Importantly, transfemoral prosthetic runners must use a passive prosthetic knee to participate in official para-sport competitions. Existing prosthetic knees for running usually have few functions such as a simple hinge joint in the stance phase. Hence, running prosthetic knees induce a higher risk of falling than daily prosthetic knees. Moreover, the supply of running prosthetic knees is low. Currently, transfemoral amputees use almost the same prosthetic knee, regardless of their running abilities or skills. For this reason, for intact and transtibial amputee runners, they can contact the ground with their leg with its knee flexed, but transfemoral prosthetic runners require full extension of the prosthetic knee to run safely during the prosthetic stance [ 1 4 ]. Several studies have shown that hip extensor muscles play a major role in the process to keep the prosthetic knee fully extended during the prosthetic stance [ 5 6 ]. In addition, transfemoral amputees who run have to acquire skills to strike their prosthetic leg on the ground with the knee extended after the swing phase. It takes a certain amount of time to acquire these skills, thus, beginner-level runners have a high incidence of unintended knee flexion during weight-bearing on the prosthetic side. If a new running prosthetic knee has the function of knee flexion lock during the stance phase, even beginners who are not sufficiently trained can run safely. Transfemoral amputees can also choose a non-articulating prosthesis to remove the risk of falls with knee flexion. However, a non-articulating knee results in a stiff-legged gait, which requires compensatory motions that significantly increase the heart rate and oxygen consumption during walking [ 7 ]. Similarly, Highsmith et al. [ 8 ] suggested that an articulating-knee prosthesis reduces the ambulatory energy costs of transfemoral prosthetic runners when compared to using a non-articulating knee prosthesis. As noted above, prosthetic knees have several advantages. Therefore, a new prosthetic knee that mechanically controls knee flexion only in the stance phase is necessary to address these situations. The purpose of the present study was to develop a prosthetic knee mechanism to prevent unintended prosthetic knee flexion during the running stance phase and hence, for transfemoral amputees to run safely.
In recent years, lower-extremity amputees have been able to choose prosthetic leg parts to fit their individual lives and abilities due to the technological advancement of healthcare and engineering. Transfemoral amputees who are amputated at the thigh segment between the knee and hip joints attach a prosthetic knee to recover the knee function. Passive prosthetic knees are widely used for transfemoral amputees [ 1 ]. However, a previous study stated that 52.4% of transfemoral amputees reported falling in the past year, whereas 49.2% reported a fear of falling [ 2 ]. Existing prosthetic knees cannot support the body when the flexion moment acts on the knee joint by external forces, such as the ground reaction force, during the stance phase. Subsequently, the prosthetic knee is unintentionally flexed, causing a fall.
The condition for the stopper to remain in the gear is given by Equation (3).therefore, if the load () is sufficiently large, the condition of Equation (3) is maintained, and slip does not occur. The load () required for the stopper to remain on the gear is given by Equation (4), derived from Equations (1)–(3).as the running speed is slower, the joint reaction force (equivalent of) is smaller [ 10 ]. It is assumed thatat any running speed is larger than that during normal walking [ 10 11 ]. In order to restrict flexion at all running speeds, we set that the relationship betweenandduring walking as the maximum requirement. Preliminary experiments were therefore conducted to investigate the relationship betweenandduring walking and running. Figure 4 shows pilot intact subject data (sex: male, age: 21 years, weight: 63 kg, height: 1.70 m) of joint reaction force (equivalent of) and muscle moment (equivalent of) during the stance phase of level walking (1.50 m/s) and running (3.29 m/s) on a 10 m straight walkway., and= 20°,= 8°,= 0.3,= 0.032 m, and= 5.087) in Equation (4) were set with reference to this preliminary experiment.depends on the link lengths (links 1, 2, and 3). The relationship betweenandof Equation (4) is also shown in Figure 4 . If the relationship betweenandis in the gray shaded area, the stopper moves on the gear and the stopper eventually separates from the gear. If the same value ofis obtained with the proposed prosthetic knee, the maximum stress acting on the stopper tooth will be 24.88 MPa. However, the kinetics and kinematics of prosthetic legs during walking and running are different from those of able-bodied and intact legs, owing to prosthetic mechanical properties [ 5 17 ]. Subsequent measurements by amputees wearing the proposed prosthetic knee are necessary.
When the load pushing shaft A () down acts on the prosthetic knee during the stance phase, the force () that pushes the stopper into the gear arises ( Figure 3 ).is calculated from the principle of virtual work and shown in Equation (1).whereis the microscopic displacement of the stopper in the direction perpendicular to the long axis of the shank part, andis the microscopic displacement of the spur gear in the longitudinal direction of the shank part.
When the gear and stopper are engaged during the stance phase and when an external flexion moment is generated around the knee joint axis (Shaft A), the gear pushes the stopper back, owing to the tooth angles of the gear and stopper. If the external flexion moment increases, the force of the gear pushing the stopper back also increases. When the pushback force exceeds the maximum static friction between the teeth of the gear and stopper, the gear starts to slip on the stopper. The stopper then separates from the gear and the flexion lock is released. For this reason, the forces on the mechanism were analyzed statically.
As the prosthesis is off the ground, its lower leg moves downward relative to the thigh because of its weight. In other words, shaft A is located at the top of the slit during the swing phase. The stopper then separates from the spur gear. Therefore, flexion and extension are not restricted ( Figure 2 b).
Shaft A and the spur gear are attached to the inner ring side and outer ring side of the one-way clutch, respectively. The one-way clutch transmits torque in only one direction between the rings. Therefore, even when the spur gear and stopper teeth are engaged, the mechanism only limits flexion and allows extension.
As the prosthesis contacts the ground where the compressive force (load) acts on the proposed prosthetic knee along the long axis of the shank part, Shaft A moves downward. Simultaneously, the stopper moves through three links (links 1, 2, and 3 in Figure 2 ), and contacts the spur gear on shaft A to limit the rotation of the prosthetic knee joint ( Figure 2 a). This mechanism prevents unintended knee flexion during the stance phase, even if the prosthetic knee joint is not fully extended.
The CAD models of the transfemoral prosthesis assembly and the proposed prosthetic knee are shown in Figure 1 9 ]. The dimensions are length 143.9 mm, width 120.0 mm, and height 157.5 mm. The weight of the proposed prosthetic knee is 5.44 kg. The proposed mechanism has a single axis (shaft A) for the flexion and extension of the knee joint axis. Shaft A, which is fixed to the thigh socket via a socket connector, can move in the slit of the housing along the long axis of the shank part. A spur gear was placed on Shaft A. The stopper (gear ruck) was moved along the linear guide mounted on the housing. The stopper moves orthogonally to the long axis of the shank part through several links in accordance the motion of Shaft A. In the present design, when shaft A moves downward by 3.00 mm, the total movement distance of the stopper is 14.36 mm.
The main function of the proposed mechanism is to limit flexion during the stance phase to prevent unintended knee flexion. Extension during the stance phase allows the amputee to move forward easily even if the prosthetic leg contacts the ground with its knee flexed. No interference occurs with the flexion and extension during the swing phase.
The curved lines in Figure 9 indicate the relationship between the load on the prosthetic knee () and the moment (equivalent to) during the stance phase. The gray shaded area shows the condition that does not satisfy Equation (4). The curved lines were almost out of the area, except at the end of the stance phase.
The gate speed was 1.05 ± 0.03 m/s, and the stance phase of the prosthetic side was 51% GC. Figure 7 a illustrates the distance between the spur gear and stopper. As the prosthetic leg contacts away from the ground, the distance between the gear and stopper rapidly decreases, and the gear and stopper began to mesh at approximately 4% GC. The distance data indicated that the teeth were fully interlocked from 10 to 45% GC. Subsequently, the teeth were separated from each other, and the gear was able to rotate freely. The prosthetic knee joint angle began to flex at 45% GC ( Figure 7 b). The load on the prosthetic knee () was compressive (minus values in Figure 8 a) from 10 to 45% GC, synchronizing with the distance between the gear and stopper. The flexion moment was generated almost throughout the stance phase, except for 12–16% GC ( Figure 8 b).
Five successful trials of level walking were collected. All results were normalized to the gait cycle of the prosthetic leg. A foot contact and next foot contact of the prosthetic leg was defined as the beginning (0% GC) and end (100% GC) of a gait cycle, respectively.
Inverse dynamic analysis was performed on the intact leg side using AnyBody (AnyBody Technology) [ 18 ]. The body was modeled with upper body (head, trunk) and intact leg side (thigh, lower leg, foot) segments. Subsequently, 192 Hill-type muscle models were added to the lower limbs. The hip joint was defined as a spherical joint. The knee and ankle joints were defined as hinge joints. The mass, length, moment of inertia, and muscle anatomy of the body segments were calculated using a template model [ 19 ]. Kinematic and kinetic data were calculated from the measured GRFs and body coordinate data, respectively. An optimization method that minimizes the sum of squares of muscle activity was used to estimate the muscle forces. The muscles were grouped according to their function, as shown in Table 1 14 ]. Joint moments and muscle forces were normalized to body weight.
The kinematics of the prosthetic knee was determined based on the marker positions. Figure 6 illustrates the definition of the distance between each tip of the spur gear and stopper in the proposed mechanism and joint angles of the lower limb. When the value of the distance is negative, the mechanism locks flexion (fully interlocked by less than approximately −5 mm). Kinetic data were computed from inverse dynamics using kinematic data and measured GRFs.
One intact person (sex: male, age: 21 years, weight: 63 kg, height: 1.70 m) who obtained informed consent performed level walking on a 10 m straight walkway. The participant attached a simulated thigh socket and the prototype of the proposed prosthetic knee (on the right limb). The ankle joint of the prosthesis was fixed, and a 1D10 (OttoBock) was used for the foot. A total of 71 retro-reflective markers were attached to the bony landmarks and the prosthetic leg ( Figure 5 ). After the practice session, five successful trials each prosthetic leg and intact leg were recorded.
The purpose of the evaluation experiment in the present study was to confirm the functions of the stance phase. We developed a rough prototype of the proposed mechanism, and a pilot evaluation experiment was conducted with an intact participant using a simulated thigh socket. Since this evaluation experiment was a trial with an inexperienced subject, only the gait measurements were conducted to ensure safety.
The present study aimed to propose and analyze a prosthetic knee mechanism to prevent flexion during the prosthetic stance phase in transfemoral amputees. We developed a rough prototype of the mechanism and conducted an evaluation experiment.
The prototype of the proposed mechanism is completely passive and theoretically has no friction; therefore, the moment about the prosthetic knee must have been generated by structural restriction. The flexion moment, which was the reaction to the extension moment by external forces, was induced by the restriction of hyperextension. An extension moment of approximately 15% GC ( Figure 8 b) was induced by the proposed mechanism. A previous study stated that unintended knee flexion during weight bearing on the prosthetic side occurs, in particular, during the initial 40% of the gait cycle in transfemoral amputees [ 20 ]. The extension moment in the present study indicated that the designed mechanism functioned appropriately and prevented unintended prosthetic knee flexion. The slight flexion at approximately 15% GC ( Figure 7 b) is attributed to backlash between the gear and stopper. The rough prototype in the present study can theoretically flex a maximum of 15° even when the teeth are fully meshed. Although we could not evaluate the running conditions, the proposed mechanism would have appropriately functioned. A greater load acts on the prosthetic knee during running than during walking. The flexion–lock condition (Equation (4)) may be easier to maintain.
As shown in Figure 9 , the gait data were in the gray shaded area immediately before toe-off, indicating a risk of falling. However, the body weight had already shifted to the intact side in this terminal prosthetic stance phase, therefore, the participant would be able to continue walking without falling. If it is necessary to mechanically reduce the risk of falls, a possible solution would be to change the pressure angle (α) of the stopper. The prototype in the present study was set up with the pilot data of an intact subject. If α changes from 20° to 18° or 17°, the load on the prosthetic knee required for the stopper to remain at rest on the gear will be significantly reduced ( Figure 13 ). The maximum stress will be 1.61–1.97 MPa at the tooth root of the stopper of the present prototype. Theoretically, α = 0 is the best for maintaining locked flexion, but a larger pressure angle is ideal for smooth meshing and releasing of the teeth.
16,The gait patterns of the intact and prosthetic limbs were asymmetric. The proportion of the stance phase in the gait cycle was greater on the intact side. In addition, the knee flexion angles of the intact limb during the stance phase were larger than those of the prosthetic legs ( Figure 8 b). These temporal and kinematic results agree with those of previous studies that analyzed the gait of unilateral transfemoral amputees by using passive knees [ 14 21 ]. However, the result of the knee moment in the intact leg disagrees with that of a previous study [ 22 ]. As shown in Figure 11 b, an extension moment occurred in the knee joint in the loading response (1–11%). Subsequently, a flexion moment was observed until the swing phase was reached. This pattern is more similar to that of intact subjects than that of amputees in a previous study. The difference between the present and previous studies may be due to the muscles around the knee. The present study showed that the VAS generated greater forces in the loading response, which resulted in the knee joint extension moment. Subsequently, the VAS force decreased, and the HAM generated a flexion moment. In contrast, a previous study showed that the HAM generated greater forces in the loading response, which resulted in the knee joint flexion moment. Subsequently, VAS and RF caused an extension moment until the swing phase. These differences may have been induced by the prosthetic knee function. Further research with a larger sample size is required.
In our future work, we will develop a polished prototype with a swing control function in which the dimensions and material must be modified to reduce the mass and downscale the size of the prosthetic knee. In addition, an evaluation experiment on running with actual amputees is necessary.
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