Method for controlling prosthetic limbs or orthotics

Abstract

The present invention relates to a method for controlling a prosthetic or orthotic device for a lower limb having an upper part (10) and a lower part (20) connected to the upper part (10) via a knee joint (1) and pivotably attached to the upper part (10) around a joint axis (15), wherein an adjustable resistance device (40) is arranged between the upper part (10) and the lower part (20), and the resistance device varies the flexion resistance (Rf) during early stance and mid-stance from after initial heel strike to mid-stance during walking based on sensor data, and after initial heel strike, the flexion resistance (Rf) increases to a value at which further flexion is stopped or at least slowed down, and the time course of the increase in flexion resistance and / or the maximum achievable flexion angle (Af) is varied depending on the ground slope or the height difference (ΔH) to be overcome.

Description

[Technical Field]

[0001] The present invention relates to a method for controlling a prosthetic or orthotic device for a lower limb having an upper part and a lower part connected to the upper part via a knee joint and attached so as to be pivotable about a joint axis relative to the upper part, wherein an adjustable resistance device is disposed between the upper part and the lower part, and the resistance device changes flexion resistance during walking from after initial heel strike to mid-stance and mid-stance based on sensor data. [Background technology]

[0002] Artificial knee joints are used in prosthetics and orthotics, as well as in exoskeletons as a special case of orthotics. An artificial knee joint has an upper and lower section pivotably mounted around a joint axis, i.e., the knee axis. In the simplest case, the knee joint is configured as a single-axis knee joint, where the individual knee axis is formed by, for example, a bolt or two bearing points arranged on a pivot axis. Knee joints that do not form a fixed rotation axis between the lower and upper sections but instead have sliding or rolling surfaces or multiple links articulated together are also known. To influence the kinematic characteristics of the knee joint and achieve a kinematic behavior of the orthosis, prosthesis, or exoskeleton that is closer to natural walking behavior, a resistance device is provided between the upper and lower sections, which allows the resistance to be changed at any given time. Purely passive resistance devices are passive dampers, such as hydraulic dampers, pneumatic dampers, or dampers that change the kinematic resistance based on the magnetorheological effect. There are also active resistance devices, such as motors or other drives, that can be connected accordingly and act as generators or energy accumulators.

[0003] Each knee joint, i.e., prosthetic or orthotic knee joint, is fixed to the patient using a respective connection means. In the case of a prosthetic knee joint, fixation is usually achieved via the thigh socket that receives the limb stump. Alternative types of fixation are possible, for example, via osseointegration connection means or straps and other devices. In the case of orthoses and exoskeletons, the upper and lower parts are fixed directly to the thigh and lower leg. Fixation devices provided therefor are, for example, straps, cuffs, shells or frame structures. The orthosis can also have a foot part for the foot or shoe to rest on. The foot part can be articulated to the lower part.

[0004] DE 102013011080 A1 relates to a method for controlling an orthopedic articulated device for the lower limbs, which has an upper part and a lower part articulated to the upper part, wherein a conversion device is disposed between the upper part and the lower part, which converts mechanical work from the relative movement during pivoting of the upper part relative to the lower part and stores it in at least one energy accumulator. The stored energy is then supplied to the articulated device again with a time lag to support the pivoting of the upper part and the lower part during the movement. The support of the relative movement is controlled. In addition to the conversion device, a separate damper in the form of an adjustable hydraulic or pneumatic damper can be provided, which allows the damper device to influence resistance in both the flexion and extension directions during walking.

[0005] Prosthetic knee joints are designed to have a knee angle of 180° at the maximum achievable extension, and hyperextension, i.e., a posterior angle greater than 180°, is usually not intended. Posterior rotation of the lower leg relative to the upper leg is called knee flexion, while forward or forward rotation is called extension. During initial contact, the foot strikes the floor at the end of the swing phase at the beginning of the stance phase. In this case, during flat walking, a so-called heel strike usually occurs, where the foot strikes the heel first. If the prosthetic knee joint remains in an extended, straight position during heel strike, this leads to direct force transmission to the pelvis, which is uncomfortable on the one hand and contrary to the natural gait pattern on the other. Therefore, in prosthetic or orthotic limbs, so-called stance phase flexion is enabled, in which the knee joint bends around the joint axis after heel strike, possibly against the resistance of a resistance device, as in normal flat walking.

[0006] From WO 2015 / 0101417 A1 is known a prosthetic knee joint having an upper and lower part pivotally attached to one another via a four-bar linkage system, the linkage system being attached to the lower part so that it can pivot from an initial position against a spring force during stance phase flexion, the line of action of the spring force being aligned to provide a moment that counteracts flexion of the linkage system. [Prior art documents] [Patent documents]

[0007] [Patent Document 1] DE102013011080A1 [Patent Document 2] WO2015 / 0101417A1 Summary of the Invention

[0008] The problem with previous control methods is that the flexion resistance during the stance phase is permanently set, which may make it difficult to provide a comfortable walking behavior in deviating walking situations.

[0009] It is therefore an object of the present invention to provide a method for controlling a lower limb prosthesis or orthosis that makes it possible to easily achieve an improved walking behavior for a user with an artificial knee joint.

[0010] According to the invention, the above problem is solved by a method having the features of the main claim. Advantageous embodiments and developments of the invention are disclosed in the dependent claims, the following description and the figures.

[0011] A method for controlling a lower limb prosthesis or orthosis having an upper part and a lower part connected to the upper part via a knee joint and pivotally attached to the upper part about a joint axis, wherein an adjustable resistance device is disposed between the upper part and the lower part, and the resistance device changes flexion resistance during walking from after initial heel strike to mid-stance based on sensor data. The method contemplates that after initial heel strike, the flexion resistance increases to a value at which further flexion is stopped or at least slowed, and the time course of the increase in flexion resistance and / or the maximum achievable flexion angle is changed depending on the ground inclination or the elevation difference to be overcome. The elevation difference to be overcome is the height of the prosthetic foot or foot or foot of a user of the artificial knee joint relative to the patient's contralateral foot or foot in the stance phase, or the elevation difference relative to the level of the prosthetic foot or orthotic foot or foot immediately preceding the stance phase during walking. Typical locomotion includes walking on flat ground, as well as slope walking (climbing up and down slopes) and stair walking, with stair descent being distinguished from slope descent. It is intended to limit the maximum possible stance phase flexion after initial heel strike to an adjustable angle. To avoid direct force transmission to the user's pelvis, stance phase flexion is permitted after initial heel strike. In this case, flexion damping increases depending on the knee flexion angle until the target angle is reached or at least not exceeded. Once the target angle is reached in the stance phase, further flexion is stopped. Until then, the flexion angle, which is manifested by a decrease in the posterior knee angle of the knee joint, or increases as the knee flexion angle increases, thereby preventing flexion from being stopped at loads and flexion moments that do not result in continued stance phase flexion up to the maximum allowable target angle. In this case, the time course of the increase in flexion resistance and / or the maximum achievable flexion angle is modified depending on the ground inclination or, in the case of stair climbing, when a height difference must be overcome. This allows the user to bend or flex the knee during the stance phase without risking excessive stance flexion and thus preventing full stance extension. There is no need for laborious inspection of the residual limb or the conscious use of remaining muscles. This allows the wearer to walk in a very relaxed and effortless manner, especially at slow speeds.

[0012] One development of the present invention contemplates that the maximum achievable flexion angle and / or the flexion angle at which the maximum flexion resistance is achieved increases as the downward slope of the ground increases. When walking downhill, the weight needs to be supported by stance-phase flexion when stepping off the wearer's foot. This is facilitated by increasing the maximum achievable flexion angle as the downward slope of the ground increases, i.e., as the descent becomes steeper, thereby providing a longer path for providing sufficient flexibility and a longer path for converting kinetic energy into heat, electrical energy, or other energy storage. Furthermore, as the downward slope of the ground increases, the maximum flexion resistance can be reduced to allow further flexion and an increase in the maximum achievable flexion angle. Such relatively large flexion angles can occur not only when walking on steep ground, but also during so-called braking processes or when descending stairs. When the maximum achievable flexion angle or maximum flexion resistance is reached, a plateau of varying degrees occurs in the time course of knee angle, as further knee flexion is prevented or becomes difficult after the target angle is reached or just before flexion stop is reached.

[0013] The flexion stop or increased flexion resistance can be maintained for a defined period in the plateau phase, followed by a decrease in the flexion resistance. A subsequent decrease in the flexion resistance over a defined period, which can also be set, for example, by achieving spatial alignment of the upper part, the lower part, or the connecting line between the upper part and the lower part, reduces the flexion damping again, for example, to the initial level of stance phase damping. In this case, in order to minimize the tilting moment generated in the upper part, the flexion damping is advantageously reduced gradually, the duration of which can depend on the ground slope or the orientation of the components relative to each other in space.

[0014] If the magnitude of the maximum lateral force in the lower leg exceeds a limit value dependent on the ground inclination, and / or the leg chord exceeds a forward tilt dependent on the ground inclination, and / or the magnitude of the hip moment first exceeds the limit value and then falls below it, the flexion resistance can be reduced after flexion stops and the maximum stance phase flexion angle is reached, or after flexion resistance increases and the stance phase flexion angle achievable with this resistance is reached. The magnitude of the lateral force in the lower leg, for example, in the lower leg support or lower leg brace, is a possible indicator of the current gait phase. This magnitude can be the lateral force itself, or it can be determined in relation to the weight, for example, known or detected by a sensor in the orthosis or prosthesis. If the lateral force or the magnitude of the lateral force exceeds a maximum value dependent on the ground inclination, a reduction in the flexion resistance can be initiated. In this case, the lateral force is a force component acting perpendicular to the longitudinal extension of the lower leg. In a vertical position with the leg fully extended, the lateral force extends in the anterior-posterior direction in the sagittal plane. The lateral force value can be measured directly by a lateral force sensor, which in this variant of the invention is the only force sensor required to carry out the method.

[0015] Alternatively or additionally, if the leg chord exceeds a forward tilt dependent on the ground inclination, the flexion resistance can be decreased again after flexion stops and the maximum stance phase flexion angle is reached, or after flexion resistance increases and the possible flexion angle is reached. The connecting line between the upper and lower parts, or two defined points on the parts following the lower part, is considered the leg chord. A preferred embodiment contemplates using the connecting line between the hip rotation center and the foot point as the leg chord. When using a prosthetic knee joint, the hip rotation center is identified by an orthopedic technician to determine the thigh or upper segment length, which is defined as the distance between the joint axis or knee axis and the hip rotation center. The lower segment length is defined by the distance between the knee axis and the foot point. For example, the foot center, the instantaneous pole of the rolling motion, the foot, the plantar level of the prosthetic foot, or the end point of a vertical line of the lower leg on the floor can be defined as the foot point; other points close to the floor are also suitable for defining the foot point. In an orthosis or exoskeleton, the foot is not required to support the remaining intact foot, so the distance from the floor to the joint axis can also be used. The position and / or length of the leg chord reliably provides information about the leg orientation and the progression of the movement. The leg chord can be calculated or estimated by multiple absolute angle sensors, one absolute angle sensor, and one knee angle sensor in combination with known segment lengths. When the leg chord exceeds a forward tilt relative to the ground, it can be inferred from this that the progression of the movement allows the release of the flexion arrest or the continued reduction of flexion resistance to be achieved in order to achieve swing and the start of the swing phase. By exceeding a forward tilt dependent on the ground inclination, mid-stance and the end of mid-stance are also detected. The ground inclination can be determined, for example, from the angle sensed at the ankle joint. However, the ground inclination can also be calculated in other ways.

[0016] Alternatively or additionally, if the magnitude of the lumbar moment first exceeds the threshold and then falls below it again, flexion can be stopped and the maximum stance-phase flexion angle reached, or flexion resistance can be increased and the flexion resistance can be reduced after the stance-phase flexion angle achievable with this resistance is reached. The magnitude of the lumbar moment can be exceeded or exceeded during a single stance phase. For example, when ascending a steep slope, a high lumbar moment is generated first and then decreases over the course of the stride, so the occurrence of a high, flexing lumbar moment during the stance phase is an indicator of ground inclination. The same is true for stair climbing. If a hip extension moment is detected, this is an indicator of slope climbing. Since the flexing lumbar moment decreases over the course of the stance phase stride, if the threshold is first exceeded and then falls below it, ground inclination can be inferred and the flexion resistance adapted accordingly. The lumbar moment can be calculated from the knee moment and known geometric relationships, from the orientation of the upper body in space, or from the orientation of the lower body in space and the knee angle. As the degree of the lumbar moment, in addition to the lumbar moment itself, a quantity related thereto, such as a value or numerical value determined depending on the spatial position of the upper part and / or body weight, can also be used.

[0017] Instead of measuring the lateral force directly, it is possible to calculate the magnitude of the lateral force from the difference between the lateral force components of the ankle moment and the knee moment. If the weight of the user of the knee prosthesis is additionally taken into account, the control and flexion resistance can be specifically adapted individually.

[0018] After increasing, the flexion resistance can decrease again once a predetermined knee flexion angle is exceeded, this decrease being reduced to a level below the stop level, for example, to the damping level of the early stance phase, but in this case, the magnitude of the increase in flexion damping depends on the ground inclination, and therefore the knee flexion angle may be exceeded, especially on relatively steep slopes.

[0019] The reduction of the bending resistance depending on the lateral force is particularly important during braking, especially on level ground, as well as when descending slopes or stairs. The angle of the leg chord depending on the slope inclination or the forward tilt of the leg chord depending on the slope inclination is crucial, especially on flat slopes or slopes with a moderate gradient, to avoid excessive hip extension and to allow timely flexion of the knee joint.

[0020] Ground tilt can be calculated as a path calculation reference from the vertical and / or horizontal distance traveled by the knee joint, particularly a reference point near the sole of the foot, during the preceding swing phase, or from the vertical and horizontal distance traveled by the knee joint, but particularly a reference point near the sole of the foot, during the preceding swing phase. For this purpose, for example, sensor signals from an inertial measurement unit are evaluated and integrated over a defined period. From this, speed and distance traveled are determined, which can be used to calculate ground tilt. Ground tilt is the ratio of the vertical distance traveled to the horizontal distance traveled. In this case, the path traveled by the point near the sole of the foot, i.e., the path traveled by the reference point, must be calculated. For this purpose, the position of the lower leg or shank is determined at the start and end of the integration, and the distance traveled by the reference point or foot relative to the inertial measurement unit (IMU) is calculated using geometric quantities and simplified angular functions.

[0021] The start of the stance phase to be controlled can be determined based on the axial force impulse, plantar flexion acceleration, and / or ankle moment. A pure axial force sensor in the foot or lower part can detect when the foot strikes the ground. After the force-free or axial force-free phase, a spontaneous increase in the axial force component is detected and used as a convincing indicator of the start of the stance phase. In the case of a prosthetic foot articulated to an articulated foot or lower part, plantar flexion acceleration can be calculated without a force sensor. Similarly, the ankle moment acting in the direction of plantar flexion can be calculated and used as the starting point of the stance phase to be controlled after the moment-free phase that causes plantar flexion.

[0022] In one variant of the invention, the ground inclination is calculated as a kinematic criterion from an evaluation of the flexion angle and the absolute angle of the upper or lower leg, or from an evaluation of two absolute angles of the upper and lower leg. The progression of the knee angle is detected and recorded together with the absolute angle of the upper or lower leg. Alternatively, the absolute angles of the upper and lower leg are used as a kinematic criterion, from which the ground inclination is calculated. After step-off or initial heel strike, depending on the ground inclination, different tangential gradients arise between the lower leg inclination and the knee joint. Knowing the respective tangential gradients, the respective ground inclinations can be estimated. For this purpose, the angular velocity of the knee and the angular velocity of the lower leg in space during walking can be calculated. From this, the quotient of the two angular velocities is calculated, and the ground inclination is calculated based on the change in the angular velocity quotient.

[0023] Such kinematic criteria or calculation of ground slope based on kinematic quantities can be used together with path calculation criteria based on calculating vertical and / or horizontal distance traveled, and each criterion can be weighted. In addition to equally weighting the slope calculated from the movement of the lower legs and thighs and the slope calculated based on path calculation data from IMU signals, for example, the kinematic criteria can be weighted less or used as an additional decision factor or error prevention measure only in certain situations or walking situations. For example, in dangerous situations when descending stairs, the kinematic criteria can be used in addition to the path calculation criteria to avoid unintentional locking or release of the knee joint.

[0024] The position and / or orientation of the ground reaction force vector relative to the prosthesis or orthosis can be used as a further control variable. Similarly, recognition of foot rolling on the edge can prevent increased damping or further reduce high resistance, which is advantageous during rolling movements of the attached leg, especially when descending stairs. The distance traveled for the path calculation criterion is calculated from the IMU values ​​of the lower part, in particular at the end of the preceding stance phase and the beginning of the stance phase to be controlled, where the IMU positions on the orthosis or prosthesis relative to the respective reference points and their spatial positions at the end of the stance phase, i.e., at toe-off, initial heel contact, or heel strike, are known. Both the path calculation criterion and the kinematic criterion can be used separately to determine the ground inclination, and the selectivity of the respective sensor signals can also be a factor for applying one criterion or another.

[0025] The start and end points of the path integral can be determined by a state machine, with various sensor signals monitored in response to various events. Such an event can be, for example, a loaded roll on the edge of a staircase, which can be recognized by detecting the axial force with simultaneous forward tilt of at least the lower or chord of the leg. Similarly, a loaded roll or lift of an orthosis or prosthesis, as well as reloading of the orthosis or prosthesis, can be used as definitive landmarks for the start and end points of the path integral.

[0026] Hereinafter, an embodiment of the present invention will be described in detail with reference to the accompanying drawings. [Brief explanation of the drawings]

[0027] [Figure 1] Schematic diagram of a prosthetic leg. [Figure 2] A diagram of the leg chords. [Figure 3] FIG. 1 is a diagram illustrating the definition of elevation difference during walking. [Figure 4] FIG. 10 illustrates the tilt-dependent adjustment of bending angle and bending resistance. [Figure 5a] FIG. 10 is a diagram showing the transition of the bending angle and bending resistance when the ground is inclined. [Figure 5b]FIG. 10 is a diagram showing another change in bending angle and bending resistance when the ground inclination is different. [Figure 5c] FIG. 10 is a diagram showing yet another progression of bending angle and bending resistance for different ground inclinations. [Figure 6] FIG. 10 is a diagram showing bending angles and roll angles when the ground inclination is different. [Figure 7] FIG. 10 illustrates kinematic criteria for different ground slopes. [Figure 8] FIG. 1 is a diagram showing geometrical references when descending a slope. [Figure 9] FIG. DETAILED DESCRIPTION OF THE INVENTION

[0028] FIG. 1 shows a schematic diagram of a knee prosthesis 1 applied to a prosthetic leg. Instead of applying it to a prosthetic leg, a correspondingly configured knee prosthesis 1 can also be used in an orthosis or exoskeleton. Instead of replacing a healthy joint, the respective knee prosthesis is then placed medially and / or laterally of the healthy joint. In the illustrated embodiment, the knee prosthesis 1 is configured in the form of a prosthetic knee joint, comprising an upper part 10 having a side, or anterior side 11, located in the forward, or walking direction, and a posterior side 12 located opposite the anterior side 11. A lower part 20 is attached to the upper part 10 so as to be pivotable about a pivot axis 15. The lower part 20 also has an anterior side 21 or front side and a posterior side 22 or rear side. In the illustrated embodiment, the knee prosthesis 1 is configured as a monocentric knee joint, but in principle, a polycentric knee joint can also be controlled accordingly. A foot 30 is located at the distal end of the lower part 20, and the foot can be connected to a fixed ankle joint as a fixed foot 30, or can be connected to the lower part using a pivot 35 to enable movement sequences that approximate natural movement sequences.

[0029] The knee angle KA is measured between the posterior side 12 of the upper part 10 and the posterior side 22 of the lower part 20. The knee angle KA can be measured directly by a knee angle sensor 25, which can be located in the area of ​​the pivot axis 15. The knee angle sensor 25 can be coupled to or can include a moment sensor to detect the knee moment about the joint axis 15. An inertial angle sensor or IMU 51 is located in the upper part 10, which measures the spatial position of the upper part 10 relative to a certain force direction, such as gravity G pointing vertically downward. An inertial angle sensor or IMU 53 is also located in the lower part 20 to sense the spatial position of the lower part during use of the prosthetic leg.

[0030] In addition to the inertial angle sensors 53, acceleration sensors and / or lateral force sensors 53 can be located in the lower part 20 or foot 30. Axial forces FA acting on the lower part 20 or ankle moments acting about the ankle joint axis 35 can be sensed by force or moment sensors 54 in the lower part 20 or foot 30.

[0031] A resistance device 40 is arranged between the upper part 10 and the lower part 20 to influence the pivotal movement of the lower part 20 relative to the upper part 10. The resistance device 40 can be configured as a passive damper, a drive device, or a so-called semi-active actuator 40, which allows kinetic energy to be stored and later released again precisely to damp or assist the movement. The resistance device 40 can be configured as a linear or rotary resistance device. The resistance device 40 is connected to a control device 60, for example, by wire or wireless connection, which is also coupled to at least one of the sensors 25, 51, 52, 53, 54. The control device 60 electronically processes the signals transmitted by the sensors by a processor, computing unit, or computer. The control device has an electrical energy supply and at least one memory unit, which stores programs and data and includes a main memory for processing the data. After processing the sensor data, an activation or deactivation command is output, which activates or deactivates the resistance device 40. By actuating the actuator of the resistance device 40, for example, a valve can be opened or closed, or a magnetic field can be generated to change the damping behavior.

[0032] The upper part 10 of the prosthetic knee joint 1 is fitted with a prosthetic socket used to accommodate the femoral stump. The prosthetic leg is connected to the hip joint 16 via the femoral stump. A hip angle HA is measured on the anterior side of the upper part 10, and this hip angle is provided on the anterior side 11 between a vertical line passing through the hip joint 16 and the longitudinal extension of the body and a connection line between the hip joint 16 and the knee joint axis 15. When the femoral stump is raised and the hip joint 16 is flexed, for example, when sitting, the hip angle HA decreases. Conversely, the hip angle HA increases during extension, for example, when standing up or similar movement sequences.

[0033] During the gait cycle when walking on level ground, the foot 30 strikes the ground first with the heel; this initial contact of the heel or foot 30 is called heel strike. Subsequently, plantar flexion occurs until the foot 30 fully contacts the floor, typically with the longitudinal extension of the lower leg 20 behind a vertical line extending through the ankle joint axis 35. Then, during level walking, the body's center of mass shifts forward, the lower leg 20 pivots forward, the ankle angle AA decreases, and toe loading increases. The ground reaction force vector shifts forward from the heel to the toe. At the end of the stance phase, toe-off, or so-called toe-off, occurs, followed by the swing phase, during which the foot 30 is moved behind the center of mass or ipsilateral hip joint with a decreasing knee angle KA, then rotates forward after reaching a minimum knee angle KA, and then typically reaches heel strike again with the knee joint 1 fully extended. Thus, the force entry point PF moves from heel to toe during the stance phase, which is shown schematically in Figure 1.

[0034] In Figure 2, the leg tendons 70 of the ipsilateral attached leg and the contralateral unattached leg are defined. The leg chord passes through the hip center of rotation 16 and forms a line to the ankle joint 35. As can be seen from Figure 2, the length of the leg chord and the orientation φ of the leg chord 70 L changes during exercise, especially for different gradients. The height difference ΔH to be overcome can be estimated and predicted or detected by the progression of the change in length and / or orientation of the leg chord 70, from which the respective control commands are then derived. Lk The opposite leg chord φ Li The directions of each are written in.

[0035] Based on Figure 3, the step height between the contralateral non-attached leg and the ipsilateral foot 30 of the attached leg can be defined. For example, the distance H1 is set from the floor to a characteristic point of the hip at the height of the standing leg, such as the hip joint 16 or the greater trochanter, and the distance H2 is the distance between the floor and the hip joint 16 or the greater trochanter on the leading side, that is, on the attached side in the illustrated example. In this case, the height difference ΔH is calculated from the difference between H1 and H2. The definition of the height difference ΔH also applies to walking on a slope.

[0036] Figure 4 shows diagrams of different settings for the bending resistance Rf and the bending angle Af. The bending resistance Rf and the bending angle Af are set as maximum values, respectively. The maximum bending attenuation or maximum bending resistance Rf remains approximately constant during slope ascent, flat walking, and flat slope. The maximum bending resistance decreases, for example, by 5%, only as the downward slope of the slope increases. In the stair descent situation, the maximum bending resistance Rf decreases to a much lower level, particularly the level of stance phase attenuation during flat walking. The maximum bending angle Af also varies depending on the ground inclination. Slope ascent and flat walking are performed at the same maximum bending angle Af during the stance phase. The maximum achievable bending angle Af increases depending on the ground inclination up to the maximum value set for steep slope walking and stair descent. Such stance phase control adapts the bending resistance Rf during early and mid-stance depending on the ground inclination, thereby also limiting the maximum bending angle within the stance phase. The maximum possible flexion angle Af can be set, and the limit value is achieved by changing the flexion resistance Rf. Depending on the ground inclination, a maximum target value is predetermined, at which the flexion resistance Rf has a high value at which further knee flexion or knee flexion is impossible. Once the maximum flexion angle Af, which depends on the ground inclination, is reached, further flexion is prevented, and different knee angle plateaus are formed during further exercise, i.e., walking on level ground, walking on a slope, or climbing stairs. If an increase in the flexion angle is recognized, for example, when walking on a moderate or steep slope or climbing stairs, the flexion resistance Rf is reduced, thereby enabling an increase in the flexion angle Af without a plateau. For example, during stair climbing, the flexion resistance Rf is reduced to the stance phase damping level.

[0037] Figures 5a to 5c show the time course of the flexion angle Af and the flexion resistance Rf for different ground inclinations. The left graph in Figure 5a shows the course of the two characteristic quantities for walking on a flat slope, the center graph in Figure 5b shows the course of the two characteristic quantities for walking on a flat slope, and the right graph in Figure 5c shows the course of the characteristic quantities for walking on a slope with a moderate slope. In Figure 5a, after initial contact with the ground, knee flexion occurs first, thereby increasing the flexion angle Af. In the left third of Figure 5a, it becomes clear that as the flexion resistance Rf increases, further flexion is prohibited, resulting in a plateau in both the flexion resistance Rf and the flexion angle Af. During further forward movement, after rolling, the flexion angle Af decreases. Subsequently, the flexion resistance Rf decreases from a certain limit value to allow further flexion at the end of the stance phase, so that a sufficiently large flexion angle Af can be achieved during the swing phase.

[0038] Figure 5b shows flat slope walking. After heel strike, flexion resistance Rf increases until the flexion angle Af cannot be further increased. Unlike Figure 5a, the plateau phase in the first third of the diagram in Figure 5b is followed by rolling, in which the knee joint flexes rather than extends. This movement process is typical for flat slope descents. Flexion resistance Rf is reduced after a defined period. This period can be determined, for example, so that the foot achieves full foot contact after heel strike and normal walking speed. To initiate flexion, statistical data on the duration of the stance phase and, therefore, the first decrease in flexion resistance Rf can be used. Flexion resistance Rf decreases, allowing further flexion and an increase in the flexion angle Af. Again, the swing phase is enabled, and flexion resistance Rf is reduced to a minimum value.

[0039] Figure 5c shows the progression of the characteristic quantities on a relatively steep surface, with a very short plateau after the increase in flexion resistance Rf. The decrease in flexion resistance Rf for the onset of the swing and swing phases, as in Figure 5b, occurs as late as possible, ensuring sufficient stability, especially during downhill walking. Each plateau in the progression of the knee angle Af is accompanied by an initial deceleration of knee flexion due to the increase in flexion resistance Rf. This allows for more ground clearance in the event of a swing on the opposite, i.e., non-wearing, side. This avoids unnecessary compensatory movements on the opposite side. Of particular importance is the progression of flexion resistance Rf between heel strike or initial heel contact and a new extension movement or during a rolling movement.

[0040] Figure 6 shows three diagrams for different ground inclinations. The top diagram is for walking on level ground, the middle diagram is for descending a slope, and the bottom diagram is for descending stairs. The left graph shows the progression of the flexion angle Af and the lower inclination angle or roll angle As. The right graph plots the roll angle As as the X component and the flexion angle Af as the Y component. Each heel strike is marked with a circle. Starting from heel strike, each diagram forms a closed two-dimensional curve with a characteristic shape. The tangent slope of the curve shape is calculated at selected times, such as at fixed discrete times. To achieve this, the ratio of the flexion angular velocity to the roll angular velocity is continuously calculated. The calculated ratio is optionally smoothed by a low-pass filter. To avoid excessive signal interference due to the initial heel strike, the filter is turned on or initialized only immediately after heel strike or heel strike. From heel strike through the onset of stance phase flexion, the calculated and filtered values ​​are assigned to the corresponding slope of the slope path by an interpolation function using the defined fulcrum. As can be seen from the diagram in Figure 6, different tangent slopes appear on the curve for different gait situations with different slopes, which can be used to determine the ground slope and to set the corresponding change in flexion resistance depending on the ground slope.

[0041] Figure 7 shows three graphs for the evaluation of kinematic criteria, namely, the lower flexion angle Af and the roll angle As. The top graph represents walking on level ground, the middle graph represents walking on a flat slope, and the bottom graph represents walking on a steep slope. The curve shape of the flexion angle Af substantially corresponds to the shape of the flexion angle in Figures 5a-5c. The solid lines represent the progression of the flexion resistance Rf, and the dashed lines represent actual measurements over time. In the top graph, the kinematic criterion Kk is consistently set to a value of 2, which corresponds to walking on level ground. A value of 1 corresponds to walking on a moderate slope, and a value of 0 corresponds to walking on a steep slope or stairs. In the top graph, the kinematic criterion Kk for walking on level ground consistently exhibits accurate values, providing correspondingly adapted control of the flexion resistance Rf during stance phase flexion with a pronounced plateau phase of the flexion angle Af. In the middle graph, the value of the kinematic criterion Kk drops to approximately 1.6, which corresponds to walking on a flat slope. As soon as the value of flexion resistance Rf increases, a self-reinforcing process begins, which again raises the kinematic criterion Kk in the direction of level walking. The minimum value of the kinematic criterion Kk reached after heel strike is therefore crucial for the control of flexion resistance Rf. In this case, this causes the flexion resistance Rf to increase until it stops at a flexion angle Af that is significantly larger than in level walking.

[0042] In the bottom diagram of Figure 7, it can be seen how the kinematic criterion Kk drops rapidly to a value of 0 when walking on a very steep slope, equivalent to walking on stairs. This prevents an increase in the bending resistance Rf, which allows further bending of the knee and an increase in the bending angle Af without a plateau phase. Therefore, each walked surface or each perceived walking situation leads to significantly different changes in the bending resistance Rf.

[0043] FIG. 8 shows a path calculation criterion, calculated based on the IMU sensor signal over a defined period. In this case, the acceleration of the knee joint 1 or lower part 20 is integrated in the forward and upward directions. The upward direction is the direction opposite to the direction of gravity, and the forward direction is forward movement from rear to front on the sagittal plane. Each velocity is calculated by simple integration with respect to time, and the vertical and horizontal distances traveled are calculated by double integration with respect to time of the IMU values. In FIG. 8, the horizontal distance traveled ΔV is shown, and the vertical distance traveled is denoted by ΔH. The vertical velocity is represented by Vv, and the horizontal velocity is represented by Vh. In FIG. 8, walking on a downhill slope is shown. The overall slope of the ground is defined as the ratio of the horizontal distance traveled ΔV to the vertical distance traveled ΔH, and the distance traveled near the sole of the foot is calculated. For this purpose, the position of the lower part 2 is detected at the start and end of the integration, and the distance traveled by the foot relative to the respective IMU can be calculated using known geometric quantities such as the lower part length, the position of the IMU in the lower part 20 or in the upper part 10, and the knee angle. The distance calculation criterion also allows the respective ground inclination to be recognized with a relatively good resolution. The value of the overall ground inclination calculated by the distance criterion can then be used as a control parameter for setting the bending resistance Rf.

[0044] FIG. 9 shows a schematic diagram of an embodiment of a brace having an upper section 10 and a lower section 20 pivotally attached to the upper section about a pivot axis 15, which can also be used to perform the method. The upper section 10 and the lower section 20 form an artificial knee joint 1, which in the illustrated embodiment is positioned laterally relative to the intact knee joint. In addition to positioning the upper section 10 and the lower section 20 on one side of the leg, the two upper and lower sections can also be positioned medially and laterally relative to the intact leg. The lower section 20 has a foot section 30 at its distal end pivotally attached to the lower section 20 about an ankle joint axis 35. The foot section 30 can have a footplate on which a foot or shoe can be placed. Fixation devices for fixing the lower leg or thigh can be disposed on both the lower leg 20 and the upper section 10. A device for fixing the foot to the foot section 30 can also be disposed on the foot section 30. The fixation devices can be formed as clasps, belts, buckles, or the like, so that the brace can be removably attached to the user's leg and removed again without destruction. A resistance device 40 is attached to the upper part 10 and is supported by the lower part 20 and upper part 10, providing adjustable resistance to pivoting about pivot axis 15. Thus, the sensors and controls described above in connection with the prosthetic embodiment are also provided in the orthosis. The inventions described in the original claims of this application are set forth below. [1] A method for controlling a prosthetic or orthotic device for a lower limb having an upper part (10) and a lower part (20) connected to the upper part (10) via a knee joint (1) and pivotably attached to the upper part (10) around a joint axis (15), wherein an adjustable resistance device (40) is disposed between the upper part (10) and the lower part (20), and the resistance device changes the flexion resistance (Rf) during early stance and mid-stance from after initial heel strike to mid-stance during walking based on sensor data, characterized in that after initial heel strike, the flexion resistance (Rf) increases to a value at which further flexion is stopped or at least slowed down, and the time course of the increase in flexion resistance and / or the maximum achievable flexion angle (Af) are changed depending on the ground slope or the height difference (ΔH) to be overcome. [2] The method according to [1], characterized in that the maximum achievable flexion angle (Af) and / or the flexion angle (Af) at which the maximum flexion resistance (Rf) is achieved increases as the downward slope of the ground increases. [3] The method according to [1] or [2], characterized in that the maximum bending resistance (Rf) decreases as the downward slope of the ground increases. [4] The method according to any one of [1] to [3], wherein the cessation of bending or the increase in bending resistance is maintained for a predetermined period of time, followed by a decrease in the bending resistance (Rf). [5] The method according to [4], characterized in that after the cessation of flexion or the increase in flexion resistance, the flexion resistance decreases when the degree of lateral force in the lower part (20) exceeds a limit value that depends on the inclination of the ground, and / or the leg chord (70) exceeds a forward tilt that depends on the inclination of the ground, and / or the degree of lumbar moment first exceeds and then falls below a limit value. [6] The method according to [5], characterized in that the degree of the lateral force is calculated by a lateral force sensor or from the difference between the lateral force components of the ankle moment and the knee moment. [7] The method according to any one of [1] to [6], characterized in that when a predetermined flexion angle (Af) is exceeded, the flexion resistance (Rf) increases and then decreases to below a stop level. [8] The method according to any one of [1] to [7], characterized in that the inclination of the ground surface is calculated as a path calculation criterion from the vertical and / or horizontal distance of a reference point, particularly near the sole of the foot, of the knee joint (1) advanced in the preceding swing phase, or from the ratio of the vertical and horizontal distance of a reference point, particularly near the sole of the foot, of the knee joint (1) advanced in the preceding swing phase. [9] The method according to [8], characterized in that the onset of the stance phase to be controlled is determined based on axial force impulse, plantar flexion acceleration and / or ankle moment.

[10] The method according to any one of [1] to [7], characterized in that the inclination of the ground is calculated as a kinematic criterion from an evaluation of the bending angle (Af) and absolute angle of the upper part (10) or the lower part (20), or from an evaluation of two absolute angles of the upper part (10) and the lower part (20).

[11] The method according to

[10] , characterized in that the knee angular velocity and lower angular velocity during walking are detected, a quotient of the two angular velocities is calculated therefrom, and the inclination of the ground is calculated based on a change in the quotient of the angular velocities.

[12] The method according to [8] and

[10] , characterized in that the path calculation criterion and the kinematic criterion are used to determine the ground slope.

[13] The method according to any one of [1] to

[12] , characterized in that the position and / or direction of the ground reaction force vector for the orthosis or prosthesis is used as the control variable.

[14] The method according to [1], characterized in that the recognition of the rollover of the foot (30) at the edge prevents the bending resistance from increasing or reduces said increased bending resistance (Rf) again.

Claims

1. A method for controlling a prosthetic or orthotic device for a lower limb, the prosthetic or orthotic device having an upper part (10) and a lower part (20) connected to the upper part (10) via a knee joint (1) and pivotably attached to the upper part (10) around a joint axis (15), wherein an adjustable resistance device (40) is arranged between the upper part (10) and the lower part (20), and the resistance device changes a flexion resistance (Rf) during walking from initial heel strike to mid-stance based on sensor data, wherein after initial heel strike, the flexion resistance (Rf) increases to a value at which further flexion is stopped or at least knee flexion is decelerated, and the time course of the increase in the flexion resistance and / or the maximum achievable flexion angle (Af) is changed depending on the slope of the ground or the height difference (ΔH) to be overcome. The method is characterized in that the maximum bending resistance (Rf) decreases as the slope of the ground increases.

2. 2. The method of claim 1, wherein the maximum achievable bending angle (Af) and / or the bending angle (Af) at which the maximum bending resistance (Rf) is achieved increases with increasing ground inclination.

3. 3. A method according to claim 1 or claim 2, characterized in that the cessation of bending or the increase in bending resistance is maintained for a defined period of time, followed by a decrease in the bending resistance (Rf).

4. 4. The method according to claim 3, characterized in that after the cessation of bending or the increase in bending resistance, the bending resistance decreases when the degree of lateral force in the lower part (20) exceeds a limit value that depends on the inclination of the ground, and / or the leg chord (70) exceeds a forward tilt that depends on the inclination of the ground, and / or the degree of lumbar moment first exceeds and then falls below a limit value.

5. 5. The method of claim 4, wherein the degree of lateral force is determined by a lateral force sensor or from the difference between the lateral force components of the ankle moment and the knee moment.

6. 6. The method according to any one of claims 1 to 5, characterized in that above a predetermined bending angle (Af), the bending resistance (Rf) increases and then decreases below a stop level.

7. 7. The method according to claim 1, wherein the ground slope is calculated as a path calculation criterion from the vertical and / or horizontal distance of a reference point, in particular near the sole of the foot, of the knee joint (1) advanced during the preceding swing phase or from the ratio of the vertical and horizontal distances of a reference point, in particular near the sole of the foot, of the knee joint (1) advanced during the preceding swing phase.

8. 8. The method of claim 7, wherein the onset of the stance phase to be controlled is determined on the basis of axial force impulse, plantar flexion acceleration and / or ankle moment.

9. 7. The method according to any one of claims 1 to 6, characterized in that the ground slope is calculated as a kinematic criterion from an evaluation of the bending angle (Af) and the absolute angle of the upper part (10) or the lower part (20) or from an evaluation of two absolute angles of the upper part (10) and the lower part (20).

10. 10. The method according to claim 9, wherein the angular velocity of the knee and the angular velocity of the lower body during walking are detected, a quotient of the two angular velocities is calculated therefrom, and the ground slope is calculated based on a change in the quotient of the angular velocities.

11. The method according to any one of claims 1 to 6, characterized in that the path calculation criterion calculated according to claim 7 and the kinematic criterion calculated according to claim 9 are used to determine the ground slope.

12. 12. The method according to any one of claims 1 to 11, characterized in that the position and / or orientation of the ground reaction force vector for the orthosis or prosthesis is used as a control variable.

13. 2. A method according to claim 1, characterized in that the recognition of the rollover of the foot (30) at the edge prevents the bending resistance from increasing or reduces said increased bending resistance (Rf) again.

Images