomer creep and drift in the electronics. A short-term loaded creep test was performed. A 220 N (fifty pound) load was applied for twenty-four hours. Change in the null point was 7 mV, or 1.8 N (.4 pounds). Again, the effect of electronics drift are included in this measurement. Bending moment isolation was evaluated by applying a varying torque of up to 35 N-m (300 inch-pounds), similar to that seen in amputee gait. Overall change in output was .047 volts, representing an average of 5.8 N (1.3 pounds) or 2.5 percent of the applied load. Linearity and hysteresis were measured by comparing the output to that of a Kistler force platform. This platform has a rated linearity of +0.5 percent of full scale output, and hysteresis of less than 0.5 percent full scale output. Different loading rates were used to determine what, if any, frequency effects were present. Figures 3 and 4 compare output of the force plate and load cell for two different loading rates. The maximum error was 2.5 percent of the largest value measured. (This test was not performed to the full design load because of the difficulty of manipulating such large weights.) One might expect significant error due to hysteresis effects of the elastomer. This error was measured under typical loading rates and found to be at most 2.5 percent of full scale, at the highest loading frequency. Overall worst-case accuracy of the load cell is estimated to be 6.5 percent of full scale.
Journal of Biomechanical Engineering
The load cell performed extremely well during six months of regular clinical use at the White 9 Physical Therapy Unit of the Massachusetts General Hospital. No failures of any kind occurred, and installation into a prosthesis was quickly and easily performed. Acknowledgments This work was supported by the National Institute of Handicapped Research, grant numbers 6008003004, 6008200048, "and 6008300074. Equipment support was also provided by Bell Laboratories of New Jersey.
References 1 Cullen, C , "Design and Evaluation of Weight and Angle Sensors for Gait Training of Above-Knee Amputees," S.M. thesis, Department of Mechanical Engineering, M.I.T., May 1984. 2 Cullen, C , and Flowers, W., " A Radio Goniometer for Use in Training Above-Knee Amputees," Proceedings of the Second International Conference on Rehabilitation Engineering, Ottawa, Canada, 1984. 3 Flowers, W., Cullen, C , and Tyra, K. P., "A Preliminary Report on the Use of a Practical Biofeedback Device for Gait Training of Above-Knee Amputees," Journal of Rehabilitation Research and Development, Vol. 23, No. 4, October 1986, pp. 7-18. 4 Craik, R., and Wannstedt, G., "The Limb Load Monitor: An Augmented Sensory Feedback Device," Proceedings: Devices and Systems for the Disabled, Philadelphia, 1975. 5 Wannstedt, G. T., and Herman, R. M., "Use of Augmented Sensory Feedback to Achieve Symmetrical Standing," Physical Therapy, 58:553-559, May 1978. 6 Wolf, S. L., and Hudson, J. E., "Feedback Signal Based on Force and Time Delay," Physical Therapy, 60:1289-1290, October 1980. 7 Ark, J. W., "Feedback Device for Lower Extremity Amputees," Proceedings of the 5th Annual Conference on Rehabilitation Engineering, Houston, Texas, 1982. 8 Wolf, S. L., and Binder-Macleod, S. A., "Use of the Krusen Limb Load Monitor to Quantify Temporal and Loading Measurements of Gait," Physical Therapy, 62:976-984, July 1982.
FEBRUARY 1990, Vol. 112/105
Downloaded From: http://biomechanical.asmedigitalcollection.asme.org/ on 05/26/2018 Terms of Use: http://www.asme.org/about-asme/terms-of-use
