U.S. patent application number 15/459565 was filed with the patent office on 2017-08-31 for biomimetic transfemoral prosthesis.
The applicant listed for this patent is BionX Medical Technologies, Inc.. Invention is credited to Christopher Eric Barnhart, Richard J. Casler, JR., Zhixiu Han, Hugh Miller Herr, Charles E. Rohrs, Christopher Williams.
Application Number | 20170250632 15/459565 |
Document ID | / |
Family ID | 47279011 |
Filed Date | 2017-08-31 |
United States Patent
Application |
20170250632 |
Kind Code |
A1 |
Herr; Hugh Miller ; et
al. |
August 31, 2017 |
BIOMIMETIC TRANSFEMORAL PROSTHESIS
Abstract
In an artificial limb system having an actuator coupled to a
joint for applying a torque characteristic thereto, a control
bandwidth of a motor controller for a motor included in the
actuator can be increased by augmenting a current feedback loop in
the motor controller with a feed forward of estimated back
electromotive force (emf) voltage associated with, the motor.
Alternatively, the current loop is eliminated and replaced with a
voltage loop related to joint torque. The voltage loop may also be
augmented with the feed forward of estimated back emf, to improve
the robustness of the motor controller.
Inventors: |
Herr; Hugh Miller;
(Somerville, MA) ; Williams; Christopher;
(Pittsburgh, PA) ; Barnhart; Christopher Eric;
(Carlisle, MA) ; Han; Zhixiu; (Acton, MA) ;
Rohrs; Charles E.; (Newton, MA) ; Casler, JR.;
Richard J.; (Lowell, MA) |
|
Applicant: |
Name |
City |
State |
Country |
Type |
BionX Medical Technologies, Inc. |
Bedford |
MA |
US |
|
|
Family ID: |
47279011 |
Appl. No.: |
15/459565 |
Filed: |
March 15, 2017 |
Related U.S. Patent Documents
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Application
Number |
Filing Date |
Patent Number |
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14355657 |
May 1, 2014 |
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PCT/US2012/063395 |
Nov 2, 2012 |
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15459565 |
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61554921 |
Nov 2, 2011 |
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Current U.S.
Class: |
1/1 |
Current CPC
Class: |
A61F 2002/7625 20130101;
A61F 2002/4667 20130101; A61F 2002/6818 20130101; A61F 2002/704
20130101; A61F 5/0111 20130101; A61F 2/54 20130101; A61F 2/60
20130101; A61F 2/64 20130101; A61F 2/6607 20130101; A61F 2/4657
20130101; A61F 2/68 20130101; A61F 2002/6827 20130101; A61F
2005/0155 20130101; A61F 2/70 20130101; A61F 2002/701 20130101;
A61F 2002/7645 20130101; A61F 5/0123 20130101 |
International
Class: |
A61F 2/38 20060101
A61F002/38; G06F 19/00 20110101 G06F019/00; A61F 2/48 20060101
A61F002/48 |
Claims
1-22. (canceled)
23. A method for controlling a motorized, artificial limb having an
actuator, the actuator comprising a motor coupled in series with an
elastic element, to apply a torque characteristic to a joint, the
method comprising the steps of: computing target motor current
required to achieve substantially a target torque characteristic at
the joint; measuring motor current and motor speed; computing a
current error as a difference between the target motor current and
the measured motor current; estimating voltage corresponding to
back electromotive force related to the motor speed; and
controlling voltage applied to motor windings based on, at least in
part, (i) the estimated voltage corresponding to the back
electromotive force, and (ii) the current error, to reduce the
current error.
24. The method of claim 23, wherein the torque characteristic
comprises at least one of a joint torque, a joint impedance, and a
joint equilibrium.
25. The method of claim 23, wherein the artificial limb comprises
at least one of a prosthetic limb and an orthotic limb.
26. The method of claim 23, wherein the joint is at least one of an
ankle joint, a knee joint, and a hip joint.
27. An artificial limb system comprising: an actuator coupled to a
joint for applying a torque characteristic thereto, the actuator
having a motor coupled in series with an elastic element; a power
source for applying a voltage to windings of the motor; a current
sensor for measuring motor current; a motor encoder for measuring
motor speed; and a controller configured to: compute a current
error as a difference between a target current and the measured
motor current; estimate voltage corresponding to back electromotive
force related to the measured motor speed; and control voltage
applied to motor windings based on, at least in part, (i) the
estimated voltage corresponding to the back electromotive force,
and (ii) the current error, to reduce the current error.
28. The system of claim 27, wherein the artificial limb system
comprises at least one of a prosthetic limb system and an orthotic
limb system.
29. The system of claim 27, wherein the joint is at least one of an
ankle joint, a knee joint, and a hip joint.
30. The system of claim 27, wherein the controller is adapted for
determining the target motor current to achieve the target joint
torque characteristic.
31. The method of claim 23, wherein the back electromotive force is
estimated using a current-loop controller.
32. The method of claim 23, wherein the back electromotive force is
estimated using a direct measurement of joint torque.
33. The method of claim 23, wherein the target torque
characteristic is controlled by a joint torque controller
comprising a forward path and a feedback path, and a resonant
spring dynamic of the artificial limb appears in the forward path
and not the feedback path.
34. The method of claim 23, further comprising estimating a speed
of the motor, and using the estimated speed to cancel the back
electromotive force.
35. The method of claim 23, wherein the target torque
characteristic is controlled by a joint torque controller that
applies a proportional-integral (PI) algorithm based on the motor
current.
36. The method of claim 23, wherein the target torque
characteristic is controlled by a joint torque controller that
applies a proportional-integral-derivative (PID) algorithm based on
a torque of the joint.
37. The system of claim 27, wherein the controller estimates the
back electromotive force.
38. The system of claim 27, wherein the back electromotive force is
estimated using a direct measurement of joint torque.
39. The system of claim 27, wherein the controller is further
configured to control the target torque characteristic using a
forward path and a feedback path, and a resonant spring dynamic of
the artificial limb appears in the forward path and not the
feedback path.
40. The system of claim 27, wherein the controller is further
configured to estimate a speed of the motor, and use the estimated
speed to cancel the back electromotive force.
41. The system of claim 27, wherein the controller is further
configured to control the target torque characteristic by applying
a proportional-integral (PI) algorithm based on the motor
current.
42. The system of claim 27, wherein the controller is further
configured to control the target torque characteristic by applying
a proportional-integral-derivative (PID) algorithm based on a
torque of the joint.
Description
RELATED APPLICATIONS
[0001] This application claims priority to and the benefit of U.S.
Provisional Patent Application Ser. No. 61/554,921 filed on Nov. 2,
2011, the disclosure of which is hereby incorporated herein by
reference in its entirety. The application also relates to U.S.
Pat. No. 8,074,633, U.S. patent application Ser. Nos. 12/551,845,
12/552,013, 12/552,021, 12/552,028, 12/552,036, 12/872,425,
13/079,564, 13/079,571, 13/347,443, 13/349,216, 13/356,230, and
13/417,949, and U.S. Provisional Patent Application Nos.
61/554,921, 61,595,453, 61,658,568, 61/659,729, 61/662,104,
61/649,640, 61/659,723, 61/679,194, and 61/691,684, the disclosures
of which are each hereby incorporated herein by reference in their
entireties.
FIELD OF THE INVENTION
[0002] This invention relates generally to systems and methods for
lower body locomotion, and more particularly to the control of
prosthetic, orthotic, and exoskeleton devices usable at various
points of the leg, including the ankle, knee, and hip.
BACKGROUND
[0003] Over 120,000 people suffer lower-extremity amputations each
year in the United States, of which transfemoral (above-knee)
amputations account for over 40%. Over 30,000 people in this
transfemoral population require a new prosthetic limb each
year--typically a passive, microprocessor-controlled knee joint,
employing hydraulic damping and a passive carbon-fiber ankle-foot
prosthesis. Such passive leg systems tend not to be biomimetic;
instead they ten to be passive-elastic during stance and can
neither perform net non-conservative work to propel the amputee
upward and forward nor deliver the temporal torque response
supplied by an intact knee and ankle joint during the gait cycle,
and hence fail to fully restore function when integrated onto the
residual limb. Researchers have hypothesized that the inability of
conventional passive-elastic ankle-foot prostheses to provide
sufficient positive power at terminal stance to limit heel strike
losses of the adjacent leg is a key mechanism for the increased
metabolic rate of walking amputees. These limitations in both
ankle-foot and knee designs contribute to the severity of clinical
problems experienced by transfemoral amputees.
[0004] Current leg prostheses ten not to provide the balance
desired by the transfemoral community. Amputees often fall,
especially while traversing rough or irregular terrain. This may be
at lease partially due because most ankle-foot prostheses fail to
actively control a zero-moment point (ZMP) at the foot-ground
interface, a balancing strategy sometimes employed in the field of
humanoid robotics. In addition to balance problems, amputees tend
to tire easier and walk slower than non-amputees. For example,
amputees can require 10-60% more metabolic energy to walk than
intact persons. The actual differences for any individual at a
particular time result from differences in walking speed, physical
fitness level, cause of amputation, level of amputations, and
prosthetic intervention characteristics. Amputees may walk at much
slower (e.g., 11-40%) slower self-selected gait speeds than do
person with intact limbs.
[0005] Integration of non-biomimetic ankle-foot and
microprocessor-controlled knee prostheses has confounded
researchers. Although some improvements in gait have been observed
with variable-damper knee designs, many problems still remain for
transfemoral amputees. For example, a variable-damper knee combined
with a passive-elastic ankle-foot prosthesis offers little to no
improvement in gait metabolism and walking speed compared to a
mechanically-passive transfemoral system. Although some powered
knee systems have been developed, these tend not to be
biomechanically-conceived, and can take hours to fit and tune to a
specific wearer. These powered knees tend to have a noisy motor and
transmission system. Other powered robotic leg systems often
exhibit three fundamental limitations: inefficient actuator design,
non-biomimetic actuator control software, and poorly executed
terrain-adaption software. Without a biologically-conceived
actuator, the motor must be made larger and heavier to deliver
necessary joint powers.
[0006] To deliver the increased power, high gear-ratio
transmissions are normally employed in powered leg prosthetic and
orthotic applications, driven by high RPM brushless motors that are
operated at currents in excess of 10.times. the rated current. The
result is often a noisy transmission that dissipates battery power
excessively, heating the motor windings instead of applying power
to the joint. As a result, batteries must be made larger than
necessary, or else the range on a battery charge is compromised
unnecessarily. Further, motor heating can be excessive when
extended periods of walking (e.g., hundreds of steps consecutively)
are applied. The useful range of the prosthesis may be constrained
by the need to "fold back" power when the motor winding get too
hot. Most robots are programmed explicitly in a position-controlled
or playback mode. Biophysically-conceived robots employ
mono-articular and biarticular bionic muscle-tendon units that
modulate joint impedance, equilibrium, torque and positive-feedback
reflex during a gait cycle. The behavior of these bionic systems
can be encoded in a relatively few parameters, implicitly defined
rather than explicitly defined. Indeed, only a few parameters need
to be changed to emulate biological behavior. In contrast, the
explicitly controlled systems often require that movement as
defined by joint angle trajectories is tuned to match biological
behavior to create a response for every special case, driven by
speed, terrain modality and wearer athleticism/payload. Even the
most experienced clinicians may be unable to set up such a system.
While many current robotic leg prostheses employ inertial
componentry to estimate terrain modality, these are usually
configured to adapt to the new terrain modality after several
steps, whereas an intact person adjusts to terrain modality within
each step. Though certain implementations of such a control may
rely on playing back a temporal response at a particular terrain
state, the desired behavior can be virtually impossible to
tune.
[0007] It is therefore desirable to provide leg devices that
provide a biomimetic response throughout a walking cycle.
SUMMARY OF THE INVENTION
[0008] Lower-extremity augmentation may rely on a muscle-tendon
architecture employing series-elasticity as a means of amplifying
available joint power, reducing motor work, and improving shock
tolerance. An intact ankle-foot for instance employs a
series-elastic actuator in the form of the calf muscle (motor)
driving through the Achilles tendon (series-elasticity). A model of
an intact human ankle progressing through various phases of a gait
cycle is depicted in FIG. 1. Elastic energy is stored in the tendon
in the controlled dorsiflexion phase and released later, like a
catapult, in powered plantar flexion to augment the power applied
by the calf muscle. A series elastic actuator (SEA) as described
herein can provide a biomimetic response, and may be capable of
amplifying power by greater than a factor of two.
[0009] A transfemoral prosthesis may include the SEA, motor
technology, neuromuscular-inspired actuator control, and intrinsic
inertial sensing to provide quiet and efficient biomimetic
mechanical behaviors across distinct walking speeds and terrains
and the transitions between these ground surfaces. The prosthesis
can provide the amputee with an enhanced metabolic economy when
using a powered, ankle-foot prosthesis compared to the metabolic
cost when using a conventional passive-elastic prosthesis,
including a normalization of self-selected walking speed and
improvement in metabolic cost-of-transport across a broad walking
speed range (0.75-1.75 m/sec) using sensing, control and
muscle-tendon unit (MTU) actuators.
[0010] The prosthesis may integrate an artificial MTU that employs
a high-torque, transverse-flux motor that reduces, in some
instances over an order of magnitude, the normalized motor copper
loss, R/k.sub.f.sup.2, in relation to conventional, radial flux
motor topologies employed in the many existing robotic augmentation
devices. The use of a direct-drive, transverse flux motor in an MTU
may result in a much lower drive transmission ration (e.g., six
times lower than a typical 200:1 gear-ratio transmission of a
typical motor), allowing for operation at lower motor RPM to
provide a more quiet and efficient device. Another result may be a
higher resonant frequency for the MTU, allowing for higher fidelity
joint torque control.
[0011] The device may also include neuromuscular-inspired control
software to emulate the biological response of an intact limb,
aiding in the normalization of metabolic walking economy and
biomechanical response. Intrinsic inertial sensing can enable
in-situ, real-time reconstruction of hip, knee and ankle trajectory
throughout the gait cycle. Pattern recognition algorithms may be
applied to these reconstructed trajectories for intra-step terrain
discrimination and adaptation, thereby enabling seamless terrain
transitions across walking speeds.
[0012] A wearer of the biomimetic prosthesis may experience and
improved walking speed, metabolic economy, gait symmetry, and gain
stability across level, sloped, and stair ground surfaces when
compared to conventional prostheses. This restored function by
objective metabolic and biomechanical measures can provide profound
improvements in the quality of life of a wearer. While the device
is often described with respect to a prosthetic for an amputee
population, the robotic augmentation platform described here may be
useful to researchers and educators focused on humanoid robots.
Open interfaces may be provided to configure, refine, and deploy
new applications.
[0013] According to one aspect, the invention relates to a method
for controlling a motorized, artificial limb having an actuator.
The actuator includes a motor coupled in series with an elastic
element to apply a torque characteristic to a joint. The method
includes applying a voltage to windings of the motor, measuring the
torque characteristic at the joint, computing a torque
characteristic error as a difference between a target torque
characteristic and the measured torque characteristic, and
controlling the applied voltage independently of motor current to
reduce the torque characteristic error.
[0014] In some embodiments, the torque characteristic is a joint
torque, a joint impedance, or a joint equilibrium. The applied
voltage may be controlled solely based on the target torque
characteristic and the torque-characteristic error. In certain
embodiments, the controlling step avoids computation of motor
current and/or computation of an adjustment to the motor current
supplied to the motor to achieve the target torque characteristic,
and may avoid measurement of motor current and using the measured
current for adjusting the motor current (at least in part).
[0015] In other embodiments, the limb control method is independent
of a resonance frequency associated with the coupling of the motor
and the elastic element. The step of measuring the torque
characteristic at the joint may include measuring an angular
position of the motor and/or an angular position of the joint. The
method may also include measuring motor speed, estimating voltage
corresponding to back electromotive force related to the motor
speed, and controlling the applied voltage based on, at least in
part, the estimated voltage corresponding to the back electromotive
force. In certain embodiments, the artificial limb is a prosthetic
limb and/or an orthotic limb. The joint may be an ankle joint or a
knee joint.
[0016] In another aspect, the invention relates to an artificial
limb system. The system includes an actuator coupled to a joint for
applying a torque characteristic thereto, the actuator having a
motor coupled in series with an elastic element. The system also
includes a power source for applying a voltage to winding of the
motor, at least one sensor for estimating the torque characteristic
at the joint, and a controller for: (i) computing a torque
characteristic error as a difference between a target torque
characteristic and the torque characteristic measured by the
sensor, and (ii) controlling the applied voltage independently of
motor current, to reduce the torque characteristic error.
[0017] In some embodiments, the torque characteristic comprises a
joint torque, a joint impedance, and/or a joint equilibrium. The
controller may be adapted for controlling the applied voltage
solely based on the target torque characteristic and the
torque-characteristic error. In other embodiments, the controller
is adapted for controlling the applied voltage without using a
computation of motor current and/or an adjustment to the motor
current supplied to the motor to achieve the target torque
characteristic. The controller may be adapted to avoid at least
partially using a measurement of motor current for an adjustment
thereof. The controller may be adapted such that a controller
response is independent of a resonance frequency associated with
the coupling of the motor and the elastic element.
[0018] In certain embodiments, the artificial limb system is a
prosthetic limb system and/or an orthotic limb system. The joint
may be an ankle joint or a knee joint. The sensor may be a joint
encoder, a torque sensor, a deflection sensor disposed on the
series elastic element, a joint angle sensor, or a motor angle
sensor. In some embodiments the system includes a motor encoder for
measuring motor speed with the controller adapted to: (i) estimate
voltage corresponding to back electromotive force related to the
motor speed, and (ii) control the applied voltage based on, at
least in part, the estimated voltage corresponding to the back
electromotive force. The system may include an observer to estimate
voltage corresponding to back electromotive force with the
controller adapted to control the applied voltage based on, at
least in part, the estimated voltage corresponding to the back
electromotive force. The system may include a linkage having a
plurality of links where the linkage is coupled to the joint, the
motor is coupled to a first link in the linkage, and the elastic
element is coupled to a second link in the linkage.
[0019] In yet another aspect, the invention relates to a method for
controlling a motorized, artificial limb. The artificial limb has
an actuator with a motor coupled in series with an elastic element
to apply a torque characteristic to a joint. The method includes
computing target motor current required to achieve substantially a
target torque characteristic at the joint, measuring motor current
and motor speed, and computing a current error as a difference
between the target motor current and the measured motor current.
The method also includes estimating voltage corresponding to back
electromotive force related to the motor speed and controlling
voltage applied to motor winding based on, at least in part, (i)
the estimated voltage corresponding to the back electromotive
force, and (ii) the current error, to reduce the current error.
[0020] In certain embodiments, the torque characteristic is a joint
torque, a joint impedance, and/or a joint equilibrium. The
artificial limb may be a prosthetic limb and/or an orthotic limb.
The joint may be an ankle joint, a knee joint, and a hip joint.
[0021] In still another aspect, the invention relates to an
artificial limb system. The system includes an actuator with a
motor coupled in series with an elastic element coupled to a joint
for applying a torque characteristic a power source for applying a
voltage to windings of the motor, a current sensor for measuring
motor current, and a motor encoder for measuring motor speed. The
system also includes a controller for computing a current error as
a different between a target current and the measured motor
current, estimating voltage corresponding to back electromotive
force related to the measured motor speed, and controlling voltage
applied to motor winding based on, at least in part, (i) the
estimated voltage corresponding to the back electromotive force,
and (ii) the current error, to reduce the current error.
[0022] In some embodiments, the artificial limb system is a
prosthetic limb system and/or an orthotic limb system. The joint
may be an ankle joint, a knee joint, or a hip joint. The controller
may be adapted to determine the target motor current to achieve the
target joint torque characteristic.
BRIEF DESCRIPTION OF THE FIGURES
[0023] Other features and advantages of the present invention, as
well as the invention itself, can be more fully understood from the
following description of the various embodiments, when read
together with the accompanying drawings, in which:
[0024] FIG. 1 is a diagram of a healthy ankle throughout the phases
of a walking cycle;
[0025] FIG. 2 is a schematic perspective view of an ankle device,
in accordance with one embodiment of the invention;
[0026] FIG. 3 is a graph depicting the power output of the ankle
device, in accordance with one embodiment of the invention;
[0027] FIG. 4A is a schematic cross-section view of a knee device,
in accordance with one embodiment of the invention;
[0028] FIG. 4B is a schematic diagram of the knee device of FIG.
4A;
[0029] FIG. 4C is a schematic side view of the knee device of FIG.
4A;
[0030] FIG. 5 is a schematic side view of a leg device, in
accordance with one embodiment of the invention;
[0031] FIGS. 6A and 6B are schematic cross-section and rear
perspective views, respectively, of a leg device in accordance with
another embodiment of the invention;
[0032] FIGS. 7A and 7B are graphs of joint trajectories for an
ankle and knee, respectively, in various modalities;
[0033] FIG. 8 is a block diagram of a control system for a leg
device, in accordance with one embodiment of the invention;
[0034] FIG. 9 is a block diagram of signals associated with a motor
and joint system, in accordance with one embodiment of the
invention;
[0035] FIG. 10 is a block diagram of a current feedback loop, in
accordance with one embodiment of the invention;
[0036] FIG. 11 is a block diagram of a system for directly
controlling joint torque, in accordance with one embodiment of the
invention;
[0037] FIG. 12 is a block diagram of an emf feedforward with
current feedback loop, in accordance with one embodiment of the
invention;
[0038] FIG. 13 is a block diagram of an emf feedforward with joint
torque feedback loop, in accordance with one embodiment of the
invention;
[0039] FIG. 14 is a block diagram of a model for using motor
current to compute a feedback signal to manipulate voltage to
control motor torque, in accordance with one embodiment of the
invention;
[0040] FIG. 15 is a block diagram of a model for using motor
current to compute a feedback signal to manipulate voltage to
control motor torque without an inner current-feedback loop, in
accordance with one embodiment of the invention;
[0041] FIG. 16 is a block diagram of a model with emf feedforward,
in accordance with one embodiment of the invention;
[0042] FIG. 17 is a block diagram of a model for using a
measurement of joint torque through a single compensator to compute
a voltage signal, in accordance with one embodiment of the
invention; and
[0043] FIG. 18 is a block diagram of a model for using a
measurement of joint torque through two compensators to compute a
voltage signal, in accordance with one embodiment of the
invention.
DETAILED DESCRIPTION
[0044] Devices for transfemoral amputees typically include at least
a foot-ankle device and a knee device. The foot-ankle and knee
devices described herein may be used together or separately,
particularly only a foot-ankle device for below-knee amputees.
While the embodiments described relate to prostheses, the concepts
contained herein may also be useful in other applications,
including orthoses and exoskeletons.
[0045] One ankle device 200 capable of restoring ankle function as
defined by objective metabolic and biomechanical measures is
depicted in FIG. 2. The ankle device 200 is capable of varying
ankle impedance during the early to mid-stance periods of walking,
emulating the quasi-static stiffness of an intact biological ankle.
In addition, the prosthesis 200 provides a sufficiently large
instantaneous power output and torque to propel an amputee upward
and forward during powered plantar flexion, while still matching
the size and weight of an intact ankle-foot complex, e.g.,
approximately 2.4% of body weight. The ankle device 200 may be a
variety of weights for different applications. For example, a
device 200 weighing approximately 4.8 lbs is appropriate for
wearers weighing approximately 190-250 lbs.
[0046] The ankle device 200 may include several different
technologies, including a series-elastic actuator (SEA) 202,
integral joint torque sensor 204, six degree-of-freedom inertial
measurement unit (IMU) 206, battery 208, and a controller with
control firmware 210 (intrinsic within the system). The SEA 202 is
capable of modulating torque, impedance and position in accordance
with the gait cycle state. Such may be nuanced as a function of
sensed walking speed and terrain modality. Through use of the
intrinsic control firmware 210, the SEA 202 may emulate the calf
muscle/Achilles tendon reflex during late stance to achieve
biomimetic operation. The SEA 202 design may be optimized to
amplify peak joint mechanical output power, e.g., 2.times..sup.+
amplification during fast walking, as illustrated in FIG. 3. The
SEA 202 may have a high-RPM brushless motor 212, a hybrid
belt/ball-screw transmission 214, and a carbon-fiber spring 216.
The transmission 214 can have an L1 design life of greater than 5
million cycles, and the SEA 202 may operate at an overall
motor-joint gear ratio of about 220:1.
[0047] The joint torque sensor 204 may be used to precisely sense
joint angle and SEA motor 212 position. The joint torque may be
computed in real-time using factory calibrated models of ankle
shell and carbon-fiber spring deflection, even without strain
measurement devices. The IMU 206 may measure angular rate and
acceleration in each of three orthogonally-opposed directions.
These measures may be used selectively to detect gait cycle state
transitions and instantaneous walking speed. The batter 208 is used
to power the device, and may be modular to aid in insertion and
removal. This is particularly helpful when the battery 208 is
replaced multiple times in a day, e.g., two or three times. Many
different types of batteries 208 may be used, including a lithium
polymer battery.
[0048] The control firmware 210 may rely on a state machine for
transitioning between various control schemes based on the phase of
gait. Sensors may be used to sense the transitions while walking.
For example, the firmware 210 can use intrinsic sensing to track
gait cycle transitions--enabling appropriate modulation of torque
(including SEA reflex), impedance and position. The device 200 may
employ non-linear, positive torque feedback to simulate muscle
response. This can provide a wearer with a natural (biomimetic)
feel. Other models, e.g., a neuromuscular muscle model, may be
used.
[0049] These component technologies may be used separately or in
conjunction with each other, with several possible alterations for
each. For example, the SEA 202 may have a transverse-flux motor 212
that eliminates the belt transmission 214, instead relying on a
direct-drive ball-screw implementation that can reduce the gear
ratio (e.g., by a factor of six). Power consumption may be reduced
(in some cases by more than 50%), and acoustic noise in the 1.5-3
kHz band may be reduce by up to, and sometimes greater than,
80%.
[0050] Powered prosthetic knees may use a variety of different
components in different arrangements. In one embodiment, a
prosthetic knee 400 depicted in FIGS. 4A-4C provides a biomimetic
response through the use of an agonist-antagonist actuator (AAA)
system 402. The AAA-based knee prosthesis 400 may emulate the
synergistic muscle activity of an intact limb by using a dual-SEA
system 402 that resembles the major mono-articular muscle groups
that flex and extend the human knee joint. Each SEA 404 includes a
torque source 406 (e.g., a motor), an elastic element 408 (e.g., a
spring) in series with the torque source 406, and a transmission
410 (e.g., a ballscrew and ballnut drive). In some embodiments, the
SEA 404, along with additional torque sensing, inertial sensing,
and intrinsic control, may be the same or similar to those used in
the ankle device 200, adapted for use in the knee 400 when
necessary. A floating point 412 may be coupled to an upper leg
attachment for the knee 400.
[0051] The mechanical architecture of the AAA knee 400 allows for
independent engagement of flexion and extension tendon-like, series
springs 408 for modulating joint impedance and non-conservative
motive power during the stance period, and joint position when the
prosthetic foot is off the ground. Furthermore, this architecture
may permit joint rotation with near zero friction via disengagement
of both SEAs 404, allowing a controller 414 (that may be intrinsic
to the system) to take advantage of the passive (ballistic)
dynamics of the system 402 in the swing phase, and thus, augment
the overall energetic economy of the prosthesis 400 and human
wearer. By matching the stiffness of the two series springs 408 to
match the biomechanical extension and flexion stiffness, modulation
of the stiffness can be accomplished with approximately
zero-electrical power by simply shorting the motor windings (to
apply a high-degree of damping and holding torque) of the
appropriate motor 406. The clutching feature can be utilized to
great advantage in a transverse-flux motor implementation due to
the transverse-flux motors' superior brake/clutch performance
resulting from its low motor resistance, R.sub.m, in relation to
the motor torque constant, R.sub.m/k.sub.t.sup.2, which defines the
motor damping with shorted leads. Certain embodiments of the knee
prosthesis 400 may have a single motor with bilateral spring
elements that can be always fully engaged as in a traditional SEA,
or can be configured to have a "dead zone" in which neither spring
is engaged, enabling agonist-antagonist "clutch" implementations. A
leg prosthesis 500 including the ankle device 200 and a knee device
502 (similar to the knee device 400) coupled together is depicted
in FIG. 5.
[0052] Another embodiment of a leg prosthesis 600 is depicted in
FIGS. 6A and 6B. The substantially fully-powered transfemoral
prosthesis 600 may have a transverse flux, series-elastic ankle 602
and knee 604 to produce biological ankle and knee mechanics for
steady-state level-ground walking with a prosthetic mass, shape and
acoustic noise output comparable to a biological leg. For example,
the prosthesis may have a mass <5% body weight, a shape smaller
than a 50% person, and an acoustic output of <45 dB one meter
from the prosthesis. Such a transfemoral prosthesis 600 may
normalize metabolic economy and self-selected walking speed during
ground surface ambulation in a variety of contexts, including
during level-ground, sloped (e.g., up to .+-.10 degrees and
beyond), stair ascent, stair descent, and transitions therebetween.
The prosthesis 600 may be able to adapt ankle and knee mechanics
within a single walking cycle to enable a transfemoral amputee to
walk across these ground surface variations in a manner that
emulates biological leg biomechanics.
[0053] The ankle device 602 may be substantially similar to the
ankle device 200, with a transverse flux motor 606 used instead of
a belt transmission. Using biomimetic, muscle-tendon-like actuators
(e.g., SEAs) that employ such high-torque transverse flux motors
606 can help reduce power consumption and acoustic noise levels by
significant amounts (e.g., over 50% and 80% respectively) when
compared to conventional (radial flux) motor transmissions to
achieve cool, quiet operation across substantially all walking
speeds. High-torque transverse flux motors 606 may be available
from Electric Torque Machines (Flagstaff, Ariz.). Another benefit
of transverse-flux motor technology is the reduction of motor
copper loss per square unit of torque production (R/k.sub.t.sup.2),
which may be over a factor of nearly 300 at the same gear ratio as
would be employed by conventional high-RPM radial flux motors like
the EC30 PowerMax motor (www.maxon.com) used in most
lower-extremity augmentation systems built by the research
community. This attenuation in loss can be used to reduce bath
transmission gear ratio (and hence motor rotational speed) and
motor temperature rise. The result is quiet and efficient operation
that extends range on a battery charge and reduces or eliminates
heating-related range limitations. Actuator design life may be
extended because friction and travel are reduced in concert with
gear ratio reduction.
[0054] The knee device (or muscle tendon unit; "MTU") 604 may
employ a "buckled-beam" SEA design with a gear ratio, in this
embodiment approximately 30:1. The buckled beam can be light-weight
and yet store up to 30 J/kg (or more) of energy, enough suitable
for running and jumping. In some embodiments, the buckled-beam SEA
may have a stiffness of approximately 4 Nm/rad/kg and energy
storage of approximately 8 J/kg for level ground walking. The
buckled-beam SEA may also use the transverse flux motor 606 as a
dynamic brake (e.g., a clutch) in early stance flexion to reduce
battery energy consumption and acoustic emissions.
[0055] The prosthesis 600 may have multiple MTU's (e.g., the ankle
device 602 and the knee device 604), each with embedded SEA power
electronics and servo controller(s) integrated onto the motor,
networked to a State Controller/IMU assembly adapted to control and
coordinate the individual MTU's in accordance with the gait cycle
and terrain modality. Connectivity among the remote components may
be accomplished through the use of wireless technologies, such as
Bluetooth.RTM. and smart WiFi (sWiFi) control modules. A
BlueTooth.RTM. port can enable remote programming by Android-based
cell phone/PDA to facilitate clinical tuning and evaluation. The
sWiFi may be used to promote 500 Hz logging of up to 50 floating
point-state variables, enabling a modification of up to 200 (and
greater) control and signal processing parameters while the
biomimetic prosthesis is in use.
[0056] The MTUs 602, 604 may be designed around biomechanical
torque-displacement trajectories to ensure the MTUs 602, 604 can
provide a biomimetic response. Through the design process, certain
targets may be set, such as battery and motor power electronic
requirements. In one instance, the components may be designed to
work across walking speeds for a typical wearer weighing 250 lbs.
Other design input parameters include motor diameter and winding
characteristics; series elasticity stiffness; ball-screw
transmission ratio; battery and power supply topology and switching
frequency. The result is a multi-variable design output space that
includes walking range; acoustic energy (torque ripple harmonics)
in the 1.5 kHz-3 kHz audio band; design life and prosthesis
weight.
[0057] A robust design is suitable for a broad range of wearer
weights, including from approximately 185 lbs. to approximately 250
lbs., as may be common amongst soldiers expected to use the device
600. However, the device 600 is usable by wearers outside of this
range. The battery may be capable of powering at least 2000 steps
for a 185 lb. wearer. The device 600 may be designed to output a
peak torque from the muscle tendon units 602, 604 of at least
approximately 150 N-m, and may output a peak power of at least
approximately 800 watts. The ankle 602 may have a range of rotation
of at least approximately 25.degree. and the knee 604 may have a
range of rotation of at least 130.degree..
[0058] The devices 200, 400, 600 described above may rely on
processing of the IMU state to detect gait cycle state transitions
and to determine intra-step estimates of walking speed, allowing
for intrinsic control of reflex strength in accordance with walking
speed. Further, intrinsic measures of torque can detect the
difference between heel-strike first and toe-strike first gait
cycles, sufficient to nuance the control for stair ascent
(toe-strike first) in relation to normal walking on sloping ground
(heel-strike first). Separate or complementary controls may be used
to control biomimetic leg systems (ankle-foot and/or knee) for
stair ascent/descent and steep sloping terrain modalities. In full
leg systems, the knee prosthesis may be controlled to define the
trajectory of the ankle-foot in multi-terrain environments.
IMU-based kinematic reconstruction algorithms may use the IMU to
determine the six degree-of-freedom (6 DOF) inertial state; the
homogeneous transformation that defines the orientation and origin
of a coordinate system with respect to ground of the lower
extremity link that connects the ankle and knee joint.
Self-calibrating algorithms, which may be considered a
"zero-velocity update" (ZVUP) that occurs at a "zero moment point"
(ZMP), can compute the inertial state within a gait cycle without a
priori information averaged from prior steps. By combining the
inertial state with ankle and knee joint encoder feedback and
knowledge of wearer shoe size and the kinematic transforms between
ankle, knee and hip, the trajectory of the toe, heel, ankle, knee
and hip can be computed in real-time.
[0059] FIG. 7A illustrates offline-kinematic reconstructions of
ankle and knee trajectory using the IMU and intrinsic sensing from
the ankle device, superimposed on stair and sloping ground terrain
modalities. In this manner, high-quality inertial state information
can be used to discriminate between different terrain modalities in
real-time. As shown in FIG. 7B, irregular knee paths may be used to
anticipate foot-strike in a wide range of terrain modalities. As
shown, the folding back of knee path is a precursor to foot strike.
Knowing when foot-strike is reached, ankle-foot trajectory can be
used to sense the stair parameters or terrain slope and the
"momentum" of the impending strike, from which appropriate
pre-strike joint mechanical impedance, equilibrium and reflex
characteristics can be nuanced.
[0060] Trial data from the walking tests may be treated as known
"test vector" inputs to the kinematic reconstruction and pattern
recognition algorithms. Terrain modality discrimination to
disambiguate between level ground, sloped and stair terrain
modalities, both in the swing phase and the stance phase, may also
be tested and validated statistically. Other reference
biomechanical inputs in addition to those described above may be
used to design ankle and knee reflex response characteristics.
Preferred muscle models may be identified using goodness of fit and
efficiency as a function of terrain slope and walking speed (e.g.,
for slopes of +10, +5, 0, -5, and -10 degrees for walking speeds of
0.75, 1.0, 1.25, 1.5 and 1.75 m/s). Similar techniques may be used
in defining the neuromuscular model as a function of polarity
(stair ascent vs. descent), stair characteristics, and stair
ascent/descent rates.
[0061] A functional test stand may be used as a bench platform for
life-testing actuators and for engineering/manufacturing
calibration of SEA torque-displacement models. The test stand may
employ a backdrive mechanism mounted on a precision, 6 DOF,
force/torque sensor to apply torque or position input/stimulus to
the prosthesis joint under test. While the ankle test configuration
is described for reference purposes, similar configurations may be
used for other components. One configuration may be used to verify
the MTU and neuromuscular control response characteristics,
including, but not limited to, joint (or motor) torque,
displacement/velocity, work impedance, power, design-life and
battery energy per cycle. In the tests of biomechanical capability,
the backdrive mechanism will "play back" a joint displacement
trajectory, .theta.(t), that emulates the trajectory in an intact
ankle. The MTU under test may deliver a joint torque, .GAMMA.(t),
or motor torque, .tau.(t), as appropriate in synchrony with the
displacement trajectory. This way, a repetitive test can be applied
corresponding to theoretical or actual gait cycles to validate
MTU/Neuromuscular performance. Some testing embodiments may include
an in-line, robotic calibration and verification of the IMU using a
robotic work cell. For example, a Staubli (Pfaffikon, Switzerland)
RX90 robot may be used to supply a calibrated motion input from
which the calibration and validation can be accomplished. The
kinematic reconstruction verification may use this method to verify
trajectory computations, including ZVUP.
[0062] Certain kinematic reconstruction validation may be
accomplished by comparing results, off-line, in relation to the
joint trajectories computed using stereo photogrammetric methods.
Raw IMU data may be recorded simultaneously with the stereo
photogrammetric acquisition during biomechanical testing. On an
off-line basis, the kinematic reconstruction results may be
validated with the stereo photogrammetric (gold standard) used as
the reference. Terrain discrimination algorithms to be deployed may
also be tested against this reference data.
[0063] The MTU joint torque and displacement may be calibrated
individually on the functional test stand, using an adaptor built
for holding the prosthesis to permit this isolated testing and
calibration. Once calibrated, the torque-displacement predictions
may be compared to the backdrive stimulus reference using special
stimulus trajectories designed for the purpose, thereby
accomplishing sensing, calibration and algorithmic verification of
the MTU torque/displacement models.
[0064] In a manner consistent with the kinematic reconstruction
validation above, MTU torque and displacement may be recorded
during clinical evaluation. The biomechanical computations arising
from the joint torque-displacement predictions from the stereo
photogrammetric and force plate data can be used as a reference
with which to validate the MTU torque-displacement predictions on
an off-line basis. Once validated, in situ measurements of
torque-displacement and mechanical work can be used to validate the
tuning of the neuromuscular-derived MTU control parameter vectors
and to compare, in the form of a dashboard, the measurements to
predefined reference measures of joint displacement, work and power
across walking speeds derived from an inverse dynamics calculation
on human kinetic and kinematic walking data. Once validated, in
situ measures of electro-mechanical efficiency may be collected in
real-time, gait-cycle-by-gait-cycle.
[0065] Sound level comparisons between the radial-flux-based and
the transverse-flux-based prostheses can be made in an area with an
ambient noise level of <40 dB. A microphone (MIC) can be placed
in a number of locations relative to the test object including:
[0066] Close Proximity: MIC attached to ankle via mounting bracket
such that separation distance <6''.
[0067] Prosthesis Wearer Perception: MIC attached to user's
head/upper-torso to accurately capture perception of user.
[0068] Observer Perception: MIC placed in fixed location at a 1 m
distance from the wearer to accurately capture perception of
observer. This step may be used to account for system specific
sound projection characteristics and the frequency dependent nature
of sound propagation.
[0069] Preprocessing can include application of an
A-frequency-weighting filter to the raw sound level measurements
(per IEC 61672). The acoustic comparison may be made within the
1.5-3 kHz frequency range according to the mean/maximum sound
pressure level for each step event and an equivalent continuous
sound Level.
[0070] The biomimetic transfemoral system 600 may also be tested in
its ability to reduce the metabolic economy to walk compared to an
intact limb population. The metabolic rate may be calculated from
measures of oxygen consumption and carbon dioxide production while
amputees and age, weight and height-matched non-amputees walk on
level ground and sloped (.+-.10 degrees) surfaces. Measures of
oxygen uptake and carbon dioxide production may be obtained using a
portable, lightweight, breath-by-breath telemetric system (e.g.,
the Cosmed K4b2 (Rome; Italy)). Study participants walk around an
indoor track next to an electric vehicle programmed to move at
their self-selected speed when using the biomimetic prosthesis,
collecting data for a minimum of 8 minutes. The final three-minute
section of the recordings may be selected after steady-state
conditions have been reached. The metabolic and walking speed data
confirm when a transfemoral prosthesis is capable of biological
ankle and knee mechanics to normalize metabolic economy and
self-selected walking speed during level-ground and sloped (.+-.10
degrees) ground surface ambulation.
[0071] To assist in measuring human and prosthetic knee and ankle
mechanics, reflective markers may be positioned on the body and
data may be recorded using a stereo-photogrammetric system. At each
gait speed, researchers measure whole body kinematics and kinetics
from the right and left sides of the body, e.g., during ten trials
per subject and condition. Reflective markers may be applied to the
skin, including on the following bony landmarks: bilateral anterior
superior iliac spines, posterior superior iliac spines, lateral
femoral condyles, lateral malleoli, forefeet and heels. Additional
markers may be rigidly attached to wands over the mid-femur and
mid-shaft of the tibia. For transfemoral amputees, the tibia,
lateral malleoli, forefeet and heel markers may be placed on the
prosthesis at appropriate locations matching the intact limb. 3D
pelvic and bilateral lower extremity joint kinematics may be
collected by means of an eight-camera motion analysis system,
sampling at 120 frames per second. 3D whole body kinetics may be
collected by means of force plates (AMTI) sampling at 1080 Hz. The
following anthropometric measures may be collected, along with
motion analysis and force platform measures, to calculate joint
kinematics and kinetics: body weight, height, leg length (measured
from medial malleolus to anterior superior iliac spine), knee
width, and ankle width.
[0072] Joint torque and power calculations may be based on the mass
and inertial characteristics of each lower-extremity segment, the
derived linear and angular velocities and accelerations of each
lower extremity segment, and ground reaction force and joint center
position estimates. For each trial, bilateral hip, knee, and ankle
joint kinematic and kinetic data during walking may be derived.
Position data recorded using the motion capture system may be
computed in three planes (sagittal, coronal, and transverse) using
standard software provided with the system. Ground reaction forces
may be measured synchronously with the kinematic data using two
embedded force platforms (AMTI). Joint kinetics in each plane may
be calculated using a full-inverse dynamics model. For the
transfemoral ambulation trials, the human model can be modified to
accurately represent the mass distribution of the prosthesis. To
evaluate the accuracy of the full-inverse dynamics model, the
intrinsic sensory measurements of prosthetic knee and ankle torque
and power will be compared to the torque and power calculated from
the inverse dynamics model. At each gait speed and test
participant, average curves for each subject's knee and ankle
position, velocity, torque and power may be obtained from the
collected trials. Human knee and ankle mechanics from the
individuals with intact limbs may then be compared to prosthetic
knee and ankle mechanics at equivalent gait speeds.
[0073] The use of neuromuscular-inspired MTU control software may
emulate the biological response of an intact limb, thereby
normalizing metabolic walking economy and biomechanical response,
as depicted in FIG. 8. The implicit control architecture may be
complementary or separate from the control architecture previously
described, including the use of non-linear positive force feedback
to achieve a reflex behavior. To create the more complex reflex
dynamics in a multi-axis leg prosthesis, the control structure may
be more generalized and aligned to real muscle activation models. A
metasensor that measures the ground reaction force and ZMP may be
created by combining the force-torque and the IMU-based kinematic
reconstruction.
[0074] Intrinsic inertial sensing, signal processing and pattern
recognition may be used to achieve intra-step terrain
adaptation--thereby enabling seamless terrain modality transitions.
The kinematic reconstruction described above may be applied to the
multi-dimensional leg system 600. Pattern recognition of both ankle
and knee trajectories may be used to discriminate between steep
slope and stair ambulation.
[0075] Closed-loop torque control algorithms may be designed using
the calibrated, SEA, torque-displacement models as the primary
feedback. The first SEA resonant frequency may fall in the range of
45-50 Hz, enabling a feed forward-enabled closed loop bandwidth of
25 Hz--sufficient to make the MTU response substantially invariant
with transmission friction variation. Other feedback systems are
described in detail below. Motor torque cogging compensation
firmware can also be designed. Various neuromuscular models may be
embedded in the MTU to test torque-reflex and impedance response
characteristics across walking speeds (e.g., at 0.75, 1.0, 1.25,
15, and 1.75 m/s).
[0076] Known prosthetic and/or orthotic devices typically employ a
current-feedback system to control the operation of the motor in an
SEA. Generally, in such a system voltage applied to the motor is
manipulated so as to produce a desired results, e.g., to deliver
the required motor torque. In order to guide the voltage
manipulation the known systems only use the motor current in a
feedback control system. FIG. 9, however, depicts other signals in
addition to motor voltage and current that may also be associated
with a feedback system for the device. These signals include motor
torque, joint torque, and motor speed. In various embodiments
described below one or more of these other signals are used instead
of or in addition to motor current to control the motor voltage so
to produce a desired torque characteristic, e.g., a joint torque,
impedance, and/or equilibrium. Various combinations of the signals
may or may not be available for measurement over a period of time.
High fidelity joint torque control may be difficult in SEAs,
particularly those where the series-elastic element is used as a
catapult to amplify motor power at the joint. In these SEAs, the
spring stiffness is relatively low (e.g., 350 Nm/rad in an ankle
joint). In such a case, the first SEA resonance, which arises from
the reflected motor inertia reacting with the series-spring, can be
as low as 8 Hz. In conventional joint torque controllers, the
bandwidth may be limited to a frequency less than this first
resonance, resulting in an often underdamped response that cannot
emulate biological muscle-tendon units.
[0077] FIG. 10 depicts one such system that controls motor current
as a surrogate for joint torque, creating a command current signal,
and applying feedback compensation that drives motor voltage so as
to make the motor torque track the command. In this open-loop joint
torque control, motor cogging, motor ripple, SEA transmission
friction, backlash, and other non-linear dynamics will influence
the degree to which the joint torque can track the commanded value.
An outer joint torque compensator may be applied to address this
issue. One known system uses joint torque in addition to motor
current to compute motor voltage. However, the manipulation of the
command current signal in such a system often makes the control
system very hard to control or even unstable as shown in the
sequel, due to resonance characteristic of the associated physical
device, allowing only a limited joint torque bandwidth for
controller operation in this configuration. These problems often
exist when the inner current loop is closed and joint torque error
due to the SEA resonance dynamics is not eliminated by the addition
of an outer joint torque loop. As described with reference to FIG.
12, the instability and/or bandwidth control problems may be
overcome by measuring motor speed and using this variable in an EMF
feedforward compensator.
[0078] FIG. 11 depicts a system where joint torque is measured and
the corresponding signal used instead of the measured current to
compute a voltage that produces desired results. In addition to
fewer limitations on the bandwidth of the compensator, this
configuration more easily enables the joint torque to follow a
joint torque command signal that produces desired results, in
contrast with the current command signal described above. The joint
torque is the most naturally commanded signal, while a current
signal tends only to be related to usual performance objective
through complicated interactions with other variables.
[0079] FIGS. 12 and 13 illustrate configurations where voltage is
computed using two components: a component that is a scaled
measurement of a measurement of the motor speed and a feedback
component. Feeding back a scaled measurement of the motor speed is
called EMF feedforward, as the signal can provide an estimate of
the electromotive force produced within the motor while canceling
the effect of this force. Measuring motor speed is one way of
producing this estimate of the EMF, though other methods for
producing the estimate that are known to those of skill in the art
are considered within the scope of this disclosure. These other
methods could include the use of Kalman-filter based observer to
estimate the EMF from motor winding voltage and phase current.
[0080] When using the EMF estimate, the resulting system may become
easier to control, such as by enabling the use of other signals in
conjunction with the EMF estimate to produce desired results. FIG.
12 depicts a current feedback loop in conjunction with the EMF
feedforward. The EMF feedforward allows an improved performance in
the current feedback loop. FIG. 13 depicts a joint torque feedback
loop as in in conjunction with the EMF feedforward. The use of
other feedback in conjunction with the EMF feedforward, or using
the EMF feedforward alone, is also contemplated.
[0081] A representative dynamic model of a typical SEA that
illustrates the challenges inherent in designing a controller that
achieves a torque loop bandwidth not limited by the first resonance
is described below. To model a foot and spring, a motor may be
attached to an imaginary member which represents the foot position
if the spring were in neutral. Thus the angle .beta. associated
with this member is the motor angle divided by the gearing
constant. k.sub.g, with k.sub.g in one embodiment being
approximately 220.
[0082] Variable and coefficient list:
[0083] .theta. Angle of the foot (rad)
[0084] .beta. Angle of the imaginary member associated with neutral
spring (rad)
[0085] k.sub.g{dot over (.beta.)} Angle of the motor shaft
(rad)
[0086] k.sub.g Gearing multiplier (approx. 220)
[0087] L.sub.f Length of foot from ankle (0.28 meters)
[0088] J.sub.f Moment of inertia of foot
[0089] In an embodiment, the foot may be assumed to weigh 1.5 lbs,
which acts as a point mass halfway down the foot. A force of 1.5
lbs corresponds to 6.67 Newtons or kg*m/(sec).sup.2. Dividing by
acceleration of gravity 9.8 m/(sec).sup.2 gives a mass of 0.68 kg,
resulting in a value of J.sub.f of 0.68*(0.14).sup.2=0:013
kg*m.sup.2 (0.14 is L.sub.f/2).
[0090] J.sub.M Moment of inertia of the motor shaft
[0091] K.sub.S The spring constant.
[0092] When .beta.<.theta. the spring is stretched and pulling
with K.sub.S=350 N*m/rad, and when .beta.<.theta. the spring is
compressed and pushing with K.sub.S=1200 N*m/rad.
[0093] .tau..sub.E External torque on the foot. If a person of
weight W steps onto the toe of the foot, .tau..sub.E=-W*L.sub.f.
The negative sign indicates the torques work to decrease .theta.. A
person weighing 225 weighs 1000 Newtons. With a foot measuring 0.28
m, the torque applied would be 280 N*m.
[0094] B.sub.f The damping in the ankle joint. Due to friction, it
creates a force opposing the angular velocity of the foot.
[0095] B.sub.M The damping in the motor due to friction.
[0096] The basic kinematic equation .tau.=j*{umlaut over (.theta.)}
applied to the foot around the ankle joint gives:
.tau..sub.E+K.sub.S(.beta.-.theta.)-B.sub.f{dot over
(.theta.)}=J.sub.f{umlaut over (.theta.)} (1)
[0097] The basic kinematic equation .tau.=J*{umlaut over (.theta.)}
is now applied to the motor shaft, representing the motor angle as
k.sub.g*.beta..
.tau. M - K S k g ( .beta. - .theta. ) - B M k g .beta. . = I M k g
.beta. ( 2 ) ##EQU00001##
[0098] The system may be changed to average-angle and difference.
New variables may be defined as follows:
[0099] .phi..sub.A=0.5(.beta.+.theta.); the average position
between the foot and the imaginary member
[0100] .phi..sub.D=0.5(.beta.-.theta.) the difference position
between the foot and the imaginary member
J=J.sub.f+(k.sub.g).sup.2J.sub.M;
B=B.sub.f.revreaction.(k.sub.g).sup.2B.sub.M;
[0101] The two equations become:
.tau..sub.E+k.sub.g.tau..sub.M-B{dot over
(.phi.)}.sub.A+(B.sub.f{dot over (-)}B.sub.Mk.sub.g.sup.2){dot over
(.phi.)}.sub.D=J{umlaut over
(.phi.)}.sub.A+(J.sub.f-(k.sub.g).sup.2J.sub.D){umlaut over
(.phi.)}.sub.D (3)
.tau..sub.E+k.sub.g.tau..sub.M-2K.sub.S.phi..sub.D-B{dot over
(.phi.)}.sub.D+(B.sub.f{dot over (-)}B.sub.Mk.sub.g.sup.2){dot over
(.phi.)}.sub.A=J{umlaut over
(.phi.)}.sub.D+(J.sub.f-(k.sub.g).sup.2J.sub.D){umlaut over
(.phi.)}.sub.A (4)
[0102] When the effects of the differential damping terms are
considered to be negligible, the general structural elements can be
seen. The dynamics of the mechanical system may be modeled as:
.tau..sub.E+k.sub.g.tau..sub.M-B{dot over (.phi.)}.sub.A=J{umlaut
over (.phi.)}.sub.A (5)
-.tau..sub.E+k.sub.g.tau..sub.M-2K.sub.S.phi..sub.D-B{dot over
(.phi.)}.sub.D=J{umlaut over (.phi.)}.sub.D (6)
The equations for the motor dynamics are:
L dt dx + Rl = V - K s k g .beta. . ( 7 ) ##EQU00002##
.tau..sub.M=T.sub.Mi (8)
[0103] Taking transforms:
Lsi+Ri=V-K.sub.sk.sub.gs.beta. (9)
.tau..sub.M=T.sub.Mi (10)
[0104] The back emf of the motor is represented by the term
K.sub.sk.sub.gsB
.tau..sub.E+k.sub.g.tau..sub.M-Bs.phi..sub.A=Js.sup.2.phi..sub.A
(11)
-.tau..sub.E+k.sub.g.tau..sub.M-2K.sub.S.phi..sub.D-Bs.phi..sub.D=Js.sup-
.2.phi..sub.D (12)
[0105] The SEA drives the joint through a spring that serves as a
catapult to amplify motor power at the joint. This form of bionic
muscle-tendon unit may serve as the basis for emulation of the
biomechanical response of an intact limb.
[0106] FIG. 14 is a block diagram of the system, as closely
approximated by the equations (9)-(12). FIG. 15 depicts a system
that achieves a similar result, i.e., the required joint torque
characteristic, but the inner current-feedback loop is
eliminated.
[0107] The open-loop joint torque system depicted in FIG. 14 relies
on use of the motor current to compute a voltage to be applied to
the motor windings to control joint torque. The controlled motor
torque must in itself account for the effects of the complicated
transfer function between the motor torque and the joint torque.
Also, the complex loop involving the back emf proportional to {dot
over (.beta.)} is not treated directly but its effect are assumed
to be negated by the disturbance rejection properties of
feedback.
[0108] However, there are significant limitations of using current
feedback in this setting. From the partially reduced system of FIG.
15, it can be seen that the poorly damped spring dynamics appear as
poles in the return transfer function in the naturally occurring
back emf feedback within the model. This means that, when the
current loop is closed, as shown in FIG. 14, the effects of the
spring element appear as a pair of zeros near the imaginary axis in
the transfer function between the voltage and the motor current.
When a current-loop controller with these dynamic limitations is
embedded in a closed-loop torque control, as shown in FIG. 18, the
closed loop torque response either exhibits a bandwidth far below
the resonance or exhibits an underdamped response, neither of which
is useful in emulation of biological muscle tendon unit
response.
[0109] Using a direct measurement of joint torque, a wider
bandwidth and more robust closed-loop control system is possible.
Here, proportional-integral compensation applied to the joint
torque tracking error will generate a stable response that extends
beyond the first resonance. In the system depicted in FIG. 15, the
resonant spring dynamics appear in the forward path of the joint
torque controller rather than the feedback path. Therefore, these
dynamics do not appear as troublesome zeros in a loop that measures
the joint torque and manipulates that signal to produce a
compensating voltage input. In general, poorly damped zeros in a
feedback control design limit the bandwidth and effectiveness of
the control loop, as described above. Indeed, since the poorly
damped poles are substantially eliminated) i.e., legitimately
cancelled by the zeros arising from the same dynamics in the
naturally occurring back emf loop), the transfer function between
the motor's voltage and the joint torque is quite benign (without
oscillatory dynamics) and amenable to a PI or PID controller.
[0110] FIG. 16 depicts a further improved system with the inclusion
of emf feedforward in a closed-loop torque controller with embedded
current loop. Here the feedforward substantially eliminates the
underdamped dynamics, enabling simple compensation in the outer
torque loop. A measurement or derived estimate of the motor speed
{dot over (.beta.)} is used in a controller to cancel the effect of
the back emf. As shown in FIG. 16, this greatly simplifies the
dynamics of the system. A feedback controller measuring and feeding
back either the current i or the joint torque can be used in
addition to the emf feedforward to reduce sensitivity to any
remnant error in the estimation of the motor speed and other
modeling errors. A simple PI loop using the motor current can
control the motor torque while a PI or PID loop using the joint
torque may produce excellent control of the joint torque.
[0111] FIG. 17 is a block diagram of a model using a measurement of
joint torque through a single compensator to compute a voltage
signal that produces a desired result. The blocks of FIG. 17 have
been filled with typical transfer functions for a typical joint
problem; one block giving a transfer function from motor voltage v
to motor current i and a second block giving, a transfer function
from motor current i to joint torque .tau..sub.j. For the single
loop, only the product of the two blocks may be considered so that
the poorly damped poll pair that originates with the spring is
exactly cancelled by the zero pair in the first block that
originates in the same physical element. The resulting dynamics
allow for the design and implementation of a compensator C(s) to
follow the command signal .tau..sub.JC. C(s) may be a PT
controller.
[0112] By contrast, a model of the known motor controllers,
described with reference to FIG. 18, includes two compensators
arranged in an inner loop using the motor current i and an outer
loop using the joint torque .tau..sub.J. The inner loop may have to
contend with a poorly damped zero pair in the first block, though
two systems may be expected to perform similarly by setting:
C 2 ( s ) = C ( s ) - s 2 + 100 s + 5000 8500 C 1 ( s ) ( 13 )
##EQU00003##
[0113] Thus, the C2 compensator needed to make the systems perform
the same may be more complex than systems requiring only a single
compensator, in addition to the complexity of making multiple
measurements and implementing multiple compensators.
[0114] It should be understood that for those skilled in the art,
the control architecture described here can be extended to bionic
ankles that employ physical or SEA-applied virtual, unidirectional
parallel elastic elements where the torque-displacement
characteristics of these can be calibrated before use. Further,
while the control architecture has been applied to leg prostheses,
the principles can be readily extended to orthotic, exoskeletal or
humanoid applications in lower-extremity augmentation of ankles,
knees and hips.
[0115] The terms and expressions employed herein are used as terms
and expressions of description and not of limitation, and there is
no intention, in the use of such terms and expressions, of
excluding any equivalents of the features shown and described or
portions thereof. In addition, having described certain embodiments
of the invention, it will be apparent to those of ordinary skill in
the art that other embodiments incorporating the concepts disclosed
herein may be used without departing from the spirit and scope of
the invention. The compositions, components, and functions can be
combined in various combinations and permutations, to achieve a
desired result. For example, all materials for components
(including materials not necessarily previously described) that are
suitable for the application are considered within the scope of the
invention. Accordingly, the described embodiments are to be
considered in all respects as only illustrative and not
restrictive. Furthermore, the configurations described herein are
intended as illustrative and in no way limiting. Similarly,
although physical explanations have been provided for explanatory
purposes, there is no intent to be bound by any particular theory
or mechanism, or to limit the claims in accordance therewith.
* * * * *