Human-device Integration: Introducing the Biotensegrity Bridge

Most of us in O&P recognize the interface as being the core of an upper- or lower-limb prosthetic system, and we are familiar with what we're told is its basic role of accepting and transferring the stresses of weight bearing, suspension, and motion to the target limb while transferring user input to its downstream components. However, a deeper dive into the role of the interface is necessary if we are to progress beyond mere device acceptance and tolerance to a point of total device integration or embodiment—our ultimate goal. With so many different design approaches, fabrication techniques, and varying skills and experience levels, is it possible to apply universal interface standards? And with such a high level of rejection for upper-limb prosthesis wearers, how can we best ensure a favorable outcome? We can certainly agree the interface should be comfortable and remain reasonably attached to its target limb while static or in motion. But if we think about it, ideally it should create a seamless, synchronized connection between user and device, should efficiently transfer energy and information to the wearer, and continuously elicit a desire to be worn and utilized. If the connection is not harmonious, proprioception, kinesthesia, energy expenditure, perceived weight, wearer confidence, as well as limb and health are compromised and will contribute to rejection of the device. From my experience fitting upper-limb prostheses users, I have seen that the greatest reason for rejection was not the myoelectric hand or wrist, but the interface—or more specifically, the interface connection or lack thereof. When discussing human-device synchronization, it is imperative to understand the underlying bone of the target limb is indeed the primary mover, and that the motion of the bone, together with information from forces both within and outside the limb, transfer information to the brain via a vast network of proprioceptors and mechanoreceptors. When we encapsulate the bone, which has now been disarticulated at its end, its motion relative to the attached device becomes critically important. When the prosthetic interface lags behind skeletal motion, or translates, rotates, or otherwise slips on the limb, the information regarding this asynchronous behavior is quickly sent to the brain via mechanotransduction, or the conversion of mechanical stimulus into electrical activity. This constant transmission of information with every movement of the arm or leg reminds the brain that something artificial is attached and something artificial is occurring, with respect to the type, extent, and quality of motion, as well as affects the users' perception of device usefulness. This disparity between skeleton and attached device has largely been ignored, either through a lack of appreciation for its adverse effects on prosthetic performance and wearer health, or of a belief that nothing can be done about it short of surgical attachment through osseointegration. While I have documented in my clinic, with our partners, and through research, that bone control provides improved benefits and outcomes, the concept of biotensegrity provides added support and a new perspective to consider. To appreciate the mechanisms needed to achieve a nonsurgical synchronous linkage, we turn to a relatively new theory of structural biology, termed biotensegrity, and how we can apply it to interface biomechanics, standards, and outcomes. The term tensegrity was coined in 1961 by Buckminster Fuller, receiving inspiration from models created by his student Kenneth Snelson, a contemporary artist and photographer who termed the concept "floating compression." The idea of tensegrity is best described as a structural principle based on a system of isolated components under compression inside a network of continuous tension and arranged in such a way that the compressed members (usually bars or struts) do not touch each other while the prestressed tensioned members (usually cables or tendons) delineate the system spatially. In other words, the tensegrity structure works through tension rather than compression, is omnidirectional, and functions independent of gravity, being structurally stable and functional in any orientation. Stephen Levin, MD, coined the term biotensegrity in the 1980s in applying the principles of tensegrity to the body, recognizing the application of tensegrity architectures requires moving away from modeling human biological systems, including bones and joints as lever systems. Whereas the conventional approach views bones as global, load-bearing structures supported by tensile structures such as tendons, ligaments, capsules, fascia, and muscles, biotensegrity views the bones as separate floatingcompressive elements embedded in and supported by a continuous, global tensile network. The body gets its strength from the supportive tension of its fascia. Inherent in this structure, a change of tension anywhere within the system is instantly signaled everywhere else in the body chemically and mechanically. A total body response occurs by mechanotransduction (proprioception/balance/touch) and the entire system re-equilibrates to adapt to such stresses as occur due to changes in posture and motion, pathological conditions (such as amputation), as well as pain or external forces as experienced via external attachments (prostheses, orthoses, exoskeletons, wearables, etc.). The biotensegrity model allows us to more greatly appreciate the impact an amputation or multiple amputations have with respect to compensation and the mechanisms behind it. A great way to visualize biotensegrity and how it differs from the traditional view of the human body is to look at the differences between the wagon wheel and its bicycle counterpart. In a wagon wheel, the ground reaction force travels directly up the strut in contact with the ground and into the hub, and thus the strut must be able to withstand the full weight of the wagon, whereas in the bicycle wheel, ground reaction forces are distributed throughout the network of spokes that are under a pulling force, or tension and no single spoke bears the burden. One concern I have with osseointegration is that it transfers ground reaction forces directly into the bone shaft, just like the wagon wheel. This is one reason I recommend a small, load-bearing and force-distributing external interface to be used in conjunction with an osseointegrated device, to better distribute forces along and throughout the soft tissue of the limb as opposed to directly into the bone shaft. The Biotensegrity Bridge If we regard the upper and lower limbs as tensegrity structures, wherein the underlying bone is the bicycle hub, the soft tissue the spokes, and the prosthetic interface as the rim of the wheel surrounding it all, it is easier to understand how important it is that the bone (hub) remains stationary with respect to the wheel, and how important fascial tension (spoke tension) would be in stabilizing the bone (hub). A traditional socket with its light, global tension is incapable of preventing large displacements of the bone shaft within the limb it surrounds, and this has far reaching implications for upper-and lower-limb prosthesis users. What is needed is an optimized interface structure that reduces significant displacement of the bone of the target limb and efficiently transfers both energy and information to the user, while still eliciting a desire to be worn and utilized. This nonsurgical, tensegrity-based structure, integrating dissimilar, nonbiological mechanical elements with biological organisms is what I have termed the Biotensegrity Bridge. Whereas tensegrity structures comprise a set of discontinuous compression components inside a tensile network, a biotensegrity bridge comprises a set of discontinuous compression components outside a tensile network. By pre-tensioning the underlying fascia via circumferentially spaced, longitudinally oriented compression zones, a tension-stabilized linkage is created. This linkage not only reduces relative motion of the underlying bone with respect to the interface but also firms up surface tension of the medium the interface relies on for its stability. It eschews the traditional approach of hydrostatic encapsulation and its reliance on lightly compressed, highly mobile soft tissue, which is responsible for the loss of stability and proprioception. In addition, the longitudinally shaped structures of the compression zones better match the shapes of the ligaments and large muscles and thus work with, as opposed to against, natural anatomy. So how does this relate to upper-limb interface strategies, clinical fittings, acceptance, and overall outcomes? First, let me address what I learned from my experience with the Luke Arm. When DEKA, the Luke Arm contract holder, awarded by the Defense Advanced Research Projects Agency (DARPA), asked me to assist with interface design, their upper-limb interface research was already well underway, but they had been unsuccessful in gaining tester acceptance. What became clear was that the limiting factor was not the prototype and advanced capabilities, but the interface design and its lack of biomechanical stability. The weight and functional capabilities of the early Luke Arms exacerbated the flaws in traditional interface approaches. Simply stated, the Luke Arm outpowered and outclassed the interface, causing wearer rejection due to pain and discomfort. I identified this was due to bone contact at the interfacial boundary while lifting, preventing testers from effectively and comfortably using the system. An altogether different approach had to be taken. To prevent skeletal strikes, I had to utilize an interface design that could greatly restrict intrinsic skeletal motion. I had already been We accomplished a 100 percent acceptance rate with the DEKA subjects while ensuring no adverse clinical effects. using a new interface design on select patients in my clinic, referred to initially as a Compression Release Stabilized (CRS) Interface with Skeletal Capture, but eventually called High-Fidelity Interface System, or HiFi. The Luke Arm project greatly accelerated testing of the limitations and capabilities of my design, especially due to our ability to assess the safety of osseostabilization with physician oversight. The DARPA project highlighted to everyone involved, especially the engineers, that superior mechanical and component design does not automatically ensure patient success and acceptance, and that a clinician's fitting approach can make or break the system. That led to biodesigns winning a Phase 2 DARPA contract focusing on the fitting system rather than the downstream components. The award was intended to create tools to enable inexperienced prosthetists to successfully fit heavy, multidextrous arms such as the Luke Arm, as well as to determine the metrics responsible for stable, effective, high-performing interfaces. The result was an effective, semi-automated, software-driven Adjustable Tether Fitting Tool (ATFT) and a Sensorized Imaging Tool (SIT) capable of ensuring correct paddle placement, correct compression-zone shape and taper, and finally, self-suspending on the limb for casting or scanning. The clinician was responsible for obtaining a proximal and distal circumference measurement and matching paddle length and width to the limb. After inserting the appropriately sized paddles into the ATFT and entering the two circumferences, the device would automatically set itself to the desired position, orientation, and taper in preparation for attaching a proximal and distal tether. The tether conformed around the limb shape but resisted radial expansion to ensure any applied compression was directed into the soft tissue and not simply forcing outward stretch of the assembly, allowing the clinician to select appropriate software-directed compression levels based on sensor input. The clinician then removed the SIT from the assembly, placed it in the proper orientation and adjusted the compression while watching the monitor, indicating when compression for each zone was in the green. This process worked equally well over plaster or when scanning directly over the limb. Our multisite study included eight subjects (three radial, three humeral, and two interscapulothoracic). At the end of the study, all the prosthetists successfully created their interim dynamic interfaces from start to finish without outside clinical assistance, relying only on the instructions and the hardware and software tools they were given. All eight subjects and participating prosthetists provided positive feedback on the tools as well as the overall fitting process. We again achieved a 100 percent acceptance rate, with all subjects stating they would wear the heavy system if they could use the interface. In summary, I am convinced, regardless of the level of amputation or amelia, if there is a limb length capable of being functionally captured with an interface, and there exists a desire for a prosthesis to be worn and utilized, then we as prosthetists should strive to create total embodiment of the chosen device. We should not simply attach adjustable straps and dials or 3D-printed aesthetic enhancements to try and make up for a biomechanically flawed method. Worse, we shouldn't congratulate ourselves on our patient's affirmation that the prosthesis is indeed comfortable and stays on. This is just the first step over the smallest of thresholds. We must do better than this. If our current standard of care designs were adequate, we would not have so many socket issues including a high degree of socket dissatisfaction, high abandonment rates, and high fall rates. While we have allowed many of our patients to achieve a level of independence, I challenge you all to seek more advanced, biomechanical, functional-based solutions. A structural design capable of synchronizing the motions of the underlying bone with those of the prosthesis while offering omnidirectional stability, enhanced proprioception, and the potential for cognitive indifference to its very existence, is a significant step in the right direction. While one can argue how best to achieve this, we are not fulfilling our responsibility to those who rely on us to help regain the function they have lost by clinging to the belief that it isn't a goal worth attaining, let alone chasing. Randall Alley, CP, is the CEO of biodesigns, Westlake Village, California. He can be contacted at ralley@biodesigns.com.

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