U.S. patent application number 11/642993 was filed with the patent office on 2007-07-12 for artificial joints using agonist-antagonist actuators.
This patent application is currently assigned to Massachusetts Institute of Technology. Invention is credited to Ken Endo, Hugh M. Herr, Lee Harris Magnusson.
Application Number | 20070162152 11/642993 |
Document ID | / |
Family ID | 46326877 |
Filed Date | 2007-07-12 |
United States Patent
Application |
20070162152 |
Kind Code |
A1 |
Herr; Hugh M. ; et
al. |
July 12, 2007 |
Artificial joints using agonist-antagonist actuators
Abstract
Artificial limbs and joints which behave like a biological limbs
and joints employ a synthetic actuator which consume negligible
power when exerting zero force, consume negligible power when
outputting force at constant length (isometric) and while
performing dissipative, nonconservative work, are capable of
independently engaging flexion and extension tendon-like, series
springs, are capable of independently varying joint position and
stiffness, and exploit series elasticity for mechanical power
amplification.
Inventors: |
Herr; Hugh M.; (Somerville,
MA) ; Magnusson; Lee Harris; (Cambridge, MA) ;
Endo; Ken; (Cambridge, MA) |
Correspondence
Address: |
CHARLES G. CALL
68 HORSE POND ROAD
WEST YARMOUTH
MA
02673-2516
US
|
Assignee: |
Massachusetts Institute of
Technology
Cambridge
MA
|
Family ID: |
46326877 |
Appl. No.: |
11/642993 |
Filed: |
December 19, 2006 |
Related U.S. Patent Documents
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Application
Number |
Filing Date |
Patent Number |
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11395448 |
Mar 31, 2006 |
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11642993 |
Dec 19, 2006 |
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11499853 |
Aug 4, 2006 |
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11642993 |
Dec 19, 2006 |
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11495140 |
Jul 29, 2006 |
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11642993 |
Dec 19, 2006 |
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11395448 |
Mar 31, 2006 |
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11642993 |
Dec 19, 2006 |
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11600291 |
Nov 15, 2006 |
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11642993 |
Dec 19, 2006 |
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11395448 |
Mar 31, 2006 |
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11642993 |
Dec 19, 2006 |
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60751680 |
Dec 19, 2005 |
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60666876 |
Mar 31, 2005 |
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60704517 |
Aug 1, 2005 |
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60705651 |
Aug 4, 2005 |
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60704517 |
Aug 1, 2005 |
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60736929 |
Nov 15, 2005 |
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Current U.S.
Class: |
623/24 ;
623/27 |
Current CPC
Class: |
A61F 2002/704 20130101;
A61F 2002/5004 20130101; A61F 2/605 20130101; A61F 2/72 20130101;
A61F 2002/7635 20130101; B62D 57/032 20130101; A61F 2002/763
20130101; A61F 2002/5003 20130101; A61F 2002/5075 20130101; A61F
2/70 20130101; A61F 2002/5033 20130101; A61F 2/6607 20130101; A61F
2002/764 20130101; A61F 2002/6657 20130101; A61F 2002/503 20130101;
A61F 2/64 20130101; B25J 19/0008 20130101; A61F 2002/6818 20130101;
A61F 2002/741 20130101; A61F 2002/7625 20130101; A61F 2/60
20130101; A61F 2002/701 20130101; A61F 2002/7645 20130101 |
Class at
Publication: |
623/024 ;
623/027 |
International
Class: |
A61F 2/48 20060101
A61F002/48; A61F 2/74 20060101 A61F002/74 |
Claims
1. An artificial limb comprising, in combination, first and second
members connected for movement relative to one another at a joint,
a flexion actuator connected between said first and second members,
said extendable flexion actuator comprising the series combination
of a first motor and a first elastic element drawing said first and
second members together to reduce the angle between said first and
second members at said joint, a extension actuator connected
between said first and second members, said extension actuator
comprising the series combination of a second motor and a second
elastic for urging said first and second members apart to increase
the angle between said first and second members at said joint, and
a controller for independently energizing said first motor and said
second motor at different times to control the movement of said
artificial limb.
Description
CROSS REFERENCE TO RELATED APPLICATIONS
[0001] This application is a non-provisional of U.S. patent
application Ser. No. 60/751,680 filed on Dec. 19, 2005.
[0002] This application is a continuation in part of U.S. patent
application Ser. No. 11/395,448 entitled "Artificial human limbs
and joints employing actuators, springs, and Variable-Damper
Elements" filed on Mar. 31, 2006 by Hugh M. Herr, Daniel Joseph
Paluska, and Peter Dilworth. application Ser. No. 11/395,448 claims
the benefit of the filing date of U.S. Provisional patent
application Ser. No. 60/666,876 filed on Mar. 31, 2005 and the
benefit of the filing date of U.S. Provisional patent application
Ser. No. 60/704,517 filed on Aug. 1, 2005.
[0003] This application is also a continuation in part of U.S.
patent application Ser. No. 11/499,853 entitled "Biomimetic motion
and balance controllers for use in prosthetics, orthotics and
robotics" filed on Aug. 4, 2006 by Hugh M. Herr, Andreas G.
Hofmann, and Marko B. Popovic. application Ser. No. 11/499,853
claims the benefit of the filing date of, U.S. Provisional patent
application Ser. No. 60/705,651 filed on Aug. 4, 2005.
[0004] This application is also a continuation in part of U.S.
patent application Ser. No. 11/495,140 entitled "An Artificial
Ankle-Foot System with Spring, Variable-Damping, and Series-Elastic
Actuator Components" filed on Jul. 29, 2006 by Hugh M. Herr, Samuel
K. Au, Peter Dilworth, and Daniel Joseph Paluska. application Ser.
No. 11/495,140 claims the benefit of the filing date of U.S.
Provisional patent application Ser. No. 60/704,517 filed on Aug. 1,
2005 and was also a continuation in part of the above-noted
application Ser. No. 11/395,448.
[0005] This application is also a continuation in part of U.S.
patent application Ser. No. 11/600,291 entitled "Exoskeletons for
running and walking" filed on Nov. 15, 2006 by Hugh M. Herr, Conor
Walsh, Daniel Joseph Paluska, Andrew Valiente, Kenneth Pasch, and
William Grand. application Ser. No. 11/600,291 claims the benefit
of the filing date of U.S. Provisional patent application Ser. No.
60/736,929 filed on Nov. 15, 2005 and is a continuation in part of
the above noted applications Ser. Nos. 11/395,448, 11/499,853 and
11/495,140.
[0006] The present application claims the benefit of the filing
date of each of the foregoing patent applications and incorporates
the disclosure of each of the foregoing applications herein by
reference.
FIELD OF THE INVENTION
[0007] This invention relates to artificial joints and limbs for
use in prosthetic, orthotic or robotic devices.
BACKGROUND OF THE INVENTION
[0008] Biomimetic Hybrid Actuators employed in
biologically-inspired musculoskeletal architectures as described in
the above noted U.S. patent application Ser. No. 11/395,448 employ
an electric motor for supplying positive energy to and storing
negative energy from an artificial joint or limb, as well as
elastic elements such as springs, and controllable variable damper
components, for passively storing and releasing energy and
providing adaptive stiffness to accommodate level ground walking as
well as movement on stairs and surfaces having different
slopes.
[0009] The above noted application Ser. No. 11/495,140 describes an
artificial foot and ankle joint consisting of a curved leaf spring
foot member that defines a heel extremity and a toe extremity, and
a flexible elastic ankle member that connects said foot member for
rotation at the ankle joint. An actuator motor applies torque to
the ankle joint to orient the foot when it is not in contact with
the support surface and to store energy in a catapult spring that
is released along with the energy stored in the leaf spring to
propel the wearer forward. A ribbon clutch prevents the foot member
from rotating in one direction beyond a predetermined limit
position, and a controllable damper is employed to lock the ankle
joint or to absorb mechanical energy as needed. The controller and
a sensing mechanisms control both the actuator motor and the
controllable damper at different times during the walking cycle for
level walking, stair ascent and stair descent.
[0010] The above noted Application G-25 describes an exoskeleton
worn by a human user consisting of a rigid pelvic harness worn
about the waist of the user and exoskeleton leg structures each of
which extends downwardly alongside one of the human user's legs.
The leg structures include hip, knee and ankle joints connected by
adjustable length thigh and shin members. The hip joint that
attaches the thigh structure to the pelvic harness includes a
passive spring or an active actuator to assist in lifting the
exoskeleton and said human user with respect to the ground surface
upon which the user is walking and to propel the exoskeleton and
human user forward. A controllable damper operatively arresting the
movement of the knee joint at controllable times during the waking
cycle, and spring located at the ankle and foot member stores and
releases energy during walking.
[0011] The additional references listed below identify materials
which are referred to in the description that follows. When cited,
each reference is identified by a single number in brackets; for
example, the first reference below is cited using the notation "{1
}." [0012] 1. Palmer, Michael. Sagittal Plane Characterization of
Normal Human Ankle Function across a Range of Walking Gait Speeds.
Massachusetts Institute of Technology Master's Thesis, 2002. [0013]
2. Gates Deanna H., Characterizing ankle function during stair
ascent, descent, and level walking for ankle prosthesis and
orthosis design. Master thesis, Boston University, 2004. [0014] 3.
Hansen, A., Childress, D. Miff, S. Gard, S. and Mesplay, K., The
human ankle during walking: implication for the design of
biomimetic ankle prosthesis, Journal of Biomechanics, 2004 (In
Press). [0015] 4. Koganezawa, K. and Kato, I., Control aspects of
artificial leg, IFAC Control Aspects of Biomedical Engineering,
1987, pp.71-85. [0016] 5. Herr H, Wilkenfeld A. User-Adaptive
Control of a Magnetorheological Prosthetic Knee. Industrial Robot:
An International Journal 2003; 30: 42-55. [0017] 6. Seymour Ron,
Prosthetics and Orthotics: Lower limb and Spinal, Lippincott
Williams & Wilkins, 2002. [0018] 7. G. A. Pratt and M. M.
Williamson, "Series Elastic Actuators," presented at 1995 IEEE/RSJ
International Conference on Intelligent Robots and Systems,
Pittsburgh, Pa., 1995. [0019] 8. Inman V T, Ralston H J, Todd F.
Human walking. Baltimore: Williams and Wilkins; 1981. [0020] 9.
Hof. A. L. Geelen B. A., and Berg, J. W. Van den, "Calf muscle
moment, work and efficiency in level walking; role of series
elasticity," Journal of Biomechanics, Vol 16, No. 7, pp. 523-537,
1983. [0021] 10. Gregoire, L., and et al, Role of mono- and
bi-articular muscles in explosive movements, International Journal
of Sports Medicine 5, 614-630. [0022] 11. Endo, K., Paluska D.,
Herr, H. A quasi-passive model of human leg function in
level-ground walking. IEEE/RSJ International Conference on
Intelligent Robots and Systems (IROS); Oct. 9-16, 2006; Beijing,
China.
[0023] As noted in references {1 }, {2 }, {3 }, and {4} above, an
artificial limb system that mimics a biological limb ideally needs
to fulfill a diverse set of requirements. The artificial system
must be a reasonable weight and have a natural morphological shape,
but still have an operational time between refueling or battery
recharges of at least one full day. The system must also be capable
of varying its position, stiffness, damping and nonconservative
motive power in a comparable manner to that of a normal, healthy
biological limb. Still further, the system must be adaptive,
changing its characteristics given such environmental disturbances
as walking speed and terrain variation. The current invention
describes a novel actuator and limb architecture capable of
achieving these many requirements.
[0024] From recent biomechanical studies described in references {1
}, {2 } and {3 } above, researchers have determined that biological
joints have a number of features. Among these are: [0025] (a) the
ability to vary stiffness and damping; [0026] (b) the ability to
generate large amounts of positive mechanical work (nonconservative
motive output); and [0027] (c) the ability to produce large amounts
of power and torque when needed.
[0028] An example of the use of more than one control strategy in a
single biological joint is the ankle. See {1 } and {2}. For level
ground ambulation, the ankle behaves as a variable stiffness device
during the early to midstance period, storing and releasing impact
energies. Throughout terminal stance, the ankle acts as a torque
source to power the body forward. In distinction, the ankle varies
damping rather than stiffness during the early stance period of
stair descent. These biomechanical findings suggest that in order
to mimic the actual behavior of a human joint or joints, stiffness,
damping, and nonconservative motive power must be actively
controlled in the context of an efficient, high cycle-life, quiet
and cosmetic biomimetic limb system, be it for a prosthetic or
orthotic device. This is also the case for a biomimetic robot limb
since it will need to satisfy the same mechanical and physical laws
as its biological counterpart, and will benefit from the same
techniques for power and weight savings.
[0029] The current state of the art in prosthetic leg systems
include a knee joint that can vary its damping via
magnetorheological fluid as described in {5}, and a carbon fiber
ankle which has no active control, but that can store energy in a
spring structure for return at a later point in the gait cycle e.g.
the Flex-Foot {4} or the Seattle-Lite {6}. None of these systems
are able to add energy during the stride to help keep the body
moving forward or to reduce impact losses at heel strike. In the
case of legged robotic systems, the use of the Series Elastic
Actuator (SEA) enables robotic joints to control their position and
torque, such that energy may be added to the system as needed. See
{7}. In addition, the SEA can emulate a physical spring or damper
by applying torques based on the position or velocity of the joint.
However, for most applications, the SEA requires a tremendous
amount of electric power for its operation, resulting in a limited
operational life or an overly large power supply. Robotic joint
designs in general use purely active components and often do not
conserve electrical power through the use of variable-stiffness and
variable-damping devices.
SUMMARY OF THE INVENTION
[0030] The following summary provides a simplified introduction to
some aspects of the invention as a prelude to the more detailed
description that is presented later, but is not intended to define
or delineate the scope of the invention.
[0031] In this specification and the claims, the following terms
have the following meanings:
[0032] actuator: see the definition of "motor" below; [0033]
agonist: A contracting element that is resisted or counteracted by
another element, the antagonist; [0034] agonist-antagonist
actuator: a mechanism comprising (at least) two actuators that
operate in opposition to one another: an agonist actuator that,
when energized, draws two elements together and an antagonist
actuator that, when energized, urges the two elements apart; [0035]
antagonist: An expanding element that is resisted or counteracted
by another element, the agonist; [0036] biomimetic: a man-made
structure or mechanism that mimics the properties and behavior of
biological structures or mechanisms, such as joints or limbs;
[0037] dorsiflexion: bending the ankle joint so that the end of the
foot moves upward; [0038] elastic: capable of resuming an original
shape after deformation by stretching or compression; [0039]
extension: A bending movement around a joint in a limb that
increases the angle between the bones of the limb at the joint;
[0040] flexion: A bending movement around a joint in a limb that
decreases the angle between the bones of the limb at the joint;
[0041] motor: an active element that produces or imparts motion by
converting supplied energy into mechanical energy, including
electric, pneumatic or hydraulic motors and actuators; [0042]
plantarflexion: bending the ankle joint so that the end of the foot
moves downward; [0043] spring: an elastic device, such as a metal
coil or leaf structure, which regains its original shape after
being compressed or extended.
[0044] For an artificial joint to behave like a biological joint, a
synthetic actuator must have the following properties:
[0045] 1) The actuator must consume negligible power when exerting
zero force. Near the equilibrium length of muscle (peak of active
tension-length curve), the passive tension is typically zero. Thus
a muscle-actuated joint goes limp when the muscles are not
electrically stimulated.
[0046] 2) The actuator must consume negligible power when
outputting force at constant length (isometric) and while
performing dissipative, nonconservative work. Muscle tissue is very
efficient for isometric and dissipative control modes.
[0047] 3) The actuator must be capable of independently engaging
flexion and extension tendon-like, series springs. Since biological
joints have at least one flexor muscle and at least one extensor
muscle, the time at which a flexor tendon becomes taught or engaged
can be independent of the time at which an extensor tendon becomes
engaged. As an example, with a muscle-actuated joint, the elastic
energy from one tendon can be released as a second tendon is being
elongated.
[0048] 4) The actuator must be capable of independently varying
joint position and stiffness. Through co- contraction between a
muscle flexor and extensor, joint stiffness can be modulated
without changing joint position. Further, joint position can be
varied while keeping joint stiffness constant.
[0049] 5) The actuator must be capable of exploiting series
elasticity for mechanical power amplification, or a "catapult"
control modality. For motion tasks that require high mechanical
power, muscle-tendon units in animals and humans often employ a
catapult control where the muscle belly stretches the series
tendon, and later that stored elastic energy is released to achieve
relatively higher joint powers than would be possible if the muscle
belly were to generate that power directly.
BRIEF DESCRIPTION OF THE DRAWINGS
[0050] In the detailed description which follows, frequent
reference will be made to the attached drawings, in which:
[0051] FIG. 1 depicts the various subdivisions of the stance phase
of walking;
[0052] FIGS. 2A, 2B and 2C show torque vs. angle plots in
level-ground walking for slow speed, normal and fast walking;
[0053] FIG. 3 illustrates human ankle biomechanics for stair
ascent;
[0054] FIG. 4 illustrates human ankle-foot biomechanics for stair
descent;
[0055] FIGS. 5A and 5B illustrate the manner in which knee angle
and knee power respectively vary during the walking cycle for level
ground walking;
[0056] FIGS. 6A and 6B are posterior and side elevational views
respectively of an agonist-antagonist actuator embodying the
invention;
[0057] FIGS. 7A and 7B are posterior and side elevational views of
an agonist-antagonist actuator mechanism implementing an artificial
ankle;
[0058] FIGS. 8A and 8B are posterior and side elevational views of
agonist-antagonist actuator mechanisms implementing an artificial
knee;
[0059] FIGS. 9A and 9B are side elevational and perspective views
of an agonist-antagonist actuator mechanism positioned on both
sides of the joint axis;
[0060] FIGS. 10A and 10B are posterior and side elevational views
of an agonist-antagonist actuator mechanism using motor and spring
combinations;
[0061] FIGS. 11A shows a model of leg prothesis employing
series-elastic clutches at the hip, knee and ankle joints;
[0062] FIGS. 11B and 11C are graphs comparing the behaviors of
biological ankle and knee joints respectively with the modeled
joints of FIG. 11A;
[0063] FIGS. 12A, 12B and 12C are plots of the mechanical power of
each model element is versus percentage gait cycle for ankle, knee
and hip, respectively;
[0064] FIGS. 13A shows the major components of the transtibial
system are shown;
[0065] FIGS. 13B shows the monoarticular ankle mechanism of FIG.
13A in more detail;
[0066] FIGS. 13C and 13D show elevational and schematic views
respectively of the bi-articular ankle knee mechanism of FIG.
13A;
[0067] FIGS. 14A and 14B shows the major components of an
artificial ankle and knee system;
[0068] FIGS. 14C is a schematic diagram of the artificial ankle and
knee system seen in FIGS. 14A and 14B; and
[0069] FIG. 14D shows the knee's variable moment arm (VMA) device
(seen at the top of FIGS. 14A and 14B) in more detail.
DETAILED DESCRIPTION
[0070] In the construction of a biologically realistic limb system
that is high performance, light weight, quiet and power efficient,
a agonist-antagonist actuator design is proposed herein comprising
a plurality of actuators and series elastic structures. Since it is
desirable to minimize the overall weight of the limb design, the
efficiency of the agonist-antagonist actuator design is critical,
especially given the poor energy density of current power supplies,
e.g. lithium-ion battery technology. By understanding human
biomechanics, the lightest, most energy efficient agonist-
antagonist actuator design can be achieved.
[0071] In the next section, the key features of biomechanical
systems are highlighted. A more complete description of
biomechanical systems is found in the patent applications cited in
the foregoing "Cross Reference to Related Applications" whose
disclosures are incorporated herein by reference.
[0072] Joint Biomechanics: The Human Ankle
[0073] Understanding normal walking biomechanics provides the basis
for the design and development of the agonist-antagonist actuator
design. Specifically, the function of human ankle under sagittal
plane rotation is described for different locomotor conditions
including level-ground walking and stair/slope ascent and descent.
In addition, the function of the human knee during level ground
walking is described. From these biomechanical descriptions, the
justifications for key mechanical components and configurations of
the actuator invention are established.
[0074] Level-Ground Walking
[0075] A level-ground walking gait cycle is typically defined as
beginning with the heel strike of one foot seen at 103 in FIG. 1
and ending at the next heel strike of the same foot seen at 113.
See {8}. The main subdivisions of the gait cycle are the stance
phase (.about.60%) and the swing phase (.about.40%) which are
illustrated in FIG. 1. The swing phase represents the portion of
the gait cycle when the foot is off the ground. The stance phase
begins at heel strike when the heel touches the floor and ends at
toe off when the same foot rises from the ground surface.
Additionally, we can further divide the stance phase into three
sub- phases: Controlled Plantar Flexion (CP), Controlled
Dorsiflexion (CD), and Powered Plantar Flexion (PP).
[0076] Detailed descriptions for each phase and the corresponding
ankle functions are described in FIG. 1. CP begins at heel-strike
103 and ends at foot-flat shown at 105. Simply speaking, CP
describes the process by which the heel and forefoot initially make
contact with the ground. In {1} and {3}, researchers showed that CP
ankle joint behavior is consistent with a linear spring being
loaded or stretched where joint torque is proportional to joint
position.
[0077] During the loading process, the spring behavior is, however,
variable; joint stiffness is continuously modulated by the body
from step to step. After the CP period, the CD phase begins. In
FIG. 2, the average torque versus angle curves are shown for 68
healthy, young participants walking on a level surface. As is
shown, during CP (from 103 to 105), the ankle behaves as a linear
spring of variable stiffness during the loading cycle, but the
torque curve does not trace back to point 1, but rather assumes
higher torque values during the early period of CD.
[0078] Ankle torque versus position during the CD period from 105
to 107 can often be described as a nonlinear spring being loaded or
stretched where stiffness increases with increasing ankle position.
It is noted that as walking speed increases, the extent to which
the ankle behaves as a nonlinear spring increases, with the CD
loading phase exhibiting distinct nonlinear behavior during fast
walking (see fast walking, FIG. 2). The main function of the ankle
during CD is to store elastic energy to propel the body upwards and
forwards during the PP phase. See {9} and {3}.
[0079] The PP phase begins at 107 after CD and ends at the instant
of toe-off shown at 109. During PP in moderate to fast walking
speeds, the ankle can be modeled as a catapult in series or in
parallel with the CD spring or springs. Here the catapult component
includes an actuator that does work on a series spring during the
CD phase and/or during the first half of the PP phase. The catapult
energy is then released along with the spring energy stored during
the CD phase to achieve the high plantar flexion power during late
stance. This catapult behavior is necessary because the work
generated during PP is more than the negative work absorbed during
the CP and CD phases for moderate to fast walking speeds as clearly
seen in FIGS. 2A-2C. See {1}, {2}, {3} and {9}.
[0080] FIGS. 2A, 2B and 2C show torque vs. angle plots in
level-ground walking for slow speed walking at 0.9 m/sec (FIG. 2A),
normal walking speed at 1.25 m/sec (FIG. 2B) and fast walking at
1.79 m/sec. Only the stance period of a single foot is shown (heel
strike to toe off). Point 1 on the charts denotes heel strike,
point 2 foot flat, point 3 peak dorsiflexion, and point 4 toe off.
Although during slow walking the loading curve (points 2-3) is
approximately equal to the unloading curve (points 3-4), for higher
walking speeds the torque assumes higher values during the
unloading, PP phase (points 3-4). Hence, for walking speeds greater
than 0.9 m/s (slow speed), the human ankle cannot be modeled as a
series of coupled springs because the positive work performed by
the ankle exceeds the negative work. It is noted that, as walking
speed increases, the degree of nonlinear behavior during CD (points
2-3) increases along with the total amount of positive work
production during PP (points 3-4), consistent with a catapult model
where the soleus muscle belly slowly elongates the series Achilles
tendon spring during CD, increasing the slope of the torque versus
angle curve and the subsequent positive power output of the
ankle.
[0081] Stair Ascent and Descent
[0082] FIG. 3 illustrates human ankle biomechanics for stair
ascent; The first phase of stair ascent is called Controlled
Dorsiflexion 1 (CD 1), which begins with foot strike in a
dorsiflexed position at 201 and continues to dorsiflex until the
heel contacts the step surface at 203. In this phase, the ankle can
be modeled as a linear spring. The second phase is Powered Plantar
flexion 1 (PP 1), which begins at the instant of foot flat (when
the ankle reaches its maximum dorsiflexion) at 203 and ends when
dorsiflexion begins once again at 205. The human ankle behaves as a
torque actuator to provide extra energy to support the body weight.
The third phase is Controlled Dorsiflexion 2(CD 2), in which the
ankle dorsiflexes as seen at 205 until heel-off at 207. For that
phase, the ankle can be modeled as a linear spring. The fourth and
final phase is Powered Plantar flexion 2 (PP 2). Here the foot
pushes off the step as seen at 207, acting as a torque actuator in
parallel with the CD 2 spring to propel the body upwards and
forwards until toe-off occurs at 209 and the swing phase
begins.
[0083] FIG. 4 illustrates the human ankle-foot biomechanics for
stair descent. The stance phase of stair descent is divided into
three sub-phases: Controlled Dorsiflexion 1 (CD1), Controlled
Dorsiflexion 2 (CD2), and Powered Plantar flexion (PP). CD1 begins
at forefoot strike seen at 303 and ends at foot-flat seen at 305.
In this phase, the human ankle can be modeled as a variable damper.
In CD2, from foot flat at 305, the ankle continues to dorsiflex
forward until it reaches a maximum dorsiflexion posture at 307.
Here the ankle acts as a linear spring in series with a
variable-damper designed to effectively control the amount of
energy stored by the linear spring. During PP, beginning at 307,
the ankle plantar flexes until the foot lifts from the step at 309.
In this final phase, the ankle releases stored CD2 energy,
propelling the body upwards and forwards. From toe off at 309 until
the next foot strike at 313, the foot in the swing phase.
[0084] Because the kinematic and kinetic patterns at the ankle
during stair ascent/descent are significantly different from that
of level-ground walking (see {2}), a description of such ankle-foot
biomechanics seems appropriate. For stair ascent, the human
ankle-foot can be effectively modeled using a combination of an
actuator and a variable stiffness mechanism. However, for stair
descent, variable damping needs also to be included for modeling
the ankle-foot complex; the power absorbed by the human ankle is
much greater during stair descent than the power released by 2.3 to
11.2 J/kg. See reference {2}.
[0085] Joint Biomechanics: The Human Knee
[0086] There are five distinct phases to knee operation throughout
a level-ground walking cycle as illustrated in FIG. 5A and 5B. See
reference {8}. FIG. 5A shows how the knee angle varies during the
walking cycle, and FIG. 5B shows how knee power varies. As seen in
FIG. 5A, the stance phase of walking can be divided into three
sub-phases: Stance Flexion, Stance Extension, and Pre-Swing. The
swing phase is divided into two phases: Swing Flexion and Swing
Extension. As seen in FIG. 5B, for level ground walking, the human
knee does more negative work than positive work.
[0087] Beginning at heel strike indicated at 403, the stance knee
begins to flex slightly. This flexion period, called the Stance
Flexion phase, allows for shock absorption upon impact as well as
to keep the body's center of mass at a more constant vertical level
throughout the stance period. During this phase, the knee acts as a
spring, storing energy in preparation for the Stance Extension
phase.
[0088] After maximum flexion is reached in the stance knee at 404,
the joint begins to extend, until maximum extension is reached as
indicated at 406. This knee extension period is called the Stance
Extension phase. Throughout the first .about.60% of Stance
Extension, the knee acts as a spring, releasing the stored energy
from the Stance Flexion phase of gait. This first release of energy
corresponds to power output indicated at 501 in the graph at the
bottom of FIG. 5B. During the last .about.30% of Stance Extension,
the knee absorbs energy in a second spring and then that energy is
released during the next gait phase, or Pre-Swing.
[0089] During late stance or Pre-Swing from 406 to 407, the knee of
the supporting leg begins its rapid flexion period in preparation
for the swing phase. During early Pre-Swing, as the knee begins to
flex in preparation for toe- off, the stored elastic energy from
Stance Extension is released. This second release of energy
corresponds to power output seen at 503 in FIG. 5B.
[0090] As the hip is flexed, and the knee has reached a certain
angle in Pre-Swing, the leg leaves the ground at 407 and the knee
continues to flex. At toe-off 407, the Swing Flexion phase of gait
begins. Throughout this period, knee power is generally negative
where the knee's torque impedes knee rotational velocity. During
terminal Swing Flexion, the knee can be modeled as an extension
spring in series with a variable damper, storing a small amount of
energy in preparation for early Swing Extension.
[0091] After reaching a maximum flexion angle during swing at 408,
the knee begins to extend forward. During the early Swing Extension
period, the spring energy stored during late Swing Flexion is then
released, resulting in power output seen at 505 in FIG. 5B. During
the remainder of Swing Extension, the human knee outputs negative
power (absorbing energy) to decelerate the swinging leg in
preparation for the next stance period. During terminal Swing
Extension, the knee can be modeled as a flexion spring in series
with a variable damper, storing a small amount of energy in
preparation for early stance (at 507). After the knee has reached
full extension, the foot once again is placed on the ground, and
the next walking cycle begins.
[0092] An agonist-antagonist actuator described below implements
these muscle-like actuation properties. The actuator comprises a
plurality of springs, mechanical transmissions, and active elements
where each spring is in series with an active element via a
transmission, and each spring-transmission-active element
combination are in parallel and capable of opposing one another in
an agonist-antagonist manner. The components of the
agonist-antagonist actuator are listed in Table 1 with their
functional purposes outlined.
[0093] The Agonist-Antagonist Actuator: An Example
[0094] In FIG. 6, one implementation of the actuator is shown as an
example. For this particular actuator form, the active element
comprises a motor in parallel with a variable damper. The flexion
and extension motors can control the position of flexion and
extension nuts, respectively, via ballscrew mechanical
transmissions. As seen in FIG. 6, two side-by-side actuators are
attached at their upper ends to a cross-rod 601 which provides a
connection point to the upper link 603 of the joint mechanism. The
upper link 603 is connected to the lower link 605 at a joint
607.
[0095] The actuator that extends along the left-hand side of the
upper and lower links 603 and 605 as seen in FIG. 6 includes an
extension nut 611 that engages with and compresses an extension
spring 613. The extension spring 613 is positioned between the
extension nut 611 and a linear bearing 617 which is attached to the
lower link 605. An extension ballscrew seen at 621 connected via a
gearbox (not shown) to the armature of an extension motor 623. An
extension nut guidance shaft 625 is attached to the case of the
motor 623 and extends downwardly from the motor 623 through the
extension nut 611 and the linear bearing 617 to a shaft endcap 629.
The guidance shaft 625 prevents the extension nut from rotating so
that, as the extension motor 623 rotates the extension ballscrew
621, the extension nut 611 moves longitudinally with respect to the
cross-rod 601 and the motor 623, varying the joint angle at which
the extension nut engages with the extension spring 613. Thus, the
extension motor 623 can compress the extension spring 613 as the
extension nut 611 is driven downward to increase the length of the
actuator and extend (increase) the joint angle.
[0096] The actuator that extends along the right-hand side of the
upper and lower links 603 and 605 as seen in FIG. 6 includes a
flexion nut 631 that engages with and compresses a flexion spring
633. The flexion spring 633 is positioned between the flexion nut
631 and a linear bearing 637 which is attached to the lower link
605. A flexion ballscrew seen at 641 connected via a gearbox (not
shown) to the armature of a flexion motor 643. A flexion nut
guidance shaft 645 is attached to the case of the flexion motor 643
and extends downwardly from the motor 643 through the linear
bearing 637 and the flexion nut 631 and the to a flexion shaft
endcap 649. The flexion nut guidance shaft 645 prevents the
extension nut from rotating so that, as the flexion motor 643
rotates the flexion ballscrew 641, the flexion nut 631 moves
longitudinally with respect to the cross-rod 601 and the flexion
motor 643, varying the joint angle at which the flexion nut engages
with the flexion spring 633. Thus, the flexion motor 623 can
compress the flexion spring- 633 as the flexion nut 631 is driven
upwardly to decrease the length of the actuator and decrease the
joint angle during flexion.
[0097] A variable damper is connected in parallel with each of the
motors. An extension variable damper seen at 651 is connected in
parallel with the extension motor 623 and a flexion variable damper
seen at 653 is connected in parallel with the flexion motor
643.
[0098] Through the independent control of flexion and extension nut
positions, the actuator length at which the flexion and extension
springs are engaged can be independently controlled (Muscle-Like
Property 3). Furthermore, the flexion and extension motors can
compress each series spring simultaneously without the joint
rotating where each spring exerts an equal but oppositely opposed
force.
[0099] If the series springs are hardening springs where spring
stiffness increases with increasing compression, joint stiffness
can be effectively controlled through this agonist-antagonist motor
action (Muscle-like property 4). After the motors co-contract and
compress the flexion and extension springs to a desired spring
deflection and a desired actuator stiffness, to maintain that
stiffness, the variable dampers can output high damping levels to
impede ballscrew rotation at low power requirements.
[0100] Since each motor is in parallel with each variable damper,
both motors can be turned off while still maintaining spring
deflection and overall actuator stiffness (Muscle-Like Property 2).
The actuator can also dissipate mechanical energy at low power
(Muscle-Like Property 2).
[0101] In the actuator form of FIG. 6, the ballscrew transmissions
are backdrivable. Hence, when an external agent compresses or
lengthens the actuator, energy can be dissipated using the variable
dampers. Since each variable damper is in parallel with each motor,
during such a dissipative action, the motors can act as generators
to store electrical power for later use. Finally, zero actuator
force can be achieved at zero power consumption (Muscle-Like
Property 1). If the motors move the ballscrew nuts away from their
respective spring element, the actuator will output zero force and
no energy is required to maintain that force.
[0102] Component Implementations
[0103] Active Element. Depending on the application, each active
element could be either a motor or a variable damper/clutch, or a
combination of these elements. If the active element includes a
variable damper/clutch, it could be implemented using hydraulic,
pneumatic, friction, electrorheological, magnetorhelogical,
hysteresis brake, or magnetic particle brake damping/clutching
strategies. The preferred mechanism for damping control is a
hysteresis brake because the zero power damping level is
negligible. This feature is important because the variable damper
is behind the mechanical transmission where any strain rate
dependent, low-end viscous or frictional effect would likely be
amplified.
[0104] If the active element includes a motor, it could be any
electric motor, brushed or brushless. It could also be a hydraulic
or pneumatic cylinder or other mechanical power-producing elements
such as artificial muscle, piezoelectrics or nitinol wire.
[0105] Spring. The springs could be implemented as linear or
torsional spring elements. They may be metal die springs, carbon
fiber leaf springs, elastomeric compression springs,or pneumatic
springs. For the preferred implementations described in this
specification, the springs are die compression springs.
[0106] Mechanical Transmission. The mechanical transmissions could
be implemented as linear or torsional transmission elements. They
could be harmonic drives, ballscrew drives, leadscrew drives, or
any other mechanical transmission known in the art. For the case
where the active element and the series spring are both linear or
both rotary elements, and no gear reduction is deemed necessary,
the transmission would simply be a material linkage, connecting
spring to active element. For example, if the active element is a
linear artificial muscle, and the spring a linear, elastomeric
element, then the spring would simply be attached directly to the
artificial muscle. For the preferred embodiments described in FIGS.
6-10, the mechanical transmissions are ballscrew transmissions.
TABLE-US-00001 TABLE 1 Mechanical components of the
Agonist-Antagonist Actuator System Component Function Spring Store
and release energy, absorb shock, provide stiffness Active Element
Control positive and negative work and power, control effective
spring equilibrium length and stiffness, generate electrical power,
clutch to engage series elasticity Mechanical Transmission Couple
spring to active element, offer gear reduction between active
element and output, convert rotary active element to linear spring
element
[0107] Sensing Implementations
[0108] For the Agonist-antagonist actuator to function properly,
there are various sensors required to measure the state of the
various actuator components. The sensors required to enable general
actuator operation and control are: [0109] 1) Position sensors
located at the biomimetic joint axis to measure joint angle (a
rotary potentiometer), and at the active element (motor/variable
damper/clutch) rotor to measure total displacement of the element's
drive shaft and additionally the active element's velocity (a shaft
encoder). [0110] 2) A force sensor (strain gauges) to measure the
actual torque borne by the joint. [0111] 3) A displacement sensor
on each spring in order to measure the amount of energy stored.
[0112] Instead of directly measuring the deflection of the series
springs (#3), sensory information from #1 can be employed. By
subtracting the biomimetic joint angle from the active element
output shaft angle, it is possible to calculate the amount of
energy stored in the motor series spring. Also, the series spring
displacement sensor can be used to measure the torque borne by the
joint because joint torque can be calculated from the series spring
output force.
[0113] Many variations exist in the particular sensing
methodologies employed in the measurement of the listed parameters.
Although preferred sensory methods have been specified, it is noted
here that what is critical is to capture the energy state of the
spring elements and the velocities of interior points.
[0114] In the remaining sections, we present embodiments of the
agonist-antagonist actuator capable of providing biologically
realistic dynamic behaviors for an artificial ankle and knee
joint.
[0115] An Agonist-Antagonist Actuator for an Artificial Ankle
Joint
[0116] Mechanical Design
[0117] The ankle design comprises flexion and extension motors for
the active elements, and corresponding flexion and extension
transmissions and springs. The flexion and extension motors provide
control of joint spring equilibrium position and stiffness, damping
and nonconservative, motive force output. In the section to follow,
we provide an example of how the agonist-antagonist actuator could
be employed as an artificial ankle.
[0118] The Agonist-antagonist actuator, as used in an artificial
ankle application, is shown in FIGS. 7A and 7B. An upper shin link
701 and a foot link 702 rotate with respect to one another about an
ankle joint 705 as best seen in the side view, FIG. 7B. Two
actuators extend in parallel alongside the shin link 701. In the
actuator seen at the left in FIG. 7A, a plantar flexion motor 711
drives a flexion ballscrew 713 that extends through a linear
bearing 715, a plantar flexion spring 717 and a plantar flexion nut
719 to an endcap 720. The flexion motor 711 is attached to a
crossrod 723 by a strut 725. The dorsiflexion actuator is seen at
the right in FIG. 7A and includes a dorsiflexion motor 731 which is
attached at its lower end by a strut 735 to the cross rod 723. A
dorsiflexion ballscrew 741 is driven by the dorsiflexion motor and
extends upwardly through a dorsflexion nut 743, a dorsiflexion
spring 747, and a linear bearing 758 to an endcap 749. The foot
link 702 is attached to a leaf spring foot plate seen at 750.
[0119] The description that follows explains how, during
level-ground walking, the joint might be controlled for the swing,
controlled plantar flexion (CP), controlled dorsiflexion (CD), and
powered plantar flexion (CP) phases of gait. In addition, the
description will explain how the joint might be controlled for
stair/slope ascent and descent.
[0120] Level-Ground Walking: Swing Phase and CP
[0121] During early swing, the plantar flexion ballscrew nut 719 is
positioned such that the ankle joint is dorsiflexed to achieve foot
clearance. During terminal stance, three distinct control methods
can be employed in preparation for heel strike and the CP phase. In
human walking, the amount of energy stored during CP increases with
increasing walking speed. To achieve this increase in energy with
speed, the total angular deflection of the ankle can be increased
with increasing speed and/or the quasi-stiffness or the actual
stiffness of the ankle can be increased. Thus, in a first control
approach, the effective spring equilibrium length of the actuator
at heel strike could be increased with increasing walking speed.
Here the spring equilibrium position of the joint is equal to the
desired heel strike ankle angle. The effect of this control would
be that more mechanical energy is stored in the dorsiflexion spring
during CP as walking speed increases. In an alternate approach,
during terminal swing both dorsi and plantar flexion motors 731 and
711 could do work on their respective series springs in a
co-contraction control scheme. If the series springs are hardening
springs (stiffness increases with increasing deflection), this
co-contraction action would effectively increase the actual
stiffness of the actuator, and the ankle joint across which the
actuator spans. Still further, in a third approach, the
quasi-stiffness of the actuator/joint could be increased or
decreased during CP. For the ankle system shown in FIGS. 7A and 7B,
the flexion and extension ballscrews are non-backdriveable. Hence,
during CP, if the desired ankle stiffness can be achieved simply by
compressing the dorsiflexion spring 747, the dorsiflexion motor 731
can be turned off to conserve power. If a lower quasi-joint
stiffness is required, the dorsiflexion motor 731 can unwind the
dorsiflexion spring 747 during CP, and if a greater quasi-joint
stiffness is required, the motor can compress the spring 747 during
CP. Depending on the terrain (smooth or uneven), walking speed, and
power consumption constraints, the control algorithm of the
artificial ankle will select the appropriate ankle spring
equilibrium and stiffness values for terminal swing/CP to achieve a
smooth heel strike to forefoot strike transition.
[0122] It is noted here that in the invention described herein,
there can be separate series spring stiffnesses for joint dorsi and
plantar flexion, and these two sets of springs 717 and 747 can be
selected to give distinct flexion and extension joint stiffnesses
at little to no power consumption. If the motors change ankle
position when minimal torques are applied to the joint, such as
during the swing phase of walking, very little electrical power is
required to change the spring equilibrium position of the joint. In
the embodiment seen in FIGS. 7A and 7B, where the ballscrews 713
and 741 are non-backdriveable, the motors need not consume any
electrical power to hold the joint's position even during ground
contact. Controlling the joint spring set point at heel strike can
be useful, for example, when the wearer switches shoes with
different heel heights or when the terrain changes character
(slopes/stairs and uneven terrain), thus changing the natural angle
of the ankle joint when the foot is resting on a flat ground
surface.
[0123] Level-Ground Walking: CD and CP Phases
[0124] During early CD in human walking, the ankle torque does not
return to point 1 in FIG. 2. Rather, the torque assumes a higher
value compared to the torque values from points 1 to 2. To achieve
this higher torque output, the plantar flexion motor 711 has to
move the plantar flexion nut 719 to reduce the gap between the nut
and the plantar flexion spring 717 as the dorsiflexion spring 747
is being compressing during CP. This repositioning of the plantar
flexion nut allows the plantar flexion spring to be engaged even
before the dorsiflexion spring has released its energy, thus
providing a higher torque during early CD than during CP.
[0125] During mid to terminal CD in human walking, the ankle torque
versus angle curve becomes increasingly nonlinear as walking speed
increases. In addition, peak ankle power and the net ankle work
during stance increases with increasing walking speed (see FIG. 2).
Thus, at 0.9 m/sec, when the human ankle, on average, stores as
much energy as it releases, the mechanical response of the
artificial ankle during CP will, on average, be dictated by the
series, plantar flexion spring. That is to say, the stiffness of
the plantar flexion spring will be tuned to correspond to the
average, quasi-stiffness (slope of the torque-angle curve) of the
human ankle during CD. To decrease the quasi- stiffness of the
artificial ankle during CP, the plantar flexion motor would be
controlled to unwind the plantar flexion spring, and to increase
quasi-stiffness, the motor would compress the spring. Thus, as
walking speed increases above 0.9m/sec, the plantar flexion motor
would compress the plantar flexion spring during CD to achieve the
following characteristics 1) to increase the quasi-stiffness of the
artificial ankle during CD and 2) to increase the power output and
the positive work performed during PP. It is noted here that to
achieve a passive, spring response during the stance period of
walking, the flexion and extension motors can be turned off to
conserve power since the ballscrews are non-backdriveable.
[0126] From {1 } {2 }, it has been shown that the maximum
dorsiflexion ankle torque during level-ground walking is in the
range from 1.5 Nm/kg to 2 Nm/kg, i.e. around 150 Nm to 200 Nm for a
1.00 kg person. Further, the maximum controlled plantar flexion
torque is relatively small, typically in the range of 0.3 Nm/kg to
0.4 Nm/kg. Because of these biomechanics, a uni-directional spring
in parallel with the agonist-antagonist actuator of FIGS. 7A and 7B
would lower the peak torque requirements of the actuator. The
uni-directional spring would engage at a small or zero dorsiflexion
angle (90 degrees between foot and shank) and would lower the peak
torque requirements of the Agonist-antagonist actuator since the
peak controlled plantar flexion torque is considerably smaller than
the peak dorsiflexion torque. Thus, additional elements could be
added to the design of FIG. 7 such as a parallel, uni- directional
spring.
[0127] Stair/Slope Ascent and Descent
[0128] For ascending a stair or slope, the dorsi and plantar
flexion motors would move the nuts to reposition the ankle joint to
an appropriate angle given the nature of the stair/slope. Once the
artificial toe is loaded at first ground contact, the plantar
flexion spring compresses and stores energy. During this CD process
the plantar flexion motor can compress the spring farther so that
additional power is delivered to the walking robot or
prosthesis/orthosis user during PP. After toe-off, the motors
control the equilibrium position of the ankle in preparation for
the next step.
[0129] During stair descent, the body has to be lowered after
forefoot contact until the heel makes contact with the stair tread.
See reference {2}. During this CD phase, the plantar flexion motor
unwinds the plantar flexion spring as the spring is compressing to
effectively dissipate mechanical energy. Once the heel makes
contact with the stair tread, the motor can be turned off so that
the plantar flexion spring begins to store energy for release
during PP. For slope descent, the ankle response is similar, except
that mechanical energy is absorbed by the dorsiflexion motor during
CP instead of during CD.
[0130] An Agonist-antagonist actuator for an Artificial Knee
Joint
[0131] The knee design comprises an extension motor and a flexion
variable damper for the active elements, and corresponding flexion
and extension transmissions and springs. The extension motor and
the flexion variable damper provide control of joint spring
equilibrium position and stiffness, damping and nonconservative,
motive force output. In this implementation of the
agonist-antagonist actuator, a flexion motor is not included in an
attempt to simplify the mechanism. Since only a flexion variable
damper is present, the flexion nut is mechanically grounded to the
linear bearing since a flexion motor is not present to actively
reposition the flexion nut. Hence, when the knee joint flexes and
extends, the flexion ballscrew rotations, but that rotation does
not introduce significant zero-power joint resistance because 1)
the flexion ballscrew is highly backdriveable and 2) the flexion
variable damper has a negligible low-end damping value. A preferred
method for the flexion variable damper is a hysteresis brake
because of its minimal low-end damping value. In the section to
follow, I provide an example of how the agonist-antagonist actuator
could be employed as an artificial knee.
[0132] The agonist-antagonist actuator, as used in an artificial
knee application, is shown in FIGS. 8A and 8B. The actuator
consists of an upper (thigh) link 801 and a lower (shin) link 803
which are rotatably connected at a joint 805. As seen at the left
of the lower link 803, an extension motor 811 drives an extension
ballscrew 813 that extends downwardly from the motor 811 through an
extension nut 815, an extension spring 817, and a linear bearing
819. An extension nut guidance shaft 825 prevents the extension nut
from rotating as the extension ballscrew 813 rotates.
[0133] The mechanism on the right side of the lower link 803 is
passive; that is, it does not include an active motor element but
rather includes a flexion variable damper 831 and a flexion spring
833. A flexion ballscrew 841 extends from the damper 831 downwardly
through a linear bearing 843, the flexion spring 833 and a flexion
nut 847. A flexion nut guidance shaft 851 prevents the flexion nut
847 from rotating as the extension ballscrew 841 rotates.
[0134] Level-Ground Walking
[0135] During level-ground walking, the joint is controlled for the
swing, early stance flexion, mid-stance extension, and pre-swing
phases of gait. In addition, as described below, the joint may be
controlled for stair/slope ascent and descent. Beginning at heel
strike, the stance knee begins to flex slightly in normal human
walking (FIG. 5). As was noted earlier, this flexion period, called
the Stance Flexion phase, allows for shock absorption upon impact
as well as to keep the body's center of mass at a more constant
vertical level throughout the stance period. During this phase, the
artificial knee shown in FIGS. 8A and 8B outputs a spring response,
storing energy in preparation for the Stance Extension phase. Here
the extension spring 817 stores energy, and then that energy is
released during the Stance Extension phase. In this implementation
of the agonist-antagonist actuator, the extension ballscrew
transmission is non-backdriveable. Thus, if the desired actuator
stiffness during Stance Flexion corresponds to the extension spring
stiffness, the extension motor need not be active, reducing
electrical power requirements. If a higher or lower quasi-joint
stiffness is desired, the extension motor 811 can compress or
unwind the extension spring 813 during early stance knee flexion,
respectively, by repositioning the extension nut 815 that acts on
the extension spring 817.
[0136] After maximum flexion is reached in the stance knee in
normal human walking, the joint begins to extend, until maximum
extension is reached. This knee extension period is called the
Stance Extension phase. Throughout the first .about.60% of Stance
Extension, the knee acts as a spring, releasing the stored energy
in the extension spring from the Stance Flexion phase of gait. This
first release of energy corresponds to power output P2 in FIG. 5B.
During the last .about.30% of Stance Extension, the artificial knee
is controlled to absorb energy in the flexion spring 833 and then
that energy is released during the next gait phase, or Pre-Swing.
Here the energy from hip muscular work and the remaining stored
energy in the extension spring 817 is then stored in the flexion
spring 833. To engage the flexion spring, the flexion variable
damper 831 outputs a high damping value, locking the flexion
ballscrew 841, and forcing the flexion nut 847 to compress the
flexion spring 833. During this energy storage, if it is desirable
to lower the effective quasi-stiffness of the joint, the flexion
variable damper 831 can output lower damping values to allow the
flexion ballscrew 841 to slip, and for energy to be dissipated as
heat. Here again, as in the artificial ankle joint of FIGS. 7A and
7B, the flexion and extension springs of the agonist antagonist
actuator of FIGS. 8A and 8B are precisely tuned such that
biological knee mechanics can be achieved while minimizing power
supply demands and overall artificial joint mass.
[0137] During late stance or Pre-Swing, a normal human knee of the
supporting leg begins its rapid flexion period in preparation for
the swing phase. During early Pre-Swing in the artificial knee
joint of FIGS. 8A and 8B, as the knee begins to flex in preparation
for toe-off, the stored elastic energy in the flexion spring 833
stored during Stance Extension is released. This second release of
energy corresponds to power output P3 in FIG. 5B. During this
process, the flexion variable damper 831 can be used to modulate
the amount of stored elastic energy in the flexion spring that is
actually released to power the knee joint.
[0138] In normal human walking, as the hip is flexed, and the knee
has reached a certain angle in Pre-Swing, the leg leaves the ground
and the knee continues to flex. At toe-off, the Swing Flexion phase
of gait begins. Throughout this period, human knee power is
generally negative where the knee's torque impedes knee rotational
velocity. In the artificial knee joint of FIGS. 8A and 8B, once the
elastic energy from the flexion spring 833 has been released and
the artificial leg has entered the swing phase, the knee joint
typically has to absorb mechanical energy to decelerate the
swinging lower leg. This can be done in two ways. First, the
flexion variable damper 831 can be used to dissipate mechanical
energy as heat and to decelerate the swinging artificial leg. In
addition, during late Swing Flexion, the extension motor 811 can
position the extension ballscrew nut 815 such that the extension
spring 817 compresses and stores elastic energy for use during
Swing Extension.
[0139] After reaching a maximum flexion angle during swing, a
normal human knee begins to extend forward. For the artificial knee
of FIGS. 8A and 8B, during the early Swing Extension period, the
elastic energy stored during late Swing Flexion in the extension
spring 817 is released, resulting in power output P4 in FIG. 5B.
This control action, once again, reduces the energy demands from
the knee's power supply. In all cases, the flexion variable damper
831 can be used to precisely modulate the amount of power delivered
to the swinging artificial leg from the stored elastic energy.
[0140] During the remainder of Swing Extension, the human knee
typically outputs negative power (absorbing energy) to decelerate
the swinging leg in preparation for the next stance period. As with
Swing Flexion, this can be done in two ways. First, the flexion
variable damper 831 can be used to dissipate mechanical energy as
heat and to decelerate the swinging artificial leg. In addition,
during late Swing Extension, the flexion variable damper 831 can
output a relatively high damping value such that the flexion spring
833 compresses and stores elastic energy for use during Stance
Flexion. Here a small amount of energy is stored in preparation for
early stance (power P1). After the knee has reached full extension,
the foot once again is placed on the ground, and the next walking
cycle begins.
[0141] In summary, the artificial knee shown in FIGS. 8A and 8B is
capable of reproducing the positive power contributions P1, P2, P3
and P4 shown in FIG. 5 for level-ground walking.
[0142] Stair/Slope Ascent and Descent
[0143] For stair/slope descent, a normal human knee performs
negative work during stance where knee torque is in the opposite
direction to knee rotational velocity. The agonist-antagonist
actuator of FIGS. 8A and 8B can perform this negative work in two
ways. First, the flexion variable damper 831 can be used to
dissipate mechanical energy as heat and to decelerate the rotating
artificial leg. In addition, during terminal stance, the extension
motor 811 can position the extension ballscrew nut 815 such that
the extension spring 817 compresses and stores elastic energy for
use later to power Swing Extension to prepare the artificial leg
for the next stance period.
[0144] For stair/slope ascent, during the swing phase the extension
motor 811 can actively control knee position to accurately locate
the foot on the next stair tread or slope foothold. Once the
artificial foot is securely positioned on the stair tread or
ground, the motor 811 can then deflect and store energy in the
extension spring 817. This stored elastic energy can then assist
the knee wearer or humanoid robot to actively straighten the knee
during the stance period, lifting the body upwards.
[0145] Finally, the agonist-antagonist actuator of FIGS. 8A and 8B
allows for the "windup" phase of a catapult style control to occur
at any desired time. This means much greater flexibility as to when
large amounts of power can be efficiently generated and used. This
flexibility is critical when designing an artificial knee that can
be used for jumping. For such a movement task, energy has to be
stored prior to the jump, and then the elastic energy has to be
released at a precise time to facilitate a jumping action.
Specifically for the agonist-antagonist actuator of FIGS. 8A and
8B, the flexion variable damper 831 would be controlled to output
high damping to effectively lock the flexion ballscrew 841.
Following this action, the extension motor 811 would slowly
compress the extension spring 817. Once high powers are deemed
necessary about the joint output, the flexion variable damper 831
would then be controlled to suddenly unlock to allow rapid rotation
of the flexion ballscrew 841 and the release of elastic strain
energy from the extension spring 817.
[0146] Alternative Configurations of the Agonist-Antagonist
Actuator
[0147] It should be understood that the agonist-antagonist actuator
described herein could be implemented in a number of different
ways. For example, an active element and transmission-spring
combination could be positioned on each side of the artificial
joint. This configuration, shown in FIG. 9, has the advantage that
when only one spring is being compressed, no off-axis bending
torques are borne by the lower link seen at 901. The lower link 901
is attached to the upper link 903 at a joint 905. A crossbar strut
907 is rigidly attached to the lower link 901. A linear bearing is
attached to each end of the crossbar strut 907 and a ballscrew, one
of which is seen at 909, extends through the linear bearing. The
ballscrew seen at 909 extends downwardly from a drive motor 911
through a variable damper 913, the linear bearing, a spring 915,
and a ballscrew nut 917 to an end cap 919.
[0148] In the agonist-antagonist actuator implementations shown in
FIGS. 6, 7 and 8, when only a single spring is being compressed,
the upper and lower links experience a bending torque because the
pair of active element- transmission-spring combinations are on the
same side of the joint axis. It should also be understood that more
than two active element-transmission-spring combinations could be
employed to actuate multiple degrees of freedom. For example, in
FIG. 10, four active element- transmission-spring combinations are
shown to actuate a two degree of freedom joint. Still further, it
should be understood that an agonist-antagonist actuation system
can include active element-transmission-spring combinations than
span two or more joints in a poly-articular architecture. The
biomechanics of poly-articular actuation is discussed in the next
section.
[0149] In the arrangement shown in FIG. 10, the joint attaches an
upper link 1001 to a lower link 1003 for rotation about two
orthogonal axes. As seen in FIG. 10B, the upper link rotates in a
first degree of freedom about an axis through the crossbar 1007
that is parallel to the long dimension of a crossbar 1009, and in a
second degree of freedom about an axis through the crossbar 1009
that is parallel to the crossbar 1007. Four different actuators are
attached from the ends of the crossbars 1007 and 1009 and all four
have a like structure illustrated by the actuator at the left in
FIG. 10A. An drive motor 1021 attached to the crossbar 1005 rotates
a ballscrew 1022 that passes through variable damper 1027 and a
linear bearing 1029 attached to the lower link 1003. The ballscrew
1022 further extends through a series spring 1031 and a ballscrew
nut 1033 to an endcap 1040. For each degree of freedom, one of the
motor-spring-damper mechanisms controls the rotation of the upper
link 1001with respect to the lower link 103 in one direction while
an opposing motor-spring-camper mechanism attached to other end of
the same crossbar controls the rotation in that degree of freedom
in the other direction, thus providing agonist-antagonist actuator
control in both degrees of freedom.
[0150] Agonist-Antagonist Actuators Spanning More than One
Joint
[0151] In the foregoing description, the agonist-antagonist
actuator mechanism contemplated by the present invention was
described and specific examples were provided as to its use in
ankle and knee actuation, and different illustrative
implementations were described. For each of these implementations,
the agonist-antagonist actuator spanned a single joint. In other
implementations, an agonist-antagonist actuator may span more than
one rotary joint. The functional purpose of poly-articular muscle
architectures in the human leg is to promote the transfer of
mechanical energy from proximal muscular work to distal joint power
generation. See reference {10}. To capture truly biomimetic limb
function, both muscle-like actuators and mono, bi, and
poly-articular artificial musculoskeletal architectures are
critical. Hence, it should be understood that the
agonist-antagonist actuator described herein could span more than
one artificial joint. For example, an active
element-transmission-spring combination could act across the hip
and knee of an artificial leg, or across the knee and ankle of an
artificial leg.
[0152] The Biomechanics of Mono and Bi-Articular Leg Actuation
[0153] In the previous sections, an agonist-antagonist actuator was
described and specific examples were provided as to its use in
ankle and knee actuation. For each of these descriptions, the
actuator was used as a mono-articular device, spanning only a
single joint. In subsequent embodiments, we describe how
mono-articular actuation strategies can be used in combination with
bi-articular actuation strategies to better replicate biological
limb dynamics and efficiency.
[0154] The functional purpose of bi-articular muscle architectures
in the human leg is to promote the transfer of mechanical energy
from proximal muscular work to distal joint power generation {10}.
To better explain how bi-articular actuation effects biological
limb energetics, we present a biomechanical model of the human
musculoskeletal architecture in FIG. 11A{ 11 }. By modeling the
human leg, we seek to understand how leg muscles and tendons work
mechanically during walking in order to motivate the design of
efficient prosthetic, orthotic, and robotic limbs.
[0155] We hypothesize that a robotic leg comprising only knee and
ankle variable-impedance elements, including springs, clutches and
variable-damping components, can capture the dominant mechanical
behavior of the human knee and ankle for level-ground ambulation.
As a preliminary evaluation of this hypothesis, we put forth a
simple leg prosthesis model, shown in FIG. 11A, that is motivated
by the human leg musculoskeletal architecture {11}. The model seen
in FIG. 11 includes a drive motor 1101 at the hip, a knee joint
1103 and an ankle joint 1105. A musculo-skeletal model of human leg
function in walking. The model comprises seven mono-articular
series-elastic clutches and four bi-articular series-elastic
clutches/variable-dampers. Only a single actuator 1101 acts at the
model's hip joint. In (B) and (C), model predictions for ankle and
knee are compared with human gait data, respectively. Here gait
data are shown for a 70 kg study participant with a 0.9 meter leg
length and a walking speed of 1.2 m/s. The model of (A) agrees well
with the human gait data, suggesting that muscles that span the
ankle and knee primarily act as variable-impedance devices during
level-ground walking. We vary quasi-passive model parameters, or
spring constants, damping levels and times when series-elastic
clutches are engaged, using an optimization scheme where errors
between model joint behaviors and normal human joint biomechanics
are minimized.
[0156] The capacity of the musculoskeletal leg model to capture
human-like ankle and knee mechanics in level-ground walking is
shown in FIGS. 11B and 11C, respectively. At each joint state
(position and velocity), the leg model is in good agreement with
experimental values of joint torque and power, suggesting that a
robotic leg can produce human-like walking dynamics through the
control of only knee and ankle impedance.
[0157] Mono-articular ankle mechanism. The ankle mechanism
comprises mono-articular dorsi and plantar flexion springs that can
be engaged or disengaged with series elastic clutch mechanisms (see
FIG. 11A). In FIG. 12A, the mechanical power for each ankle
component is plotted versus percent gait cycle. At heel strike (0%
cycle), the clutch for the ankle dorsiflexion spring is engaged,
causing the spring to stretch and store energy during early stance
plantar flexion. When the tibia begins rotating forwardly after
forefoot contact, the ankle plantar flexion spring is engaged and
continues to store energy throughout the controlled dorsiflexion
phase, and then that stored energy is released to contribute to
ankle powered plantar flexion at terminal stance. Mechanical power
output for each component of the human leg model of FIG. 11A. In
(A), (B) and (C), the mechanical power of each model element is
plotted versus percentage gait cycle for ankle, knee and hip,
respectively. Here the gait cycle begins at heel strike (0%) and
ends with the heel strike of the same leg (100%).
[0158] Mono-articular knee mechanism. The knee mechanism comprises
mono-articular flexion and extension springs that can be engaged or
disengaged with series elastic clutch mechanisms (see FIG. 11A). In
FIG. 12B, the mechanical power for each knee mono-articular
component is plotted versus percent gait cycle. At heel strike (0%
cycle), the clutch for the knee extensor spring is engaged, causing
the spring to stretch during early stance knee flexion. Here the
knee extensor spring inhibits the knee from buckling. As the knee
extends from a flexed posture, the knee flexor spring is engaged at
the point of maximum knee extension velocity, storing energy that
is subsequently used during terminal stance to help lift the lower
leg from the ground surface.
[0159] Ankle-Knee Bi-Articular Mechanism. The leg model's
ankle-knee bi-articular mechanism comprises a spring that can be
engaged or disengaged with two clutch mechanisms (see FIG. 11A). A
first clutch, or the distal clutch, attaches the series spring to a
point between the ankle and knee joint, and a second clutch, or the
proximal clutch, attaches that same spring to a point above the
knee axis. After heel strike in human walking, the knee typically
undergoes a flexion period. During that phase of gait, both the
proximal and distal clutches are disengaged, and the bi-articular
spring does not apply a force to the prosthesis skeleton. However,
as the knee begins to extend (.about.10% cycle), the proximal
clutch engages, and the bi-articular spring stretches. When the
knee is fully extended, the distal clutch changes from a disengaged
state to an engaged state, and the proximal clutch disengages.
Engaging the distal clutch mechanically grounds the bi-articular
spring below the knee rotational axis, changing the ankle-knee
mechanism from a bi-articular to a mono-articular device. As a
consequence of this action, all the energy stored in the
bi-articular spring is used to power ankle plantar flexion during
terminal stance. Thus in summary, the ankle-knee mechanism allows
energy from hip muscular/actuator work to be transferred to the
ankle for late stance powered plantar flexion.
[0160] Knee-Hip Bi-Articular Mechanism. The leg model's knee-hip
bi-articular mechanisms comprise a spring that can be engaged or
disengaged with either a clutch or variable-damper mechanism (see
FIG. 11A). There are three knee-hip bi-articular mechanisms. The
clutch of the knee-hip flexor is engaged during swing phase knee
extension and begins storing energy its series spring. As a result
of this control action, the lower leg is decelerated smoothly prior
to reaching full knee extension. In addition, elastic energy is
stored in the knee-hip flexor spring that is later released during
the early stance period. The knee-hip flexor also undergoes an
energy storage/release sequence that begins during stance knee
extension. The stored energy is then released to power rapid knee
flexion movements at terminal stance to lift the foot and lower leg
from the ground surface. The clutch of the knee-hip extensor is
engaged during terminal stance, storing energy that is later
released to enhance knee extension. Finally, the iliotibial tract
series-elastic variable-damper applies an extensor knee torque to
offset the knee flexor torque applied by the ankle-knee
bi-articular mechanism. During stance knee extension, the
ankle-knee bi-articular spring is elongated, exerting a torque
about the knee. At the same time the iliotibial tract series spring
is elongated thereby applying an extensor torque at the knee. Thus,
through the action of the iliotibial tract mechanism, the effect of
the ankle-knee bi-articular mechanism on net knee torque is
minimized.
[0161] In the human leg, the functional purpose of bi-articular
muscle is to promote the transfer of mechanical energy from
proximal muscular work to distal joint power generation {10}. Using
the biomimetic architecture shown in FIG. 11A, the robotic leg can
achieve ankle powered plantar flexion without the requirement of
powering a large motor located at the ankle joint. Approximately
ten joules of net work are transferred to the ankle from the knee
and hip in the modeling results shown in FIGS. 11 and 12.
[0162] In subsequent embodiments, we motivate the design of
prosthetic, orthotic and robotic leg structures using the leg model
of FIG. 11.
[0163] Mono and Bi-Articular Actuation for a Transtibial Prosthetic
Leg System
[0164] The prosthetic leg model of FIG. 11A suggests that leg
prostheses could produce human-like joint mechanics during
level-ground ambulation if a musculoskeletal leg architecture and a
variable-impedance control paradigm were exploited. However, the
proposed biomimetic leg prosthesis does not eliminate the need for
knee and ankle actuators, but the model does suggest that
nonconservative joint actuator work need not be performed during
normal, steady state walking. For some situations, positive joint
actuator work is required. For example, for uphill locomotory
function, some positive actuator work would be necessary,
especially at the knee. Furthermore, ankle and knee torque control
would be necessary to reject large whole-body force disturbances
that threaten balance. Although joint actuation is still necessary,
the proposed biomimetic design will increase the time between
battery recharges or power supply refueling, and will reduce
robotic limb noise production during level-ground walking.
[0165] In FIG. 13, the design of the proposed transtibial
ankle-foot system with mono and bi-articular mechanisms is shown.
In FIG. 13A, the major components of the transtibial system are
shown, including the mono-articular ankle mechanism at 1303, the
bi-articular ankle-knee mechanism at 1305, and a flexible nylon
cord at 1307. FIG. 13B shows the monoarticular ankle mechanism in
more detail. This mechanism comprising two motors 1313, mechanical
transmissions and series dorsiflexion springs at 1317, and series
plantar flexion springs at 1319. FIG. 13C shows the bi-articular
mechanism (1305 in FIG. 13A) and FIG. 13D shows a schematic of the
bi-articular mechanism, including two uni-directional clutches seen
at 1321 and 1351 and a series spring at 1324. The limb architecture
largely reflects the leg model shown in FIG. 11A, except the
mono-articular knee mechanism has been excluded as this basic
musculoskeletal structure is still intact in transtibial
amputees.
[0166] The ankle mechanism 1305 seen in FIG. 13B comprises two
agonist-antagonist, series-elastic actuators acting across the
ankle joint. The foot-ankle design is similar to that described
earlier in FIG. 7. Each actuator has a small electric motor 1313 in
series with one of the die springs 1317 or 1319. Each series spring
is a nonlinear hardening spring where spring stiffness increases
with increasing spring compression. A non-backdriveable leadscrew
1331 is employed to covert rotary motor movement into linear
movement of a leadscrew nut 1332. A slider mechanism is seen at
1334 and a guide rod at 1335. By re-positioning the leadscrew nut
1332, each motor 1313 can independently vary the position of the
ankle joint at which the series spring 1317 or 1319 becomes
engaged. Such an ankle spring equilibrium control is important for
many prosthesis functions, including slope and stair ascent and
descent. The mono-articular ankle mechanism can also change ankle
spring stiffness. During the swing phase each motor can
simultaneously compress each nonlinear spring using a
co-contraction control. Since spring stiffness increases with
increasing deflection, the more the motor system compresses the
springs, the stiffer the ankle joint becomes. Since the mechanical
transmission is non-backdriveable, once a desired ankle stiffness
has been achieved, the motors can be turned off to save electrical
power. The foot-ankle design is similar to that described earlier
in FIG. 7.
[0167] In FIGS. 13C and 13D, the bi-articular ankle-knee mechanism
and schematic are shown, respectively. The mechanism comprises two
uni-directional clutches seen at 1351 and 1321 and a spring at
1324. Each clutch is formed by two opposing cams (see 1353) that
press against a shaft that directly connects to the spring. At the
bottom of FIG. 13C, a foot assembly is seen at 1327 and the ankle
axis is at 1328. The ankle joint connection is seen at 1329. In an
engaged state, the cam configuration only allows for shaft movement
in one direction. As can be seen in FIG. 13D, if both
uni-directional clutches A and B are in the disengaged state, with
each cam pair rotated outwardly with a small cam motor, the ankle
and knee can freely rotate without the bi-articular spring exerting
a force. When the ankle dorsi and plantar flexes in this disengaged
state, the lower floating cam-clutch assembly 1321 translates on
the linear guide rail 1367. Furthermore, when the knee flexes and
extends, the entire spring assembly translates on the linear guide
rail 1367. In distinction, when both clutches are in their engaged
state, both ankle dorsiflexion and knee extension cause the
bi-articular spring 1324 to stretch and store energy. Since the
flexible nylon cord 1307 can resist tension but not compression,
once the knee has reached full extension during the stance phase,
knee flexion throughout terminal stance is not restricted by the
bi-articular assembly, and all the stored energy in the
bi-articular spring augments ankle powered plantar flexion.
[0168] Sensors for Active Ankle-Foot Prosthesis
[0169] For the active transtibial prosthesis to function properly,
there are various sensors required to measure the state of the
various system components and the intent of the amputee user. The
additional sensors required to enable general prosthesis operation
and control are: [0170] 4) position sensors located at the knee and
ankle axes to measure joint angles (rotary potentiometers), and on
each motor shaft to measure total displacement and velocity of each
motor (a shaft encoder); [0171] 5) an inertial measurement unit
(IMU) to determine the absolute position of the prosthesis in
space; [0172] 6) a displacement sensor on each spring in order to
measure the amount of force borne by a spring and the torque borne
by the ankle joint; and [0173] 7) electromyographic (EMG) sensors
to determine residual limb muscle activity.
[0174] Series spring displacement sensors can be used to determine
the torque borne by the ankle joint because joint torque can be
calculated from the agonist-antagonist spring output forces.
[0175] Control for Active Ankle-Foot Prosthesis
[0176] Local Prosthesis Control. A critical advantage of the
human-like musculoskeletal prosthesis is that it allows the amputee
user to directly control ankle powered plantar flexion. Because of
the bi-articular ankle-knee mechanism, the extent of midstance knee
extension defines how much energy is transferred to the prosthetic
ankle for powering ankle plantar flexion at terminal stance. Since
transtibial amputees generally have direct control over their knee,
the biomimetic transtibial prosthesis allows for direct control
over ankle power output.
[0177] The point in the gait cycle where the prosthesis series
spring elements are engaged will largely be defined by joint state
(position and velocity) and foot-ground interaction forces. The
spring equilibrium angle for the ankle mono-articular mechanism
will be equal to the ankle angle at first heel strike. Here heel
strike will be detected using ankle torque sensing. For level
ground ambulation, the heel strike ankle angle will be kept largely
invariant with walking speed, but will be modulated from step to
step for slope and stair ambulation.
[0178] The uni-directional clutch devices in the bi-articular
mechanism will be controlled in a speed invariant manner. After
heel strike in walking, the knee typically undergoes a flexion
period. During that phase of gait, both bi-articular clutches will
be disengaged, and therefore the bi-articular spring will not apply
a force to the prosthesis skeleton. However, as the knee begins to
extend (.about.10% cycle), both clutches will be engaged, causing
the bi-articular spring to stretch. Once the prosthesis enters the
swing phase as detected by zero ankle torque, the bi-articular
clutches will be disengaged so as to allow unrestricted knee and
ankle movement throughout the swing phase.
[0179] Electromyographic (EMG) Control of Prosthetic Ankle
Stiffness. The residual anatomy will allow amputees to voluntarily
control joint stiffness via activation of the muscles in the
residual limb. When walking on a rigid ground surface, the amputee
user can select a low ankle stiffness, whereas when walking on a
compliant terrain, the amputee can exploit a relatively high ankle
stiffness.
[0180] Within the human body, such voluntary changes in joint
stiffness are modulated by muscular co-activation. When antagonist
muscles are simultaneously recruited, the net torque produced about
the joint is related to the difference between the forces generated
by the activated muscles, while the joint stiffness is related to
their sum. Thus, activity from residual muscles is a natural
control source for specifying the desired level of ankle stiffness.
Since EMG provides a measure of muscular effort, it can be used in
a "natural" manner to control stiffness of a joint. For a
transtibial amputee, the muscles of the anterior and posterior
compartment of the leg form the natural location from which to
derive stiffness control signals.
[0181] A joint stiffness control signal is derived from the sum of
the plantar flexion and dorsiflexion EMG amplitudes. The stiffness
control signal will be related to stiffness via a straight line
relationship with a zero-level control signal signifying the
minimum available stiffness level and the maximum-level control
signal signifying the maximum available stiffness level. Thus,
limited muscle effort results in a low ankle stiffness while high
muscular effort results in a high ankle stiffness. Using this
control strategy, stiffness can be volitionally controlled by the
amputee in a natural manner.
[0182] Although the device of FIGS. 13A-13D was described as a
transtibial prosthesis, the mechanism could also be used as an
orthosis or exoskeleton. The mechanism would be useful as an
orthosis for an individual that suffers from an ankle pathology but
generally has normal knee and hip function. For such an
application, the mechanism would be placed in parallel with the
human leg to augment ankle mechanics as a permanent assistive
device.
[0183] Mono and Bi-articular actuation for an Artificial Ankle and
Knee System
[0184] Description
[0185] A proposed artificial ankle and knee system is shown in FIG.
14. The mechanism could be employed for a transfemoral prosthesis,
orthosis, leg exoskeleton, or robotic leg. The mono-articular
ankle-foot and knee designs are identical to the structures
described in FIG. 13B and FIG. 8, respectively. However, the
ankle-knee bi-articular mechanism is different from that proposed
in FIGS. 13C and 13D. The bi-articular device of FIG. 13 has to be
attached above the knee axis. In distinction, the bi-articular
device of FIGS. 14A-14D attaches adjacent to the knee axis.
[0186] The bi-articular ankle-knee mechanism of FIGS. 14A-14D
comprises a motor 1411, non-backdriveable mechanical transmission
1413, screw nut 1414, series spring 1417, a knee bi-articular
connection 1421, an ankle bi-articular connection 1431, and a knee
variable moment arm (VMA) device 1441 (seen in more detail in FIG.
14D).
[0187] During level-ground walking, we describe how the ankle-knee
bi-articular mechanism would be controlled for the swing, early
stance flexion, mid-stance extension, and pre-swing phases of
gait.
[0188] During the swing phase and early stance knee flexion, the
screw nut 1414 is moved away from the series spring 1417 so that
ankle and knee joint movements do not cause the spring to compress.
However, when stance knee extension begins (18% gait cycle), the
lead screw nut 1414 is moved by the motor 1411 until it engages the
series spring 1417. As a consequence of this control action, both
knee extension and ankle dorsiflexion contributes to spring
compression. Once the knee has reached full extension, the VMA
device 1441 then minimizes the moment arm that the knee
bi-articular connection makes with the knee axis of rotation.
Because the knee moment arm is minimized, most of the strain energy
stored in the bi-articular spring contributes to ankle powered
plantar flexion at terminal stance. Generally, the knee moment arm
1441 can be controlled to effectively modulate the amount of energy
release that occurs through the knee joint.
[0189] The VMA device comprises a small motor 1451 plus gear train
1455, non-backdriveable lead screw 1459, lead screw nut 1461, and
variable moment arm pin 1466. A shin tube mount is seen at 1457.
When the motor 1451 rotates, the lead screw nut 1461 moves the
variable moment arm pin 1466 across the variable moment arm slot
1471. The pin is attached to the knee bi-articular connection.
Thus, the VMA motor can actively control the perpendicular
distance, or moment arm, between the knee bi-articular connection
and the knee axis.
SUMMARY
[0190] Several agonist-antagonist actuator variations comprising a
plurality of active element transmission-spring combinations acting
in parallel have described. These actuator embodiments combine
active and passive elements in order to achieve high performance
with minimal mass. In addition, the use of agonist-antagonist
actuators as mono and poly-articular linear elements has been
described. The combination of biologically-inspired musculoskeletal
architectures and agonist-antagonist actuation strategies as
described above provide novel, low mass, efficient and quiet
biomimetic artificial limbs. These artificial limb structures may
be used to advantage to provide improved orthotic and prosthetic
devices and legged robotic mechanisms.
CONCLUSION
[0191] It is to be understood that the methods and apparatus which
have been described above are merely illustrative applications of
the principles of the invention. Numerous modifications may be made
by those skilled in the art without departing from the true spirit
and scope of the invention.
* * * * *