RU2348380C2 - Prosthesis foot with adjusted functional characteristics - Google Patents

Abstract

FIELD: medicine; orthopaedics and traumatology.

SUBSTANCE: prosthesis inferior limb contains foot, ankle joint, fixed leg and artificial muscle. Fixed leg is bent with convexity forward from lower end of fixed leg to upper end of fixed leg. The fixed leg is longitudinally bendable during step to accumulate and release energy for enhancement of prosthesis dynamic reaction during step. The artificial muscle accumulates energy during application of power load to prosthesis within active human take-off step phase and position phase during step at later stages, release of said energy to intensify taking-off driven limb and human body. Method for generating kinetic energy for taking-off force in elastic prosthesis foot involves the stages as follows: two concavity located in saggital plane are extended including backward concavity of fixed leg located in saggital plane thus to accumulate energy in prosthesis during application of power load to prosthesis within active human take-off step phase. During application of power load to prosthesis within active human take-off step phase, additional energy is accumulated in artificial muscle. At later stages of position phase during step, additional energy is released to intensify taking-off driven limb and human body additionally.

EFFECT: improved dexterity of sportsmen with the amputated limbs.

29 cl, 54 dwg

Description

FIELD OF THE INVENTION

The present invention relates to a prosthetic foot with high functional characteristics, providing improved dynamic response capabilities to the extent that these capabilities depend on the applied mechanics of the forces.

Technology Overview

A jointless artificial foot for a foot prosthesis is proposed in US Pat. No. 5,897,594 to Martin et al (Martin et al.). Unlike previous solutions in which the artificial foot has a rigid structure equipped with a hinge to simulate the function of the ankle joint, Martin et al., The jointless artificial foot, uses an elastic foot insert that is located inside the molded foot. The insert has a longitudinal section approximately C-shaped, open back, accepts the prosthetic load with its upper C-lobe and with its lower C-lobe transfers this load to a flat spring connected to the insert. In the bottom view, the flat spring has a curved shape and extends approximately parallel to the sole area forward of the foot insert into the area of the tip of the foot. The invention of Martin et al. Was created to improve the articulated artificial foot in terms of shock absorption of the heel, elasticity, rolling from heel to toe when walking and lateral stability, thereby giving the user the ability to walk naturally, while the goal is to provide the user with the opportunity to walk normally and do physical exercises and play sports games. However, the dynamic response characteristics of this known artificial foot are limited and do not imitate the human biomechanical function of the human foot, ankle joint and lower leg and soft supporting tissue. Martin’s artificial foot and co-authors and other well-known foot prostheses that use the ankle joint design mentioned above and a rigid pylon of a foot prosthesis imitating the lower leg are not able to accumulate enough elastic deformation energy to generate normal dynamic power during the step in the sagittal plane of the ankle joint. Tests have shown that known foot prostheses with the mentioned structures produce during the step only about 25% of the normal kinetic energy in the sagittal plane of the ankle joint.

Van L. Phillips has proposed other prosthetic foot prostheses, which he claims to provide the amputee with maneuverability and mobility when engaged in various activities that were excluded in the past due to design limitations and the corresponding functional characteristics of known prostheses. According to the inventor, these well-known feet steadily withstand running, jumping and other activities and, according to available data, allow the user to engage in the same activities as a regular foot. See, for example, US Pat. Nos. 6,071,313; 5,993,488; 5,899,944; 5,800,569; 5,800,568; 5,728,177; 5,728,176; 5,824,112; 5,593,457; 5,514,185; 5181932 and 4822363. These prostheses contain a foot, ankle joint and lower leg made of composite material, while the ankle joint is mechanically facing back with a convex bend. Tests have shown that known prostheses of this design produce about 40% of normal kinetic energy during the step in the sagittal plane of the human ankle joint. There is a need for a prosthesis with higher functional characteristics, which can increase the amputee's motor performance in activities such as walking, running, jumping and jerking.

SUMMARY OF THE INVENTION

In order to increase the motor capabilities of an athlete with an amputated leg, a foot prosthesis with high functional characteristics and improved applied mechanics is required, and it is necessary that the foot can exceed the functional characteristics of the human foot, as well as known prosthetic feet. An athlete with an amputated leg is interested in such a prosthetic foot with high functional characteristics, improved applied mechanics, combined high / low dynamic response and the ability to adjust the relative position, which can be fine-tuned to improve the horizontal and vertical components of the movement, which can be specialized in accordance with the character specific natural task.

A prosthetic foot in accordance with the present invention is intended to satisfy the aforementioned needs. In accordance with an illustrative embodiment described herein, a prosthetic foot according to the present invention comprises a longitudinally extending foot keel with a portion of the forefoot at one end, a portion of the hindfoot at the opposite end and a relatively long portion of the midfoot extending between the forefoot portions department of the foot and hindfoot and curved upward from these areas. A tibial stand is also provided, comprising a lower end curved downwardly by a bulge. An adjustable fastener attaches the curved lower end of the calf shank to the arcuate upwardly curved portion of the midfoot belonging to the foot keel to form the ankle joint of the prosthetic foot.

Adjustable mounting device allows you to adjust the location of the legs of the leg and foot keel relative to each other in the longitudinal direction of the foot keel to adjust the functional characteristics of the prosthetic foot. By adjusting the location of the upward arcuate section of the middle part of the foot belonging to the foot keel and opposing it with the lower end of the lower leg strut curved by convexity downward relative to each other in the longitudinal direction of the foot keel, changes in the dynamic response characteristics and the resulting foot movements according to a specific task are achieved depending on the required / desired horizontal and vertical linear speeds. A universal prosthetic foot is proposed, which has the capabilities of high and low dynamic response, as well as two-plane kinematic characteristics, and which increases the resulting functional capabilities of amputees involved in sports and / or recreational activities. A description of a prosthetic foot specifically designed for sprinting is also provided.

The prosthetic foot may also include a device designed to limit the extent of movement of the upper end of the calf in response to the application and removing the load on the calf during the use of the prosthetic foot. In one embodiment, the device is a piston assembly attached between the upper and lower ends of the calf shank and containing at least one pressurized fluid to limit travel and to absorb energy that is accumulated or released during compression and expansion racks of a shin. In other embodiments, the back shin device accumulates its own potential energy during application of the force load on the prosthesis and returns the accumulated energy during removal of the force load in addition to the possibility of accumulating the total energy of elastic deformation, which increases the dynamic power of the pushing force produced by the prosthetic foot during the step . In still other embodiments, an artificial muscle is provided on at least the foot, ankle and prosthesis stance to increase the potential energy of the prosthesis and the dynamic power in the sagittal plane of the ankle during the step. The artificial muscle on the prosthesis can also extend to an adjacent support structure on the user's stump, for example, to the receiving sleeve or to the knee clip of the prosthesis, or the said muscle can be attached to the proximal end of the human knee or prosthesis.

Mentioned and other objectives, features and advantages of the present invention are apparent from a consideration of the following detailed description of embodiments of the present invention and the accompanying drawings.

Brief Description of the Drawings

Figure 1 is a schematic representation of two adjacent radii of curvature R ₁ and R ₂ pressed against each other by the foot keel and lower leg of the prosthetic foot in accordance with the present invention, which provides dynamic response and the resulting movement of the foot during the step in the direction of arrow B, which is perpendicular to the tangent line A connecting the two radii.

FIG. 2 is a view similar to that of FIG. 1, but showing the relative position of the two radii, changed in the prosthetic foot in accordance with the present invention, in order to increase the horizontal component and reduce the vertical component of the dynamic response and the resulting foot movement during the step, so that arrow B ₁ , perpendicular to the tangent line A ₁ , is directed closer to the horizontal than in the case shown in figure 1.

Figure 3 is a side view of a prosthetic foot in accordance with an example embodiment of the present invention with an adapter of a pylon of a prosthetic leg and a pylon of a foot prosthesis attached to this adapter to secure the foot to the amputee stump.

FIG. 4 is a front view of a foot prosthesis with an adapter of a foot prosthesis pylon and a foot prosthesis pylon shown in FIG.

5 is a top view of the embodiment depicted in FIGS. 3 and 4.

6 is a side view of another foot keel in accordance with the present invention, in particular for sprinting, which can be used in the prosthetic foot in accordance with the present invention.

Fig.7 is a top view of the foot keel shown in Fig.6.

Fig. 8 is a bottom view of the foot keel of the prosthetic foot of Fig. 3, which provides combined high / low dynamic response characteristics, as well as two-plane mobility.

Fig.9 is a side view of an additional stop keel in accordance with the present invention for a prosthetic foot, especially suitable for sprinting, an amputee who underwent an operation to isolate the foot of Saimaa.

Figure 10 is a top view of the foot keel shown in figure 9.

11 is an additional version of the foot keel for a prosthetic foot in accordance with the present invention, intended for an amputee with articulation of the Saimaa foot, where the foot keel provides the prosthetic foot with combined high / low dynamic response characteristics, as well as the possibility of two-plane mobility.

Fig.12 is a top view of the foot keel shown in Fig.11.

13 is a side view of the foot keel in accordance with the present invention, in which the thickness of the keel narrows, for example, gradually narrows from a portion of the midfoot to a portion of the rear portion of the keel.

Fig. 14 is a side view of another in the shape of the foot keel, in which the thickness narrows from the midfoot to both the forefoot and the hindfoot of the keel.

Fig is a side view and a little top and front of a parabolic bent leg of the prosthesis of the foot in accordance with the present invention, the thickness of the leg of the leg tapering to its upper end.

Fig. 16 is a side view resembling that of Fig. 15, but depicting another tibia, tapering from the middle to its upper and lower ends.

Fig. 17 is a side view of the C-shaped leg of the prosthesis of the foot prosthesis, wherein the thickness of the leg of the leg is narrowed from the middle to its upper and lower ends.

Fig. 18 is a side view of another example of a C-shaped leg support of a prosthetic foot, wherein the thickness of the leg support gradually decreases from its middle portion to its upper end.

Fig. 19 is a side view of an S-shaped leg of the lower leg of the prosthetic foot, wherein the thickness of the leg is gradually reduced to both ends from its middle.

FIG. 20 is a further example of an S-shaped tibia, which tapers in thickness only at its upper end.

FIG. 21 is a side view of a J-shaped shank strut narrowed at each end for a prosthetic foot in accordance with the present invention.

Fig. 22 is a view reminiscent of that of Fig. 21, but depicting a J-shaped leg of the lower leg, which gradually narrows in thickness only to its upper end.

FIG. 23 is a side view and a little top view of a metal alloy or plastic joint used in an adjustable fastener in accordance with the present invention for attaching a calf stand to a foot keel, as shown in FIG.

24 is a side view and a little front view of the adapter of the pylon of the foot prosthesis used on the prosthetic foot shown in FIG. 3-5, and also suitable for the foot, shown in FIGS. 28 and 29, for attaching the foot to the pylon of the foot prosthesis, which should be attached to the cult of the amputee.

Fig - side view of another prosthetic foot in accordance with the present invention, similar to the prosthetic foot shown in Fig.3, but depicting the use of a connecting part with two removable fastening means, longitudinally spaced and connecting the part with the leg calf and foot keel, respectively.

Fig. 26 is a side elevational view of the connecting part shown in Fig. 25.

Fig. 27 is a side elevational view of the lower leg of the prosthetic foot of Fig. 25.

FIG. 28 is a side view of a further exemplary embodiment of a foot prosthesis similar to the embodiments of FIGS. 3 and 25, in which a travel restricting shock absorber is attached between respective ends of the calf shank in order to limit the travel distance of the upper end of the calf shank in response to the application and relieving the power load on the tibia using a prosthetic foot.

Fig.29 is a front view of the prosthesis when viewed from the left to the prosthetic foot, shown in Fig.28, which shows a longitudinal slot in the ankle of the foot of the foot.

Fig.30 is a rear view of the prosthetic foot when viewed from the right to the prosthetic foot, shown in Fig.28.

Fig. 31 is a bottom view of the prosthetic foot of Fig. 28.

Fig. 32 is a side view of the calf shank and foot keel of the prosthetic foot shown in Fig. 28, representing an example of displacement of the upper end of the calf shank as a result of applying and removing the load on the calf shank during use of the prosthetic foot.

Fig. 33 is a side view of yet another additional exemplary embodiment of a prosthetic foot similar to the embodiment shown in Figs. 28-32, except that a flexible tape is used to limit only the extent of movement of the extension of the upper end of the calf shank.

Fig. 34 is a side view of another embodiment of a prosthetic foot with a centering connecting device located on an adapter connected to the upper end of the calf, for attaching the foot to the prosthesis receiving sleeve mounted on the leg of the amputee, wherein the centering connecting device allows medial lateral and anterior - Rear sliding adjusting movement of the foot relative to the prosthesis receiving sleeve.

Fig. 35 is a front view of the prosthesis of the foot shown in Fig. 34, when viewed from the left to the foot, shown in Fig. 34.

Fig. 36 is a rear view of the prosthesis of the foot shown in Fig. 34, when viewed from the right to the foot, shown in Fig. 34.

Fig. 37 is a plan view of another stop keel for a prosthetic foot in accordance with the present invention, wherein the rear end of the foot is parallel to the frontal plane, for example perpendicular to the longitudinal axis A-A of the foot, and the longitudinal axis F-F of the proximal concavity of the rear foot is also perpendicular to the longitudinal axis A-A.

Fig. 38 is a side view of the foot keel shown in Fig. 37 when viewed in the direction from the side of the foot keel.

Fig. 39 is a plan view of an additional stop keel in accordance with the invention similar to the foot keel shown in Figs. 37 and 38, but the longitudinal axis F′-F′ of the proximal concavity of the hindfoot forms an obtuse angle Δ ′ with the longitudinal axis AA , which makes the lateral stressed supporting portion of the hindfoot actually longer and more flexible than the medial stressed supporting portion in order to facilitate eversion of the foot when resting on the heel during the step.

FIG. 40 is a side view of the foot keel shown in FIG. 39 when viewed from the side of the foot keel.

Fig. 41 is a side view of an additional embodiment of a prosthetic foot containing an elastic posterior lower leg device attached between the upper portion of the lower leg strut and a connecting part connecting the lower leg to the foot keel, the device accumulating energy in the device springs when applying a power load during the step and returns the accumulated energy when removing the power load in order to increase the dynamic power of the pushing force produced by the prosthetic foot during the step.

Fig. 42 is a rear view of the prosthesis shown in Fig. 41.

Fig. 43 is a side view of an additional embodiment of a foot prosthesis containing a rear shin device to increase the dynamic power of the pushing force produced by the foot prosthesis during a step, wherein an adjustable length tape included in the device is stretched between the upper portion of the shin leg and the front end of the foot keel.

Fig. 44 is a rear view of the prosthesis shown in Fig. 43.

FIG. 45 is a bottom view of the prosthesis shown in FIG. 43 and 44, which depicts a tension cable connected to each side of the foot keel and continuing backward.

FIG. 46A-46D are side views of respective examples of artificial muscle constructions for use with the prosthesis in accordance with the invention to increase the potential energy of the prosthesis and dynamic power in the sagittal plane of the ankle joint during a step.

FIG. 47A-47E are side views of respective different artificial muscle models for use with the prosthesis of the invention.

Figa is a side view of a long artificial muscle, in which the cross-sectional area is 12 linearly located units of area.

Figv is a side view of a shorter, but wider artificial muscle, in which the cross-sectional area contains the same amount, 12 units of area, as in the artificial muscle shown in figa.

Fig. 49 is a vertical projection of a side and bottom view of a sheath or foot covering for a prosthesis according to the invention, the foot sheath comprising an artificial muscle connecting the front and rear plantar sections of the foot sheath.

Fig. 50 is a side view of the foot keel for the prosthesis according to the invention, the foot keel provided with an artificial muscle connecting the front and rear plantar portions of the foot keel.

Fig. 51 is a side view of the leg stand and connecting part of the prosthesis in accordance with the invention, comprising an artificial muscle, a cam for adjusting the artificial muscle tension and a sensor for detecting force in the prosthesis during a step and controlling the cam and artificial muscle tension in response to the detected force and depending on him.

Figa-52D are side views of the respective pads of different sizes, and the pads are suitable for selective installation between the prosthesis and the artificial muscle, so that the user can adjust the preliminary tensile load on the muscle before applying the prosthesis and, thereby, change the potential energy of the prosthesis.

Fig. 53 is a vertical projection of a side and rear view of the prosthesis according to the invention, which shows an artificial muscle mounted on the prosthesis and connecting the receiving sleeve for the user's stump and the lower end of the lower leg of the prosthesis.

Fig. 54 is a vertical projection of a side and rear view of a receiving sleeve for a stump of a human leg, wherein the receiving sleeve is applicable with a prosthesis in accordance with the invention and comprises a chamber containing a pressurized fluid mounted on this sleeve for use between the receiving sleeve and artificial muscle not shown to adjust muscle tension as a preload before use.

Detailed Description of Embodiments

As shown in the drawings, the prosthesis 1 of the foot according to the embodiment of FIGS. 3-5 contains a longitudinally extending foot keel 2 with a portion 3 of the forefoot at one end, a portion 4 of the rear foot at the opposite end and an arcuate upwardly curved portion 5 of the middle foot continuing between sections of the forefoot and hindfoot. In an embodiment, the midfoot portion 5 is curved upwardly convexly along its entire longitudinal extension between the forefoot and hindfoot portions.

The protruding leg 6 of the foot of the foot 1 is attached to the proximal rear surface of the keel of the middle part of the foot 5 with a removable fastening means 8 and a connecting piece 11. In the embodiment, the fastening means 8 is the only bolt with a nut and washers, but may also be a detachable clip or other fastening means for securely positioning and securing the calf shank on the foot keel when securing The tool has been tightened.

A longitudinally elongated hole 9 is made in the proximal posterior surface of the keel portion 5 of the midfoot, as shown in FIG. A longitudinally elongated hole 10 is also made in the curved lower end 7 of the lower leg 6, for example, as shown in FIG. The removable fastening means 8 passes through the openings 9 and 10 so that the tibia and foot keel can be adjusted to the correct position relative to each other in the longitudinal direction AA in FIG. 5, when the fastening means 8 is loosened or disconnected to adjust the functional characteristics of the prosthetic foot to a specific task. Thus, the fastening means 8, the connecting part 11 and the longitudinally elongated holes 9 and 10 constitute an adjustable fastening device for attaching the tibia to the foot keel to form the ankle joint of the prosthetic foot.

The result of adjusting the relative position of the lower leg 6 and foot keel 2 is obvious from a consideration of FIG. 1 and 2, on which, with two radii R ₁ and R ₂ , one directly next to the other, are shown adjacent, facing each other, convex or convex-curved surfaces of the middle portion 5 of the foot keel and leg 6 of the lower leg. From the consideration of two given radii, one directly next to the other, it follows that the movement can be perpendicular to the tangent line A in Fig. 1, A ₁ in Fig. 2, drawn to these two radii. The interconnected arrangement of these two radii determines the direction of the resulting displacements. As a consequence, the application of force determined by the dynamic response of the foot 1 depends on the interconnected arrangement mentioned. The larger the radius of the concave surface, the greater the possibility of a dynamic reaction. However, the steeper the radius, the faster the reaction.

The ability to adjust the relative position of the tibia and foot keel in the prosthetic foot in accordance with the present invention allows radii to be shifted so as to affect the horizontal or vertical linear foot speeds during sports. For example, to increase the maximum possible horizontal linear speed of the prosthesis of 1 foot, you can change the relative location to affect the interconnected location of the radius of the tibia and radius of the foot keel. Namely, in order to increase the horizontal linear velocity characteristic, the lower radius R _{2 of the} foot keel is shifted distal to its initial position, as shown in FIG. 2 in comparison with FIG. 1. This changes the characteristics of the dynamic reaction so that the resulting movements of the foot 1 should be directed closer to the horizontal, and, as a result, a higher horizontal linear speed can be achieved with the same applied forces.

An amputee can, in practice, find a setting for each sport that meets his / her needs when these needs are related to horizontal and vertical linear speeds. A jumper and a basketball player need, for example, a higher vertical climb than a sprint runner. The connecting part 11 is a mobile connecting part made of plastic or a metal alloy (see Figs. 3, 4 and 23), inserted between the fastened foot keel 2 and the leg 6 of the lower leg. A removable fastening means 8 passes through the hole 12 in the connecting part. The connecting part extends along the attached portion of the tibia and the proximal posterior surface of the midfoot portion 5 belonging to the keel.

The curved lower end 7 of the lower leg 6 has a parabolic shape with a minimum radius of curvature of the parabola located at the lower end, and continues up and first forward along the contour of the parabola. The concave surface facing back is formed by the bend of the tibia, as shown in FIG. The parabolic shape is advantageous because, due to the relatively large radii of the proximal extremity of the parabola, this shape has improved dynamic response characteristics in terms of developing a greater horizontal linear velocity and, at the same time, has a lower end with a smaller radius of curvature for the formation of characteristics that determine fast reactions . The large radii of curvature at the upper end of the parabolic shape allow the tangent line A, described with reference to figures 1 and 2, to stay closer to the vertical when the relative position is changed, which ensures the development of a greater horizontal linear velocity.

The tibial stance of a parabolic shape reacts to efforts in the initial phase of contact during a person’s step by compression or twisting. Thus, the radii of the parabolic curve become smaller, and, as a result, the compression resistance is less. On the contrary, when the tibia of the parabolic shape responds by expanding on the support reaction forces (GRF) when the heel is torn off during a person’s step, the radii of the parabolic curve increase, and as a result, the resistance turns out to be much larger than the aforementioned compression resistance. Resistance data is associated with the function of the anterior and posterior muscles of the lower leg of a person during a person's step. From the initial phase of contact to the phase of support on the full foot during a person’s step, the smaller front group of the calf muscles responds to the support reaction forces (GRF) with an eccentric contraction to lower the foot to the support, and a moment of back flexion is created. From the full-foot support phase to the toe-off phase, the large posterior group of the calf muscles responds to the support reaction forces (GRF) with an eccentric contraction, and a more powerful moment of flexion of the sole is created. The magnitude of this moment depends on the difference in size of the anterior and posterior groups of the leg muscles. As a result, the resistance of the prosthetic calf to the moments of the back flexion and flexion of the sole during a person’s step is simulated, and a normal step is provided. The variable resistance of parabolic curves imitates the function of the muscles of the human lower leg during a person’s step and runs and jumps, and, as a result, the effectiveness of the prosthesis is ensured.

A man walks at a speed of about three miles per hour. A mile runner with a result of 4:00 minutes runs at a speed of 12 miles per hour, and a 100-meter sprinter with a result of 10 seconds runs at a speed of 21 miles per hour. The ratio of speeds in this case is 1: 4: 7. In each case, the horizontal component is greater, the greater the speed of the sport. Therefore, it is possible to set the radii of the leg of the prosthesis. A pedestrian needs a parabolic curved shank with shorter radii than a mile runner and sprinter. The sprinter needs a parabolic curved stance of the lower leg, which is seven times larger. This relationship shows how to determine the radii of parabolas for pedestrians, runners and sprinters. The above is of great importance, since sprinters need to be moved on a large scale, and their shank racks must be stronger to withstand the higher loads associated with this type of exercise. A wider or larger parabolic tibia will be bent along a relatively flatter curve, which corresponds to greater structural strength with an increased range of movement.

The adapter 13 of the leg prosthesis pylon is attached to the upper end of the leg stand 6 by the fasteners 14. The adapter 13, in turn, is attached to the lower end of the leg prosthesis 15 by the fasteners 16. The leg prosthesis 15 is attached to the lower limb of the amputee with a support structure (not shown) attached to the stump of the foot.

In an exemplary embodiment, the sections of the forefoot, midfoot and hindfoot of the foot keel 2 are made of one piece of elastic material. For example, you can apply a solid part of a material that is plastic in nature, which has the ability to maintain shape when deflected by the reaction forces of the support. More precisely, the foot keel, as well as the tibia, can be made of a layered composite material containing a reinforcing fiber laminated with a polymer matrix material. In particular, high-strength graphite laminated with thermosetting epoxy resin or extruded plastic known under the Delran trademark or degassed polyurethane copolymers can be used to form the foot keel, as well as the tibia. The functional qualities inherent in these materials make it possible to ensure high strength with low weight and minimal deformation. Layers of thermosetting epoxies are applied under vacuum in accordance with industry standards for prosthetics. Polyurethane copolymers can be poured into negative forms, and extruded plastic can be processed on a machine. Each material used has its own advantages and disadvantages. It has been found that in a preferred embodiment, the multilayer composite material for the foot keel and the tibia can also be a thermoformed multilayer composite material (prepreg) made in accordance with industry standards, with a reinforcing fiber and a thermoplastic polymer binder to achieve the highest mechanical tensile characteristics. A suitable commercially available composite material of the type described is CYLON® from Cytec Fiberite Inc., Havre de Grace, Maryland (m. Maryland). The tibia and foot keel can also be elastic metal elements made, for example, of spring steel, stainless steel, a titanium alloy or an alloy of another metal.

All physical properties of an elastic material related to stiffness, flexibility, and strength are determined by the thickness of the material. A thinner material will bend more easily than a thicker material of the same density. The material used, as well as the physical properties, are associated with the stiffness-flexibility characteristics of the foot keel and the lower leg of the prosthesis. In the exemplary embodiment of FIGS. 3-5, the foot keel and tibia are uniform in thickness or symmetrically distributed, but the thickness may vary along the length of these components, as described below, for example, when performing areas of the hindfoot and forefoot thinner and more sensitive to bending in the midfoot.

To make it easier to give the prosthesis of the 1st foot the ability to combine high / low dynamic reactions, section 5 of the middle part of the foot is made in the form of a longitudinal arch such that the medial side of the longitudinal arch has a higher dynamic response than the lateral side. For this purpose, according to an embodiment, the medial side of the concave surface of the longitudinal arc has a larger radius than the lateral side of this surface.

The relationship between the values of the medial and lateral radii of the concave surface of the longitudinal arch of section 5 of the middle foot is additionally determined, like the front and back loaded areas of the plantar surface of the foot keel 2. The line T ₁ -T ₂ in the front section 5 of Fig. 8 represents the front loaded region of the plantar surface. Line P ₁ -P ₂ represents the back plantar laden surface. Plantar-laden surfaces on the lateral side of the foot would be represented by the distance between T ₁ -P ₁ . Plantar-laden surfaces on the medial side of the foot 2 are represented by the distance between P ₂ -T ₂ . The distances represented by T ₁ -P ₁ and P ₂ -T ₂ determine the radii, and, as a result, the relationship between high / low dynamic reactions is determined, and it may be affected by the convergence or discrepancy between these two T ₁ -T ₂ lines and P ₁ -P ₂ . As a result, the combined high / low dynamic response can be specified when designing a structure. The plantar-laden forefoot surface T ₁ -T ₂ may deviate only 5 ° from the normal to the longitudinal axis AA of the foot keel to create the aforementioned high / low dynamic response, see FIG.

The rear end 17 of the portion 4 of the hindfoot has the shape of an upwardly curved arc that responds to the reaction forces of the support during compression heel shock to absorb shock. The heel, formed by the section 4 of the hindfoot, is made with a posterolateral angle 18, which is shifted back and to the lateral side from the medial angle 19, to facilitate eversion of the hindfoot in the initial phase of contact during the step. The front end 20 of the forefoot portion 3 is shaped like an upwardly curved arch to simulate the back flexion of the fingers of the human foot in the raised heel position and repelled by the toe at the end of the stance of the prosthesis during the step. The forefoot and hindfoot are equipped with rubber or foam pads 63 and 64 from below as shock absorbers.

Improved two-plane mobility of the prosthetic foot is achieved through the medial and lateral holes 21 and 22 of the compensation grooves passing through section 3 of the forefoot from the back to the plantar surface of this section. The compensation grooves 23 and 24 extend forward from the corresponding holes to the front edge of the forefoot portion to form the medial, middle and lateral compensatory tension elements 25-27, which provide improved two-plane mobility of the forefoot portion of the stop keel. The holes 21 and 22 of the compensation grooves are located on the line BB, shown in figure 5, in the transverse plane, which extends at an angle α equal to 35 ° to the longitudinal axis AA of the foot keel, while the medial hole 21 of the compensation groove is shifted forward compared to lateral hole 22 of the compensation groove.

The angle α between the line BB and the longitudinal axis AA in FIG. 5 can be as little as 5 °, and even then a high / low dynamic response will be provided. With a change in this angle α, the angle Z of the line T ₁ -T ₂ in FIG. 8 should also change. In the projection onto the sagittal plane, the openings 21 and 22 of the compensation grooves are inclined at an angle of 45 ° to the transverse plane, with the back side of the holes shifted forward relative to the plantar side. With this arrangement, the distance from the removable fastening means 8 to the lateral opening 22 of the compensation groove is shorter than the distance from the removable fastening means 8 to the medial opening 21 of the compensation groove so that the extension of the toes of the foot in the lateral portion of the prosthesis 1 foot is shorter than in the medial portion, for creating the possibility of high / low dynamic response of the midfoot. In addition, the distance from the removable fastening means 8 to the lateral plantar bearing surface indicated by the line T ₁ is shorter than the distance from the removable fastening means to the medial plantar bearing surface indicated by the line T ₂ , so that the fingers are shorter on the lateral portion of the foot 1 of the prosthesis than on the medial, to create the possibility of high / low dynamic response of the middle part of the foot.

The front part of the portion 4 of the posterior portion of the foot belonging to the foot keel 2 further comprises an opening 28 of the compensation groove passing through the portion 4 of the posterior portion of the foot from the back surface to the plantar surface of this portion. The compensation groove 29 extends in the rear direction from the opening 28 to the trailing edge of the rear foot portion to form compensated stressed elements 30 and 31. These stressed elements provide improved two-plane mobility of the rear foot portion. In an embodiment, each of the compensation groove openings 28, 21, and 22 may be provided with another small opening 28A, 21A, and 22A, see FIG. 8, drilled through an area adjacent to the compensation groove to perform the function of the discharge opening. An additional small hole redirects the current wave spectrum and slows down the process of cracking and / or destruction of the foot keel.

On the back side, the midfoot portion 5 and the forefoot portion 3 belonging to the foot keel 2 form an upwardly concave surface 32, as shown in FIG. 3, so that this surface functionally mimics the fifth degree of human foot mobility. Namely, the concave surface 32 has a longitudinal axis CC, which is oriented at an angle β from 5 ° to 35 ° to the longitudinal axis AA of the foot keel with the medial direction in front compared to the lateral direction, so as to provide a fifth degree of mobility during the step, like a slightly inclined axis of the second - the fifth metatarsal bones of the human foot.

The value of two-plane mobility can be estimated in the case when the amputee is walking on uneven terrain or when the athlete sharply changes direction to the medial or lateral side when resting on the foot. The support reaction force vector changes direction from sagittally oriented to the direction containing the component in the frontal plane. The support will push away in the medial direction, opposite to the lateral push of the foot. As a result of this, the tibia strut tilts in the medial direction, and the weight is applied to the medial construction of the foot keel. In response to the pressure data, the medial compensation tension elements 25 and 31 of the foot keel 2 are bent backward (tilted back) and turned inward, and the lateral compensated tensioned elements 27 and 30 are bent in the plantar direction (tilted down) and turned outward. This movement tends to place the plantar surface flat on a support (plantar stage).

In the prosthetic foot in accordance with the present invention, it is possible to apply mainly another foot keel 33 for sprinting, see FIGS. 6 and 7. The center of gravity of the body in the sprint becomes oriented exclusively in the sagittal plane. The prosthetic foot does not need a low dynamic response. As a result, the longitudinal axis of the concave surface of the forefoot and midfoot does not need to be oriented by turning outward at an angle of 5 ° -35 °, as in the foot keel 2. The longitudinal axis DD of the concave surface would be more correctly oriented parallel to the frontal plane, as shown in FIG. . 6 and 7. This orientation forces the sprint foot to respond only in the sagittal direction. In addition, the openings 34 and 35 of the expansion joints in the sections of the forefoot and hindfoot are oriented along the line E-E parallel to the frontal plane, i.e. the lateral opening 35 is shifted forward to align with the medial opening 34 along a line parallel to the frontal plane. The front end 36 of the foot keel 33 is also made parallel to the frontal plane. The calcaneal region 37 at the posterior end of the foot keel is also parallel to the frontal plane. The described modifications adversely affect the universal capabilities of the prosthetic foot. However, its functional characteristics become aimed at solving a specific problem. Another change in the foot keel 33 for sprinting takes place in the area of the toes in the forefoot, where the angle of 15 ° of the back bend of the foot keel 2 is increased to 25-40 ° of the back bend of the foot keel 33.

Figures 9 and 10 show an additional stop keel 38 in accordance with the present invention for a prosthetic foot, particularly useful for sprinting an amputee who has undergone the isolation of the Saima foot. To this end, the midfoot portion belonging to the foot keel 38 comprises a rear concave surface 39 facing upward, on which the curved lower end of the foot rack is attached to the foot keel with a removable fastening means. This foot keel can be used by all amputees with amputation of the lower extremities. Foot keel 38 contains a longer stump, corresponding to amputations such as the isolation of the foot of Saimaa. Its functional characteristics make noticeably faster dynamic reactions possible. This keel is not specialized for use at a particular amputation level. It can be used for all amputations at the tibia and femur. The foot keel 40 according to the embodiment shown in FIGS. 11 and 12 also has a concave surface 41 for an amputee with articulation of the Syme foot, and this foot keel provides the prosthetic foot with a combined characteristic with high and low dynamic responses, as well as two-plane motion capabilities similar to capabilities in the embodiment depicted in figures 3-5 and 8.

The functional characteristics of several foot keels for a prosthetic foot of 1 foot are due to the peculiarities of shape and design, since these characteristics depend on concave surfaces, convex surfaces, radii, stretching, compression and physical properties of the material, and all the mentioned properties determine the relationship with the reaction to support forces during classes in walking, running and jumping.

The foot keel 42 in FIG. 13 is similar to the foot keel according to the embodiment shown in FIGS. 3-5 and 8, except that the thickness of the foot keel tapers from a portion of the midfoot to the posterior end of the hindfoot. The thickness of the foot keel 43 in FIG. 14 gradually decreases or tapers toward its both front and rear ends. Similar thickness changes are shown at lower leg 44 in FIG. 15 and lower leg 45 in FIG. 16, which can be used in a prosthetic foot 1. Each design of the foot keel and lower leg forms a different resulting functional characteristics, since these resulting functional characteristics are associated with horizontal and vertical linear speeds, which are specifically assigned to increase functional characteristics for different sports tasks. The possibility of several configurations of the calf shafts and adjustments when setting the relative location of the foot keel and calf shafts form such an interdependence between the prosthetic foot and the calf shank that gives the amputee and / or prosthetist the ability to adjust the prosthetic foot to achieve the highest possible functional characteristics when performing the intended exercise from a variety of sports - recreational activities.

Other lower leg racks for the prosthetic foot 1 are shown in FIGS. 17-22 and include C-leg legs 46 and 47, S-leg legs 48 and 49 and J-legs 50 and 51. The upper end of the calf shank may also contain a straight vertical end portion, to the proximal end of which is attached a pyramidal mounting plate. The male pyramid can be attached to said vertical end of the tibia with bolts passing through the vertical end of the tibia, as shown in FIGS. 28-30 and 33-36. Another embodiment of proximal rack attachment may be the rear and / or front sides of the amputee receptacle and / or other components of the prosthesis. In the elongated holes at the proximal and distal ends of the calf shafts, plastic or aluminum inserts can also be provided to fit the proximal male pyramid and distal stop keel. The prosthetic foot in accordance with the present invention is a modular system, structurally made preferably from standard nodes or with standard sizes for universal and diverse use. An exemplary pyramidal fastening bar attached to the proximal extremity of the calf shank is indicated at 88 in the embodiment shown in FIG. 28.

All actions while running on a treadmill are accompanied by counterclockwise movement. Another additional feature of the present invention is to take into account the forces acting on the foot moving forward along such a path curve. Centripetal acceleration acts in the direction of the center of rotation when the object moves along a curved path. Newton’s third law applies to force. Namely, an equal and opposite reaction takes place. Therefore, for each “centripetal” force, there is a “centrifugal” force. The centripetal force acts toward the center of rotation, and the centrifugal force, the opposing force, acts from the center of rotation. If the athlete runs along a closed curve along the treadmill, then the centripetal force pushes the runner to the center of the curve, and the centrifugal force pushes away from the center of the curve. To counter the centrifugal force that tends to tilt the runner out, the runner leans inward. If the runner’s movement on the turn of the treadmill is always counterclockwise, the left side is inside the treadmill. As a result, in accordance with a feature of the present invention, the left side of the leg supports of the right and left foot prostheses can be made thinner than the right side, and functional characteristics can be improved when the amputee runs along a curve.

Each of the foot keels 2, 33, 38, 42 and 43 according to several embodiments has a length of 29 cm with the proportions of the foot 1 shown on a scale in FIGS. 3, 4 and 5 and in several views of different legs of the lower leg and foot keels. However, one skilled in the art can readily understand that the specific dimensions of the prosthetic foot may vary depending on the size, weight and other characteristics of the amputee wearing the foot. The length of the tibial strut and its modulus of elasticity form its potential capabilities and the ability to accumulate elastic strain energy. This stored energy of elastic deformation is transformed by means of a mechanical structure into dynamic power, which provides a directed force with the characteristics of direction and magnitude. Therefore, the longer the strut, the greater the pushing force. The point of proximal fastening of a rack for athletes with the highest functional indicators should be provided as more proximal as the components of the prosthesis allow.

The following describes the functioning of the prosthetic foot 1 in the phase positions of the walking cycles when walking and running. The kinematics of foot movement 2 are based on the three laws of Newtonian mechanics, which are the laws of inertia, acceleration, and action-reaction. From Newton’s third law, the law of action-reaction, it is known that the support pushes the foot with a force equal in magnitude and opposite in direction to the force with which the foot presses on the support. The forces mentioned are known as reaction forces of support. A lot of scientific research has been carried out regarding the actions of a person during a step, while running and jumping. Studies on force plates show how Newton's third law acts during a step. As a result of these studies, the authors determined the direction of pressure of the support on the foot.

The position phase during walking and running can be further divided into phases of braking and acceleration. When the prosthesis of the foot touches the support, the foot presses the support forward, and the support pushes in the opposite direction with equal and opposite forces, i.e. the support pushes the prosthetic foot back. This force makes the prosthetic foot move. An analysis of the phase of the positions when walking and running starts from the point of contact, which is the posterolateral angle 18, see FIG. 3 and 8, which is shifted back and to the lateral side compared with the medial side of the foot. At the beginning of the contact, this displacement forces the foot to turn outward and the tibia to bend to the plantar side. The tibia rack always tends to a position in which the body weight is transmitted through the rack, for example, the tibia rack tends to set its long vertical element in the position to counter the reaction forces of the support. That is why the stance moves backward - it bends in the plantar direction to counteract the reaction force of the support, which pushes the foot backward.

The reaction forces of the support force the legs 44, 45, 46, 47, 50 and 51 to contract, which is accompanied by a backward movement of the proximal end. In the uprights 48, 49, the distal portion, which is 1/2 of the lower leg, will shrink depending on the orientation of the distal concave surfaces. If the distal concavity is compressed in response to the support reaction forces (GRF), then the proximal concavity will expand and the entire tibia will move in the backward direction. The reaction forces of the support force the tibia to contract with the proximal end moving backward. The lower lower radius of the tibia is compressed, simulating the bend of the human ankle joint, and the front of the foot is lowered by pressing on the support. At the same time, the back side of the foot keel, represented by the portion 4 of the hindfoot, indicated at 17, is compressed upward due to compression. Both of these compressive forces have a shock absorbing effect. Shock absorption is further improved by a laterally offset lateral heel 18, which forces the foot to turn outward, which also has a shock absorbing effect after the tibia has stopped bending in the plantar direction and the support pushes the foot in the rear direction.

Then, the compressed elements of the foot keel and tibia begin to unload, namely, they tend to their original shape, and the stored energy is released, which forces the proximal end of the tibia to move forward with acceleration. When the tibia strut reaches its vertical initial position, the reaction forces of the support change from pressing backwards to pressing the foot vertically upwards. Since the foot prosthesis contains the front and back loading areas of the plantar surface, and these areas are connected by an unloaded long arcuate middle section, vertically directed forces from the prosthesis force the long arcuate middle section to be loaded by straightening. The rear and front loading surfaces diverge. These vertically directed forces accumulate in the long arcuate midfoot of the foot as the reaction forces of the support move from the natural vertical direction to the forward direction. The tibia stance is straightened, simulating the back flexion of the ankle joint. This forces the prosthetic foot to turn from the front loaded sole surface. As the weight load is removed, the longitudinal arc of section 5 of the midfoot and the straight leg stand are changed in shape by leaving the straightened state and tend to their original shape, which simulates the trajectories and the resulting movements of the group of flexor muscles of the sole. As a result, the mechanical structures of the prosthesis release the stored energy of elastic deformation in the form of dynamic power.

The long arch of the foot keel and the stance of the lower leg resist the straightening of their respective structures. As a result, the forward movement of the tibia is delayed, and the foot begins to turn from the front loading area of the plantar surface. The straightening of the midfoot portion of the foot keel has the ability to react high and low in the case of the foot keels according to the embodiments shown in FIGS. 3-5 and 8, FIGS. 11 and 12, FIG. 13 and FIG. 14. Since the transitional region between the midfoot and forefoot of these foot keels deviates 15 ° -35 ° to the outside of the long axis of the foot, the medial long arc is longer in length than the lateral long arch. This is important because in the normal foot, the medial side of the foot is used during acceleration or deceleration.

A longer medial arch of the prosthetic foot has a greater dynamic response characteristic than the lateral arch. Shortened lateral extension of the toes of the foot is used when walking or running at low speeds. The body's center of gravity moves in space along a sinusoidal curve. It moves in the medial, lateral, proximal and distal directions. When walking or running at low speeds, the center of gravity of the body makes wider medial and lateral movements than when walking or running at high speeds. In addition, less torque or inertia and less ability to cope with a higher dynamic response. A prosthetic foot in accordance with the present invention is made in accordance with these principles of applied mechanics.

In addition, in the middle position of the human step cycle, the center of gravity of the body is at the extreme lateral point of its movement. When moving from a mid-position to a toe-off point, the body's center of gravity (BCG) moves from a lateral position to a medial position. As a result, the center of gravity of the body crosses the lateral side of the foot keel 2. First (low gear) and as the center of gravity of the body (BCG) moves forward, it moves in the medial direction on the foot keel 2 (high gear). As a result, the stop keel 2 of the prosthesis acts like an automatic transmission. That is, he begins to move in low gear and switches to high gear at every step that the amputee does.

Since the reaction forces of the support push the prosthesis of the foot forward, which presses the support back, when the heel begins to rise, the front section of the long arc of the midfoot section has such a profile that the mentioned backward forces are applied perpendicular to the plantar surface of this section. This method is the most effective and efficient for the application of the mentioned forces. The same can be said about the posterior portion of the hindfoot of the prosthetic foot. It also has such a shape that the backward reaction forces of the support at the beginning of the contact are opposite to the plantar surface of the foot keel, perpendicular to the direction of application of the mentioned forces.

At the last stages of the heel lifting and pushing back with the toe during walking and running, the area of the toes of the foot in the anterior section of the foot is curved to the back at an angle of 15 ° -35 °. This vertically extending arc allows the forward reaction forces of the support to compress this portion of the foot. Resistance to said compression is less than straightening, and a smooth transition to the free swing phase occurs during step and run on the prosthetic foot. In the last stages of the prosthesis stance phase during the step, the rectified leg calf and the straightened long arc of the middle part of the foot release the energy stored in them in addition to the forward and upward push of the driven limb and the center of gravity of the amputee body.

One of the main motive mechanisms during a person’s step is called the active jerk phase. When the heel rises, the weight of the body passes in front of the supporting limb, and the center of gravity falls. When the body weight passes through the forefoot roll shown in Fig. 5 by the C-C line, downward acceleration takes place, which results in the maximum vertical force perceived by the body. The acceleration of the leg forward from the ankle joint, coupled with the rise of the heel, results in a backward shift in relation to the support. When the center of pressure moves forward to the axis of rotation of the heads of the metatarsal bones, the back bending moment is further enhanced. Thereby, a state of full fall forward is created, which produces a significant forward force used when walking. Signs of effective performance of the function of the ankle joint during an active push are the heel lift, minimal joint movement and an almost neutral position of the ankle joint. A stable midfoot is necessary for a normal sequence of movement when lifting the heel.

The rear side of the hindfoot and the forefoot portion of the foot keel comprise openings of the compensation grooves and compensatory stressed elements in several of the previously mentioned embodiments. For a given orientation, the openings of the compensation grooves act as an oblique hinge, so the possibilities of two-plane movement become more perfect and thereby increase the overall contact characteristics of the plantar surface of the foot when walking on uneven terrain.

Shown in Figs. 9-12, the foot keels for amputees with the isolation of the Saima foot noticeably differ in terms of dynamic characteristics, since these possibilities relate to actions when walking, running and jumping. These foot keels differ in four distinct features. These signs include the execution of the proximal posterior portion of the midfoot with a concave surface for more convenient placement of the distal stump when isolating the Saima foot than in the case of a flat surface. In addition, this concave surface reduces the height of the foot keel, which accommodates a longer stump, which is typical for amputants with articulation of the Saimaa foot. The installation concave surface requires that the corresponding front and rear radii of the arched middle portion of the foot keel are more visible and smaller. As a result, all the radii of the long arch of the midfoot and the radii of the hindfoot become steeper and smaller. This greatly affects the characteristics of dynamic reactions. Smaller radii produce less potential energy for a dynamic reaction. However, the prosthetic foot responds faster to all of the aforementioned reaction forces of the support when walking, running and jumping. As a result, a faster foot with less dynamic response is obtained.

Improving the motor performance of an athlete performing a specific task can be achieved by changing the relative position when using the prosthetic foot in accordance with the present invention, since the said changes in the relative position affect the vertical and horizontal components of the force and movement for each task. The human foot is a universal apparatus that can walk, run and jump. On the contrary, the design of the tibia, which imitates the human tibia and fibula, is not a universal apparatus. The mentioned design is a simple lever that exerts its efforts in the implementation of actions related to walking, running and jumping, parallel to its long proximal-distal direction. The stand is an incompressible structure and is not able to accumulate energy. On the other hand, the prosthetic foot in accordance with the present invention has dynamic response capabilities, since these dynamic response capabilities are associated with components of horizontal and vertical linear speeds of the athlete during walking, running and jumping and are superior in functional characteristics to the tibia and fibula of a person. As a result, it is possible to improve the functional characteristics of an amputee involved in sports. To this end, in accordance with the present invention, the fastening means 8 weaken and adjust the relative position of the tibia and foot keel relative to each other in the longitudinal direction of the foot keel. This change is shown in connection with FIGS. 1 and 2. Then, the tibia leg is fixed to the foot keel in the adjusted position by the fastening means 8. During this adjustment, the screw of the fastening means 8 is shifted relative to one or both of the opposite, relatively longer, longitudinally elongated holes 9 and 10, respectively, in the foot keel and tibia.

A change in the relative position, which causes an increase in the functional characteristics of the runner who makes initial contact with the support on the full foot, like a sprinter landing on the middle part of the foot, for example, is that the foot keel is shifted forward relative to the tibia and the foot is bent into the plantar side on the tibia. This new relative arrangement enhances the horizontal component of the run. Namely, when the tibia has a bend in the plantar side to the foot and the foot comes into contact with the support when the foot is flat, as opposed to the initial contact with the heel, the support immediately pushes the foot back, which presses the support in the forward direction. This effect forces the tibia to quickly move forward (due to straightening) and down. The forces of dynamic reactions are formed due to straightening, which resists the direction of the initial movement of the tibia. As a result, the foot rotates relative to the loaded area of the plantar metatarsal surface. This forces the portion of the midfoot belonging to the foot keel to straighten, and the resistance to this straightening is stronger than compression. The total effect of straightening the tibia and straightening the midfoot is that there is resistance to the further forward movement of the tibia, which allows the knee extensors and extensors of the hip joint in the user's body to move the center of gravity of the body in the anterior and proximal directions more efficiently (i.e. with increased horizontal speed). In this case, it is more forward than upward, in contrast to the case of a runner moving from heel to toe, when the forward movement of the tibia forward is less resistant due to the fact that the tibia begins with a larger back (vertical) flexion than case with a runner lowering to the full stop.

To analyze the sprint foot during operation, a change was made in the relative location of the tibia and foot keel. Advantages are provided by the foot keel, in which the longitudinal axes of all its concave surfaces are oriented parallel to the frontal plane. The tibia is bent in the plantar direction and pushed back on the foot keel. This reduces the distal circles even more than the runner with lowering to a full foot with a universal stop keel similar to the keel shown, for example, in Figures 3-5 and 8. As a result, the horizontal movement potential is even higher, and the dynamic reaction is directed to improve the capabilities of this movement horizontally.

Sprinters are characterized by increased range of motion, forces and moments (inertia), and the moment is the main mover. Since their braking phase in the stance phase of the prosthesis is shorter than their acceleration phase, increased horizontal linear velocities are achieved. This means that in the initial contact phase, when the toe touches the support, the support pushes the foot backward and the foot pushes the foot forward. The tibia, which is characterized by increased forces and moment, is forced to even more bend and downward movement than a runner with initial contact on a full foot. As a result of the action of these forces, the long arched concave surface of the foot is stretched, and the tibia is stretched. The resistance exerted by the given tensile forces exceeds the resistance to all the other previously mentioned forces associated with running. As a result, the dynamic response of the foot is proportional to the applied force. The reaction of the tibial strut, which imitates the human tibia and fibula, is only related to the energy potential of the forces, since this strut has a straight structure and cannot accumulate energy. The mentioned tensile forces in the prosthetic foot according to the present invention during sprinting are larger in size than all the other previously mentioned forces associated with walking and running. As a result, the possibility of dynamic response of the foot is proportional to the applied forces, and therefore, the motor performance of an athlete with an amputated leg can be increased in comparison with the functional indicators of the human body.

The prosthetic foot 53 shown in FIG. 25 resembles that of FIG. 3 with the exception of an adjustable fastener between the tibia and foot keel and the design of the upper end of the tibia for connecting to the lower end of the leg prosthesis. In this exemplary embodiment, the foot keel 54 is adjustably connected to the lower leg 55 by a plastic or metal alloy connecting part 56. The connecting part is attached to the foot keel and lower leg with corresponding removable fastening means 57 and 58, which are spaced from each other in the connecting part in the direction coinciding with the longitudinal direction of the foot keel. The fastening means 58 connecting the connecting part to the tibia is offset backward with respect to the fastening means 57 connecting the foot keel to the connecting part. Due to the increase in the effective length of the tibial strut, the possibilities for the dynamic response of the tibial strut itself are increased. Changes in the relative location are performed in conjunction with the longitudinally elongated holes in the calf and foot keel, as in other embodiments.

At the upper end of the leg stand 55, an elongated hole 59 is made to accommodate the pylon 15 of the leg prosthesis. The foot prosthesis pylon inserted into the hole can be firmly clamped in the lower leg by wrapping the bolts 60 and 61 to tighten the free sides and lower legs of the leg leg 62 and 63 along the length of the hole. The described connection of the foot prosthesis pylon can be easily adjusted, for which the bolts are loosened, the foot prosthesis pylon is moved to a predetermined position relative to the lower leg, and the foot prosthesis pylon is clamped again in the adjusted position by wrapping the bolts.

The foot prosthesis 70 shown in FIGS. 28-32 is similar to the foot prostheses shown in FIGS. 3-5, 8, 23 and 24 and in FIGS. 25-27, but further comprises a leg swing range limiter and a shock absorber 71 on foot to limit the extent of movement of the upper end of the tibia by applying and removing the load on the tibia while using the amputee. The mentioned distinguishing feature is especially useful in the prosthetic foot, containing a relatively long stance of the lower leg, when the user plans to engage in activities such as running and jumping, during which the leg develops efforts that are many times greater than the user's body weight, for example, when running - 5-7 times more body weight, and when jumping - 11-13 times more body weight. On the contrary, forces developed during walking, are only 1-1 _^1/2 body weight.

The device 71 in an exemplary embodiment is a double-acting piston assembly into which fluids, gas, such as air, or hydraulic fluid are supplied under pressure through fittings 73 and 74. The device contains two smooth regulators, one for compression and one for extension, which allow the adjustment of the permissible extent of movement of the upper end of the tibia rack 72 both during compression and during straightening of the tibia rack during application and unloading. The device 71 also absorbs energy accumulated or released during compression and straightening of the tibia. The opposite ends of the piston device 71 are attached to the upper end of the calf shafts, the lower portion of the foot, and, preferably in an exemplary embodiment, to the corresponding ends of the calf shafts in articulated joints 75 and 76, which are preferably spherical joints.

The movement of the upper end of the tibia of the tibia of the foot 70 during compression and straightening of the tibia is shown in Fig. 32. The tibia is generally parabolic in shape so that the upper end of the tibia can move longitudinally with respect to the foot keel 77, with the lower end of the tibia attached to the foot keel, for example, in the direction AA shown in FIGS. 5 and 32, squeezing and straightening the tibia during application and removal of the load on the tibia. Therefore, in the exemplary embodiment shown in FIGS. 28-32, enhanced characteristics of the prosthetic foot dynamic responses are retained.

The device 71 is not limited to the above-described piston assembly, but may be another speed control device and / or movement restriction. For example, it can be imagined that the shock-absorbing device 71 for limiting the backward movement limit used on the shank of the prosthetic foot of the foot can be a microprocessor-based hydraulic unit for controlling the phases of compression and extension, for example, the type that is currently used to control movement in artificial knee joints. In this case, built-in sensors are provided that read the movements of a person and adjust to them. When using special software and a personal computer (PC), you can make precise adjustments to the microprocessor-controlled hydraulic unit to adjust it to the amputee. Moments can be measured at a pace up to 50 times per second, to ensure that the step dynamics are as identical as possible with natural walking. Thanks to the responses of the hydraulic unit, it is suitable for a wide variety of amputees without lower limbs. A lithium-ion battery charged into the device provides enough energy to operate the hydraulic unit for a full day. Compression resistance is adjustable independently of straightening adjustments. Several built-in sensors provide a data stream for step analysis to the built-in microprocessor, which automatically adjusts the rack and changes the characteristics of the device in step phases at a rate of 50 times per second.

This microprocessor-controlled hydraulic unit 71 is more sensitive than a mechanical hydraulic unit. An electrically operated compression stroke valve (flexion to the side of the sole) adjusts 50 times per minute. The compression stroke valve in the block automatically fully opens in the phase before the step. As a result, the block compresses very easily when bent at low speeds, in tight spaces, and under similar conditions. The speed of the servomotor of the unit allows it to very quickly close the valves of the compression stroke (bending towards the sole) and the straightening stroke with rear bending according to the microprocessor commands sent 50 times per second. When the valves are almost closed, the damping force of the unit becomes very large, which allows you to walk quickly and even run. The uniquely determined dynamic coefficient, regulated by the prosthetist, allows you to optimize the hydraulic unit in accordance with the characteristics of all walks, from slow to energetic, with high speeds and fast movements when walking. This possibility of "tuning" the hydraulic unit with microprocessor control in accordance with an individual uniquely defined gait characteristic allows you to get a wide range of changes in the pace of the step in the prosthetic foot with highly efficient and comfortable walking. That is, the use of a microprocessor hydraulic unit as a device 71 improves the adjustable pace required when the prosthetic foot is used by active amputants.

A longitudinally extending foot keel 77 of the foot prosthesis 70 shown in FIGS. 28-32 contains portions of the forefoot, midfoot and hindfoot similarly to the foot keels shown in FIGS. 3 and 25. The tibia foot 72 is attached to the foot keel a connecting part 78 with two longitudinally spaced removable fastening means 79 and 80 securing the connecting part to the tibia and foot keel, respectively, as in the exemplary embodiment shown in FIGS. 25-27. The lower leg 72 comprises a longitudinally extending compensation groove 81 between the ends of the lower leg. The holes 82 and 83 of the compensation groove are located at the ends of the compensation groove. The parts of the forefoot and hindfoot belonging to the stop keel are also made with corresponding compensation grooves, as can be seen from Figs. 29, 30 and 31.

A prosthetic sleeve connected to the bottom of the stump of the amputee is attached to the upper end of the lower leg 72 through an adapter 85 fixed to the upper end of the lower leg by mounting means 86 and 87, as shown in the drawings. The adapter includes a mounting unit 88 in the form of an inverted pyramid connected to a mounting plate fixed to the upper surface of the adapter. The pyramidal unit is housed in a sleeve-shaped unit having a reciprocal form on the dependent prosthesis sleeve, for connecting the prosthetic foot to the prosthetic sleeve. A compound of this type is shown in the embodiment of FIGS. 34-36.

The travel limiting and damping device 71 in the exemplary embodiment shown in FIGS. 28-32 limits the travel distance of the upper end of the calf shank both during compression and when the calf shank is straightened, however, a similar device that limits the travel distance of the upper end of the strut can be used shins only with any one stroke from the compression and straightening moves. A travel limiting and damping device 84, restricting only the straightening of the upper end of the calf shank when applying and removing a power load, is shown in the exemplary embodiment shown in FIG. The device 84 in this figure is a flexible tape that allows limited elastic stretching and, thus, the straightening of the tibia, but does not limit the movement of the upper end of the tibia while applying a compressive load on the tibia. This resilient device 84 can be tensioned when applied, whereby the resilient device predisposes the proximal end of the strut to rearward movement. The device 84 does not need to be in the form of a tape, but may have a different shape, for example, a coil spring or a plurality of coil springs can be used as a travel limiting and damping device that allows limited elastic tension and which accumulates energy during application of a power load and releases the accumulated energy when removing power load during a step.

On Fig-36 presents the leg 90, made with the ankle joint as a monolithic elastic element, this leg can be used with the foot keel 77 of the prosthetic foot, shown in Fig-32, or with one of the other foot keels. The tibia rack 90 has a generally parabolic shape with the smallest radius of curvature of the rack at its lower end and continues upward and initially forward with a change to relatively large radii at the proximal end of the rack. The concave surface facing back is formed by the bend line of the tibial strut shown in FIG. 34. The distal end of the calf shank contains a longitudinally extending hole 91, which, together with the connecting part 78, removable fastening means 79 and 80, and a longitudinally extending hole in the foot keel, allows adjustment of the shin and foot keel position relative to each other in the longitudinal direction when the fastening means 79 or 80 weaken or disconnect to adjust the functional characteristics of the prosthetic foot in accordance with a specific task.

The distal end of the leg stand 90 is bent more sharply, for example, has smaller radii of curvature than that of the leg stand 72, shown in FIGS. 28-32, and extends up and down for a smaller longitudinal distance. The shape of this lower leg makes a more favorable external impression. That is, the distal end of this rack is more precisely located in the area of the ankle joint, where the medial and lateral ankles of the outer sheath of the prosthetic foot, made in the shape of a human foot, should usually be located. The foot stand is better embedded in the outer shell of the prosthetic foot. The functional characteristics of the rack allow this rack to respond more quickly to the reaction forces of the support in the initial contact phase, but with a more stringent dynamic response characteristic than in the case of the lower leg with a wider parabola, for example, the longer curvature radii described above. Thus, active individuals who run and jump on the prosthetic foot will benefit from the use of a wider parabola or radius of curvature, which makes it possible to achieve greater horizontal speed.

The shin leg 90 shown in FIGS. 34-36 further comprises a centering connector 92 located between a plastic or metal adapter 93 connected to the upper end of the leg by fastening means 94 and 95, and the lower end of the prosthesis sleeve 96 is attached to the user's stump. The user may be, for example, an amputee with amputation of the legs above the knee or below the knee. The alignment connecting device comprises a pair of slide guides 97 and 98 located at right angles to each other, in planes parallel to the support. The relative arrangement of the components of each sliding guide can be adjusted by loosening the threaded fasteners 99 to regulate the respective sliding guides 97 and 98 with a change in the relative orientation of the prosthesis sleeve relative to the lower leg and foot keel of the foot prosthesis. The top of the adapter 93 supporting the device 92 is preferably parallel to the leg in the step phase with the foot prosthesis.

The top of the upper sliding guide 98 of the device 92 comprises an inverted pyramid assembly 101 mounted on this sliding guide, which is clamped with adjustment in the corresponding fastening assembly 102 on the prosthesis sleeve 96 by means of threaded fasteners 103. This connection of the nodes 101 and 102 to one another allows the change of angle - bending / straightening and abduction / adduction - between the prosthesis sleeve and the foot. The slideways of the device 92 allow for medial lateral and anteroposterior linear adjustments. Thus, the device 92 is a centering device that allows the prosthesis sleeve to be moved in all directions, which affects how the reaction forces of the support respond to the effects of the mechanical structures of the tibia and foot keel.

Foot keel 110, shown in Figs. 37 and 38, and foot keel 120, shown in Figs. 39 and 40, are further exemplary embodiments of the foot keels that can be applied to the prosthetic foot in accordance with the present invention. Foot keels are designed for the right foot and have identical designs except for the portion of the hindfoot. The medial and lateral sides of the two foot keels have the same shape. In the foot keel 110, a sagittal cut is made in the region of the hindfoot with the formation of identical lateral and medial compensation tension elements 111 and 112, separated by a longitudinally extending compensation groove or slot 113. The heel region 114 at the posterior end of the foot keel 110 is parallel to the frontal plane, for example, perpendicular to the longitudinal axis AA foot keel. Similarly, the back concave surface 115 of the hindfoot of the foot keel is characterized by its own longitudinal axis F-F directed parallel to the frontal plane, for example, at right angles to the longitudinal axis A-A, i.e. the angle Δ is 90 °.

The foot keel 120, in contrast to the foot keel 110, does not contain a sagittal cut in the hindfoot, but contains its own back concave surface 121 of the hindfoot with such a cut that the longitudinal axis F′-F ′ of the concave surface is laterally inclined to the frontal plane, for example, makes an obtuse angle Δ ′ with the longitudinal axis AA, preferably 110-125 °, with the lateral side protruding further forward compared to the medial side. Due to this orientation of the back concave surface, the lateral compensation tensile element 122 has a smaller thickness along a greater length than the medial compensation tensile element 123, and as a result, in fact, has a greater length and greater flexibility than the stressed element 123. The aforementioned greater flexibility makes the back of the foot predisposed to responding to the reaction forces of the support in the initial contact phase by eversion, which serves as a shock absorption mechanism. This contributes to the effective transmission of the center of gravity of the body through the back of the foot of the foot keel during gait to provide a more normal step characteristic.

The prosthetic foot 124 shown in FIGS. 41 and 42 comprises a foot keel 165, a tibia leg 126 and a rear tibia device 125 for storing additional energy while moving the upper end of the tibia leg forward during a step. That is, in the active jerky phase of the step, the force load on the elastic prosthesis straightens the concavity of the leg 126 located in the sagittal plane, formed by the convex-curved section of the leg leg facing forward, which leads to the forward movement of the upper end of the leg leg. The flexible tape 128 of the device 125 is attached to the upper portion of the tibial leg and the lower portion of the prosthetic foot, namely the connecting part 129 that connects the tibial leg and foot keel in the manner described above. The length of the flexible tape, which may be elastic and / or non-elastic, is stretched during the step and can be adjusted using a slide adjuster 130 located between the overlapping sections of the tape length.

The two springs 131 and 132 are adjustablely secured by their bases on the upper end of the calf post between the calf post and an adapter 133 attached to the calf post by fasteners 134. The lower free ends of the springs are set to interact with the flexible tape. When the tape is pulled, the springs change the direction of the longitudinal extent of the tape. Moving the upper end of the tibial strut during a step pulls / additionally pulls (if a tension load has been previously applied to the tape) the tape and loads / additionally loads the springs to accumulate energy when a force load is applied to the prosthetic foot during the step. This accumulated energy is returned by the springs when removing the load on the prosthetic foot to increase the dynamic power of the pushing force produced by the prosthetic foot during the step.

When the tape 128 is shortened to the initial prestress of the tape before applying the foot prosthesis, the tape tension serves to support the rearward movement of the upper end of the elastic element, as well as to regulate the forward movement of the lower leg when using the prosthesis. Support for backward movement can be useful to ensure a quick reaction of the flat foot of the prosthetic foot at the moment of heel strike in the initial phase during the step, similar to what happens in the foot and ankle of a person during the step of the heel strike, when plantar flexion of the foot occurs.

Both the enhanced rearward movement of the upper end of the elastic lower leg strut and the adjustable forward movement of the upper end of the elastic lower leg strut when using the prosthesis with the rear lower leg device 125 can effectively change the transformation coefficient of the ankle torque of the prosthetic foot during the step by influencing the change in the characteristic of the bend in the sagittal plane , which determines the longitudinal movement of the upper end of the tibia, in response to the application and the removal of power load when applied by the user m prosthesis of the foot. According to previously published data, the natural physiological coefficient of transformation of ankle torques in a person’s foot during a step is determined as a result of dividing the maximum ankle moment of the back bend, which occurs at the end of the stand on the prosthesis during the step, by the ankle moment of the plantar bend, created at the time of the initial reaction to the load of a flat foot after hitting the heel during the step is 11.33 to 1. The purpose of changing the characteristics of the bend in the sagittal plane, about redelyayuschey longitudinal movement of the upper end of the tibia with a rear rack device tibia 125, it is to increase the transformation coefficient ankle prosthesis twisting moments for simulating factor which occurs in the human foot during a step. The mentioned goal is important for ensuring the correct gait on the prosthesis and, for persons with one natural foot and one foot of the prosthesis, to ensure a symmetrical gait. In a preferred embodiment, by adjusting the forward movement and possibly supporting the backward movement with the back of the lower leg device 125, the transformation coefficient of the ankle torque of the prosthesis is increased so that said maximum ankle moment of the back bend that occurs in the prosthesis is an order of magnitude greater than the mentioned the ankle moment of the plantar bend in the same prosthesis. In a more preferred embodiment, the ankle torque transformation ratio is increased to about 11: 1, which is comparable to the well-known natural ankle torque transformation ratio of 11.33: 1.

An additional goal of the posterior calf device is to increase the effectiveness of the foot prosthesis during the step by accumulating additional elastic deformation energy in the springs 131 and 132 of the device during application of the force load to the prosthesis and returning the accumulated elastic deformation energy during removal of the force load to increase the dynamic power of the pushing force, produced by the prosthetic foot during the step. We can assume that the device 125 performs the same function in the prosthetic foot that the muscles of the lower leg in the human foot, ankle joint and lower leg perform during the step, namely, it effectively generates a pushing force acting on the human body during the step, using potential energy produced in the body during the application of power load on the foot, and the conversion of this potential energy into kinetic energy to create a pushing force during the removal of power load from the foot. Achieving or even exceeding the performance indicators of the human foot in the prosthetic foot in accordance with the present invention with the rear shin device is important, for example, to restore the “normal functioning” of the amputee.

Controlling forward movement of the upper end of the calf shank 126 by the rear shin device 125 helps to limit the forward movement of the upper end of the calf shank, as in the previous embodiments shown in FIGS. 28-33. The stop keel in the prosthesis 124 feet, when straightening its elastic longitudinal arc, also contributes to the accumulation of energy during the application of power load during the step. This potential energy is returned in the form of kinetic energy to create a pushing force during the unloading during the step. In this embodiment, the combined high / low dynamic response is ensured as a result of the fact that the portion of the midfoot belonging to the stop keel is made in the form of a longitudinal arc, which differs on the medial side by a large radius and a relatively higher dynamic response characteristic than on the lateral side of the arc, as described above in connection with the embodiments shown in FIGS. 3-5 and 8. However, the foot keel for the sprint, such as shown in FIGS. 6 and 7, or the foot keel for For Saima-dysfunctional amputants described with reference to Figures 9 and 10, 124 prostheses can be used in the prosthesis.

The prosthetic foot 135 shown in FIGS. 43-45 comprises a rear shin device 136 similar to the device in the embodiment shown in FIGS. 41 and 42, except that a flexible tape of adjustable length 137 is secured between the upper end of the shank and the front the end of the foot keel using a connecting cable 138. The ends of the cable 138 are attached to the respective ends of the medial and lateral front tension elements 139 and 140, separated by a compensation groove 141. The cable continues back and then up through the pulleys 142 and 143, installing lenny on the connecting part 144, to the semicircular revolving element 145, attached to the distal end of the tape 137. The elastic arch of the foot keel, as well as the spring 146, mounted on the upper end of the leg calf and connected to the tape, as in the embodiment shown in Fig.41 and 42, are used to accumulate and return energy to increase the pushing force developed by the prosthetic foot during the step in the manner described above.

The adapter 133 in FIG. 43 is a male pyramidal adapter, while the adapter in FIGS. 41 and 42 is a female adapter of the present invention comprising a square socket with rounded corners at its proximal end to fit with a clearance of a square protrusion of the reciprocal shape on the receiving sleeve on the bottom limbs or other component on the cult of the amputee. See the dashed lines in FIG. Four screws, not indicated by positions, one in the middle of each side wall of the square socket, can be screwed in and out of engagement with the protrusion to attach the prosthesis to the supporting structure on the stump of the amputee. The gap between the protrusion and the socket and the ability to adjust the location of the four screws of the female adapter allow front-back and medial-lateral adjustment, as well as adjusting the angle or inclination of the prosthesis relative to the supporting structure. In accordance with an additional feature of the female adapter, the threaded fastener connects with a connector the upper element containing the socket, the adapter element with the bearing surface of the adapter. The top of the threaded fastener, which is open for access in the lower surface of the adapter socket, contains a hexagonal recess for receiving the Allen key so that the element containing the adapter socket on the bearing surface can be released so that the element can be rotated relative to the bearing surface and prosthesis. Thus, the adapter provides the ability to rotate in the transverse plane; those. a distinctive feature that makes it easier to bring in and out the foot within the necessary limits, for example, within 1/8 of an inch.

A prosthesis that has higher functional characteristics and develops a higher dynamic power in the sagittal plane of the ankle joint during a step, reaching the power of a normal human ankle joint, is obtained by using an additional distinguishing feature of the present invention, according to which an artificial muscle is proposed for at least least one of the foot, ankle joint, and prosthetic leg. An artificial muscle stores energy during the application of a force load to the prosthesis in the active jerk phase of the user's step, and during the last stages of the position phase during the step releases stored energy to enhance the push of the guided limb and the user's body. The artificial muscle is preferably made of a viscoelastic material, for example rubber or polymer, or a composite containing a viscoelastic material. You can apply a hybrid design that combines a viscoelastic material and a mechanical element or elements to form an artificial muscle. A useful highly elastic biological material that can be used is resilin. The aim of the invention is that the artificial muscle or muscles provided on the prosthesis in accordance with the invention increase the dynamic power in the sagittal plane of the ankle joint during the step on the prosthesis to at least normal dynamic power in the sagittal plane of the human ankle joint during step.

The artificial muscle may be a continuous elastic tape, for example, such as that indicated at 170 in FIG. 50, where the muscle is placed on the foot keel 171 of the prosthesis in accordance with the present invention. Alternatively, several layers of tapes, similar to those indicated at 172, 172 in FIG. 49, can be used to form artificial muscle 173. Muscle 173 is placed on the sheath 174 of the prosthetic foot in accordance with the present invention. The materials and / or elastic properties of the respective tapes may be different. Artificial muscles in exemplary embodiments are passive materials, but it is contemplated that active artificial muscles, such as artificial muscles based on electroactive polymers, can be used. When used to copy the function of the posterior tibia muscle group during a step, the artificial muscle in exemplary embodiments is positioned continuously with a proximal-distal orientation. The artificial muscle on the foot has anteroposterior orientation. However, similarly to the human foot, artificial muscles can also have a medial-lateral orientation and other orientations, or combine different orientations.

The shape characteristics and the mass of the viscoelastic artificial muscle dictate specific motion characteristics. In the case of passive artificial muscles, in exemplary embodiments, the muscle mass reflected by the product of the muscle cross-sectional area by its length determines its possible power. An artificial muscle of a larger mass will have a higher possible power than a muscle of a smaller mass. Two artificial muscles 180 and 181 of equal mass are shown schematically in FIGS. 48A and 48B, where each muscle is depicted with a cross-sectional area of 12 square units, indicated by numbers 1-12 in the drawings. The artificial muscle 180 in FIG. 48A is characterized by a long and narrow cross-section of the tape with a uniform distribution of the cross-sectional area along its length. The muscle 181 in FIG. 48B is shorter and has a wider section in the middle between the ends of the muscle. For a given viscoelastic material, a long artificial muscle 180 with a narrow cross-section will have a greater range of movement compared to a shorter artificial muscle 181 with a larger cross-sectional area. Artificial muscle 181 with a larger cross-sectional area and shorter length creates higher tension values for a given longitudinal extension in a shorter time interval than a long narrow artificial muscle 180 of the same mass. The described characteristics of muscle 181 are preferable for the artificial muscle placed on the prosthesis in accordance with the present invention, for the most accurate copy of the functional characteristics of the human foot, ankle joint and lower leg when walking.

Artificial muscles of different shapes can be placed in layers one on top of the other to obtain, as a result, different movements. Each resulting displacement has the characteristics of tension during displacement, unloading and time interval. By varying the density and hardness of the viscoelastic materials of the artificial muscle, these characteristics can be further adjusted. An example of a multilayer artificial muscle 190 with layers 190A and 190B is shown in FIG. 51, where the proximal and distal ends of the artificial muscle are attached to the leg stand 191 in accordance with the present invention at the upper and lower ends of the leg leg. The fastener at the lower end of the calf shank is located on or near the prosthesis connecting part 192, which connects, as described above, the shank leg with a stop keel not shown. The ends of the artificial muscle can be attached to the prosthesis by clamping them between the abutting elements of an adjustable length fixture not shown in the drawing, which are bolted to the connector. Between the ends of the artificial muscle and the prosthesis, a fastening device with a dovetail joint or another type can also be used.

The layers 190A and 190B of artificial muscle 190 in FIG. 51 together form a layered structure that has a cross-sectional area wider in the middle section along the length of the muscle, as shown in cross-section along the longitudinal axis in FIG. 51, to increase the tension and reduction of reaction time, as described above. Examples of artificial muscles of other forms with the aforementioned similar favorable distinguishing features are indicated at 220-224 in FIGS. 47A-47E. The artificial muscles 222 and 224 are provided with double-ended cutouts 222A and 224A. Mentioned cutouts allow you to create a greater relative elongation of the muscle before stretching the artificial muscle under power load. Four additional configurations of artificial muscles for use in the prosthesis in accordance with the present invention are indicated at 150, 151, 152 and 153 in FIGS. 46A-46D. The artificial muscle 150 in Fig. 46A is made with a central viscoelastic ball 154, which is surrounded in a circle by a plurality of cables 155, elastic or inelastic, which extend along a 180 ° arc around the ball between the upper and lower muscle strands 156 and 157. The tension of the artificial muscle in the direction of its longitudinal axis causes elastic compression of the ball for energy storage.

The artificial muscle 151 in Fig. 46B is made with many cylindrical segments 158 of viscoelastic material, the ends of which are attached to the support plates 159 and 160. The segments are intertwined, as shown in Fig. 46B. The configuration 152 in Fig. 46C is similar to the configuration in Fig. 46B except that the cylindrical segments 158 are not interlaced, but parallel and surrounded by a ring of viscoelastic material 161, which elastically stretches when the muscle extends in the direction of its longitudinal axis under tension. The cylindrical segments 162 in FIG. 46D are parallel and extend between the common support plates 163 and 164 and are attached thereto.

The foot sheath 174 in FIG. 49 is made of an elastic material, such as plastic or polyurethane. The shell of the foot is made with the possibility of tight coverage of at least the lower extremity of the prosthesis in accordance with the present invention to create the appearance of the human foot, for lining and protecting the prosthesis from weathering. The artificial muscle 173 on the prosthesis sheath is fixed at its ends to the prosthesis sheath to connect the posterior and anterior portions of the sole of the prosthesis sheath. The flexion of the body of the foot / shell of the foot during the step stretches the artificial muscle with the accumulation of energy, which is released later during the step to support the effect of the absence of air flow and pushing the foot and limb forward. A similar result is obtained by means of artificial muscle 170, which is attached to the front and rear ends of the foot keel 171 of the prosthesis in accordance with the present invention shown in Fig. 50. The artificial muscle 190 on the tibia strut 191 in FIG. 51 is stretched by moving forward the upper end of the tibia strut during the step to conserve and release energy, as described in connection with the rear tibia devices shown in FIGS. 33 and 41-45.

The potential energy of the prosthesis with at least one artificial muscle in accordance with the present invention can be further increased to increase the kinetic energy and dynamic power in the sagittal plane of the ankle joint during the prosthesis step by preloading the artificial muscle in tension. This can be done by the user of the prosthesis in accordance with the present invention using a gasket, cam or chamber containing compressed air or other fluid, or by combining the above devices. On figa-52D shows four gaskets 230-233 of different sizes. The selected pad can be placed between the prosthesis structure and its artificial muscle to tension the artificial muscle by a value proportional to the size of the pad. For example, when the artificial muscle is located on the back side of the leg of the prosthesis, as shown in FIG. 51, but without the cam shown in the configuration of this figure, a pad could be placed between the leg of the leg and the muscle.

The cam 192 shown in FIG. 51 is pivotally mounted on a cylindrical axis 193 mounted on a cam support structure 194, dependent on its fastening on the upper end of the tibia, in front of the artificial muscle 190. System 195 with a pneumatic cylinder, hydraulic cylinder and / or by an electric drive (electromagnet), the cam rotates adjustable around the axis 193 to tension the muscles. The said system on cylinders (electromagnet) is switched on by movement performed in the tibia during physical activity of the user in the middle position of the position phase during the step. When the user applies a force load to the prosthesis, the distal end of the calf shank comes into contact with the pusher 196 of the lower cylinder (electromagnet), which forces the pusher of the upper cylinder (electromagnet) to come into contact with the cam. During the execution of the action, the forward-directed force increases, the pressure on the lower cylinder increases proportionally, which proportionally moves the upper cylinder into contact, which forces the cam 192 to come into contact with the muscle with the development of a proportional tensile load of the artificial muscle. An increase and / or decrease in power load similarly affects the tension of the artificial muscle. Thus, the forward forward variable load on the foot keel causes the alternating tension of the artificial muscle, and the resulting movements of the prosthesis vary depending on the type of user activity. In a modification of this embodiment, instead of an electromagnet for cam drive 192, a device with a worm gear or a simple screw for adjusting the cam and muscle tension depending on the measured force is applicable.

The chamber 210 of FIG. 54, containing compressed fluid or air, is mounted on the stump receptacle 211. The camera is located between the artificial muscle 213 shown in Fig. 53, which extends between the receiving sleeve and the prosthesis in accordance with the invention and is attached to them, as shown in Fig. 53. The user can change the pressure of the fluid in the chamber with a valve 214 for an adjustable pre-load of the muscle tensile before playing sports to optimize the functional characteristics of the prosthesis. In a variation of this embodiment, the proximal end of the artificial muscle may be attached to the proximal end of the popliteal sleeve or to the body of the artificial knee prosthesis. In a further modification of the embodiment, the artificial muscle is attached from the proximal side to the human knee or knee prosthesis. In this last modification, the artificial muscle will act on two joints, the knee and the ankle.

Here, a description of exemplary embodiments is completed. Although the present invention is described with reference to a number of illustrative embodiments, it should be understood that those skilled in the art will be able to create many other modifications and embodiments that will not go beyond the essence and scope of the principles of the present invention. For example, the lower end of the ankle of the foot prosthesis according to the present invention is not limited to a parabolic shape or a substantially parabolic shape, but can be made in a different curved shape convex down to obtain a given resulting foot movement when attached to the foot keel to create an ankle area joint of the foot. The distinguishing features of the various embodiments may also be used in combination with each other. In particular, justified changes and modifications of components and / or layouts in the system with a combination of features within the scope of the above description and drawings and the attached claims are possible without departing from the scope of the invention. In addition to changes and modifications to the components and / or layouts, alternative applications will also be apparent to those skilled in the art.

Claims (29)

1. A method of producing kinetic energy for pushing force in an elastic prosthesis of the lower limb containing an elastic foot, an elastic ankle joint and an elongated protruding, upward elastic leg of the leg above the ankle joint, while the elastic leg of the leg is a curved bulge forward from the lower end of the leg to the upper the end of the tibia, so that the concavity, which is backward located in the sagittal plane, is formed due to the bent tibia of the tibia, and the method comprises the following steps:
at least two concavities located in the sagittal plane of concavity are straightened, including the concavity of the lower leg, back of the sagittal plane, of the elastic prosthesis during the application of a force load on the prosthesis in the active push phase of a person’s step to accumulate energy in the prosthesis,
wherein the prosthesis is capable of releasing said energy stored in said straightened concavity located in the sagittal plane at later stages of the position phase during the step to amplify the push of the driven limb and the human body,
moreover, during the application of the said force load on the prosthesis in the active jerk phase of the step, additional energy is accumulated in the artificial muscle provided on at least one of the foot, ankle joint and leg of the prosthesis, and, at the aforementioned late stages of the phase of the positions during the step release said additional energy to further enhance the push of the driven limb and the user's body. 2. The method according to claim 1, wherein said straightening step includes straightening a concave surface formed by an arcuate upwardly curved middle portion of the foot of said curved foot. 3. The method according to claim 1, in which the integral elastic element of said prosthesis forms said ankle joint and said lower leg post. 4. The method according to claim 1, in which the accumulation of additional energy in an artificial muscle includes tensioning a viscoelastic material provided on at least one of the foot, ankle joint and the leg of the prosthesis. 5. The method according to claim 1, comprising the step that the ability of the prosthesis to store energy is regulated by applying a tensile load to the artificial muscle before using the prosthesis. 6. The method according to claim 5, in which said application of the preload includes the use of at least one of the cam, gasket and chamber containing pressurized fluid to tension the artificial muscle. 7. The method according to claim 1, comprising providing artificial muscle on the foot of the prosthesis. 8. The method according to claim 7, in which said foot contains a foot keel, while said artificial muscle connects the rear and front plantar sections of the foot keel. 9. The method according to claim 7, in which said foot comprises a sheath of the foot over the lower extremity of the prosthesis, said artificial muscle connecting the posterior and anterior plantar portions of the foot sheath. 10. The method according to claim 1, comprising attaching said prosthesis to a receiving sleeve on a stump of the human body, and wherein said artificial muscle provided on the prosthesis continues from said prosthesis to said receiving sleeve. 11. The method according to claim 1, in which the integral elastic element of said prosthesis forms an ankle joint and ankle support, and at the same time, an artificial muscle is provided on said elastic element. 12. The method according to claim 11, further comprising providing artificial muscle on said prosthetic foot. 13. The method according to claim 1, comprising performing said artificial muscle using a viscoelastic material selected from the group consisting of rubber and polymer. 14. The method according to claim 1, further comprising, during said application of the force load to the prosthesis, measuring the force developed by the prosthesis and adjusting the ability of said artificial muscle to accumulate energy during said application of the force load depending on the measured force. 15. The method according to claim 1, comprising pre-loading the artificial muscle in tension to the aforementioned application of power load during the step, while the artificial muscle is configured to increase the potential energy of the prosthesis. 16. An elastic prosthesis of the lower limb, containing
a foot extending in the longitudinal direction;
elastic ankle joint;
elongated protruding upward elastic leg tibia above the ankle joint,
wherein the ankle joint and the tibia are made as an elastic element, and the tibia is a curved bulge forward from the lower end of the tibia to the upper end of the tibia, so that the concavity, which is backward located in the sagittal plane, is formed due to the bent tibia, and the strut the lower leg is capable of bending in the longitudinal direction during the step to accumulate and release energy to improve the dynamic response of the prosthesis during the step;
an artificial muscle provided on at least one of the foot, ankle joint and ankle of the prosthesis, the artificial muscle being configured to store energy during application of a force load on the prosthesis in the active jerky phase of the human step and in later stages phase position during the step, the release of the aforementioned energy to enhance the shock of the driven limb and the human body. 17. The prosthesis according to clause 16, in which said artificial muscle is pre-loaded with tension and made with the possibility of increasing the potential energy of the prosthesis. 18. The prosthesis according to clause 16, further comprising a means for controlling the ability of the prosthesis to store energy by applying an adjustable tensile preload to the artificial muscle. 19. The prosthesis of claim 18, wherein said adjusting means is selected from the group consisting of a cam, a gasket, and a chamber containing pressurized fluid. 20. The prosthesis according to clause 16, containing artificial muscle on the foot of the prosthesis. 21. The prosthesis according to claim 20, in which the foot contains a foot keel and said artificial muscle on the foot connects the rear and front plantar sections of the foot keel. 22. The prosthesis according to claim 20, in which the foot comprises a sheath of the foot over the lower extremity of the prosthesis and said artificial muscle on the foot connects the rear and front plantar sections of the foot sheath. 23. The prosthesis according to claim 16, wherein said artificial muscle extends between the prosthesis and said receiving sleeve on a stump of the human body and connects them when the prosthesis is in use. 24. The prosthesis according to clause 16, in which said elastic element is an integral element, which, in the ankle joint, is facing forward by a convex bend. 25. The prosthesis according to paragraph 24, in which said artificial muscle is provided on said elastic element. 26. The prosthesis of claim 25, wherein the artificial muscle is also provided on said foot. 27. The prosthesis of claim 16, wherein said artificial muscle is made at least partially of a viscoelastic material selected from the group consisting of rubber and polymer. 28. The prosthesis according to clause 16, further comprising a sensor for measuring the force exerted by the prosthesis during said application of the force load on the prosthesis, and a means responsive to the measured force by adjusting the ability of the artificial muscle to store energy during said application of the force load during said application power load depending on the measured effort. 29. The prosthesis according to clause 16, in which said artificial muscle has a wider cross-sectional area in the intermediate section along the length of the muscle.

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