There are over 40 million limb-deficient individuals worldwide, most of whom do not have access to any form of prosthetic care [1
]. The World Health Organization (WHO) has estimated that among this population, upper-limb (UL) loss or absence accounts for about 16% [2
]. Global estimates for congenital UL differences vary widely and range from 4 to 5/10,000 to 1/100 live births [3
]. Acquired UL amputations can be due to a wide variety of aetiologies, with the primary cause often being trauma. Unfortunately, most UL prosthetics currently available to patients are often neither affordable nor applicable in more challenging environments. The issue of appropriate prosthetics is even more pressing for children and adolescents, as there are fewer options available to them [4
]. There is immense variability between and within countries regarding how children with UL differences are treated [7
]. Active prostheses such as myoelectric devices are rarely fitted to the skeletally immature (this is especially true before the age of four) due to cost, weight, and/or muscle strength constraints [8
]. Most children are, therefore, typically provided with a passive device or a body-powered (BP) prosthesis as their first active device [6
]. These systems play a vital role in improving their gross motor development [10
]. Passive devices have been generally associated with higher rejection rates in children because of a preference for function over cosmesis [5
]. Providing suitable active devices to children and adolescents—and, consequently, improving their quality of life (QOL) as well as prosthetic outcomes in the long term—should thus be a priority.
The reported benefits of body-powered (BP) prostheses include silent action, lightweight, moderate cost, increased reliability, reduced training period, simple maintenance, proprioceptive feedback, and a simple operational mechanism that uses the body motions of the user to voluntary open or close the terminal device [11
]. Furthermore, a voluntary closing BP prosthesis provides the user with extended physiological proprioception [12
]. It extends the proprioceptive feedback to tools that users might engage with.
Various prosthetic options are available based on an individual’s level of UL difference [13
]. However, the most widely-used and affordable UL prosthesis across different settings and decades remains a Bowden cable-driven BP system [9
]. The earliest recorded model of a BP prosthesis is the ‘Ballif arm’, which dates back to 1818 [20
], making it now more than two centuries old. This purely mechanical technology still dominates much of the market despite the great efforts taken to offer the improved provision of artificial hands for the user population [18
]. There has been little progress in developing new approaches specifically for powerand control of BP devices compared to their myoelectric counterparts [11
]. As detailed earlier, BP prostheses have many advantages over externally-powered (EP) devices (e.g., myoelectric arms). However, despite these advantages, BP devices are rejected as frequently as EP devices for reasons like comfort (especially of the harness). The cost of even a standard (entry-level) BP device can be prohibitively expensive to own and maintain in low-resource settings. Moreover, BP prostheses usually require extensive fitting procedures from a high-skilled professional. Traditional BP devices rely on a harness system for their control that results in a restrictive operational workspace [26
]. Unfortunately, current BP prostheses are also associated with higher rates of device repair and maintenance compared to other device types [27
]. In addition, they often include cables and harnesses [30
] that many children and female users find uncomfortable or cumbersome [31
]. Users rejecting BP prostheses frequently describe the poor comfort and cosmesis associated with the traditional figure-of-eight and figure-of-nine harnesses [15
Current BP prosthetics typically necessitate high operating forces [34
], which could result in pain and fatigue during and/or after operation [33
]. BP prostheses are reported to be mechanically inefficient [34
] and offer a limited pinch force despite necessitating high cable operation forces from the user [34
]. The high operation forces for BP prostheses contribute to their relatively high rejection rates [15
]. Apart from fatigue and pain as primary drawbacks, high cable operation forces have also been found to deteriorate pinch force control accuracy in voluntary closing (VC) prostheses [38
]. Notably, Hichert et al. [39
] report that users of BP prostheses perceive and control low operation forces better than high forces. Consequently, their main advantage of offering feedback to the user is overshadowed, and the high operating forces negatively influence the comfort.
UL prostheses—and specifically BP devices—have generally witnessed poor outcomes, as the functional gain is limited compared to the disadvantages felt by the users [15
]. UL prostheses’ low success rates underscore the vast need for further improvements in this area [45
]. In a large-scale epidemiological study, Atkins et al. [49
] reported that users sought improvements in prosthetic control mechanisms, with the option to reduce the amount of visual attention needed. Ease of control and elimination of the harness altogether has remained a hope for BP device users [33
]. According to a 2015 survey [51
], participants generally still preferred novel techniques for UL prosthetic control that are not surgically invasive, wherein they expressed the highest interest in basic prosthesis features (e.g., opening and closing the hand slowly), compared to sophisticated features (e.g., touch sensation offered by some highly surgically-invasive techniques for bionic limbs) [52
]. Choosing the path between invasive and non-invasive UL prosthetics has remained an ongoing topic for general discussion in the field [56
]. BP prosthesis users’ design priorities and needs have remained virtually unchanged for decades [33
], and little attention has been given to broadening the design choices available to them. Recently, the “Self-Grasping Hand” was developed (for adults with hand absence) at the Delft University of Technology [57
] to address the need for a purely mechanical device that does not require harnessing. However, the Self-Grasping Hand is a passive-adjustable device that does not provide the user with continuous control over grasping movement compared to active UL prostheses. Besides, as opposed to an active prosthesis, most passive-adjustable devices require some involvement from the contralateral hand or the surrounding environment during the ‘grasp’ and ‘release’ phase. Another alternative is the “Wilmer Elbow Control”, which is a harness-free, elbow-controlled BP prosthesis (that uses the elbow’s flexion-extension movements as a control input instead of a traditional harness) [59
]. However, the obvious downside to this approach is that the elbow’s movement is coupled to the TD actuation. Hence, any unintended elbow motion leads to TD opening or closing, thereby restricting the user’s desirable operational space. Efforts are still underway in the quest for a more appropriate or fit-for-purpose BP prosthesis for both adults and children [6
Principally, one of the most significant challenges in this field has been the control of a UL prosthesis [43
]. Identifying novel, intuitive, and non-invasive control options for UL prostheses is an ongoing topic of interest for researchers, academics, and clinicians alike [43
]. The human body can generate a wide variety of control signals that could potentially be used to operate a prosthetic hand. Most practical control inputs typically originate from muscular activity around the arm and shoulder. The primary sources of control for BP prostheses are biomechanical. When body power is insufficient or undesirable, EP components may be utilised. The source of the “external power” comes from outside the human body. Contemporary versions are battery-powered electronic devices but can also rely on, for instance, pneumatic or hydraulic sources [16
]. Despite its numerous benefits, this ”external power” approach limits the operating time required for continuous use and increases the weight, complexity, safety requirements, and/or device cost. Furthermore, improvements in control are still sought [47
]. This is an issue in particular for paediatric users or in low-resource settings. Alternatively, in a BP prosthesis, motion from the musculoskeletal (MSK) system is usually harnessed to power the technology. Still, conflicts in control can arise when the MSK system is both the power generator and the system being augmented/supplemented. Thus, an independent modality that can drive this kind of assistive technology could create a new robust alternative source for device power and control.
The exciting idea is to provide a novel customised BP device that allows the user to disconnect the positioning of the artificial hand in the user’s workspace from the control and power requirements (which are currently imposed by the harness and cable). This implies that the user can freely place their prosthetic hand anywhere in space without it affecting the device’s operation or function. This is particularly promising for UL-deficient children and adolescents, as their prosthetic operating space should not be limited as they grow and develop. Prosthesis use in children has been found to affect the development of their brain, as well as their motor control strategies [72
]. This could have implications for long-term prosthetic outcomes. It is argued that a device powered and controlled by breathing input could expand the user options and address specific requirements that are difficult to meet with current BP prostheses. We can regulate our exhalation by forcing the diaphragm upward, and the controllable airflow can subsequently be used to power a small (custom-built and optimised) Tesla turbine [77
], which can then help to achieve accurate control of the artificial hand. This breathing-powered device provides a patient-specific prosthetic option that can be used without limiting any of the user’s body movements, compared to the traditional BP devices. This technical feasibility paper presents a novel type of BP prosthetic arm that relies on the user’s respiratory system to power and control the TD. It introduces a breathing-powered prosthetic hand design comprising a purpose-built Tesla turbine and a transmission system suitable for paediatric users.
This is the first proof-of-concept study demonstrating the feasibility of a novel breathing-powered UL prosthesis. We used virtual prototyping to explore the design parameters and optimise the design before physical prototyping. Furthermore, the demonstration with a volunteer shows the potential functionality of the proposed concept. This novel way of powering and controlling the device allows it to compete with the traditional Bowden cable-driven BP devices while simultaneously overcoming several limitations of a cable-driven approach.
Most activities of daily living (ADLs) necessitate fast speed and low grip force (e.g., typing, gesturing) [25
]. Tözeren suggests that a
s closing time is sufficient for prosthetic hands [95
], although Dechev et al. [96
] states a slightly slower 1.0–
s closing time is adequate for conducting ADLs. The demonstration model has a relatively longer closing time, but minor changes to the gearbox can create the closing times suggested above. Ideally, a clutch-based gearbox could be developed in the future to provide the option to select between appropriate operation speed and grip strength delivery. This stage of development can and should be driven by further user input.
One of the main limitations of the current device is the lack of optimisation in extracting power from the turbine, while careful consideration was taken in the turbine’s design to optimise its performance, there is still significant room for improvement. Further work should focus on building a suitable analytical/numerical model for the whole turbine assembly (rotor plus the inlet plenum chamber) validated by experimental investigations. Additionally, the experiments should ensure that there is proper control of the inlet mass flow rate and the shaft’s speed, as current studies are limited by pressure control [88
In addition, the manufacturing process of the rotor blades is reaching the edge of current machinery and conventional methods capabilities. Minimisation of the turbine’s size and weight, as well as the maximisation of the turbine’s power, relies on manufacturing very slender discs and spacers to a high precision – in this study, a 50 mm by mm (diameter × height) disc with a 10 mm by mm spacer for the gap were used. Further optimisation would need even thinner discs (to the order of 50 μm) while still keeping structural integrity under high centripetal loads.
The weight of the overall design is another important aspect that needs to be considered. The human hand on average weighs 400 g [113
% of the total body weight for men and
% for women [95
]. Although, prosthetic TDs of an equivalent weight have been described as being ‘too heavy’ by users [15
]. Since the forces from the prosthesis are borne by the soft tissue instead of the skeleton, the perceived weight of the TD is increased. Besides the overall device weight, the mass distribution affects the perceived weight of the system; hence, due consideration will be given to this aspect in the future. A range of 350–615 g is seen in current commercial prostheses and 350–2200 g in research-based hands [25
]. Within the prosthetics field, no set specification exists for the maximum weight of the prosthesis; the only agreed-upon specification is to minimise weight in general [25
]. The weight of the prosthetic device is a key contributor to socket interface discomfort and user fatigue. The prototype shown in the preliminary user testing weighs 429 g, with the wrist adaptor weighing 137 g. It should be noted that in this phase of the device development, the turbine is situated exterior to the TD, which adds mass to the overall system. This modular concept is useful for design exploration but should not be part of the final implementation. This study also used a gear module of
mm/teeth to reduce the transmission size. However, later we intend to explore a much smaller module in the order of 0.2–
mm/teeth to help further miniaturise the transmission. Essentially, there will be a research interest in reducing the current gearbox’s dimensions and weight.
In general, future work should focus on miniaturising the turbine and transmission to reduce the weight further. Ideally, the turbine will be integrated into the palm of the TD. This will directly reduce the weight of the overall system whilst also allowing for more aesthetically pleasing and anthropomorphic designs. In addition, future work could aim to use appropriate silicone gloves or conforming fingertip/palmar pads (i.e., friction pads on the TD’s volar surface to (i) increase the coefficient of friction and (ii) achieve compliance), which has been highly recommended in the mechanical design of prosthetic hands [25
The use of 3D printing and associated manufacturing techniques for our device has been extremely vital to iterate across designs and achieve a working prototype in a cost-effective and timely manner. Nevertheless, additive manufacturing for creating prosthetics needs to be carefully considered, as 3D printed prosthetics are still lagging behind conventional prosthetics in terms of, e.g., clinical evidence [110
]. Other manufacturing techniques should therefore be considered to scale the proposed system.
At the moment, the demonstration model has not been tested beyond the preliminary use of two volunteers. The volunteers were two boys aged nine and 11 who were asked to engage with the prototype without any acclimation time. The next steps will need to consist of validation of the device clinically by selecting (subjective and objective) outcome measures set within the WHO International Classification of Functioning, Disability, and Health (WHO-ICF) framework [116
] and the recommendations by Upper Limb Prosthetic Outcome Measures (ULPOM) Group [117
]. For example, the function can be tested by exploring the accuracy of patient’s control of a TD, dexterity, grasp, and speed.
The lightweight tube used for input is stretchable, unlike the Bowden cable. This means that the user can look away without issues. Initial positive responses have been gathered with regards to the use of this approach from a small group of patient representatives. However, the mouthpiece does still need to be placed in the mouth for operation, and there are several ways in which this can be placed near the mouth for easy access. Further studies will be needed to explore the most appropriate design from a user perspective.
Another essential user consideration is the hygiene and sanitation of these kinds of devices. Users will need to be made aware that periodically cleaning the device might be required. Similar hygiene issues also exist in the control of motorised wheelchairs using the ‘sip-and-puff’ technology. This technology has been around for several decades and is also safely applied in other assistive technology. The application of appropriate valves will further help to ensure the safe use of these systems over time. These aspects will need to be explored in clinical user trials.
It should be noted that the device presented in this paper helps decouple aperture control from the control of the position of the prehensor in space. The presented design can open or close the hand without the need for contact or the requirement to hold/move a particular body part (e.g., shoulder) in a certain position.
Micera et al. [119
] highlight that reliable and intuitive control of UL prostheses to obtain adequate dextrous manipulation requires sufficient feedback of prosthetic finger positions and pinch forces applied to objects. One of the problems with myoelectric devices is that they do not offer direct feedback of the forces or opening of the TD, as offered by a conventional harness and (Bowden) cable in BP prostheses. Although the absence of a harness is in some respects a great advantage of the myoelectric devices, the harness feedback ultimately enables the experienced user to feel the prehensor as a natural extension of their body.
Current BP prostheses offer the benefits of proprioceptive feedback [12
], and it has been shown that this proprioceptive feedback is superior to visual and tactile feedback [120
]. Conversely, a (myo)electric device requires its user to rely predominantly on visual feedback to estimate the end-effector position. However, auditory and tactile feedback from motor vibrations serves as a rough estimate of grasping forces, which is inferior to BP prostheses’ proprioceptive feedback. It will be interesting to explore whether or not the breathing-powered device provides users with suitable feedback. Nevertheless, like an electric device, this device currently offers visual and auditory (in the form of noise from the gearbox and turbine rotor) feedback during device operation. However, understanding these aspects better, or perhaps, incorporating another feedback mechanism in the device [121
] if required should be considered as future work.
Finally, it should be noted that the operation of the breathing-powered prosthesis does not seem to require any meaningful training time. There was no real acclimation time given in the user demonstrations included in this paper. The volunteers came in and used the device straight "out of the box" without any practice. They were just asked to explore the device, while recordings were made of their interactions. The users were able to directly operate the device due to the comprehensive responsiveness to a simple, intuitive, and controllable input. Reducing or even (almost) eliminating training times will positively influence user acceptance rates and provide a valuable directionality for what can be achieved with novel prosthetic designs.