Edited by: Hyung-Soon Park, Korea Advanced Institute of Science & Technology (KAIST), South Korea
Reviewed by: Ming Liu, North Carolina State University, United States; Sang Hoon Kang, Ulsan National Institute of Science and Technology, South Korea
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The CYBERLEGs Beta-Prosthesis is an active transfemoral prosthesis that can provide the full torque required for reproducing average level ground walking at both the knee and ankle in the sagittal plane. The prosthesis attempts to produce a natural level ground walking gait that approximates the joint torques and kinematics of a non-amputee while maintaining passively compliant joints, the stiffnesses of which were derived from biological quasi-stiffness measurements. The ankle of the prosthesis consists of a series elastic actuator with a parallel spring and the knee is composed of three different systems that must compliment each other to generate the correct joint behavior: a series elastic actuator, a lockable parallel spring and an energy transfer mechanism. Bench testing of this new prosthesis was completed and demonstrated that the device was able to create the expected torque-angle characteristics for a normal walker under ideal conditions. The experimental trials with four amputees walking on a treadmill to validate the behavior of the prosthesis proved that although the prosthesis could be controlled in a way that allowed all subjects to walk, the accurate timing and kinematic requirements of the output of the device limited the efficacy of using springs with quasi-static stiffnesses. Modification of the control and stiffness of the series springs could provide better performance in future work.
Current transfemoral prostheses are most often passive, modular systems that cannot generate joint work. From fully passive ankles such as the SACH foot (Staros,
Providing positive work is an important aspect of the biological joint and there are new robotic designs that are capable of delivering it. Devices such as the OttoBock emPOWER (previously iWalk, BionX, BiOM Au and Herr,
A major drawback of all of these active propulsion systems is the high electrical demand and increased actuator complexity of providing such work. This has led to a large variety in the designs of robotic prostheses, mainly differing from how passive compliance elements are used within the system in order to reduce total energy consumption and motor requirements. Representing the fully active design principle, the Vanderbilt prosthesis only contains one spring, a 6 Nm/deg parallel spring (Lawson et al.,
These devices all have added mechanical complexity and require additional control techniques to accurately detect gait and activate control when compared to their passive equivalents. The RIC Hybrid Knee Prostheses avoid some of this actuator complexity by removing the influence of the actuator on the normal gait cycle. The devices consists of a passive mechanical knee that is used for most normal walking conditions, but also have a motor that can be engaged and disengaged using a clutch (Lenzi et al.,
Here we have designed an active prosthesis, the CYBERLEGs Beta-Prosthesis (Figure
The CYBERLEGs Beta-Prosthesis. Left is a CAD model with all of the relevant components labeled. The front of the knee shows the components of the knee drive including the carriage and series elastic springs. The WA and ET mechanisms can be found at the back of the knee. Right is the realized prosthesis with the electronics and Energy Transfer module connected in the locked position.
The knee torque angle characteristics were divided into three gait regions, and using separate systems that were optimized for particular portions of the gait cycle. The main knee actuator is a highly compliant series elastic actuator, which has the ability to change knee energy during all periods of the gait cycle. There is a Weight Acceptance (WA) mechanism to efficiently handle stance flex in the knee directly after heel strike. Connecting the knee and ankle is a special Energy Transfer (ET) mechanism, a system built with the intention of utilizing captured work from the knee to assist the ankle motor in driving the ankle. This has been explored before in passive devices (Matthys et al.,
Here we discuss the development of the CYBERLEGs Beta-Prosthesis, the design of which requires four major systems to work together to produce the desired joint torque/angle output. This device was first tested on the bench and found to reduce motor energy consumption while generating expected torque-angle characteristics for a normal walker under ideal conditions.
In amputee trials, four individuals were able to walk over level ground with the prosthesis using a simple state machine based controller. Because of the requirements of the complex relationships of the four major systems, the output kinematics, and the behavior of the person using the device, the position based control technique was not capable of producing the desired output kinematics rendering the ET system ineffective. In general, the use of the quasi-stiffness to determine series actuator spring stiffness with a controller using position setpoints can reduce the motor electrical energy consumption as long as the output kinematics, and therefore the joint torques, are near normal. In actual use, people do not find a way to use the device in a way that provides these natural kinematics, and therefore the position targets must be changed allowing people to walk, but reducing the efficacy of the series elastic actuators due to the compensation for deviations in normal torque/angle characteristics.
This paper defines each of the different systems that are contained within the CYBERLEGs prosthesis, first describing the design rationale, desired behavior, and solutions (section 2). The experiments run on the bench and in subject trials are described in section 3. Results of the behavior of the prosthesis during bench (section 4.2) and amputee validation testing (section 4.3) are then presented. We then discuss the results (section 5) and conclude with future work planned for the prosthesis system (section 6).
The Beta-Prosthesis is a transfemoral prosthesis that contains an active drive in both the knee and the ankle that are both capable of net power output on each of the joints in the sagittal plane. The design began as a passive knee/active ankle system in the Alpha-Prosthesis (Flynn et al.,
The CYBERLEGs Prosthesis was created as a part of the CYBERLEGs FP7-ICT Project, which combines a prosthesis system to replace a lost limb in parallel with an exoskeleton to assist the sound leg (Giovacchini et al.,
The CYBERLEGs Beta-Prosthesis consists of four major systems (Figure
The ankle can be fully represented by the schematic shown in Figure
Beta-Prosthesis ankle actuator schematic. Configuration of the selected MACCEPA using rigid linkages. Note the Beta-Prosthesis includes a parallel spring system with a predetermined rest position as well as a manual screw to change the MACCEPA pretension (P).
The ankle of the device is a Mechanically Adjustable Compliance and Controllable Equilibrium Position Actuator (MACCEPA) series elastic architecture (Van Ham et al.,
A parallel spring system was added to the ankle to reduce the energy consumption by reducing the necessary holding torque required by the motor and increase the velocity of ankle actuation, as shown in Figure
An example of how the torque output of the actuator is affected by two different parallel spring configurations is shown in Figure
Two examples of adding a parallel spring to modify the Torque/Angle Characteristics of the ankle. By subtracting the torque from the parallel spring (red) from the required ankle torque (black), the required motor torque is determined. In the left example, which was chosen to minimize the peak torque using two linear springs, the peak torque of the actuator is reduced from 130 to 50 Nm. In the right example, the spring was chosen to assist as much as possible without the motor needing to work against the parallel spring during the gait cycle. The peak torque is reduced to 80 Nm.
The left side of Figure
The knee is comprised of three major systems that, when used in combination, can approximate the total knee torque of normal walking with low electrical cost. These systems are the KD, the WA, and the ET system. The roles of each of these systems are outlined in Figure
Torque/Angle Characteristics of a 80 kg individual showing the behavior of the Baseline Spring (blue) and the torque during WA (green). The gait cycle begins and ends at the heel strike, progressing to the Weight Acceptance phase where the WA system provides the majority of the torque. After the WA phase, the ET system provides the necessary extension torque to keep the knee from collapsing during pushoff by pulling on the ankle. After full flexion, the ET system is disengaged and the BL spring provides flexion torque to arrest the end of swing phase and is adjusted by the KD system. In this example there is no pretension on the carriage and the carriage is placed in the nominal position (zero torque at 60 degree knee flexion).
The knee architecture schematic. The left side of the diagram shows the main knee carriage, as well as the baseline (BL) and extension (EX) springs. The BL spring provides the breaking torque during knee extension during normal walking while the EX spring provides the torque during high power extension operations. The carriage moves the rest position of the two springs. This figure also shows the relationship of the WA with the main knee drive. The right figure shows the kinematic definitions used in determining knee actuator torque, as defined in section Knee Actuator Kinematics (see
The front of the knee houses the KD actuator, as in Figure
There are two series elastic springs held within the carriage. The Baseline Spring (BL) provides the flexion torque of the knee that is shown in Blue in the Figure when the knee carriage is held at a constant position, corresponding the neutral position at approximately 60 deg. The torque created by this spring, approximately 0.3 Nm/deg flexion, can be modified while under load during all phases of the gait cycle. It is of note that this is much softer than the estimated physiological stiffness seen in Pfeifer et al. (
The right half of Figure
Directly after heel strike is the weight acceptance stage of the gait cycle. This is characterized by a spring-like loading and unloading period of the knee joint, as seen in Figure
The WA requires a lock that is capable of providing high forces when compressed, but requires little energy to position the spring under low to zero load. The WA does not need to actively provide joint work above the passive storage and recoil of the spring. Also the WA should not unlock when there is a high force on the spring, because if the spring is loaded, it usually means the user of the prosthesis is loading the knee for stability. One way to satisfy these requirements including infinite locking positions, high locking torque, and low continuous locking power is a friction based, non-backdrivable gearing (Plooij et al.,
A non-backdrivable screw was chosen as the main drive mechanism, where the rest position of the spring is modified by a small 30 W motor (Maxon EC16) through a gear drive so that as the knee flexes it compresses the spring at a desired angle, as shown in Figure
The WA system
This mechanism is attached behind the knee by a moment arm of length
During normal walking the knee primarily acts as a dissipative system, which is one reason why active damping systems can provide relatively good behavior at the knee joint. This necessity of removing energy from the joint provides the opportunity to capture this energy to return to the system at another point of time in the gait cycle. This has been previously done in systems such as the passive WalkMECH (Unal et al.,
The ET system of the Beta-Prosthesis intends to capture two major sources of negative work of the gait cycle, work done to stop the lower leg swinging forward at the end of swing phase and negative work done at the knee during late stance that is required to keep the knee from collapsing under the individual while beginning flexion for the swing phase. To do this, there is a cable that directly connects the ankle to the knee when it is locked, and allows the cable to go slack when it is unlocked. A schematic of this system can be found in Figure
The ET system schematic showing the direct connection of the ankle to the knee joint. Here the knee joint
The Beta-Prosthesis was designed to integrate with the CYBERLEGs control and orthosis module systems and therefore does not have onboard control systems or power. The CYBERLEGs control and power system is a completely autonomous system consisting of a wearable backpack housing the control unit and batteries for power (Giovacchini et al.,
The sbRIO computer is connected to the prosthesis through a tether cable that carries all of the control signals from the purpose made driver board mounted on the prosthesis. This driver board uses four Maxon 50/5 module driver boards, one for each of the motors in the prosthesis. The tether also transmits the 36 V supply voltage to be compatible with the other components of the CYBERLEGs system.
The knee and ankle systems were tested on a custom designed test bench to verify that the system could produce the desired torques and the passive behavior of the knee creates a good approximation of the desired knee torque, as seen in Figure
The prosthesis was clamped in a rigid frame with a motor connected to the output of the joint to be tested. This figure shows the prosthesis as it would be positioned in knee actuator or ET system tests. The frame has a large DC motor attached to the prosthesis through a torque sensor, so that the output kinematics of the joint can be varied at the same time as the prosthesis is actuated.
Low level control of the system of the prosthesis is handled by the ESCON drivers, running a tuned current and velocity feedback loop. The velocity signals used by the ESCON are generated by the FPGA of the sbRIO running a control loop at 1,000 Hz. Position commands are calculated by the real time system at 100 Hz and fed to the FPGA. Position commands that are tracked by the real time system are calculated by the Top Level Control.
Control methods for the Beta-Prosthesis utilized a modified Intention Detection (ID) system and Wearable Sensory Apparatus (WSA) controller with a finite state machine as discussed in Ambrozic et al. (
Normal level ground gait cycle (adapted from Cuccurullo,
The CYBERLEGs Walking State Machine. The CYBERLEGs WSA was used to create the triggers for state transitions, using the angular velocity, ω(
It should also be stated that these position setpoints (locked, unlocked, half flex, etc.) were tuned for the individual's self selected speed and were used only for level ground walking.
The male subjects (
The ankle was placed in the test bench described in section 2.2 with the ankle output connected to the torque transducer. The output motor was used to drive the ankle angle along the Winter trajectory at 1.5 s/stride while the ankle moment arm was commanded to provide the ankle torque for the given time in the gait cycle. Although the ankle is capable of providing 130 Nm of torque during the gait cycle, the test bench motor was only capable of driving the output to 90 Nm due to limitations in the current the external motor driver was able to provide. Therefore a Winter torque/angle trajectory was selected with a peak torque of around 90 Nm. Results can be found in Figure
Ankle actuator torque using moment arm control
In this experiment the knee was locked at a fixed angle and connected to the test bench torque transducer. The carriage was commanded to deliver the maximum joint torque. This was repeated three times for both flexion and extension every 10° of knee flexion, from full extension to full flexion. The actual joint torque was recorded and maximum measured torque was compared to estimated maximum torque. Results can be found in Figure
The prosthesis was placed in the test bench setup with the output motor of the test bench connected to the knee joint. The ET system was then commanded to lock and unlock from approximately 43% to 64% of the gait cycle, corresponding to the time of negative work of the knee joint. The ET cable tension was measured with a load cell while the knee and the ankle were commanded to track respective position trajectories. The ankle motor was commanded in output position mode, pinning the kinematics of the ankle to the Winter trajectory regardless of the torque required.
As part of the preliminary validation of the prosthesis system, four amputees were subject to a 3 min treadmill experiment. Each subject attached the prosthesis to their existing socket through a standard prosthesis pyramid adapter and a visual alignment was performed by a prosthesis technician. After a short session of approximately 15 min for tuning of the prosthesis state machine parameters from section 2.3.2 over a level ground 10 meter long catwalk, the subject was asked to walk at a self selected pace on the treadmill for 3 min to investigate the behavior of the knee joint. This self selected speed was determined by slowly increasing the speed of the treadmill in 0.2 km/h increments until the subject felt it was too fast to sustain. The speed was then reduced until the subject was comfortable. Subjects were asked to minimize the use of the handrails, although to not remove their hands from the rails in case of an emergency. Subjects for the experiments were selected and approved by the Ethical Committee of the Don Gnocchi Foundation (FDG), Florence Italy, and experiments were conducted under FDG supervision. Data was recorded from the prosthesis during the treadmill testing period.
During the validation with subjects, the ET system was not used for testing. This was because the control of the ET system would not allow for kinematics that varied greatly from the desired Winter kinematics, leading to high forces as well as ineffective gait. This was mitigated though the use of a modified KD trajectory, allowing the subjects to walk. This is discussed in greater detail in section 5.2.1.
A list of prosthesis design values and prosthesis characteristics are found in Table
Selected prosthesis characteristics used in simulation and for final design.
Active DOF | 2 | knee, ankle |
System Voltage | 36 | VDC |
Ankle Moment Arm Length ( |
50 | mm |
Ankle Linkage Length ( |
50 | mm |
Ankle Actuator Spring Constant ( |
130 | N/mm |
Ankle Actuator Spring Weight | 113.4 | g |
Ankle Parallel Spring Constant | 94.20 | N/mm |
Ankle Parallel Spring Weight | 60 | g |
Ankle Gear Ratio | 320:1 | - |
Ankle Winding Voltage | 48 | V |
Ankle Max Torque | 130 (@15A) | Nm |
Ankle Continuous Torque (no parallel spring) | 30.6 | Nm |
Ankle Range of Motion | –30 to 20 | deg |
Shoe Size | 42 | EU |
Knee Gear Ratio | 5.8:1 | - |
Knee Ball Screw Ratio | 2 | mm/turn |
Knee Max Torque | ~70 | Nm |
Knee Continuous Torque | ~55 | Nm |
Knee Range of Motion | 0 to 95 | deg |
Baseline Spring Constant ( |
10.7 | N/mm (each spring) |
Baseline Spring Length | 64 | mm |
Baseline Spring Diameter | 16 | mm |
Baseline Spring Mass | 19.7 | g |
Extension Spring Constant ( |
89.1 | N/mm (each spring) |
Extension Spring Length | 32 | mm |
Extension Spring Diameter | 16 | mm |
Extension Spring Mass | 15.6 | g |
Weight Acceptance Spring Constant | 300 | N/mm |
Weight Acceptance Spring Rest Length | 38 | mm |
Weight Acceptance Spring Diameter | 32 | mm |
Weight Acceptance Spring Weight | 90 | g |
Prosthesis Overall Mass | ~5 | kg |
- Knee Module Mass | 1,736 | g |
- WA Module Mass | 440 | g |
- Ankle Module Mass | 1,850 | g |
- Electronics | 300 | g |
Prosthesis Height (Overall, longest) | 500 | mm |
Knee to Ankle Length (Shortest) | 355 | mm |
Knee to Ankle Length (Longest) | 395 | mm |
Knee to Pyramid Top Length | 30 | mm |
The results of the ankle torque/angle characteristics from the bench testing device can be found in Figure
Figure
Figure
Bench test results of the prosthesis with ET system.
A representative example of a subject walking on the treadmill for 126 strides at a speed of 2.2 km/h and average stride duration of 1.76 s/stride (
The experimental mean torque-angle characteristics of the ankle
Figure
Knee angle, Knee Drive carriage position, and WA position during treadmill walking. The desired position of the WA and KD carriage are shown in red, here the gait state machine determines when the position setpoint should be changed, the position of which was determined experimentally through tests over level ground. The state changes do not happen at exactly the same time during the gait cycle on every step because of variances in the signals given to the WSA, resulting in curved averaged state values. Δ
Figure
Figure
Figure
Figure
From previous power consumption measurements, the total device consumes about 65 J/stride. Of this, the WA requires about 10 J/stride regardless of the speed of the gait. Electrical model estimations show that during this trial the ankle uses about 19 J/stride while the knee uses about 23 J. These values seem to be considerably less than older versions of the Vanderbilt prosthesis (91 J/stride), although this was at 5.1 km/h and 87 steps/min with a higher joint work output (Sup et al.,
Testing of the device was divided into two different sections, bench testing and the patient trials. Bench testing of the device shows the device can work as designed when the external kinematics are imposed, with good ankle and knee torque approximation of the biomechanical data at speeds close to the actual speeds of the tests. Patient trials showed that all four patients were able to walk with the prosthesis using the CYBERLEGs WSA over level ground and on the treadmill with minimal (< 1 h) training. Overall these tests show it is possible to find walking gaits that allow for ambulation with actuators that utilize soft series elastic springs, although the overall behavior was not the same as averaged normal gait as defined by the Winter biological data. An example of the typical gait can be seen in the
From the patient experiments, it was apparent that ET design and control was insufficient for gaits greatly differing from the average kinematics, leading to the loss of ET function and in some cases interfering with the gait cycle. Due to this loss of ET function, the knee actuator control was highly modified compared to the original design. This resulted in gait that was closer in behavior to a passive knee than the average biomechanical data. This was exacerbated by short training periods that were insufficient to allow the needed level of familiarity with a device that operates dramatically different from their currently prescribed prostheses.
Even though the tests were slightly limited in speed and magnitude compared to the original torque and angle targets, we have been able to show that the power consumption of the knee actuator alone (i.e., without the ET system) does reduce the energy consumption of the motors compared to a direct drive system (Geeroms et al.,
The ankle has a bit of hysteresis primarily coming from friction in the parallel spring mechanism, which can be seen in Figure
The right side of Figure
Using the force in the cable and the moment arm of the ET on the ankle, the torque on the ankle can be computed. Multiplying this with the displacement of the ankle gives the energy transferred from the knee to the ankle. In this example the value is 8.2 J, which is about 80% of the available energy from the knee and about half of the required energy for pushoff, a considerable savings if it can be replicated in walking individuals.
The top graph in Figure
The knee trajectory was slightly modified to allow the best fit between knee and ankle torques. This also had the effect of shifting the knee power a bit earlier which allowed the energy transfer to align better with the ankle pushoff. In normal walking, the negative work from the knee is slightly after the pushoff of the ankle, as shown in Figure
Four subjects were able to walk on level ground after a short training session. Two of the four were able to increase their self selected walking velocity by approximately 0.2 km/h over their own prosthesis. Figure
The ankle was able to inject energy into the gait cycle, but this amount was small during these tests. The ankle used a very small pushoff angle, at the request of the subjects. On average the ankle performed about 2.8 J of work per step, which is far from the 17 J of the target walking, but better than the slightly negative total work provided by passive devices.
The question then arises if users settle into these ideal torque and angle characteristics if the prosthesis is able to do so on the test bench. From the initial patient trials with the prosthesis, it was clear that the subjects did not have a Winter-like average gait torque/angle progression, with joint torque and angle characteristics varying greatly from the original targets.
The largest deviation from the original design was the inability to utilize the ET system. The ET system was designed with specific required output kinematics to correctly harvest knee negative work and deliver that work to the ankle. Deviations from these kinematics caused a dramatic rise in the cable system past the ET design criteria. For example under normal conditions the ET could expect to see approximately 600 N of tension during walking (Heins et al.,
Removing the ET modified the behavior of the KD because the original torque angle behavior of the prosthesis depended on this system to provide extension torque during the late stance. This modification of the KD required the actuator to move much further distance than originally intended to provide this extension torque, which limited the peak velocity of the knee.
It is possible that if proper feedback control for the ET mechanism was created, then it might be possible to safely use the system by limiting the tension in the cable, although because the ET was not designed to actuate under load, this was not implemented. In addition if the ET is then a fully active system rather than a clutched system as designed, there must be a controller that can watch the knee and ankle joints and predict periods of time when the knee would dissipate power and the ankle provide kinematics that were suitable to receive this energy, if those periods exist. Another way of solving this issue is to strictly control the output kinematics, which would guarantee the knee and ankle relationships. This method would not necessarily result in a reduction of motor electrical consumption or a suitably stable gait. Modifications to the ET system must be made to further examine this aspect of the design in walking trials.
The control system of the prosthesis was designed to change the configuration of the prosthesis in order for the output torque to match some quasi-static torque target assuming that the output kinematics of the system would then converge to the normal gait kinematics. This method makes a of assumptions, most importantly first that if the average joint torques are obtained, the person would naturally have kinematics which are about normal and second that deviations away from normal joint torques caused by external disturbances would be sufficiently handled by the natural impedance of the prosthesis and the control of the person using the prosthesis. The control method was not designed to be a classical impedance or torque based system using output feedback to generate a specific stiffness or trajectory. On the bench this works well because the output kinematics of the system are constrained by the output motors, and therefore the joint torques and kinematics are as expected.
When both the torque and kinematics are unconstrained, the person tends to walk very differently than expected, particularly at the knee. While ankle kinematics and torques were somewhat normal for the low input power that was desired, the knee behavior was far different. To find gait cycles that were capable of safely walking, the ET needed to be disabled. Because of the lack of the ET system during the walking trials, the control of the system was modified so that the KA could replace the functionality of the WA system. This included changing the state machine to follow a different trajectory than that with the ET enabled, ultimately implementing a very simple flexion/extension motion for the knee, as was determined from feedback from the subjects. For each of the 4 states of the gait cycle, the subjects were asked what behavior they would like to have from the prosthesis and the position threshold for the state was determined.
The Beta-Prosthesis was designed with the intention of providing the normal torque and kinematics of a leg, as determined by average healthy gait. With this particular design and controller there is an implicit assumption that the prosthesis has a similar mass and moment of inertia as an average human leg, because the target torques are dependent on these aspects. This prosthesis has been built with this as a constraint, but tends to lead to a relatively heavy prosthesis and associated problems, such as socket pistoning. From discussions with the subjects during the trials, while walking while powered this extra weight is not noticed until the prosthesis performs poorly or is not actuated, and then the weight is highly detrimental.
It should be noted that even without the normal biological torque/angle joint progression the patients were able to walk at speeds equal to or above those while using their every day prosthesis. So how necessary is it that the prosthesis really track the normal gait characteristics? Indeed some extremely fast transfemoral amputee sprinters find that the design of their passive prostheses may not need a knee joint at all, relying on the prosthesis design to generate the pushoff and using the hips to provide ground clearance.
Volumetric oxygen measurements with almost all current prostheses are generally 10–30 percent higher than normal walking, and 50–100 percent higher at maximal speeds of walking (Genin et al.,
Although symmetric gaits may not be optimal for energy consumption, there is an increased chance of osteoarthritis in the sound leg in transtibial and transfemoral amputees, and there are reasons to believe that increasing joint kinematic symmetry generally leads to reduced detrimental loading, particularly peak force and peak knee external adduction in the contralateral limb (Morgenroth et al.,
Even though natural kinematics and torques do not necessarily minimize metabolic energy consumption or minimize injury, one thing that is certain is that the more normal and symmetric the gait kinematics the more natural and unassuming it looks, which is a large part of the functionality of a daily worn prosthesis. It also provides a familiar starting target for the design of prosthetic limbs which are designed to replace normal limbs. Possibly designs based on non-anthropomorphic principles will allow the discovery of other solutions in the future (LaPrè and Sup,
Regardless, the current control of the prosthesis does not attempt to force the Winter kinematics output at the knee and results in an asymmetric gait. This was shown to increase the metabolic rate of the participants (10±9%) when compared to their conventional prostheses. As the subjects become more familiar with the device, the control becomes more refined, and we are able to better apply torques with more accurate timing, we expect an increase in gait symmetry and to eventually reduce metabolic consumption (Malcolm et al.,
Figure
Comparison of the prosthesis behavior vs. the target Winter and C-Leg data (Segal et al.,
The current prosthesis control uses motor position setpoints which change the position of the motor side of the SEA based on a heuristic rule-based state machine. This method requires that the dynamic and contact forces of the user are somewhat near to the normal values from which the targets were derived because both the generated kinematics and torques of the joints are completely dependent on these external forces.
There is no feedback of the output trajectory, output impedance, or output torque to compensate for deviation from the target torque/angle in this method. This is actually similar to a rest position microcontroller controlled system, where the rest position of a spring is changed during different phases in the gait cycle, although here the position can be changed while loading and unloading. It was theorized that if the position of the motor side of the spring was placed close enough to the correct position, the loading characteristics of the output could be slightly modified by the walker and they would find the best way to walk with the device, resulting in near normal kinematics and joint torque. In this way, neither the kinematics or the torque are fully constrained. Results show that this tends to work well in the ankle, the users seem to be able to load and unload the ankle in a biologically similar fashion, albeit with reduced energy injection, but with the modified knee control, the knee did not prove to be as well-behaved.
Because the ET was designed to utilize a very specific knee/ankle torque and kinematics relationship, the lack of constraint in the control of the kinematics allowed the device to attain unsafe conditions, and could not be used as designed. It is possible that in a system that is designed to retain angle relationships between the knee and ankle the ET system would work as designed, although because of the actuator effort to keep kinematic accuracy, it isn't guaranteed that this would be useful toward prosthesis energy reduction. Another option would be to determine a gait suitable for an individual without the use of the ET and then adding back ET capability if the gait allows for negative energy of the knee to be transmitted.
When it was decided that the ET portion of the knee could not be used, a new trajectory for the position of the knee carriage was created based on feedback from the people doing the trials, without full regard to the actual torque/angle characteristics of the knee. The subjects also did not use the stance flex WA system, and preferred to utilize the end stop of the knee as much as possible during stance. This behavior may, in part, be to the way conventional sockets are set and how people are trained to use prostheses. Knee hyperextension is often used for knee “stability” during the stance phase using conventional prostheses and it is possible with a modified socket alignment this tendency could be reduced. Control and setup were the main reasons that the behavior of the prosthesis resembled that of a passive prosthesis.
It is clear that a refined, better tuned, control system with clearer goals in system constraints will be required to produce more normal knee torque/angle characteristics. The top level state machine system was not the most adaptable system that could have been chosen for this task, although it was sufficient to obtain preliminary walking gait. Improvements to this will need to include much more training of the user to utilize the WA correctly as well as adding in a better position trajectory of the knee carriage to provide expected knee torque. For topics such as gait symmetry and metabolic consumption, better performance is needed than this control provided. For a fairly complete discussion on different control methods in prostheses and exoskeletons, readers should refer to Tucker et al. (
We have created a new, active, combined ankle-knee prosthetic system which achieves as much as possible with a passive approach, using springs that were chosen to match the biological quasi-stiffness of normal gait. These springs are locked and unlocked during the gait cycle and combined with an energy harvesting system to passively provide the majority of the required torque angle characteristics during normal walking, while maintaining versatility by providing active actuation. Under ideal conditions the prosthesis worked on the bench as designed, which also showed a lower motor electrical cost than most current designs. In particular, the capability of the energy transfer system to reduce both the knee and ankle motor work was considerable. In the ideal case the prosthesis performs quite well compared to the current array of powered prostheses.
In reality there are two major issues with the prosthesis. First these savings are all but eliminated by the implementation, where it takes approximately 10 J/stride to capture a similar amount of work in both the WA and ET systems. More efficient, and in the case of the ET more controllable, locking mechanisms should be found to better utilize these systems. Second is that the people wearing the prosthesis do not seem to be able to find walking patterns that utilize the average torque/angle characteristics. Because the prosthesis does not impose either torque or trajectory upon the user, they tend to find gait patterns that are very different from the average biomechanical data. This may be due to training and unfamiliarity with the prosthesis, it may have to do with the nature of the socket interface, inaccuracy of the control timing, or a combination of other reasons. When the user deviates from the average biomechanical trajectories, the energy saving functions of the prosthesis are reduced and the device functions similarly to other powered prostheses under evaluation today.
We conclude with a summary of points learned while developing this prosthesis:
– Bench testing showed the quasi-static stiffness based prosthesis can reproduce average walking knee and ankle joint torques when the output of the prosthesis was constrained with external motors. Under these conditions the ET was found to be capable of transferring energy from the knee to the ankle and a considerable energy consumption reduction of the motors was found. – The prosthesis was used in a preliminary validation experiment with four amputee subjects and through modification of the main actuator behavior, the prosthesis was able to create a stable gait cycle with all subjects. – Using the quasi-stiffness estimations from average biomechanical data for the stiffness of the ankle springs creates behaviors that resemble the average biological data in walking trials. – Using the quasi-stiffness estimations for the stiffness of the knee springs did not provide sufficiently average kinematics and torques during walking trials. Even though it is possible to generate average torque and kinematics in the prosthesis it does not mean the person using it will choose to walk with average torque and kinematics without stabilizing constraints. – Energy transfer from the knee to the ankle is possible under ideal conditions. – Because of deviations in the knee and ankle joint kinematics during walking tests, the tests had to be run without the use of the ET system. These mismatches stem from a combination of the prosthesis control, which does not constrain the kinematics, the ET control, which was treated as a locked or unlocked clutch in this implementation, as well as the way the subjects interact with the prosthesis, preferring behaviors that were not like average biomechanical data. – In order to overcome differences in kinematics, the motor must actuate in a different manner than average biomechanical data would suggest which reduces the efficacy of the quasi-stiffness approach in reducing energy consumption, particularly in the knee. – The use of low stiffness springs in the knee determined by quasi-stiffness trajectories limit the ability of the actuator to modify the behavior of the knee due to low actuator bandwidth, although solutions can be found that provide stable gait. – A simple state machine system with a number of experimentally tuned variables to set thresholds for actuation and timing was implemented and work sufficiently to provide basic gait functions. These thresholds were primarily determined by feedback from the patients, and resulted in a powered ankle actuation that was similar to biological ankle function at a reduced amplitude and knee behavior similar to current microcontroller devices. – It was determined that a much longer training period must be allowed for the users before measurements. Because the prosthesis behaves quite differently to a standard prosthesis, the user must learn to have high trust the device and they must have a detailed understanding of the behavior of the device and how it can be utilized. Training alone may improve kinematics, although it is not the only issue. – New gait detection and control methods should be able to better utilize the passive aspects of the prosthesis, but how this can best be accomplished is a focus of future work.
This study was carried out in accordance with the recommendations of Fondazione Don Carlo Gnocchi. Ethical approval for the protocol was obtained through FDG. All subjects gave written informed consent in accordance with the Declaration of Helsinki.
LF designed and constructed the hardware, conducted experiments, analyzed data, and wrote the manuscript. JG designed and constructed the hardware, conducted experiments, analyzed data, and proofread the manuscript. RJ-F supervised experiments at VUB. SH helped with data analysis and derived governing torque equations. BV supervised experiments at VUB and proofread the manuscript. MM provided the WSA system and assisted with top level control. RM presided over the experimental sessions at FDG, gained ethical approvals, and lead patient recruitment. NV supervised the experimental sessions. DL supervised the hardware design and proofread the manuscript.
The authors declare that the research was conducted in the absence of any commercial or financial relationships that could be construed as a potential conflict of interest.
We would like to thank Marnix De Boom and Marc Luypaert for their work building the Beta-Prosthesis. Thank you to Carlos Rodriguez-Guerrero for assistance in proofreading the manuscript. We would also like to thank our test subjects for their participation in our project.
The Supplementary Material for this article can be found online at: