EP4687775A1 - Method for controlling an orthopaedic joint arrangement - Google Patents

Abstract

The invention relates to a method for controlling an orthopaedic joint arrangement of a lower extremity having an upper part (10) and a lower part (20), which are hinged to one another so as to be pivotable about a pivot axis (15), using an actuator (30) which is coupled to the upper part (10) and the lower part (20) and influences a movement state of the upper part (10) and/or lower part (20), wherein the actuator (30) is coupled to a control device (40) which is coupled to at least one sensor (50) and, on the basis of sensor values from the at least one sensor (50), activates, deactivates or modulates the actuator (30), wherein at least one movement speed of at least one part of the orthopaedic joint arrangement is detected from the sensor values and the actuator (30), on the basis of the movement speed in the standing phase, is activated, deactivated or modulated, wherein the movement speed in the standing phase and the resistance at least for part of the movement are inversely correlated.

Description

Method for controlling an orthopedic joint device

The invention relates to a method for controlling an orthopedic joint device of a lower extremity with an upper part and a lower part, which are pivotably mounted to one another about a pivot axis, with an actuator which is coupled to the upper part and the lower part and influences a state of movement of the upper part and/or lower part, wherein the actuator is coupled to a control device which is coupled to at least one sensor and activates, deactivates or modulates the actuator on the basis of sensor values of the at least one sensor, wherein at least one movement speed of at least one part of the orthopedic joint device is determined from the sensor values and the actuator is activated, deactivated or modulated on the basis of the movement speed in the stance phase.

Orthopedic joint devices are in particular orthoses, exoskeletons or prostheses that have an upper part and a lower part that is articulated to it. In orthoses and exoskeletons, the upper part and the lower part are attached to a limb that is still present, for example by means of shells, belts, straps, cuffs or other fastening devices. Orthoses and exoskeletons can be used to guide movements, limit pivoting around a joint axis, prevent pivoting movements or support or determine the alignment of limbs with respect to one another. In polycentric joints, the pivot axis is the instantaneous center of rotation, which shifts depending on the pivoting movement. In addition, orthoses can be provided with damping devices to dampen a pivoting movement around the joint axis. The damping devices can be provided with a control so that a changed damping in the flexion direction and/or extension direction can be provided depending on sensor data. It is also known to assign energy storage devices to the upper or lower part, so that movement support can be achieved by releasing the stored energy from the energy storage device.

Prostheses replace a missing or no longer existing limb and are used to provide functionality that is as close as possible to the functionality of the natural limb. In addition, prostheses are used to provide the most natural appearance possible for the prosthetic user. A prosthetic upper part is designed, for example, as a prosthetic shaft or as a component attached to a prosthetic shaft, where the prosthetic shaft is used to attach it to a limb or a limb stump. The prosthetic joint, for example a prosthetic knee joint or a prosthetic ankle joint, connects the upper part to a lower part, which in turn can have further prosthetic components, for example a lower leg tube or a prosthetic foot.

Particularly in the case of orthoses, exoskeletons and prostheses of the lower extremities, but also of the upper extremities, dampers, in particular hydraulic dampers or other resistance devices, are arranged between the upper part and the lower part, which provide different resistances in individual states or movement situations based on sensor data. Such resistance devices are often designed as linear actuators that provide a defined resistance to a flexion movement and/or extension movement. The resistance is changed, for example, by changing the position of valves. When the flow cross-section is reduced, the corresponding resistance to a movement increases. Passively damped, in particular passively hydraulically damped prostheses or orthoses work purely dissipatively. Energy is taken from the movement of the upper part relative to the lower part, whereby very high moments or forces can be generated. At the same time, passive damping in an open state, for example when no valves are closed or throttles are activated, has only a very low resistance. The working range of such an orthosis or prosthesis is limited in that no energy can be directed into the movement to support it or To actively counteract movement or to make a change from a static state.

In addition, orthoses, exoskeletons and prostheses with motor drives are known from the state of the art, so-called active orthoses or prostheses, in which a movement is initiated, supported or slowed down by activating, deactivating or modulating the drive. For this purpose, stored electrical energy from a battery or accumulator is converted in the actuator. The motor drives also serve to influence the movement behavior between the components of the orthosis or prosthesis, for example to slow down a pivoting movement. Motor drives can be operated in braking mode or as part of a generator circuit.

Purely passive devices such as dampers, as well as semi-active devices such as energy storage devices and motor drives influence the movement behavior of the upper part and/or the lower part and are actuators that influence a state of motion of the upper part and/or the lower part. The actuators can initiate a movement, reverse a movement, support a movement or resist a movement. An influence on a state of motion of the upper part and/or the lower part also occurs when a load is counteracted, a static state is maintained or a change in a state of motion is prevented or stopped due to external forces. This can happen, for example, if a uniform pivoting movement is to be maintained and external forces act in the direction of movement or act against the direction of movement.

EP 2 869 792 B1 discloses a method for controlling an orthopaedic joint device of a lower extremity with an upper part and a lower part articulated thereto, between which an energy conversion device and/or a storage device is arranged, via which kinetic energy from the relative movement between the upper part and the lower part is converted and/or stored during walking. This energy can be fed back into the joint in order to compensate for the relative movement to support, whereby kinetic energy is converted and/or stored within a movement cycle of the joint device and is fed back as kinetic energy in a controlled and time-delayed manner within the same movement cycle of the joint device. The conversion rate and/or storage rate of the energy conversion device or storage device is inversely proportional to the pivoting speed of the lower leg.

WO 2016/169 850 A1 relates to a method for controlling a damping change in an artificial joint of an orthosis, an exoskeleton or a prosthesis of a lower extremity with a resistance unit between an upper part and a lower part, which are pivotally attached to one another. The resistance is changed via a resistance unit when a sensor signal from a control unit assigned to the adjustment device activates the adjustment device. The resistance is changed depending on the position and/or length of the leg tendon or its temporal derivatives.

WO 2016/169 848 A1 also relates to a method for controlling a damping change of an artificial knee joint, in which the flexion resistance is reduced for the swing phase. During walking or standing, the course of at least one load characteristic that acts on an orthosis or prosthesis to which the artificial knee joint is attached is recorded. If a maximum of the load characteristic course is determined during the stance phase or standing and a threshold value of the load characteristic below the maximum is subsequently detected, the flexion damping during the stance phase is reduced to a swing phase damping level.

For example, in the stance phase of orthopedic devices for the lower extremities, microprocessor-controlled knee joints use the resistance to movement in the direction of flexion to enable controlled bending in different movement situations. Depending on the respective movement situation, for example walking on a level surface with stance phase flexion, walking down ramps or walking down stairs, the level of resistance or the course of the resistance is changed. During stance phase flexion, there is usually a resistance that is higher than the resistance when standing. If the orthopaedic device is at rest for a certain time, for example if the user is standing for a longer period of time, the joint is locked and when movement occurs, the lock is released again.

The object of the present invention is to provide a method for controlling an orthopedic joint device of the lower extremity, which enables the user to use the orthopedic device comfortably and safely with as little physical effort as possible.

This object is achieved by a method having the features of the main claim. Advantageous embodiments and further developments of the invention are disclosed in the subclaims, the description and the figures.

In the method for controlling an orthopedic joint device of a lower extremity with an upper part and a lower part, which are pivotably mounted to one another about a pivot axis, with an actuator which is coupled to the upper part and the lower part and influences a state of movement of the upper part and/or lower part, wherein the actuator is coupled to a control device which is coupled to at least one sensor and activates, deactivates or modulates the actuator on the basis of sensor values of the at least one sensor, wherein at least one movement speed of at least one part of the orthopedic joint device is determined from the sensor values and the actuator is activated, deactivated or modulated on the basis of the movement speed in the stance phase, the movement speed in the stance phase and the resistance are inversely correlated with one another for at least part of the movement, for example, as the movement speed in the stance phase increases, the resistance is reduced. This results in an adjustment of the resistance of a relative movement of the upper and lower parts around the pivot axis depending on the speed of movement, whereby a reduced resistance is provided in the standing phase at an increased speed of movement. The speed of movement is also particularly the Speed is viewed as the speed of movement; in addition, translational or rotational speeds of the upper and/or lower parts can also be used to determine the speed of movement. It is also possible to determine the speed via the temporal change of forces, moments, lever arms and/or force application points, for example via the rate of change with which a force application point on the sole of the foot moves from the forefoot towards the heel or vice versa, or via the rate of change of the ankle moment. The resistance can be both flexion resistance and extension resistance. The resistance can be passive resistance, such as that provided by a passive damping device, a locking device or a brake, or a force applied to the movement by an actuator, for example a drive or a force accumulator. A moment or force applied by the actuator can counteract a movement in one direction of movement and influence the movement as resistance, and actively support a movement and do work in an opposite direction of movement. For example, a force can counteract knee flexion and represent a flexion resistance and actively support knee extension and act as a drive or support.

In a further development of the method, the resistance is increased when the movement speed in the stance phase decreases, so that when the walking speed decreases, the flexion resistance in the stance phase increases. The slower the walking speed, the longer the stance phase lasts and the smaller the stride length. Accordingly, it is advantageous to increase the resistance to movement, in particular the resistance to flexing of the knee joint, in order to make the sinking slow so that the person using it does not have to raise their center of gravity excessively at the end of the stance phase flexion in order to complete the step. By increasing the resistance with decreasing speed, the extent of a movement, for example flexing of a knee joint, can be reduced or completely prevented and/or the extent of movement can be increased or movement can be permitted as the speed increases. In one embodiment of the method, a change in resistance only occurs above a specified threshold value and/or only up to a specified threshold value of the movement speed. The movement speed is in particular the walking speed, whereby in addition to a forward progression, the speed in the backward or sideways direction can also be used. Alternatively, the speed is a rotational movement about an axis. Preferably, the movement speed is determined in the respective stance phase itself, the determination is carried out as up-to-date as possible and only takes into account the speeds in a previous step or in a previous movement phase as boundary conditions, since the movement speed can change at any time and a change in resistance is best adapted to the current conditions and not to the step or movement at a previous point in time. Changing the resistance only above a specified threshold value of the change in the movement speed means that changes at very low and/or high speeds are disregarded and the control effort as well as the energy consumption are reduced. Alternatively or additionally, the change in resistance only occurs from or up to a specified threshold value of the change in speed, which ensures that smaller changes in speed are disregarded and also reduces the control effort and energy consumption.

In one embodiment, the change in resistance in at least one speed range is non-linear, in particular the resistance initially remains almost constant at low speeds and only drops slightly with a small increase in speed and is reduced with increasing speed, with saturation occurring in particular at a very high movement speed and the resistance to flexion or extension approaching a limit value. Alternatively or additionally, the resistance in at least one speed range decreases proportionally to a power of the movement speed, in particular the square of the movement speed. The reduction in resistance with increasing speed can also be designed in one step or several discrete steps or levels. The resistance can be changed by changing the resistance level, i.e. a uniform increase in resistance, for example, over the entire swivel range. The resistance level or resistance profile can change over the swivel range, for example, a resistance adapted to the load can be set depending on an angular position. The resistance curve can also be changed to change the resistance, for example, starting from a standard resistance curve for a certain movement speed, if the speed is reduced, the resistance can be increased in certain areas over the swivel angle or over the duration of the movement, and if the movement speed is increased, the resistance can be reduced more.

The change in resistance can also occur depending on forces, relative and/or absolute angles and/or moments of the position of components or limbs in relation to one another or to the environment and/or their temporal progressions as well as temporal derivatives determined by sensors or calculated from sensors. These quantities can relate to the supplied side or the contralateral side. The change in resistance can also depend directly on time. The resistance can therefore also be adjusted based on other determined quantities, for example a knee angle, a segment angle, a leg tendon angle, the ground reaction force vector, a moment, a force, a lever arm, a size of the contralateral side or a combination of one or more of these quantities or factors. The temporal derivatives can also be used to adjust the resistance. It is also possible that the properties of the environment, for example the inclination, geometry or condition of the ground, are determined from sensor data and used to adjust the resistance. In this case, the relationships between the quantities or factors with the resistance may also depend on the speed of movement and may vary together with the speed. For example, a resistance may be increased more strongly or earlier at a low speed with a comparatively small knee angle than at a higher speed or a larger knee angle. In one embodiment, the resistance curve and/or the resistance level are changed separately for each stance phase, so that the correct resistance level or the correct resistance curve is provided for the current stance phase.

A further development provides that the speed of movement, in particular the walking speed in orthopedic devices of the lower extremities, in particular with artificial knee joints, is calculated or determined using a determined position and/or length of a leg tendon and/or its rates of change and is used as the basis for the change in resistance as the speed of movement. The leg tendon is in particular the connecting line between a hip pivot point and a foot point, whereby the foot point is defined in particular at the end of the extension of a lower leg part up to the sole of the foot. The foot point can also be defined as the pivot point of the foot part relative to the lower leg or as a rolling point or instantaneous center of rotation of the foot part. The length of the leg tendon thus changes due to a change in the knee angle, the position and location of the leg tendon changes due to a pivoting around the foot point in the stance phase or the hip pivot point in the swing phase. Both the leg tendon of the treated side and the contralateral side can be used. Based on the known segment lengths, i.e. the distance from the knee joint axis to the hip pivot point and the foot point, and the determined position of the upper part of the knee joint to the lower part of the knee joint, it is possible to calculate the length of the leg tendon. The position of the leg tendon is determined, for example, using spatial position sensors and angle sizes as well as segment lengths, for example on the basis of an absolute angle of a lower leg and a knee angle as well as the length of the lower leg and thigh. The change in position can be determined using the temporal derivative of the position or using speeds and segment lengths. The speed of movement is calculated from the position and/or length of the leg tendon and, if applicable, its changes over time and used as the basis for the change in resistance. In the stance phase, the speed of the hip, torso and/or body center of gravity can be determined from the leg tendon and its change in position and/or length. A further development provides that the speed of movement is determined by means of temporal integration of accelerations recorded by sensors, for example the translational speed of the lower leg at the height of the knee joint axis or that of the torso. Using kinematic chains and their degrees of freedom, which can be determined using sensors, it is possible to deduce the speed of one point, for example on the sole of the foot or on the lower leg at the height of the knee axis, from that of another point, for example from the foot to the hip. It is also possible to deduce a speed based on sensor signals, for example using a force sensor and an absolute angle sensor on a lower leg to deduce that a foot is in the middle stance phase and therefore almost at a standstill.

Alternatively or additionally, the speed of movement can also be determined relative to the ground and/or the environment, for example using environmental sensors such as LIDAR, radar, Doppler radar or ultrasound or navigation systems such as indoor navigation or global navigation systems such as QZSS or Galileo.

It is possible to use the absolute speed and/or one or more components, for example the components parallel to the ground or the horizontal component, as the speed of movement.

A further development provides that at least a rotational speed of a thigh, a lower leg and/or an angular speed of an ankle joint, knee joint and/or hip joint is used as the movement speed. For example, the rotational speed of the lower leg in the standing phase can be used as the movement speed for the change in resistance.

The change in resistance occurs in real time, whereby an estimate can be made on the basis of the previous data or movement profiles as to what the future movements, speeds or loads might look like, so that an expected value can be preset and compared with can be compared with parameters measured in real time. Preferably, the speed and the other parameters or sizes are determined in the respective movement phase itself and as up-to-date as possible. The change in resistance depending on the movement speed can take place in preparation for the stance phase and/or during the stance phase. The adjustment can take place once per step or movement cycle, or the resistance can be adjusted multiple times or continuously. The adjustment of the resistance with the movement speed can be limited, in particular the adjustment from step to step or from movement cycle to movement cycle, or the rate of change over time. In one embodiment, a resistance level is increased with decreasing movement speed by reducing resistance later and/or more slowly. A later and/or slower reduction in resistance allows slower and more controlled movements to be carried out. Alternatively or additionally, an increase in resistance can take place earlier or more quickly with decreasing speed.

In a further development, the resistance is changed in such a way that when the speed is reduced by increasing the resistance, a movement is stopped earlier, a movement amplitude is reduced or a movement reversal is achieved earlier.

In one embodiment, the resistance, which is reduced with decreasing speed and/or increased with increasing speed, has a supporting effect in one direction of movement. For example, an actuator can apply an extension moment around a knee joint axis. This extension moment acts as a resistance against a bending movement and has a supporting effect during an extension movement. A springy or elastic resistance, which is generated, for example, by a force storage device or an active actuator such as an electric motor, also has a supporting effect in the opposite direction of movement. If the resistance is increased with decreasing speed and/or reduced with increasing speed, the supporting effect is also increased or reduced. The walking speed in particular can be used as the speed. In order to achieve a harmonious In order to realize a movement sequence at different movement speeds, it is sensible not only to change the resistance to the movement depending on the speed, but also to support a movement alternatively or in addition. For example, in the stance phase flexion and extension it is advantageous to first apply resistance to the flexion movement in order to achieve controlled flexion, and then to support the extension movement after the movement is reversed. At higher movement speeds the person using the exercise often generates higher forces and moments via the remaining limbs and joints. For example, a higher hip extension moment is applied in the stance phase to generate propulsion. The hip extension moment also influences the moment acting on the knee joint axis, in particular the external flexion moment is reduced and the extension moment is increased in the middle and late stance phase. Accordingly, a supporting knee extension moment in a middle stance phase can be reduced compared to a slower movement speed. The support can be adjusted, for example, by changing an internal stretching torque provided by an actuator such as an electric motor, or by adjusting a spring stiffness, a spring point, a gear ratio, a lever arm, additional damping, or by switching a force accumulator on or off. A lower spring stiffness, damping, or a lower lever arm, for example, leads to a reduction in resistance. It is also possible to change the resistance level with the speed by changing the resistance, for example a motor torque or spring stiffness, earlier or later depending on the speed. This also influences the level of support. In particular, in a movement phase that includes a reversal of movement, a resistance or force or moment is provided that acts against a movement and supports a movement in the opposite direction; the resistance provided during the movement phase is increased as the speed decreases and/or reduced as the speed increases. The resistance can be varied over the course of the movement.

In one embodiment, the state of a resistance device, which influences the resistance, is determined by the control as a function of the speed adjusted. With a hydraulic throttle valve, the flow cross-section can be changed, with a magnetorheological brake, the applied magnetic field, or with a friction brake, a braking force. It is possible that for a state of the resistance device, the resistance to a movement is not constant, but depends on the movement, as is the case with a throttle valve, for example, whose flow resistance depends on the flow rate. The resistance therefore results from both the characteristics of the resistance device itself and from its state, which is changed via the control system depending on the speed.

In one embodiment, the change in resistance occurs in at least one movement or movement phase that differs from the standing phase of level walking. In particular, a change in resistance in the standing phase is useful when walking up and/or down slopes, when walking up and/or down stairs and/or over individual steps or ledges, when walking backwards and/or sideways or even when making turning movements. Such a change in resistance is also advantageous when stopping and starting. Last but not least, a change in resistance can be used in a special mode that is suitable for special movements and sports such as cycling, skiing or scootering, especially when under load. For this purpose, the speed of movement is determined, in particular the walking speed or speed of travel, and if the speed of movement is reduced under load, the resistance is increased, or if the speed of movement is increased, the resistance is reduced. The standing phase can be a movement phase in which the orthopaedic device is loaded with part or all of the body weight.

In a further development, the resistance adjustment when walking downhill on an inclined surface and/or when overcoming height differences such as steps is carried out in such a way that a correlation of forward progression and downward movement in the standing phase of the treated side is adapted to the surface inclination or height difference recorded by sensors or depends on this. At a faster walking speed, it is advantageous to sink in more quickly or to move the body downwards at a given inclination or height difference. move. Based on a determined correlation of forward and downward speed, which is advantageous for the determined ground gradient, the resistance is increased as the speed of movement decreases, or reduced as the speed of movement increases. In addition to the correlation of the speeds, the correlation of the distances covered in forward and downward directions over a certain period of time can also be used to adjust the resistance. The quotient or another functional relationship can be used as a correlation of the speeds or distances covered.

The change in resistance can depend on the type of movement, the phase of movement, the surface, the environment and/or the current operating mode of the aid. The adjustment of the resistance can, for example, be designed differently for climbing stairs than for walking down slopes. It is also possible for the change in resistance and speed to be individually adjustable for the user, for example to adapt to body weight or personal preference. Such individual adjustment can be made via controls or an app. The type and/or extent of the change in resistance can also be adapted autonomously by the aid from step to step or during a movement in order to adapt to the user or to components of the orthopedic device, for example a shoe or cosmetics.

In one embodiment, the resistance is a force or moment characteristic depending on the movement and/or position of the actuator or components of the orthopaedic device, such as friction, damping, elasticity, spring force or the like. It can be either a linear or a non-linear behavior. A change in the resistance can be achieved by changing a friction value, damping, stiffness, zero point and/or the like. One or more parameters of such characteristics can also be adjusted in the sense of a change in resistance, for example the progressivity of a non-linear stiffness or damping. A resistance can also be achieved by a combination of several Characteristics are generated, for example a springy and a damping behavior, which act in series or in parallel. The characteristics, such as stiffness and damping, can also relate to degrees of freedom other than those between the upper and lower parts, for example the lower part angle in relation to the direction of gravity. For example, a knee-extending moment can be linearly related to the lower leg inclination in relation to the direction of gravity, which corresponds to a linear spring behavior, and the spring stiffness can be changed with the speed of movement. A resistance can also be a moment or force applied by an actuator; this moment or force can be changed depending on the speed of movement.

With an active electromechanical or piezoelectric actuator, resistance to a movement can be applied or movements can be supported using currents and voltages. Voltages and/or currents can therefore be changed depending on the speed of movement. In addition to applying moments and moment curves, control algorithms and sensor information can be used to track trajectories, control target variables or emulate system properties in the sense of impedance or admittance control via the pivoting movement of the upper and lower parts. For example, the behavior of a linear or non-linear spring, the behavior of a damper or inertia can be emulated, or a combination of several properties, whereby the state of movement of the upper and/or lower parts can be influenced. Such control offers a high degree of flexibility. Swiveling movements can also be actively supported using such controls. For example, if the actuator emulates a spring characteristic, a movement against the spring is initially resisted and if the movement is reversed, the movement is supported. For example, as the speed of movement increases, the stiffness of this spring characteristic can be reduced, thereby reducing both the resistance to a movement and the degree of support in the opposite direction of movement. Other control strategies can also be adjusted with the speed of movement so that the resistance decreases as the speed of movement increases. Actuators can also activate and deactivate energy storage devices, for example a hydraulic spring accumulator, change the gear ratios of drives and/or couple or lock them. These types of actuation can also be used to influence the swivel movement and change the resistance.

In a further development, the sensor signals are superimposed on a biosignal to change the resistance, whereby the biosignal is recorded by a human-machine interface (MMI) and transmitted to the control device. This makes it possible to adjust the resistance, in particular the flexion resistance, at will. The control via the MMI can also take place subconsciously and thus be integrated into the natural control of body movement by the central nervous system. The resistance can thus be determined both by the voluntary and arbitrary signal of the human-machine interface and by the sensor signals with regard to the speed of movement and, if applicable, the additional sensor signals. In one embodiment, the human-machine interface is a sensor device via which electromyographic signals are recorded and transmitted to the control device. By tensing the muscles in the area of the electrodes, the resistance to movement of the orthopedic device, for example the prosthesis or orthosis, is modulated, in particular increased. This can involve the tensing of a single muscle, several muscles, or one or more muscle groups. It can also be the intention of one or more muscle tensions or the influence of a movement, which is determined, for example, by recording nerve impulses in the central or peripheral nervous system. The deformation of tissues, the conductivity of tissues, the absorption, reflection and/or propagation of sound and/or electromagnetic waves in tissues, electromagnetic fields, electrical potentials, substance concentrations, as well as electrochemical gradients in tissues can also be used as biosignals, in particular this can be used to determine the control of one or more muscles. It is also possible that the MMI applies the method of pattern recognition, signal processing, classification and/or artificial intelligence to the recorded biosignals and the discrete and/or continuous values calculated from them to the control as biosignals. Different biosignals can be used in different contexts, environments, operating modes, movements and/or movement phases.

In one embodiment, the biosignal is superior to the movement speed signal in the control device in at least one movement phase, so that without the biosignal there is no change in resistance with speed. The biosignal thus serves as a so-called trigger for a change in resistance. The type of change is then changed and adapted by evaluating the movement speed, possibly in conjunction with other sensor variables and variables derived from sensor data. It is also possible for the movement speed to determine the extent to which the resistance can be influenced or modulated via the biosignal. At low speeds, the resistance can be adjusted by the biosignal up to a complete block, while at increasing movement speeds only a small increase is possible. If a threshold value for the movement speed is exceeded, the resistance can also be completely prevented by the biosignal. Alternatively or additionally, the sensitivity of the resistance change can be changed depending on the biosignal based on the speed, in particular in the sense of lower sensitivity at higher speeds. The biosignal is particularly suitable for defining the time and/or duration of the change in resistance, while the movement speed determines the type and/or extent of the change in resistance, possibly together with other sensor values and indicators. In one embodiment, the biosignal triggers a resistance adjustment and maintains this over a certain period of time or movement phase, even if control via the biosignal no longer occurs. In addition to continuous modulation, the biosignal can also be used to discretely adjust a resistance level, which is superimposed on a continuous or discrete adjustment with the movement speed. Depending on the situation, the movement sequence, the movement phase or the operating mode, control via the biosignal can be designed differently or deactivated. In one embodiment, the movement speed is averaged over a movement phase and the averaged movement speed is used for control.

The translational speed in one or more directions of at least one component of the orthopedic device, the torso, the center of gravity and/or the contralateral side can be determined from sensor values and used as the speed of movement, in particular a speed component parallel to the ground or in the horizontal direction.

The resistance is applied to a movement in at least one movement phase, whereby in one embodiment the resistance in at least one movement phase counteracts a bending movement and/or supports a stretching movement. The resistance in at least one movement phase actively supports a movement when the direction of movement changes while the direction of action remains the same. The resistance can be adjusted in a movement phase with a reversal of movement, whereby the resistance counteracts a movement in a first direction of movement and supports a movement in the opposite direction of movement.

An adjustment of the resistance with the speed of movement can take place in particular when walking down one or more steps, when walking down slopes and/or when walking down a height difference. In one embodiment, the resistance is also adjusted with the speed of movement when braking and/or stopping after a movement.

The resistance can be adjusted when overcoming differences in height, in particular when climbing one or more steps and/or climbing on inclined surfaces.

In one embodiment, the resistance is increased with decreasing movement speed until a lock or until the movement of the upper part and lower part relative to each other stops. In one embodiment, the resistance is a linear or non-linear, elastic and/or damping behavior depending on the pivoting movement between the upper part and the lower part and/or on movements and/or loads detected by sensors.

A change in resistance can be made by adjusting one or more parameters of an elastic and/or damping behavior. The resistance can be changed with the speed of movement so that the horizontal and vertical speed of movement and/or a horizontal and vertical path traveled correlate with each other in at least one movement phase, in particular in a ratio that depends on the slope of the ground determined by sensors or a height difference to be overcome. The correlation of the resistance with the speed of movement can be changed depending on the operating mode, the movement, the movement phase and/or the ground.

The embodiments and further developments of the method can be combined with one another and/or with other control methods and embodiments, for example with control methods known from the prior art. Examples of the invention are explained in more detail below with reference to the attached figures. They show:

Figure 1 - a schematic representation of a prosthetic leg;

Figure 2 - a schematic movement sequence and a speed determination;

Figure 3 - a resistance curve versus speed;

Figure 4 - a resistance and velocity curve over time;

Figure 5 - different resistance curves over time;

Figure 6 - a relationship between biosignal and resistance curve;

Figure 7 - a variant of Figure 6;

Figure 8 shows a design of the control system; and

Figure 9 - different control characteristics of an actuator

Figure 1 shows a schematic representation of a prosthetic knee joint as part of a prosthesis. The prosthetic knee joint has an upper part 10 and a Lower part 20, which are pivotably mounted on one another about a pivot axis 15. A prosthetic foot 60 is arranged on the lower part 20 at the distal end. In the design as a prosthetic leg according to Figure 1, a prosthetic shaft or another device for receiving a thigh stump or for securing to a person is arranged or formed on the upper part 10. A resistance device 30 as a linearly acting hydraulic actuator is arranged between the upper part 10 and the lower part 20. In the exemplary embodiment shown, the hydraulic actuator 30 is designed with a hydraulic chamber or a cylinder, which is arranged or formed in a housing or base body 31. A piston 32 is displaceably mounted in the cylinder. The piston 32 can be displaced along the longitudinal extent of the cylinder and is fastened to a piston rod 33, which protrudes from the housing or base body 31. The piston 32 divides the cylinder into chambers that are fluidically connected to one another via a hydraulic line. The base body 31 or the housing can be pivotably mounted on the lower part 20 at a fastening point 23 in order to prevent the piston 32 from tilting during a pivoting movement of the upper part 10 relative to the lower part 20. The end of the piston rod 33 facing away from the piston 32 is attached to the upper part 10, in the embodiment shown on a boom to increase the distance to the pivot axis 15, at an upper fastening point 21. During flexion, the piston 32 is pressed downwards so that the volume of a flexion chamber decreases, and correspondingly the volume of an extension chamber increases, reduced by the volume of the retracting piston rod 33. An electric motor can be arranged in the housing 31 to generate a pressure within one of the chambers, which drives a pump (not shown) in order to apply pressure to the hydraulic fluid within one of the two chambers and thereby move the piston 32 within the cylinder in one direction or the other. This causes a flexion movement or an extension movement of the orthopedic device in the form of the prosthetic leg. The electric motor for driving the pump is an option that can be used in one embodiment in combination with the hydraulic actuator 30. In principle, a drive or motor is not necessary for a passive prosthetic knee joint. An alternative design to the design of the actuator 30 as a passive linear damper, in particular linear hydraulics, is a Rotational damper, in particular a rotational hydraulic system, a magnetorheological resistance device or an electric motor, in particular in combination with a gear or a spindle drive. The electric motor can be operated in generator mode. A combination of several of the aforementioned resistance devices as an actuator is also implemented in one embodiment.

A drive 34 is arranged inside the housing 31 or on the housing 31 and is coupled to at least one control valve 35, via which the hydraulic resistance in the actuator 30 can be changed. The actuator 30 and in particular the drive 34 is coupled to a control device 40 which activates, deactivates or modulates the drive 34 on the basis of sensor values in order to be able to provide an adapted resistance through the passively designed actuator 30. When the actuator 30 is designed as a magnetorheological resistance device, the resistances are changed by activating, deactivating or modulating a magnetic field, the drive 34 is then the electromagnet or the magnetic coil. When the actuator 30 is designed as an active drive with an electric motor, the resistance is changed by activating, deactivating or modulating voltages which influence the torque generated by the electric motor.

At least one sensor 50 for detecting the spatial orientation of the lower part 20 or the upper part 10 is arranged on both the upper part 10 and the lower part 20. In particular, the sensor 50 for detecting the spatial orientation is only arranged on the upper part 10. This sensor 50, which can be designed as an IMU (inertial measurement unit), for example, is used to determine the spatial angle or the absolute angle to a fixed spatial orientation, for example the direction of gravity, during use of the prosthetic knee joint. Instead of being designed as an IMU for detecting spatial positions, the corresponding sensor 50 can also detect other status data, in particular status data relating to the artificial knee joint. In particular, positions, angular positions, speeds, accelerations, forces and their progressions or changes are recorded as status data. The determined spatial angle of the upper part 10 and/or the lower part 20 or another status variable is compared with a threshold angle. When a threshold value is reached or exceeded, the is stored in a control device 40 for the respective sensor value or a variable derived therefrom, the drive 34 is modulated, activated or deactivated in order to otherwise change the flow resistance in the actuator 30 in the design as a hydraulic damper, the viscosity, the braking force, the torque, the stiffness or the force counteracting the flexion movement.

The actuator 30 in an artificial knee joint is usually used to modulate a flexion movement and an extension movement in order to generate or support an appropriate or desired movement sequence. An extension movement is supported if necessary and advantageously braked shortly before reaching maximum extension in order to avoid a hard impact. A flexion movement is braked or prevented in the stance phase and in the swing phase in order to ensure that the flexion is limited. In order to be able to drive the drive 34 to actuate the control valve 35, an energy storage device, in particular in the form of an accumulator, is also assigned to the drive 34. The energy storage device can be arranged directly next to the drive 34 or at another location in the orthopedic device where more space is available or where this appears advantageous due to the weight distribution.

Furthermore, the control device 40 and at least one angle detection device as a sensor 50 are arranged on the prosthesis or orthosis. The angle detection device 50 detects the angle between the upper part 10 and the lower part 20 and is designed, for example, as a direct angle sensor that detects the angle directly. Alternatively, the angle between the upper part 10 and the lower part 20 can be determined by evaluating the sensor data from two spatial position sensors 50. Both methods can also be used simultaneously or in addition to one another. All sensors arranged on the prosthesis or orthosis are coupled to the control device 40 and their sensor values serve as the basis for controlling the drive 34 of the actuator 30 if this is designed as a damper, or as input signals for a motor control if the actuator 30 is designed as a motor. In the case of magnetorheological damping, the sensor values are used to control the magnetic field or its change. Based on the sensor data, in particular the spatial positions and/or angular positions as well as position data and data on the load, orientation, acceleration and/or deformation of other components, the drive 34 is controlled in order to reduce or increase a swivel resistance by the actuator 30.

Furthermore, a human-machine interface 100 is assigned to the prosthesis, via which a biosignal can be transmitted from the human to the control device 40. The human-machine interface (MMI) 100 as a switch or sensor can be housed in a separate component or, for example, be part of the prosthesis. The MMI transmits a biosignal to the control device 40 wirelessly or via a wire and thereby triggers, for example, an increase in the flexion resistance at the beginning of the stance phase, whereby the increase in the stance phase is modulated by the evaluation of the sensor values of the sensors 50 or at least one of the sensors 50. The MMI 100 therefore only gives the initial signal, which is then adjusted sensor-supported and autonomously as the movement progresses.

Figure 2 shows a schematic representation of a prosthetic leg with an upper part 10 in the form of a prosthetic shaft with a joint component arranged thereon and a lower leg part as a lower part 20, which are arranged in an articulated manner to one another. An MMI 80 in the form of a sensor arrangement for detecting myoelectric signals is housed within the upper part 10. If, for example, contractions of thigh muscles are measured, this biosignal, which is detected by the MMI 80, is transmitted to the control device 40 and serves to trigger the change in resistance. Figure 2 also shows a leg tendon 70 which extends proximally from the hip pivot point 71 to a foot point 72. The change in resistance via the actuator is carried out depending on the walking speed in the stance phase. The walking speed is the movement speed of the entire body in the stance phase from the heel strike, which is shown at the time tO, to the end of the stance phase. The time t1 represents a time in the terminal stance phase. The speed of movement is not necessarily constant in the stance phase, but can vary over time. In this case, the hip or the hip pivot point 71 moves forward in the walking direction, while the prosthetic foot remains on the ground and thus the foot pivot point 72 also essentially remains stationary. The walking speed can can be determined from the distance traveled between two points in time, for example the distance Ax in the time between tO and t1 . In addition to the average speed over a longer period of time via the secant S, the instantaneous speed can also be determined from the movement via the tangent T, which is shown in the right-hand illustration of Figure 2.

An adjustment of the resistance to a displacement of the lower part 10 relative to the leg 20, in particular to flexion in the stance phase, takes place in particular as a function of the speed at which the leg is moved. The adjustment takes place in particular as a function of the walking speed. Resistance to flexion can be a passive resistance in which, for example, kinetic energy is converted into thermal energy via damping, or a force exerted against the movement by an active actuator. A resistance acting against a displacement in one direction of movement can have an actively supporting effect in the opposite direction of movement. In addition to the speed of movement, in particular the walking speed, other translational or rotational speeds, for example of the upper part 10 or the lower part 20, can also be used to characterize a speed of movement. Such other speeds of movement can also occur when a movement is carried out that deviates from walking.

The flexion resistance in the knee joint in the stance phase is reduced with increasing movement speed. Figure 3 shows the relationship between the average level of movement resistance R' and the movement speed v. The average level of movement resistance R' is reduced with increasing walking speed v, whereby at low walking speed v and a comparatively high walking speed v saturation occurs and the curve approaches a limit value. Reducing the flexion resistance in the stance phase with increasing speed is also useful when walking down ramps, for example. The slower the walking speed, the longer the stance phase lasts and the smaller the step length. The same applies to walking on level ground. Accordingly, it is advantageous to reduce the flexion resistance when walking. When walking slowly, it is best to choose a high resistance so that the descent is slow and the user of the prosthesis or orthosis does not have to laboriously raise the body's center of gravity. When walking quickly and in a short stance phase, it is advantageous to reduce the flexion resistance so that the user can sink deeper, e.g. to absorb the body's momentum. When walking on level ground at a high walking speed, reducing the flexion resistance not only takes into account the shorter stance phase, but also the stronger driving hip moment. When braking, it is also advantageous to choose a lower resistance at higher speeds, for example when the lower leg and/or leg tendon are rotating quickly forwards, as the body's high kinetic energy must be absorbed. At high speeds, it is advantageous to brake the body over a greater distance, i.e. over a larger flexion angle, with less resistance, while at lower speeds, faster, more precise braking is advantageous. Since the speed is continuously reduced when braking, this can lead to a gradual increase in resistance if the resistance changes dynamically with the speed. The feedback between resistance and speed with an increase in resistance at a reduced speed then leads to a progressive braking and a progressive resistance curve in the movement sequence, which can also smoothly transition into a blocking or into the standing function with a very strongly increased flexion resistance. This can enable relaxed standing with the joint bent.

Figure 4 shows an example of the movement speed v and the resistance R in a dynamic coupling during a braking step. Starting from an initial speed v and an initial resistance R, the body is braked. This reduces the speed v, which in turn leads to an increase in the resistance R, resulting in a progressive reduction in the speed v and thus a progressive increase in the resistance R, until the position used comes to a standstill or the joint is locked, for example.

Raising the resistance level R' depending on the speed of movement can be done in different ways, two of which are shown in Figure 5. In the left-hand illustration, the resistance curve is purely increased with decreasing speed v, whereby the characteristic or the curve as such remains essentially unchanged. The dependence of the resistance level as a function of the speed can be selected similarly to the curve shown in Figure 3. In the right-hand illustration, the shape of the curve of the average resistance level R' is changed; in the exemplary embodiment shown, the rate of change and the time of the resistance change are changed. The change in the resistance level can depend on an operating mode of the orthopedic device, the type of movement or the movement phase. It is also possible that a change in the resistance level is only made in certain operating modes, movements or movement phases. In Figure 5 on the right, different dependencies of the resistance level R' on the speed v are shown for two different movement phases, marked as R1'(v) and R2'(v), whereby in a middle movement phase, no adjustment of the resistance level with the speed is made.

In addition to walking on level ground, walking on ramps and inclined surfaces, over steps or when braking, adjusting the flexion resistance depending on the speed of movement of either the person or a part of the orthosis or prosthesis can be useful in situations that deviate from walking, for example when kneeling, sitting down or squatting. For example, the shift in the body's centre of gravity or the speed of a swivelling or rolling movement forwards or backwards can be used as the speed of movement and an increase in the speed of movement can initiate a reduction in the flexion resistance.

The adaptation of the resistance R can also refer to a maximum, an average value, instantaneous or integral values or another characteristic value. If a resistance level is changed with the speed, the resistance does not have to be constant over the course of the movement, but can follow a course that depends in particular on other sensor values. For example, an increase in the resistance depending on the speed of movement can occur earlier or later in the The movement sequence can be adjusted to be faster or slower. By changing the ratio of areas with lower resistance to areas with higher resistance, the overall level is changed. It is also possible to increase the flexion resistance to the point where the joint is locked, whereby the point in time or threshold at which a lock is triggered can be made dependent on the speed of movement. In addition to a lock, an actuator can apply a moment or force that is opposite to and equal to the external moment or force, so that a balance of forces is achieved and movement between the upper and lower parts is stopped or prevented.

In addition to the speed, other determined variables can be used to optimally adapt the resistance to the movement situation, for example the joint angle, a segment angle, a leg tendon angle relative to a reference plane, a ground reaction force vector, a parameter of the contralateral side, the time derivatives of these variables or a combination thereof. For example, a resistance can be increased earlier at an increasingly lower movement speed, i.e. at a small knee angle. Both the level and the time or rate of change in the resistance can be changed with the movement speed. The adjustment of the level or the course can be dynamic, so that the resistance can be changed continuously within a step in terms of both the level and the course. Alternatively, only one adjustment can be made per step, so that the level does not change within a step. Alternatively, the resistance can only be changed in an initial phase of the movement and kept constant in a later phase.

The speed of movement is preferably the walking speed of the person using the device. This can be the speed of the trunk, upper body, center of gravity or hips. In addition to a forward speed, a speed in a backward or sideways direction can also be used. The speed of a rotational movement around the longitudinal axis or the rate of change a direction of progression can be used as the speed. Alternatively or in addition, other reference points or reference elements can be used to record the speed of movement. The speed of movement, in particular the walking speed, can be calculated or estimated from one or more segment angular velocities, for example from the forward rotation of the lower leg in the stance phase. Movement speeds can also be calculated from angular velocities, for example based on the sensors in the joints, and the known segment lengths, such as the length of the lower leg or the thigh part. The movement speed of the leg tendon and the translational speed of the torso and the hip can also be calculated or estimated in this way. It is possible to use only one component of the speed as the movement speed, for example the horizontal movement component or the component that is oriented parallel to the ground. If a time interval is relevant for a movement, it can be based on typical events during the movement or during a movement sequence; for example, when walking, this would be the heel strike or initial contact, the toe-off or the knee break. For example, by determining the hip position via the leg tendon at the time of the initial contact and at the current time as well as the elapsed time, an average walking speed in the stance phase can be determined. The movement speed is therefore the average walking speed in the stance phase. The walking speed can also be estimated from a relative angular velocity, for example the maximum knee angular velocity in the previous swing phase. The speed in the swing phase can be estimated by integrating accelerations. This is particularly useful if the resistance is to be adjusted to the walking speed at the time of the initial contact or the heel strike, or is adjusted during the stance phase flexion. Due to the high dynamics of the initial contact, an estimate from the kinematic parameters recorded in the sensors is difficult at this time. To avoid abrupt transitions in the determined speed, it can be filtered or smoothed between different methods. Another possibility is to estimate the speed of movement or walking speed from the cadence or other gait parameters such as stance phase duration, swing phase duration, stride length or stride duration.

In addition to adjusting the resistance based on the speed of movement, a biosignal B from an MMI 80, 100 can also be incorporated into the control system. Such a combination is shown in Figure 6. In the upper diagram, the resistance curve R is plotted against time, and in the lower diagram, the curve of the biosignal B is plotted against time. If there is no change in the biosignal B or if there is no biosignal B at the control device 40, as shown by the dashed curve in the lower curve, for example because a muscle is not tensed, this is considered a signal not to change the resistance depending on the speed of movement, which is shown by the dashed, straight-line curve of the resistance R in the upper curve. The resistance R is therefore at a constant level in the control system shown and does not change across the different speeds. However, the resistance curve can also be changed based on other sensor signals. However, if a change in the biosignal B is generated via the MMI 80, 100, which is indicated by the increase in the solid line of the biosignal B, this is an input variable for the control device 40 in order to change the resistance R, in particular to increase it. The basic increase in the resistance R is triggered by the biosignal B, the level of the increase in resistance is determined by the respective speed v of the patient, the prosthesis, the orthosis or one of the components of the orthosis or prosthesis. An increase in resistance takes place between tO and t1. At a high speed v1, the resistance R is increased less than at a lower speed v2. Without the biosignal B, no speed-dependent adjustment of the resistance curve takes place in the embodiment shown. It is particularly expedient to generate the biosignal B via electrodes that are arranged on the patient and record a muscle contraction or pick up nerve signals. In one variant of the control, the resistance level is modulated via the at least one biosignal, for example via a proportional change in the resistance depending on the biosignal, whereby the amplitude of the modulation or the sensitivity of the control depends on the speed of movement, as in the solid resistance curves in Figure 6. Alternatively or additionally, the maximum increase in resistance by the biosignal can also depend on the speed of movement, thereby limiting the increase in resistance. Up to the respective maximum resistance or resistance level, the resistance can be controlled proportionally by the biosignal or a corresponding modulation can take place. In addition, the resistance level or resistance curve can also be adjusted based on the speed of movement without control by the MMI or without the biosignal, whereby the type and manner of resistance adjustment can be designed differently depending on the speed of movement with and without control by the MMI, or depending on the biosignal.

Figure 7 shows a design of a control system in which the biosignal B acts as a trigger for the increase in resistance and takes precedence over the adjustment of the resistance level or resistance R. If no biosignal is present, which is shown by the dashed line, the resistance R is at an initial level, which is also shown by the dashed line. If control is via the MMI and the biosignal B exceeds a threshold value BO or there is a minimum level of control via the MMI, the resistance R is adjusted depending on the speed of movement. This is shown in the range between tO and t1. If the biosignal is below the threshold value BO or falls below it again, for example because control is no longer taking place via the MMI, the resistance R remains at the initial level or is reduced to it again. In the variant shown, a variation of the biosignal above the threshold value BO, shown by the solid line in the range tO to t1, does not cause any further modulation of the resistance R. The resistance level in this range is determined by the movement speed and may also depend on other sensor information, for example a knee angle and an axial force. When the movement speed increases from v2 to v1, the resistance level is reduced and when the movement speed decreases, it is increased.

The variants of control via an MMI shown in the figures are carried out at In the presence of a biosignal and a sufficiently low speed, the resistance increases compared to the initial level. However, it is also possible that the resistance level is reduced compared to the initial level by the biosignal, which is superimposed with an adjustment of the resistance level to the speed of movement, or that the extent to which the resistance level is adjusted by the biosignal depends on the speed of movement.

Figure 8 shows a design of the control in a stance phase flexion and stance phase extension, for example when walking on level ground, walking uphill on an incline or climbing stairs. The curve of the knee angle cpK and the internal knee moment MK is shown, both for a first movement speed v1 and a lower movement speed v2. In the movement sequence shown, the low movement speed is also accompanied by a longer movement duration. An increasing knee angle corresponds to knee flexion, a positive knee moment MK to an internal flexion moment. In the special design, the knee moment is generated by an active actuator. The actuator is controlled in the range from tO to t1, or from tO to t2 at reduced speed v2, so that the knee moment is changed as a function of the knee angle according to a linear torsion spring characteristic that has its neutral point at time tO, whereby no moment is generated at this time. Such a characteristic can be achieved, for example, by an electric motor based on a knee angle signal from the control device. During the stance phase flexure, the extension moment generated by the actuator increases according to the spring characteristic and provides resistance to the flexure movement. After the movement is reversed, the moment is reduced again until the neutral point is reached, which is the case at times t1 and t2. In the embodiment shown, a further extension is counteracted by the actuator with a bending moment in order to harmoniously stop the extension movement. As the speed of movement decreases, the stiffness of the spring characteristic, which is realized via the actuator, increases, so that a higher stiffness results for the speed v2 than for v1. This can be seen from the fact that at speed v2, despite a lower maximum knee flexure angle, a higher maximum extension moment results. This means that the Resistance increases with decreasing movement speed. The moment generated by the actuator according to the spring characteristic stored in the control system, which acts as resistance to the bending, has an actively supportive effect after the movement is reversed. By adapting the resistance in the form of a torsion spring characteristic to the movement speed, the extent of the support is also adapted. Adapting the resistance to the movement speed in this way enables a particularly harmonious movement sequence.

Figure 9 shows different control characteristics of an actuator in the form of a rotary electromechanical drive, which act as resistance and influence a pivoting movement of the upper part 10 to the lower part 20 in a standing phase. The characteristics are shown as the relationships between a degree of freedom or a sensor signal <p of the orthopedic device, here in the form of the knee angle, or its rate of change w, and the moment T generated by the actuator, based on the pivoting angle between the upper part 10 and lower part 20. An increasing knee angle corresponds to knee flexion, a positive moment counteracts a flexion movement. Shown on the left is a linear elastic relationship between the generated knee moment and the knee angle. The resistance to flexion is reduced with increasing movement speed by shifting the zero crossing on the horizontal coordinate axis towards a larger knee angle. With the same knee angle, this results in a lower extension moment or even a flexion moment. In the middle, a linear-elastic relationship is also shown, whereby the gradient, or stiffness, is reduced with increasing movement speed, which, for example, reduces the resistance to bending. The linear-elastic characteristic enables the actuator to support a stretching movement. On the right, a non-linear damping characteristic is shown, whereby the damping coefficients are reduced with increasing movement speed, which results in less resistance to movement when the knee angle speed w is kept constant compared to a slower movement speed. In addition to implementation via an electromechanical actuator, such characteristics can also be implemented via one or more springs, hydraulic or pneumatic dampers, magnetorheological resistance devices or the like as well as their combination.

The statements regarding prostheses also apply to orthoses, particularly those that span the knee joint. The control makes walking at different walking speeds easier and more comfortable for the user.

Claims

patent claims 1. Method for controlling an orthopedic joint device of a lower extremity with an upper part (10) and a lower part (20), which are pivotably mounted to one another about a pivot axis (15), with an actuator (30) which is coupled to the upper part (10) and the lower part (20) and influences a state of movement of the upper part (10) and/or the lower part (20), wherein the actuator (30) is coupled to a control device (40) which is coupled to at least one sensor (50) and activates, deactivates or modulates the actuator (30) on the basis of sensor values of the at least one sensor (50), wherein at least one movement speed of at least part of the orthopedic joint device is determined from the sensor values and the actuator (30) is activated, deactivated or modulated on the basis of the movement speed in the stance phase, characterized in that the movement speed in the stance phase and the resistance at least for part the movement are inversely correlated. 2. Method according to claim 1, characterized in that with a decreasing movement speed the resistance is increased and with an increasing movement speed in the standing phase the resistance is reduced. 3. Method according to claim 1 or 2, characterized in that a change in the resistance only occurs from a specified threshold value and/or up to a specified threshold value. 4. Method according to one of the preceding claims, characterized in that the change in resistance is non-linear. 5. Method according to one of the preceding claims, characterized in that the change in resistance is carried out by changing the resistance level with a constant or changed resistance curve. 6. Method according to one of the preceding claims, characterized in that the change in resistance is changed as a function of forces, angles, positions and/or moments which are determined by sensors (50) or which are calculated from sensor values. 7. Method according to one of the preceding claims, characterized in that a change in the resistance curve and/or resistance level takes place separately for each stance phase. 8. Method according to one of the preceding claims, characterized in that the speed of movement, in particular the walking speed, is calculated via at least one measured angular velocity and a known leg tendon length and is used as the speed of movement of the change in resistance. 9. Method according to one of the preceding claims, characterized in that the change in resistance occurs in real time. 10. Method according to one of the preceding claims, characterized in that a biosignal is superimposed on the sensor signals for changing the resistance, which biosignal is detected by a human-machine interface (80,100) and transmitted to the control device (40). 11. Method according to claim 10, characterized in that the biosignal is superior to the movement speed in at least one movement phase and without a biosignal no change in the resistance as a function of the movement speed occurs. 12. Method according to claim 10 or 11, characterized in that the time of the change in the resistance and the The type of change in resistance is determined by the speed of movement. 13. Method according to one of the preceding claims, characterized in that the movement speed is averaged over a movement phase and the averaged movement speed is used for the control. 14. Method according to one of the preceding claims, characterized in that the translational speed in one or more directions of at least one component of the orthopedic device, the torso, the body's center of gravity and/or the contralateral side is determined from sensor values and is used as the movement speed, in particular a speed component parallel to the ground or in the horizontal direction. 15. Method according to one of the preceding claims, characterized in that the resistance is applied in at least one movement phase of a movement 16. Method according to one of the preceding claims, characterized in that the resistance counteracts a bending movement in at least one movement phase and/or supports a stretching movement 17. Method according to one of the preceding claims, characterized in that the resistance actively supports a movement in at least one movement phase 18. Method according to one of the preceding claims, characterized in that the resistance is adjusted in a movement phase with reversal of movement. 19. Method according to claim 18, characterized in that the resistance is opposed to a movement in a first direction of movement and a movement in the opposite direction of movement is supported. 20. Method according to one of the preceding claims, characterized in that the resistance is adapted to the speed of movement when walking downhill over one or more steps, when walking downhill on slopes and/or when walking downhill over a height difference. 21. Method according to one of the preceding claims, characterized in that an adaptation of the resistance with the speed of movement takes place when braking and/or stopping from a movement. 22. Method according to one of the preceding claims, characterized in that an adaptation of the resistance takes place when overcoming height differences, in particular when climbing one or more steps and/or when climbing on inclined ground. 23. Method according to one of the preceding claims, characterized in that the resistance is increased with decreasing movement speed until a blockage or until the movement of the upper part (10) and lower part (20) relative to each other stops. 24. Method according to one of the preceding claims, characterized in that the resistance is a linear or non-linear, elastic and/or damping behavior depending on the pivoting movement between the upper part (10) and the lower part (20) and/or on movements and/or loads detected by sensors (50). 25. Method according to one of the preceding claims, characterized in that a change in the resistance takes place by adjusting one or more parameters of an elastic and/or damping behavior. 26. Method according to one of the preceding claims, characterized in that the resistance is changed with the speed of movement such that the horizontal and vertical speed of movement and/or a horizontal and vertical path traveled correlate with one another in at least one phase of movement, in particular in a ratio which depends on the ground slope determined by sensors (50) or a height difference to be overcome. 7. Method according to one of the preceding claims, characterized in that the correlation of the resistance with the movement speed is changed depending on the operating mode, the movement, the movement phase and/or the surface.