JP2012516717A - Actuator-powered knee prosthesis with antagonistic muscle action - Google Patents

Abstract

The artificial knee includes an actuator muscle-antagonist muscle array composed of two series elastic actuators in parallel, a knee joint, a bending and extension actuator connected to the knee joint and in parallel with a leg member, and the actuator And a controller for independently supplying energy to control movement of the knee joint and the leg member. The bending actuator includes a series connection of a bending motor and an elastic element for bending, and the extension actuator includes a series connection of an extension motor and an elastic element for extension. The sensor provides feedback to the controller. The bending actuator and the extension actuator can be unidirectional, in which case the bending elastic element and the extending elastic element are series springs. Alternatively, the extension actuator can be bidirectional, in which case the extension elastic element is a set of pre-compressed series springs. Alternatively, the bending elastic element can be a non-linear softening spring and the extension elastic element can be a non-linear softening spring.

Description

This application claims the benefit of US Provisional Patent Application No. 61 / 148,545, filed Jan. 30, 2010, the entire disclosure of which is incorporated herein by reference. Shall.

  This application is a co-pending application, US patent application Ser. No. 12 / 608,627, filed Oct. 29, 2009, which was filed on Dec. 19, 2006, but abandoned. No. 60 / 751,680 filed on Dec. 19, 2005, but has already expired, and is a continuation-in-part application. U.S. Patent Application Nos. 11 / 395,448, 11 / 495,140, 11 / 600,291, and 11 / 499,853 (which were patented as U.S. Pat. No. 7,313,463, August 4, 2005). And is a continuation-in-part application of US Patent Application No. 11 / 395,448. The entire disclosure of these US patent applications is hereby incorporated by reference.

  This application is also a U.S. Patent Application 11 / This is a continuation-in-part of 395,448. U.S. Patent Application No. 11 / 395,448 includes the benefit of the filing date of U.S. Provisional Patent Application No. 60 / 666,876 filed on March 31, 2005, and the U.S. Provisional Patent filed on August 1, 2005. Claims the benefit of the filing date of application 60 / 704,517.

  This application is also filed by Hugh M. Herr, Samuel K. Au, Peter Dilworth and Daniel Joseph Paluska on July 29, 2006, “An Artificial Ankle-Foot System with Spring, Variable-Damping, and Series-Elastic Actuator Components. Is a continuation-in-part of US patent application Ser. No. 11 / 495,140. U.S. Patent Application No. 11 / 495,140 claims the benefit of the filing date of U.S. Provisional Patent Application No. 60 / 704,517, filed on Aug. 1, 2005, and U.S. Patent Application No. 11 / 395,448. Is a continuation-in-part application

  This application is also a U.S. Patent Application 11 / This is a continuation-in-part of 600,291. U.S. Patent Application No. 11 / 600,291 claims the benefit of the filing date of U.S. Provisional Patent Application No. 60 / 736,929 filed on November 15, 2005, and U.S. Patent Application No. 11 / 395,448. No. 11 / 499,853 and 11 / 495,140.

  This application claims the benefit of the filing date of each of the above patent applications, and incorporates the disclosure of each of the above applications by reference.

STATEMENT REGARDING FEDERALLY SPONSORED RESEARCH AND DEVELOPMENT This invention was made with US government support under grant number VA241-P-0026 awarded by the US Veterans Assistance Agency. The US government has certain rights in this invention.

  The present invention relates to a prosthesis, a prosthetic limb for use in a prosthesis, an exoskeleton, a corrector, or a robotic device, in particular, a powered prosthetic knee joint.

  Most commercial prosthetic and lower limb orthoses are passive and cannot provide positive mechanical force to replicate joint biomechanics (joint biomechanics) during the gait cycle. Existing approaches to the design of powered knee systems have focused primarily on using a single motor-transmission system that is directly coupled to the joint. However, such a direct drive configuration fully emulates the mechanical behavior (or mechanical behavior) of a biological knee joint, even when walking on flat ground. ) To consume a lot of power. One reason for this energy inefficiency is that it does not fully utilize the passive dynamic properties of the legs, elastic energy storage, and tendon-like structure restoration.

  Knee prostheses for thigh amputees can be divided into three main groups: passive, variable-damping, and powered. A passive artificial knee (joint) does not require a power source (power source) for operation, but generally has less adaptability to environmental disturbances than a variable damping prosthesis. A variable-damping knee (joint) requires a power source (power source), which is only needed to change the damping (braking) level. On the other hand, a powered prosthetic knee can perform the non-conservative positive work of the knee.

  A variable-damping knee offers several advantages, including improved knee stability and adaptability for different walking speeds when compared to a mechanically passive configuration. Although a variable-damping knee offers several advantages over a completely passive knee mechanism, it cannot generate positive mechanical forces and is therefore (from a chair, etc.) The positive work phase of the human knee joint cannot be reproduced due to movements such as standing up, walking on a flat ground, and walking up stairs / slopes. Of course, when a thigh amputee uses a variable-damping knee technique, for example, the walking pattern is asymmetrical and walking speed is reduced compared to those who do not cut the thigh, and Faced with clinical problems such as the need for large metabolic energy.

  Current approaches to the design of powered prostheses, orthotics, exoskeletons, and robotic leg systems are primarily focused on using a single motor-transmission system directly coupled to the joint. . Such a direct drive configuration consumes a large amount of power to fully emulate the mechanical behavior (or mechanical behavior) of a human leg. The biomimetic knee presented here utilizes passive dynamics and uses elastic energy storage and tendon-like structure reversion to minimize power requirements. The knee can reproduce the same mechanical characteristics of a knee as a human while walking on a flat ground with low power consumption from an on-board power supply (eg, a power supply on a substrate).

  In one aspect, the present invention is an artificial knee (a knee prosthesis) including an actuator muscle-antagonist muscle structure (or an actuator-antagonist structure, hereinafter the same) of two series elastic actuators arranged in parallel. is there. This knee prosthesis design was motivated by a variable impedance knee prosthesis model that included two series elastic clutch mechanisms and a variable damper. Using human gait data, the model joints were controlled to move biologically. The parameters of the model are obtained using an optimization scheme that minimizes the sum of the squares of the differences between the model's knee torque and biological knee (joint) values over time. The optimized values are used to specify the mechanical and finite state control configuration of an artificial knee having an agonist-antagonist muscle structure. Two preferred embodiments have been developed.

  Due to its architecture, the knee according to the invention is controlled to behave as a variable elastic damper during the stance phase of the gait cycle so that it behaves as a series elastic clutch element of the acting muscle-antagonist type. This provides an energy efficient artificial knee device when walking on flat ground. The knee embodiments are fully powered (e.g., motorized) with a series elastic force detection function, and therefore, knee joint torque can be increased from walking up stairs or slopes or standing up from a sitting position. Direct control over high energy cost tasks. Therefore, the knee architecture is designed to accommodate non-conservative mechanical movements and still provide a very efficient walking mode on flat ground.

  In one aspect, the disclosed invention is a knee joint that is rotatable with respect to a prosthetic leg (prosthetic leg) member and that can be coupled to the prosthetic leg member, in parallel with the prosthetic leg member. A series-elastic flexion actuator connected to the knee joint and a series-connected to the knee joint in parallel with the prosthetic leg member on the opposite side of the prosthetic leg member as viewed from the bending actuator Powered with an elastic extension actuator and a controller for independently supplying energy (electric power) to the flexing motor and the extension motor at different times to control the movement of the knee joint and the combined prosthetic leg member The bending actuator is for bending a prosthetic leg member by applying a force for rotating the knee joint, and the extension actuator is used for the knee joint. The prosthetic leg member is extended by applying a force to rotate the arm. The bending actuator includes a series connection of a bending motor and an elastic element for bending, and the extension actuator includes a series connection of an extension motor and an elastic element for extension. In a preferred embodiment, at least one sensor is used to provide feedback to the controller. The sensor is preferably an angular displacement and acceleration of the knee joint, torque at the knee joint, compression of the elastic element for bending, compression of the elastic element for extension, rotation of the bending motor, rotation of the extension motor, and / or Sensors that respond to contact with the walking surface are included, but are not limited to these.

  In a preferred embodiment, the bending actuator and the extension actuator are unidirectional, and the bending elastic element and the extending elastic element are series springs. In another preferred embodiment, the bending actuator is unidirectional, the extension actuator is bidirectional, the bending elastic element is a series spring and the extension elastic element is a set of Pre-compressed series springs (pre-compressed series springs). In another preferred embodiment particularly suitable for variable speed walking, the flexural elastic element is a non-linear softening spring and the extension elastic element is a non-linear hardening spring. .

  Other aspects, advantages and novel features of the invention will become more apparent from the following detailed description of the invention and the accompanying drawings.

A visual representation of typical knee biomechanical characteristics when walking on a horizontal ground, with a subject whose knee joint has not been damaged walking on the horizontal ground at a self-selected speed FIG. 5 is a plot of the knee joint angle, torque and power curves against the percentage (%) of the walking cycle. FIG. 7 is a diagram visually showing typical biomechanical characteristics of a knee when walking on a horizontal ground, and plotting knee joint torque against knee joint angular positions indicating five walking stages (walking periods). It is a thing. 2 is an exemplary embodiment of a variable impedance artificial knee model in accordance with one aspect of the present invention. FIG. 2B is a plot of optimization results from the model of FIG. 2A versus the biological torque data from FIGS. 1A and 1B. 1 is a schematic mechanical diagram of a powered agonist-antagonist knee according to one aspect of the present invention. FIG. 1 is a side view of a structural design (mechanical design) of an exemplary embodiment of an active knee prosthesis according to the present invention. FIG. 2 is a cross-sectional view in the sagittal plane of the structural design (mechanical design) of an exemplary embodiment of an active knee prosthesis according to the present invention. FIG. FIG. 5 is a rear view of a structural design (mechanical design) of an exemplary embodiment of an active knee prosthesis according to the present invention. 4D is an exploded view of the major components of the exemplary embodiment of the active knee prosthesis of FIGS. 4A-4C. FIG. 4D is another exploded view of the major components of the exemplary embodiment of the active knee prosthesis of FIGS. 4A-4C. FIG. Schematic diagram of a finite state controller for walking on a horizontal ground implemented to reproduce the behavior of an undamaged knee shown in FIGS. 1A and 1B in accordance with one aspect of the present invention. It is. FIG. 7A visually shows the transition of the finite state control of the knee when the limb amputee is walking on the horizontal ground. A through E show the results obtained from the preliminary walking evaluation of the powered prosthesis when walking on the horizontal ground at a self-selected speed in accordance with one aspect of the present invention. FIG. 6 is a side view of the structural design (mechanical design) of another exemplary embodiment of an active knee prosthesis according to the present invention. FIG. 6 is a cross-sectional view in the sagittal plane of the structural design (mechanical design) of another exemplary embodiment of an active knee prosthesis according to the present invention. FIG. 7 is a rear view of the structural design (mechanical design) of another exemplary embodiment of an active knee prosthesis according to the present invention. 9C is an exploded view of the major components of the exemplary embodiment of the active knee prosthesis of FIGS. 9A-9C. FIG. FIG. 4 shows the result of an optimized nonlinear polynomial fitting for force versus displacement behavior of an active knee series elastic element according to the present invention. FIG. 3 illustrates an exemplary embodiment of a variable impedance artificial knee model for variable speed walking in accordance with one aspect of the present invention. It is the graph which showed the output torque of an artificial knee model compared with biological knee (joint) torque data about the 1st walking speed using the model of Drawing 11B. It is the graph which showed the output torque of an artificial knee model compared with biological knee (joint) torque data about the 2nd walking speed different from the 1st walking speed using the model of Drawing 11B. 11B is a graph showing the output torque of the artificial knee model compared with biological knee (joint) torque data for a third walking speed different from the first and second walking speeds using the model of FIG. 11B. .

  A variable impedance artificial knee according to the present invention has two series elastic actuators arranged in parallel in an agonist-antagonist muscle configuration. The artificial knee model includes a variable damper and two in-line elastic clutch units that straddle the knee joint. The variable impedance control arrangement provides the same mechanical characteristics as a human knee during steady walking on a horizontal ground. Due to the variable impedance characteristics of the artificial knee, it is possible to moderately reduce the required power during walking, thereby providing an energy-efficient powered knee (for example, an electric knee). In one application, the variable impedance knee prosthesis according to the present invention has the advantage that it can be used as part of a biomimetic robot leg that does not require tethering (connection to the outside for control or operation).

  As used herein, the following terms expressly include, but are not limited to the following:

  “Actuator” means a motor of the type described below.

  “Agonist muscle” means a contractile element that is resisted or counteracted by another element, an antagonist muscle.

  “Actuator-antagonist actuator” means a mechanism (device) that comprises (at least) two actuators that operate in opposition to each other, where one actuator causes two elements to interact with each other when energized. An actuating muscle actuator that attracts, another actuator is an antagonist actuator that acts to pull the two elements apart when energy is supplied.

  “Antagonist muscle” means an inflatable element that is resisted or counteracted by another element, the working muscle.

  “Biomimetic” means an artificial structure or mechanism that mimics (imitates) the characteristics and behavior of biological structures such as joints and limbs and biological mechanisms.

  “Dorsal flexion” means bending the ankle joint so that the tip of the foot moves upward.

  "Elasticity" means the ability to recover to its original shape after deformation due to stretching or compression.

  “Extension” means a bending motion around the joint of the limb that increases the angle between the bones of the limb at the joint.

  “Bend” means a bending motion around the joint of the limb that reduces the angle between the bones of the limb at the joint.

  “Motor” means an active element that creates or imparts motion by converting supplied energy into mechanical energy (or mechanical energy), and includes an electric motor / electric actuator, a pneumatic motor / pneumatic actuator, Or it includes a hydraulic motor / hydraulic actuator.

  “Planar flexion” means bending the ankle joint so that the tip of the foot moves downward.

    “Spring” means an elastic device such as a metal coil or leaf structure that returns to its original shape after compression or extension.

  The biomechanics of the human knee joint when walking on a level ground. 1A and 1B are diagrams representing the biomechanical characteristics of a typical human knee when walking on a horizontal ground. In FIG. 1A, the power curve of the knee (joint) of a male subject (weight is 81.9 kg) is self-selected on the horizontal ground (1.31 m / power curve). It is plotted against the rate (%) of the walking cycle when walking in seconds. Average values (solid line, N = 10 walking experiment) and 1 standard deviation (dashed line) are plotted. In FIG. 1B, the knee joint torque is plotted against the angular position of the knee joint showing five walking phases. Important gait events that separate the five gait phases are heel contact (HS) 105, stance phase maximum flexion (MSF) 110, stance phase maximum extension (MSE) 115, toe off (TOe off: TO) 120 and the maximum flexion (MWF) 125 in the swing phase.

  As can be seen from FIG. 1A and FIG. 1B, the biomechanical characteristics (biomechanics) of the knee (joint) when walking on the horizontal ground using five distinct stages, ie the gait phase, are described. Yes. These walking steps will be described next.

  (1) Starting from heel contact (HS) 105, the knee (joint) in the stance phase begins to bend slightly (about 15 °). It is possible to absorb the impact at the time of the collision (to the ground) by the stage of the stance stance flexion 130. During this phase, the knee (joint) can be modeled as a spring that stores energy in preparation for the Stance Extension phase 135 (linear torque versus tilt angle; see FIG. 1B).

  (2) After reaching maximum flexion (MSF) 110 in the stance phase, the knee joint begins to extend (at 15% of the gait cycle) and reaches the maximum stance phase (at 42% of the gait cycle). Reach extension (MSE) 115. This knee (joint) extension period is referred to as the stance extension stage 135. During the stance extension 135, the knee (joint) operates as a spring with a stiffness similar to that of the stance phase flex 130 (linear torque versus tilt angle; see FIG. 1B). Here, stiffness is not actual joint stiffness, but quasi-static stiffness defined as the slope of the torque versus angle curve.

  (3) During the late stance phase, that is, during the pre-swing period (Pre-Swing) 140 (42% to 62% of the walking cycle), the knee (joint) of the supporting leg is in the swing flexion stage 145. In preparation, it begins the period when it bends rapidly. During the front swing phase 140, when the knee (joint) begins to flex in preparation for the toe off ground 120, the knee (joint) acts like a spring (linear torque versus tilt angle; see FIG. 1B). However, the rigidity is smaller than that in the stance phase bending 130 and the stance phase extension 135.

  (4) When the hip joint is bent, the leg is separated from the ground, but the knee (joint) continues to bend. At the time of the toe off 120, the swinging leg stage flexing stage 145 of walking starts. Over this period (62% to 73% of the walking cycle), the power of the knee (joint) is generally negative because the knee (joint) torque slows down the rotational speed of the knee (joint). (See FIG. 1A). Accordingly, during the swing leg flexion 145, the knee (joint) can be modeled as a variable damper.

  (5) When the maximum flexion angle (about 60 °) is reached during the swing period 145, the knee (joint) starts to extend forward. During Swing Extension 150 (73% -100% of the walking cycle), knee (joint) power is generally negative to slow down the free leg for the next stance phase. . Therefore, the knee (joint) during the period of free leg extension 150 can be modeled as a variable damper in the same way as the free leg flexion 145. When the knee (joint) is fully extended, the foot is again placed on the ground and the next walking cycle begins.

  Quasi-passive artificial knee model and optimization. Given the described knee (joint) biomechanical characteristics, it is possible to generate knee mechanical characteristics similar to humans during steady walking on a horizontal ground, with damping ( It is clear that there is a need for a variable impedance knee prosthesis that can change both braking and stiffness. Examples of such prostheses include two antagonizing single-joint series elastic clutches 205, 210 (for modeling knee mechanics during stance) and (for modeling mechanics during swing) FIG. 2B is an exemplary knee model shown in FIG. 2A with one variable damping element 215. In this model, each of the series springs 220, 225 can be engaged (or connected; the same applies hereinafter) by actuating the respective clutch 205, 225, and can be disconnected by opening the clutch. . Due to model constraints, each clutch can only be connected once during each walking cycle. In addition, when the clutch is connected, the series spring releases all of its energy and the clutch can only be disconnected when the force on the clutch is zero.

であった。ここで、τⁱ _bioとτⁱ _simは、それぞれ、生物学的トルクデータと膝モデルから得られる、歩行サイクルのｉ番目の割合（％）において膝関節のまわりに加えられる角トルク（angular torque）であり、τ^max _bioは、歩行サイクル中の関節における最大の生物学的トルクである。費用関数（１）は、伸筋ばねが、踵接地時（θ_E＝０）に常に係合するという制約によって最小化された。この制約は、肢喪失者の安全対策として、踵接地時における膝（関節）の曲がりを制限するために適用された。 The parameters of the series elastic clutch model were changed to match the biomechanical behavior of the knee joint. The model parameters are two spring constants (k _E , k _F ) corresponding to the stiffness of the extension spring and the stiffness of the flexion spring, and the relative knee (joint) extension angle and knee (joint) flexion angle ( θ _E and θ _F ), and with such parameters, the extension spring and the bending spring engage during the stance phase. Traditionally, extension springs (also called extensor springs) tend to extend (extend) the knee joint when engaged, while flexure springs (also called flexor springs) bend the knee (joint) ( Tend to bend). The knee model was fitted to the biomechanical data using an optimization scheme that minimizes the sum of the squared differences between the model's knee and biological knee torque values over time. More specifically, the cost function used for optimization is

Met. Where τ ⁱ _bio and τ ⁱ _sim are angular torques applied around the knee joint at the i-th percentage of the walking cycle, obtained from biological torque data and the knee model, respectively. Τ ^max _bio is the maximum biological torque at the joint during the gait cycle. The cost function (1) was minimized by the constraint that the extensor spring always engages when the heel is in contact (θ _E = 0). This restriction was applied to limit the bending of the knee (joint) during heel contact as a safety measure for people with limb loss.

  The determination of the desired global minimum of the cost function (1) is first performed using a genetic algorithm to find the region containing the global minimum, and then the constraint (or The exact value of the global minimum was determined by a gradient optimizer without the condition. After optimization of the cost function (1) by changing the parameters of the series elastic clutch element, the artificial knee model in the region where the series elastic element cannot sufficiently absorb the negative mechanical force using the variable damper of the model The torque value and the biological torque value were completely matched. Biological knee (joint) torque values are dynamics and kinematics obtained from 10 walking experiments on a healthy subject with a weight of 81.9 kg and a height of 1.87 m walking at 1.31 m / sec. Obtained from inverse dynamics calculation using dynamic data.

In FIG. 2B, the optimization results obtained from the model of FIG. 2A are plotted against the biological torque data from FIGS. 1A and 1B. As shown in FIG. 2B, in the upper graph 245, the optimized net torque output 250 of the knee model is compared to the torque profile of an undamaged human knee joint, averaging 255 And 1 standard deviation 260, 265 are shown (N = 10 walking experiment). The biological data taken from FIGS. 1A and 1B is for a subject (body weight = 81.9 kg) whose limb walking at a self-selected speed (walking speed = 1.31 m / sec) is not damaged. . The lower graph 270 shows the torque contribution from the series elastic clutch element extension spring 275 and bending spring 280 and variable damper 285. The optimization tool is for extension spring stiffness equal to k _E = 160 Nm / radian, bending spring stiffness equal to 137 Nm / radian, and for bending equal to 0.27 radians (15.46 °) Gives the knee engagement angle (knee engagement angle).

  The torque output of the model agrees well with the experimental value. As constrained by the optimization procedure, the extension springs engage during heel contact and store energy during flexion of the knee (joint) in the early stance phase. As the knee (joint) begins to stretch, the flexing spring engages and accumulates energy as the extending spring releases its energy. During the swing phase, the model's variable damper exactly matches the biological torque value in the negative power region. In the middle and end of the swing phase, power is positive in biological data, so the dampers output zero torque in their walking area.

  A simplified mechanism diagram of an agonist-antagonist knee according to one aspect of the present invention is shown in FIG. As shown in FIG. 3, an artificial knee 300 having an agonist-antagonist action includes two unidirectional series elastic actuators, a series elastic extension actuator 305, and a series elastic bending actuator 310. . Each of the unidirectional actuators 305 and 310 of the artificial knee 300 includes motors 315 and 320 and series springs 325 and 330 connected via transmissions (power transmission devices) 335 and 340. The extension motor 315 and the bending motor 320 can be used independently to control the knee (joint) angle at which the series springs 325 and 330 are engaged. The knee joint 345 between the upper part 350 and the lower part 355 of the leg is connected by a cable-driven transmission to a linear carriage (cart) that can move freely along the longitudinal direction of the device. The carriage can be engaged on either side by an extension spring 325 and a bending spring 330 positioned by a ball screw driven by an electric motor. In a preferred embodiment, both unidirectional series elastic actuators 305, 310 feature transmissions 335, 340 composed of a 2: 1 belt drive coupled to a ball screw (Nook industries, 10 × 3 mm). To do.

  The principles of the exemplary knee model of FIG. 2A are embodied in two exemplary preferred physical embodiments. In one preferred embodiment, as shown in FIGS. 4A-4C, an active knee prosthesis that performs an agonist-antagonist-like operation comprises two unidirectional serial elastic actuators. 5 and 6 are exploded views of the major components of the exemplary embodiment of the active knee prosthesis of FIGS. 4A-4C.

  As shown in FIGS. 4A to 4C, 5, and 6, the unidirectional actuators are the extension actuator 402 and the bending actuator 404. The artificial knee extension actuator 402 includes an extension motor 406 and a series spring 408 connected via a transmission. The extension actuator is near (eg, adjacent to) the knee joint 410. The extension transmission is comprised of a timing pulley 412 and a timing belt (or belt) 414 drive system coupled to a precision ball screw 416 drive. The artificial knee bending actuator 404 includes a bending motor 418 and a series spring 420 connected via a transmission. The bending transmission is comprised of a timing pulley 422 and timing belt (or belt) 424 drive train coupled to a precision ball screw 426 drive. The extension actuator 402 and the bending actuator 404 can be used independently to control the angle of the knee joint 410 that engages each of the series springs 408, 420.

  Powered element of series elastic extension actuator 402 (extension electric motor 406), brushed DC motor (such as Maxon RE40 motor) or brushless DC motor (such as Maxon EC-powermax 30) It can be. This extension motor directly drives the timing pulley-belt drive 412, 414 mechanism. This mechanism has a 1: 2 speed transmission ratio (transmission ratio). Timing pulley-belt drive mechanisms 412, 414 rotate a ball screw 416 (such as Nook Industries, 10 x 3 mm). When the ball screw 416 of the extension actuator 402 rotates, the coupled ball nut housing 428 is linearly displaced. The ball nut housing 428 is directly attached to the extension series spring cage 430. The extension series spring cage 430 safely accommodates the extension spring 408. Thus, when the coupled ball nut housing 428 is linearly displaced, the extension series spring cage can produce a linear displacement. The ball nut housing moves along two straight precision steel guide rails 432 with minimal friction due to the linear bearings incorporated in the ball nut housing. Each of the two precision guide rails 432 is attached to a corresponding internal sidewall of the knee main frame (knee body) 434.

  Powered elements of the series elastic bending actuator 404 (electric motor 418 for bending), brushed DC motors (such as Maxon RE40, RE30 motors) or (such as Maxon EC-powermax 30 and 22) ) Can be a brushless DC motor. This bending motor directly drives the timing pulley-belt drive 422, 424 mechanism. This mechanism has a 1: 2 speed transmission ratio (transmission ratio). The pulley-belt drive mechanisms 422, 424 rotate a ball screw 426 (such as Nook industries, 10 × 3 mm). When the ball screw 426 of the bending actuator 404 rotates, the coupled ball nut housing 436 is linearly displaced. The ball nut housing 436 is directly attached to the bending series spring cage 438. The bending series spring cage 438 safely houses the bending spring 420. Thus, when the combined ball nut housing 436 is linearly displaced, the bending series spring cage can be linearly displaced. The ball nut support 436 moves along two straight precision steel guide rails 432 with minimal friction due to the linear bearings incorporated in the ball nut housing. Each of the two precision guide rails 432 is attached to a corresponding internal sidewall of the knee main frame (knee body) 434.

  The rotation of the knee joint 410 is coupled to the linear displacement of the linear carriage 442 attached to the cable driven transmission. The cable driven transmission is composed of two steel cables (steel cables) 440. The two ends of each steel cable 440 are attached to the knee joint U link support 444. Each cable wraps around a corresponding joint pulley 446 located on each side of the knee. Each of the side joint pulleys 446 has a shaft attached to the lower adapter 448. The linear carriage 442 is supported by two steel precision rail guides 432 and guided on the guides. Each rail extends along the inner side wall of the knee main frame 434. Low friction between the precision rail and the linear carriage 442 is obtained by a linear bearing built into the linear carriage 442. The steel cable 440 allows the linear displacement of the linear carriage 442 to be coupled to the rotational movement of the knee joint 410.

  The linear carriage 442 can be individually engaged using the linear motion of the extension spring 408 and the bending spring 420. Both springs are positioned independently by the drive action of their corresponding series elastic actuators. Each of the series elastic actuators 402, 404 can provide sufficient power for both walking on a horizontal ground and tasks that require more energy, such as climbing stairs. .

  All actuation mechanisms are fully supported by an aluminum structure corresponding to the knee main frame 434 having a structure resembling the anatomical skin of the lower limb. The main frame of the knee has a lower adapter 448 that allows a conventional high performance robotic short leg brace to be attached to the frame. By connecting the U-link support 444 of the knee joint and the main frame of the knee, a degree of freedom of rotation corresponding to the knee joint 410 is created. The knee joint U-link support 444 allows attachment of a standard prosthetic pyramid adapter 450 so that the prosthetic knee can be attached to a normal thigh socket and worn by a person with a limb loss It becomes possible to do.

Table 1 shows the design parameters of the knee (joint) size, angle range (unit deg = degree), and maximum torque value of the embodiments of FIGS. 4A to 4C, 5 and 6 based on such functional requirements. It is described in.

Real-time feedback information to the onboard controller is provided by onboard specific sensors. These sensors utilized by the embodiments of FIGS. 4A-4C, 5 and 6 are listed in Table 2.

  The angular displacement of the knee (joint) is indirectly measured by a digital linear encoder 452 fixed to the left outer surface of the main frame 434 of the knee. The rotational movement of the knee joint 410 is coupled to the displacement of the linear carriage 442 via the steel cable 440 transmission. The degree of compression of each of the series elastic actuator springs 408, 420 is measured by corresponding Hall effect sensors 454, 456. These sensors measure the change in magnetic field that occurs when the proximity of the magnets attached to each of the spring cages 430, 438 changes while the springs 408, 420 are compressed. The rotation of each motor is measured by a motor digital encoder 458, 460 attached to the back of the corresponding motor 406, 418. Interaction with the ground is measured with a force sensitive footpad (not shown). With this foot pad, it is possible to detect when the leg is on the walking surface and when it is away from the walking surface, and the controller determines which walking stage the robot artificial knee worn by the user is in. To help.

  All electronic circuits are implemented in a single on-board microcontroller-based system. The motor is driven by an H-bridge controller, its speed is determined by 20 kHz pulse width modulation (PWM), and the power is DC (such as a 6-cell lithium polymer battery (rated 22.2V)) Supplied by a battery (DC battery). The analog sensor was read by a 10-bit analog-to-digital converter (ADC). The system is controlled by an on-board microcontroller (such as AVR) and can be monitored by USB and / or Bluetooth. All processing is performed onboard and power is supplied by a relatively small battery (mass = 0.15 kg). The prototype (prototype) is completely self-contained and does not require tethering (an external connection for control or operation).

  Finite state control method. A finite state controller for walking on a horizontal ground was implemented to reproduce the behavior of an intact knee (joint) shown in FIGS. 1A and 1B. This state machine is shown in FIG. 7A. Three states are shown with control action and transition conditions. The three states of the controller are a standing stance initial state (Early Stance: ES) 710, a front swing leg state (Pre-Swing: PW) 720, and a swing leg state (SW) 730. A quasi-passive equilibrium point control was implemented for states ES 710 and PW 720, and variable damping control was utilized during SW state 730. Transitions between states were mainly determined by three measurements: heel-to-ground contact, toe (toe) -to-ground contact, and knee (joint) angle.

For state transition and its identification, the system used the following variables.
踵 Contact (H). H = 1 indicates that the heel is in contact with the ground, and H = 0 indicates a state where the heel is away from the ground.
Toe contact (T). T = 1 indicates that the toes are in contact with the ground, and T = 0 indicates that the toes are away from the ground.
The knee (joint) angle (θ) is the relative angle of the knee joint. All knee angles are flexion angles. The angles θ _E and θ _F define angles at which the extension spring and the bending spring are engaged during the stance phase, respectively. Further, θ ₊ is the angle of the knee joint when flexing during the swing phase, and θ ₋ is the angle of the knee joint when extending during the swing phase.

Next, each state will be described. The stance initial state (ES) 710 starts from heel contact (HS). When the heel touches the ground (H = 1), the extension motor shaft is locked using high PD gain control, in which case the desired axial speed is zero. Next, the extension spring engages at a spring equilibrium angle θ _E equivalent to the knee position at the time of heel contact. Next, in preparation for the extension of the knee (joint), the extension spring stores energy while the knee (joint) is bent in the initial stage of the stance. While the knee (joint) is flexing, the flexion spring equilibrium point θ _F uses position control to accurately track the linear carriage coupled to the knee output joint. Servo-controlled (servo control is control by a servo mechanism). When the knee (joint) is bent to the maximum in the initial state of the stance, the knee (joint) starts to extend and the shaft of the bending motor is locked. Next, the bending spring is engaged at a spring equilibrium angle θ _F equivalent to the position of the knee (joint) when the knee (joint) has the maximum bending. The energy is first released from the extension spring and then stored in the bending spring.

The front swing leg state (PW) 720 starts when the heel leaves the ground (H = 0) and the knee (joint) angle decreases (θ <3 °). In this state, the balance point θ _E of the extension spring is position controlled under zero load and precisely tracks the linear carriage as the knee (joint) bends in preparation for the swing phase. The bending spring maintains its current equilibrium position θ _F (the motor shaft remains locked). Therefore, when the knee (joint) is bent over the entire PW, the energy stored in the bending spring is released.

The free leg state (SW) 730 starts with the toe off (T = 0). The balance angle θ _E of the extension spring continues to track the knee (joint) angle under no load. When the knee (joint) bends beyond 20 ° (θ ₊ > 20 °), the bending spring's equilibrium angle θ _F is servo controlled under no load to a position corresponding to θ = 15 °. . When the knee (joint) bends beyond 60 ° (θ ₊ > 60 °), the extension spring engages. The low gain damping control for the extension motor shaft backdrives the extension motor and transmission to act as a variable damper to reduce knee (joint) hyperflexion. When the knee (joint) begins to extend and reaches an angle less than 60 ° (θ ₋ <60 °), the balance angle θ _E of the extension spring again tracks the linear carriage under no load. Follow the middle knee (joint) while extending. When the knee (joint) continues to extend beyond 15 ° in the late swing phase (second half of the swing phase), the bending spring engages. Also in this case, the low-gain damping control for the bending motor shaft back-drives the bending motor and the transmission to smoothly decelerate the free leg until the bending spring equilibrium angle θ _F reaches 3 °. Operate as a variable damper. When the knee (joint) extends beyond 5 ° (θ ₋ <5 °) during the swing leg, the balance angle θ _E of the extension spring is used in preparation for engagement and energy storage at the time of heel contact in the subsequent walking cycle. Servo controlled to 3 °.

  A finite state control diagram showing the transition of the state machine of FIG. 7A is shown in FIG. 7B. The graph of FIG. 7B illustrates the state transition of the controller for three consecutive walking cycles on the horizontal ground according to the control scheme implemented in the embodiments of FIGS. 4A-4C, 5 and 6. is doing. The control states in the case of walking on the horizontal ground are the stance initial state (Early Stance: ES = state 1) 750, the front swing leg state (Pre-Swing: PW = state 2) 760, and the swing leg state ( SW = state 3) 770 is defined. The system went through state sequence 1-2-3 (ES-PW-SW) in each walking cycle. The controller steadily transitioned to state 1-2-3 through three consecutive walking cycles.

  The results obtained from the preliminary walking evaluation of the powered prosthesis during walking on the horizontal ground at a self-selected speed (0.81 m / sec) are shown in FIGS. The knee joint angle of the artificial knee (A in FIG. 8), the net torque (B in FIG. 8), and the power (C in FIG. 8) are walking on the flat ground (weight of the subject with limb loss = 97 kg, walking speed) = 0.81 m / sec) is plotted against the rate of walking cycle (% for the walking experiment with N = 10, the mean is shown as a solid line and one standard deviation is shown as a dashed line). In FIG. 8D, the torque contribution from the extension spring (lower line) and the bending spring (upper line) of the series elastic actuator is plotted. E in FIG. 8 plots the knee joint torque against the knee joint angular position and shows two characteristic stiffnesses during the stance phase.

  These artificial knee values are qualitatively consistent with those of an intact knee mechanism shown in FIG. 1A that has not lost a limb whose weight and height are in harmony. The artificial knee is similar to the kinematics (target characteristic) of an intact knee shown in FIG. 8A. In the early stage of stance, the knee (joint) is flexed (maximum). Bending angle of about 14.5 °). At the end of stance, the artificial knee performs rapid knee (joint) flexion in preparation for the swing phase. The knee bends to a bending angle of about 61 ° at the maximum during the swing phase, and then extends forward before heel contact. As shown in FIGS. 8B and 8C, the knee joint torque and power (power) of the artificial knee are negative during knee flexion in the initial stage of stance. In the undamaged knee (joint) data shown in FIG. 1A, knee joint torque and power are initially positive after heel contact but rapidly become negative as the knee continues to flex. As shown in FIG. 8B, the artificial knee showed a similar behavior at the start of the initial stance. The reason for this behavior was that a person with a limb lost his knees first when he touched the heel and engaged the bending spring, but immediately continued to bend the knee and engaged the extension spring. is there. Compressing the extension spring immediately after heel contact limits the knee so that it does not suffer excessive buckling, thereby ensuring greater safety for those with limb loss It was helpful. During the extension of the knee (joint) in the stance phase, the torque is initially negative and then positive, and the knee (joint) power is initially positive and then negative for both intact and artificial knees. become. Furthermore, during the previous swing phase, torque and power (power) are initially positive and then negative in preparation for the swing phase. During the swing phase, for both undamaged knees and artificial knees, power limits the maximum flexion (negative torque) of the knee (joint) and then during extension during the swing phase (positive torque) Generally, it is negative to smoothly slow down the free leg.

  In FIG. 8D, the torque contribution from each of the unidirectional series elastic actuators is plotted against the walking cycle percentage. The flexion spring is temporarily engaged immediately after heel contact due to the short extension of the knee (joint) of the limb-losing person. As soon as knee flexion occurs in the stance phase, the extension spring engages in the same manner as the artificial knee model. During flexion in the stance phase, the flexion spring loses its energy rapidly when closely tracking a linear carriage coupled to the knee output joint. Thereafter, the bending spring re-engages when the knee (joint) has a maximum flexion during the initial stage of stance, and stores energy during extension in the stance phase and during the forward swing phase. In the swing phase, the extension motor when the knee flexion is maximum, and the flexion motor at the end of stance limit the maximum flexion of the knee and smoothly decelerate the swing leg at the end of stance. Each is actively backdriven. In FIG. 8E, the knee (joint) torque of the artificial knee is plotted against the angular position of the knee (joint). Like an undamaged human knee (FIG. 1B), an artificial knee has two characteristic stiffnesses during the stance phase. The stiffness of the knee (joint) during the flexion and extension stages in the initial stage of stance is greater than that during the front swing stage.

  The artificial knee embodiments of FIGS. 4A-4C, 5 and 6 utilize an integrated series elastic element in combination with variable impedance control to minimize electrical energy during walking. According to this method, knee behavior is matched between an artificial knee and an intact knee mechanism. When the mechanical design architecture and control are combined, the knee prosthetic motor does not perform positive work on the knee joint during walking on the horizontal ground, resulting in moderate power requirements.

  The artificial knee (artificial knee) variable impedance control during walking on the horizontal ground, the required power of the artificial knee is small during steady walking experiments with an average walking speed of 0.81 m / sec (8 Watts electrical) ). Researchers have used a step count monitor device to find that an active person who has lost one leg walks 3060 ± 1890 steps a day. To estimate the required power of the motor (power (current) = current x voltage), measure the current of each motor directly using on-board motor current detection, as well as estimate the motor voltage, motor speed, motor speed Constants (motor speed constant) and motor resistance values were used. Assuming that a person who has lost a limb walks 5000 steps at a moderate walking speed, the size of the on-board battery can be estimated. For example, a 0.13 kg lithium polymer battery (with an energy density of 165 watt hours / kg) would allow 5000 steps of powered walking (8 watts x 1.95 seconds / cycle x 5000 cycles = 78 kg). Jules). This battery mass is 1/5 or less of that required for other commercially available powered knees.

  In a second preferred embodiment shown in FIGS. 9A-9C, the active knee prosthesis comprises two actuators arranged to form an agonist-antagonist architecture. FIG. 10 is an exploded view of the major components of the exemplary embodiment of the active knee prosthesis of FIGS. 9A-9C. In this embodiment of an active knee prosthesis according to the present invention, one of the two actuators is an extension series elastic actuator 902 and the other is a bending actuator 904. The extension actuator 902 is bidirectional, and the bending actuator 904 is unidirectional. An extension actuator 902 near the knee joint 906 of the knee prosthesis is comprised of an extension motor 908 connected through a transmission and a set of pre-compressed series springs 910. The extension transmission consists of a timing pulley set 912 and a belt 914 drive system coupled to a precision ball screw 916 drive. An artificial knee unidirectional bending actuator 904 includes a bending motor 918 and a series spring 920 connected via a transmission. The bending transmission is comprised of a timing pulley set 922 and a belt 924 drive system coupled to a lead-screw 926 drive. The extension actuator 902 and the bending actuator 904 can be used independently to control the angle of the knee joint 906 with which the series springs 910, 920 can be engaged.

  The electric motor 908 of the extension actuator 902 can be a brushed DC motor (such as Maxon's RE40 motor) or a brushless DC motor (such as Maxon's EC-powermax30). This extension motor directly drives the timing pulley-belt drive 914, 928 mechanism. This mechanism has a 1: 2 speed transmission ratio (transmission ratio). Timing pulley-belt drive mechanisms 914, 928 rotate a ball screw 916 (such as Nook Industries, 10 × 3 mm). When the ball screw 916 of the extension actuator 902 rotates, the combined ball nut housing 930 is linearly displaced. The ball nut housing 930 is directly attached to the extension series elastic spring cage 932. The series elastic spring cage 932 for extension secures a set of springs 910 consisting of two identical pre-compressed passive mechanical springs (the stiffness of which is the same as that of the model extension actuator). To house. Therefore, when the coupled ball nut housing 930 is linearly displaced, the extension series elastic spring cage 932 is linearly displaced. The ball nut housing 930 moves along two straight steel guide rails 934. Each of these two rails is attached to a corresponding side wall 936 of the housing. The spring cage 932 moves along a guide rail supported by roller bearings 938 incorporated therein. Because the extension actuator 902 is attached to the steel cable drive system 940, it is directly coupled to the rotational movement of the knee joint 906.

  The electric motor 918 of the bending actuator 904 may be a brushed DC motor (such as Maxon's RE40, RE30 motor) or a brushless DC motor (such as Maxon's EC-powermax 30 or 22). This bending motor directly drives the timing pulley-belt drive 922, 924 mechanism. This mechanism has a 1: 2 speed transmission ratio (transmission ratio). Pulley-belt drive mechanisms 922, 924 rotate a lead screw 926 (such as Nook Industries, 10 × 3 mm). When the lead screw 926 of the bending actuator 904 rotates, the combined ball nut housing 942 is linearly displaced. The ball nut housing 942 is directly attached to the bending series spring cage 944. The bending series spring cage 944 securely houses the bending spring 920. Therefore, when the coupled ball nut housing 942 is linearly displaced, the bending series spring cage can be linearly displaced. Ball nut housing 942 moves along two straight steel guide rails 934 with minimal friction due to rollers 938 incorporated in the ball nut housing. The flexing actuator 904 is not directly coupled to the rotational motion of the knee joint 906, but can bend during operation and backdrives the series elastic spring cage 932 for extension.

  The knee joint 932 is coupled to a pair of two steel cable drives 940 that are connected to an extension series elastic spring cage 932. The series elastic cage is supported and guided by two precision guide rails 934 made of steel. Steel cable 940 allows the linear displacement of series elastic spring cage 932 to be coupled to the rotational movement of knee joint 906.

  The two ends of each of the steel cables 940 are attached to a knee drive hub 944. Drive hub 944 is supported by knee joint housing 946. Each cable wraps around a corresponding joint pulley 948 located on each side of the knee remote from the knee joint 906. Each of the side joint pulleys 948 has a shaft attached to a corresponding side wall 936. Each cable drive 940 can be tensioned independently by adjusting the corresponding cable tensioner 949 and adjusting the side joint pulley 948. Each of the series elastic actuators 902, 904 can provide sufficient power for both walking on a horizontal ground and tasks that require more energy, such as climbing stairs. .

  All actuation mechanisms are fully supported by an aluminum structure corresponding to the knee sidewall 936, the upper pyramid adapter 928, and the lower pyramid adapter 950 assembly. This structure provides a support frame that resembles the anatomical skin of the lower limb. The lower pyramidal adapter 950 allows a conventional high performance robotic short leg brace to be attached to an artificial knee. A standard prosthetic upper pyramid adapter 928 allows the prosthetic knee to be attached to a normal thigh socket and worn by the thigh amputee. The prosthetic knee configuration has removable side and front covers that allow easy access to the drive actuators and mechanisms, thus facilitating maintenance of the prosthetic knee.

Table 3 shows design parameters of the knee (joint) size, angle range (unit: deg = degree), and maximum torque value of the embodiment of FIGS. 9A to 9C and FIG.

これらのセンサーは、歩行中の、膝関節９０６の角変位、各直列弾性要素の変位及び変形、及び、膝と環境（地面）との力／接触相互作用をモニタする。 The knee's inherent sensory system (sensor system) provides feedback to the on-board control electronics. These sensors are listed in Table 4.

These sensors monitor the angular displacement of the knee joint 906, the displacement and deformation of each series elastic element, and the force / contact interaction between the knee and the environment (ground) during walking.

  The angular displacement of the knee is directly measured by an absolute encoder 952 located at the knee joint 906. This sensor 952 is mounted on an encoder housing 953 that interlocks with the rotation of the joint. The degree of compression of the series elastic springs 910, 920 of each actuator is measured by corresponding Hall effect sensors 954, 956. These sensors measure the change in magnetic field that occurs when the proximity of the magnets attached to each of the spring cages 932, 944 changes while the springs 910, 920 are compressed. The rotation of each motor is measured by a motor digital encoder 958, 960 attached to the back of the corresponding motor 908, 918. Interaction with the ground is measured with a force sensitive footpad (not shown). With this foot pad, it is possible to detect when the leg is in contact with the walking surface, and the controller can determine which walking stage the robot artificial knee worn by the user is in. Another method used to measure the interaction with the ground is the strain that is mounted on the knee frame (specifically, in the shin cover 962) and penetrates the attached lower pyramid adapter 950. Is to use a gauge. This sensor information provides information on the forces and torques that interact with the knee and can provide information for calculating the torque at the knee joint. This information helps the controller determine the walking stage. Another sensor considered for this embodiment is to use an inertial measurement unit (not shown) attached to the main frame of the knee, which allows the controller to determine orientation and acceleration during the walking cycle. It becomes possible to do.

  A set of electronic devices that provide autonomy (control without tethering) to the artificial knee is built in. The electronic devices are mounted on five printed circuit boards (PCBs) that form a group assembled to the side wall of the knee and the rear wall (back wall). Two of these boards are allocated for actuator control, the other board is responsible for monitoring the entire knee control scheme, and the other board is an external monitoring PC / lap Assigned for connection and communication with the top system, the last one board is responsible for processing the data of the inertial measurement unit installed on the knee. The electronics set is powered by a 6-cell lithium polymer battery with a rated voltage of 22.2V, based on PIC microcontroller technology. The motor board includes a brushless controller capable of controlling up to 20 KHz. System behavior can be monitored and updated using USB or by wireless communication via Wi-Fi interconnection.

  Finite state control method. Using five stages in the walking cycle, a knee control scheme for horizontal ground walking has been established that utilizes a semi-passive knee model to mimic the knee biomechanics with minimal energy consumption.

  The knee in the stance phase starts from the heel contact and begins to bend slightly (about 15 degrees). This stance phase flexion phase enables shock absorption during a collision. During this phase, the knee extension actuator engages its series elastic element by maintaining a spring equilibrium position as energy is stored during knee flexion. In this case, the motor operates as an engagement clutch (or meshing clutch).

  When flexion is maximized in the stance phase, the knee joint begins to extend (with a walking cycle of about 15%) and reaches a maximum extension (MSE) state in the stance phase (with a walking cycle of about 42%). This knee extension period is called the stance extension phase. During the stance extension phase, the flexing actuator positions its series spring so that energy is stored when the knee begins to extend, thereby allowing the energy of the extension spring to be transmitted. This energy transfer makes it possible to adjust the stiffness during the stance phase. Since the bending actuator transmission has a lead screw, as soon as the actuator positions a linear spring, the motor no longer provides positive power, thus minimizing overall energy consumption. It is important to mention that. This effectively uses the clutch engagement behavior of the model in the “normally closed (normally closed)” setting state.

  During the late stance phase or the anterior swing phase (about 42% to about 62% gait cycle), the knee of the support leg begins a period of rapid flexion in preparation for the swing phase. During the front swing phase, the knee begins to flex in preparation for toe separation, and the energy stored in the flexing spring is released to assist the user before toe separation.

  The leg continues off the ground and the knee continues to flex. When the toes leave, the flexion phase of walking begins. Over this entire period (62% to 73% walking cycle), knee power is generally negative because knee (joint) torque slows the rotational speed of the knee. Therefore, during the swing phase flexion phase, the knee extension actuator can be used as a regenerative element (or regenerative element), and when the actuator is back-driven, the energy reduced during the swing phase is turned on. The battery on the board can accumulate and the legs are repositioned to begin a new walking cycle.

  When the maximum flexion angle (about 60 °) is reached during the swing phase, the knee begins to extend forward. During the swing phase extension phase (about 73% to about 100% walking cycle), the knee power is generally negative in order to decelerate the swing leg for the next stance phase. At this stage, the extension actuator is used as an element to provide energy regeneration (or energy regeneration) to improve the energy efficiency of the knee prosthesis. During this phase, the flexing motor repositions the linear spring in preparation for a new walking cycle.

  Artificial knee for variable speed walking. The described active knee prosthesis embodiments are intended to provide the ability to walk around at the K3 level (i.e., the ability to walk or have the potential to vary at various paces) for those who have lost their limbs. In order to provide a knee prosthesis that is adaptable to changes in the speed of the limb loser and that can maintain an optimal energy efficiency level, both series elastic elements should be non-linear, And it is necessary to adjust to the weight of the person with limb loss.

  In order to evaluate speed adaptability, using the optimization method described above, a variable impedance knee model is injured walking at speeds of 1.0 m / sec, 1.3 m / sec, and 1.6 m / sec. Adapted to biomechanical data from subjects who did not receive it. The optimization results resulted in linear stiffness values (at the respective walking speeds) for the series elastic elements as well as the engagement angles of both series elastic elements during the walking cycle. For each spring and walking speed, the maximum force-displacement data pair was first selected. By plotting all data pairs in the same force versus displacement space, it is now possible to estimate the optimal non-linear spring function (or non-linear spring function) for the bending spring and the extension spring. . FIG. 11A is a plot of optimized nonlinear polynomial fitting results for force versus displacement behavior of a series elastic element of an active knee according to the present invention. Specifically, a second order polynomial fitting was performed on the model extension spring 1105 and a piecewise polynomial fitting was performed on the bending spring 1110.

  As suggested by the data in FIG. 11A, in the case of variable speed walking applications, the extension series elastic element of FIG. 2 is replaced with a non-linear hardening elastic element, and the flexing element of FIG. It is necessary to replace the element with a softening spring. FIG. 11B illustrates an exemplary embodiment of a variable impedance artificial knee model for variable speed walking according to one aspect of the present invention. As shown in FIG. 11B, the model includes two single-joint series elastic clutches 1120, 1125, two nonlinear springs 1130, 1135, and a variable damping element 1140.

Using these nonlinear spring functions (or nonlinear spring action), the output torque of the artificial knee model was compared with biological knee (joint) torque data. 11C-11E show this comparison made for three different walking speeds using the model shown in FIG. 11B. In FIGS. 11C-11E, the net torque output of the optimized knee models 1150, 1155, 1160 is compared to the torque profile of an undamaged human knee joint, with an average of 1165 for each speed. Both 1170, 1175 and 1 standard deviation 1180, 1182, 1184, 1186, 1188, 1190 (N = 10 walking experiment) are shown. The goodness of fit of the model is provided by a coefficient of determination R ^{2 at} each walking speed. Biological data are for subjects with no limb injury walking at three different speeds: 1.0 m / sec (FIG. 11C), 1.3 m / sec (FIG. 11D), 1.6 m / sec (FIG. 11E). Data obtained for (body weight = 66.2 kg). For each selected speed, a range of walking speed ± 5% was allowed.

  An artificial knee model for variable speed walking can be physically implemented in a manner similar to the embodiment described for the model of FIG. 2, using components and materials readily apparent to those skilled in the art; And it will be apparent to those skilled in the art that any variations and modifications suitable for the embodiment of FIGS. 4A-4C and 9A-9C are also suitable when implemented using the model of FIG. 11B. I will.

Moreover, while preferred embodiments of the invention have been disclosed, many other examples will occur to those skilled in the art, and all of these examples are within the scope of the invention. Each of the various embodiments described above can be combined with the other described embodiments to provide a number of features. Furthermore, although a number of individual embodiments of the apparatus and method of the present invention have been described herein, what has been described herein is merely illustrative of the application of the principles of the present invention. Accordingly, other sequences (configurations), methods, modifications, and substitutions by those skilled in the art are also considered to be within the scope of the present invention, and the present invention is not limited except by the appended claims. .

Claims (9)

A powered artificial knee,
A knee joint rotatable with respect to the prosthetic limb member and connectable to the prosthetic limb member;
A series elastic bending actuator for bending a prosthetic limb member by applying a force for rotating the knee joint, connected to the knee joint in parallel with the prosthetic limb member, and a bending motor A series elastic bending actuator including a series connection of a bending elastic element and
A series elastic extension actuator for extending the prosthetic limb member by applying a force for rotating the knee joint, connected to the knee joint and on the opposite side of the prosthetic limb member as viewed from the bending actuator A series elastic extension actuator that is in parallel with the prosthetic limb member and includes a series connection of an extension motor and an extension elastic element;
Power comprising a combination of controllers for controlling the position, impedance, and non-conservative torque of the knee joint and the combined prosthetic limb member by independently supplying energy to the flexing motor and the extension motor at different times Artificial knee with.   The knee prosthesis of claim 1, further comprising at least one sensor configured to provide feedback to the controller.   The sensor is responsive to at least one of angular displacement of the knee joint, compression of the elastic element for bending, compression of the elastic element for extension, rotation of the bending motor, rotation of the extension motor. 2 artificial knees.   The artificial knee according to claim 1, wherein the bending actuator and the extension actuator are unidirectional.   The artificial knee according to claim 4, wherein the elastic element for bending and the elastic element for extension are series springs.   The artificial knee of claim 1, wherein the bending actuator is unidirectional and the extension actuator is bidirectional.   The knee prosthesis of claim 6 wherein the bending elastic element is a series spring and the extension elastic element is a set of pre-compressed series springs.   The artificial knee according to claim 1, wherein the elastic element for bending is a non-linear gradually soft spring, and the elastic element for extension is a non-linear gradually soft spring. The knee prosthesis of claim 1, wherein the bending actuator comprises a transmission that cannot be back-driven.

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