In the next section, it is explained how the investigation was performed for both patents and scientific communications. The Result Section presents a discussion about a new prosthesis classification according to this investigation, main authors, countries, and keywords analyzed. In the discussion Section, findings and other designs of prosthesis designs are disclosed.
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Active prostheses are considered state-of-the-art prostheses due to the use of actuators, microcontrollers, or other electronic devices; usually, these work using ESAR foot systems combined with some external elements such as actuators or other electronic components. These prostheses have better control and stability during a walk cycle [ 6 ].
CESR prostheses aim to capture the energy that is dissipated during a gait impact. On the loading phase of stance, energy is stored by a spring and locked. Then, this energy is timely released during the terminal stance of walking using microelectronic components [ 5 ].
ESAR, also known as ESR, was developed in the 1980s. This type of prosthesis uses a foot-modeled plate (usually carbon fiber made) that stores elastic potential energy and progressively releases it as kinetic energy [ 5 ].
Ankle-cushion heel (SACH-foot): This was developed in the 1950s and incorporated a compressible heel that dampens the impact on the ground while emulating a plantarflexion movement. This type of prosthesis is used for its relatively low cost and weight [ 4 ].
In order to improve and develop ankle/foot prostheses, it is necessary to know and understand present-day solutions to walking and running for BKA patients (and the people behind those solutions), so our designs meet both user and technical requirements. A state-of-the-art analysis of BKA prostheses is performed in this research.
Understanding the functioning of these prostheses is necessary to identify the foot movements: internal–external axial rotation, eversion–inversion, dorsiflexion (DF), and plantarflexion (PF), as shown in Figure 1 . The forces acting on the human foot are distributed with 60% towards the heel and 40% towards the phalanges. The loads are distributed between the heel and the metatarsals to the fourth and fifth phalanges and towards the big toe to the second and third phalanges [ 3 ].
In scientific documents, there is wide confusion with the terms prosthesis, prosthetic, and prostheses; prosthetic is the process to manufacture an artificial member (AM), prosthesis a component of the AM, and prostheses are all the components that make up an AM. From patents and scientific document searches, the term prosthesis is more commonly used; in this paper, prostheses and prosthesis will be used interchangeably.
Below-knee amputation (BKA) is a surgical procedure that mainly originates from trauma, diabetes, and peripheral vascular diseases [ 1 ]. While it is estimated that an average person walks about 6500 steps per day, current trends suggest that 10,000 steps per day represent a healthy lifestyle [ 2 ] for which a suitable prosthesis is necessary for a BKA patient in order to achieve a complete user reintegration to his/her pre-amputation activities. These designs should adapt to different patient’s activities.
For the literature analysis, the same keywords as for the patents’ search were applied in the Web of Science (WOS), obtaining 406 documents related to foot/ankle prostheses. The first filter was performed directly on the website, removing undesired keywords for a total of 136 documents. Subsequently, a second filter was applied, deleting repeated and undesired results. An individual document selection was made, resulting in 97 results. Finally, a bibliometric analysis was performed using data recovery software (R studio ® ) and a complement for bibliometric analysis (Bibliometrix ® ).
Based on a third filter, the results of all databases were merged, and keywords such as knee, orthosis, and tibia were eliminated. Duplicated results were filtered, and the remaining patents were individually analyzed for a total result of Derwent analytics (70), Espacenet (12), Google patents (19), Patentscope (72), and The Lens (151), resulting in 324 patents directly related to ankle and foot prostheses. From Figure 2 , it can be observed that although Google patents and Patentscope were the ones with more results, these contained a higher number of duplicates or undesired data.
Subsequently, data cleaning was performed using Open refine ® . The second filter was applied to eliminate duplicates, IPC categories that did not correspond, and keywords such as heart, valve, elbow, Arthroplasty, and Orthosis. An individual selection of the patents was made, and the unwanted results were eliminated. The remaining patents were as follows: Patentscope (369), Google patents (390), Espacenet (55), Derwent analytics (546), and The Lens (309), resulting in 1669 patents.
An initial filter was applied directly to the search engines where undesired categories and keywords were removed, in addition to a manual selection of patents directly on the website.
For the patent search, five different search engines were used, of which four were free-source, and one was paid. The databases were Derwent analytics (842 results), Espacenet (86 results), Google patents (5539 results), Patentscope (2281 results), and The Lens (778 results), with a total of 9526 results (see Figure 2 ).
For the patents and scientific communications searches, the following boolean operations were used under the International Patent Classification (IPC) A61F2 belonging to artificial substitutes or replacements for parts of the body: ((Ankle OR foot) AND (prosthetic OR prosthesis OR artificial)). Dates ranges were set from 2014 to 2020. For the patent analysis, 9526 documents were found. A scientific communications search provided 406 results. Figure 2 shows the results filtered on different search engines and the total number of documents obtained in every stage, among which The Lens was the most effective.
The United States (US) is the most productive country (46 documents), followed by Belgium (seven documents) and China (five documents). Some documents showed multiple country collaborations ( Figure 7 ). There is a clear relation between authors, journals, and countries. For example, most of the documents submitted in the US are from IEEE magazines and Plos One; meanwhile, Europe tends to apply to Prosthetic and Orthotic international and the American society of mechanical engineers (ASME).
On the authors’ part, Lefeber D. and Vanderborght B. are the top authors (11 articles each). Nevertheless, Hugh M. Herr is the most cited author in this field, with five of the most cited articles.
From the information obtained by the scientific documents, several aspects must be considered when designing a new prosthesis, such as aesthetics, which allows empathy between the users and their prosthesis [ 1 ], a size that permits the use of footwear, a mass corresponding to 2.5% of bodyweight [ 160 ] (literature shows an average of 2.5 kg for a 75 kg person), an ankle torque corresponding to 100–140 Nm, an ankle power between 250-300 W, and a device capable of storing and releasing energy (5–9 J)
After the final filter was applied, 98 scientific documents directly related to ankle/foot prostheses were selected; results are shown in Table 2 . Keywords were analyzed resulting in the top 10: gait (frequency = 15 articles), prosthesis (frequency =14 articles), prosthetics (frequency =13 articles), amputation (frequency =11 articles), biomechanics (frequency =11 articles), ankle (frequency = eight articles), transtibial (frequency = eight articles) prosthetic foot (frequency = seven articles), powered prosthesis (frequency = six articles), and gait analysis (frequency = five articles). This means there is a major trend in developing prostheses devices compared with gait studies or the creation of new methodologies.
Applicants and inventors in the databases were considered. Otto Bock Health Co. and Clausen Arinbjorn V. are the main applicants with ten and eight patents, respectively, from 2014 to 2020. Figure 5 shows the main applicants for BKA prostheses.
Among the results, 95 refer to foot prostheses, 65 to ankle prostheses, and 48 to a combination of both, of which 182 are removable, and 26 are osseointegrated. In this investigation, only removable prostheses are considered. Table 1 shows the selected patents, the technology used, and the type of prostheses. Among removable prostheses, 135 are mechanical or propelled with the body, hydraulic (18 results), and electronic or active (29 results). These results are distributed among ESAR, CESR, active, and hybrid (which did not match any of the aforementioned technologies or they are a combination of two or more categories). From Figure 4 , it can be observed that for electronic prostheses, 17 are active, three are CERS (use a controlled energy return without the use of complex devices), one is ESAR, and eight are hybrid. For hydraulic prostheses, four use electronic components, three are based on CERS, three on ESAR, and eight are a combination of three or more categories. For mechanical prostheses, 94 use ESAR systems exclusively, 26 combine different technologies (but mostly are mechanical), 13 are CERS (energy return is controlled using only mechanical devices), and two use actuators to release the energy.
Among the 324 results obtained, 208 results match prostheses designs, 51 match prosthetic mechanisms (motion blocking systems, aids to align prostheses, etc.), 22 match sockets, 11 match aesthetic covers, and 10 match joints. In total, 22 results are associated with methodologies (manufacturing methods, design methods, tests). Figure 3 shows these results; the number of prosthesis designs suggests a high interest in the development of new solutions for BKA amputees.
To compare the effectiveness during a walk cycle on uneven terrain, prostheses A, B, and C were analyzed using the same velocity and loads. Figure 17 shows a clear advantage of (C) over the other two models, thanks to the uneven deformation on its divided footplates, as shown for the displacement colored in red.
According to the structural analysis, B tends to offer major elastic energy compared to A and C, as shown in the instep colored in green/blue.
Most of the active prostheses use ESAR foot to generate enough power to initiate the gait cycle. From the patentometric and scientometric analysis, it is evident that types A, B, and C are the most used (see Figure 8 ). A structural analysis was performed to make a comparison between these types. Carbon-fiber footplates and a concrete floor were used. A load of 785 N was applied on the prosthesis upper faces obtaining a maximum deformation on the Y-axis of 0.63, 0.33, and 0.67 mm for types A, B, and C, respectively (see Figure 16 ). Meanwhile, deformations on A and C mostly occur on the ankle; B shows major flexibility along the foot. The red color shows maximum displacements on the foot connection with the body, but blue shows no deformation.
Another case is the robotic foot prosthesis made by Lapre [ 229 ]. This device aims to actively align the foot during different stances of the gait cycle using a four-bar linkage system to rotate and translate the foot with the use of a single actuator. It works using an ESAR foot and a DC motor (MaxonEC-30 200 W) that moves a Ball screw transmission via a belt drive. As this actuator system (motor and ball screw) contracts, it extends and shifts the foot center (see Figure 15 ).
Some designs do not correspond to the categories previously described. These designs are the pneumatic foot prosthesis by Huang et al. [ 189 ] (see Figure 14 A), where DF and PF are managed by two artificial muscles each, so stiffness and PF torque are easier to control. It is capable of emulating 3 DOF and is controlled via a desktop computer. Another design is the two DOF cable-driven ankle–foot prosthesis by Ficanha et al. [ 213 ], where instead of using pneumatic systems, it uses pulleys and Bowden cables that are externally controlled by two motors (Maxon EC-4), see Figure 14 B. Both systems have an external power source and are capable of emulating foot eversion and inversion movements.
CAS prostheses (see Figure 13 ) are mainly based on an ESAR foot (D), and in some cases complemented with a Cushion (E). The main goal of this prosthesis is the modulation of the stiffness during different stages of a gait cycle. This is granted by moving a Slider (G) along the length of the foot. Depending on the gait cycle, this slider moves forward and backward, providing the necessary stiffness to adapt to different situations such as walking, running, or climbing stairs, and it is controlled by a DC motor (C). A linkage system could be provided by a Ball screw transmission (F) or pulleys and belts. Motor (C) could be programmed to adapt to different activities. Housing (B) provides support for all the components and allows one degree of freedom (DOF) for the foot. The pyramid adapter (A) provides a connection between the transtibial components and the prosthesis.
Another powered prosthesis design is the LPP shown in Figure 12 . It aims to reduce the necessary power required by the actuators. It contains different Footplates (G and C), which in some designs (similar to the AMP Foot 2.1 [ 199 ]) are merged into a single plate. In another case such as the VSPA Foot [ 245 ], footplates (G) are individually controlled, allowing eversion–inversion movements; the DC motor (A) is located in a Housing (J) and rotates the Ball screw transmission (B), which moves the Footplate (C) up or down, allowing plantarflexion and dorsiflexion movement. Heel (D) may be composed of a flexible plate; ankle stiffness is provided by Springs (H) and (E). Depending on the model, two Springs (H) are used when there are individually controlled Footplates, and Spring (E) is used when (G) and (C) are merged. In this case, Spring (E) is attached directly to Footplate (C). Spring (E) is elongated using a Pulley system (F) connected to the Footplate (C). The pyramid adapter (I) provides a connection between the transtibial components and the prosthesis. Designs for this model use an external power supply that is not integrated into the main prosthesis body.
Active prostheses can be categorized by the components they use into three types: Multi-Array Prostheses (MAP), Low Powered Prostheses (LPP), and Controlled Adaptative Stiffness (CAS). For MAP, the form is similar to the one shown in Figure 11 . It uses an ESAR composite foot (E), and a DC motor (A), usually a 200 W Maxonconnected to a ball-screw transmission (C) that moves the linkage system (D) upward/downward and converts motor rotary motion into linear motion. In some cases, the motor is located instead of the spring (G) and connected to (C) using a timing belt. The linkage system (D) is in charge of connecting different mechanisms and allows plantarflexion and dorsiflexion movements; it may be composed of cables and/or pulleys, a bar mechanism, or crank sliders. F and G, depending on the prostheses, represent springs or actuators (pneumatic, electric, or hydraulic), for which torque varies from 100 to 140 Nm. Sometimes a parallel spring is aggregated due to the demanding torque requirements, and it aims to reduce the loads supported by the linkage system. Spring (G) saves energy during plantarflexion and dorsiflexion and supplements it during the swing phase. Housing (B) allocates all the electronic systems and provides stability to the system. The pyramid adapter (H) provides a connection between the transtibial components and the prosthesis. Some models have a lock mechanism, so the prosthesis could be used in a passive mode. See Figure 11 Figure 14 and Figure 15
For CERS prosthesis, the model by Endo Ken [ 129 ] (see Figure 10 ) considers a locking mechanism that preserves the energy storage in the spring. This energy is released upon the foot movement during the terminal stance. This impulse, in combination with the ESAR foot, provides necessary torque during the walk cycle.
There are some variations for ESAR prostheses that use a simple plate arrangement to adjust the return of energy (see Figure 9 A). Other designs use a single spring bar that regulates the energy storage/release (see Figure 9 B).
From the previous analyses, it can be determined that the general form for ESAR prosthesis is similar to the one illustrated in Figure 8 A and mostly differs in form; sometimes, a single talon plate is aggregated, or the disposition of the plates may vary. In other cases, as in Figure 8 B, the center of mass is moved, and the plates are rearranged. In the variation represented by Figure 8 C, the foot plates are divided, so the prosthesis emulates eversion and inversion movements. In Figure 8 D, some polymeric cushions are aggregated, replacing the use of extra plates. Figure 8 E shows the usage of different types of damping systems (springs, actuators, etc.) that replace some plates. All of these designs use pyramid adapters as a connection between the prosthesis and transtibial components.
ESAR prostheses are categorized into five different designs (see Figure 8 ). CERS and active categories are merged and divided into five different categories. There are some unique designs whose components cannot be grouped; these will be discussed individually.
From the selected patents and scientific documentation, a new ankle/foot prosthesis classification has been created besides ESAR, CERS, and active, based on its components and prosthesis functions.
The number of results per database does not reflect the effectiveness of each search engine. For this research, priority was given to search engines that provided useful data such as direct links to patents, the inventor’s name, and IPC codes. Nevertheless, there are some difficulties with some of them, such as the lack of options for filtering results or IPC categories, among others. Besides, some applicants may be included in the name of their companies (for example, Herr Hugh in Massachusetts Institute of Technology); this is because some search engines only show the applicant/owner’s name instead of the inventor. In some cases, there is a lack of consistency between the author’s names in different patents (for example, Smith Keith and Smith, Keith, B.); these kinds of inconsistencies were clustered, but still, results could not be entirely precise.
The United States has 56% of patent applications and 34% of scientific documents registered. These results do not necessarily display that they produce most of the knowledge on this topic, but because of the language, most of the search engines are capable of accessing the data, unlike languages such as Spanish, Chinese, or languages spoken in India. Therefore, some designs could remain undiscovered for this investigation.
Based on the obtained results, it can be established that for this study, the effectiveness per search engine is as follows: Derwent 8.4%, Google patents 0.34%, Patentscope 3.2%, The Lens 19.9%, and Espacenet 13.95%.
The classification of the 208 prosthesis patents related to prostheses designs was obtained according to the main technology used; results show that the ESAR mechanical prosthesis is the main patent object by 44%, although claims are different for each one. All of them can be classified based on the five ESAR categories presented in this document. Outcomes also show a tendency for the use of ESAR regardless of the technology used. For 151 removable foot/ankle patent prostheses analyzed, 53% use only ESAR-type prosthesis, and 90% use ESAR in its components. From these, the more commonly used were selected and compared using Ansys, with no major differences between A and C, but for B, results show a more elastic foot thanks to its mass-centered design.
The significant trend in the use of ESAR prostheses may be because of their lower cost and greater energy efficiency. Different designs are used according to the user’s lifestyle.
The minimum amount of components found for designing an active prosthesis is a DC motor, housing, a power transmission unit, a composite foot or equivalent, an energy storage device (springs, locking systems), a linkage system, an energy power supply, and a prosthesis/socket connector. From these components, most prostheses use a Maxon® Brushless motor between 12 and 200 W. Power variations are mostly due to the gear ratio used (the more power, the lower the gear ratio), springs with stiffness between 60–445 kNm, and a Li-ion battery between 12–24 V. From these components it is especially important to consider when designing a BKA prosthesis the linkage system that needs to support most of the necessary loads, and it must be capable of tolerating at least 2 kN (for an 80 kg patient) without any failure.
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Materials also play a vital role in supporting loads with 4000/5000 duty cycles per day; that is why aluminum, carbon fiber, and other composites are used in fabrication, and sometimes load reduction along the system is necessary and archived using a parallel spring arrangement.
The current development of batteries allows active prostheses to obtain enough power and charge duration without adding extra mass and weight, but for hydraulic and pneumatic prostheses, power supply currently is a problem because most of these systems are connected externally and the mass could reach up to 15 kg. Nevertheless, these systems are more efficient in mimicking human ankle movements.
For BKA prostheses, continuous growth in the development of active ones is estimated. Even though actual prostheses are capable of emulating three degrees of freedom, there is space for a complete body-integrated ankle/foot prosthesis.
Robotic leg prostheses promise to improve the mobility and quality of life of millions of individuals with lower-limb amputations by imitating the biomechanics of the missing biological leg. Unfortunately, existing powered prostheses are much heavier, bigger, and have shorter battery life than conventional passive prostheses, severely limiting their clinical viability and utility in the daily life of amputees. Here, we present a robotic leg prosthesis that replicates the key biomechanical functions of the biological knee, ankle, and toe in the sagittal plane while matching the weight, size, and battery life of conventional microprocessor-controlled prostheses. The powered knee joint uses a unique torque-sensitive mechanism combining the benefits of elastic actuators with that of variable transmissions. A single actuator powers the ankle and toe joints through a compliant underactuated mechanism. Since the biological toe dissipates energy while the biological ankle injects energy into the gait cycle, this underactuated system regenerates substantial mechanical energy and replicates the key biomechanical functions of the ankle/foot complex during walking. A compact prosthesis frame encloses all mechanical and electrical components for increased robustness and efficiency. Preclinical tests with three individuals with above-knee amputation show that the proposed robotic leg prosthesis allows for common ambulation activities with close to normative kinematics and kinetics. Using an optional passive mode, users can walk on level ground indefinitely without charging the battery, which has not been shown with any other powered or microprocessor-controlled prostheses. A prosthesis with these characteristics has the potential to improve real-world mobility in individuals with above-knee amputation.
Here we present a robotic leg prosthesis designed to replicate the key biomechanical functions of the biological knee, ankle, and toe joints in the sagittal plane while matching the weight, size, and battery life of microprocessor-controlled prostheses. As we will show, this level of lightness, performance, and efficiency is enabled by bioinspired actuation designs combined with a holistic design approach( 52 ). Using biomechanical analysis, analytical models, and dynamic simulations, we will show that a unique torque-sensitive actuator can replicate key biomechanical features of the biological knee joint by combining the benefits of series elastic actuators with that of variable transmissions. We will also show that an actuator based on a compliant underactuated mechanism can replicate the ankle and toe joint functions while regenerating substantial mechanical energy during walking. Finally, we will also show the ability to walk on level ground indefinitely without charging the battery. A prosthesis with these characteristics has the potential to improve real-world mobility in individuals with above-knee amputations.
Biomechanical studies of nonamputee gait suggest that the metatarsal (toe) joint plays an important function during gait ( 46 ). Individuals with metatarsophalangeal arthrodesis (fusion of the toe joint) have decreased step length and reduced plantarflexion moment in the affected side( 47 ). Studies with a passive prosthesis emulator show that the stiffness of the toe joint has a substantial effect on ankle power, center of mass, and push-off work during walking( 48 ). Moreover, a previous study has shown improvements in the metabolic cost of walking of individuals with below-knee amputations using a passive prosthesis with midfoot and metatarsophalangeal joints ( 49 ). Despite this evidence, only one research robotic ankle/foot prosthesis has been designed with a powered toe joint ( 50 , 51 ). Unfortunately, the added toe function comes at the cost of a substantial increase in prosthesis weight due to the addition of a dedicated toe actuator. Thus, we need new designs that can replicate the biomechanical function of the toe joint in a lightweight prosthesis.
An alternative design strategy for robotic leg prostheses consists of powering only a subset of activities ( 32 – 35 ), avoiding net-positive energy injection, or adjusting the mechanical behavior of the prosthetic joint without actively controlling movements or injecting energy( 36 – 39 ). By relaxing the actuation speed and torque requirements, these semi-active and quasi-passive prostheses can be made lighter, smaller, and still achieve longer battery life than fully powered prostheses. For example, designing the knee actuator to power only stair ambulation ( 33 , 34 ) or only the swing phase of gait ( 32 , 40 ) leads to a lighter and smaller prosthesis. Similarly, lighter and smaller ankle/foot prostheses can be developed by avoiding net-positive energy injection ( 41 ), limiting the active control of movements to non-weight bearing activities ( 38 , 42 ), or to adjusting the mechanical stiffness of the prosthetic joint ( 37 ). More recently, switching gears between different ambulation activities has been proposed to bridge the gap between semi-active and fully powered prostheses. However, this design solution comes at the cost of functionality, as the prosthesis cannot provide torque while switching gears( 43 – 45 ). Thus, semi-active and quasi-passive prostheses are typically lighter, smaller, and have longer battery life than fully powered devices but cannot replicate key biomechanical functions of the missing biological leg.
Robotic leg prostheses promise to improve the ambulation ability of individuals with lower-limb amputation by imitating key biomechanical functions of the missing biological leg with onboard actuation systems, sensors, and power supplies ( 10 ). Researchers have proposed different advanced actuation designs to efficiently provide the wide ranges of speed and torques needed to imitate the biological leg ( 6 , 7 ). Powered knee prostheses can be actuated using springs in series ( 11 ) and parallel to a motor ( 12 ), multi-joint actuators ( 13 ), antagonistic actuators ( 14 ), or high-torque density motors ( 15 ). Similarly, powered ankle prostheses can be actuated using four-bar mechanisms and polycentric designs ( 16 – 18 ), using a spring in parallel or in series to the motor ( 19 ), inductive charging ( 20 ), or implementing clutches and brakes in combination with motors ( 21 , 22 ). These advanced actuation systems have enabled two powered prostheses to reach the market. The Ottobock Empower ankle/foot prosthesis uses a series/parallel elastic actuator ( 23 ), whereas the Ossur Power Knee uses a clutchable series-elastic actuator ( 24 ). Unfortunately, after several years on the market, powered prostheses have failed to achieve clinical success ( 25 – 28 ). Although this negative outcome likely results from a combination of factors, there are key design limitations that affect function and usability of existing powered prostheses. These powered devices are much heavier, bigger, and have shorter battery life than their passive counterparts. Increasing the prosthesis weight affects both biomechanics and clinical outcomes negatively. During walking, larger prosthesis weight has been correlated to increased metabolic energy cost ( 29 ), stance-time and swing-time asymmetries ( 30 ) hip effort ( 31 ), and reduced socket stability. Increasing the prosthesis build height and the distance between the knee center of rotation and the top of the pyramid has limited the number of people that can be fitted to the prosthesis. Moreover, powered prostheses have much shorter battery life than their passive counterparts—a few hours versus a few days—which has a negative effect on usability in real life. We need to decrease the weight and size and extend the battery life of robotic leg prostheses to improve their clinical effect.
To date, most prostheses available to individuals with above-knee amputation are passive devices that cannot replicate key biomechanical functions of the missing biological leg. Most ankle/foot prostheses consist of a carbon fiber plate enclosed in a rubber foot shell ( 1 ). Some ankle-foot prostheses have an actual ankle joint actuated by passive elements such as springs and dampers. Virtually all available ankle/foot prostheses rely on the flexibility of the rubber foot shell to emulate the movement of the metatarsal joint (toe joint) and only one ankle-foot prosthesis available on the market has an articulated, passive toe joint ( 2 ). For above-knee patients, the ankle/foot prosthesis is connected to a prosthetic knee, which may have a single-joint or polycentric design, passively actuated by springs and dampers ( 3 ). In microprocessor-controlled prostheses, the mechanical impedance of the ankle and knee joint can be actively adjusted during gait to facilitate walking at a variable cadence while improving stability and reducing the risk of falls ( 4 ). However, they cannot actively generate movements or inject net-positive energy into the gait cycle ( 5 ), which are key biomechanical functions of biological legs ( 6 , 7 ). Prosthesis users compensate for these deficiencies with their residual limb and contralateral leg resulting in a slower, less efficient, and less stable gait compared to nonamputees ( 8 ). Since these passive prostheses cannot actively generate knee torque, climbing stairs and ramps or transitioning between sitting and standing is much more challenging for individuals with above-knee amputations than nonamputees. As a result, most individuals with an above-knee amputation are not able to ambulate in the community ( 8 ). Improvements in prosthetic technologies are necessary to address the unmet needs of the millions of individuals living with lower-limb amputation ( 9 ).
The proposed powered prosthesis—namely Utah Bionic Leg—consists of independent knee and ankle/foot modules ( ). The knee module uses a unique torque-sensitive actuator that works as a variable transmission to change the torque ratio passively, continuously, and quickly in response to varying knee extension torque by following a specific curve defined by the design geometry. The ankle/foot module uses a compliant, underactuated mechanism to power both the ankle and the toe joint. This compliant, underactuated mechanism transfers mechanical energy from the toe to the ankle joint, improving efficiency, while storing and releasing energy in a lightweight spring assembly. The knee and the ankle/foot modules have independent onboard power supplies and embedded electronic systems ( ). Both modules are designed based on ISO standards, facilitating their use outside the laboratory. The powered knee module of the Utah Bionic Leg has similar weight and size to the Ottobock C-leg—a microprocessor-controlled knee prosthesis (53). The powered ankle/foot module has similar weight and size to the Ottobock Meridium—a microprocessor-controlled ankle/foot prosthesis (54). Our powered knee module is slightly lighter (40 g, 2.3%) and has a shorter pyramid-to-knee joint length than the C-Leg Genium (3 mm, 11.5%). Our ankle module is slightly taller (4 mm, 5%) and heavier (70 g, 4%) than the Ottobock Meridium, although our powered device includes a custom force/torque sensor (55), which is not present in the Meridium. Most importantly, our powered knee and ankle/foot prosthesis is substantially smaller and lighter than the Ossur POWER KNEE™ (24) and the Ottobock Empower™ (23) —the only powered devices on the market. Similarly, powered prostheses designed by research laboratories that have not yet been marketed are between ~50% and ~100% heavier than the Utah Bionic Leg, which has an additional powered toe joint and embeds all the mechanical and electrical components (11, 56–58). Both commercially-available and research powered prostheses have a larger joint axis to pyramid length compared to the Utah Bionic Leg. Although battery life depends on many factors, the Utah Bionic Leg can be operated in passive mode, enabling the user to walk without needing to charge the battery. In this passive mode, the knee joint behaves primarily as a damper whereas the ankle joint behaves primarily as a spring, and they imitate the mechanical behavior of microprocessor-controlled knee and ankle/foot prosthesis. This functionality has never been shown by any existing powered or microprocessor-controlled prostheses.
Open in a separate windowOur torque-sensitive actuator is inspired by the biomechanical analysis of the biological knee joint. This analysis shows that the knee generates almost four times the torque in extension than flexion (1.21 vs. 0.3 Nm/kg) and the peak extension torque is more than double in stair ascent compared to level-ground walking (1.21 vs. 0.51 Nm/kg). In contrast, the knee extension velocity is three times higher during walking than stair ascent (339 vs. 99 °/s), although the peaks of the knee flexion velocity are similar between the two activities. The stance phase (foot on the ground) requires much greater torque and lower velocity than the swing phase (foot off the ground) for both level-ground walking and stair ascent ( ). Thus, the biological knee joint produces wide ranges of torques and speeds during ambulation, but the peak of knee torque and speed are not simultaneous. Providing wide ranges of torques and speeds with a small and lightweight electrical actuator is challenging because mechanical power output and electrical efficiency decrease sharply outside narrow ranges of torque and speed (59, 60). Our torque-sensitive actuator aims to address this issue by altering the torque ratio in response to knee extension torque.
Open in a separate windowDynamic simulations show how the passively variable torque ratio affects the motor speed/torque requirements by comparing the performance of a torque-sensitive actuator to that of the same actuator fixed at the maximum or minimum of the torque ratio range ( ). These simulations suggest that using the actuator at a fixed high-torque ratio would not satisfy the winding voltage limit, primarily due to the high velocity required in the swing phase ( ). Moreover, the simulations show that using the actuator at a fixed low-torque ratio would exceed the maximum current limit during stair ascent, primarily due to the high torque required in stance ( ). In contrast, the torque-sensitive actuator can satisfy both motor torque and speed requirements by adapting the torque ratio during ambulation. In stance, when the knee torque is high and the speed is low, the torque ratio quickly increases from its minimum to its maximum, lowering the required motor torque for both walking and stair ascent ( ). In the swing phase, when the knee torque is low and the speed is high, the torque ratio stays close to its minimum, reducing the required motor speed as well as the inertial torque ( ). Because the knee speed is low when the torque is high ( ), increasing the torque ratio proportionally to the knee extension torque does not cause the motor to reach the winding limit even though the motor velocity is higher than it would be without the torque-sensitive actuator during stance (Figure S4). Thus, the simulations show that by leveraging the non-simultaneous peaks of torque and speed, the torque-sensitive actuator can reduce the motor speed/torque requirements, enabling a small, lightweight motor to efficiently provide the wide ranges of torques and speeds required for a knee prosthesis.
Our compliant, underactuated mechanism is inspired by the biomechanical analysis of the biological ankle/foot complex. This analysis ( ) shows that during walking, the toe torque is nearly proportional to the ankle torque for a large part of the stance phase (20–60% Stride), although the peak torque is much lower for the toe than the ankle (0.12 Nm/kg versus 1.35 Nm/kg). Moreover, the velocities of the two joints are comparable in magnitude and opposite in direction, peaking at −218 °/s and 280 °/s for the ankle and toe joint, respectively. As a result, the toe dissipates power whereas the ankle generates power. Thus, during walking, the ankle and toe torque are nearly proportional, and the combined ankle and toe power is smaller than the power of the ankle alone. This analysis suggests that a single actuator could power both the ankle and toe joint, requiring fewer mechanical and electrical components than using two separate actuators. Moreover, this analysis shows that the combined ankle and toe power is smaller than the power of the ankle alone. Powering both the ankle and toe joints in a prosthesis is challenging due to the stringent weight and size requirements. Our compliant, underactuated mechanism aims to address this issue by enabling a single actuator to efficiently power both the ankle and the toe joints.
Dynamic simulations show the function of the compliant, underactuated system during walking by comparing its performance to that of an equivalent actuator powering the ankle joint only. The energy flow analysis ( ) shows that, the ankle-only design requires 14.4 J/stride of electrical energy to produce 6.3 J/stride of mechanical energy at the ankle, achieving an overall efficiency of 43.8%. In comparison, the underactuated design requires 8.2 J/stride of electrical energy, achieving an overall efficiency of 76.8%. Thus, the electrical energy consumption per stride is 43.0% lower in the underactuated design ( ). The energy flow analysis also shows that the reduced electrical energy consumption is primarily due to the toe regenerating 4.5 J/stride of mechanical energy. Finally, the underactuated system shows lower energy losses than the ankle-only design (−1 J/stride of Joule heating and −0.5 J/stride of friction). These lower energy losses are due to the lower velocity and acceleration of the linear actuator in the underactuated design, which result in lower inertial torque and mechanical power output at the motor ( ). Thus, the simulations show that by leveraging the concurrent torque generation at the biological toe and ankle joints, a compliant underactuated design can enable a single actuator to power both the ankle and toe joint and also reduce electrical energy consumption.
The knee closed-loop step position tests ( ) showed rise times between 37 ms and 56 ms, depending on the position step size. Thus, the −3 dB bandwidth of the knee position controller was between 9.5 Hz and 6.2 Hz, which exceeded the position bandwidth of the biological knee (61). Output impedance tests showed that the minimum-backdriving torque measured by the external 6-axis load cell with the motors off was 0.3 Nm ( ), which was lower than other powered knee prostheses (3.2 Nm (58), 2.6 Nm (62), 3–5 Nm (56)). Based on system identification of the output impedance ( ), the damping and inertia at the output joint were 0.43 Nms/rad and 0.04 kgm2, respectively. The low minimum backdriving torque and output impedance of the knee module was essential to allow for passive swing motion (Video S3). Also, based on the reading of an external 6-axis load cell, the knee torque step response showed steady-state errors lower than 1.2% and a rise time of ~32 ms irrespective of torque direction, resulting in a −3 dB bandwidth of ~11 Hz ( ), which was greater than that of the biological knee (61). Since the torque-sensitive joint acted in extension only, and the torque step response was nearly identical in extension and flexion, we conclude that the torque-sensitive joint did not affect the bandwidth of the open-loop torque controller ( ). The motor current during the torque step response test was 18–26% lower in extension than in flexion, because the torque-sensitive joint only changed the torque ratio in response to extension torque ( ). The continuous motor current increased from 4.21 A (the rated value) to 9.91 A (a 135% increase), as the contact between the motor and the frame provided enhanced thermal dissipation (35) ( ). This enhanced thermal dissipation was critical for the powered knee module to provide biomechanically appropriate function without overheating (see Supplementary Methods). Locking the torque-sensitive joint removed the ability of the knee module to increase the torque ratio ( ). With the torque sensitive joint locked at the bottom end, the Joule heating losses increased by 140%, so the knee could not provide more than 90 Nm without overheating. This torque level was 40% lower than the 150 Nm limit with the torque-sensitive unlocked and is not sufficient for most people in the United States (75 percentile adult male and 46 percent female) (63) to naturally climb stairs or stand up from a seated position. Locking the torque-sensitive joint at the top end can address this torque limitation. However, it lowered the maximum knee joint velocity substantially due to voltage saturation at the motor, and the knee module could not perform natural swing phase movements in waking and stair ascent ( ), which may cause the user to scuff, stumble, and fall. Thus, the torque-sensitive joint was essential for the knee module to satisfy basic torque and speed requirements for ambulation.
Open in a separate windowThe ankle closed-loop step position tests ( ) showed rise times between 30 ms and 61 ms, resulting in −3 dB bandwidths between 12 Hz and 6 Hz. The sine-wave position test of the underactuated ankle mechanism showed that if the toe joint moves in phase with the ankle joint the peak velocity of the linear actuator reduced from ~25 mm/s to 1 mm/s, a 95% decrease ( ). Thus, the underactuated mechanism could substantially reduce the required motor speed when the toe and ankle joint moved in phase. The ankle torque step response ( ) showed rise times between 32 ms and 25 ms, depending on the torque direction. These differences were consistent with the measured stiffness of the series spring ( ) in tension (829 N/mm) and compression (1273 N/mm). For the tested conditions, the bandwidth of the ankle torque controller was between 14 Hz and 11 Hz, which exceed the torque generation bandwidth of the biological knee and ankle (61). The continuous current increased from 7.58 A (the rated value) to 10.30 A, a 36% increase ( ), as the motor was in contact with the frame. Thus, the thermal dissipation was necessary for the ankle module to provide biomechanically appropriate function without overheating (see Supplementary Methods). Output impedance tests ( ) with an external 6-axis load cell showed that the minimum-backdriving torque with the motors off was 1.2 Nm, which was about half of other powered prostheses tested under similar conditions (3.2 Nm (58), 3–5 Nm (56)). System identification estimated the damping and inertia at the output joint to be 2.18 Nms/rad and 0.35 kgm2, respectively.
Open in a separate windowClinically relevant ambulation goals derived from nonamputee biomechanics (64–67) provided a reference to assess the performance of the Utah Bionic Leg during ambulation with three above-knee amputee participants ( ). provides an in-depth assessment of the performance and main results are summarized hereafter.
Open in a separate windowDuring walking in standard mode, the knee joint was slightly flexed at heel strike (9.3±3.1°, mean ± S.D.) and stance knee flexion peaked at 13.7±5.0°. However, there were substantial differences between participants. For Participants 1 and 3, the stance knee flexion was minimal. So, the knee extension torque was generally low, and the torque-sensitive joint stayed close to its minimum. In contrast, Participant 2 showed physiological stance knee flexion and knee extension torque. The torque-sensitive joint extended as expected from simulations, causing the knee torque ratio to increase proportionally. Also, in standard mode, the ankle plantarflexion torque at push-off reached 1.50±0.10 Nm/kg, which was within 2% of normative value and within 5% of normative timing. The maximum angle and torque of the toe joint were 36.3±1.1° and 0.12±0.02 Nm/kg, respectively, which were close to biomechanical goals of 39.6° and 0.12 Nm/kg. In the swing phase, both the knee and the ankle joint trajectory closely followed the normative data. Walking in standard mode with the toe joint locked had a visible effect on the ankle kinetics and kinematics (Figure S10). The peak motor velocity increased by 47%, resulting in a 56% increase in the peak motor mechanical power. Due to the motor approaching the winding limitation, the ankle mechanical energy decreased by ~18%, from 0.17 J/kg with the toe unlocked to 0.14 J/kg with the toe locked. Despite the lower ankle mechanical energy, the electrical energy consumption increased by 24% with the toe locked. In passive mode, there was no active ankle push-off or visible stance knee flexion, similar to a passive prosthesis. Both the maximum ankle plantarflexion torque and dorsiflexion angle of the ankle were within 10.0% of the passive prosthesis reference. In the swing phase, the ankle angle is near 0°, and little active dorsiflexion movement is shown. Also, in the swing phase, the knee achieved 63.8±1.3° showing an extension timing within 0.7±0.2% of passive prosthesis reference, which allowed for proper foot clearance. Thus, the Utah Bionic Leg enabled the amputee participants to walk safely both in standard and passive mode, even though the standard mode better satisfied the clinically relevant ambulation goals. Moreover, locking the toe reduced electrical efficiency and mechanical power output of the ankle module.
When climbing stairs one step at a time ( ), the powered knee generated biomechanically accurate torque, peaking at 1.08±0.08 Nm/kg, within 1.4±1.4% of able-bodied timing. However, the prosthetic knee joint extended faster than the biological knee, which resulted in a peak mechanical power 28.5% higher than the physiological value (3.25±0.19 W/kg). In contrast, the maximum ankle plantarflexion torque (0.62±0.07 Nm/kg) was considerably lower than the normative value (1.27 Nm/kg). When climbing stairs two steps at a time, the knee provided up to 1.67±0.13 Nm/kg of extension torque and 4.71±0.36 W/kg of positive mechanical power, which corresponded to a 54.6% and 44.9% increase, respectively, compared to climbing stairs one step at a time. In the swing phase, the knee and the ankle module closely followed the able-bodied reference ( ), safely clearing the step and positioning the prosthetic foot in preparation for the next step to be taken. In the stance phase, the knee torque ratio increased proportionally to the knee extension torque until it reached its maximum for all stair climbing conditions. In the swing phase, the knee extension torque is low, so the torque ratio stayed close to its minimum ( ). Thus, the Utah Bionic Leg provided appropriate torque, clearance, and foot placement to enable climbing stairs both one and two steps at a time, which was not possible with microprocessor-controlled prostheses.
Open in a separate windowIn stair descent, the powered knee generated up to 1.05±0.12 Nm/kg of extension torque, which was within 25% of nonamputee data (1.34 Nm/kg). The ankle dorsiflexed up to 19.3±0.7°, slowing down the stair descent movement. Also, the toe showed a peak extension of 36.0±1.0°, enabling the prosthesis to maintain stable contact with the step. In the swing phase, the prosthetic knee flexed to safely clear the step, with timing similar to able-bodied data (1.5±1.7%). Also, in the swing phase, the ankle plantarflexed up to 9.9±0.2°, enabling the foot to contact the subsequent step with the toe first. Consistently with the controller in use (68), there were visible differences between the ankle kinetics and the nonamputee reference. Similar to stair ascent, the knee torque ratio reached its maximum during stance when the knee extension torque was high and it stayed close to its minimum in swing ( ). The Utah Bionic Leg provided sufficient torque and range of motion to enable descending stairs while placing the whole foot on the step.
The actuation performance during ambulation was assessed by estimating mechanical and electrical energy consumption based on the actuator models and the embedded sensors (see Supplementary Methods). These estimates showed that, when walking in standard mode, the knee joint absorbed 0.27±0.04 J/kg of mechanical energy per stride, while the ankle joint generated 0.23±0.03 J/kg and the toe absorbed 0.04±0.01 J/kg. After accounting for Joule heating and friction losses (Supplementary Table S6), the knee motor was estimated to regenerate 0.03±0.06 J/kg of electrical energy, whereas the ankle/toe motor was estimated to consume 0.35±0.05 J/kg of electrical energy. Thus, the total electrical energy consumption estimate in standard mode was 0.32±0.03 J/kg per stride. When walking in passive mode, the knee, ankle, and toe joints absorbed 0.15±0.01 J/kg, 0.03±0.02 J/kg, and 0.03±0.01 J/kg, respectively, of mechanical energy per stride. After accounting for Joule heating and friction losses (Table S6), the knee motor was estimated to regenerate 0.11±0.01 J/kg of electrical energy, and the ankle/toe motor was estimated to expend 0.04±0.01 J/kg of electrical energy per stride. Thus, the combined knee and ankle/foot prosthesis was estimated to regenerate 0.07±0.02 J/kg of electrical energy per stride.
Based on these estimates and accounting for the ~2 W electrical power consumption of the embedded electronic system, in standard mode, the proposed powered prosthesis would provide up to 7,730±905 strides and 15,460 steps with the current battery (2400 mAh). This battery life exceeded the 10,000 (69) and 1,500 (70)steps that nonamputee and amputee participants, respectively, take on average in a day. Moreover, in passive mode, the battery was estimated to recharge at the rate of 2.0±0.6 J per stride. Additional experiments with a single amputee participant confirmed the estimates of electrical energy by directly measuring the battery voltage and current in standard and passive mode (Figure S11). Specifically, the electrical energy consumption directly measured at the battery was 0.27 J/kg per stride in standard mode, 22% smaller than the estimated 0.35 J/kg value. In passive mode, the electrical energy regeneration measured at the battery was 0.1 J/kg per stride (Figure S12). As expected, the battery voltage decreased in standard mode and increased in passive mode (Figure S13). Thus, the Utah Bionic Leg could regenerate electrical energy during walking while providing similar kinematics and kinetics to a microprocessor-controlled prosthesis. This function could enable prosthesis users to walk on level ground indefinitely without needing to charge the battery.
Based on modeling, in step-over-step and double stair ascent, the knee motor injected 0.49±0.01 J/kg and 1.28±0.09 J/kg of mechanical energy per stride, respectively. Subtracting the energy losses due to Joule heating and friction, the knee motor required 0.89±0.10 J/kg and 2.52±0.29 J/kg of electrical energy to climb one and two steps at the time, respectively. In contrast, the ankle motor regenerated small amounts of electrical energy (Table S6). Thus, the combined knee and ankle/foot prosthesis requires 0.88±0.17 and 2.55±0.29 J/kg of electrical energy per stride for stair ascent and double stair ascent, respectively. In stair descent, the knee motor absorbed 1.06±0.05 J/kg of mechanical energy per stride, regenerating 0.17±0.20 J/kg of electrical energy. The ankle motor absorbed 0.12±0.09 J/kg of mechanical energy per stride and regenerated 0.07±0.05 J/kg of energy. Thus, in stair descent, the combined knee and ankle/foot prosthesis regenerated 0.24±0.17 J/kg of electrical energy per stride.
Providing wide ranges of torque and speed is necessary for powered knee prostheses to replicate the biological knee biomechanics during ambulation ( ). Satisfying these requirements with a lightweight and compact actuator is challenging because efficiency and mechanical power output of electrical motors decrease sharply outside of a small operating range. Our analytical models show that a torque-sensitive actuator can address this problem by working like a passive variable transmission (see Supplementary Methods). Bench-top experiments confirm the model predictions by showing that when increasing the torque ratio proportionally to the knee extension torque, the required motor torque and current were substantially reduced ( ). Notably, the motor torque decreased proportionally to the torque ratio, but the Joule heating losses decreased quadratically ( ), which enabled a small motor (22-mm diameter, 170 g) to efficiently provide large torque at the knee joint. Although the required motor speed tends to increase as the motor torque decreases (Figure S4), the winding voltage limit is not violated because, during ambulation, the knee velocity is generally low when the knee torque is high ( ). The torque-sensitive actuator can generate physiological swing trajectories, which was not possible when the torque-sensitive joint was locked at the top end ( ). Thus, without the torque-sensitive joint, amputee users would not be able to safely ambulate with the proposed knee ( ).The knee torque bandwidth measured experimentally was the same in flexion (non-sensitive to torque) and extension (sensitive to torque) ( ), confirming that, different from series-elastic actuators (71), the compliance of the torque-sensitive actuator did not have a negative effect on the torque-control bandwidth. By combining a small, lightweight motor (22-mm diameter, 170 g) with a relatively low transmission ratio, the torque-sensitive actuator achieved low backdriving torque (0.3 Nm) and reflected inertia (0.04 kg m2) ( ). The low output impedance of the knee joint was essential for achieving electrical energy regeneration in walking. Notably, these performance improvements were obtained with a small and lightweight mechanism (~40 g) that fits into an extremely compact knee prosthesis (70 mm max width, 255 mm build height, 23 mm joint-pyramid distance, ). Thus, this study showed that the proposed torque-sensitive actuator was key to enabling lightweight and efficient knee prostheses to replicate key biomechanical functions of the biological knee joint during ambulation.
An articulated toe joint is necessary for a powered prosthesis to replicate key biomechanical functions of the missing biological foot (46, 47). However, adding a dedicated actuator for the toe joint inevitably increases the overall prosthesis weight, which has well-known negative effects on gait (72). Our experiments showed that powering the ankle and the toe joint with a single actuator could provide close to normative toe and ankle biomechanics in walking. Not surprisingly, using a single actuator results in a lighter and smaller design compared to using two actuators (51). More importantly, experiments showed that an underactuated design could be highly efficient because substantial mechanical energy was transferred from the toe joint to the ankle joint during ambulation (~3.3 J/stride, Table S6). In accordance with the simulations ( ), experiments with one amputee participant showed that locking the toe joint resulted in a substantial increase in the required motor velocity (Figure S10) and mechanical power, and consequently higher electrical energy consumption. Moreover, the mechanical energy at the ankle joint decreased, likely due to the motor reaching the winding limit. Thus, locking the joint had a substantial negative effect on the performance of the proposed ankle/foot prosthesis by spending more electrical energy to generate less mechanical energy at the ankle. Like other powered prostheses, the proposed underactuated mechanism uses a spring in series with the motor ( ). The proposed spring assembly is uniquely integrated with the ballscrew nut to achieve a lightweight (~125 g) and compact design (30-mm outer diameter) that fits into the narrow foot shell used for our prototype (40-mm width). The proposed compliant, underactuated mechanism is key to enable a lightweight and efficient powered ankle/foot prosthesis to replicate key biomechanical functions of the biological ankle and toe joint during ambulation.
Preclinical tests with three participants with an above-knee amputation showed that the Utah Bionic Leg provided gait kinematics and kinetics similar to nonamputee individuals ( , ). Of note, the toe joint worked as expected from simulation during walking, and it achieved an even larger range of motion in stair descent, which was not simulated. There were some visible differences between participants in early stance phase of walking. Participant 2 showed physiological stance knee flexion with kinematics and kinetics profiles closely matching the nonamputee references ( ). In contrast, Participant 1 showed a relatively small and constant stance knee flexion angle, and Participant 3 kept the knee fully extended against the mechanical end stop throughout early stance. Abnormal stance-knee flexion is often observed in prosthetics (56–58), and is due to habitual compensatory movements used by individuals with an amputation to walk with their prescribed prosthesis rather than the controls or mechanics of the Utah Bionic Leg. Furthermore, there were visible differences in the biomechanics of the Utah Bionic Leg and that of nonamputee individuals. In stair ascent, the powered ankle prosthesis did not actively push-off, leading to visible differences in ankle kinetics ( ). These differences were due to the controller in use (68) which did not implement ankle push-off, rather than the powered prosthesis mechanics. In stair descent, the powered prosthetic ankle provided lower torque than the biological ankle ( ). The ankle resistive torque was set during pilot tests based on the feedback received from the participants. Thus, the observed difference was due to subjective preference rather than the prosthesis mechanics. Despite these limitations, the preliminary clinical validation showed the potential of the proposed powered prosthesis to replicate the key biomechanical functions of the biological leg during walking, stair ascent, and stair descent.
In standard mode, the Utah Bionic Leg was expected to allow for 15,460 steps on single battery charge. This number was higher than the average steps taken by individuals with lower-limb amputation (1,500 (70)) and nonamputee individuals (7,500–10,000 (69)). Thus, the experiments suggest that the Utah Bionic Leg could support multiple days of use on a single battery charge, like microprocessor-controlled prostheses. However, it is worth noting that kinematics, kinetics, and net-energy injection change with walking speed in nonamputee individuals (73). This speed-dependent behavior could be replicated in a powered prosthesis (74) and may have a substantial effect on electrical energy consumption and battery life. When the Utah Bionic Leg was set to replicate the biomechanical functions of a passive prosthesis with a microprocessor-controlled knee (75) ( ), the mechanical energy dissipation at the knee joint resulted in the Utah Bionic Leg regenerating 2.0 J of electrical energy per stride. In this passive mode, a user could walk indefinitely even if the battery was depleted. This functionality is essential to real-world viability because users may not have access to a charger or may forget to charge the prosthesis. Thus, the Utah Bionic Leg offers functionality that is not currently available to individuals with above-knee amputation.
Simulations show that increasing the range of motion of the torque-sensitive actuator would have resulted in better performance. However, in the current knee implementation, the range of motion of the torque-sensitive joint is limited to keep the overall knee dimensions comparable to a microprocessor-controlled knee, resulting in suboptimal performance ( ). To reduce the negative effects of the suboptimal range of motion, we selected a relatively compliant spring, which causes the torque-sensitive joint to saturate before the maximum knee torque is achieved. Due to the saturation, the spring deflection cannot be used to measure the knee torque. Therefore, we could not implement closed-loop torque control. Bench-top tests against an external load cell show that the performance of the open-loop knee torque controller is adequate for a powered prosthesis. However, this result likely depends on the low friction and inertia of the current knee implementation and may not generalize. Also, the performance of the open-loop torque controller may degrade with time as the system friction and damping change.
The analytical model also shows that a torque-sensitive actuator can store and release energy during ambulation, theoretically reducing the required mechanical power at the motor. However, in the current implementation, we use a small, lightweight spring compared to series-elastic actuators with a similar output torque (20 g versus 200–500 g (11, 56)). As a result, energy storage and release are negligible (Figure S9). Another limitation is that the proposed torque-sensitive actuator increases the torque ratio only in response to extension knee torque ( ). This unidirectional torque sensitivity does not have a negative effect on the knee performance because the biological knee flexion torque is substantially smaller than the knee extension torque during ambulation ( ). However, there is no obvious way to achieve bidirectional torque sensitivity with the proposed design. Therefore, the benefits of the proposed torque-sensitive actuator are more substantial in joints like the biological knee or the ankle, which show a marked bias in torque. Thus, there are limitations related to the dimensions, energy efficiency, torque sensing, and sensitivity that can be achieved with the proposed torque-sensitive actuator. These limitations should be considered in future studies aiming to use the proposed torque-sensitive actuator.
Open in a separate windowThe proposed underactuated toe-ankle mechanism enables the development of a lighter and more efficient ankle/foot prosthesis compared to using two separate actuators. However, it cannot provide independent control of the ankle and toe joint. Thus, a key limitation of the underactuated design is that the ratio between the ankle and toe torque is fixed and cannot be changed based on the user’s needs or preference. Moreover, even with an underactuated design, adding a toe joint increases the prosthesis weight (approximately 100 g). Comparative tests with and without the toe joint or with different ankle/toe torque ratios are necessary to assess the effect of the toe joint on clinical outcomes. These clinical studies are necessary to justify the use of an articulated toe joint. Moreover, simulations show that a more compliant spring would have improved dynamic performance and electrical efficiency. However, it would have required a longer spring, which would have reduced the range of motion of the ankle and toe joints. These limitations are inherent to the proposed underactuated toe-ankle mechanism and should be considered in future designs.
Similar to most microprocessor-controlled and powered ankle/foot prostheses, the proposed ankle design does not have actuation in the frontal plane. Adding the frontal plane actuation is likely to increase the size and weight of the prosthesis. However, it may also improve clinical outcomes, especially when walking on inclines and rough terrains. Powered emulators (76, 77) and prostheses (78, 79) with passive or active frontal plane actuation have been developed to study the effectiveness of frontal plane actuation. Future work should consider implementing frontal plane actuation based on the outcome of these studies.
This design-validation study shows that the Utah Bionic Leg has the potential to replicate the key biomechanical functions of the missing biological leg for participants spanning large ranges of height (160–191 cm) and body weight (59–91 kg without the prosthesis). Our results suggest that the performance of the Utah Bionic Leg is appropriate for clinical studies assessing biomechanics and clinical outcomes in a statistically significant number of participants. Leveraging the modularity of our powered prosthesis, these clinical studies should assess the contribution of the three powered joints—the knee, the ankle, and the toe—to amputee mobility and subjective preference.
By replicating key biomechanical functions of the missing biological leg, robotic leg prostheses have the potential to improve ambulation for millions of people living with an above-knee amputation (9, 80). However, excessive weight, size, and short battery life have prevented existing robotic leg prostheses to achieve clinical success. In this paper, we show that a torque-sensitive actuator can enable a small, lightweight motor to efficiently provide the wide ranges of torques and speeds required for a knee prosthesis. We also show that a compliant, underactuated system can concurrently power the toe and ankle joint, and also enable substantial mechanical energy regeneration. Combined, these design solutions enable a robotic leg prosthesis to replicate the key biomechanical functions of the biological knee, ankle, and toe in the sagittal plane and match the weight, size, and battery life of microprocessor-controlled prostheses.
By enabling the development of prostheses that are both lightweight and powered, this study provides a scientific tool to study amputee gait mechanics and improve the mobility of individuals with above-knee amputations in real life. The Utah Bionic Leg can enable scientists to study both the effects of energy injection and the effect of active control on amputee gait mechanics without the confounding effect of prosthesis weight. Moreover, leveraging its lightweight design, future studies using the Utah Bionic Leg could include elderly and dysvascular participants, who lack the strength and balance required to use heavier powered devices. Moreover, the Utah Bionic Leg satisfies basic requirements for use at home. Thus, it may enable researchers to conduct studies outside the laboratory space, expanding the landscape of powered prosthesis research.
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