U.S. patent application number 13/079564 was filed with the patent office on 2011-12-01 for controlling power in a prosthesis or orthosis based on predicted walking speed or surrogate for same.
Invention is credited to Chris Barnhart, Richard J. Casler, Gary Girzon, Zhixiu Han, Hugh M. Herr.
Application Number | 20110295384 13/079564 |
Document ID | / |
Family ID | 44065277 |
Filed Date | 2011-12-01 |
United States Patent
Application |
20110295384 |
Kind Code |
A1 |
Herr; Hugh M. ; et
al. |
December 1, 2011 |
CONTROLLING POWER IN A PROSTHESIS OR ORTHOSIS BASED ON PREDICTED
WALKING SPEED OR SURROGATE FOR SAME
Abstract
In some embodiments of a prosthetic or orthotic ankle/foot, a
prediction is made of what the walking speed will be during an
upcoming step. When the predicted walking speed is slow, the
characteristics of the apparatus are then modified so that less
net-work that is performed during that step (as compared to when
the predicted walking speed is fast). This may be implemented using
one sensor from which the walking speed can be predicted, and a
second sensor from which ankle torque can be determined. A
controller receives inputs from those sensors, and controls a
motor's torque so that the torque for slow walking speeds is lower
than the torque for fast walking speeds. This reduces the work
performed by the actuator over a gait cycle and the peak actuator
power delivered during the gait cycle. In some embodiments, a
series elastic element is connected in series with a motor that can
drive the ankle, and at least one sensor is provided with an output
from which a deflection of the series elastic element can be
determined. A controller determines a desired torque based on the
output, and controls the motor's torque based on the determined
desired torque.
Inventors: |
Herr; Hugh M.; (Somerville,
MA) ; Casler; Richard J.; (Franklin, MA) ;
Han; Zhixiu; (Acton, MA) ; Barnhart; Chris;
(Watertown, MA) ; Girzon; Gary; (Sudbury,
MA) |
Family ID: |
44065277 |
Appl. No.: |
13/079564 |
Filed: |
April 4, 2011 |
Related U.S. Patent Documents
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Application
Number |
Filing Date |
Patent Number |
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61320991 |
Apr 5, 2010 |
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61422873 |
Dec 14, 2010 |
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61432083 |
Jan 12, 2011 |
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Current U.S.
Class: |
623/24 ;
623/50 |
Current CPC
Class: |
A61F 2/70 20130101; A61F
2002/5003 20130101; A61F 2002/6685 20130101; A61F 2002/5018
20130101; A61F 2002/745 20130101; A61F 2002/607 20130101; A61F
2005/0155 20130101; A61F 2002/7635 20130101; A61F 2002/7645
20130101; A61F 2002/7695 20130101; A61F 2/6607 20130101; A61F
5/0127 20130101; A61F 2002/701 20130101; A61F 2002/704 20130101;
A61F 2002/5007 20130101; A61F 2002/503 20130101; A61F 2/68
20130101; A61F 2002/7625 20130101; A61F 2002/5033 20130101 |
Class at
Publication: |
623/24 ;
623/50 |
International
Class: |
A61F 2/48 20060101
A61F002/48; A61F 2/66 20060101 A61F002/66 |
Claims
1. An ankle-foot prosthesis or orthosis apparatus comprising: a
shank member; a foot member that is operatively configured with
respect to the shank member so as to supporting walking and permit
the foot member to plantarflex and dorsiflex with respect to the
shank member; a motor configured to plantarflex the foot member
with respect to the shank member; a series elastic element
connected between at least one of (a) the motor and the shank
member and (b) the motor and the foot member; at least one first
sensor having an output from which a walking speed of an upcoming
step can be predicted; at least one second sensor having an output
from which ankle torque can be determined; and a controller
configured to control the motor's torque, based on the output of
the at least one first sensor and the at least one second sensor,
so that the motor's torque for slow walking speeds is lower than
the motor's torque for fast walking speeds.
2. The apparatus of claim 1, wherein the motor is also configured
to dorsiflex the foot member with respect to the shank member.
3. The apparatus of claim 1, wherein the at least one first sensor
comprises at least one of an angular rate sensor and an IMU.
4. The apparatus of claim 1, wherein the controller controls the
motor's torque based on the output of the at least one first sensor
immediately before a reflex occurs.
5. The apparatus of claim 1, wherein the controller is configured
to (i) determine, based on the output of the at least one first
sensor, a control gain that varies with walking speed, wherein the
control gain at slow walking speeds is lower than the control gain
at fast walking speeds, (ii) determine a desired motor torque based
on the control gain and a determined ankle torque, and (iii) drive
the motor to achieve the desired motor torque.
6. The apparatus of claim 1, wherein the controller is configured
to (i) determine, based on the output of the at least one first
sensor, an angular rate .omega..sub.x of the shank, (ii) determine
a control gain Kv(.omega..sub.x) that is a function of the angular
rate, wherein the control gain at low angular velocities is lower
than the control gain at low angular velocities, (iii) determine a
desired motor torque based on the equation Motor
torque=Kv(.omega..sub.x).times.pff.times.(normalized_Torque).sup.n,
where pff is a constant and n is between 2 and 7, and (iv) drive
the motor to achieve the desired motor torque.
7. The apparatus of claim 6, wherein Kv(.omega..sub.x)=0 when
.omega..sub.x=0, Kv(.omega..sub.x)=1 when .omega..sub.x exceeds a
threshold .omega..sub.TH, and Kv(.omega..sub.x) is a monotonically
increasing function between .omega..sub.x=0 and .omega..sub.TH.
8. The apparatus of claim 7, wherein the motor is also configured
to dorsiflex the foot member with respect to the shank member.
9. The apparatus of claim 1, wherein the at least one second sensor
measures ankle torque directly.
10. The apparatus of claim 1, wherein the at least one second
sensor has at least one output from which a deflection of series
elastic element can be determined, and the controller computes the
torque based on the at least one output.
11. The apparatus of claim 1, wherein the at least one second
sensor comprises a sensor that senses a position of the motor and a
sensor that senses an angle of the foot member with respect to the
shank member, and the controller computes the torque based on the
sensed position of the motor and the sensed angle.
12. The apparatus of claim 1, wherein the at least one second
sensor comprises a sensor that senses a position of the motor and a
sensor that senses an angle of the foot member with respect to the
shank member, and the controller determines a torque component
.GAMMA..sub.S based on the sensed position of the motor, the sensed
angle, and a torque vs. deflection characteristics of the series
elastic element.
13. The apparatus of claim 12, further comprising a bumper that is
compressed when the foot member is sufficiently dorsiflexed with
respect to the shank member, wherein the controller determines a
torque component .GAMMA..sub.B based on the sensed angle and a
torque vs. deflection characteristics of the bumper, and wherein
the controller determines a total torque based on .GAMMA..sub.S and
.GAMMA..sub.B.
14. A method of modifying characteristics of an ankle-foot
prosthesis or orthosis apparatus, the method comprising the steps
of: predicting what a walking speed will be during an upcoming
step; and modifying a characteristic of the apparatus during the
upcoming step in situations when the predicted walking speed is
slow, wherein the modification of the characteristic results in a
reduction in net non-conservative work that is performed during the
upcoming step as compared to the net non-conservative work that is
performed when the predicted walking speed is fast.
15. The method of claim 14, wherein the step of modifying a
characteristic comprises reducing a power control gain in
situations when the predicted walking speed is slow.
16. The method of claim 14, wherein the predicting step comprises
predicting what a walking speed will be during an upcoming step
based on a shank angular rate measurement during a controlled
dorsiflexion phase that directly precedes the upcoming step.
17. The method of claim 16, wherein the shank angular rate
measurement is made at foot-flat.
18. An apparatus comprising: a proximal member; a distal member
that is operatively connected with respect to the proximal member
by a joint so that an angle between the distal member and the
proximal member can vary; a motor configured to vary the angle
between the distal member and the proximal member; a series elastic
element connected between at least one of (a) the motor and the
proximal member and (b) the motor and the distal member; at least
one first sensor having an output from which a walking speed of an
upcoming step can be predicted; at least one second sensor having
an output from which a joint torque can be determined; and a
controller configured to control the motor's torque, based on the
output of the at least one first sensor and the at least one second
sensor, so that the motor's torque for slow walking speeds is lower
than the motor's torque for fast walking speeds.
19. The apparatus of claim 18, wherein the at least one first
sensor comprises at least one of an angular rate sensor and an
IMU.
20. The apparatus of claim 18, wherein the controller controls the
motor's torque based on the output of the at least one first sensor
immediately before a reflex occurs.
21. The apparatus of claim 18, wherein the controller is configured
to (i) determine, based on the output of the at least one first
sensor, a control gain that varies with walking speed, wherein the
control gain at slow walking speeds is lower than the control gain
at fast walking speeds, (ii) determine a desired motor torque based
on the control gain and a determined joint torque, and (iii) drive
the motor to achieve the desired motor torque.
Description
CROSS REFERENCE TO RELATED APPLICATIONS
[0001] This application claims the benefit of U.S. Provisional
Applications 61/320,991 filed Apr. 5, 2010, 61/422,873 filed Dec.
14, 2010, and 61/432,083 filed Jan. 12, 2011, each of which is
incorporated herein by reference.
BACKGROUND
[0002] US published patent applications 2010/0174384 ("the '384
application") and 2006/0249315, each of which is incorporated
herein by reference, describe that the gait cycle for walking can
be divided into five phases: controlled plantarflexion, controlled
dorsiflexion (CD), powered plantarflexion (PP), early swing, and
late swing, as depicted in FIG. 1.
[0003] The '384 application also discloses a number of embodiments
of lower-extremity prosthetic and orthotic systems in which the
reflex torque generation during PP is achieved via non-linear,
positive feedback between the series elastic element (SEE) motor
torque and ankle torque. More specifically, the reflex action
involves behaving like a non-linear spring during CD and like a
torque source during PP. This reflex action can be implemented by
driving the motor using the following equation:
Motor Torque=pff.times.(normalized_Torque).sup.n Eq. 1
Where, pff is the power control gain tuned for high walking speed;
normalized_Torque is the ankle torque, .GAMMA..sub.A, normalized by
a torque, .GAMMA..sub.0, (strongly related to users' weight); n is
the power exponent, typically in the range of between 3 and 5 for
level-ground walking. Note that pff has units of N-m, and the value
of pff controls the magnitude of the level of the torque reflex
during fast walking. Once the desired motor torque is determined,
the drive current can be computed based on the equation Motor
Current=Motor Torque/kt, where kt is the motor torque constant.
While using Equation 1 does provide good results, the results
provided by the control approach described below are significantly
better.
SUMMARY OF THE INVENTION
[0004] One aspect of the invention is directed to an ankle-foot
prosthesis or orthosis apparatus. The apparatus includes a shank
member and a foot member that is operatively configured with
respect to the shank member so as to supporting walking and permit
the foot member to plantarflex and dorsiflex with respect to the
shank member. A motor is configured to plantarflex the foot member
with respect to the shank member, and a series elastic element is
connected between at least one of (a) the motor and the shank
member and (b) the motor and the foot member. There is at least one
first sensor having an output from which a walking speed of an
upcoming step can be predicted, and at least one second sensor
having an output from which ankle torque can be determined. The
apparatus also includes a controller configured to control the
motor's torque, based on the output of the at least one first
sensor and the at least one second sensor, so that the motor's
torque for slow walking speeds is lower than the motor's torque for
fast walking speeds.
[0005] Another aspect of the invention is directed to a method of
modifying characteristics of an ankle-foot prosthesis or orthosis
apparatus. The method includes the steps of predicting what a
walking speed will be during an upcoming step and modifying a
characteristic of the apparatus during the upcoming step in
situations when the predicted walking speed is slow. The
modification of the characteristic results in a reduction in net
non-conservative work that is performed during the upcoming step as
compared to the net non-conservative work that is performed when
the predicted walking speed is fast.
[0006] Another aspect of the invention is directed to an apparatus
that includes a proximal member and a distal member that is
operatively connected with respect to the proximal member by a
joint so that an angle between the distal member and the proximal
member can vary. A motor is configured to vary the angle between
the distal member and the proximal member, and a series elastic
element is connected between at least one of (a) the motor and the
proximal member and (b) the motor and the distal member. There is a
least one first sensor having an output from which a walking speed
of an upcoming step can be predicted, and at least one second
sensor having an output from which a joint torque can be
determined. The apparatus also includes a controller configured to
control the motor's torque, based on the output of the at least one
first sensor and the at least one second sensor, so that the
motor's torque for slow walking speeds is lower than the motor's
torque for fast walking speeds.
[0007] Another aspect of the invention is directed to an ankle-foot
prosthesis or orthosis apparatus that includes a shank member and a
foot member that is operatively configured with respect to the
shank member so as to supporting walking and permit the foot member
to plantarflex and dorsiflex with respect to the shank member. A
motor is configured to plantarflex the foot member with respect to
the shank member, and a series elastic element is connected between
at least one of (a) the motor and the shank member and (b) the
motor and the foot member. The apparatus also includes at least one
sensor having an output from which a deflection of the series
elastic element can be determined, and a controller configured to
determine a desired torque based on the output, and to control the
motor's torque based on the determined desired torque.
[0008] Another aspect of the invention is directed to a method of
controlling an ankle-foot prosthesis or orthosis having a foot
member and shank member, with a motor configured to plantarflex the
foot member with respect to the shank member and a series elastic
element in series with the motor. The method includes the steps of
sensing a position of the motor, determining a deflection of the
series elastic element while the motor is at the position sensed in
the sensing step, and controlling the motor's torque based on the
motor position sensed in the sensing step and the deflection
determined in the determining step.
[0009] Another aspect of the invention is directed to an apparatus
that includes a proximal member, a distal member that is
operatively configured with respect to the proximal member so that
an angle between the distal member and the proximal member can
vary, and a motor configured to vary the angle between the distal
member and the proximal member. A series elastic element is
connected between at least one of (a) the motor and the proximal
member and (b) the motor and the distal member, and at least one
sensor having an output from which a deflection of the series
elastic element can be determined. The apparatus also includes a
controller configured to determine a desired torque based on the
output, and to control the motor's torque based on the determined
desired torque.
BRIEF DESCRIPTION OF THE DRAWINGS
[0010] FIG. 1 is a schematic illustration of the phases of a user's
gait cycle when walking on level ground.
[0011] FIG. 2A depicts the statistic range of net non-conservative
work vs. walking speed for healthy human ankles.
[0012] FIG. 2B depicts the statistic range of peak-power vs.
walking speed for healthy human ankles.
[0013] FIG. 2C shows the net non-conservative work vs. walking
speed when two different equations are used to control a motor.
[0014] FIG. 2D shows peak-power vs. walking speed when two
different equations are used to control a motor.
[0015] FIG. 3A depicts the relationship between walking speed of
the upcoming step and the shank angular rate.
[0016] FIG. 3B depicts what shank angular rate is used in FIG.
3A.
[0017] FIG. 4A depicts one suitable gain function for use in
controlling the motor.
[0018] FIG. 4B depicts another suitable gain function.
[0019] FIG. 5A is a block diagram of an embodiment that relies on
torque sensing.
[0020] FIG. 5B depicts a mechanical configuration for the FIG. 5A
embodiment.
[0021] FIG. 6A is a block diagram of an embodiment that relies on
deflections and torque vs. deflection characteristics.
[0022] FIG. 6B depicts mechanical configuration for the FIG. 6A
embodiment.
[0023] FIG. 6C depicts a section view of the FIG. 6B
configuration.
[0024] FIG. 7 depicts a test fixture for measuring torque vs.
deflection characteristics.
[0025] FIG. 8A is a graph from which a spring rate can be
determined.
[0026] FIG. 8B is a graph depicting changes in a torque component
over time.
[0027] FIG. 9 depicts the torque vs. deflection characteristics for
a series elastic element.
[0028] FIG. 10 is a .GAMMA.-.THETA. plot for the stance-phase
torque-angle response of an intact ankle.
DESCRIPTION OF THE PREFERRED EMBODIMENTS
[0029] In healthy humans, the ankle-foot normally creates the
positive net-work and peak-power on each stride that the body needs
to achieve ordinary walk with metabolic efficiency. The net-work
and peak-power in the ankle during the stance of gait is highly
related to walking speed. FIGS. 2A and 2B depict this relationship.
More specifically, FIG. 2A shows the statistic range (+1 sigma
bounds) of net non-conservative work vs. walking speed, which lies
between the lines 11, 12. FIG. 2B shows the estimated statistic
ranges (+1 sigma bounds) of the peak-power vs. walking speed as
lines 16, 17. FIG. 2B also shows the mean value of peak-power vs.
walking speed (as measured in a study) as line 18, which lies
between lines 16 and 17.
[0030] The data points depicted by stars in FIG. 2C shows the net
non-conservative work vs. walking speed when Equation 1 above is
used to control the motor current. Note that net non-conservative
work can be determined by calculating the loop area, over one cycle
of ankle-torque vs. ankle angle (e.g., as seen in FIG. 10, starting
at point 1, passing through points, 2, 3, and 4 in sequence, and
returning to point 1. It can be seen that the net non-conservative
work is higher than the statistic range bounded by lines 11, 12 for
intact ankles, and the deviation from that range is larger at
slower walking speeds than it is at faster walking speeds.
Similarly, the data points depicted by stars in FIG. 2D show the
peak power vs. walking speed when Equation 1 above is used to
control the motor current. It can be seen that the peak power is
higher than the mean value line 18 for intact ankles. The net work
is also higher, and is wasted, causing extra heat and reduction in
battery life.
[0031] To more closely mimic the human ankle-foot biomechanics for
ordinary walk across a wide range of walking speeds, the
embodiments disclosed in the '384 application may be modified by
using the power control approach described herein so as to deliver
net-work and peak-power on each stride that more closely matches
the statistic ranges bounded by the lines 11, 12 in FIG. 2A, and
the mean line 18 in FIG. 2B. In this approach, a prediction of the
walking speed for the upcoming step is made, and that predicted
walking speed is used to set the ankle control parameters
(including setting of the power control gain) for the upcoming
step.
[0032] One way to predict the walking speed of the upcoming step is
based on the shank (pitch) angular rate .omega..sub.x based on the
relationship depicted in FIG. 3A. These two velocities are highly
linearly correlated such that the peak angular rate in stance phase
serves as an excellent prediction of the walk speed of the up
coming step. The correlation between walking speed and the shank
angular rate is present at various times during the stance and
swing phase, but it is preferable to minimize the latency between
the walking speed estimate and when it will be applied. One way to
accomplish this is to sample the shank angular rate at the very
start of controlled dorsiflexion (i.e., at foot-flat), immediately
before the reflex begins. This reduced latency ensures that a
reflex is not applied in certain situations, such as when the user
is stopping. If, on the other hand, a stale walking-speed
prediction were used, (e.g., by estimated walking speed from the
shank angular rate at the prior toe-off), the estimate might be
invalid (e.g., in situations where the user decides to stop
suddenly).
[0033] The shank angular rate may be measured by any suitable
means, such as an inertial measurement unit (IMU) or an angular
rate sensor (ARS). The IMU or ARS may be placed onto the top part
of the prosthesis or orthosis that is rigidly connected to a socket
such that shank angular rate, as depicted in FIG. 3B, can be
measured. In alternative embodiments, it could be mounted on the
foot structure. An example of a suitable angular rate sensor is the
Invensense IDG-300. In one preferred embodiment, the IMU can be
made from three orthogonally-aligned angular rate sensors such as
the Analog Devices ADXRS610, and three orthogonally-aligned
accelerometers such as the Freescale MMA7360L.
[0034] An advantage of using the angular rate sensing technique is
that it provides an instantaneous measure of angular rate just
prior to invoking the reflex control. More specifically, the
maximum angular rate in the stance phase can be calculated and
employed to adjust the reflex torque response during the controlled
dorsiflexion and powered plantar flexion phases of a step. This
reflex is largely responsible for generating the net-work and
peak-power that meet human ankle-foot needs for ordinary
walking.
[0035] The reflex torque generation is achieved via non-linear,
positive feedback between the series elastic element (SEE) motor
torque and ankle torque by controlling the motor using the
following equation:
Motor
Torque=Kv(.omega..sub.x).times.pff.times.(normalized_Torque).sup.n
Eq. 2
where Kv(.omega..sub.x) is a power control gain function related to
the maximum angular rate, an example of which is depicted in FIG.
4A; pff is the power control gain tuned for high walking speed;
normalized_Torque is the ankle torque, .GAMMA..sub.A, normalized by
a torque, .GAMMA..sub.0, (strongly related to users' weight); and n
is the power exponent, typically in the range of between 3 and 5
for level-ground walking. This is similar to Equation 1 above,
except that the right side of the equation is multiplied by a gain
function Kv(.omega..sub.x) that is selected to reduce the motor
torque for lower angular velocities, which correspond to slower
walking speeds. Note that the companion equation for converting a
desired motor torque to a drive current for the motor remains the
same for all embodiments described herein (i.e., Motor
Current=Motor Torque/kt, where kt is the motor torque
constant).
[0036] One suitable gain function Kv(.omega..sub.x) is depicted in
FIG. 4A, which starts at 0 when the angular rate is zero, and
increases linearly to 1 at an angular rate .omega..sub.TH that
corresponds to a fast walking speed. Above that threshold angular
rate .omega..sub.TH, the gain function Kv(.omega..sub.x) remains at
1. A suitable setting for the threshold .omega..sub.TH is an
angular rate that corresponds to a fast walking speed (e.g., an
angular rate that corresponds to a walking speed of between 1.5 and
1.75 meters per second). In some embodiments, the threshold point
may be settable by a prosthetist, preferably constrained to some
legal range (e.g., to an angular rate that corresponds to a walking
speed of between 1.25 and 2 meters per second). In other
embodiments, provisions for adjusting the .omega..sub.TH set point
within a legal range may even be made available to the end
user.
[0037] The result of multiplying the right side of Equation 2 by
Kv(.omega..sub.x) is that the motor will be driven by lower
currents for slower walk speeds. That will result in less torque at
slower walk speeds (as compared to when Equation 1 is used). When
this approach is used to control a prosthetic or orthotic ankle,
during the flat-foot portion of the gait the torque will initially
be zero. The ankle torque .GAMMA..sub.A will start to increase at
the end of the controlled dorsiflexion phase. In response to the
rising .GAMMA..sub.A, the controller will drive the motor based on
Equation 2, which will increase the torque further in a positive
feedback reflex response. This positive feedback continues until
prior to toe-off as the lower leg begins to lift the foot off the
ground. At this point the positive feedback is diminishing, so the
torque starts to drops off. The positive feedback is quenched at
toe-off because at that point there is nothing to push against,
which makes the torque fall off rapidly. In addition, the state
machine that controls the application of the reflex also
transitions to the swing phase where position control is used. Note
that operation of the state machine is described in the '384
application, which is incorporated herein by reference.
[0038] The speed based power control method of Equation 2 has been
implemented and tested on an iWalk.TM. Powerfoot.TM. BiOM.TM.
prosthetic ankle/foot. When Equation 2 was used to control the
motor, the net non-conservative work vs. walking speed is depicted
by the circle data points in FIG. 2C. A comparison between the
circle data points and the star data points (discussed above) in
FIG. 2C reveals that the net non-conservative work is closer to the
statistic range bounded by lines 11, 12 when Equation 2 is used.
Similarly, the circle data points in FIG. 2D show the peak power
vs. walking speed when Equation 2 above is used to control the
motor current. It can be seen that the peak power when Equation 2
is used is much closer to the mean value line 18 than when Equation
1 is used (indicated by the star data points in FIG. 2D). This
experiment result was obtained from a patient with weight of 240 lb
and shank length of 53 cm. The walk speed was measured using IMU
systems, and ranged from 0.8 m/s to 1.5 m/s. The system provided
smooth transitions of power when users randomly changed their
walking velocities.
[0039] In alternative embodiments, gain functions with other shapes
may be used instead of the ramp depicted in FIG. 4A. Preferably,
all such functions start at 0 when .omega..sub.x=0, end at 1, and
are monotonically increasing. Examples of suitable shapes for the
gain function include shapes that resemble (a) the first quadrant
of a sine curve; or (b) the third and fourth quadrants of a cosine
curve (scaled and offset so as to start at 0 and end at 1). Other
transition shapes, including smooth shapes and shapes with abrupt
changes, may also be used. For example, the curve depicted in FIG.
4B would operate to keep the power low for low walking speeds
(which would be suitable in certain situations like a classroom),
and increase it only if the speed goes over a threshold
.omega..sub.TH2. Optionally, the gain function may also be
operative for negative velocities to control the reflex response
when walking or running backward. For this reason, negative
velocities are included in FIG. 4B. If desired, the maximum gain
for negative velocities may be lower than 1, so as to provide a
smaller power boost when walking backwards In some embodiments, the
gain function could also be made to be a function of velocity when
side-stepping or hopping sideways.
[0040] In some embodiments, a user interface may be provided to
give the prosthetist control over the value of n in Equation 2,
preferably constrained within some legal range (e.g., between 2 and
7). Set points of between 3 and 5 have been found to be preferable.
Since normalized_Torque is .GAMMA..sub.A normalized by
.GAMMA..sub.0, when n is high (e.g., around 5), the current will
not rise until .GAMMA..sub.A gets closer to .GAMMA..sub.0. This
delays (in time) the onset of the positive feedback. Conversely,
when n is lower (e.g., around 3), the current will start to
increase before .GAMMA..sub.A gets too close to .GAMMA..sub.0. This
advances (in time) the onset of the positive feedback. When the
system is configured to give the prosthetist control over n, n can
be adjusted (e.g., based on verbal feedback from the end user) to
maximize the user's comfort. In other embodiments, a user interface
may be provided to give the end user control over n (within a legal
range).
[0041] In alternative embodiments, the reflex torque generation
equation may be modified to be as follows:
Motor
Torque=Kv(.omega..sub.x).times.pff.times.(normalized_Torque).sup.n-
f(.omega..sup.x.sup.) Eq. 3
Equation 3 is very similar to Equation 2, except that in Equation
3, the exponent n of the normalized_Torque is multiplied by a
function of the angular rate .omega..sub.x. The function
f(.omega..sub.x) is preferably selected so that the resulting
exponent is larger at higher angular velocities than it is at lower
angular velocities. This would operate to advance the onset of
reflex (in time) when the user is walking faster, with respect to
the timing when the user is walking slower.
[0042] Note that in the embodiments described above, the system
does not explicitly make a prediction of the walking speed for the
upcoming step. Instead, the system relies on the angular rate
.omega..sub.x of the shank (which, as described above, is
correlated to the predicted walking speed). In this case, the
angular rate .omega..sub.x of the shank serves as a surrogate for
the walking speed. In alternative embodiments, instead of relying
on the angular rate .omega..sub.x of the shank, other parameters
may be used to predict the walking speed. The ankle power would
then be adjusted accordingly based on the predicted walking speed
based on these alternative sensors. For example, the angular rate
of the leg section above the knee, or the knee linear moving
velocity in stance phase may be used to predict the walking speed
of the upcoming step. The Cartesian trajectory of the ankle or
knee, tracked using an IMU, could also be used to predict the
walking speed of the upcoming step.
[0043] In other embodiments, the equations may implemented so as to
explicitly compute the estimated walking speed as an intermediate
result, and then adjust the various parameters based on that
intermediate result to control the power and net non-conservative
work (e.g., by replacing Kv(.omega..sub.x) with Kv(speed) in
Equation 2).
[0044] Preferably, the system includes special-event handing to
modify the power level when it determines that a special walking
environment exists. For example, the power may be increased for
upstairs/up-ramp walking, even though the walk speed is low. Or the
power may be decreased for down stairs or down ramp walking even
though the walk speed is high. Note that the ankle trajectory or
knee trajectory (determined, for example, using an IMU) may be used
as a discriminator to determine if a special walking environment
exists, so that the characteristics of the ankle (including the
reflex) can be adjusted for the special walking environment.
[0045] The system described above provides users improved net-work
and peak-power to achieve normal biomechanics for ordinary walking
across a range of walking speeds. The system also uses reduced
motor current at low walking speeds, which is the case for the
majority of walking in most people's routines. This may help keep
the motor temperature low, save energy, and reduce the frequency of
recharging batteries and the need to carry spare batteries. Lower
currents also reduce the stress and fatigue on the drive
transmission, including the series-spring, and can increase the
design life of various components in the device.
[0046] The embodiments described above rely on the ankle torque
.GAMMA..sub.A as an input to the equations that ultimately control
the motor current during controlled dorsiflexion and powered
plantar flexion. This ankle torque .GAMMA..sub.A may be determined
by a number of approaches. One such approach, which is described in
the '384 application, is to actively measure the ankle torque
.GAMMA..sub.A using, for example, strain gauges arranged in a
Wheatstone bridge configuration to measure the torque applied by
the socket attachment at the top of the ankle prosthesis.
[0047] FIG. 5A is a system block diagram for this embodiment. The
prosthetic or orthotic ankle/foot includes a shank member 52 and a
foot member 54 operatively connected to permit plantarflexion and
dorsiflexion, e.g., by a joint 53. A motor 56 is affixed to the
shank member 52, and a series elastic element 58 sits between the
shank member 52 and the foot member 54, so that it will be in
series with the motor, as explained in U.S. Pat. No. 5,650,704,
which is incorporated herein by reference. Driving the motor in one
direction or the other will plantarflex or dorsiflex the foot
member 54 with respect to shank member 52. In alternative
embodiments (not shown) the positions of the motor 56 and the
series elastic element 58 could be swapped, in which case the motor
would be mounted to the foot member 54.
[0048] A torque sensor 66 measures the ankle torque .GAMMA..sub.A
and send an output that represents that torque to the controller
68. The controller 68 is programmed to control the motor 56 by
implementing Equation 2. In alternative embodiments, analog
circuitry configured to implement Equation 2 may be used in place
of the controller 68. The power driver 60 contains the drive
circuitry needed to convert the low level signals from the
controller 68 into the high power signals needed to drive the motor
56.
[0049] FIG. 5B depicts a practical mechanical configuration for
implementing the architecture shown in the FIG. 5A embodiment. In
FIG. 5B, the torque sensor 1732 (which corresponds to ref. # 66 in
FIG. 5A) is positioned at the very top of the shank member 1716
(which corresponds to ref. # 52 in FIG. 5A).
[0050] Another approach for determining the ankle torque
.GAMMA..sub.A is to break that torque up into its constituent
components, and analyze the torque of each of those components
separately. For example, in the design depicted in FIG. 6A-C, there
are two components that contribute to the total torque: the torque
applied by the series elastic element (.GAMMA..sub.S) and the
torque applied by the bumper (.GAMMA..sub.B). The bumper is
positioned between the shank portion of the ankle and the foot
portion, and can also be considered a hardstop when the stiffness
is high. In alternative embodiments, a spring may be used instead
of a bumper. Note that the .GAMMA..sub.B component only comes into
play during bumper engagement (i.e., during dorsiflexion, when the
shank member presses against a bumper that is affixed to the foot
member, or, in alternative embodiments, when the foot member
engages a bumper that is affixed to the shank member).
[0051] If each of the contributing components is known, the total
ankle torque can be determined by vector-adding .GAMMA..sub.S and
.GAMMA..sub.B (i.e., .GAMMA..sub.A=.GAMMA..sub.S+.GAMMA..sub.B). In
the design depicted in FIG. 6B, both .GAMMA..sub.S and
.GAMMA..sub.B can be determined as a function of displacement as
measured by position sensors that are distributed throughout the
design, like a motor encoder that detects the position of the motor
and an ankle angle encoder that detects the angle of the ankle
pivot.
[0052] We begin with .GAMMA..sub.S. In FIG. 6C, the motor 1B-102
drives a ballscrew 1B-106, and a digital encoder 1B-104 mounted on
the motor measures the ballscrew extension p. If the foot were to
be operated unloaded (e.g., when it is up in the air), for every
given value of ballscrew extension p, the ankle joint 1B-108 would
move to an angle .beta.(p). The .beta.(p) function can be
determined empirically by lifting the device in the air so that it
is unloaded, then driving the motor through its entire operating
range, and measuring the resulting angle of the ankle joint 1B-108
at each value of p. Alternatively, .beta.(p) could be calculated
based on the known geometry of the device. The .beta.(p) function
is stored in a memory that is accessible by the controller 78
(shown in FIG. 6A) in any suitable format (e.g., as an equation or
a lookup table).
[0053] During normal operation, the device will be loaded, and the
actual angle .theta. of the ankle joint 1B-108 can be determined
(e.g., by a high-resolution encoder, not shown, mounted on the
ankle joint). In addition, the actual ballscrew extension p can be
determined based on the output of the digital encoder 1B-104. The
controller inputs p from the motor encoder and retrieves the
unloaded angular position .beta.(p) from memory. It then inputs the
actual angle .theta. from the ankle joint angle encoder and
subtracts .beta.(p) from .theta. (i.e., the controller computes
.theta.-.beta.(p)). That difference is the angular deflection of
the SEE 1B-110. In some embodiments, a "single-turn" motor
controller can be used. At power on, its absolute position within
one motor turn and the absolute joint position can be used together
to determine the absolute displacement of the ballscrew in relation
to the end-of-travel in the plantarflexion direction.
[0054] After the deflection has been determined, the torque
.GAMMA..sub.S can be found because torque is a function of the
deflection. In a simple model, the torque vs. deflection
characteristics can be modeled as a linear function (Hooke's Law),
so that .GAMMA..sub.S=k.sub.S.times.deflection, where k.sub.S is
the spring rate for the SEE. FIG. 9 depicts the torque vs.
deflection characteristics for the series elastic element 1B-110
(shown in FIG. 6B). From these characteristics, a measured
deflection can be used to determine .GAMMA..sub.S. Note that
relying on an equation involving a spring constant k.sub.S is just
one of many possible ways to determine the torque from a
deflection, and alternative models and approaches for determining
the torque vs. deflection characteristics may also be used (e.g., a
lookup table, polynomial curve fitting, or non-linear
estimation).
[0055] We turn next to the .GAMMA..sub.B component. During
dorsiflexion, the shank member 1B-111 pushes towards the foot
member 1B-114, and a bumper 1B-112 that sits between those two
members (and could be affixed to either member) is compressed.
During testing of the previous generation designs, which used a
relatively soft plastic for the bumper 1B-112, the inventors
recognized that there is observable compliance in the bumper during
engagement, in the range of 0.25.degree. of deflection per 85 Nm
peak reference load for a 250 lb amputee. When harder plastics are
used (e.g., EPDM, with a 95 A durometer), there is much less
deflection (e.g., 0.1.degree. of deflection per 85 Nm peak
reference load for a 250 lb amputee), and the force-deflection
characteristic of this compliance became more stable and more
easily modeled. Note that the metal shells that house the ankle
mechanism will also flex measurably, and so can the foot structure
and the member that contacts the bumper. When the flexural
displacements are measured empirically for a particular design or
sample of a design (e.g., using a test fixture), all of those
flexures would be automatically accounted for.
[0056] The variation of .GAMMA..sub.B with the compression of the
bumper can be determined empirically for a given design or a
particular instantiation of a design. One way to do this is to bolt
a sample ankle/foot 250 into a test fixture 200, like the one shown
in FIG. 7. The test fixture 200 preferably uses a six
degree-of-freedom force-torque sensor 210 that simultaneously
measures force and torque along and about three orthogonal axes
(e.g., made by JR3, Inc.), with a backdrive ballscrew actuator 220
installed between the foot portion 252 of the ankle/foot 250 and
the JR3 210. In this test fixture 200, the ankle/foot 250 is driven
until the foot portion 252 makes initial contact with the bumper
(shown in FIG. 6B) on the shank portion 254 of the ankle/foot 250.
The angle of initial contact is defined as .theta..sub.I. Then,
using the backdrive ballscrew actuator 220, the foot portion 252 is
further driven to an angle .theta..sub.C. The angle .theta..sub.C
can be measured by the ankle encoder 1B-108 on the ankle/foot
prosthesis (shown in FIG. 6C). As .theta..sub.C increases, the
compression of the bumper increases, and the forces as determined
by the JR3 210 are stored for every possible angle
.theta..sub.C.
[0057] The Z (vertical) and Y (Horizontal) forces measured by the
JR3 210 are summed using vector mathematics to determine the force
along the backdrive screw axis. The ankle torque is then calculated
by multiplying the axial force by the perpendicular moment arm,
after subtracting any torque contribution from the SEE. The ankle
torque versus ankle angle is plotted for a number of cycles (e.g.,
10 cycles) for every possible angle .theta..sub.C and a least
squares best fit line is calculated, assuming a linear relationship
.GAMMA..sub.B=K.sub.S.times.(.theta..sub.C-.theta..sub.I), where
K.sub.S is the rotational spring rate for the bumper 1B-112. The
slope of the resulting best-fit line is the spring rate K.sub.S of
the bumper in Nm/rad as shown in FIG. 8A. In alternative
embodiments, instead of using this linear relationship to model the
bumper, alternative models and approaches for determining the
torque vs. deflection characteristics in the design may also be
used (e.g., a lookup table, polynomial curve fitting, or non-linear
estimation).
[0058] Note that when increasing the torque (i.e., when the foot
portion is being driven into the bumper and is compressing the
bumper), the relationship of the ankle torque to ankle angle
deflection is very linear. However when returning back to zero
(decreasing torque), the curve is different. This discrepancy is
due to the effect of the energy absorbing properties of the bumper.
It is preferable to use the slope of the least squares best fit
line for the increasing torque portion to determine the spring rate
K.sub.S of the bumper.
[0059] FIG. 8B depicts the .GAMMA..sub.B component of torque that
is determined using this approach over time in a situation where
the bumper is increasingly compressed for about half a second
(until the torque reaches -90), and then released. The quantized
nature of the .GAMMA..sub.B torque is a function of the encoder
resolution. This quantization can be minimized by utilizing higher
resolution encoders. In one preferred embodiment, a 13 bit encoder
(8196 counts/360 degrees) manufactured by Renishaw Inc (P/N
RMB13BC1) is used. The Renishaw encoder employs a custom
Hall-effect IC that measures the field angle arising from a
single-pole, cylindrical magnet mounted on the foot structure in
relation to the orientation of the IC affixed to a printed circuit
assembly embedded in the ankle shell. Filtering of the angle
measurement, using a FIR Low-Pass filter executing in a dedicated
DSP, has been shown to extend the effective resolution to between
15-16 bits.
[0060] Once the torque vs. deflection characteristics of a
bumper/ankle shell has been modeled (e.g., as explained above), the
.GAMMA..sub.B contribution at any given instant during operation of
the prosthesis can be determined by measuring .theta..sub.C and
plugging the result into the equation
.GAMMA..sub.B=K.sub.S.times.(.theta..sub.C-.theta..sub.I), or into
an alternative model that models .GAMMA..sub.B as a function of
.theta..sub.C. Thus, from a measured angular deflection
.theta..sub.C, the second torque component .GAMMA..sub.B can be
determined. In alternative embodiments, other ankle angle encoding
means could be employed to determine how far the bumper has been
compressed, including optical, magneto-restrictive and inductive
sensors.
[0061] At this point, both the .GAMMA..sub.S and .GAMMA..sub.B
components are known. .GAMMA..sub.S can now be added to
.GAMMA..sub.B to arrive at .GAMMA..sub.A, and the resulting
.GAMMA..sub.A is used as an input to Equation 2 to control the
motor.
[0062] FIG. 6A is a system block diagram for implementing this
approach by determining .GAMMA..sub.S and .GAMMA..sub.B separately
and then adding those components to arrive at .GAMMA..sub.A.
Elements 52-60 are the same as the correspondingly numbered
elements in FIG. 5A. Angular position sensors 76 measure the motor
displacement p and the ankle joint displacement .theta., and send
outputs representing those displacements to the controller 78. The
controller 78 is programmed to convert those displacements to
torque .GAMMA..sub.S as explained above. In addition, the
controller 78 is programmed to convert the ankle joint displacement
.theta. to torque .GAMMA..sub.B as explained above. The controller
78 then vector-adds .GAMMA..sub.S to .GAMMA..sub.B to determine
.GAMMA..sub.A. The controller 78 then controls the motor 56 (with
the assistance of the power driver 60, as in the FIG. 5A
embodiment) by implementing Equation 2.
[0063] As mentioned above, n in Equation 2 can be tuned to make the
device more comfortable for the user. Other parameters may also be
similarly tuned, such as pff and the threshold angular rate
.omega..sub.TH, which affects the Kv(.omega..sub.x) function in
Equation 2.
[0064] Referring now to FIG. 10, which is a .GAMMA.-.THETA. plot
for the stance-phase, body-mass-normalized torque-angle, response
of an intact ankle, additional parameters can be found that may be
tuned in a prosthesis or orthosis to try to better mimic the intact
ankle and thereby improve comfort and performance. Examples
include, modulating impedance as the ankle-foot transitions from
controlled plantar flexion (the slope of K.sub.1-2), through
controlled dorsiflexion (the slope of K.sub.2-3), to powered
plantarflexion (the slope of K.sub.3-4). The initial values of
these three impedances, and the initial value of .theta. at toe-off
(.theta.*.sub.TOE-OFF) can be derived from the mean .GAMMA.-.THETA.
response of intact ankles, and those initial values can then be
tuned to suit the activity level, limb length, body-mass
distribution and preferences of an individual user.
[0065] In the above-described embodiments, a single motor is used
to implement both plantarflexion and dorsiflexion. But in
alternative embodiments, that motor could be replaced by one motor
for implementing plantarflexion, and another component for
implementing dorsiflexion. In other alternative embodiments, a
plurality of motors may be arranged in parallel to perform both
plantarflexion and dorsiflexion. In still other embodiments, the
electric motors described above can be replaced with other types of
motors (e.g., hydraulic motors), in which case the controller and
the power driver will have to be adjusted accordingly.
[0066] Note that while the concepts described above are explained
in the context of prostheses, they can also be applied in the
context of orthoses. In addition, while the embodiments described
above all relate to ankles, the above-described concepts can be
applied in other prosthetic and orthotic applications, such as
hips, torso, and arms, in which case suitable modification should
be made that will be appreciated by persons skilled in the relevant
arts. For example, in the context of a knee, where the reflex
occurs right during toe-off, the walking speed prediction would use
"fresh" shank speed measurement just prior to toe-off. In those
other contexts, the shank member can be generalized as a proximal
member, the foot member can be generalized as a distal member, and
dorsiflexion/plantarflexion can be generalized as varying the angle
between the distal member and the proximal member. The
above-described concepts can also be applied in the context of
humanoid robots.
[0067] While the present invention has been disclosed with
reference to certain embodiments, numerous modifications,
alterations, and changes to the described embodiments are possible
without departing from the sphere and scope of the present
invention, as defined in the appended claims. Accordingly, it is
intended that the present invention not be limited to the described
embodiments, but that it has the full scope defined by the language
of the following claims, and equivalents thereof.
* * * * *