Neurorehabilitation Technology 3031089944, 9783031089947

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Neurorehabilitation Technology David J. Reinkensmeyer Laura Marchal-Crespo Volker Dietz Editors Third Edition

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Neurorehabilitation Technology

David J. Reinkensmeyer ● Laura Marchal-Crespo ● Volker Dietz Editors

Neurorehabilitation Technology Third Edition

123

Editors David J. Reinkensmeyer Department of Mechanical Aerospace Engineering and Department of Anatomy and Neurobiology University of California Irvine, CA, USA

Laura Marchal-Crespo Delft University of Technology Delft, The Netherlands

Volker Dietz Spinal Cord Injury Center University Hospital Balgrist Zürich, Switzerland

ISBN 978-3-031-08994-7 ISBN 978-3-031-08995-4 https://doi.org/10.1007/978-3-031-08995-4

(eBook)

1st edition: © Springer-Verlag London Limited 2012 2nd edition: © Springer International Publishing 2016 3rd edition © The Editor(s) (if applicable) and The Author(s), under exclusive license to Springer Nature Switzerland AG 2022, corrected publication 2023 Chapter “Spinal Cord Stimulation to Enable Leg Motor Control and Walking in People with Spinal Cord Injury” is licensed under the terms of the Creative Commons Attribution 4.0 International License (http://creativecommons.org/licenses/by/4.0/). For further details see license information in the chapter. This work is subject to copyright. All rights are solely and exclusively licensed by the Publisher, whether the whole or part of the material is concerned, specifically the rights of translation, reprinting, reuse of illustrations, recitation, broadcasting, reproduction on microfilms or in any other physical way, and transmission or information storage and retrieval, electronic adaptation, computer software, or by similar or dissimilar methodology now known or hereafter developed. The use of general descriptive names, registered names, trademarks, service marks, etc. in this publication does not imply, even in the absence of a specific statement, that such names are exempt from the relevant protective laws and regulations and therefore free for general use. The publisher, the authors, and the editors are safe to assume that the advice and information in this book are believed to be true and accurate at the date of publication. Neither the publisher nor the authors or the editors give a warranty, expressed or implied, with respect to the material contained herein or for any errors or omissions that may have been made. The publisher remains neutral with regard to jurisdictional claims in published maps and institutional affiliations. This Springer imprint is published by the registered company Springer Nature Switzerland AG The registered company address is: Gewerbestrasse 11, 6330 Cham, Switzerland

Introduction to the Third Edition

When I want to discover something, I begin by reading up everything that has been done along that line in the past—that’s what all these books in the library are for. I see what has been accomplished at great labor and expense in the past. I gather data of many thousands of experiments as a starting point, and then I make thousands more. Attributed to Thomas Edison

The aim of this book is to provide a comprehensive overview of the ongoing revolution in neurorehabilitation technology. World leaders have taken the time to step back from their work, evaluate the state of the art in their field, and trace the development of their own work in creating this state of the art. We wish to provide a cutting-edge resource for those seeking to use, evaluate, and improve these technologies. There are four unique features of the book. First, we have attempted to ground the discussion of neurorehabilitation technology on neurorehabilitation science. Thus, you will find less information about the details of mechanical design or low-level machine controllers than information about the physiology of sensorimotor impairments, strategies for human-machine interaction, and the results of clinical testing. Second, we have chosen to emphasize movement rehabilitation after stroke and spinal cord injury, thereby focusing on the leading causes of disability and the largest user groups of neurorehabilitation technology. However, we note that many of the design principles discussed can transfer to a broader user group and that several chapters cover applications for people with neuromuscular disease, cognitive impairment, and cerebral palsy as well. Third, we have chosen to dedicate a greater amount of attention to robotic therapy than other approaches. This is because robotics therapy technologies have experienced the greatest growth over the past 40 years. Yet we recognize that other technologies—including neural stimulation, sensor-based devices, passive mechanical devices, exoskeletons, virtual reality, and mobile devices —show promise and are increasingly playing greater roles in the clinical delivery of rehabilitation, supplanting “traditional” robotic therapy in some cases. Therefore, we have expanded this third edition to include a greater amount of discussion of other emerging technologies. Finally, we have focused on therapeutic rather than assistive technology. That is, the emphasis here is on technologies that assist the motor system in v

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Introduction to the Third Edition

improving its intrinsic capacity to respond through training. We note however that the line between therapeutic and assistive devices is blurring because of emerging technologies like legged exoskeletons and spinal stimulation systems. We have organized the book into seven sections, which can roughly be divided into two halves. The first half of the book (Sects. 1–4) focuses on the design and implementation of neurorehabilitation technology. Section 1 contains three chapters that explain the relevant principles of neuroplasticity, motor learning, and sensorimotor recovery that can be used to inform neurorehabilitation technology development. Section 2 contains a set of chapters that exemplify how neurorehabilitation technology can be implemented to treat specific aspects of movement pathophysiology. Section 3 overviews principles of interactive rehabilitation technology—including issues of optimal challenge, psychophysical interaction, error manipulation, haptic interactions, expectations of end-users, and device implementation in a pediatric context. Section 4 contains three chapters on assessment technology and predictive modeling, all of which suggest working toward a precision medicine approach in rehabilitation. The second half of the book concentrates on specific technologies. Section 5 surveys a broad scope of neurorehabilitation technologies: spinal cord stimulation, functional electrical stimulation, virtual reality, wearable sensors, brain-computer interfaces, passive devices, mobile technologies for cognitive rehabilitation, and telerehabilitation. Sections 6 and 7 then provide detailed overviews of upper extremity and lower extremity robotics technologies. The book concludes with an Epilogue in the form of a debate over the current efficacy and ongoing potential of neurorehabilitation robotics. New chapters selected for the third edition include the neurophysiological basis of rehabilitation, the role of challenge for optimizing motor performance, the role of haptic interactions in promoting motor learning, computational neurorehabilitation, precision rehabilitation, spinal cord stimulation to enable walking, wearable sensors for stroke rehabilitation, mobile technologies for cognitive rehabilitation, telerehabilitation, body weight support devices, and the debate on the current and future value of rehabilitation robotics. Other chapters published in the third edition have also been substantially updated and reorganized to reflect the ongoing revolution. We hope that this book will inspire the next generation of innovators— clinicians, neuroscientists, and engineers—to move neurorehabilitation technology forward, thus benefiting the next generation of people with a neurologic impairment. Irvine, USA Delft, The Netherlands Zürich, Switzerland

David J. Reinkensmeyer Laura Marchal-Crespo Volker Dietz

Contents

Part I 1

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Learning in the Damaged Brain/Spinal Cord: Neuroplasticity . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Andreas Luft, Amy J. Bastian, and Volker Dietz

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Movement Neuroscience Foundations of Neurorehabilitation . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . Robert L. Sainburg and Pratik K. Mutha

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Recovery of Sensorimotor Functions After Stroke and SCI: Neurophysiological Basis of Rehabilitation Technology . . . . . Volker Dietz, Laura Marchal-Crespo, and David Reinkensmeyer

Part II 4

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Basic Framework: Motor Recovery, Learning, and Neural Impairment

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From Movement Physiology to Technology Application

Use of Technology in the Assessment and Rehabilitation of the Upper Limb After Cervical Spinal Cord Injury . . . . . . José Zariffa, Michelle Starkey, Armin Curt, and Sukhvinder Kalsi-Ryan Implementation of Impairment-Based Neurorehabilitation Devices and Technologies Following Brain Injury . . . . . . . . . Julius P. A. Dewald, Michael D. Ellis, Ana Maria Acosta, M. Hongchul Sohn, and Thomas A. M. Plaisier

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The Hand After Stroke and SCI: Restoration of Function with Technology. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 113 Mohammad Ghassemi and Derek G. Kamper

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Neural Coordination of Cooperative Hand Movements: Implications for Rehabilitation Technology . . . . . . . . . . . . . . . 135 Volker Dietz and Miriam Schrafl-Altermatt

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Contents

Robotic Gait Training in Specific Neurological Conditions: Rationale and Application . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 145 Markus Wirz, Jens Bansi, Marianne Capecci, Alberto Esquenazi, Liliana Paredes, Candy Tefertiller, and Hubertus J. A. van Hedel

Part III 9

Principles for Interactive Rehabilitation Technology

Designing User-Centered Technologies for Rehabilitation Challenge that Optimize Walking and Balance Performance . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 191 David A. Brown, Kelli L. LaCroix, Saleh M. Alhirsan, Carmen E. Capo-Lugo, Rebecca W. Hennessy, and Christopher P. Hurt

10 Psychophysiological Integration of Humans and Machines for Rehabilitation . . . . . . . . . . . . . . . . . . . . . . . 207 Vesna D. Novak, Alexander C. Koenig, and Robert Riener 11 Sensory-Motor Interactions and the Manipulation of Movement Error . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 223 Pritesh N. Parmar, Felix C. Huang, and James L. Patton 12 The Role of Haptic Interactions with Robots for Promoting Motor Learning . . . . . . . . . . . . . . . . . . . . . . . . . 247 Niek Beckers and Laura Marchal-Crespo 13 Implementation of Robots into Rehabilitation Programs: Meeting the Requirements and Expectations of Professional and End Users . . . . . . . . . . . . . . . . . . . . . . . . . 263 Rüdiger Rupp and Markus Wirz 14 Clinical Application of Rehabilitation Therapy Technologies to Children with CNS Damage . . . . . . . . . . . . . . 289 Hubertus J. A. van Hedel, Tabea Aurich Schuler, and Jan Lieber Part IV

Assessment Technology and Predictive Modeling

15 Robotic Technologies and Digital Health Metrics for Assessing Sensorimotor Disability . . . . . . . . . . . . . . . . . . . 321 Christoph M. Kanzler, Marc Bolliger, and Olivier Lambercy 16 Computational Neurorehabilitation . . . . . . . . . . . . . . . . . . . . . 345 Nicolas Schweighofer 17 Precision Rehabilitation: Can Neurorehabilitation Technology Help Make It a Realistic Target? . . . . . . . . . . . . . 357 W. Zev Rymer and D. J. Reinkensmeyer

Contents

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Part V

General Technological Approaches in Neurorehabilitation

18 Spinal Cord Stimulation to Enable Leg Motor Control and Walking in People with Spinal Cord Injury . . . . . . . . . . . 369 Ismael Seáñez, Marco Capogrosso, Karen Minassian, and Fabien B. Wagner 19 Functional Electrical Stimulation Therapy: Mechanisms for Recovery of Function Following Spinal Cord Injury and Stroke . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 401 Milos R. Popovic, Kei Masani, and Matija Milosevic 20 Basis and Clinical Evidence of Virtual Reality-Based Rehabilitation of Sensorimotor Impairments After Stroke . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 429 Gerard G. Fluet, Devraj Roy, Roberto Llorens, Sergi Bermúdez i Badia, and Judith E. Deutsch 21 Wearable Sensors for Stroke Rehabilitation . . . . . . . . . . . . . . 467 Catherine P. Adans-Dester, Catherine E. Lang, David J. Reinkensmeyer, and Paolo Bonato 22 BCI-Based Neuroprostheses and Physiotherapies for Stroke Motor Rehabilitation . . . . . . . . . . . . . . . . . . . . . . . . 509 Jeffrey Lim, Derrick Lin, Won Joon Sohn, Colin M. McCrimmon, Po T. Wang, Zoran Nenadic, and An H. Do 23 Passive Devices for Upper Limb Training . . . . . . . . . . . . . . . . 525 Marika Demers, Justin Rowe, and Arthur Prochazka 24 Mobile Technology for Cognitive Rehabilitation . . . . . . . . . . . 549 Amanda R. Rabinowitz, Shannon B. Juengst, and Thomas F. Bergquist 25 Telerehabilitation Technology . . . . . . . . . . . . . . . . . . . . . . . . . . 563 Verena Klamroth-Marganska, Sandra Giovanoli, Chris Awai Easthope, and Josef G. Schönhammer Part VI

Robotic Technologies for Neurorehabilitation: Upper Extremity

26 Forging Mens et Manus: The MIT Experience in Upper Extremity Robotic Therapy . . . . . . . . . . . . . . . . . . . 597 Hermano Igo Krebs, Dylan J. Edwards, and Bruce T. Volpe 27 Three-Dimensional Multi-Degree-of-Freedom Arm Therapy Robot (ARMin) . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 623 Tobias Nef, Verena Klamroth-Marganska, Urs Keller, and Robert Riener

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Contents

28 Upper-Extremity Movement Training with Mechanically Assistive Devices . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 649 David J. Reinkensmeyer, Daniel K. Zondervan, and Martí Comellas Andrés Part VII

Robotic Technologies for Neurorehabilitation: Gait and Balance

29 Technology of the Robotic Gait Orthosis Lokomat . . . . . . . . . 665 Laura Marchal-Crespo and Robert Riener 30 Using Robotic Exoskeletons for Overground Locomotor Training . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 683 Arun Jayaraman, William Z. Rymer, Matt Giffhorn, and Megan K. O’Brien 31 Beyond Human or Robot Administered Treadmill Training . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 701 Hermano Igo Krebs, Conor J. Walsh, Tyler Susko, Lou Awad, Konstantinos Michmizos, Arturo Forner-Cordero, and Eiichi Saitoh 32 A Flexible Cable-Driven Robotic System: Design and Its Clinical Application for Improving Walking Function in Adults with Stroke, SCI, and Children with CP . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 717 Ming Wu 33 Body Weight Support Devices for Overground Gait and Balance Training . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 745 Andrew Pennycott and Heike Vallery 34 Epilogue: Robots for Neurorehabilitation—The Debate . . . . . 757 John W. Krakauer and David J. Reinkensmeyer Correction to: Spinal Cord Stimulation to Enable Leg Motor Control and Walking in People with Spinal Cord Injury . . . . . . . Ismael Seáñez, Marco Capogrosso, Karen Minassian, and Fabien B. Wagner

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Index . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 765

Part I Basic Framework: Motor Recovery, Learning, and Neural Impairment

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Learning in the Damaged Brain/Spinal Cord: Neuroplasticity Andreas Luft, Amy J. Bastian, and Volker Dietz

Abstract

Keywords

Neuroplasticity refers to the ability of the central nervous system (CNS) to undergo persistent or lasting modifications to the function or structure of its elements. Neuroplasticity is a CNS mechanism that enables successful learning. Likely, it is also the mechanism by which recovery after CNS lesioning is possible. The chapter gives an overview of the phenomena that constitute plasticity and the cellular events leading to them. Evidence for neural plasticity in different regions of the brain and the spinal cord is summarized in the contexts of learning, recovery, and rehabilitation therapy.

Recovery Rehabilitation cord injury Brain lesion

A. Luft (&) Department of Neurology, University of Zurich, Frauenkliniksrasse 26, 8091 Zurich, Switzerland e-mail: [email protected] A. J. Bastian Department of Neuroscience, The Johns Hopkins School of Medicine, Kennedy Krieger Institute, 707 N. Broadway, Baltimore, MD 21205, USA e-mail: [email protected] V. Dietz Spinal Cord Injury Center, University Hospital Balgrist, Forchstr. 340, 8008 Zurich, Switzerland e-mail: [email protected]

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1.1

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. Stroke . Spinal . Plasticity

Learning in the CNS

Rehabilitation technologies that support movement recovery make use of different brain and body mechanisms, one of which is the brain’s ability to learn. Likely, the learning in the lesioned brain that mediates functional recovery is not identical to learning in the healthy state. Nevertheless, there are certain mechanisms on the cellular and the systems level, that can be broadly called neuroplasticity mechanisms, which are shared by healthy learning and recovery. Clearly, the main behavioral determinants of healthy learning of novel movements, activity, and repetition are also important in recovery. Hence, movement recovery may depend in part on motor learning. Motor learning is a general term that encompasses many different processes. Distinct behavioral and neural mechanisms are engaged depending on the level of complexity of the movement to be learned and the stimulus driving learning. A few different forms of motor learning are briefly reviewed. Motor adaptation is a type of motor learning that acts on a time scale of minutes to hours in

© The Author(s), under exclusive license to Springer Nature Switzerland AG 2022 D. J. Reinkensmeyer et al. (eds.), Neurorehabilitation Technology, https://doi.org/10.1007/978-3-031-08995-4_1

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order to account for predictable perturbations to a movement. Adaptation occurs on a trial-by-trial basis, correcting a given movement from one trial to the next. It is driven by sensory prediction errors, which represent the difference between the brain’s estimate of the sensory consequences of movement and the actual sensory feedback [1]. Once a movement has been adapted, it can be de-adapted when the predictable perturbation is reversed or removed. Discontinuation of training also leads to “forgetting” the adaptations over relatively short periods of time [2]. Associative learning can also occur on a time scale of minutes to hours. Classical conditioning is perhaps the most studied form of associative learning. It links two previously unrelated phenomena to improve behavior. For example, in eye-blink conditioning, a “conditioned” stimulus like a sound or tone can be repeatedly paired with a second, slightly delayed “unconditioned” stimulus like a puff of air to the eye [3]. Early in the learning process, the eye blinks in response to the puff of air (i.e., unconditioned response). However, with repeated exposure, the eye begins to blink when the tone is presented, therefore anticipating the air puff by closing the eye (i.e., conditioned response). This type of conditioning can be used to make associations between many types of behaviors. Motor learning can also be driven by feedback. Feedback can be given in close temporal association with the movement (knowledge of performance) and is used to adjust movement parameters on a trial-to-trial basis. It may be given later, e.g., after several repetitions, and can reflect the average outcome (knowledge of result) [4]. Performance feedback may be associated with a rewarding or discouraging element thereby inducing reinforcement learning [5] or avoidance learning [6], respectively. These learning processes can occur on short- or long time scales depending on the type and complexity of the movement. The rewarding or discouraging feedback can be explicit (e.g., earning money) or implicit [7], i.e., small improvements after repeating a novel movement, e.g., when learning to play a piano piece, are often not

A. Luft et al.

obvious or consciously perceived. Unconscious rewarding feedback may play a role. The conscious reward of playing the piece well typically comes late and temporally unrelated to the movement (e.g., the audience applauds). Thus, implicit motor learning may be mediated through use-dependent or Hebbian-like plasticity rather than reinforcement mechanisms. All these forms of motor learning rely on networks of neural structures rather than single areas, but there are some key regions that seem to play especially important roles in each. Adaptation is known to be cerebellum-dependent [8]. Classical conditioning can involve the cerebellum and hippocampus depending on the specific timing between stimuli [3]. Reward and avoidance learning are dependent on basal ganglia circuitry [9]. Use-dependent learning likely occurs at many levels of the nervous system, including spinal cord, brain stem, and cerebral (cortical and subcortical) structures. Complex motor skill learning induces plasticity in the motor cortex, especially during the consolidation of the learned movement [9–11]. Importantly, all forms of motor learning are dependent on cellular mechanisms of plasticity including gene/protein expression, synaptic and fiber growth, and functional changes such as long-term potentiation and long-term depression. As such, these mechanisms are reviewed below.

1.2

Mechanisms of Neuroplasticity in Learning and After Lesions

1.2.1 Gene Expression Learning a motor skill requires gene expression in the primary motor cortex (M1) [10, 11]. If this expression is pharmacologically blocked, learning is inhibited. Gene and subsequent protein expression is a common requirement of various learning processes [12, 13] as well as for cellular equivalents of learning, i.e., the changes in neuronal structure [14] and synaptic strength in the form of long-term potentiation (LTP) and depression (LTD) [15]. For motor skill learning,

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Learning in the Damaged Brain/Spinal Cord: Neuroplasticity

proteins are not only expressed during training but also thereafter while the subject is resting [10]. This delayed synthesis can be regarded as reflecting intersession consolidation processes [16]. The genes induced by learning are manifold, including immediate early genes (IEG, transcription factors). Expression of the IEG Arc in M1 was shown to occur specifically during skill learning but not during movement without learning [17]. Dopamine-related gene expression has been shown to be related to motor skill learning in mice [18]. Gene expression is induced by ischemia, especially in the peri-infarct cortex [19]. Some of these genes could also promote cellular plasticity offering the potential for stroke-induced plasticity as self-healing mechanism of the brain. Genes and proteins induced by ischemia include axonal growth stimulators while growth inhibitors are suppressed [20, 21].

1.2.2 Cellular Plasticity Long-term potentiation (LTP) and depression (LTD) are commonly seen as cellular equivalents of the brain’s learning abilities [22]. Either by repetitive stimulation, seen as the equivalent to repetitive training, or by synchronizing two signals that converge at one neuron, potentially reflecting associative learning phenomena, an increase in synaptic strength is induced that lasts from hours to days, termed LTP [23]. LTD is induced by low-frequency stimulation and leads to a lasting reduction in synaptic strength [22]. Both LTP and LTD have been described in various brain regions including the primary motor cortex (M1) [24]. The observation that the ability of M1 neurons to undergo LTP and LTD is reduced after training provides indirect evidence for the hypothesis that primary motor cortex LTP/LTD is consumed by training, and hence can be considered as one candidate mechanism of motor skill learning [25]. Two months after a skill has been learned in a 2 week training period and is well remembered, the synaptic strengthening that is observed in M1 shortly after training persists. But, the ability to

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undergo LTP has recovered and is now expressed on a higher level of synaptic strength [24]. In addition to changes in synaptic strength, the structure of neuronal networks is reorganized in association with motor skill learning. Apical and basal dendrites expand in association with skill training [26, 27]. This expansion is specific to the neurons involved in the control of the muscles used in the trained movement but not in the other musculature [28]. It remains open whether these changes are permanent or reflect a temporary expansion of M1 connectivity. Changes in dendritic spines, in contrast, were shown to be temporary and return to baseline 1 week after training has ended [29]. Microglia also play a role in plasticity by promoting synapse formation via a BDNF (brainderived neurotrophy factor)-mediated mechanism [30]. In the context of recovery after a brain or spinal cord injury, the role of LTP and LTD is unclear. LTP is facilitated in the peri-infarct cortex [31]. This result may be incompatible with the hypothesis that LTP is consumed during recovery as it is after healthy skill learning; hence, LTP would be reduced in the peri-infarct cortex not facilitated. But, the study lacks information about the recovery of function or lesion size, so a valid comparison to healthy learning is impossible, and the issue of LTP utilization during recovery is left unanswered. In the hippocampus, short-term ischemia leads to a disruption of LTP formation [32]. In humans, preliminary evidence indicates that LTP-like phenomena elicited in M1 of the lesioned hemisphere (cortical or subcortical lesions) by repetitive transcranial magnetic stimulation (TMS) predict good recovery in 6 months [33]. Paired associative stimulation (peripheral muscle and TMS stimulation of M1)—a potential human equivalent of associative LTP—can be elicited in the affected hemisphere M1, especially in those patients with limited deficits [34]. LTP-like phenomena are enhanced by serotonin [35] possibly explaining the beneficial effect of serotonin reuptake inhibitors in stroke recovery [36]. Hence, the ability of the lesioned cortex to undergo LTP may contribute to recovery.

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1.2.3 Systems Plasticity in the Brain Plasticity phenomena not only exist on the level of single neurons or networks but also in distinct functional systems of the brain. The input–output organization and the somatotopy of M1 undergo persistent changes during motor skill learning. Skill learning leads to an expansion of the cortical representation of the trained limb [37, 38]. Longitudinal motor cortex mapping experiments in rats show that this expansion is transient and is reversed after training ends although the skill is maintained [39]. In humans who continuously train new motor skills, e.g., professional pianists, task-related activation is smaller in the area and more focused [40, 41]. Musicians also have enlarged gray matter volumes in various areas of the cortex including the motor cortices [42]. The M1 of musicians contains memory traces of practiced skills that can be probed by TMS [43]. Representations in the primary motor cortex are also modified while recovering from a stroke. Initially, large areas of motor and adjacent cortices are recruited in the attempt to accomplish a movement as detected by functional magnetic resonance imaging (fMRI) [44, 45]. If M1 itself is lesioned, expanded activation is found in the peri-infarct cortex [46] or the premotor cortex [47]. As subjects recover, this hyper-activation is reduced, and movement-related activity focuses on the ipsilesional hemisphere contralateral to the moving limb [48–50]. If recovery is unsuccessful, cortices remain hyper-activated in the lesioned as well as the non-lesioned hemisphere which has been interpreted as a sign of a continuous attempt to initiate recovery [47]. But recovery is not only accompanied by cortical activation changes. Larger activation in the cerebellum ipsilateral to the moving limb [48] and smaller activation in the contralateral cerebellum are associated with better recovery [49]. Connectivity between different cortical regions in the brain is impaired after stroke, not only in areas in the vicinity of the lesion but also in the intact hemisphere [51]. There is reduced interhemispheric connectivity after stroke, especially between primary motor cortices [52].

A. Luft et al.

While movement-related activation observed with functional imaging methods demonstrates the brain areas that are involved in the control of the movement performed during imaging, TMS can directly assess the output efficacy and the viability of descending pathways in the lesioned hemisphere. Larger motor evoked potentials in response to TMS and the absence of ipsilateral responses to stimulation of the intact hemisphere are correlated with good functional recovery [53, 54]. Studies in animals and humans emphasized the importance of the GABAergic inhibition and the balance between excitation and inhibition in the peri-infarct cortex. Although the evidence from different studies in animals and humans using different methodologies is difficult to unify, it is assumed that inhibition is abnormally high early after stroke, then is reduced below the levels of healthy controls and returns to normal afterwards [55]. In addition, the interhemispheric (transcallosal) interactions of contralesional and ipsilesional motor cortices are altered after stroke and change during recovery [56], but it is yet unclear whether these observations are a cause or a consequence of (un)successful recovery [57].

1.2.4 Plasticity in Spinal Cord There is convincing evidence in cats with a transected spinal cord that use-dependent plasticity of neuronal circuits within the spinal cord exists [58, 59]. When stepping is practiced in spinal cat, this task can be performed more successfully than when it is not practiced, but the standing duration is not improved and vice versa, i.e., training of standing has only an effect on this task [60, 61]. The training effects of any motor task critically depend on the provision of sufficient and appropriate proprioceptive feedback information to initiate a reorganization of neural networks within the spinal cord. This is usually achieved by functional training. In contrast, following a complete SCI, when locomotor networks are no longer used, a neuronal dysfunction develops below the level of the spinal cord lesion in humans [62] and rodents [46].

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Learning in the Damaged Brain/Spinal Cord: Neuroplasticity

1.2.4.1 Spinal Reflex Plasticity The isolated spinal cord can exhibit some neuronal plasticity. Evidence for such plasticity at a spinal level has been obtained for the relatively simple monosynaptic reflex arc [62]. Monkeys could either be trained to voluntarily increase or decrease the amplitude of the monosynaptic stretch reflex in response to an imposed muscle lengthening [62], as well as to its analogue, the H-reflex [46]. The fact that the training effects persist after spinal cord transection [60] indicates that learning by spinal neuronal circuits is possible. Similarly, humans can be trained to change the gain of the monosynaptic stretch reflex ([63]; for review, see [64]). The idea that neuronal circuits within the spinal cord can learn is also supported by studies of spinal reflex conditioning. Simple hind limb motor responses to cutaneous or electrical stimulation are enhanced in animals with a transected spinal cord by reflex conditioning (i.e., pairing the stimulus with another stimulus that evokes a stronger motor response) [65]. These reflex responses become enhanced within minutes of conditioning indicating that synaptic efficacy along the reflex arc has changed, possibly through long-term potentiation [65]. 1.2.4.2 Task-Specific Neural Plasticity Today, it is obvious that there is also considerable task-specific plasticity of the sensorimotor networks of the adult mammalian lumbosacral spinal cord (for review, see [58, 59, 66]). The detailed assessment of the modifiability of neuronal network function is reflected in the research on central pattern generators (CPGs) underlying stepping movements [65, 67–69]. Observations made in spinal cats indicate that the lumbosacral spinal cord obviously can execute stepping or standing more successfully if that specific task is practiced. If the training of a motor task is discontinued and no similar task is subsequently trained, then the performance of the task previously trained is degraded [58]. Consequently, plasticity can be exploited for rehabilitative purposes using specific training approaches following a spinal injury.

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In the cat, recovery of locomotor function following spinal cord transection can be improved using regular training, even in adult animals [70, 71]. The provision of an adequate proprioceptive input to the spinal cord during training is essential to achieve an optimal output of the spinal neuronal circuitry with the consequence of an improved function. This aspect of training could meanwhile also be demonstrated for the locomotor training of subjects suffering an SCI [72]. Furthermore, in association with hind limb exercise, the reflex activity becomes normalized in adult rats following spinal cord transection [73]. Exercise obviously helps to normalize the excitability of spinal reflexes. Several neurotransmitter systems within the spinal cord (glycinergic and GABAergic systems) are suggested to be involved in the mediation of plastic changes following repetitive task performance [58]. In cats with a spinal cord transection, stepping can be induced by the administration of the noradrenergic agonist clonidine, which enhances the activity in spinal neuronal circuits that generate locomotor activity [74–76]. However, the application of Dopamine in patients with an SCI has no effect on the outcome of function [77]. Furthermore, serotonin seems to be involved in the production of locomotor rhythms [78]. Training paradigms of stepping and standing can modify the efficacy of the inhibitory neurotransmitter, glycine [58]. For example, when glycine is applied to a chronic spinal cat that has acquired the ability to step successfully, there is little change in its locomotor capability. If it is administered to a stand-trained cat, it becomes able to successfully step with body support [58, 66]. These findings suggest that the effect of glycine is so far specific in its action as it enables spinal networks to integrate proprioceptive input by reducing inhibition [75, 76].

1.2.5 Subcortical Contributions to Movement Learning The cerebellum is thought to use adaptive learning mechanisms to calibrate internal models

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for predictive control of movement. Such models are needed because sensory feedback is too slow for movements that need to be both fast and accurate; without internal models, corrections would be issued too late. Instead, the brain generates motor commands based on internal predictions of how the command would move the body [79]. This feedforward control requires stored knowledge (i.e., models) of the body’s dynamics, the environment, and any object to be manipulated. For example, recent work has demonstrated that cerebellar damage causes a bias in the brain’s representation of limb inertia relative to actual inertia, which results in characteristic patterns of reaching dysmetria, i.e., over- or under-shooting targets [80]. This specific deficit was confirmed in simulation and in behavioral studies of control subjects reaching with their limb while inertia is unexpectedly changed via an exoskeleton robot. Perhaps most importantly, this work also demonstrated a way of correcting this mismatch using cerebellar patient-specific compensations rendered by an exoskeleton robot. This suggests that there may be ways to compensate for biases in internal model representations using robotics. Unfortunately, cerebellar patients cannot learn to correct their internal model biases due to a loss of a cerebellum-dependent learning process often referred to as adaptation. Many studies have shown that the cerebellum is essential for adapting a motor behavior through repeated practice—it uses error information from one trial to improve performance in subsequent trials. It is important to note that cerebellum-dependent motor learning is driven by errors directly occurring during the movement (knowledge of performance), rather than knowledge of results after the movement is completed (e.g., hit or miss). Studies have suggested that the type of error that drives cerebellum-dependent learning is not the target error (i.e., “How far am I from the desired target?”), but instead what has been referred to as a sensory prediction error (i.e., “How far am I from where I predicted I would be?”) [1]. Damage to the cerebellum impairs the ability to adapt many types of movements, including reaching [81], walking

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[82], balance [83], and eye movements [84]. To date, there has been no systematic way to substitute or compensate for deficits in this form of learning. The microcircuit involved in cerebellar adaptation was first proposed by Marr [85], Albus [86], and Ito [87]. These works continue to provide the basis for many of the current theories of cerebellar function. Central to the idea of cerebellar involvement in learning was the discovery that Purkinje cell output can be radically altered by climbing fiber induction of long-term depression (LTD) of the parallel fiber-Purkinje cell synapse [88]. Hence, climbing fiber inputs onto Purkinje cells can be viewed as providing a unique type of teaching or error signal to the cerebellum. It has been shown that the climbing fiber may not simply be an all-or-none signal indicating error [89]. Instead, the duration of climbing fiber bursts is predictive of the magnitude of plasticity and learning, making it a graded instructive signal for adaptation. In addition to the climbing fiber-dependent LTD, there are many other sites of plasticity in the cerebellar cortex and deep nuclei that involve LTP and nonsynaptic plasticity (for review, see [90]). Thus, there are multiple avenues for activity-dependent plasticity to occur within the cerebellum over relatively short time scales. It is presumed that the plastic changes in cerebellar output are responsible for changing motor behavior during the process of adaptation which is processes that occur on relatively fast time scales. Purkinje cells are organized in microzones that either increase or decrease the output activity during learning (for review, see [91]). These zones interact with widespread regions of the cortex, thereby influencing learning processes across different domains of cognition and motor behavior. While the cerebellum operates on faster time scales, the brain’s reward system encodes the outcome of a behavior, i.e., the knowledge of the result, and can thereby influence learning (reinforcement learning) (for review, see [92]). Dopaminergic neurons in the substantia nigra (SN)/ventral tegmental areas (VTA) complex encode reward prediction errors and feed this information to the cortex. Especially neurons in

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Learning in the Damaged Brain/Spinal Cord: Neuroplasticity

9

Fig. 1.1 Schematic representation of the integration of reward circuits in the sensorimotor network. Via a dopamine (DA) signal encoding a form of immediate “implicit” reward, this signal can directly modulate synaptic plasticity in sensorimotor cortex synapses enabling motor skill learning

the VTA project to the primary motor cortex and are involved in motor skill learning. Ipsilateral dopaminergic projections from VTA to M1 [93] are specifically necessary for acquiring but not for performing a skill once acquired. Elimination of dopaminergic terminals in M1 [94] or destruction of dopaminergic neurons in VTA impairs the acquisition of a reaching skill in rats [95]. Dopamine modulates the excitability of M1 [96] and S1 [97] and, more importantly, is necessary for the formation of LTP in layer II/III synapses [94] that link M1 and S1 via horizontal connections. Plasticity at these synapses is involved in skill learning—as evidenced by the fact that the capacity of these synapses undergoing LTP is reduced after skill learning [17]. It seems plausible that the VTA-to-M1 projection relays signals of the same nature as compared to those that activate dopaminergic neurons from VTA to nucleus accumbens and prefrontal cortex, i.e., information about reward value and expectedness [98]. In the complex interplay of these circuits that include the basal ganglia (esp. the ventral striatum) and the cerebellum, it may be that dopaminergic neurons signal two forms of reward, one that reflects the outcome of

an action or goal attainment (explicit reward), the second a more immediate result of a single movement component [99]. This latter form is not consciously experienced but rather causes a “good feeling” about the ongoing movement and may be termed “implicit reward” (Fig. 1.1).

1.3

Learning and Plasticity During Rehabilitation Therapy

1.3.1 Lesions of Cortex and Descending Pathways Rehabilitative training is associated with neurophysiological alterations that are related to the improvement in motor function observed in individual stroke survivors [100]. Although correlation is not proof of causation, these studies provide an argument for neuroplasticity being one possible mechanism by which rehabilitative training can operate effectively. While bilateral arm training was associated with an increase in premotor cortex activation in both hemispheres that correlated with functional improvement in

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the Fugl-Meyer and Wolf tests [101, 102], conventional physical therapy (based on Bobath exercises) did not show altered brain activation despite being equally effective [102]. Conventional physical exercise may have utilized a mechanism other than those detectable by fMRI, e.g., by inducing changes in muscle, peripheral nerves, or spinal cord. Lower extremity repetitive exercises in the form of aerobic treadmill training likely utilize yet another form of brain reorganization to improve gait. As compared with stretching exercises, improvements by treadmill training were related to increased activation of cerebellum and brainstem as detected with fMRI of paretic knee movement [103]. Interestingly, the areas recruited in the cerebellum and brainstem corresponded to regions that control spinal pattern generators (cerebellar and midbrain locomotor region). These regions may have compensated for the loss of corticospinal projections that were injured by the stroke. It has also been shown that individuals with a cerebral stroke can improve walking symmetry using adaptive mechanisms of learning on a split-belt treadmill [104–106]. Repeated split-belt training over a 1 month time resulted in improvements in step length symmetry in chronic stroke survivors [107]. Importantly, the split-belt treadmill was used to augment the step asymmetry errors that the stroke survivors produced. This was done to drive a cerebellum-dependent learning process that would correct their error. Stated simply, making their error bigger drove the nervous system to learn how to correct it. After training, when they walked over ground, they had learned to correct their step length asymmetry. Training over 4 weeks led to improvements that lasted (and even improved further) for 3 months post training. Here again, the hypothesis is that intact cerebellar mechanisms are responsible for this form of motor learning. Hence, subcortical reorganization may be the mechanism to target in lower extremity, and particularly walking, rehabilitation. The availability of treatments that operate through distinct mechanisms may provide the rehabilitation clinician with many tools to

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individualize therapy for the patient. It seems likely that different patients with different brain injury and lesion profiles will require different therapeutic approaches.

1.3.2 Cerebellar Lesions Recovery from a first ischemic cerebellar stroke is often very good, with minimal to no residual deficits in up to 83% of patients [108–110]. On the other hand, individuals with degenerative cerebellar disorders tend to have persistent or progressively worsening clinical signs and symptoms [111]. One study has shown that people with damage to the deep nuclei do not recover as well as those with damage to only the cerebellar cortex and white matter [112]. Thus, the etiology of the lesion and the extent of damage are major indicators in recovery. There is a growing body of literature on the effectiveness of rehabilitation interventions for individuals with primary cerebellar damage. There are studies on the effects of rehabilitation interventions in this patient population, but most are noncontrolled small series (e.g., [113]) or case studies (e.g., [114]). Most work has been done on walking rehabilitation with common interventions including combinations of exercises targeting gaze, static stance, dynamic stance, gait, and complex gait activities [113, 114]. Dynamic balance activities in sitting, kneeling, and quadruped have also been advocated [113]. Individuals with acute cerebellar stroke seem to recover similarly regardless of whether they participated in a 2 week treadmill training intervention [115]. Further, individuals with superior cerebellar artery infarcts tend to show more severe ataxia than those with posterior inferior cerebellar artery infarcts early on, but both groups tend to recover to the same extent after 3 months. People with degenerative disorders tend to benefit more from rehabilitation training. Ilg and coauthors found that 4 weeks of an intensive coordination training followed by 8 weeks of home exercise could improve walking coordination and static and dynamic balance

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scores. It has also been shown that a 6 week home balance exercise program can improve balance and walking measures in people with cerebellar degeneration [116]. In that study, it was shown that the difficulty of the balance exercise is what predicted the best outcomes, with more challenging balance activities resulting in the greatest improvement. It was also shown that the effects of home exercise lasted for a month after therapy. In all these studies, it is not known whether such changes translate to improved real-world function. Locomotor training over ground and on treadmills, and with and without body weight support, has also been used with some success in single case examples [117, 118]. It is not clear how imbalance is corrected in the body weight support environment, however. With all gait and balance activities, it seems critical that the exercise be sufficiently and increasingly challenging, to facilitate plasticity in other intact areas of the nervous system [119, 120]. Two controlled trials investigated the use of cerebellar direct current stimulation on ataxia and found a positive effect in patients with neurodegenerative forms of ataxias (spinocerebellar ataxia, Friedreich’s ataxia, and cerebellar form of multiple system atrophy) [121, 122].

1.3.3 Spinal Lesions 1.3.3.1 Plasticity of Spinal Neuronal Circuits: Rehabilitation Issues Based on the knowledge gained from animal experiments, the aim of rehabilitation after stroke or SCI should be focused on the improvement of function by taking advantage of the plasticity of spinal and supraspinal neuronal circuits and should less be directed to the correction of isolated clinical signs, such as the reflex excitability and muscle tone. For the monitoring of outcome and the assessment of the effectiveness of any interventional therapy, standardized functional tests should be applied.

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1.3.3.2 Functional Training in Persons with a Spinal Cord Injury The coordination of human gait seems to be controlled in much the same way as in other mammals [123]. Even arm swing during locomotion represents a residual function of quadrupedal coordination of bipedal gait [124]. Therefore, it is not surprising that in persons with complete or incomplete paraplegia, due to a spinal cord injury, locomotor EMG activity and movements can be both elicited and trained similarly as in the cat. This is achieved by partially unloading (up to 80%) the patients who are standing on a moving treadmill ([72, 125, 126] for review, see [127]). In severely affected patients, the leg movements usually must be assisted externally, especially during the transmission from stance to swing. In addition, leg flexor activation can be enhanced by flexor reflex stimulation of the peroneal nerve during the swing phase [128]. The timing of the pattern of leg muscle EMG activity recorded in such a condition is like that seen in healthy subjects. However, the amplitude of leg muscle EMG is considerably reduced, corresponding to the paresis, and is less well-modulated. This makes the body unloading necessary for locomotor training. There are several reports about the beneficial effect of locomotor training in incomplete paraplegic patients (for review, see [71, 129, 130]), and patients who undergo locomotor training have greater mobility compared to a control group without training [131]. The neuronal networks below the level of an SCI can be activated to generate locomotor activity even in the absence of supraspinal input [75, 76, 132]. The analysis of the locomotor pattern induced in complete paraplegic patients indicates that it is unlikely to be due to rhythmic stretches of the leg muscle because leg muscle EMG activity is, as in healthy subjects, equally distributed during muscle lengthening and shortening phases [133]. In addition, recent observations indicate that locomotor movements induced by a robotic device in patients who are completely unloaded do not lead to a significant leg muscle activation

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[134]. This implies that the generation of the leg muscle EMG pattern in these patients is programmed at a spinal level and requires appropriate proprioceptive input from load and hip joint signaling receptors [123]. During daily locomotor training, the amplitude of the EMG in the leg extensor muscles increases and becomes better modulated during the stance phase, and inappropriate leg flexor activity decreases. Such training effects are seen both in complete and incomplete paraplegic patients [135]. The training effects lead to a greater weightbearing function of the leg extensors, i.e., body unloading during treadmill locomotion can be reduced during training. This indicates that even the isolated human spinal cord has the capacity not only to generate a locomotor pattern but also to show some neuroplasticity which can be exploited by a functional training [136–139]. However, only persons with incomplete paraplegia benefit from the training program in so far as they can learn to perform unsupported stepping movements on solid ground [135]. Neuroplastic changes also occur in elderly SCI subjects. This becomes reflected in a recovery of neurological deficits like that of young subjects. However, the translation of this improvement into a better function is significantly worse in elderly subjects [140]. Therefore, it is required to develop and apply specific and focused rehabilitation procedures in elderly subjects. In complete paraplegic patients, the training effects on leg muscle activation become lost after the training has been stopped [132]. Furthermore, after about 1 year after injury, complete paraplegic patients develop a neuronal dysfunction below the level of injury, especially in the long flexor muscles [141]. According to rodent experiments, this dysfunction is thought to be due to undirected sprouting within neuronal circuits [46].

1.3.3.3 Prerequisites for a Successful Training The spinal pattern generator must be activated by the provision of an appropriate proprioceptive feedback that leads to a meaningful muscle activation associated with plastic neuronal

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changes and consequently to an improvement of function [134]. Proprioceptive input from receptors signaling contact forces during the stance phase of gait is essential for the activation of spinal locomotor centers [134, 142–145] and is important to achieve training effects in paraplegic patients [135] (Fig. 1.2). Furthermore, hip joint-related afferent input seems to be required to generate a locomotor pattern [134, 146]. In addition, for a successful training program for stroke and SCI subjects, spastic muscle tone must be present as partial compensation for paresis [147]. Only in patients with moderately impaired motor function, a close relationship between motor scores (clinical assessment of voluntary muscle contraction: ASIA motor score) and locomotor ability exists. More severely affected

Fig. 1.2 Schematic demonstration of proprioceptive input during locomotor training in SCI subjects. The input from load and hip joint afferents was shown to be essential to achieving training effects

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SCI subjects require a threshold motor score which allows performing stepping movements. During the course of a locomotor training, subjects can achieve an improved locomotor function without or with little change in motor scores [138, 148, 149]. In these cases, a relatively low voluntary force level in the leg muscles associated with an automatic synergistic muscle activation leads to an improved ability to walk. A considerable degree of locomotor recovery can be attributed to a reorganization of spared neural pathways ([150]; for review, see [151]). It has been estimated that if as little as 10–15% of the descending spinal tracts are spared, some locomotor function can recover [152, 153]. In addition, by a training approach with the provision of appropriate proprioceptive input, directed, meaningful sprouting within neural circuits takes place below the level of lesion with the consequence of an improved recovery of function in the rat [46]. The improvement of locomotor activity might be attributed to a spontaneous recovery of spinal cord function that can occur over several months following a spinal cord injury [151, 154]. However, several observations indicate that the increase of leg extensor EMG activity also occurs independently of the spontaneous recovery of spinal cord function, as assessed by clinical and electrophysiological means [126, 137, 149, 151, 155]. Thus, functional training effects on spinal locomotor centers most likely contribute to an improvement of locomotor function in incomplete SCI subjects [126, 155], even in a chronic stage [149]. However, part of the recovery in locomotion corresponding to observations in the rat [58] might also be attributed to changes in muscle properties that occur during the training period.

1.4

Conclusion

Neuroplasticity mechanisms and training methods can improve function in patients with cerebral, cerebellar, and spinal cord injuries. However, patients with complete or almost complete hemi- or paraplegia do not, yet, profit from training because they cannot actively train.

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In the future, in these patients a combination of regeneration-inducing therapy and exploitation of neuronal plasticity possibly by using novel training devices could have a beneficial effect on the recovery of function. In this aspect, the research in spinal cord regeneration appears to be quite encouraging (for review, see [156]). Novel training devices (often referred to as rehabilitation robots) become increasingly important and popular in clinical and rehabilitation settings for functional training and standardized assessments if neurophysiological requisites are met [146]. Such devices allow a prolonged training duration, increased number of repetitions of movements, improved patient safety, and less physical demands for the therapists. Supportive therapies that enhance the brain’s potential to undergo plastic changes could supplement the training itself. For all these developments, testing in clinical trials will be required to prove efficacy and optimize the treatment for various disease and lesion types. Acknowledgements This work was supported by the P&K Pühringer Foundation.

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Movement Neuroscience Foundations of Neurorehabilitation Robert L. Sainburg and Pratik K. Mutha

Abstract

Research into the neural control of movement has elucidated important principles that can provide guidelines to rehabilitation professionals for enhancing the recovery of motor function in stroke patients. In this chapter, we elaborate on principles that have been derived from research on neural control of movement, including optimal control, impedance control, motor lateralization, and principles of motor learning. Research on optimal control has indicated that two major categories of cost contribute to motor planning: explicit task level costs, such as movement accuracy and speed, and implicit costs, such as energy and movement variability. Impedance control refers to neural mechanisms that modulate rapid sensorimotor circuits, such as reflexes, in order to impede perturbations that cannot be anticipated prior to movement. Research on motor lateralization has indicated that different aspects of motor control have been specialized to the two

R. L. Sainburg (&) Kinesiology and Neurology, Penn State University, University Park, PA 16801, United States e-mail: [email protected] P. K. Mutha Department of Biological Engineering and Centre for Cognitive Science, Indian Institute of Technology Gandhinagar, Ahmedabad, Gujarat 382424, India e-mail: [email protected]

cerebral hemispheres. This organization leads to hemisphere-specific motor deficits in both the ipsilesional and contralesional arms of stroke patients. Ipsilesional deficits increase with the severity of contralesional impairment level and have a substantial effect on functional independence. Finally, motor learning research has indicated that different neural mechanisms underlie different aspects of motor learning, such as adaptation vs skill learning, and that learning different aspects of tasks can generalize across different coordinates. In this chapter, we discuss the neurobiological basis of these principles and elaborate on the implications for designing and implementing occupational and physical therapy treatment for movement deficits in stroke patients. Keywords

.

Rehabilitation Motor control learning Motor lateralization

.

2.1

. Motor

Introduction

Deficits that result from strokes in sensory and motor regions of the brain represent a major impediment to the recovery of function in activities of daily living for stroke survivors. Such deficits most commonly include hemiparesis, a syndrome encompassing unilateral motor

© The Author(s), under exclusive license to Springer Nature Switzerland AG 2022 D. J. Reinkensmeyer et al. (eds.), Neurorehabilitation Technology, https://doi.org/10.1007/978-3-031-08995-4_2

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dysfunction on the side of the body opposite to the brain lesion, and spasticity, characterized by abnormally high muscle tone and atypical expression of reflexes. Occupational and physical therapy interventions often focus on reducing motor impairment, following stroke, by exposing patients to a range of movement activities, with a major focus on repetitive experience or practice. In general, the amount of practice corresponds to improvements in motor function, as measured by a variety of scales [1]. Unfortunately, gains made during therapy often show the limited translation to activities of daily living (ADLs) and carry over to the home environment. Over the past decade, rehabilitation approaches have incorporated technological innovations that can provide more cost-effective means of achieving higher intensity practice over longer periods of time. These computer-based and robotic technologies [2–5] have been shown to match, or even exceed the efficacy of traditional therapy in promoting improvements in motor performance [6]. However, these interventions hold greater promise than simply replicating traditional therapy, by providing therapists with an unprecedented ability to specify and measure movement features such as speed, direction, amplitude, as well as joint coordination patterns. As these technologies become more readily available in the clinic, the most pressing question is how therapists can best utilize them to accelerate recovery of function. In this chapter, we will discuss principles that have been derived from research in motor control and learning that could be applied to training strategies using computer-based movement interventions.

2.2

Principle 1: Optimal Control

While most therapists recognize that practice and repetition of motor activities lead to improvements in motor performance, a systematic identification of which movements should be practiced is often lacking. This is partly because the question of what defines a desirable movement has yielded no clear answer. Traditionally, a common guiding principle employed in occupational and physical therapy

has been to make movements more “normal”. Thus, the goal is to develop movement patterns that are similar to those exhibited by non-impaired individuals. This idea emerged from the observation that certain characteristics of movements made by healthy individuals are fairly similar within a given task, and even across tasks. For example, when reaching for an object in space, movement trajectories across healthy individuals appear fairly straight and smooth [7]. Such reliability of motor behavior is particularly interesting because of the abundance of possible solutions to most movement tasks and the variety of environments we move in. For example, when reaching for a cup of coffee in front of us, we have the choice of using one arm or both arms, standing up or remaining seated, leaning the trunk forward or reaching further with the selected arm(s), twisting our trunk to require shoulder abduction, or keeping it straight to require shoulder flexion, among other options (see Fig. 2.1). In addition, the relative motions between our body segments can produce a wide variety of curved trajectories of the hand, in order to procure the cup. Each possible motion can also be achieved at a variety of speeds, as well as a variety of possible muscle activation patterns. There are literally infinite solutions to this simple task. Regardless of these vast possibilities, people tend to display movement patterns that are consistent across different instances of the same movement or even across different movements, whether made by the same or different individuals. These similarities are often referred to as “invariant characteristics” of movement. Many studies have shown that when different people make reaching movements, invariant characteristics include approximate straightness of the hand trajectory and smooth bell-shaped velocity profiles (see Fig. 2.2) [7–11]. How do different people arrive at similar solutions within and across tasks despite the extensive redundancy in the musculoskeletal system and the diversity and uncertainty of the environments we move in? One way to arrive at the “best” solution when confronted with many different options is to employ optimization strategies when planning the movement. Optimization procedures have

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Movement Neuroscience Foundations of Neurorehabilitation

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Fig. 2.1 Different ways of picking up a coffee cup starting from the same initial posture. The left pose involves shoulder flexion and elbow extension. The middle pose involves flexion of the trunk, slightly less

shoulder flexion, and more elbow extension. The right pose shows some trunk flexion, shoulder abduction, elbow flexion, and forearm pronation

been developed for use in engineering applications, and seek the minimum or maximum for a given “cost function”, subject to a set of constraints. For example, we can find the minimum price of a pound of coffee (function) for all the stores within a 10-mile radius of our house (constraint). Whereas this particular problem may be quite trivial, optimization routines are typically employed to find values for more complex problems, such as might be applied to human movement. Researchers have tested various cost functions that make sense heuristically and have shown that optimization of these costs reproduces many invariant characteristics observed in human motion. For example, Flash and Hogan [9] tested the idea that the smoothness of hand trajectories might reflect an important cost in the planning of reaching movements and proposed a model that minimized the jerkiness of the hand trajectory (mathematically defined as the derivative of acceleration with respect to time). Their simulation predicted straight movements with symmetrical, single-peaked, bellshaped velocity profiles. However, under several experimental conditions, minimum jerk trajectories and experimentally observed hand paths diverged, which led researchers to examine other plausible cost

functions. For example, some researchers speculated that mechanical aspects of movements might reflect important costs for planning movements. Such cost functions have included mean squared torque change [10], peak work [12], or muscle energy [13, 14] among others. These models accounted for some experimental observations that could not be accounted for by optimizations based on kinematic parameters [11]. While minimization of cost functions such as smoothness or torque change accurately predicted average behavior, Wolpert and colleagues [8] also accounted for the small, yet important trial-to-trial variability seen during repetitions of the same task. They proposed that motor commands are corrupted with variability-inducing noise, and in the presence of such noise, the CNS seeks to minimize the variance of the final arm position. This model also predicted many observed invariant characteristics of movements such as trajectory smoothness and the tradeoff between movement accuracy and speed. Two important inferences can be drawn from studies that have attempted to explain movement patterns based on optimization principles: (1) the nature of the costs associated with different tasks are often different and (2) costs such as endpoint variability and mechanical energy do not reflect

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Fig. 2.2 Some “invariant characteristics” of point-topoint movements. The top panel shows fairly straight hand trajectories for multiple movements. The bottom panel shows fairly similar bellshaped velocity profiles for six different movements at different speeds. When timenormalized, the bell-shaped velocity profiles closely overlap. (Adapted from Atkeson C.G. and Hollerbach J.M. Kinematic Features of Unrestrained Vertical Arm Movements. Journal of Neuroscience 1985, Voll. 5, No. 9. pp. 2318–2330. Copyright 1985 Society for Neuroscience)

variables that we tend to have conscious awareness of, yet they appear to be accounted for during the process of motor planning. In other words, the planning of movements not only entails explicit performance criteria that are associated with successful task performance, such as getting hold of a cup of coffee, but also entails implicit criteria that we don’t consciously consider, such as making energetically efficient and reliable movements. An important aspect of the models discussed above is that optimization of a single cost function yields a desired trajectory that is then simply executed in an open-loop manner, once it is

planned. The role of sensory feedback mechanisms in these models is simply to correct deviations from the planned or desired trajectory, regardless of whether these deviations resist or assist in task completion. Thus, the output of feedback circuits is not incorporated in the optimization phase. More recently, the idea that the determination of an optimal “control policy” incorporates knowledge about the “state” of the body and the environment, as relayed by feedback circuits and mechanisms that predict sensory consequences of motor commands, has gained prominence. According to this idea, the optimal solution is the best possible

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Movement Neuroscience Foundations of Neurorehabilitation

transformation from the current state to the motor commands that aid in achieving the task goal [15]. Not too surprisingly then, this optimal feedback control scheme yields task-specific cost functions that often represent a hybrid mix of explicit task-level variables that relate to performance goals, such as movement precision, as well as implicit mechanically related costs that correspond to muscle force or effort. For example, in a task that examined corrections to target displacements that occurred late in the movement, Liu and Todorov [16] showed that subject performance could be best described using a composite cost function that optimized for movement duration, accuracy, endpoint stability, and energy consumption. More importantly, subjects implicitly changed the relative contribution of these costs as the accuracy and stability requirements of the task were changed. Thus, rather than adopt a fixed policy across task conditions, subjects were able to flexibly adapt their control strategy in order to ensure maximum task success. These ideas of flexible control strategies and hybrid cost functions that include task-related and intrinsic biomechanical variables have important implications for designing therapy regimes. Implications for Rehabilitation. It is important to recognize that damage to the CNS from stroke and the associated secondary changes in the musculoskeletal system could induce changes in the set of possible solutions as well as the costs associated with any given task. Therefore, patients may arrive at solutions to a motor task that may not look “normal”, but may be “optimal” given physiological and biomechanical pathologies [17] Thus, rather than simply attempting to make movements look more “normal”, it is important to understand the biomechanical costs associated with different tasks. Most importantly, if movements of the hemiparetic arm elicit energetic costs that are substantially higher than those of the ipsilesional arm, it is very unlikely that the hemiparetic arm use will be spontaneously integrated into activities of daily living. As the technologies discussed in this volume become available in the clinic, assessment of biomechanical variables, such as

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joint power will also become available. While most clinical assessments of function include either the ability to perform certain ADL tasks (Functional Independence Measure—FIM [18, 19], or the ability to perform simulated ADL tasks in particular times (Jebsen-Taylor Hand Function Test—[20] we suggest that direct analysis of biomechanical costs may provide an important supplement to these tests, as an indicator of energetic efficiency. It may also be important to assess one’s subjective sense of effort, which does not always accurately reflect measures of biomechanical cost [21] This should provide a valuable addition to therapeutic assessment because even when ADLs are completed independently if they are not performed within reasonable energetic costs, one might expect minimal carry over into the patient’s spontaneous behavior. It should be stressed that the role of task-level costs is also important for determining optimal control strategies for a given task. Such costs might include the accuracy and duration of movements. Computer-based technologies allow therapists to modify feedback to stress particular performance criteria, so as to emphasize certain costs. For example, in a targeted reaching task, one could provide reward based on duration, when focusing on improving movement time. However, if movement direction and straightness need to be stressed, visual feedback can be modified to amplify errors perpendicular to the desired trajectory while reducing errors in the direction of the desired movement. Such changes would penalize deviations from the desired movement path while allowing errors in the direction of movement. This approach would assign different costs to errors that contribute to task success versus those that don’t. In fact, Ballester et al. [22] recently reported exactly this manipulation, using a virtual reality environment to train reaching hemiparetic stroke patients. The movements of a virtual representation of the patients' paretic limb were amplified in only the dimension parallel to the target direction. Following virtual reality training, the authors reported that the probability of using the paretic limb during a subsequent real-world task was

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increased by the reinforcing experience of seeing the virtual limb reach the target during training. These types of capabilities are now becoming available in the clinic, due to the increasing availability of computer-based robotic and virtual reality technologies.

2.3

Principle 2: Impedance Control

Optimal feedback control theory emphasizes that the derivation of the optimal control signal incorporates knowledge about the state of the body and the environment. If the state changes unexpectedly due to an external perturbation, or random noise, what should its influence be on the control strategy? For example, when a passenger in a vehicle drinks a cup of coffee, what should the control system do when the movement of the cup is unexpectedly perturbed by a bump in the road? Ideally, the components of the perturbing forces that assist in bringing the cup to the mouth smoothly should not be impeded. However, the components of the forces that resist in the achievement of the task goal, such as accelerating the cup too rapidly, or in the outward or downward directions, should be compensated. According to the principle of minimal intervention proposed by optimal feedback control, the central nervous system “intervenes” only when errors are detrimental to goal achievement. Such a selective compensation of errors might explain why people allow slight variability in their performance as long as the overall goals of the task are satisfied. This type of selective modulation of feedback gains is consistent with evidence that even the simplest feedback circuits, reflexes, can be modulated based on task demands. The stretch reflex represents the simplest and most ubiquitous feedback circuit in the mammalian system. The typical response to a stretch of a muscle includes a characteristic three-phase response [23, 24], measured in the electromyogram (EMG) as shown in Fig. 2.3: the shortest latency response often referred to as M1 occurs within some 20–50 ms following perturbation onset, and reflects circuitry contained within the spinal

cord. Following this, a medium latency response, M2, is observed some 60–80 ms following the perturbation onset and is thought to reflect longer latency spinal as well as transcortical circuits. This is followed by M3, a longer latency reaction that is thought to reflect a voluntary corrective process. Studies examining how these responses are modulated have shown differential effects of different task conditions on the early and later phases of the reflex. Early studies in which subjects were instructed to resist or to not resist a perturbation showed that M1 was not modified by such commands, while M2 could be greatly attenuated by the instruction to not resist, and M3 could actually be completely eliminated by this instruction [25]. More recent studies have shown that M2 can be modulated by spatial conditions in a task, such as when subjects are told to allow their hand to displace toward a particular target: when the arm is pushed toward the target, the later phases (M2, M3) of the stretch reflex that resist the perturbation are reduced. However, when the arm is pushed away from the target, the gains of these responses are increased. More importantly, this modulation varies with both the direction and the distance of the target [18]. This demonstrates that feedback circuits such as reflexes can be modulated in accord with task goals through implicit mechanisms. In fact modulation of reflexes appears to be a fundamental mechanism that our nervous system employs to control limb impedance and thus resist perturbations. An elegant example of such reflex modulation was provided by Lacquaniti and colleagues for a ball-catching task [26]. This study demonstrated not only modulation but also reversal of the stretch reflex, in response to ball impact. Both the amplitude and expression of the stretch reflex were modulated in a systematic way as the ball dropped toward the hand. The result of this reflex modulation was to generate impedance to the forces imposed by ball impact, thereby generating a smooth and effective catching response. Why is active impedance control through reflex modulation important for motor performance? During everyday tasks, many environmental perturbations cannot be predicted prior to

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Fig. 2.3 Typical reflex response to muscle stretch. An example of the wrist extensor being stretched using a motor is shown on the left. The right panel shows the typical components of the electromyographic response to

muscle stretch: the short-latency component M1 and the longer latency components M2 and M3. (Adapted from Matthews PB. The human stretch reflex and the motor cortex. Trends Neurosci. 1991 Mar;14(3):87–91.)

movement. In the example of a passenger drinking coffee in a moving vehicle, changes in vehicle acceleration due to bumps and breaking can rarely be anticipated. One can increase overall arm stiffness by co-activating muscles, but this uses a great deal of metabolic energy and interferes with the ability to bring the cup to the mouth. Franklin [27] and colleagues directly tested how subjects might selectively modify impedance without interfering with the coordination of the intended movement. In this study, subjects performed reaching movements with the arm attached to a robotic manipulandum that imposed unstable force fields that had components directed perpendicular to the required movement (see Fig. 2.4a). With practice, the participants were able to adapt to the novel dynamics and produce straight trajectories. They achieved this adaptation by selectively increasing stiffness in the direction of the instability, but not along the movement direction (see Fig. 2.4b). Remarkably, at the joint level, this impedance modification was achieved without changing baseline force and torque profiles (see Fig. 2.4c): the coordination strategy remained kinetically efficient, even though subjects were also able to effectively impede the imposed perturbations. These authors concluded that the nervous system

is able to simultaneously maintain stability through impedance control and coordinate movements in a manner consistent with optimized energy expenditure. We recently showed that such selective modification of limb impedance occurs through continuous modulation of short- and long-latency reflexes [28]. In our study, participants reached a visual target that occasionally jumped to a new location during movement initiation, thus changing the task goal during the course of motion. Unpredictable mechanical perturbations were occasionally applied, 100 ms after the target jump. Our results showed that reflex responses were tuned to the direction of the target jump: response amplitudes were increased or decreased depending on whether the perturbation opposed or assisted the achievement of the new task goal, respectively. We also showed that this reflex modulation resulted in changes in limb impedance to the perturbations. However, under conditions in which the movements were not mechanically perturbed, no changes in EMG or joint torque occurred at reflex latency relative to movements made with mechanical perturbations. These findings supported those of Franklin and colleagues by confirming that limb impedance is controlled without interfering with optimal

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Fig. 2.4 Modulation of limb impedance. a. The typical setup and the perturbing force field. The field acts to push the arm perpendicular (along X-axis) to the direction of motion (Y-axis). b. An increase in limb stiffness along the X-, but not Y-axis for all subjects. c. Shoulder and elbow

joint stiffnesses were independent of the respective joint torques. Adapted from Franklin DW, So U, Kawato M, Milner TE. Impedance control balances stability with metabolically costly muscle activation. J Neurophysiol. 2004 Nov;92(5):3097–105

coordination, by selectively modulating the expression of short- and long-latency reflex responses. The studies discussed above point to the remarkable ability of the nervous system to determine optimal responses to unpredictable situations. Such control policies appear to mediate the modulation of limb impedance through the regulation of feedback circuits such as reflexes to ensure that unexpected

perturbations are countered in a task-specific manner. Reflexive resistance to a perturbation is increased when it is inconsistent with the task goal but decreased when the perturbation is congruent with the goal of the task. These findings agree with the “minimum intervention principle” within the optimal feedback control framework. Thus, controlling limb impedance in a task-specific manner appears to be an integral component of the motor control process.

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Implications for Rehabilitation. The research summarized above indicates that the central nervous system invokes at least two aspects of control to achieve coordinated movements. First, the commands are specified that result in optimal coordination patterns that satisfy both costs associated with task performance and energetic costs. In addition, the nervous system appears to set control policies that modulate sensorimotor circuits such as reflexes, to account for perturbations from unexpected changes in environmental or internal conditions. The importance of recognizing both of these features of control in clinical environments is fundamentally important because brain damage due to stroke can have differential effects on these two aspects of coordination. For example, Beer et al. [29] showed that hemiparesis disrupts optimal intersegmental coordination, resulting in inefficient coordination that fails to account for the dynamic interactions between the segments. This deficit does not appear to depend on the extent of hemiparesis. Traditional therapeutic strategies, as well as more recent robot-aided rehabilitation strategies, tend to target the optimal control process by practicing fairly consistent patterns of coordination, and reinforcing task success. While this type of practice is critical for improving coordination and voluntary control, focusing on repetitive movements under consistent environmental conditions should only be the first step in rehabilitation training. In itself, this training may improve voluntary control of optimal coordination patterns, but is unlikely to train impedance control mechanisms. Because of this, patients may become adept at the training protocols, but show limited transfer to activities of daily living. We suggest that as patients improve their movement patterns under predictable conditions, training protocols should progressively incorporate unpredictable conditions. Such conditions might include random changes in target positions and varying force perturbations, thereby training patients to impede variations in environmental conditions that interfere with task performance.

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Principle 3: Motor Lateralization

As discussed thus far, both optimal control and impedance control are component mechanisms underlying the control of voluntary movements. Our recent work has suggested that these two mechanisms are lateralized to the left and right brain hemispheres, respectively. The seminal research of Sperry and Gazzaniga [30] on disconnection syndrome in split-brain patients first established neural lateralization as a fundamental principle of the cerebral organization. Gazzaniga proposed that distributing different neural processes across the hemispheres was a natural consequence of developing complex functions during the course of evolution. His research provided elegant support for this view of cerebral lateralization as a neural optimization process. Interestingly, early research on hemispheric lateralization was largely limited to cognitive and perceptual processes, with little attention to the motor systems. We introduced the dynamic dominance hypothesis of motor lateralization [31], based on left- and right-arm advantages in reaching performance in healthy adults, and expanded this hypothesis based on computational modeling studies [32, 33] and studies in patients with unilateral brain lesions [34–39]. The dynamic dominance model proposes that the left hemisphere, in right-handers, is specialized for predictive processes that specify smooth and efficient movement trajectories under mechanically stable environmental circumstances, while the right hemisphere is specialized for impedance control mechanisms that confer robustness to movements performed under unpredictable and mechanically unstable environmental conditions. In fact, this type of division of labor between the two sides of the brain appears to predate humans by half a billion years [40]. Rogers and colleagues have proposed a single organizing principle that might account for the large array of emotional, language, perceptual, and cognitive asymmetries that have been described across the evolutionary spectrum of vertebrates. While the left hemisphere appears “specialized for control

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of well-established patterns of behavior, under ordinary and familiar circumstances”, the right hemisphere is specialized for “detecting and responding to unexpected stimuli in the environment” [41]. The dynamic dominance model provides the movement analog to Roger’s model, and thus places handedness in the context of a larger array of neurobehavioral asymmetries across the animal kingdom [42]. An important feature of these models is that both hemispheres are recruited for their complimentary contributions to integrated functional activities. Thus, during the movement of a single arm, both hemispheres contribute their specific aspects of control.[43]. Because each hemisphere contributes specialized processes to control each arm, unilateral brain damage actually produces hemisphere-specific movement deficits in the non-paretic, ipsilesional arm, as well as the contralesional arm. Remarkably, this is the arm that is usually considered unaffected by unilateral brain damage. The idea that each hemisphere contributes to motor coordination of both arms is an important implication of ipsilesional, nonparetic arm motor deficits. While the role of contralateral motor areas in controlling limb movements is well-understood [44] the role of the ipsilateral hemisphere has more recently been implicated by the robust occurrence of ipsilesional motor deficits in both animal models of unilateral brain damage [45–47] as well as human stroke survivors [34, 36, 39, 48–58]. In addition, both electrophysiological and neural imaging studies have shown that unilateral arm and hand movements recruit motor-related areas in both cerebral hemispheres [43, 59–61]. Thus, it is the loss of the contributions of the ipsilateral hemisphere to movement control that gives rise to motor deficits in the non-paretic arm of stroke patients. Most importantly, these deficits can substantially limit functional performance [51, 54], a particularly concerning phenomenon, given that patients with severe contralesional paresis depend on the ipsilesional arm for the majority of their activities of daily living. Our recent studies have examined the specific nature of the ipsilesional movement deficits that result from left or right brain damage, shedding

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light on motor lateralization. These studies have confirmed that right and left sensorimotor strokes produce predictable deficits in impedance control or optimal control, respectively [51, 62]. For example, Schaefer et al. [51] compared reaching movements in the ipsilesional arm of hemisphere-damaged patients with those of healthy control subjects matched for age and other demographic factors. Subjects performed targeted reaching movements in different directions within a workspace to the same side of the midline as their reaching arm. The left hemisphere-damaged group showed deficits in controlling the arm’s trajectory due to impaired interjoint coordination but showed no deficits in achieving accurate final positions. In contrast, the right hemisphere-damaged group showed deficits in final position accuracy but not in interjoint coordination. These findings are exemplified in the hand paths shown in Fig. 2.5a. While control subjects made relatively straight and accurate movements, patients with left hemisphere damage made movements that were very curved, but nevertheless were accurate in the final position. In contrast, patients with right hemisphere damage made straight movements with poor final position accuracies. This double dissociation between the type of error (trajectory or final position) and the side of hemisphere damage (right or left) is emphasized in Fig. 2.5b, which shows the variance in hand positions during the initial trajectory phase (cross), or the final position phase (circle) of the movement. The ratio of errors at these two points in movement (peak velocity, movement termination) is quantified across subjects in the bar graphs, revealing that RHD patients had the greatest variance in final position, while LHD patients had the greatest variance in trajectory. Thus, these results indicate the distinct lateralization of optimal trajectory control and impedance-mediated final position control to the left and right hemispheres, respectively. It should be emphasized that these errors were associated with functional impairments in the ipsilesional arm, as measured by the Jebsen–Taylor Hand Function Test (JHFT). Thus, motor lateralization leads to deficits that depend on the side of the stroke and can lead to

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Fig. 2.5 Lateralization of motor deficits after stroke. 6A shows typical hand paths for healthy control subjects performing with their right or left arm (top panel) and left and right hemisphere-damaged stroke patients performing with their ipsilesional arm (bottom panel). 6B shows hand locations at peak velocity (crosses) and movement end (circles) for a typical left and right hemisphere-damaged stroke patient (top panel). Ellipses represent 95%

confidence intervals. The bottom panel shows the mean ratio of variable errors at peak velocity to variable error at movement end across all subjects for the control and stroke groups. (Adapted from Schaefer SY, Haaland KY, Sainburg RL. Hemispheric specialization and functional impact of ipsilesional deficits in movement coordination and accuracy. Neuropsychologia. 2009 Nov;47 (13):2953–66.)

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and included 18 participants each. The left hemispheredamaged group comprised 22 stroke survivors, whereas the right hemisphere-damaged group comprised 29 stroke survivors. On the X-axis, these groups are stratified by severity of contralesional arm paresis

significant deficits, as tested with clinical assessments, such as the JHFT. Figure 2.6 shows data from 72 age and gender-matched control subjects, 22 left hemisphere-damaged stroke survivors, and 29 right hemisphere-damaged stroke survivors. The Y-axis represents the JHFT score, taken as a percentage of the right dominant arm function in our control group. Thus, 100% is the mean for the right hand of 36 of the control subjects (those who used their right hand). The JHFT is a rather thorough assessment of unilateral arm function that includes a large range of tasks that elicit the coordination requirements of functional daily activities, such as writing, turning pages, placing large and small objects on a table, stacking checkers, and feeding. The left column (control) shows the difference between healthy subjects performing with the left arm and right arm. The data are stratified on the X-axis by both hands (right/left: in the case of stroke survivors this is only the ipsilesional arm) and severity of contralesional paresis, as measured by the upper limb component of the Fugl–Meyer [63] assessment of motor impairment (mild >= 55, moderate > 35, Severe 6 months post-injury), severe SCI despite intensive rehabilitation. However, a recent wave of clinical studies has reported unprecedented outcomes in people with both incomplete and complete SCI, independently demonstrating the long-term recovery of voluntary motor function in the chronic stage after SCI. These studies utilized a combination of intensive rehabilitation and electrical spinal cord stimulation (SCS), which was delivered via epidural multi-electrode arrays implanted between the vertebral bone and the dura mater of the lumbosacral spinal cord. SCS has a long history of applications in motor control, which started soon after its first applications as interventional studies in pain management. To date, SCS has been applied in thousands of individuals with neuromotor disorders ranging from multiple sclerosis to SCI. However, even though the motor-enabling effects of SCS were first observed about half a century ago, the lack of a coherent conceptual framework to

© The Author(s) 2022, corrected publication 2023 D. J. Reinkensmeyer et al. (eds.), Neurorehabilitation Technology, https://doi.org/10.1007/978-3-031-08995-4_18

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interpret and expand these clinical findings hindered the evolution of this technology into a clinical therapy. More importantly, it led to substantial variability in the clinical reports ranging from anecdotal to subjective descriptions of motor improvements, without standardized methods and rigorous statistical analyses. For several decades, these limitations clouded the potential of SCS to promote long-term recovery in individuals with SCI. In this chapter, we present the historical background for the development of SCS to treat motor disorders and its evolution toward current applications for neurorehabilitation in individuals with SCI (Sect. 18.1). We then provide an overview of the conjectured mechanisms of action (Sect. 18.2), and how this collective knowledge has been used to develop SCS into a promising approach to treat motor paralysis after SCI, ranging from tonic stimulation to more sophisticated spatiotemporal protocols (Sect. 18.3). Finally, we open up this review to the recent development of non-invasive methods to deliver SCS, namely transcutaneous SCS, and its comparison with epidural SCS in terms of functional effects and underlying mechanisms (Sect. 18.4). Keywords

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Spinal cord injury Spinal cord stimulation Neuromodulation Locomotion Motor control Rehabilitation

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The Origins of SCS: From Pain to Motor Control

18.1.1 The Rise of SCS for Chronic Pain Management SCS was originally developed for the treatment of chronic, intractable pain based on neurophysiological studies suggesting the possibility to inhibit pain fiber input in the spinal cord by stimulation of the larger-diameter sensory fibers [1, 2]. The spinal cord dorsal columns, which contain the longitudinal ascending continuations

of cutaneous fibers from many spinal cord segments, appeared as an optimal target for inhibiting pain, in particular because of their easy surgical access from the dorsal aspect of the spinal cord. Subdural electrical stimulation of these structures proved successful in managing pain in cats [3], followed by the first human application of SCS in a patient with cancer [4]. The evolution from subdural to less invasive epidural electrodes, and the development of fully implantable commercial systems led to the FDA approval of epidural SCS for chronic pain management in 1989. Today, SCS accounts for about 70% of all neuromodulation treatments [5].

18.1.2 First Evidence of Improved Motor Function During SCS: From Multiple Sclerosis to SCI The ability of SCS to restore function after neuromotor disorders was first unexpectedly observed in 1973 when a subject with motor deficits as a consequence of multiple sclerosis (MS) received SCS therapy to treat intractable pain. Investigators unexpectedly observed that the subject regained volitional control of nearly normal strength of the lower limbs, facilitation of sitting, standing, and ambulation when SCS was active [6]. Four additional implanted MS participants without pain reported a feeling of lightness of the legs during movement, increased endurance during ambulation, as well as the recovery of some voluntary motor function and improved bladder control. Cook improved the technique of electrode placement in the epidural space and implanted more than 200 additional patients with MS within the next few years [7, 8]. Following this first example of SCS-mediated motor improvements, Illis and colleagues introduced the use of SCS for motor disorders to Europe by replicating Cook’s methods. Their first study demonstrated marked improvements in motor, sensory, and bladder function in two participants with MS receiving SCS [9]. The first participant, who had signs of an upper motor neuron lesion and presence of spasticity, regained the ability to walk independently

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with SCS within the first 24 h of continuous stimulation, while the second participant had marked improvements in sensory function. Follow-up large-cohort studies by the same group revealed high variability in outcomes that reduced the initial enthusiasm for SCS. Indeed, although motor and sensory improvements surpassed those ever achieved by any other method at the time, it was only 5 out of 19 individuals that experienced these types of improvements [10]. A subsequent study on 90 participants confirmed that these striking improvements in motor function were infrequent and observed only in a few individuals [11]. An objective effect was only seen on bladder dysfunction and limb spasticity. Similarly, studies by other groups showed remarkable effects of SCS on altered motor function after MS, but with variability across subjects. Siegfried and colleagues tested 111 patients with MS with temporary SCS and considered only about 33% of them as responders [12]. Davis and colleagues also reported on 69 MS patients with full implantations [13]. 64% showed improvements in gait, endurance, and muscle strength. Among these patients, nine who were wheelchair users could walk again with SCS. In spite of the variable outcomes of these initial studies, the clinical importance of the potential effects led to a cascade of studies assessing the off-label use of SCS in subjects with a wide variety of motor disorders, including amyotrophic lateral sclerosis, cerebral palsy, traumatic brain injury, dystonia, torticollis, and a few individuals with SCI. Overall, these studies reported improvements in strength, balance, walking, coordination, speech, swallowing, eye movements, and bladder function [12–16]. Initial investigations on SCI focused primarily on the management of spasticity, but also explored secondary effects on autonomic function including bowel control and sexual function, as well as motor capacity [17–21]. More specifically, Waltz and colleagues observed improved motor control in 65% of 303 participants with SCI who received SCS of the cervical region [22]. These improvements in motor control were initially ascribed to a reduction in spasticity enabled by SCS [17]. However, further studies such as the one by Barolat in 1986 revealed that voluntary

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control of paralyzed muscles was strictly dependent on SCS and stopped immediately when SCS was turned off, independently of changes in spasticity [23] (Fig. 18.1). This effect is discussed in more details later (Sect. 18.3.1.1). In his observations after treating 1336 individuals with a broad range of disorders over a quarter of a century Waltz reported marked or moderate improvements in a majority of them (Table 18.1) [24]. However, despite the initial unprecedented improvements in motor function observed through SCS, the lack of a conceptual framework kept these observations anecdotal. There was no identified physiological criterion to predict responders to SCS, and a lack of agreement on the optimal electrode implantation site, which ranged from high-cervical (C2) to lowthoracic (T10) vertebral levels. These two factors significantly contributed toward the observed inconsistency in clinical outcomes [12, 13, 25], and led to a declining interest in SCS for motor disorders during the 1990s [26].

18.1.3 Standardization of SCS Location for Leg Motor Control Following Motor Disorders A critical contribution to current applications of SCS came from Dimitrijevic and colleagues, who observed that the optimal electrode placement to control leg spasticity was at the T11-L1 vertebral levels, which corresponds to the site of innervation of leg motoneurons in the lumbosacral spinal cord [19]. Indeed, tonic SCS, delivered with higher stimulation amplitudes and lower frequencies than those used for spasticity, could elicit rhythmic, step-like activity in paralyzed muscles of six subjects with chronic, complete SCI [27]. The multi-segmental muscle activity showing consistent rhythmicity and clear alternation between antagonistic muscles was interpreted as the most direct evidence at the time for the existence of central pattern generators (CPGs) in the human spinal cord [27, 28]. These important observations in humans were further supported by investigations using

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Fig. 18.1 First discovery that SCS enables voluntary movement after SCI in 1986. Pictures of one individual with SCI and with some residual motor control over the left toes and ankle, during an attempt to perform a voluntary left knee extension in the absence (left panel) or presence (right panel) of SCS. The knee angle and EMG activity of quadriceps and hamstrings muscles are represented. During SCS, EMG signals from the hamstrings are affected by stimulation artifacts. Red rectangles indicate the 3 repetitions of the flexion–extension movement, highlighting the EMG activity in the quadriceps, which is facilitated by SCS but voluntarily triggered. Adapted with permission from [23]. All rights reserved Table 18.1 Improvements in 1336 individuals treated with SCS (from [24]). All rights reserved

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simulations of the biophysical interactions between electrical fields and neural structures, which led to the understanding that SCS recruits

primarily proprioceptive afferents [29–31]. Proprioceptive afferents then convey mono and poly-synaptic excitatory potentials to spinal

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motoneurons. To maximize motoneuron activation, SCS electrodes must be therefore placed in positions that favor the direct activation of posterior spinal roots that carry these afferents [32– 34]. In the case of lower limb movements, these correspond exactly to the optimal location observed by Dimitrijevic. Currently, all scientific investigations that aim at improving lower extremity function via SCS use the same level for epidural electrode placement: below the injury and over the posterior aspect of the lumbar and upper sacral spinal cord, regardless of the level of SCI [34–39].

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nerve fiber models utilizing the Hodgkin-Huxley formalism enable to calculate membrane potentials in response to external currents [44]. Such models were used to estimate the sites of maximum depolarization and activation thresholds for individual neurons or their substructures. The electrically activated neurons that lead to motor effects were also investigated by neurophysiological studies in rats and individuals with SCI [28, 29, 38, 45–47]. The physiological effects that are caused by electrically activated neurons into their post-synaptic targets involve local and potentially suprasegmental circuits. The underlying mechanisms are not completely clear and still a topic of active research.

Potential Mechanisms of SCS for Motor Control

18.2.1 General Principles

18.2.2 Electrically Activated Neural Structures

The generation and modulation of spinal motor activity by SCS results from biophysical phenomena that occur upon the application of electrical fields to the spinal cord, which in turn causes physiological effects in the spinal circuitry. Epidural electrodes are positioned in direct contact with the dura mater. During each stimulation pulse, most of the generated ionic currents flow between the active contacts through the dural sac that contains the spinal cord and roots, owing to the relatively high electrical conductivity of the cerebrospinal fluid [29–31, 40–42]. Depending on the current flow and the relative localization and orientation of the neural structures within the dural sac, specific subsets of neural structures are depolarized to a level where action potentials are generated according to an all-or-nothing principle. These immediate electrical effects of the stimulation were studied using computer models that are able to solve established physics in arbitrary complex geometrical structures, such as the spinal cord, to calculate the electric current flow and electric potential distribution [29, 31, 40–43]. This method allows to calculate currents and voltages but does not provide per se an estimation of the response of electrically active neural tissue to an electrical field applied externally. To this end,

The neuronal substructure that is the most excited by external electrical stimulation is the axon. In particular, myelinated and large-diameter axons have the lowest thresholds to stimulation [48]. The axonal depolarization is proportional to the second-order spatial derivative of the electric potential along the axonal path [49, 50]. The stimulation-induced currents within the dural sac primarily flow within the relatively well conductive cerebrospinal fluid, in which the spinal roots are bathed, with maximum current densities in the vicinity of the active electrodes and rapid attenuation when entering the spinal cord [29, 31, 42]. Because of their small size, the lack of myelin sheath, and the fact that the induced current poorly enters into the spinal cord, direct electrical activation of spinal interneurons is highly unlikely to happen with SCS [29]. On the other hand, strong depolarizations are caused in the sensory axons of the longitudinally running lumbar and upper sacral posterior roots of passage that are the closest to the active cathode location [31]. Additional low-threshold sites are created along the sensory axons at the posterior rootlet-spinal cord interface, owing to anatomical inhomogeneities, i.e., electrical conductivity boundaries and changes of the orientation of the axon paths with respect to the electric field [31,

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51]. Based on their fiber diameters, electrical activation is likely limited to a subset of posterior root fibers ranging from group Ia muscle spindle afferents to group II afferents, both proprioceptors and cutaneous mechanoreceptors [42]. Upon entering the spinal cord through their posterior rootlets, the sensory axons bifurcate to rostral and caudal projections in the posterior columns. Collaterals connect to locally confined spinal neurons or to relay neurons with ascending axons. There are anatomical differences in the rostral projections between the Ia muscle spindle afferents and the cutaneous afferents that are relevant for their recruitment by SCS [38, 40, 52]. While the rostral projections of the cutaneous afferents can ascend the entire length of the posterior columns, those of the Ia muscle spindle afferents from the lower extremity largely terminate in the upper lumbar and lower thoracic segments to synapse with the relay neurons of Clarke’s column, and in doing so, they occupy deep positions in the white matter [53–55]. Therefore, electrical activation of neural structures within the posterior columns—that is limited to its most superficial layers—is essentially restricted to a small population of ascending projections of cutaneous afferents [56]. On the other hand, a large proportion of the total number of the large-diameter proprioceptive afferents can be in theory recruited in the posterior roots or rootlets [31], especially in the lumbosacral and cervical enlargements where the roots cover the entire cord with rootlet projections [31, 40, 42]. All neurophysiological studies in humans and animals to date support the notion that the motor effects evoked by SCS are triggered by posterior root fiber stimulation. Responses tested by paired pulses applied epidurally with a step-wise decrease of interstimulus intervals clearly demonstrated post-stimulation depression [38, 47, 57], a hallmark of monosynaptic reflexes evoked in proprioceptive afferents [58–60]. Complete suppression of responses with repetitive stimulation also rules out the direct electrical activation of motor axons in the anterior roots [38, 40, 49, 61]. Finally, the order of recruitment of lower- and upper extremity muscles with different segmental innervation, either from

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different segmental cathode positions, with different cathode–anode combinations, or with graded stimulation amplitudes, can be entirely explained by stimulation of the respective posterior roots [34, 38, 40]. Typically, moving the stimulating cathode in the rostral direction away from the segmental posterior root entries increases the response threshold for muscles innervated by motoneuron pools located in these segments, in accordance with earlier predictions by computational models [29–31, 40]. For instance, S1innervated lower leg muscles cannot be recruited by clinically applicable stimulation amplitudes with a cathode located medially over the upper lumbar segments [38]. This can only be explained if the effects of the stimulation are mediated by the posterior roots and not by the longitudinal sensory fiber projections within the posterior columns. Overall, the notion that SCS can be configured to selectively activate proprioceptive afferents of subsets of posterior roots can be used to direct SCS effects toward specific motoneurons to define more targeted stimulation strategies [32–34, 39, 62] (Fig. 18.2).

18.2.3 Evidence for Post-synaptic Activation of Neural Circuits The directly activated proprioceptive afferents make connections with multiple classes of interneurons and motoneurons in the spinal cord. Therefore, the direct recruitment of proprioceptive afferents will generate synchronized volleys of excitatory post-synaptic potentials into all the neurons directly connected by these fibers. This means that the effects of SCS can occur in multiple neural circuits, having effects that can propagate to both spinal and supraspinal structures. A major goal of current research in SCS is to elucidate the contribution of these activated circuits and their relevance to the effects observed in human clinical trials. In this section, we will discuss only a few of the most studied circuits underlying the effects of SCS, while acknowledging that many more may contribute to the restoration of voluntary motor control.

Spinal Cord Stimulation to Enable Leg Motor … b Epidural multielectrode array over lumbar spinal cord

a Anatomy of the lumbosacral spinal cord Spine Spinal Segmental model cord model innervation model (cm)

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Fig. 18.2 Segmental recruitment of posterior roots and innervated muscles during SCS. a Anatomical representation of the spine, (defined by vertebral body and intervertebral disc heights), lumbosacral spinal cord (aligned with the spine), and segmental innervation probabilities (0%–100%) of lower extremity muscles, reflected by the opacity of their respective colors. These representations correspond to an average over thousands of individuals published in the literature. Adapted with permission from [38]. (Figure published under a Creative Commons license: Fig. 18.2a, licensed under CC-BY 4.0). b Left: mid-sagittal MRI image of the spine and spinal cord of a study participant with SCI, with indications of the vertebral levels, estimated level of the tip of the spinal cord (conus medullaris, yellow line) and electrode array location in this subject (red line). Middle: reconstruction of the location of the electrode array based on MRI and Computed Tomography (CT) scan. Right: Computational model showing the location of the epidural electrode array, epidural fat (yellow), cerebrospinal fluid (dark blue), gray and white matter (gray), and posterior roots (light blue). c Responses of rectus femoris and tibialis anterior to 2‐Hz SCS with incremental amplitudes show the presence of late EMG components in tibialis anterior, but not in rectus femoris, with EES amplitudes greater than two times the response threshold. Adapted with permission from [66]. (Figure published under a Creative Commons license: Fig. 18.2c, licensed under CC-BY 4.0). d Electrophysiological recordings were used to determine optimal electrodes and amplitudes for targeting specific spinal cord regions. EMG responses when delivering single-pulse EES at increasing amplitudes are shown (gray traces). Motor neuron activation maps correspond to optimal amplitudes (black traces). Circular plots report EMG amplitude (in gray scale) at increasing amplitudes (radial axis). White circles show optimal amplitudes; polygons quantify selectivity at this amplitude. c–d: adapted with permission from [34]. All rights reserved

18.2.3.1 Recruitment of the Monosynaptic Reflex In previous paragraphs, we briefly anticipated that most of the motor effects of SCS can be largely explained by assuming the direct recruitment of large proprioceptive afferents. Indeed, Ia-afferents form strong monosynaptic excitatory connections to spinal motoneurons, especially with those innervating extensor muscles. In consequence, each pulse of SCS will generate strong synchronized volleys of

excitatory post-synaptic potentials on motoneuron membranes, which can either lead to their direct activation (H-reflex-like responses also known as posterior root-muscle reflexes), or to an increase in their membrane excitability depending on how many afferents are recruited. Since the number of recruited afferents is directly proportional to the strength of the electrical fields, one can argue that stimulation amplitude will then determine whether SCS is in “subthreshold” (i.e., modulating motoneuron membrane excitability) or “supra-threshold” (i.e., directly

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inducing action potentials in the motoneurons). Therefore, supra-threshold SCS pulses can induce single, distinct H-reflex-like responses in the muscles with a segmental recruitment order. Stimulation of the more rostral roots will mainly induce responses in rostrally innervated muscles, whereas stimulation of the more caudal roots will elicit responses in the more caudally innervated muscles [29, 33, 34, 38, 62]. This property is remarkably robust across animal species and humans as well as the spinal cord region (cervical vs lumbar). This large body of evidence demonstrates that the monosynaptic reflex is certainly involved in the generation of motor outputs during SCS and likely contributes to the facilitation of voluntary motor control by modulating motoneuron excitability.

18.2.3.2 Recruitment of Excitatory Spinal Circuits The group I and group II afferent fibers activated by SCS have rich synaptic connections to spinal motoneurons, to first-order interneurons of spinal circuits with a pivotal role in the control of locomotion, as well as to supraspinal circuits— given that residual longitudinal connectivity exists following a trauma [60, 63, 64]. Studies so far are limited to local spinal effects that have been indirectly deduced in individuals with SCI, using trains of stimulation with various frequencies (2–100 Hz) and intensities [57]. As introduced above, the most prominent events are stimulus-triggered, short-latency responses that are predominantly monosynaptic in nature— called posterior root-muscle reflexes—and are most evident at lower stimulation frequencies (i.e., 2–10 Hz) [28, 46, 47, 57, 65]. Lowfrequency SCS was also shown to evoke crossed-reflexes in thigh muscle groups [66], posterior root-muscle reflexes with superimposed, delayed-latency electromyographic (EMG) components in flexor muscles (by *8 ms with respect to the monosynaptic response) [66], and complex long-latency responses (>50 ms) [66, 67] with characteristics reminiscent of the late flexion reflex observed in individuals with chronic SCI [60, 66, 68]. These observations hence suggest that SCS can activate

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commissural neurons as well as interneurons specific to flexor-related oligo- and poly-synaptic pathways, i.e., interneuron types shown to be essential components of the mammalian lumbar locomotor circuitry [66, 69]. Indeed, previous studies demonstrated that tonic lumbar SCS at frequencies around 30 Hz applied in individuals with chronic, motor complete SCI examined in the supine position could generate periods (10– 60 s) of rhythmic EMG activities in the paralyzed lower extremity muscles [27, 47]. The generation of rhythmicity with cycle frequencies compatible with slow to fast walking speeds and various patterns of muscle recruitment, including reciprocity between antagonistic muscles [70] by a sustained stimulation with constant parameters, were interpreted as evidence for the activation of a human CPG for locomotion [27, 71]. It should be noted that the generation of these rhythmic EMG activities required a minimum frequency of 22.5 Hz and rather high SCS amplitudes, about three times the posterior root-muscle reflex threshold of the mid-lumbar-innervated quadriceps muscle group [70]. Such stimulation likely recruits a large proportion of group I and II afferent fibers within the lumbar posterior roots. The antidromic action potentials carried along these afferents toward the periphery likely cancel a large part of the naturally generated, orthodromically traveling action potentials from proprioceptors [71, 72]. The generation of rhythmic activity independent from phasic peripheral feedback (largely canceled following antidromic collision) is essential in the demonstration of centrally generated rhythmicity and significant from a neuroscientific point of view. However, the blocking of proprioceptive feedback required for generating adaptive movements limits the applicability of such high stimulation intensities in neurorehabilitation approaches [72, 73].

18.2.3.3 Recruitment of Inhibitory Spinal Circuits Despite their historical use in spinal spasticity [6, 21, 74], the recruitment of inhibitory spinal circuits by SCS has received less attention. A phenomenon observed with low SCS frequencies sheds some light on the activation of inhibitory

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interneurons [49]. Stimulation at around 16 Hz can produce specific modulation patterns of repetitively evoked posterior root-muscle reflexes, with every other response being attenuated, resulting in simple periodic patterns. The rhythmic attenuation of responses without changes in stimulation amplitude and the reciprocal organization of these patterns between antagonistic muscles suggests the recruitment of inhibitory circuits involving Renshaw cells and Ia inhibitory interneurons [49]. Indeed, similarly to monosynaptic connections to the motoneurons, Ia-inhibitory interneurons are known to have direct inputs from Ia-afferents [75, 76]. These interneurons inhibit the motoneurons that innervate antagonistic muscles. Therefore, the potentiation of these inhibitory cells, or even their direct post-synaptic activation by SCS, is likely involved in the alternation of agonist and antagonist muscle activations, which can, in turn, contribute to the production and modulation of coordinated movements.

18.2.4 Lessons from Animal Studies 18.2.4.1 Recruitment of Different Reflex Circuits by SCS in Animal Studies In intact and spinalized rats, single-pulse SCS of the lumbosacral spinal cord evoked composite EMG responses in the hindlimb muscles with a succession of stimulus-triggered, partially overlapping potentials [29, 45, 77]. Three physiologically different components were suggested, and distinguished according to the relative latencies of their major EMG peaks: an early, middle, and late-latency response. The early response is a direct, M-wave-like response [60], elicited by direct electrical stimulation of motor axons within the ventral roots—a response type not evoked in humans with electrodes placed over the midline (see above). The middle response is well documented to be a monosynaptic reflex, likely corresponding to the

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monosynaptic posterior root-muscle reflex in humans [47]. The late response occurs with an additional delay of 4–5 ms with respect to the middle response, distinguishing it as an oligosynaptic spinal reflex [38]. There is the uncertainty of the origin of the late response, but it was suggested to involve group II muscle afferents [29] or the flexor reflex afferents [45, 77].

18.2.4.2 Recruitment of Locomotor Circuits by SCS in Animal Studies Early animal studies employing spinal cord and/or root stimulation were conducted in cats with the aim to investigate the intrinsic capability of the lumbar spinal cord to generate the rhythm and pattern underlying hindlimb locomotion. Stimulation of bilateral pairs of lumbar dorsal roots, typically with 30–50 Hz, induced rhythmic alternating activity in spinal preparations after elimination of phasic inputs from the hindlimbs (through deafferentation or curarization), and hence demonstrated the existence of a CPG [78]. Similar patterns were observed in acutely spinalized cats pretreated with L-DOPA and in chronic spinal cats acutely decerebrated but without administration of drugs. Another classical study demonstrated that subdural and epidural stimulation over the posterior aspect of the lumbar spinal cord could elicit stepping in acutely spinalized cats suspended over a moving treadmill without pharmacological manipulation [79]. Electrodes positioned over the spinal cord midline or laterally over the dorsal root entry zones were both effective. This animal study was perhaps one of the first to suggest the value of SCS for locomotor rehabilitation, as it could be suitable for “a degree of “exercise” and perhaps provide sufficient force to propel the subject as long as postural support is provided” [79]. Later studies in rodents and non-human primates provided further ground for the development of SCS into a tool for enabling locomotion after SCI.

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The Evolution of SCS into a Neuroprosthetic Technology and a Neurorehabilitation Therapy

In the previous sections, we discussed the early history of human applications of SCS in motor disorders, how empirical observations highlighted the potential of SCS to enable voluntary motor control, and its potential mechanisms of action. In this section, we bring all this information together to discuss how the concept of SCS shifted from a neurophysiological tool to activate CPGs to a therapy capable of amplifying residual voluntary motor control [80]. Currently, the application of SCS from the epidural space located over the posterior aspect of the lumbosacral spinal cord represents a state-of-the-art neuromodulation technique for facilitating lower limb motor control after SCI. Here, we report the technological innovations that enabled this transition toward a neuroprosthetic solution for motor impairments and a neurorehabilitation therapy for long-term recovery. In particular, we discuss different approaches for delivering SCS, from tonic stimulation to more sophisticated spatiotemporal stimulation protocols, which can be either applied at a pre-defined pace, triggered by external events, or adjusted in closed loop.

18.3.1 Tonic SCS for the Recovery of Voluntary Motor Control in People with SCI 18.3.1.1 The Initial Discovery that SCS Enables Voluntary Motor Control The very first discovery that SCS may be used as a motor-enabling neuromodulation tool after SCI is attributed to Barolat and colleagues in 1986 [23]. This case study showed that one subject with incomplete SCI regained voluntary motor control in the presence of SCS after several months of stimulation (Fig. 18.1). Specifically,

SCS allowed the subject to perform a full knee extension against gravity, despite the complete absence of voluntary activity in the thigh muscles in the absence of SCS. This effect was present only when the stimulation was turned on, illustrating for the first time an immediate effect of SCS for enabling motor function after SCI. Importantly, this first observation of improved motor control did not involve the triggering of automatic stepping patterns, which was the main focus of animal research in the field. Instead, SCS enabled single-joint voluntary movements, which is still the main clinical outcome of modern clinical trials.

18.3.1.2 SCS as a Tool to Trigger Movement Primitives Several years later, Dimitrijevic and colleagues demonstrated in six subjects with complete SCI that SCS delivered over the lumbar spinal cord (T11-L1 vertebrae) in supine position produced rhythmic EMG responses and flexion–extension leg movements resembling stepping [27]. Stimulation was applied at relatively high amplitudes (5–9 V) and for frequencies between 25 and 60 Hz. Beyond its potential prosthetic implications, this discovery provided indirect evidence for the existence of a CPG in humans, defined as a circuit within the spinal cord capable of generating rhythmic outputs in response to a tonic input (i.e., a continuous pulse train at a given frequency without rhythmic modulations). In addition to the rhythmic movements induced by stimulation of the CPG, other simpler movements could be produced. For example, Jilge and colleagues showed in five subjects that tonic SCS at frequencies between 5 and 15 Hz induced bilateral extension of the lower limbs [81]. Minassian and colleagues next demonstrated in ten subjects that tonic SCS at different frequencies can switch the functional state of spinal circuits between distinct functional units that produce muscle synergies associated with either bilateral extension of the lower limbs (for frequencies of 5–15 Hz) or rhythmic steppinglike movements (for frequencies of 25–50 Hz) [47]. This latter study was the first one to coin the

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term “neuroprostheses” in the context of SCS: “This study opens the possibility for developing neuroprostheses for activation of inherent spinal networks involved in generating functional synergistic movements using a single electrode implanted in a localized and stable region”. However, it is important to highlight that the recruitment of CPGs or other movement primitives should not be taken as a goal to produce automatic, non-voluntary stepping (which may not be relevant as a clinical outcome), but rather as a demonstration that SCS can engage the spinal circuitry that is necessary to produce complex motor patterns such as locomotion. For this reason, subsequent studies investigated the combination of SCS with physical training as a means to improve the capacity of residual descending inputs to regain control over spared spinal circuits.

18.3.1.3 The Combination of SCS with Training and the Re-Discovery that SCS Enables Voluntary Motor Control The first combination of SCS with partial weightbearing locomotor training was performed by Herman and colleagues in a subject with incomplete SCI [82, 83], and later expanded to a second subject [84]. Both participants had incomplete SCI with no independent ambulatory function and were graded as AIS C with low lower extremity motor scores—following the American Spinal Injury Association (ASIA) Impairment Scale (AIS). Treadmill stepping alone improved gait performance on the treadmill but did not improve overground walking capabilities. Instead, the addition of SCS led to the immediate facilitation of walking, as well as further training-related gait improvements and a reduced sense of effort. Despite remarkable improvements in walking capability, there was no change in muscle strength or lower extremity motor scores when the stimulation was turned off, suggesting that this approach was insufficient to trigger neuroplasticity mechanisms in the tested subjects.

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The next milestone in the application of SCS to neurorehabilitation came from a case study by Harkema and colleagues in 2011, which aimed at promoting standing and assisted treadmill stepping in a subject with motor-complete, sensoryincomplete SCI (AIS-B) [85]. SCS enhanced rhythmic EMG activity during assisted treadmill stepping, evoked sustained activation patterns in lower extremity muscles, and allowed independent, full weight-bearing standing after 80 sessions of intensive training. An additional major outcome of the study was the incidental rediscovery of the so-called motor-enabling effect of SCS, initially observed by Barolat in 1986. Indeed, the participant reported that SCS enabled him to perform voluntary movements of paralyzed muscles, including toe extension, ankle dorsiflexion, and leg flexion. As in 1986, this effect was present only when the stimulation was on. The investigators focused on this motorenabling effect in a subsequent study involving intensive training of voluntary leg movements under SCS in four participants with motor complete SCI (two graded AIS-A, and two AIS-B), including the participant from the original study [35]. All new participants were able to voluntarily induce movement when SCS was applied from the very first day and without any training, even for the two sensory and motor complete subjects (Fig. 18.3). With training enhanced by SCS, all participants improved over time and became able to control movements based on visual or auditory cues. Three of them could generate graded levels of force in at least one leg, and two could modulate EMG activity during assisted treadmill stepping and SCS by thinking about moving the legs. Additionally, all participants became able to perform full weight-bearing standing with SCS after 80 sessions of stand training [86]. A followup study found that rehabilitation in the presence of SCS needs to be task-specific, and that stand and step training leads to different functional outcomes [87]. After completion of the study, one of the participants (AIS-B) was enrolled to receive additional activity-based training with

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I. Seáñez et al. b Graded force generation with SCS

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Fig. 18.3 Tonic SCS immediately enables voluntary motor control. a Lower extremity EMG activity during voluntary movement attempts (ankle dorsiflexion) without and with SCS in four individuals with clinically determined motor complete SCI. Electrode representations show cathodes (gray) and anodes (black). Stimulation frequency: 25– 30 Hz. Muscles, surface EMG: intercostal sixth rib (IC), tibialis anterior (TA), soleus (SOL); fine wire EMG: iliopsoas (IL), extensor digitorum longus (EDL), extensor hallucis longus (EHL). Gray highlighted: active ‘flexion/extension’ period. b Left leg force and iliopsoas, vastus lateralis and intercostals EMG activity generated during a low (20%), medium (60%), and high (100%) effort of hip/knee flexion with SCS from patient 3. Gray shading: force duration. c Volitional modulation in EMG activity by an individual with motor complete SCI during manually assisted stepping (40% body weight support, 1.07 m/s) in the presence of SCS. Initial steps show EMG pattern while the subject (patient 3) is not thinking about stepping. Section within the red dashed lines show the period of steps while the subject is consciously thinking about stepping and facilitating each step (with voluntary intent). Adapted with permission from [35]. All rights reserved

SCS, and his voluntary leg motor control progressively improved throughout the 3.7 years of training, to a level such that he could produce voluntary leg movement and standing even with SCS turned off [88]. With the goal to confirm the motor-enabling effect of SCS on participants with severe SCI, an independent group at the Mayo Clinic conducted a 2-week, 8-session study on a participant with chronic, complete SCI (AIS-A) attempting volitional control of leg movements with SCS [89]. Stimulation enabled voluntary knee flexion, initiation and termination of rhythmic leg movements, full weight-bearing standing, and voluntary generation of step-like movements while stationary in an upright position with body-

weight support. These abilities were present only when the stimulation was turned on.

18.3.2 Spatiotemporal SCS for Neuroprosthetics and Neurorehabilitation in Animal Models of SCI In parallel with these investigations utilizing tonic SCS in humans, several studies in animal models of SCI were laying the groundwork for a new stimulation paradigm that would combine the ability to promote voluntary motor control and overt synergistic movements during functional tasks: spatiotemporal SCS. So far, we

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described applications in which SCS was applied tonically, i.e., stimulation parameters such as amplitude, frequency, pulse width, and electrode configurations (choice of anodes and cathodes) were set manually by the experimenter at the beginning of a trial and were kept constant across consecutive steps. In real-life situations, however, the fine control of gait requires the ability to modulate muscle activity and kinematic outputs depending on the environment, task requirements, or levels of fatigue, which motivated the development of spatiotemporal SCS.

18.3.2.1 Spatiotemporal SCS Controlled by Residual Kinematics During SCS-enabled locomotion, this adjustment can be done artificially by linking the task requirements and the observed kinematics with the stimulation parameters used to control SCS in real time. In a first pioneering study, Wenger and colleagues established the first proof-of-concept closed-loop SCS in rat models of SCI [90]. They identified a linear relationship between SCS frequency and the elicited step height, which they exploited as part of a closed-loop proportionalintegral (PI) controller. This controller adjusted SCS frequency in real-time and for each step based on the desired and observed step heights, which depended on the task requirements and environment. For example, climbing a staircase required a higher step height than overground walking on flat surfaces. The observed kinematics were obtained by means of a motion capture system that monitored the 3D position of infrared-reflective markers placed on the joints of the hindlimb. This technological solution shows the possibility to develop intelligent systems that can support an individual in modulating motor output by adjusting SCS parameters in real time. In a subsequent study, the same group introduced a new concept of real-time control of SCS parameters, which selected spatially-distinct sets of electrodes during different phases of the gait cycle. This protocol aimed at independently activating specific muscle synergies such as leg flexion or extension at the appropriate time, a concept termed spatiotemporal SCS [62].

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Specifically, they aimed at replicating the dynamics of motoneuron activation underlying gait in non-injured animals, which was indirectly inferred from the EMG activity of several key hindlimb muscles innervated at various segments within the lumbosacral spinal cord. The obtained spatiotemporal maps of motoneuron activation during bipedal locomotion highlighted two spatially distinct activation patterns associated with major muscle synergies, and with, respectively, flexion and extension movements of the hindlimb, hereafter referred to as “hotspots”. To reactivate these hotspots in animals with SCI, spatially specific electrode configurations were first selected to recruit the dorsal roots projecting to the corresponding spinal segments. Next, kinematic gait events corresponding to the time when the animal places the foot on the ground (“foot strike”) and lifts it off (“foot off”) were extracted in real-time using the motion capture system described previously. Extracting these gait events was possible thanks to the residual kinematics of the animals with incomplete SCI placed in a body-weight support system. In this new paradigm, SCS was not applied tonically to the lumbosacral spinal cord, but as short pulse trains (lasting a few hundreds of milliseconds) triggered by these extracted gait events. Specifically, electrode configurations able to activate the upper lumbar and sacral segments of the spinal cord (associated with hindlimb flexion and extension, respectively) were triggered by the “foot off” and “foot strike” gait events, respectively. In addition, SCS was delivered at an amplitude sufficiently high to elicit motor outputs through the generation of powerful spinal reflexes, which supported the animals in movement execution. This study demonstrated that spatiotemporal SCS allows the facilitation of specific movement phases through targeted activation of the appropriate motoneuron pools through spinal circuits during gait, which can be achieved using epidural multi-electrode arrays. This approach also enables to reduce muscle co-activations observed during tonic SCS, to use higher stimulation amplitudes that can provide better weight-bearing capacity, and to vary stimulation

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frequency throughout the gait cycle depending on the desired functional outcomes.

18.3.2.2 Spatiotemporal SCS Controlled by Brain Signals The principles underlying spatiotemporal SCS were then tested in non-human primates, which represent the most suitable animal model for the translation to humans because of the unique organization of the corticospinal tract in these species [91, 92]. In their study, Capogrosso and colleagues extracted motor signals from intracortical recordings in macaque monkeys with SCI to control spatiotemporal SCS, thereby pioneering the concept of a “brain-spine interface” [33]. Specifically, a 96-channel intracortical microelectrode array (Utah array) was surgically implanted into the hindlimb area of the primary motor cortex, which sends motor commands down to the spinal cord. Thanks to a state-of-the-art wireless neuronal amplifier (able to amplify and broadcast wirelessly 96 channels of neuronal data at a sampling rate of 20 kHz), they recorded the spiking activity across the 96 electrodes while animals performed treadmill and overground locomotion. This neuronal activity was used as an input to a machine learning algorithm (a discrete classifier) able to identify neural states associated with flexion or extension of the hindlimb. At the spinal cord level, the animals were implanted with an epidural multi-electrode array designed specifically to cover the lumbosacral segments of the macaque spinal cord. The spinal electrode array was in turn connected to a modified version of an implantable pulse generator (IPG) clinically approved for deep brain stimulation. Modifications to the IPG firmware provided real-time control over stimulation parameters such as electrode configurations, amplitude, and frequency. This technological framework enabled the implementation of brain-triggered spatiotemporal SCS in freely moving non-human primates. The same principles as used in rats to optimize SCS

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were extended to non-human primates [32], and stimulation protocols facilitating flexion or extension of the leg were extracted. After a first proof-of-concept in intact animals, two macaque monkeys received a unilateral corticospinal tract lesion that left one hindlimb completely paralyzed. As early as a few days after the experimental lesion, the brain-spine interface enabled to reestablish both treadmill and overground locomotion, with the paralyzed hindlimb moving in coordination with the three other limbs. This study brought important advancements, both technologically and scientifically. On the technological side, it demonstrated that it is possible to implement brain-triggered spatiotemporal SCS with currently available technologies that are ready for human use. From a scientific standpoint, it demonstrated that spatiotemporal SCS immediately restored weightbearing locomotion as early as six days postinjury in non-human primates, which bears substantial clinical relevance. The concept of braincontrolled SCS for rehabilitation was further developed by Bonizzato and colleagues who linked cortical ensemble activity to the amplitude of SCS in rat models of SCI and showed a more pronounced and faster recovery of locomotion when training with brain-controlled SCS compared to tonic SCS [93]. In summary, this series of studies laid the technological and scientific premises for the application of spatiotemporal SCS to humans with SCI. First, spatiotemporal SCS provides a way to activate different spinal cord locations and thus different muscle synergies at different phases during the gait cycle or any other motor task. Next, these different stimulation protocols can be triggered based on residual kinematics or neuronal signals to enable a smooth integration into the ongoing locomotor activity. Finally, closedloop control policies can be additionally used to adjust various parameters such as stimulation amplitude or frequency in real-time to adapt to task requirements and environmental constraints.

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18.3.3 Tonic and Spatiotemporal SCS Combined with Intensive Rehabilitation Restore Independent Overground Walking in People with SCI

and in the presence of tonic SCS, this participant was able to stand, step on a treadmill without body-weight support, and walk overground with a walker and assistance of a physiotherapist for hip stability, for the first time in a participant graded AIS-A in the chronic state of SCI. In a recent follow-up study, Gill and colleagues also showed that maximizing participants’ intention to walk and minimizing body-weight support during training with tonic SCS improved independence and decreased the need for external assistance by a physiotherapist (Gill et al. 2020). Conversely to the earlier study by Rejc and colleagues [87], dynamic training combining the repetition of different motor tasks found positive effects on both stand and gait performance simultaneously. Clinical studies by these two groups illustrated the potential of tonic SCS combined with several months of intensive rehabilitation for restoring motor function after SCI.

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The year 2018 marked a milestone for the application of SCS in people with SCI. For the first time, three groups in parallel demonstrated in a total of six subjects with chronic, severe SCI that SCS, delivered with either tonic or spatiotemporal protocols and combined with intensive rehabilitation, could enable independent overground walking [34, 37, 94].

18.3.3.1 Tonic SCS At the Kentucky Spinal Cord Injury Center and the University of Louisville, Angeli and colleagues enrolled four participants with motorcomplete SCI (two AIS-A, two AIS-B) who performed training sessions for standing, treadmill stepping with body-weight support and manual assistance, and overground walking when possible, all in the presence of tonic SCS [94]. All four participants achieved assisted standing and improved trunk stability in the sitting position in the presence of SCS and after several weeks of training. Most importantly, the two participants with motor-complete, sensoryincomplete SCI (AIS-B) achieved the ability to walk overground with tonic SCS after 278 and 81 training sessions respectively, over a period of 85 and 15 weeks. Walking only occurred when SCS was turned on, and while the participant consciously intended to walk. After 147 sessions, the second participant was able to walk independently with a walker and with SCS, which was an unprecedented level of recovery for a person with motor-complete SCI. In parallel, at the Mayo Clinic, Gill and colleagues enrolled an individual with chronic motor- and sensory-complete SCI (AIS-A), who had previously trained to perform step-like movements with SCS in a side-lying position [89] to receive additional motor task training with tonic SCS [37]. After 43 weeks of training

18.3.3.2 Spatiotemporal SCS Meanwhile, at the Lausanne University Hospital and Ecole Polytechnique Fédérale de Lausanne (EPFL) in Switzerland, Wagner and colleagues pioneered the use of spatiotemporal SCS in three participants with chronic, incomplete but severe SCI (one AIS-D and two AIS-C, including one with motor scores of 0 in all key leg muscles but with remaining sphincter control) [34] (Fig. 18.4). They demonstrated both an immediate facilitation of body-weight-supported walking and long-term recovery of motor function even in the absence of SCS. This strategy leveraged the IPG with real-time control capabilities previously tested in non-human primates [33], which was connected to the same 16electrode array as used in the other clinical studies cited above (Specify 5-6-5, Medtronic, clinically approved for the treatment of chronic pain). Taking inspiration from their previous methodology in rodents and non-human primates [32], they developed a stimulation protocol that alternated between the swing, weight acceptance, and propulsion phases of the right and left legs at appropriate times and amplitudes during the gait cycle. Each functionality was associated with a

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stimulation pattern consisting of a spatially specific set of anodes and cathodes optimized to recruit the associated posterior roots, and with a stimulation amplitude and frequency that further maximized the activation of the desired muscle synergy. For example, stimulation frequencies between 40 and 120 Hz tended to better promote a whole-leg flexion synergy (at the hip, knee, and ankle joints simultaneously), while frequencies of 20–30 Hz preferentially recruited functional knee, and ankle extensors. Finally, the alternation of stimulation patterns could either be delivered automatically at a pre-defined sequence and pace, or they could be triggered in real time by residual kinematics for people with sufficient residual control. Movement feedback used to trigger SCS was initially obtained from an infrared-based 3D motion capture system as previously shown in rodents [62], and later by wearable sensors containing inertial measurement units (IMU) placed on the subjects’ feet. Such sensors, along with appropriate algorithms, enabled to extract the foot inclination angle as the participant attempted to initiate movement and to trigger a stimulation pattern that enabled flexion of the corresponding leg. This spatiotemporal SCS paradigm, combined with a cable-based robotic body-weight support system, allowed a wide variety of locomotor tasks both on a treadmill and overground in the three participants with chronic SCI at the cervical level. One participant (AIS-C) had complete motor paralysis on the left leg but residual activity on the right, the second (AIS-D) had paralysis in the leg flexor muscles, and the third (AIS-C, based on the presence of sphincter contraction) had motor-complete paralysis in both legs. Spatiotemporal SCS immediately (i.e., without training) facilitated EMG activity underlying locomotion in otherwise inactive or poorly active leg muscles and enabled participants to walk overground with assistive devices and body-weight support. Participants could voluntarily modulate the effect of the stimulation by exaggerating step elevations, could walk at different speeds and could cover distances of up to 1.2 km on a treadmill without deterioration of kinematics or muscle activity. The first two

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participants regained the ability to transition from sitting to standing and to walk independently with crutches without SCS or body-weight support. Neurological recovery, tested according to clinical standards and without SCS, was observed to different degrees in all three participants. The first participant improved from AIS-C to AIS-D and gained 16 points in his lower extremity motor scores (from 14 to 30, maximum of 50). The second participant gained 11 points (from 25 to 36). The third participant gained four points (0–4). Although he was not able to perform voluntary movements against gravity in the absence of SCS, the researchers observed an increase in the maximum isometric torques that the participant was able to produce in the presence of SCS. Although the investigators did not attempt to train participants with tonic SCS, they performed a comparison of the immediate facilitation of locomotor activity with tonic versus spatiotemporal SCS. In the three reported participants, tonic SCS created an important co-activation of antagonistic muscles preventing smooth locomotion. Additionally, it disrupted the residual proprioceptive inputs to the spinal cord and the brain, thereby blocking important feedback cues to the spinal locomotor circuitry as well as impairing the conscious perception of the lower limbs in space [72]. In a recent study by the same group, Rowald and colleagues expanded their approach to people with motor-complete SCI (two AIS-A, one AIS-B), who were implanted with a new 16electrode array specifically designed for the rehabilitation of both leg and trunk motor function after severe SCI [39]. Their approach also involved the development of personalized computational models of the spinal cord derived from structural and functional magnetic resonance imaging (MRI). This study demonstrated that spatiotemporal SCS immediately enables (i.e., within a week of using SCS) powerful facilitation of walking even in motor-complete participants, whereas similar functional outcomes could only be achieved after several months of intensive training using tonic SCS [37, 94]. Furthermore, it provides a path forward in the refinement of neurotechnologies for delivering SCS, ranging

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Fig. 18.4 Spatiotemporal SCS immediately enables independent walking. a Top: the three muscle synergies underlying human walking, and which can be targeted by SCS. Bottom: typical sequence of spatiotemporal SCS and associated parameters for immediate facilitation of walking after SCI. b Chrono-photography, tibialis anterior EMG activity and foot vertical position during overground walking with body-weight support and walking sticks while SCS is switched on, then off, then on in a subject with severe SCI. c Overground walking when a subject with incomplete SCI but completely paralyzed left leg is asked to perform first normal and then exaggerated step heights. d Consecutive values of step height and EMG activity over 60 min of walking with EES (1 km). BWS: body-weight support. Adapted with permission from [34]. All rights reserved

from new electrodes arrays and personalized computational models of the spinal cord to versatile software platforms for configuration and use of spatiotemporal SCS by non-experts. The future deployment of such technologies in widespread clinical practice will require the additional development of new implantable neurostimulators and automated pipelines for optimizing SCS parameters.

18.3.3.3 Limitations of Locomotor Rehabilitation Facilitated by SCS Improvements in motor function mediated by SCS require high-intensity neurorehabilitation sessions, spread over a time period that is much longer than provided in current clinical practice and covered by insurance. For SCS to become a clinically accepted method for augmenting

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rehabilitation outcomes, the duration of the rehabilitation phase should be therefore considerably shorter. This could be achieved for example by starting neuromodulation therapies in the sub-acute phase after the injury, which would leverage the intrinsic capacity of the spinal cord to reorganize. In individuals living with a chronic SCI, combining SCS with pharmacological interventions will likely further improve and accelerate rehabilitation outcomes [95, 96]. Administration of pharmacological agents would thereby mimic the effects of neurotransmitters such as serotonin and dopamine, which are essential for locomotion. These neurotransmitters are synthesized in the brainstem and the posterior hypothalamus, but their descending axons become partially separated from the lumbar spinal cord after SCI. Finally, all subjects with severe SCI who achieved overground walking with SCS required their arms to maintain balance using either a walker or crutches. This means that their arms cannot easily serve other purposes, such as reaching for an object and carrying it from one spot to another, potentially limiting the usability of this technology in certain daily life situations. To improve dynamic balance, SCS protocols will need to target additional muscle groups involved in hip/trunk movement and stabilization, such as leg abductors and adductors, as well as the quadratus lumborum and the paraspinal muscles. In fact, there is early demonstration that multi-electrode arrays placed over the low-thoracic and lumbosacral spinal segments, combined with activity-specific stimulation programs, can augment the control of both trunk and leg movements in individuals with chronic, motor complete spinal cord injury [39].

18.3.4 Other Recent Studies of SCS for Improving Motor and Autonomic Functions After SCI Following the three seminal studies from 2018, which focused on overground walking, several groups sought to improve a wider range of motor

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functions and additionally target autonomic functions. In terms of motor functions, trunk stability turned out to be a key element to improve in motor- and sensory-complete SCI, as already shown by Angeli and colleagues in 2018 [94]. Later, Gill and colleagues also demonstrated in two participants with motor- and sensorycomplete SCI that SCS could improve seated reaching distance [97]. On the technological side, Rowald and colleagues showed that a longer electrode array could target both trunk and lower limb motor functions in subjects with complete SCI and considerably improve several daily living and leisure activities that critically require trunk stability [39]. At the University of Minnesota, Darrow and colleagues showed in two female participants with chronic motor- and sensory-complete SCI (AIS-A) that tonic SCS could immediately enable volitional leg movements [98]. In a follow-up study, Pena Pino and colleagues studied the effect of long-term exposure to tonic SCS without intensive neurorehabilitation [99]. After one month of optimization of various stimulation programs for volitional motor control, spasticity, and autonomic functions, participants were allowed to use SCS at home as much as 24 h a day during their daily living activities. Out of seven participants with motor-complete SCI, four of them (all graded AIS-A) recovered the ability to perform voluntary movements even without SCS after a period ranging from 3 to 13 months. Importantly, these movements were not present at every clinical visit, showing variability over time in these motor effects. Even more importantly, higher levels of spasticity seemed to correlate positively with the recovery of voluntary movements. These results add up to the recovery without SCS observed by Rejc and colleagues [88], and independently by Wagner and colleagues [34]. In terms of autonomic functions, Darrow and colleagues showed that SCS improved bowelbladder synergy in their two participants, with SCS, and cardiovascular function in one of the two participants who had otherwise drops in blood pressure during tilt-table tests [98]. This same participant also reported the ability to

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achieve orgasm during sexual intercourse when SCS was on or immediately after it was turned off. Although a thorough review is beyond the scope of this chapter, we would like to highlight that SCS has been shown to improve bladder function [21, 98, 100], body composition, and metabolism [101], and blood pressure [98, 102– 104]. Targeting such autonomic functions is of tremendous importance for improving the quality of life of people with SCI.

18.3.5 Comparison Between SCS and Functional Electrical Stimulation (FES) Functional Electrical Stimulation is an established technology that targets efferent axons innervating specific muscles to produce a desired movement, using electrical stimulation applied at the surface of the skin [105–107] or with leads implanted in the periphery [108, 109]. FES has important clinical applications in hemiplegia, used for example as a commercially available foot-drop stimulator, and for the rehabilitation of upper extremity motor function. Moreover, it has been extensively used in research applications for SCI [107, 110–112], but did not become a standard clinical practice for this condition. In this paragraph, we discuss the conceptual and practical differences between SCS and FES.

18.3.5.1 Conceptual Differences: Stimulation of Muscles Versus Spinal Circuits Functional Electrical Stimulation aims at generating force and movement by recruiting the efferent axons that innervate muscle fibers via pulses of electrical stimulation. This direct recruitment of muscles enables a high degree of controllability because each targeted muscle can be independently stimulated. However, the stimulation patterns required to coordinate a functional movement can be extremely complex and are gravity-dependent [110]. Therefore, a specific set of parameters can only work for a pre-determined movement but can hardly be generalized [108]. This aspect significantly

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increases the complexity of FES systems, as they must be specifically tuned for each task. Conceptually, SCS works very differently than FES, as SCS engages surviving spinal circuits below the lesion via their input fibers, the excitatory sensory afferents. SCS, therefore, overcomes some of the limitations of FES because it requires simple stimulation protocols that leverage existing neural architectures to perform complex movements of a whole limb [33, 34, 113]. Indeed, excitatory spinal circuits producing synergistic movements receive rich innervation from the primary afferents stimulated by SCS [62, 114] and a single Ia afferents connects to all the motoneurons of the homonymous muscle and up to 60% of synergistic motoneurons even at different joints [115]. Another key difference is that FES imposes a specific movement according to a preprogrammed pattern, irrespective of the subject’s voluntary intention. On the other hand, SCS protocols are thought to synergistically act and enhance residual voluntary inputs. Motor outputs can then be modulated and naturally shaped by movement-specific feedback [75] as well as volitional contributions [34, 35]. Concerning the ability to produce large forces, FES suffers from the “inverse recruitment effect”. Since large axons have a lower threshold to electrical stimulation than smaller diameter fibers [48], FES systems first recruit large motor axons instead of smaller axons [116]. Large motor axons recruit muscle fibers that generate large forces but are not resistant to fatigue. This is the opposite of what happens with a natural movement, during which larger fibers are only activated when substantial forces are required. This artificial inverse recruitment rapidly leads to the generation of fatigue, making it technically challenging to produce and sustain large forces [117, 118]. By contrast, SCS does not recruit spinal motoneurons directly. The activation of spinal motoneurons by means of pre-synaptic recruitment of primary afferents leads to a natural recruitment order that is resistant to fatigue and can produce forces capable of sustaining the whole body weight for extended periods of time [72, 90].

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Finally, FES applications that rely on the stimulation of motor axons in the peripheral nerves bypass the spinal cord and consequently cannot directly lead to neuroplasticity of spinal circuits. By contrast, such neuroplasticity is believed to mediate the neurological recovery observed during neurorehabilitation facilitated by SCS [119, 120].

18.3.5.2 Practical Differences: Assistance Versus Therapy Because of the difficulty to coordinate complex activations of muscles, both implantable and non-invasive FES systems can almost exclusively be used in controlled environments. For this reason, FES therapy is applied during laboratory or clinic sessions of physical therapy. In this sense, an FES system works to assist physical exercise with a therapeutic goal. It is not a wearable assistive system that supports daily living activities in community settings. By contrast, epidural SCS is a fully implantable system that is seamlessly integrated in patient’s lives. Therefore, SCS can be used both to assist physical therapy as well as support activity of daily livings in community settings [34, 39]. Outside the laboratory or clinics, patients are able to use their fully-implanted systems similarly to what patients with Parkinson’s disease do with a DBS implant. In this regard, SCS addresses needs of assistance that cannot be addressed with modern FES devices.

18.3.6 Conclusion: SCS, a Promising Neuroprosthetic Technology and Neurorehabilitation Therapy After SCI In this section, we have described how early human studies using tonic SCS in people with SCI laid the groundwork for its subsequent use as a neurorehabilitation technique, in particular its

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motor-enabling effect which dates back to 1986 [23], and the discovery of stepping-like movements in humans with SCI [27]. Next, the recent re-discovery of the motor-enabling effect of SCS in 2011 led to its integration into intensive rehabilitation programs for volitional control of joint movements and standing [85]. In parallel with these clinical studies, the development of new stimulation paradigms for delivering SCS to the spinal cord, called spatiotemporal SCS, restored locomotion in rodent and non-human primate models of SCI [33, 62]. Finally, we showed that all these scientific and technological advances converged to the demonstration that SCS combined with intensive rehabilitation can support the recovery of voluntary motor control and overground walking in participants with severe and chronic SCI [34, 37, 94]. Current studies now aim at expanding the range of motor and autonomic functions enabled by SCS (e.g., [39, 98]). Overall, SCS has shown promises both as neuroprosthesis, i.e., an assistive technology that replaces a lost function and provides immediate relief to a particular deficit, and a rehabilitation tool, i.e., a means to train people with SCI to improve their motor functions when combined with intensive physiotherapy (Fig. 18.5). Although both tonic and spatiotemporal SCS bear great potential in terms of neurorehabilitation, spatiotemporal SCS leads to faster functional outcomes in the absence of training, hence a good indication as an effective neuroprosthesis. Another key aspect to consider is the emergence of the long-term recovery of motor functions in the absence of SCS [34, 88], which likely relies on neuroplastic mechanisms triggered by prolonged use of SCS [119, 120]. In conclusion, although clinical studies have uncovered a formidable potential of SCS both for neuroprosthetics and neurorehabilitation, technological developments and larger multicentric clinical trials are required to assess its safety and efficacy for the millions of people living with SCI worldwide.

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Fig. 18.5 SCS combined with rehabilitation leads to long-term motor improvements. a Functional outcomes that can be reached after tonic or spatiotemporal SCS combined with intensive rehabilitation over several months. For each picture, the severity of injury of the depicted individual is indicated. Adapted with permission from [34, 37, 94]. b Neurological recovery observed in two subjects with incomplete SCI after six months of intensive rehabilitation combined with SCS. Left: plots reporting changes in 6-min and 10-m walk tests. Tests were performed without bodyweight support, following clinical standards. Middle: evaluations of isometric torque production for each joint, quantified before surgery and after rehabilitation without SCS. Right: changes in lower limb motor and sensory scores after rehabilitation. Changes in motor and sensory scores on abbreviated injury scale for all levels below injury are summarized (motor scores; 0: total paralysis, 1: palpable or visible contraction, 2: active movement, gravity eliminated, 3: active movement against gravity, 4: active movement against some resistance, 5: active movement against full resistance). Adapted with permission from [34]. All rights reserved

18.4

Transcutaneous SCS as a NonInvasive Complement to Epidural SCS

18.4.1 Non-Invasive SCS: Stimulating Posterior Roots via Transcutaneous Electrodes Fifty years after the development of epidural SCS, transcutaneous SCS was developed as a non-invasive method to activate similar neural structures as epidural SCS, i.e., the posterior roots, but from outside the skin [121]. This study was also the first to suggest that “continuous transcutaneous posterior root stimulation represents a novel, non-invasive, neuromodulative

approach for individuals with different neurological disorders”. Inspired by an earlier method of high-voltage percutaneous electrical stimulation of the anterior roots to assess afferent peripheral nerve conduction [122, 123], transcutaneous SCS utilizes skin-surface electrodes with one electrode over the spine at the thoracolumbar junction overlying the lumbosacral spinal cord, and a much larger return electrode placed over the lower abdomen or the iliac crests. Computational modeling of epidural SCS has predicted low-threshold sites along proprioceptive fibers at the posterior rootlet-spinal cord interface [31]. In fact, the high angles in fiber orientation and the crossing of the electrical conductivity boundary between the cerebrospinal fluid and the spinal cord make the recruitment of posterior roots possible also by skin-surface

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electrodes, albeit with lower root specificity than epidural SCS [30]. Because of the high external voltages required to elicit sufficiently strong voltages inside the vertebral bones, transcutaneous SCS requires external stimulators similar to those typically used for traditional functional electrical stimulation (FES) or transcutaneous electrical nerve stimulation (TENS). With this approach, it is possible to directly stimulate proprioceptive afferent fibers within the posterior roots and evoke activity in extremity muscles through spinal reflexes that resemble those elicited by epidural SCS [121, 126, 127]. The sharing of the same low-threshold sites along the posterior root afferent fibers between epidural and transcutaneous SCS found by studies in computational modeling [30, 128], combined with the nearidentical reflex responses evoked by both stimulation methods in humans with SCI (Fig. 18.6 a), suggest that both epidural and transcutaneous SCS recruit common neural structures through similar mechanisms [46]. However, because of the larger distance from the spine and the reduced focality of the electrical field, the specificity in muscle recruitment is lower with transcutaneous SCS compared to its epidural counterpart [32, 34, 40, 129].

18.4.2 Transcutaneous SCS for Generating Locomotor-Like Movements Despite reduced specificity, transcutaneous SCS remains an attractive solution for applications that do not seek to achieve highly selective muscle recruitment. Indeed, it does not require any surgical procedure, thereby significantly reducing the risks and costs of the intervention. It also provides a potentially inclusive and affordable solution to obtain, at least in part, motor improvements comparable with those achieved by epidural SCS.

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Therefore, multiple studies have investigated the possibility to augment EMG activity and movements (Fig. 18.6b), as well as functional movements during active treadmill stepping in individuals with chronic, motor-incomplete SCI (AIS D) [49, 130, 131]. As a robotic gait orthosis moved the legs of individuals with clinically complete SCI (AIS A) in a walking pattern over a treadmill, a small number of muscles exhibited electromyographic responses [125]. Adding 30 Hz transcutaneous SCS increased the number of activated leg muscles, which had rhythmic activity during the different phases of gait (Fig. 18.6c). Moreover, transcutaneous SCS alone could produce rhythmic activation of leg muscles without the need for triggered stimulation. Four individuals with incomplete SCI (AIS D) were able to voluntarily modify the generated lower limb muscle activity [131, 132]. By consciously modifying their augmented muscle activity according to the gait phase, participants were able to improve the quality of their stepping kinematics, the range of hip and knee movements, and their stride length. As in epidural SCS, turning the stimulation off would immediately cause degradation of walking quality and muscle activity. The current interpretation of these results is that mechanisms of transcutaneous SCS are at least partially similar to those of epidural SCS. In particular, this interpretation implies that transcutaneous SCS interacts with the flow of stepinduced proprioceptive feedback to modulate muscle activity during locomotion through spindle feedback circuits [72]. However, the rhythmic activation without peripheral feedback suggests that transcutaneous SCS could also engage CPGs [27, 70, 71, 133] in addition to proprioceptive feedback circuits [29, 63, 75]. Moreover, a summation process between residual supraspinal inputs (via clinically silent translesional neural connections that survived the injury), and the increased spinal circuits excitability enhanced by transcutaneous SCS, may be a likely explanation for the voluntary

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control of movement enabled in otherwise paralyzed muscles (Fig. 18.6b) [125].

18.4.3 Stimulation Parameters for Transcutaneous SCS Two general methods for stimulation pulse shape configuration are currently used to deliver

transcutaneous SCS: conventional and “Russian” current stimulation. Conventional currents consist of mono- or biphasic charge-balanced pulses with rectangular pulses of 1 ms width, delivered at frequencies ranging from about 5 to 100 Hz. Stimulation at 30 Hz is commonly applied to elicit muscle activity or movement, while 50 Hz pulses are used to improve spasticity [130, 134– 136]. The “Russian current” is composed of 1 ms

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bursts filled with 10 kHz pulses [137]. Several studies have suggested that it may be possible to use these currents to reduce the potential discomfort of transcutaneous SCS directly below the stimulating surface electrodes [138, 139], while maintaining recruitment of deep neural structures to elicit comparable electrophysiological responses as conventional transcutaneous SCS. The hypothesis is that the temporal summation of graded potentials created by the rapid depolarization and repolarization of highfrequency stimulation may raise the membrane potential of larger fibers enough to be activated without depolarizing unmyelinated C-fibers [140]. Additionally, there is some neurophysiological evidence suggesting C-fibers are less likely to fire in response to high-frequency stimulation compared to large-diameter fibers [141, 142]. However, a recent study comparing the tolerance and responses elicited by the “Russian current” and conventional transcutaneous SCS protocols contradicted this view. While participants could indeed tolerate significantly higher levels of stimulation amplitude with Russian currents, both protocols produced the same amount of discomfort when the amplitudes were adjusted to obtain the same spinallyevoked muscle responses [143]. Therefore, the apparent higher comfortability may be due to the fact that it takes higher currents with 10 kHz burst to accumulate enough charge to stimulate the same deep structures in the spinal cord as with standard protocols.

18.4.4 Functional Recovery by LongTerm Transcutaneous SCS and Activity-Based Training Long-term activity-based training with transcutaneous SCS has also been investigated as a method to induce functional recovery in the chronic phase after SCI. An 18-week training strategy involving voluntary modulation of

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transcutaneous SCS-generated locomotor movements in combination with buspirone, an orally active serotonergic agonist, was tested in five individuals with chronic motor-complete, sensory-incomplete (AIS B) individuals [137]. The ability of individuals to voluntarily modulate the step-like movements induced by transcutaneous SCS improved with training, and participants gained the ability to generate voluntary movements in previously paralyzed muscles, even without stimulation. With stimulation, their motor function was equivalent to that of individuals with AIS C impairments. In a subsequent single-case study to evaluate the contribution of buspirone [124], a 4-week training paradigm in robot-assisted overground training found that transcutaneous SCS alone or in combination with buspirone, but not buspirone alone, resulted in the lowest dependence on robotic assistance. Moreover, the participant could perform voluntary knee flexion when in the supine position after a 1-week training with transcutaneous SCS alone, but not after 1-week training with the drug alone. Sayenko and colleagues [139] subsequently demonstrated the ability of transcutaneous SCS to enable standing without previous training in 15 individuals with chronic SCI (AIS A, B, C), and a 12-session training program followed by six participants improved their upright balance control and reduced their dependence on external assistance. A post-training clinical evaluation revealed an increase in muscle tone of all participants. Notably, the results of this study are quantitatively similar to those seen in previous investigations with SCS [85, 86]. Although a single session of transcutaneous SCS can facilitate residual voluntary control of single joints and modify the excitability of cortical, corticospinal, and spinal reflexes [144– 146], it is not sufficient to observe statistically significant improvements in walking performance [145]. Recent studies by several groups highlight the importance of combining SCS therapy with activity-based training to enable consistent improvements in walking [147, 148] and sit-to-stand ability [149].

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18.4.5 Recent Advances to Improve Muscle Recruitment Selectivity in Transcutaneous SCS Transcutaneous SCS presents a promising noninvasive approach to enable movement, locomotor function, and long-term recovery after SCI. However, the low specificity of transcutaneous SCS in muscle recruitment compared to its epidural counterpart may limit the type of movements and exercises that can be supported and trained during rehabilitation. Nevertheless, recent studies by Calvert, Krenn, and colleagues have demonstrated that positioning surface electrodes in lateralized (Fig. 18.7a) and rostro-caudal locations (Fig. 18.7 b) can target specific mediolateral and rostro-caudal spinal cord circuitry toward improved muscle recruitment selectivity [150, 151]. These observations suggest that future advances in electrode configurations, as well as stimulation amplitudes, frequencies, and timing, may further enhance the potential of transcutaneous SCS in non-invasive rehabilitation approaches.

18.4.6 Conclusion: Transcutaneous SCS is Less Specific Than Epidural SCS but Provides an Inclusive Access to Advanced Healthcare In summary, transcutaneous SCS is a promising tool to investigate the combination of SCS and rehabilitation for applications that do not require high specificity. However, the high levels of current required to produce substantial muscle activity to support the body weight in people with severe motor paralysis may be uncomfortable and cause large contractions of the back and abdominal muscles. Nevertheless, we believe that transcutaneous SCS should not necessarily be seen in opposition with epidural SCS, but they can be seen as having unique advantages. While less invasive, transcutaneous SCS requires electrode application for every use and requires an

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Fig. 18.7 Limitations in muscle recruitment selectivity by transcutaneous SCS can be partially overcome by electrode configuration. a Elicitation of spinal reflexes by lateral and midline transcutaneous SCS. Transcutaneous SCS applied *2 cm lateral to the midline of the lumbosacral spinal cord can selectively activate ipsilateral spinal sensorimotor circuits and thus ipsilateral lower extremity muscles. Modified with permission from [150]. b Elicitation of posterior root-muscle reflexes by stimulation from different rostro-caudal sites along a multielectrode array. The multi-electrode array was 152 mm in length with row D positioned over the T11-T12 vertebra. Stimulation of rostral electrodes (row A) elicits responses in quadriceps but not in triceps surae. In contrast, stimulation of the caudal-most row elicits a large response in triceps surae and a small response in quadriceps. Modified with permission from [151]. MG, medial gastrocnemius; MH, medial hamstrings; SOL, soleus; TA, tibialis anterior; VL, vastus lateralis; L, left; Q, quadriceps; TS, triceps surae. All rights reserved

external stimulator connected to the electrodes, and may thus suffer from limited repeatability and portability. Therefore, transcutaneous SCS is very well suited to study and perform SCSenhanced physical training and rehabilitation protocols in the hospital, and perhaps at home

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with appropriate guidance, but does not currently serve as a neuroprosthetic intervention that sustains motor activity in daily living. Instead, while more invasive, epidural SCS is fully implantable and therefore, by definition portable, thus sustaining motor activities for people with SCI not only in the clinic but in their daily life. These differences should be considered in the evaluation of the risk–benefit ratio, as it is possible that the interactions between descending voluntary input and SCS are limited when the stimulation is turned off. This would reduce the effectiveness of transcutaneous SCS compared to epidural SCS, which can instead be always on. Nevertheless, transcutaneous SCS represents an affordable tool to amplify the outcome of rehabilitation in many centers, for example, in rural areas where access to neurosurgery expertise and expensive invasive devices may be limited.

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Spinal Cord Stimulation to Enable Leg Motor … individuals with incomplete spinal cord injury. J Spinal Cord Med. 2013;37(2):202–11. Minassian K, Hofstoetter U, Tansey K, Rattay F, Mayr W, Dimitrijevic M. Transcutaneous stimulation of the human lumbar spinal cord: facilitating locomotor output in spinal cord injury. 2010. Minassian K, Hofstoetter US, Rattay F. Transcutaneous lumbar posterior root stimulation for motor control studies and modification of motor activity after spinal cord injury. In: Restorative neurology of spinal cord injury. 2011. p. 226–55. Wiley J. The human central pattern generator for locomotion: does it exist and contribute to walking? (Unpublished). The Neuroscientist. 2010. Estes SP, Iddings JA, Field-Fote EC. Priming neural circuits to modulate spinal reflex excitability. Front Neurol. 2017;8:17. Hofstoetter US, Freundl B, Lackner P, Binder H. Transcutaneous spinal cord stimulation enhances walking performance and reduces spasticity in individuals with multiple sclerosis. Brain Sci. 2021;11(4):472. Hofstoetter US, Freundl B, Danner SM, Krenn MJ, Mayr W, Binder H, et al. Transcutaneous spinal cord stimulation induces temporary attenuation of spasticity in individuals with spinal cord injury. J Neurotrauma. 2020;37(3):481–93. Gerasimenko Y, Gorodnichev R, Moshonkina T, Sayenko D, Gad P, Edgerton VR. Transcutaneous electrical spinal-cord stimulation in humans. Ann Phys Rehabil Med. 2015;58(4):225–31. Gad P, Lee S, Terrafranca N, Zhong H, Turner A, Gerasimenko Y, et al. Non-invasive activation of cervical spinal networks after severe paralysis. J Neurotrauma. 2018;35(18):2145–58. Sayenko DG, Rath M, Ferguson AR, Burdick JW, Havton LA, Edgerton VR, et al. Self-assisted standing enabled by non-invasive spinal stimulation after spinal cord injury. J Neurotrauma. 2019;36 (9):1435–50. Ward AR, Chuen WLH. Lowering of sensory, motor, and pain-tolerance thresholds with burst duration using kilohertz-frequency alternating current electric stimulation: part II. Arch Phys Med Rehabil. 2009;90(9):1619–27. Joseph L, Butera RJ. High-frequency stimulation selectively blocks different types of fibers in frog sciatic nerve. IEEE Trans Neural Syst Rehabil Eng. 2011;19(5):550–7.

399 142. Torebjörk HE, Hallin RG. Responses in human A and C fibres to repeated electrical intradermal stimulation 1. J Neurol Neurosurg Psychiatry. 1974;37(6):653. 143. Manson GA, Calvert JS, Ling J, Tychhon B, Ali A, Sayenko DG. The relationship between maximum tolerance and motor activation during transcutaneous spinal stimulation is unaffected by the carrier frequency or vibration. Physiol Rep. 2020;8(5): e14397. 144. Kumru H, Flores Á, Rodríguez-Cañón M, Edgerton VR, García L, Benito-Penalva J, et al. Cervical electrical neuromodulation effectively enhances hand motor output in healthy subjects by engaging a use-dependent intervention. J Clin Med. 2021;10 (2):195. 145. Meyer C, Hofstoetter US, Hubli M, Hassani RH, Rinaldo C, Curt A, et al. Immediate effects of transcutaneous spinal cord stimulation on motor function in chronic, sensorimotor incomplete spinal cord injury. J Clin Med. 2020;9(11):3541. 146. Parhizi B, Barss TS, Mushahwar VK. Simultaneous cervical and lumbar spinal cord stimulation induces facilitation of both spinal and corticospinal circuitry in humans. Front Neurosci. 2021;15:615103. 147. Estes S, Zarkou A, Hope JM, Suri C, Field-Fote EC. Combined transcutaneous spinal stimulation and locomotor training to improve walking function and reduce spasticity in subacute spinal cord injury: a randomized study of clinical feasibility and efficacy. J Clin Med. 2021;10(6):1167. 148. Shapkova EY, Pismennaya EV, Emelyannikov DV, Ivanenko Y. Exoskeleton walk training in paralyzed individuals benefits from transcutaneous lumbar cord tonic electrical stimulation. Front Neurosci. 2020;14:416. 149. Al’joboori Y, Massey SJ, Knight SL, Donaldson N de N, Duffell LD. The effects of adding transcutaneous spinal cord stimulation (tSCS) to sit-to-stand training in people with spinal cord injury: a pilot study. J Clin Med. 2020;9(9):2765. 150. Calvert JS, Manson GA, Grahn PJ, Sayenko DG. Preferential activation of spinal sensorimotor networks via lateralized transcutaneous spinal stimulation in neurologically intact humans. J Neurophysiol. 2019;122(5):2111–8. 151. Krenn M, Hofstoetter US, Danner SM, Minassian K, Mayr W. Multi-electrode array for transcutaneous lumbar posterior root stimulation. Artif Organs. 2015;39(10):834–40.

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Open Access This chapter is distributed under the terms of the Creative Commons Attribution 4.0 International License (http://creativecommons.org/licenses/by/4.0/), which permits use, duplication, adaptation, distribution and reproduction in any medium or format, as long as you give appropriate credit to the original author(s) and the source, a link is provided to the Creative Commons license and any changes made are indicated. The images or other third party material in this chapter are included in the work’s Creative Commons license, unless indicated otherwise in the credit line; if such material is not included in the work’s Creative Commons license and the respective action is not permitted by statutory regulation, users will need to obtain permission from the license holder to duplicate, adapt or reproduce the material.

19

Functional Electrical Stimulation Therapy: Mechanisms for Recovery of Function Following Spinal Cord Injury and Stroke Milos R. Popovic, Kei Masani, and Matija Milosevic

Abstract

Electrical stimulation is a tool that applies low-energy electrical pulses to artificially generate muscle contractions. If electrical stimulation is used to enable functional movements, such as walking and grasping, then this intervention is called functional electrical stimulation (FES). When FES is used as therapy instead of being used as an orthosis, it is called FES therapy or FET. In this chapter, we introduce recent findings and advances in the field of FET. The findings to date clearly show that FET for reaching and grasping is a therapeutic modality that should be imple-

M. R. Popovic (&) . K. Masani Institute of Biomedical Engineering, University of Toronto, Toronto, ON, Canada e-mail: [email protected] K. Masani e-mail: [email protected] M. R. Popovic . K. Masani KITE Research Institute, Toronto Rehabilitation Institute – University Health Network, Toronto, ON, Canada M. R. Popovic . K. Masani CRANIA, University Health Network & University of Toronto, Toronto, ON, Canada M. Milosevic Graduate School of Engineering Science, Osaka University, Toyonaka, Osaka, Japan e-mail: [email protected]

mented in every rehabilitation institution that is treating individuals with stroke and Spinal Cord Injury (SCI). There is also considerable evidence to support the use of FET as a therapeutic modality to treat drop-foot problem in stroke and incomplete populations. Although phase I randomized control trials have been completed with chronic SCI population using this new FET technology and preliminary findings are encouraging, further research and development are required before the multichannel FET for walking will be ready for clinical implementation. Finally, emerging evidence for the beneficial use of brain-computer interface (BCI) combined with FET (BCI-FET) for improving upper and lower limb function will also be presented. Keywords

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Functional electrical stimulation (FES) Therapy FES therapy Spinal cord injury Stroke Rehabilitation Neuroprosthesis Neurorehabilitation Neuromodulation Restoration of voluntary function Brain-computer interface (BCI)

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Functional electrical stimulation (FES) is a technology one can use to artificially generate body movements in individuals who have paralyzed muscles due to injury to the central nervous

© The Author(s), under exclusive license to Springer Nature Switzerland AG 2022 D. J. Reinkensmeyer et al. (eds.), Neurorehabilitation Technology, https://doi.org/10.1007/978-3-031-08995-4_19

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system. More specifically, FES can be used to generate functions such as grasping and walking in individuals with paralyses such as stroke and spinal cord injury (SCI). This technology was originally used to develop neuroprostheses that were implemented to permanently substitute impaired functions such as bladder voiding, grasping, and walking. In other words, a consumer would use the device each time he wanted to generate a desired function. In recent years FES technology has been used to deliver therapies to retrain voluntary motor functions such as grasping, reaching, and walking. In this embodiment, FES is used as a short-term therapy for several weeks to months, with the objective to restore voluntary function and not lifelong dependence on the FES device, hence the name FES therapy or FET. In other words, FET is used as a short-term intervention to help the central nervous system of the consumer to relearn how to execute impaired functions instead of making the consumer dependent on neuroprostheses for the rest of her/his life. In this chapter, we introduce recent findings and advances in the field of FET.

19.2

Functional Electrical Stimulation (FES)

19.2.1 Definitions Individuals with stroke and SCI have injuries that prevent the central nervous system from generating a desired motor command and/or transmitting the desired motor command to the parts of the peripheral nervous system that innervate muscles. As a result, these individuals are frequently unable to voluntarily move different body parts and perform functions such as sitting, standing, reaching, grasping, and bladder voiding. However, as long as the peripheral nerves innervating the muscles, the muscles themselves, and the joints and soft tissues supporting the muscle-joint structures are intact, the electrical stimulation can be used to generate joint movements by contracting the muscles that actuate them. The electrical stimulation used for this

purpose is called neuromuscular electrical stimulation (NMES). An organized and patterned NMES that aims to generate coordinated limb or body movements such as grasping, standing, and walking, instead of isolated muscle contractions is called functional electrical stimulation (FES). In such a context, the FES technology is used as a prosthetic/orthotic device. In the literature, this use of FES technology is referred to as a neuroprosthesis or neuroprosthetics.

19.2.2 Physiology In nerve cells, information is coded and transmitted as a series of electrical impulses called action potentials, which represent a brief change in cell electric potential of approximately 80– 90 mV. These nerve signals are frequency modulated; that is, the number of action potentials that occur in a unit of time is proportional to the intensity of the transmitted signal. The typical frequency of action potentials is between 8 and 20 Hz [1]. An electrical stimulation can artificially elicit this action potential by changing the electric potential across nerve cell/axon membranes by inducing electrical charge in the immediate vicinity of the outer membrane of the cell (Fig. 19.1). In most of the applications, FES activates the nerves. However, in some applications, FES can be used to directly stimulate muscle fibers if their peripheral nerves have been severely damaged (i.e., denervated muscles) [2]. The majority of the FES systems used today to stimulate the nerve trunk or the nerve ending at the neuromuscular junction. The main reason is the fact that direct muscle stimulation requires considerably more energy to generate contractions (at least three orders of magnitude more [3]), which makes these systems more challenging to implement at home and in clinical settings. Nevertheless, it should be noted that an electric stimulator that has been purposefully designed to generate contractions in denervated muscles is currently commercially available. Its name is Stimulette edition5, and it is manufactured by Dr. Schuhfried, Medical Technology, Austria (www.

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Fig. 19.1 A schematic representation of the surface functional electrical stimulation (FES) system. The FES system causes a muscle contraction by electrically stimulating the motor axons that are connected to the muscles. The electrical stimulation generates action potentials in the motor neurons, which propagate along the motor neurons toward the muscle. When the action potentials reach the muscle, they cause the muscle to contract

schuhfriedmed.at). In the remainder of this document, we will only discuss FES systems that have been developed to stimulate innervated muscles. In some FES applications, the stimulation electrode is located on muscle bellies, while the stimulation electrode is located over superficial nerve trunk in other FES applications. While differences exist in how nerve trunk and muscle belly stimulation affect the recruitment of motor (efferent) and sensory (afferent) nerves [4], applications of FES on the muscle belly can produce selective muscle contractions to generate functional movements. In most FES applications, the electrical charge can stimulate both motor nerves (efferent nerves —descending nerves from the central nervous system to muscles) and sensory nerves (afferent nerves—ascending nerves from sensory organs to the central nervous system). In some applications, the nerves are stimulated to generate localized muscle activity, i.e., the stimulation is aimed at generating muscle contraction orthodromically via motor nerves. In other applications, stimulation is used to activate simple or complex reflexes via sensory nerves. In other

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words, the sensory nerves are stimulated to evoke a reflex, which is typically expressed as a coordinated contraction of one or more muscles. Notably, activation of the sensory system (i.e., the reflex pathway) is believed to be one of the important contributing factors for the recovery of motor function after FES [5]. When a nerve is stimulated, i.e., when sufficient electrical charge is provided to a nerve cell, a localized depolarization of the cell wall occurs resulting in an action potential that propagates toward both ends of the axon. Typically, one “wave” of action potentials will propagate along the axon toward the muscle (orthodromic propagation), and concurrently, the other “wave” of action potentials will propagate toward the cell body in the central nervous system (antidromic propagation). While the direction of propagation in case of the antidromic stimulation and the sensory nerve stimulation is the same, i.e., toward the central nervous system, their end effects are very different. The antidromic stimulus has been considered an irrelevant side effect of FES. According to Rushton [6], repeated antidromic stimulation through Hebb-type processes may over time enable sparse supraspinal commands to activate anterior motor neuron and produce desired muscle contraction(s). Typically, FES is concerned with orthodromic stimulation and uses it to generate coordinated muscle contractions. In the case where sensory nerves are stimulated, the reflex arcs are triggered by the stimulation of sensory nerve axons at specific peripheral sites. One example of such a reflex is the flexor withdrawal reflex. The flexor withdrawal reflex occurs naturally when a sudden, painful sensation is applied to the sole of the foot. It results in flexion of the hip, knee, and ankle of the affected leg and extension of the contralateral leg in order to get the foot away from the painful stimulus as quickly as possible. The sensory nerve stimulation can be used to generate desired motor tasks, such as evoking flexor withdrawal reflex to facilitate walking in individuals following stroke, or they can be used to alter reflexes or the function of the central nervous system. In the latter case, the electrical

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stimulation is commonly described by the term neuromodulation.

19.2.3 Technology Nerves can be stimulated using either surface (transcutaneous) or subcutaneous (percutaneous or implanted) electrodes. The surface electrodes are placed on the skin surface above the nerve or muscle that needs to be “activated.” They are non-invasive, easy to apply, and generally inexpensive. To generate muscle contraction, the impedance, location, size, and orientation of the electrodes need to be optimized to maximize the current flow and muscle recruitment. Placing a smaller cathode electrode closer to the target nerve with the larger anode away from the cathode can be used to generate more accurate stimulation under the cathode while allowing a larger area of the skin under the anode to be used to close the electrical circuit and minimize discomfort. Empirically, we know that there are motor point locations where muscles are most sensitive to electrical stimulation. Placement of electrodes on these motor points also plays an important role in generating strong muscle contractions (for a review of upper and lower limb motor points, see Bersch et al. [7] and Botter et al. [8], respectively). Typically, smaller muscles tend to have one motor point [7], while larger muscles are known to have several (e.g., quadriceps group has seven motor points [8, 9]). Until recently the common belief in the FES field has been that due to the electrode–skin contact impedance, skin and tissue impedance, and current dispersion during stimulation, much higher-intensity pulses are required to stimulate nerves using surface stimulation electrodes as compared to the subcutaneous electrodes. This statement is correct for all commercially available stimulators except MyndMove® stimulator (Fig. 19.2), which is manufactured by a Canadian company MyndTec (www.myndtec.com). MyndMove® has implemented a new stimulation pulse that allows the stimulator to generate muscle contractions using electrical pulses, which steady-state amplitudes are 10–15 times

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lower in intensity than those required by other transcutaneous electrical stimulation systems. The key aspects of this new technology are stimulation pulses that have very a fast slew rate, which is the time that a circuit needs to go from 0 to targeted amplitude within a single pulse (US Patent 20130090712), and are able to rapidly engage Aa efferent nerve fibers (i.e., descending nerves from the central nervous system to muscles) using very low stimulation amplitudes and at the same time minimize engagement of afferent Ad and C nerve fibers responsible for the transmission of pain sensation [10]. This new technology not only reduces the intensity of stimulation, but it also reduces discomfort during stimulation, which is a common problem with commercially available transcutaneous electrical stimulation systems [10]. Typical FES systems could use different types of waveforms to stimulate the muscles. Common stimulation pulses are balanced biphasic impulses that ensure that the residual charge in the tissues is removed, and they also generate a discharge under both the anode and the cathode. Most FES systems use asymmetric balanced biphasic impulses to ensure that the muscle contractions occur only under the stimulating cathode electrode. The magnitude of muscle contractions can be varied by changing the stimulating pulse amplitude, pulse width, or pulse frequency. The pulse amplitude, or intensity, is related to the depolarizing effect, with higher amplitudes inducing a stronger depolarization. Increasing the amplitude results in the additional recruitment of smaller fibers near the electrode and larger fibers farther from the electrode [11]. The pulse width, or pulse duration, required to achieve adequate depolarization and cause muscle to contractions is typically around 200–500 ls. The frequency during FES determines the rate of action of the motor and sensory pathways. Depending on the application, a variety of frequencies can be used to generate contractions using FES, with most in the range between 20–50 Hz. Such frequencies are needed because electrical stimulation activates muscle fibers synchronously, which requires higher firing rates to generate tetanic contractions. Lower

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Fig. 19.2 Use of MyndMove® to retrain upper arm voluntary movements (Photo courtesy MyndTec Inc., Toronto, ON, Canada)

frequency stimulation (70 Hz), such as high- gamma band, and the relative immunity to movement artifacts can provide a potential boost in BCI control accuracy. Wang et al. [38] demonstrated that high-density ECoG signals can be used to decode 6 DOF from the upper extremity. Wang et al. [39] demonstrated that ECoG signals could be used by an individual with tetraplegia to control a robotic arm for a reaching task. Furthermore, Wang et al. [40] demonstrated highly accurate decoding of hand grasping from ECoG signals. In the BrainGate clinical trials, intracortical microelectrode arrays implanted in tetraplegic individuals enabled control of a 6-DOF robotic arm [18, 34]. Collinger et al. [41] utilized two microelectrode arrays implanted into the motor cortex of a patient with severe tetraparesis due to spinocerebellar degeneration to successfully control a 7-DOF robotic arm. Aflalo et al. [42] utilized a microelectrode array to record movement trajectory intention in the posterior parietal cortex and enable a person with tetraplegia to control a robotic arm. Despite BCI advances enabling many motor functions as above, the means by which feedback is received from these prosthetic limbs are still primitive. Real-life movement invariably involves continuous interaction with external objects and the environment and therefore sensory feedback is critical. For example, this is important in the case of grasping a delicate object, where sensory feedback is necessary to guide the grip strength to hold the object without crushing it. Another example is during

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ambulation, in which continuous knowledge of feet and leg position is necessary to maintain balance. The loss of somatosensation is known to cause deficits in motor control [43–45]. Moreover, the theory of optimal feedback control [46] corroborates that humans rely on cost and rewards [47, 48], internal models [49, 50], optimal feedback-driven policy [51], and state estimation [51, 52], all of which demand somatosensory feedback as a crucial component of normal motor control. Therefore, the important challenge for BCI development is to realize a BD-BCI system with the capability to convey sensory information back to the brain. To date, there have been reports of BD-BCIs using either ECoG-based sensory stimulation or intracortical microstimulation (ICMS). Evoked sensory percepts via ECoG electrodes placed over the cortical surface have been evaluated as a viable feedback interface [20, 21, 53] which can potentially guide control of the prostheses. A number of BD-BCIs utilizing ICMS-driven sensory feedback have demonstrated improved prosthetic arm motor control in grasping and transporting objects [22, 54] compared to when stimulation was disabled. This sheds light on the prospect that mimicking known biological control principles could result in task performance approaching that of the able-bodied human. However, existing BD-BCI systems operate in a constrained laboratory setting, since bulky, nonmobile workstation computers, data acquisition systems, and commercial stimulators prevent the untethered, mobile, everyday use of these systems. Practical implementation of BD-BCIs hinges on the integration of the above components into a special purpose and compact form factor with full programmability. Specifically, an envisioned breakthrough in BD-BCI development would likely come in the form of an embedded system small enough to enable a fully implantable BD-BCI that simultaneously restores limb movement and sensation in persons with neurological injuries. It is envisioned that the function of the aforementioned systems can be extended to those

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with hemiplegia due to stroke. For example, distal upper extremity weakness is a common clinical outcome of stroke, and systems such as those by Pfurtscheller et al. [35], may help restore hand movement in stroke survivors. In addition, since as many as 15% of stroke survivors lose their ability to ambulate, BCIcontrolled lower extremity prostheses could help restore walking. At the time of publication, there are no BCI-controlled neuroprostheses that have undergone definitive clinical trials for safety and efficacy. In addition, none have been FDA approved for marketing, and hence no clinical recommendations can be made regarding the application of neuroprosthetic BCIs for stroke. Before these devices reach the point where they can be widely used and adopted, several outstanding problems must be addressed. First, these systems are not yet sufficiently accurate to robustly restore movement to paralyzed limbs. While some literature reports that a performance accuracy of 70% may be sufficient for a BCI user to feel that they have reliable control [55–58], even the most accurate BCIs (those that differentiate between moving and idling states) can only reach 95% accuracy, which may still translate into operation errors that are potentially frustrating or dangerous to the user. Second, they typically require full-sized computers and bulky amplifier arrays, which limit their portability. Significant engineering effort will need to be invested in order to miniaturize the requisite electronic components such that they are wearable, esthetically acceptable, as well as constantly available, and easy to use. Lastly, non-invasive systems require EEG caps which are tedious and time-consuming to don and doff. In order to address these challenges, it may be necessary to develop implantable BCI systems, including electrodes, amplifiers, and special-purpose microcomputers. Ultimately, clinical trials will need to be conducted to determine whether these systems are safe (e.g., do not cause seizures or nervous tissue injuries) and effective (reduce disability).

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22.4

BCI Systems for Physiotherapy

22.4.1 Review of Existing BCI Systems for Stroke Rehabilitation and Underlying Mechanisms The general consensus in stroke motor rehabilitation is that the most effective practices employ repetitive, high-intensity, goal-oriented movement of the impaired limb (such as constraintinduced movement therapy) to overcome learned disuse [59]. Additionally, it is recommended that patients execute these movements as naturally as possible [60, 61]. However, severely disabled individuals may be unable to participate in active movement therapies, and hence BCIs may be applied as novel therapies to facilitate compliance with these rehabilitative guidelines. This is supported by the work of Vouvopoulos et al. [62], which demonstrated that individuals with more severe motor impairments may benefit most from EEG-based neurofeedback compared to EMG-based feedback. While it can be hypothesized that BCI therapy, when used in conjunction with conventional physiotherapies, may also improve motor function in those with moderate or mild impairment due to stroke, these individuals with largely intact sensorimotor pathways may derive the most benefit from EMG-based feedback [62]. Stroke survivors are still able to modulate EEG without performing any physical movement [60, 63–68], and this can be exploited by BCIs for stroke rehabilitation. Moreover, MI- and MEbased BCI therapies can be carried out in a repetitive [66] and goal-oriented [59] manner that ensures intense focus on the motor function task [60, 64, 69, 70]. These BCIs may facilitate neuroplastic cortical changes similar to repetitive movement practice [61], possibly through operant conditioning. For example, the BCI output could provide feedback to the user about his/her cortical state, and the user could then attempt to

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modulate his/her SMRs to achieve maximum control of the BCI. This learning process may lead to subsequent beneficial neural changes [63, 67–69, 71–74] and, in turn, to improved motor function (see Fig. 22.3). For example, MI and ME have been shown to strengthen visuospatial [75], primary/associated motor [61, 66, 74–76], and primary somatosensory [60] networks. Many previous studies have demonstrated that MI and ME generate robust changes in EEG signals that are suitable for BCIs, even in the post-stroke cortex, making BCI-based stroke rehabilitation a possibility. Mohapp et al. [77] used MI and ME on 10 stroke patients with an average BCI classification accuracy between 61.5 and 79.0% (depending on which limb and hemisphere were used). Bai et al. [78] investigated both MI and ME tasks with a BCI that utilized beta-rhythm SMR and found that, without extensive training, the classification accuracy in stroke subjects was comparable to healthy

subjects ( ⩾ 80%). Buch et al. [79] reported successful control of a MEG-based BCI by eight stroke patients who utilized mu-rhythm ERD during both ME and MI tasks to control an orthosis attached to the plegic hand. Six out of the eight patients achieved a significant increase in BCI classification accuracy by the 20th session, with a median accuracy of 72.48% across subjects at the final session. Prasad et al. [80] evaluated five chronic stroke patients undergoing combined physical practice and MI-based BCI and found that BCI classification accuracy was 70% on average. McCrimmon et al. [81] investigated an ME-based FES therapy in nine chronic stroke patients and found that the average classification accuracy for these subjects was 80%. From these studies, it is clear that stroke survivors can successfully control MI- and MEbased BCIs. However, the true relationship between the features used in MI- and ME-based BCIs and functional motor recovery remains

Fig. 22.3 Hypothesized mechanisms of post-stroke motor recovery using BCI systems. Here, the lower motor neuron (LMN) output and subsequently the muscle output are severely impaired. a MI- and ME-based BCIs may elicit cortical changes (yellow square) between the primary motor cortex (M1) and the supplementary motor area (SMA), premotor cortex (PM), and even the prefrontal and posterior parietal areas (not shown).

b MI- and ME-based BCIs that provide robotic assistance may additionally stimulate dorsal sensory pathways and subsequently facilitate neuroplastic changes between the primary somatosensory cortex (S1) and the non-primary motor areas. c BCIs that deliver MI- or ME-controlled FES may also promote changes in the anterior horn of the spinal cord at the level of the antidromically activated LMN

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elusive, as studies such as [82] suggest that robust motor recovery can occur even without changes in these neurophysiological features. Adding a proprioceptive feedback mechanism to MI- and ME-based BCIs may further enhance functional recovery in stroke survivors (see Fig. 22.3). More specifically, proprioceptive BCIs pair motor intention (motor and visuomotor activation) with the movement of the paretic limb (e.g., through robotic assistance). This may facilitate Hebbian-like learning and neural reorganization [61, 66, 67, 83] and ultimately improve motor recovery [68, 69, 84]. It is likely that these plastic changes occur at the level of the cortex and primarily affect motor planning and initiation since synaptic changes directly between upper motor neurons and sensory fibers are unlikely to occur. Lau et al. [85] observed that BCI coupled with proprioceptive robotic hand feedback promoted changes in the ipsilesional and contralesional cortices (such as the primary motor, pre-motor, supplementary motor, and parietal areas) and that these changes were endured for at least 6 months. Additionally, any improvement in motor function results in a subsequent increase in proprioceptive feedback, creating a positive feedback loop of further CNS changes [61]. This proprioceptive BCI concept has been successfully realized in several studies. For example, Broetz et al. [72] and Caria et al. [83] trained a hemiplegic patient with no active finger extension with a BCI that drove an orthosis attached to his paralyzed arm. The patient used mu-rhythm modulation to control the orthosis and underwent goal-directed physiotherapy training over the course of one year. GomezRodriguez et al. [84] evaluated BCI-robotic armassisted physiotherapy in three chronic stroke patients. Patients attempted either elbow flexion or extension or MI, while the BCI would detect their intention to move and then initiate active robotic assistance. Ang et al. [86] studied the effect of MI-based BCI with haptic feedback in 21 chronic stroke patients in a controlled trial. Subjects participated in 18 therapy sessions in which grasping and knob manipulation tasks were carried out using MI-BCI with robotic

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assistance. A recent study by RamosMurguialday et al. [87] showed that using an MI-based BCI with sham robotic feedback (that was not based on the user’s EEG) led to significantly poorer hand motor recovery after stroke compared to EEG- triggered robotic feedback, both immediately after and 6 months after the intervention. Note that all of the above studies were conducted using non-invasive EEG-based BCIs. In addition to delivering proprioceptive feedback, BCI can also control functional electrical stimulation (FES) systems as another potential means to drive neuronal plasticity processes (see Fig. 22.3). Using FES with MI- and ME-based BCIs not only activates afferent sensory pathways but also lower motor neurons. Compared to proprioceptive feedback alone, this mechanism may further enhance neural plastic changes, especially in sensorimotor areas [74, 76, 88, 89]. Specifically, the coincident activation of upper and lower motor neurons may induce Hebbian learning via long-term potentiation at their synapse in the spinal cord [89, 90]. Initial evidence from Hara et al. [91] supports the use of therapies that coactivate upper and lower motor neurons. Here, it is suggested that an EMGcontrolled FES therapy, in which motor intention is coupled with FES, may be more beneficial than either attempted movement or FES alone. Several preliminary studies have demonstrated the feasibility of BCI-FES systems for physiotherapy. Daly et al. [60] combined BCI and FES for motor learning in a single chronic stroke patient who was unable to perform isolated finger movements. Visual cues were provided to the subject to relax or move her paretic fingers, and ME- or MI-based motor intention (via the BCI) triggered FES-induced index finger extension. After a small number of training sessions, volitional motor control over the index finger was obtained. McCrimmon et al. [81] utilized a similar paradigm for lower extremity rehabilitation. Nine chronic stroke subjects with foot drop each participated in 12 sessions, in which they followed visual cues and attempted either ankle dorsiflexion or relaxation, while FES was either supplied or withheld, respectively.

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Gait function improved in several subjects (details discussed further below). More recently, Lee et al. [92] showed that BCI-FES promoted significant changes in sensorimotor rhythms and functional recovery of the upper extremities after stroke compared to conventional physical therapy. Chung et al. [93] showed that BCI-FES promoted significantly increased functional recovery of gait compared to FES alone (where both therapies were given over an identical duration). In addition, Biasiucci et al. [94] demonstrated that sham-FES, which avoids coactivation of upper and lower motor neurons, is significantly worse than BCI-FES at promoting recovery, and may even promote maladaptive behavior given a decrease in the Modified Ashworth Scale from baseline in these subjects.

22.4.2 BCI-Based Stroke Physiotherapy in Clinical Applications Recently, there have been an increasing number of interventional clinical studies that examined the feasibility and efficacy of BCI-based therapies for stroke rehabilitation. At the time of this review, clinical trials in various stages of completion have been reported. Studied modalities include pairing EEG-based BCI with robotic orthoses [69, 86, 87, 95–99], exoskeletons [100], FES [81, 93, 95, 101–103], repetitive transcranial magnetic stimulation (rTMS) [104], animations on a computer screen [80, 105], or virtual reality [104], transcranial direct current stimulation Table 22.1 Samples of Phase 0/I (safety and feasibility) and noncontrolled studies with at least five participants. Studies already in Table 22.2 are not listed again here. NMES = Neuromuscular electrical stimulation

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(tDCS) [82, 106] or action observation therapy (AOT) [92, 107] for feedback or treatment. Targeted body areas typically are the upper extremities (e.g., finger/wrist extension and upper arm movements) and lower extremities (e.g., ankle dorsiflexion and gait). For these studies, outcome measures typically include Fugl–Meyer Assessment (FMA), Action Research Arm Test (ARAT), Barthel Index (BI, MBI), Wolf Motor Function Test (WMFT), various strength and range of motion (ROM) assessments, various brain connectivity outcomes, and various patient-reported outcome measures. A representative list of Phases 0 and I (safety and feasibility) trials can be found in Table 22.1. At this stage of a clinical trial, a small number of participants (both stroke and able-bodied) are recruited to ensure that the BCI device and regimen produce no harmful effects and that improvement of a measurable outcome is at least feasible. Results from Phases 0 and I trials are typically used to estimate the effect size and refine the study protocol when designing larger confirmatory clinical trials. A representative list of Phase II (efficacy) randomized controlled trials (RCT) can be found in Table 22.2. This list only includes studies that have at least 20 participants, and only includes stroke subjects (studies involving able-bodied participants as control were excluded). Studies split participants into experimental and control arms in various ways, such as BCI versus conventional physiotherapy, or BCI-controlled robotic orthosis versus conventionally

Study

Feedback

Targeted Limb

Kawakami et al. [95]

Robotic orthosis and NMES

Hand

Cantillo-Negrete et al. [96]

Robotic orthosis

Hand

McCrimmon et al. [81]

FES

Foot

Jang et al. [101]

FES

Shoulder

Sebastian-Romagosa et al. [103]

FES

Hand

Hong et al. [106]

tDCS and MIT-Manus

Hand

Hu et al. [82]

tDCS and MIT-Manus

Hand

Prasad et al. [80]

Computer screen

Hand

518

J. Lim et al.

Table 22.2 Some published Phase II (efficacy) randomized controlled clinical studies testing the use of closed-loop BCI for stroke rehabilitation. ‘nmFU’ = n-month Follow-up. FMA = Fugl–Meyer Motor Assessment. ‘*’ denotes a significant efficacious outcome for BCI. ‘**’ denotes significant BCI outperformance over conventional therapy methods. For this claim to be valid, the trial must test the BCI treatment group against a non-BCI treatment group for equal treatment time, such as a 1-h BCI and 1-h non-BCI versus 2-h non-BCI. ‘†’ denotes an ongoing clinical trial with incomplete data analysis. MAL = Motor activity log. MBI = Modified Barthel index. ROM = range of motion. WMFT = Wolf motor function test. MI = Motricity Index of the arm. Dyn = Dynamometer. SIS = Stroke impact scale Study

Feedback

Targeted Limb

Sample Size

Primary Outcome Measures

Ang et al. [86]

Haptic knob robot

Hand and wrist

21

FMA**

Ang et al. [97]

MIT-Manus robot

Elbow and forearm

26

FMA*

Frolov et al. [98]

Hand exoskeleton

Hand

74

FMA*, ARAT**

Lyukmanov et al. [100]

Hand exoskeleton

Hand

55

FMA*, ARAT**

Mizuno et al. [99]

Robotic orthosis

Hand

40†

FMA

Ramos-Murguialday et al. [69]

Robotic orthosis

Hand

32

FMA*

Ramos-Murguialday et al. [87]

Robotic orthosis

Hand

30 (6mFU)

FMA*

Remsik et al. [102]

FES

Hand

21

ARAT**

Chung et al. [93]

FES

Foot

25

Gait velocity**, Cadence**

Kim et al. [107]

AOT and FES

Hand and wrist

30

FMA*, MAL*, MBI*, ROM*

Lee et al. [92]

AOT and FES

Hand and wrist

26

FMA**, WMFT**, MAL*, MBI*

Sanchez-Cuesta et al. [104]

VR and rTMS

Hand, wrist, arm

42†

MI, FMA, Dyn, SIS

Mattia et al. [105]

Computer screen

Hand

48†

FMA

controlled robotic orthosis. A BCI therapy regimen would be considered efficacious if its primary outcome measure is at least comparable to conventional physiotherapy. All BCIs reported in Table 22.2, other than those in ongoing trials, are efficacious. In addition, BCI was significantly better than conventional physiotherapy for at least one of the outcome measures in six controlled studies [86, 92, 93, 98, 100, 102]. Also, Ramos-Murguialday et al. [87] report that participants treated with BCI-controlled robotic orthosis retained positive outcomes at a similar level to those in the conventional physiotherapy control group after 6 months. Finally, no Phase III (effectiveness) trials or studies involving invasive brain signal acquisition for rehabilitation purposes have been reported at the time of this review. It should be noted that some invasive trials exist for prosthetic applications as discussed in the section above. Given the absence of any definitive Phase III clinical trials regarding the effectiveness of BCIbased physiotherapy, it is currently not possible

to determine whether any of these approaches can be recommended for the stroke population at large. There are several factors that may contribute to why there are still no Phase III studies for BCI therapy despite the multitude of successful Phase II trials (Table 22.2). These include the prohibitive cost of the BCI systems as well as extensive training and setup time. Although not prohibitive, a lack of understanding of the mechanisms underlying BCI therapy may also hinder optimal delivery of the BCI therapy. Several studies have received funding from national agencies such as the National Institute of Health and the National Science Foundation in the United States (Table 22.3) to address these outstanding concerns. Some aim to simplify the hardware and regimen for at-home rehabilitation [108, 109] and long-term use [109]. Some studies aim to elucidate the underlying mechanism of any improvements seen in stroke patients who underwent BCI-based physiotherapies [110]. In summary, the current BCI clinical trial literature provides evidence that BCI-based

22

BCI-Based Neuroprostheses and Physiotherapies for Stroke …

519

Table 22.3 Some ongoing BCI-based stroke rehabilitation studies currently funded by national agencies in the United States. NIH = National Institute of Health. VA = Veteran Affairs. NSF = National Science Foundation Principal investigator

Targeted areas

Active years and funding agency

K Bhugra (Neurolutions Inc.) [108]

At home, hand rehabilitation

2021–2022 by NIH

DJ Lin (Providence VA) [111]

Arm rehabilitation

2020–2022 by VA

A Do (Univ. of California Irvine) [110]

Gait rehabilitation with FES

2019–2024 by NIH

J Contreras-Vidal (Univ. of Houston) [109]

Long-term, at-home, robotassisted

2018–2022 by NSF

physiotherapies may be safe and promising enough to warrant large-scale clinical investigations. Simultaneously, fundamental questions still need to be answered, such as understanding the mechanism of action and how brain physiology change over the course of treatment. Furthermore, practical questions will need to be answered. Are there particular characteristics of stroke patients who will respond best to BCIbased physiotherapy? What kind of closed-loop feedback is the best? How can BCI be incorporated into at-home rehabilitation? The answers to these questions have implications for the future justification of BCI-based therapies, particularly if existing dose-matched therapies turn out to be cheaper and just as efficacious. In addition to the above scientific questions, many practical issues related to the implementation of BCI-based physiotherapies need to be addressed. Since these therapies currently require extensive setup, it is unclear how they will be efficiently and effectively delivered in clinical practice and at home. Are there ways to drastically reduce the setup time of such BCI systems? Additionally, will BCI-based therapies be provided by the physical and occupational therapists in the community? Will they be time and resource-efficient? Will the associated equipment and training costs be acceptable to practicing clinicians? Will patients be interested in such therapies? Will medical insurance providers reimburse or support their clinical use? Since this research field is still in early development, the medical device industry has yet to streamline these systems. As the field matures, it can be expected that BCI-based physiotherapy studies will transition towards large clinical trials. At that

time, significant research and development must be performed to understand and address these market issues that may ultimately affect the success of BCI-based physiotherapies.

22.5

Conclusion and Future Directions

In recent years, BCIs have garnered increasing interest as a means of substituting for lost motor functions or for improving post-stroke motor outcomes. Neuroprosthetic BCIs have been designed primarily for SCI, but can be extended to post-stroke paralysis. Significant engineering challenges must still be overcome before these systems can be used in the clinic in a robust and practical manner. BCI systems for physiotherapy may be applied as a novel means of facilitating Hebbian learning mechanisms, which can be elicited by two major strategies. One involves providing BCI-controlled proprioceptive sensory feedback to upregulate the connection between sensory and motor cortices and subsequently cause increased motor output to the lower motor neurons. The second strategy employs BCIcontrolled electrical stimulation to simultaneously activate the post-stroke motor areas and the lower motor neurons, thereby increasing their connectivity over time. Both of these strategies can potentially promote motor recovery. In fact, these strategies have already been realized in BCI-controlled robotic and FES therapies. Existing early phase clinical trials suggest that these strategies are promising. However, definitive clinical trials still need to be performed, and many questions still remain regarding the safety

520

and efficacy of BCI-based physiotherapies and whether they can be practically applied in the clinical setting.

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and cortical perfusion in stroke patients. Clin Neurophysiol. 2013;124(10):2008–15. Lee SH, Kim SS, Lee BH. Action observation training and brain-computer interface controlled functional electrical stimulation enhance upper extremity performance and cortical activation in patients with stroke: a randomized controlled trial. Physiother Theory Pract. 2020;1–9. Available from: https://www.tandfonline.com/doi/full/10.1080/0959 3985.2020.1831114. Chung E, Lee BH, Hwang S. Therapeutic effects of brain-computer interface-controlled functional electrical stimulation training on balance and gait performance for stroke: a pilot randomized controlled trial. Medicine. 2020;99(51):e22612. Available from: https://journals.lww.com/10.1097/MD. 0000000000022612. Biasiucci A, Leeb R, Iturrate I, Perdikis S, AlKhodairy A, Corbet T, et al. Brain-actuated functional electrical stimulation elicits lasting arm motor recovery after stroke. Nat Commun. 2018;9 (1):2421. Available from: http://www.nature.com/ articles/s41467-018-04673-z. Kawakami M, Fujiwara T, Ushiba J, Nishimoto A, Abe K, Honaga K, et al. A new therapeutic application of brain-machine interface (BMI) training followed by hybrid assistive neuromuscular dynamic stimulation (HANDS) therapy for patients with severe hemiparetic stroke: a proof of concept study. Restor Neurol Neurosci. 2016;34(5):789–97. Cantillo-Negrete J, Carino-Escobar RI, CarrilloMora P, Rodriguez-Barragan MA, HernandezArenas C, Quinzan˜os-Fresnedo J, et al. Braincomputer interface coupled to a robotic hand orthosis for stroke patients’ neurorehabilitation: a crossover feasibility study. Front Hum Neurosci. 2021;15:293. Available from: https://www. frontiersin.org/article/10.3389/fnhum.2021.656975. Ang KK, Chua KSG, Phua KS, Wang C, Chin ZY, Kuah CWK, et al. A randomized controlled trial of eeg-based motor imagery brain-computer interface robotic rehabilitation for stroke. Clin EEG Neurosci. 2015;46(4):310–20. Available from: http:// journals.sagepub.com/doi/10.1177/1550059414522229. Frolov AA, Mokienko O, Lyukmanov R, Biryukova E, Kotov S, Turbina L, et al. Poststroke rehabilitation training with a motor-imagerybased brain-computer interface (BCI)-controlled hand exoskeleton: a randomized controlled multicenter trial. Front Neurosci. 2017;11:400. Mizuno K, Abe T, Ushiba J, Kawakami M, Ohwa T, Hagimura K, et al. Evaluating the effectiveness and safety of the electroencephalogram-based brain-machine interface rehabilitation system for patients with severe hemiparetic stroke: protocol for a randomized controlled trial (BEST-BRAIN trial). JMIR Res Protoc. 2018;7(12):e12339. Available from: https:// www.researchprotocols.org/2018/12/e12339.

524 100. Lyukmanov RK, Aziatskaya GA, Mokienko OA, Varako NA, Kovyazina MS, Suponeva NA, et al. Post-stroke rehabilitation training with a braincomputer interface: a clinical and neuropsychological study. Zhurnal nevrologii i psikhiatrii im SS Korsakova. 2018;118(8):43. Available from: http:// www.mediasphera.ru/issues/zhurnal-nevrologii-ipsikhiatrii-im-s-s-korsakova/2018/8/downloads/ru/ 1199772982018081043. 101. Jang YY, Kim TH, Lee BH. Effects of braincomputer interface-controlled functional electrical stimulation training on shoulder subluxation for patients with stroke: a randomized controlled trial: bci-controlled fes improved shoulder subluxation in stroke survivors. Occup Ther Int. 2016;23(2):175– 85. Available from: https://onlinelibrary.wiley.com/ doi/10.1002/oti.1422. 102. Remsik AB, Dodd K, Williams L, Thoma J, Jacobson T, Allen JD, et al. Behavioral outcomes following brain computer interface intervention for upper extremity rehabilitation in stroke: a randomized controlled trial. Front Neurosci. 2018;12:752. Available from: https://www.frontiersin.org/article/ 10.3389/fnins.2018.00752/full. 103. Sebastia´n-Romagosa M, Cho W, Ortner R, Murovec N, Von Oertzen T, Kamada K, et al. Brain computer interface treatment for motor rehabilitation of upper extremity of stroke patients–a feasibility study. Front Neurosci. 2020;14:591435. Available from: https://www.frontiersin.org/articles/ 10.3389/fnins.2020.591435/full. 104. Sa´nchez-Cuesta FJ, Arroyo-Ferrer A, Gonza´lezZamorano Y, Vourvopoulos A, Badia SBi, Figuereido P, et al. Clinical effects of immersive multimodal BCI-VR training after bilateral neuromodulation with rtms on upper limb motor recovery after stroke. A study protocol for a randomized controlled trial. Medicina. 2021;57(8):736. Available from: https://www.mdpi.com/1648-9144/57/8/736.

J. Lim et al. 105. Mattia D, Pichiorri F, Colamarino E, Masciullo M, Morone G, Toppi J, et al. The promotoer, a braincomputer interface-assisted intervention to promote upper limb functional motor recovery after stroke: a study protocol for a randomized controlled trial to test early and long-term efficacy and to identify determinants of response. BMC Neurol. 2020;20(1):254. Available from: https://bmcneurol.biomedcentral.com/ articles/10.1186/s12883-020-01826-w. 106. Hong X, Lu ZK, Teh I, Nasrallah FA, Teo WP, Ang KK, et al. Brain plasticity following MI-BCI training combined with tDCS in a randomized trial in chronic subcortical stroke subjects: a preliminary study. Sci Rep. 2017;7(1):9222. Available from: http://www.nature.com/articles/s41598-017-08928-5. 107. Kim T, Kim S, Lee B. Effects of action observational training plus brain-computer interface-based functional electrical stimulation on paretic arm motor recovery in patient with stroke: a randomized controlled trial: effects of aot plus BCI-FES on arm motor recovery. Occup Ther Int. 2016;23(1):39–47. Available from: https://onlinelibrary.wiley.com/doi/ 10.1002/oti.1403. 108. Bhugra K. A fully remote telehealth brain computer interface and assessment system for motor rehabilitation of chronic stroke; 2021. https://reporter.nih. gov/. 109. Contreras-Vidal J, Feng J, Shedd B. PFI-RP: braincontrolled upper-limb robot-assisted rehabilitation device for stroke survivors; 2018. https://www.nsf. gov/awardsearch. 110. Do A, Cramer SC, Nenadic Z, Reinkensmeyer DJ. Brain-computer interface-functional electrical stimulation for stroke recovery; 2019. https:// reporter.nih.gov/. 111. Lin DJ. Targeting neuroplasticity with brain computer interfaces to maximize motor recovery for veterans with stroke; 2020. https://reporter.nih.gov/.

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Passive Devices for Upper Limb Training Marika Demers, Justin Rowe, and Arthur Prochazka

Abstract

Arm and hand motor impairments are frequent after a neurological injury. Motor rehabilitation can improve hand and arm function in many cases, but in the current healthcare climate, the time and resources devoted to physical and occupational therapy after injury are inadequate. This represents an opportunity for technology to be introduced that can complement rehabilitation practices, provide motivating task training and allow remote supervision of exercise training performed in the home. Over the last decades, many research groups have been developing robotic devices for exercise therapy, as well as other methods such as electrical stimulation of muscles or vagus nerve stimulation. Robotic devices tend to be expensive and recent studies have raised some doubt as to whether assistance to movements is always preferable as it

M. Demers (&) Division of Biokinesiology and Physical Therapy, University of Southern California, Los Angeles, CA, USA e-mail: [email protected] J. Rowe Flint Rehabilitation Devices, Irvine, CA, USA e-mail: [email protected] A. Prochazka University of Alberta, Edmonton, AB, Canada e-mail: [email protected]

can reduce salience and engagement. This chapter reviews the evidence for spontaneous recovery, the means and mechanisms of conventional rehabilitation interventions, the advent of affordable passive devices and other treatment modalities that can be used in combination with passive devices. It is argued that task practice on passive devices, in some cases remotely supervised over the internet or augmented with functional electrical stimulation (FES), is now an affordable and important modality of occupational and physical therapy. Passive devices offer numerous opportunities in the field of neurological rehabilitation to support arm and hand motor recovery. Keywords

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Stroke Spinal cord injury Multiple sclerosis Upper extremity Tele-rehabilitation Health technology

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Introduction

Neurological disorders are a leading cause of a disability worldwide [1]. One frequent impairment after a neurological insult is a deficit in movements of the arms and hands. This can range from paresis (weakness) to paralysis (plegia). Arm and hand paresis is characterized by muscle weakness, changed muscle tone, decreased sensation, and impaired voluntary

© The Author(s), under exclusive license to Springer Nature Switzerland AG 2022 D. J. Reinkensmeyer et al. (eds.), Neurorehabilitation Technology, https://doi.org/10.1007/978-3-031-08995-4_23

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movement control resulting in slow, imprecise, and uncoordinated movement [2–4]. Among stroke survivors which are estimated to represent 7.0 million individuals in the United-States (2.5% of the population) [5], approximately 65% experience residual arm motor impairments despite intensive and prolonged rehabilitation [6]. Unilateral or bilateral arm motor impairments impact approximately 30% of individuals with traumatic brain injury [7], 60% of individuals in the first year following the diagnosis of multiple sclerosis [8], and individuals with cervical spinal cord injury, the most common site of spinal cord trauma [9, 10]. Due to the important contribution of the arm and hand to everyday activities, motor impairments often lead to activity limitations and participation restriction [11–14]. Rehabilitation interventions can help remediate arm motor impairments and regain lost function. Common interventions include taskoriented training and repetitive task practice, constraint-induced movement therapy, mental imagery, mirror therapy and virtual reality [15]. However, many health system constraints limit the delivery of neurological arm and hand rehabilitation. The decreasing rehabilitation length of stay, general lack of reimbursement for therapy, disparities in access to rehabilitation care, limited available treatment time, and competing rehabilitation priorities, such as the early focus on improving lower limb mobility and gait, are examples of frequent challenges with neurological rehabilitation [16–18]. Therefore, people with neurological disorders may not receive adequate rehabilitation interventions for arm motor recovery. Moreover, people with neurological disorders are often discharged home with limited opportunities to continue home exercises and engage in evidence-based interventions to drive recovery after therapy has ended. At home, longterm adherence to exercise programs is often low [19], due to low motivation, cognitive impairments, lack of caregiver support, frustration, pain, and musculoskeletal issues [20]. This represents an opportunity for technology to be

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introduced to address key challenges to neurological rehabilitation. Passive devices offer an affordable option to continue rehabilitation outside clinical settings. Passive rehabilitation devices are defined by what they do not do. Unlike robotic rehabilitation devices, passive devices do not provide active assistance during motor rehabilitation. Although passive devices do not use actuators such as electric motors or pneumatic cylinders, some may provide postural support using energy storage devices such as springs or moving masses. Others provide no assistance, but rather simplify exercises, making them more approachable by removing degrees of freedom from the task. Still others create a motivational environment to support motor practice. The term ‘passive devices’ should be distinguished from the term ‘passive’ by rehabilitation clinicians to describe exercises or movements made without efforts from the patient. This chapter reviews the mechanisms for functional recovery, the means and mechanisms of conventional rehabilitation interventions, the rationale behind passive devices, and current evidence supporting different types of passive devices.

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Mechanisms of Functional Recovery: The Significance of Compensatory Strategies

Spontaneous mechanisms of recovery at the cellular, molecular, and systems levels often follow a neurological insult. However, the degree of spontaneous recovery varies between individuals and neurological conditions, and is generally incomplete [21]. Some of the spontaneous recovery in motor function is evidently a result of the recovery of central nervous structures temporarily inactivated by the injury, or the adaptation of uninjured nervous networks to take over functions of neighboring injured networks, a process called plasticity [22–24]. Various means of early prediction of the extent of recovery have been identified [25–28]. In this regard, the

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concept of compensation should be distinguished from recovery [29]. Specifically, individuals with neurological disorders may adopt compensatory strategies to accomplish daily tasks, such as the use of alternative motor strategies (i.e., shoulder elevation or trunk flexion) to compensate for lost motor patterns in the elbow and shoulder [3]. At the body function and structure level of the International Classification of Functioning (ICF) [30], compensation is defined as the performance of a movement in a new manner, seen as the appearance of alternative movement patterns during the accomplishment of a task. At the activity level, compensation refers to a successful task completion using different techniques [29]. The adoption of compensatory strategies may be considered maladaptive if compensations limit recovery of independent movements of the most affected arm, contribute to secondary complications such as pain, joint contracture, and discomfort [31], and lead to a pattern of learned maladaptive behavior impeding long-term functional motor recovery [32, 33]. The use of maladaptive compensatory strategies may limit one’s ability to generalize movements to a wider array of tasks [34] and contribute to incipient decline after the end of active therapy [35]. Evidence supports the effectiveness of neurological rehabilitation to remediate arm motor impairments and improve function by fostering neuroplastic changes, which often rely on mechanisms similar to those observed during spontaneous recovery [15, 21, 34, 36, 37]. An understanding of the mechanisms through which recovery is achieved after a neurological disorder is important to guide effective rehabilitation interventions. Neuronal plasticity plays a crucial role in neurologic recovery. From the work done in animal models, repetition, intensity, and salience have been identified as critical to drive experience-dependent neuroplasticity [33]. Other factors influencing plasticity include the provision of progressive and optimally adapted rehabilitation interventions tailored to one’s capability and the environmental context [38, 39]. In animal stroke models, hundreds of

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repetitions of motor tasks are needed to induce lasting neural changes [24, 40]. In humans, the critical threshold of rehabilitation intensity needed to engage plastic mechanisms is unknown. A recent review concluded there was a positive relationship between the time scheduled for exercise therapy and the outcome with large doses of exercise therapy leading to clinically meaningful improvements [41]. The authors pointed out that time scheduled did not necessarily equate to the amount of task practice actually performed. They recommended that instead of reporting scheduled time, future studies should report active time in therapy or repetitions of an exercise. The notion that more is better was challenged in recent large, randomized control trials (RCT). Specifically, Winstein et al. [42] compared four dosages of personalized arm task-oriented training in chronic stroke survivors. Higher dosage of training led to greater gains in quality of arm use measured with the Motor Activity Log, but no changes on functional capacity measured with the Wolf Motor Function Test were noted. In another RCT, Lang et al. [43] compared four doses of task-specific training on arm and hand function (i.e., 3,200, 6,400, 9,600, or individualized maximum repetitions). The results showed no evidence of a dose–response effect of task-specific training on arm and hand functional capacity in stroke survivors. However, the number of movement repetitions provided during this trial was far superior to the amount of movement practice normally provided during conventional stroke rehabilitation [44]. Similarly, the most common dosage of home exercises prescribed for adults with neurological conditions is estimated to be 16–30 min per day with a greater focus on fine motor activities, active range of motion, active assistive range of motion, and whole or partial activity of daily living tasks [45]. Despite the lack of consensus on optimal dose, interventions for arm and hand motor impairments may not always be delivered at the most beneficial intensities and may not include enough repetitions to optimize neuroplasticity [46].

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The Role of Rehabilitation to Restore Arm and Hand Function

A variety of rehabilitation interventions can be used to improve arm and hand motor recovery, but the level of evidence available varies depending on the modality and by neurological populations. The Bobath technique and proprioceptive neuromuscular facilitation, two rehabilitation approaches based on neurophysiological principles, were widely adopted in the 1970s with strong adherents in each camp. However, current evidence does not support the superiority of neurodevelopmental techniques over other types of interventions [16]. A RCT that compared these two approaches with conventional rehabilitation concluded that there were no significant betweengroup differences in improvement of the patients’ performance of activities of daily living [47]. Looking specifically at evidence-based interventions for different neurological populations, task-oriented training, resistance and endurance training, constraint-induced movement therapy, and some types of robot-supported training may improve arm and hand function in individuals with multiple sclerosis [36]. However, the evidence for one approach over another is not clear due to large variability between studies and small sample sizes. Similarly, motor training in individuals with cervical spinal cord injury or traumatic brain injury, which can include task practice and functional electrical stimulation (FES), may reduce arm and hand motor impairments and improve function. However, there were wide differences between studies in the types of patients, training, methodology and outcome parameters [37, 48, 49]. For individuals with stroke, a meta-study concluded that sensorimotor training, motor learning training with the use of imagery, electrical stimulation, and the repetitive performance of novel tasks, could all be effective in reducing motor impairment after stroke [16]. Moderate-quality evidence suggests that constraint-induced movement therapy, mental practice, mirror therapy, interventions for sensory impairment, virtual reality and task-

specific training may be effective in improving arm and hand function [15, 50]. One of the most investigated interventions to remediate arm and hand motor impairments in neurology is constraint-induced movement therapy (CIMT), a particular form of intensive and supervised task practice [51]. This approach was developed based on experiments in monkeys in which sensory input in one arm was abolished by de-afferentation. Binding of the other, nonaffected arm, led to forced use of the deafferented arm, which was associated with improvements in its motor function [52, 53] and brain reorganization [54, 55]. In humans, CIMT or the modified versions of CIMT include three key components: (1) constraint of the less affected arm for up to 90% of the waking hours to promote the use of the more affected arm, (2) delivery of intensive graded practice of the more affected arm in functionally meaningful tasks (i.e. task shaping), (3) adherence-enhancing behavioral methods designed to transfer the gains to the real-world environment (i.e., transfer package) [56–58]. CIMT and modified CIMT have been shown to have robust, clinically meaningful impacts on stroke survivors’ outcomes for arm and hand motor impairment and function, making CIMT one of the most effective interventions for the paretic arm post stroke [59– 61]. However, the use of CIMT in clinical practice is limited. In a survey of 92 therapists working in clinical neurorehabilitation in the USA, 75% reported that it would be difficult or very difficult to administer CIMT in their clinics [62]. Challenges with the delivery of CIMT or modified CIMT include the difficulty to achieve the recommended training intensity, stringent inclusion criteria, lack of resources and high delivery costs that may not be reimbursed by third-party payers [62–65].

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Robotic Exercise Devices

Monitoring movement performance and quality is often challenging in clinical practice, as common objects used for task practice (e.g., blocks,

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cones, therapy putty, peg boards, resistive prehension benches, etc.) do not have sensors to quantify movement kinematics. The supervision by a therapist can prevent the use of maladaptive compensatory movements. However, the supervision by therapists is costly, and in most cases, restricted to clinics, which in turn limits access mainly to subacute patients. Robotic devices have been developed with the aim of delivering high-intensity, repetitive and adaptive training [66]. Robotic devices can be used to provide standardized exercises, take over some supervisory functions and provide quantitative outcome measures to reduce the tedium of conventional rehabilitation. The treatment paradigm for robotics is based on the provision of physical assistance to complete desired motions of the arm and hand in combination with computer games or virtual reality presented on a screen [67]. Robotic assistance is considered advantageous because it (1) allows the challenge level of the task to be adjusted to better suit the needs and abilities of the user, and (2) allows users to move through a larger range of motion—thereby providing a large afferent response that is time correlated with the user’s efferent motor intent. Evidence from systematic reviews and meta-analyses support the use of robotic-based training for sensorimotor rehabilitation of stroke survivors and people with spinal cord injury and multiple sclerosis [36, 66, 68–71]. Research on robot-assisted movement therapy has rapidly increased in recent years, as the potential for robotic therapy after a neurological insult remains enormous [67]. Some of the drawbacks of robotic devices are the high cost of the devices and the lack of salience of the tasks (i.e., tasks may not be meaningful or engaging to the users). Robotic devices incorporate actuators and complex control systems, which makes them expensive. Retail prices can vary between $400USD per month for a powered wrist splint (e.g., Hand Mentor, Motus Nova, Atlanta, GA, https://motusnova.com/hand/) to tens of thousands of dollars for exoskeleton robots, such as the KINARM Exoskeleton Robot (Kinarm, Kingston, Ontario, https://kinarm.com/) or the Armeo Power (Hocoma, Volketswil, Switzerland, https:// www.hocoma.com/us/solutions/armeo-power/).

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The high price point is a barrier to adoption in clinical practice and home use. Another criticism towards robotic devices is that the tasks may not be meaningful to the users, solicit intrinsic motivation or active participation. Ideally, robotic assistance would allow users to practice at their ideal challenge rate, allowing users to be motivated by their success and learn from their occasional failures [72]. However, over-assisting can be counterproductive [73]. Although much work has been devoted to controllers that assist only as needed, the problem of finding the ideal assistance level remains challenging. Since movements are often restricted to a 2D-plane, robotic devices are limited in the variety of movements or tasks practiced, which may not reflect the range of arm and hand movements used for everyday object interaction. The high-cost and the lack of salience of robotic devices, along with the aforementioned challenges with the delivery of neurorehabilitation stress the need for affordable and effective solutions to harness neuroplasticity. Passive devices are more affordable than robotic devices and unlike CIMT, which has stringent inclusion criteria, they are accessible to individuals with a wide variety of impairments. They can provide motivating task practice to minimize the lack of adherence with home programs and offer the opportunity to deliver salient and intensive task practice outside clinical settings. Because passive devices are defined by what they are not, the remaining field is understandably broad and inclusive. In this chapter, we divide the field of passive devices into 5 groups: (1) passive gravity support systems, (2) tabletop therapy systems, (3) linear rail systems, (4) tone compensating orthoses, and (5) serious game controllers (Fig. 23.1).

23.5

Passive Gravity Support Systems

Mobile arm supports and balanced forearm orthoses have been a part of rehabilitation practice since at least the mid 1960s [74–77]. Although the embodiment has varied, most are either chair-mounted or desk-mounted orthotics

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Fig. 23.1 a 5 types of passive devices for neurorehabilitation; b rehabtronics hand stimulator activated by voluntary head-nods detected by a wireless earpiece

that support the arm at the wrist using either a rigid planar linkage [78, 79], or in many cases, a spring loaded mechanism tuned to balance the weight of the arm. Initially, mobile arm supports and balanced forearm orthoses were used as assistive devices and evaluated based on whether they enabled users with motor impairments to perform activities of daily living that they could not perform otherwise [80]. Early devices were also limited in their degrees of freedom and in the range of motion that they could support [81]. Although chair- or desk-mounted mobile arm support assistive devices have continued to evolve [82], much attention has been given to a newer class of devices designed specifically for rehabilitation. The Armeo Spring (based on the T-WREX, Hocoma, Volketswil, Switzerland) [83], is a counter-balanced multi-segment arm support with six instrumented and lockable degrees of freedom and an instrumented gripper. Positional data from the joint sensor and force data from the gripper are used both to (1) control a suite of games simulating activities of daily living and (2) to quantify features of the user’s motor impairment. In early testing, participants with chronic stroke that trained with T-WREX improved their motor capacity (as measured by the Fugl Meyer Assessment) significantly more than participants that trained with tabletop exercises [81]. The gains themselves were modest, but the T-WREX group also showed significantly better retention. Numerous studies have demonstrated that training in T-WREX and Armeo

Spring increases the range of motion—both immediately, while in the orthosis, and to a lesser extent, persistently [81, 82, 84]. Critically, participants prefer training with gravity assistance and assign high value to the exercise [85]. The efficacy of Armeo Spring for subacute recovery is less clear. Recently, a large clinical trial in subacute participants compared therapy in Armeo Spring to dose-matched stretching and basic active exercises. The Fugl-Meyer scores of both groups increased significantly, but the differences between the groups was not significant at either the 4-week assessment or the 12-month follow-up [86]. Although the multi-segment linkage used by Armeo Spring does an impressive job of supporting the weight of the arm without placing unwanted restrictions on its range of motion, adjusting the linkages and the springs for each user can be time consuming. Freebal (sold by Hocoma under the name Armeo Boom, Fig. 23.2), by contrast, is a sling-based gravity support system that does not provide as much freedom of movement as Armeo Spring but is much simpler and requires less adjustment. Like Armeo Spring, Freebal extends the range of motion in which users are able to train [87–90]. Although the effectiveness of Freebal has not been studied as extensively as Armeo Spring, early pilot testing suggested that it had similar and lasting effects on motor capacity and range of motion [90]. The high cost of both devices remains a barrier for clinical or home use.

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Fig. 23.2 The Armeo®Booom, a passive gravity support system

Armeo Spring and Armeo Boom are both stationary devices. There is also a developing class of wearable gravity support exoskeletons that offload the weight of the arm but are not constrained to a particular location [91–93]. The recently developed SpringWear system, for example, is lightweight, and increases the active workspace of its users. However, the exoskeleton did not consistently improve the wearer’s ability to complete functional tasks [91, 94].

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Tabletop Therapy Devices

Tabletop therapy systems are similar to passive gravity support systems in that they allow users to practice without supporting the weight of their arm. However, instead of using springs to compensate for gravity, tabletop systems rely on the surface of the table to support the weight. This simplification restricts movements to a single

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plane, but it makes the devices more cost effective and appropriate for home use. Normally these systems include some mechanism for reducing frictional forces between the arm and the support service such as omnidirectional wheels [95], or a 2D gantry [96]. Like the nonplanar gravity support systems, most tabletop therapy systems include some mechanisms (e.g. cameras [95], encoders [97], or instrumented tracks [96]) to monitor the position of the arm and couple the movements with engaging serious games. Some systems allow users to increase the difficulty of their exercises by tilting the table [98, 99] or by introducing friction [97]. The Rutgers Arm II is noteworthy for detecting unwanted compensatory movements from the shoulder [98], and the Rapael Smart Board (Neofect, San Francisco, CA) is noteworthy for being a commercially available device. Evidence of the effectiveness of tabletop therapy systems is somewhat limited. Most devices are validated using uncontrolled pilot studies, which indicate that the devices are safe, promising, and well received by their users [95, 98, 100]. The main exception is the Rapael Smart Board which was validated by a RCT for chronic stroke survivors. Participants that practiced with the Smart Board in addition to standard care increased their Wolf Motor Function scores significantly more than participants that practiced with a double dose of standard care [96].

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Linear Track System

In much the same way that tabletop therapy devices are a simpler, more constrained alternative to systems like Armeo Spring or Armeo Boom, linear track systems are a simpler and more constrained alternative to tabletop therapy systems. Linear tracks support gravity and reduce unwanted friction, but they also resist unintentional movements caused by imbalanced motor synergies. As long as there is some forward or backward component to the forces that the user applies to the slider, it will progress along the track, making the exercise very forgiving. The three most prominent linear track systems are the

BATRAC [101] which was sold for a time under the trade name “Tailwind”, the Reha-Slide [102], and the SMART Arm [103]. Both the BATRAC and the Reha-Slide promote bimanual exercises. All three devices include game-like elements, but the SMART Arm is the only device built around an actual computer gaming system. The Rehaslide has been used as an input to a gaming system with a backend for telerehabilitation, but the commercial version does not yet support these features [104]. Training with any of the three devices has been shown to significantly improve motor capacity in chronic stroke, but none of the devices have proven to be significantly better than dose-matched conventional treatment [105–110].

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Tone Compensating Orthoses

In much the same way that Armeo Spring uses springs and passive mechanisms to cancel the effects of gravity, tone compensating orthoses like the Hand Spring Operated Movement Enhancer (HandSOME), SCRIPT, and EXTEND devices use springs and carefully designed mechanisms to create tunable force profiles that do an impressive job of compensating for unwanted joint torques caused by wrist and finger flexor hypertonia [111–113]. While wearing the HandSOME orthosis, stroke survivors with finger flexor hypertonia were able to both move through a significantly larger range of motion and outperform their unassisted Box and Blocks scores [111]. Similar results were observed for the EXTEND and SCRIPT Orthoses [112, 114]. The stated goal of all three devices is to enable stroke survivors to exercise more and with greater success. However, the efficacy of exercising with a tone-compensating device is far from clear. In an uncontrolled pilot study, the training with the HandSOME orthosis was shown to significantly, but not persistently improve motor capacity [115]. The EXTEND Orthosis has been tested in a controlled, at home clinical trial, but participants in the control group (who performed exercises from a book) improved more than participants that played

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immersive video games with assistance from the EXTEND Orthosis [116]. Saebo also sells a commercial glove called SaeboFlex, which uses springs and cables to apply forces to resist flexion contractures in the wrist and hand. Unlike the HandSOME system, the springs are not tuned to truly compensate for the unwanted flexor activity, but the device makes up for this in its practicality. It is marketed as an assistive device.

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Serious Games for Home Use

Nearly all the passive devices discussed above are instrumented so that they can be used as inputs to control motivating custom and noncustom video games. The difficulty, repetitiveness, and delayed gratification of motor rehabilitation can make it very challenging for people with a neurological disorder to invest themselves. Serious games can hide the difficulty and repetition inherent to motor rehabilitation by providing a rich and motivating training environment and embedding game-playing elements. The level of difficulty of virtual tasks can often be scaled in ways that purely physical tasks cannot, allowing players to practice at higher success rates. Serious games offer the advantages of providing immediate and enhanced feedback, and can dynamically adjust task difficulty. Many systems offer the opportunity to record and monitor performance and progress, which can be very useful for patients and therapists to monitor performance and track improvements over time. Serious games also offer the opportunity to incorporate motor learning principles, such as motivation, repetitive practice, and enhanced feedback, into rehabilitation interventions [117]. However, the standard input devices (mice, keyboards, and gamepads) used to control commercial video games are poorly suited to games used for rehabilitation because they are designed to require very small, efficient movements, not the types of movements normally prescribed during motor rehabilitation. This has created an interest in unique input devices for rehabilitative serious games. Two of the earliest and most

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readily adopted gaming systems used for rehabilitation were the Nintendo Wii (Nintendo, Tokyo, Japan) and the Kinect System (Microsoft, Richmond, WA) [118, 119]. In a survey of 1071 practicing PTs and OTs, 41% reported having access to a Wii in a clinical setting, whereas 10% had the Xbox Kinect [118]. None of the commercial games, such as the Wii and the Kinect, were deliberately designed for rehabilitation, it is not surprising that they have shortcomings. Movement performance and quality may be diminished by the attributes of the virtual environment (e.g., viewing environment, visual, tactile, auditory, and other sensory cues, etc.), which can consequently be detrimental to motor learning [120, 121]. None of the games involve dexterous tasks requiring grasp/release, pronation/supination, pinch-grip/release or picking up and transferring objects. Nevertheless, many of the Wii or Kinect games are considered to be suitable for rehabilitation [122, 123]. Specifically, exercising with the Wii has also proven to be safe, the system is easy to use, and there is some evidence that participants in trials involving the Wii are less likely to drop out [124]. Results from a recent Cochrane review about the effectiveness of virtual rehabilitation after stroke demonstrated that virtual reality and interactive video gaming have a significant but modest effect on improving arm and hand impairments (measured by the Fugl-Meyer Assessment) when used in addition to usual care to increase overall therapy time (standardized mean difference = 0.49, 95% confidence interval 0.21 to 0.77, 210 participants) [125]. The potential benefits of virtual rehabilitation on improving function in everyday activities, quality of life, and reducing participation restrictions were also identified [125–127]. While well adopted, the use of commercially-available serious gaming systems was not found to be superior to usual care [125]. This suggests that these systems should not be viewed as alternatives to usual care, but rather as useful supplements. More recently, a cohort of low to moderate cost input devices specifically designed for rehabilitation have become available. The most

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noteworthy of these being the Neofect Smart glove, the AbleX rehabilitation system (AbleX, Auckland, New Zealand), Pablo (Tyromotion, Grz, Austria), the MusicGlove and FitMI systems (Flint Rehabilitation devices, Irvine, CA), the NeuroFenix gameball (NeuroFenix, London, UK), and the Rehabilitation Joystick for Computerized Exercise (ReJoyce) workstation marketed by Saebo (Charlotte, NC). The Neofect smart glove is an instrumented data glove designed to be easy to don, doff, and clean. It uses inertial sensors to detect movement of the wrist and hand, and resistive bend sensors to detect movement of the fingers. As such, it can facilitate distal grasp-related exercises in addition to some proximal upper limb exercises. In a RCT (N = 13), chronic stroke participants (defined in the study as >4 months post-stroke) who practiced with the smart glove for 15 30-min sessions in addition to 15 30-min sessions of conventional therapy improved significantly more, as measured by the Wolf Motor Function Test, than participants who performed 30 30-min sessions of conventional therapy [128]. The AbleX system consists of two different input devices: the first is an arm skate with a repositionable button that can be used to detect extension of any of the fingers and the second is an inertial measurement unit (IMU)-based controller that can be used both unimanually and bimanually. The AbleX system has not yet been evaluated in a RCT, but early pilot testing suggests that it is safe, motivating, and potentially effective [129]. The Pablo system by Tyromotion includes a Wii-mote-like orientation and force sensor called the “handle”, an adapter that allows the handle to be used for bimanual exercises and a ball adapter that allows the handle to be held in a different orientation. The software allows therapists to adjust orientation and force thresholds [130]. The MusicGlove is a distally focused data glove for hand rehabilitation designed and priced to be appropriate for home use. The glove can detect opposition of the thumb to all 5 fingers as well as pincer grip and key pinch grip. It is coupled to a game similar to Guitar Hero in which players hit notes by completing the hand grips

M. Demers et al.

indicated by the game [131]. The MusicGlove has been evaluated for both in-clinic [132] and inhome use [131]. In both trials, therapy with the MusicGlove was compared to conventional therapy as a control. The control group performed tabletop exercises in the in-clinic study and therapy guided by a booklet of exercises for the in-home study. In both studies, both groups improved significantly and sustained their improvements (as indicated by changes in Box and Blocks scores), but the MusicGlove groups did not improve significantly more than their corresponding conventional therapy groups [132, 133]. In the in-home study, participants in the MusicGlove group improved significantly more on the Motor Activity Log scores than those of the conventional therapy group. Notably, participants in the MusicGlove group also significantly and voluntarily intensified their dose from an average of 213 ± 301 grips per week during the first week to 466 ± 641 grips per week in weeks 2 and 3 [133]. One of the main limitations of the MusicGlove is that it can only be used for distal exercises. To address this limitation, Flint Rehabilitation Devices released FitMI, a more generic serious game controller with libraries of exercises that support proximal and distal upper limb exercises in addition to lower limb exercises and core strengthening/stretching exercises. The FitMI system consists of two wireless puckshaped controllers that can detect both forces and movement and can supply both haptic and visual feedback. The FitMI System is currently being evaluated in a RCT. The NeuroFenix ball is a round ball that can measure movement and orientation changes. It can be strapped to one hand, used bimanually, or secured in a dock that restricts it to orientation changes only. The shape, size, and straps hold the hand in a favorable position and make the device easier to hold than many other comparable devices [134]. The Rehabilitation Joystick for Computerized Exercise (ReJoyce: Rehabtronics.com; Fig. 23.3) comprises a spring-loaded, segmented arm that presents the user with a variety of spring-loaded attachments representing activities of daily

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Fig. 23.3 a Tele-coaching of an in-home exercise therapy session using the rehabilitation joystick for computerized exercise (ReJoyce) system; b participant

using ReJoyce workstation to play computer games; c movements performed, d selected games

living, such as a doorknob, key, gripper, jar lid and peg. Sensors in the arm and the attachments provide signals that are used by the system’s software to control serious games that exercise specific types of hand movement. The system incorporates an automated, quantitative arm and hand function test which takes about 5 min to complete and provides an overall numerical score that correlates well with the Action Research Arm Test and the Upper Extremity Fugl-Meyer Assessment [135]. It also provides scores for specific tasks such as grasp strength, whole-arm range of motion, pronation-

supination, pinch-grip and manual dexterity and can be performed in the clinic or remotely. Once a user has done the test, the system automatically suggests games and difficulty levels that match their abilities. This is achieved by an algorithm that considers the user’s score on each of the components of the test. If, for example, the user has good ranges of motion but poor pinch-grip strength, games that incorporate pinch-grip are excluded from the suggestion list, and games involving range of motion and grasp-release are included, with difficulty levels corresponding to the relevant test scores.

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The ReJoyce system also facilitates remote tele-coaching. A RCT was completed involving 13 tetraplegic participants who had sustained a spinal cord injury more than a year previously [136]. Participants were block-randomized into two groups, both performing exercise therapy at home with tele-coaching for 1 h/day, 5 days/week for 6 weeks. The control group played computer games played with a trackball and 20 min/day with therapeutic electrical stimulation. The treatment group played serious games on a ReJoyce workstation. Voluntary, hand grasp and release were augmented with functional electrical stimulation (FES) triggered by a wireless earpiece that detected small voluntary tooth clicks. The study demonstrated the feasibility of delivering tele-coached functional electrical stimulation-assisted exercises over the Internet. The treatment group showed clinically important improvements in arm and hand function that significantly exceeded those of the control group. Participants commencing with intermediate functional scores improved the most [137]. The ReJoyce system was designed to be affordable for clinics and, through short-term rental, by individual users who could receive tele-supervised treatment in their homes.

23.10

Therapeutic and Functional Electrical Stimulation

The simple and interactive nature of passive devices makes them a natural complement to functional electrical stimulation (FES) and nonmotor specific processes like Vagus nerve stimulation (VNS). This is particularly true for passive devices that serve as an input to a computer since inputs from the devices and events from the games could both ostensibly be used to trigger stimulation. FES refers to intentionally-triggered electrical stimulation of the motor neurons in a targeted muscle group to assist in a functional task. Although often used as a component of an assistive orthosis, it is also used therapeutically; exercising with FES in addition to standard care

has been shown to improve motor capacity moderately, but significantly, more than standard care alone [138]. Although FES systems are more commonly controlled by switches [136] or electromyography (EMG) [139], they can also be triggered by passive devices designed to detect movement intent [140]. While useful, FESenabled passive devices systems are not without their challenges. Although feedback controlled, multi-joint systems do exist [141], creating coordinated multi-joint movements vias FES is challenging, and nearly all systems rely on pre-programmed stimulation profiles targeting one or two muscle groups (e.g. elbow and forearm extensors) [138]. Furthermore, the motor unit recruitment order obtained by FES is difficult to control and leads to premature fatigue [142]. Vagus Nerve Stimulation (VNS) has been proposed as an adjunct to exercise training in neurorehabilitation. The proposed mechanism of VNS is not through muscle activation, as is the case for FES, but rather, through plastic changes in the central nervous system (CNS). Although the vagus nerve, which innervates autonomic organs such as the heart and gastrointestinal tract, might seem like an unlikely player in neurological motor recovery, it is known that vagal afferents project to neuromodulatory networks in the CNS. Neuromodulatory networks are groups of neurons that are influenced by neurotransmitters such as acetylcholine and noradrenaline and that modulate the activity of other CNS centers. It has been suggested that the cholinergic neuromodulatory network affects motor control and that the noradrenergic neuromodulatory network affects awareness and responsiveness to stimuli. There is evidence from animal studies that VNS can promote CNS neuroplasticity [143]. It has been suggested that VNS applied at the right time during motor practice will increase the neuroplastic response to that practice [144–146]. In the only human clinical trials to date, the timing of VNS was controlled manually by a physical therapist. In principle, the timing could be controlled with movement feedback from a passive rehabilitation device.

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23.11

Telerehabilitation

From all the above, it is clear that the emerging technologies to deliver task practice have the potential greatly to improve arm and hand function in daily life but providing sufficient support after participants leave rehabilitation clinics is problematic. Although the users may benefit from the devices in the clinic, and initially use them daily at home, in the absence of continuing supervision, usage tends to drop off. This transition is a well-known hurdle in rehabilitation [147]. We reasoned that if participants could only perform regular supervised exercise after discharge, they would benefit much more. However, clinics are not ideal locations for outpatients to perform regular training sessions. Travel is often problematic and stressful, limiting the frequency of attendance. In recent years, telerehabilitation, a form of telemedicine, has slowly gained popularity. One of the early systems used for telerehabilitation was the ReJoyce system used for at-home tele-coaching (Fig. 23.2a). In the study mentioned above, Internet-connected ReJoyce workstations were deployed in the homes of 13 tetraplegic participants, located over a wide geographic region in western Canada. Participants were tele-coached daily by a small team of therapists and students. The logistic challenges that were overcome are detailed in a book chapter [148]. A similar study followed on chronic stroke patients in Canada and the UK [149]. Other studies have also shown that telerehabilitation can be convenient and effective for both therapists and patients [148, 150, 151]. Telerehabilitation promotes flexibility, allows greater access to care and continuity of care, and can help decrease racial and economic disparities in health care [152, 153]. With the recent health care crisis induced by the global COVID-19 pandemic, rapid technological changes have followed and telerehabilitation quickly became widely adopted by rehabilitation services across the world [154]. Key barriers that limited the adoption of telerehabilitation previously, such as reimbursement and clinicians’ preference for hands-on

537

interactions, were partly overcome in response to the global pandemic [155]. Resources and guidelines from professional associations were also developed to support clinicians in the delivery of remote rehabilitation (for example, [156, 157]). A recent trend in stroke rehabilitation is the concept of early supported discharge. Early supported discharge is a multidisciplinary team intervention that facilitates earlier discharge from hospital with rehabilitation care provided in the community [158]. Evidence from meta-analyses supports appropriately-resourced early-supported discharge services delivered by a multidisciplinary team to reduce disability and shorten hospital stays in a selected group of stroke survivors [158, 159]. The use of passive devices is particularly well-suited for home use, earlysupported discharge and remote supervision using telerehabilitation.

23.12

Clinical Adoption of Passive Devices

Healthcare professionals play a pivotal role in their patients’ access to novel technologies, and they are ultimately the ones who use health technologies, such as passive devices, in their clinical practice or recommend them for home use [160]. Understanding clinicians’ perspectives and the factors affecting clinical adoption is crucial to enable clinical changes and better widespread use of passive devices in neurological rehabilitation. The clinical adoption of passive devices remains low and is not yet commonplace in clinical rehabilitation settings, which is not different from most health technologies [155, 161]. The results from a survey of 1326 healthcare professionals suggest that clinical decisions to acquire and use new technology devices are multifaceted and are based on the benefits for their patients, the technology’s appropriateness for the setting and logistical practicality within the service delivery system [162]. Patient characteristics, available financial resources,

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technology cost, experience with technology and time demands are key factors shown to impact clinical practice patterns [161–163]. The potential benefit of technology to facilitate positive health outcomes, mainly through the provision of meaningful and objective feedback, and repetitive and independent practice, was identified as a main driver to adoption [155]. Clinicians usually have very busy schedules, with little time to deal with new technology. It is therefore vital to provide equipment that is affordable and simple to use, with highly intuitive computer interfaces that do not require procedural memorization from one session to the next [155, 164]. Since barriers to adoption are multifactorial, financial and administrative support from the leadership, and training of clinicians are essential to ensure clinical adoption and use of health technologies [165]. Future technology development should consider following the stepwise approach and conditions for successful implementation of technology in daily clinical practice, as outlined in a recent systematic review [166]. In the later stages of technology development, the incorporation of user-centered design methods and involvement of clinicians and users are important to facilitate technology adoption [167].

23.13

Perspectives and Conclusions

There is general agreement in the field that the time is ripe for physical and occupational therapy to take advantage of new technologies. It is time to move beyond simple equipment currently used in clinics worldwide, to passive devices that

provide task-specific, motivating games that can also be performed in the participant’s home environment, supervised remotely over the Internet. The advantages of this approach are many: increased compliance, task-specific training on a variety of customized activities, quantification of performance and perhaps most compelling, the ability to provide continuing inhome therapy after acute care in clinics, in a manner that avoids the need for participants to travel, yet retains the important component of one-on-one supervision by enabling therapists to treat participants at times that suit them all. A crucial factor is cost, especially given that costs are generally covered by patients. The cost of passive devices presented in this chapter spans from a few hundred dollars for serious games to a few thousand dollars for tabletop therapy systems, with many options offered below $1,000USD. This chapter has made the case for affordable passive exercise devices that provide entertaining exercises involving full range-ofmotion and manual dexterity, with optional telecoaching and electrical stimulation (summarized in Table 23.1). Task training for arm and hand function on passive devices, with the option of FES-assistance or perhaps VNS, is now an affordable and effective modality of occupational and physical therapy. Passive devices offer numerous opportunities in the field of neurological rehabilitation to support arm and hand motor recovery. Future research could focus on the identification of who might benefit the most from the use of passive devices to guide clinical decision making and maximize the use of scarce health-care resources.

Commercial product

Yes

Yes

No

Yes

Previously

Yes

No

No

No

No

Yes

Yes

Yes

Device

Armeo Boom

Armeo Spring

SpringWear

Rapael SmartBoard

Tailwind

Reha-Slide

SMART Arm

HandSOME

SCRIPT

EXTEND

SaeboFlex

Wii

Kinect

Input device Input device

Whole body—software limited

Tone Compensation

Shoulder, arm, forearm, wrist, thumb (pushbutton)

Fingers, thumb extension

Tone Compensation

Tone Compensation

Wrist, fingers, thumb

Fingers, thumb

Tone Compensation

Linear track

Linear track

Linear track

Tabletop

Fingers, thumb

Shoulder, arm

Shoulder, arm

Shoulder, arm

Shoulder, arm

Gravity support

Gravity support

Shoulder, arm, forearm, wrist, hand grasp-release (optional attachment)

Shoulder, arm forearm

Gravity support

Device type

Shoulder, arm, forearm

Movement target(s)

Table 23.1 Summary of passive exercise therapy devices discussed in this chapter

Yes

Yes

No

Yes

Yes

Yes

Yes

Research only

No

Yes

Yes

Yes

Yes

Computer gaming?

No

No

No

No

No

No

No

No

No

No

No

Yes

No

Validated upper limb function test?

Software dependent

No

No

No

No

No

No

Research only

No

No

No

No

No

Integrated Telerehab/Telecoaching

[123, 170] (continued)

[122, 124]

None

[112]

[113, 169]

[111, 115]

[103, 110]

[102, 104]

[101, 105, 106]

[96]

[91, 94]

[87–89]

[79, 81– 83, 85, 168]

Recent studies

23 Passive Devices for Upper Limb Training 539

Commercial product

Yes

Yes

Yes

Yes

Yes

Yes

Device

Neofect SmartGlove

AbleX

MusicGlove

FitMI

Pablo

ReJoyce

Table 23.1 (continued)

Shoulder, arm, forearm, wrist, hand grasp-release, finger-thumb pinch

Input device

Input device

Input device

Individual finger and thumb movement

Wrist flexion–extension, pronationsupination, grasp, finger-thumb pinch

Input device/tabletop therapy

Arm, wrist, fingers, thumb

Input device

Input device

Wrist individual fingers, thumb

Ankle, core, shoulder, arm, forearm, wrist, hand, power grip, key pinch

Device type

Movement target(s)

Yes

Yes

Yes

Yes

Yes

Yes

Computer gaming?

Yes

No

No

No

No

No

Validated upper limb function test?

Yes

No

Research only

No

No

No

Integrated Telerehab/Telecoaching

[135–137, 150]

None

[171, 172]

[131– 133]

[129]

[128]

Recent studies

540 M. Demers et al.

23

Passive Devices for Upper Limb Training

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Mobile Technology for Cognitive Rehabilitation Amanda R. Rabinowitz, Shannon B. Juengst, and Thomas F. Bergquist

Abstract

Many barriers currently prevent persons living with acquired brain injury (ABI) from receiving an adequate “dose” of cognitive rehabilitation therapy via the traditional clinic-based service delivery model. The application of mobile technology to cognitive rehabilitation is an emerging area of great interest. Advances in mobile technology provide new opportunities for cognitive rehabilitation, with the potential to improve access to care, as well as increase patients’ opportunities to practice and apply skills in their everyday environments—a notion referred to as ecologically valid treatment. The aim of this chapter is to provide an overview of mobile technology in the context of cognitive rehabilitation, with a special focus on applications and interventions targeted towards patients with ABI. We provide a general background on cognitive rehabilitation as a

A. R. Rabinowitz (&) Moss Rehabilitation Research Institute, 50 Township Line Rd., Elkins Park, PA 19027, USA e-mail: [email protected] S. B. Juengst TIRR Memorial Hermann Hospital, Houston, TX 77030, USA e-mail: [email protected] T. F. Bergquist Mayo Clinic, Rochester, MN, USA e-mail: [email protected]

treatment model, discuss opportunities and considerations for developing mobile rehabilitation solutions for users with cognitive impairment, and provide a general overview of the state of the research on mobile cognitive rehab treatments. The vast majority of mobile health apps are not evidence based at present (Ramey et al. in Phys Med Rehabil Clin 30(2):485–97, 2019), however, a small but growing literature is now speaking to the use of smart technologies to both help improve functioning and monitor functioning for persons with ABI. We highlight several areas for future development needed to move towards developing clinical practice guidelines for integrating mobile technology into cognitive rehabilitation. Keywords

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Cognitive rehabilitation Mobile health (mHealth) Stroke Traumatic brain injury (TBI) Apps Technology Rehabilitation

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Introduction

Many barriers currently prevent persons living with acquired brain injury (ABI) from receiving an adequate “dose” of cognitive rehabilitation therapy via the traditional clinic-based service delivery model. This critical service gap exists for a number of reasons, including insufficient clinical resources, poor access to care, and

© The Author(s), under exclusive license to Springer Nature Switzerland AG 2022 D. J. Reinkensmeyer et al. (eds.), Neurorehabilitation Technology, https://doi.org/10.1007/978-3-031-08995-4_24

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limited insurance coverage [1]. Some patients with cognitive disabilities are sent home from acute care without any outpatient cognitive rehabilitation at all. Those fortunate enough to access services often receive coverage that is fragmented and front-loaded, with the majority of therapy provided in the weeks and months after a traumatic brain injury (TBI) or stroke. Support dwindles as patients stabilize and return to the community. Studies suggest that the majority of individuals receive inadequate support after discharge from inpatient rehabilitation, at which point caregivers report a progressive decline in both the quality and quantity of services [2]. The short duration of care is a concern. For example, the field increasingly acknowledges that moderate-severe TBI is a chronic, and even progressive disease process [3, 4], for which ongoing rehabilitation is needed to prevent increasing disability. The evidence from the stroke literature strongly suggests that greater intensity therapy over a longer duration, at least within the first year post-stroke, is associated with greater functional improvements [5, 6]. Without sustained access to services many patients are likely to fall short of achieving their optimal level of functioning, and may even be at risk of experiencing further cognitive decline years later [7]. Given all of this, there is a compelling need for remote delivery of cognitive rehabilitation services to address the needs of those living with acquired brain injury (ABI). The application of mobile technology to cognitive rehabilitation is an emerging area of great interest. Advances in mobile technology provide new opportunities for cognitive rehabilitation, with the potential to improve access to care, as well as increase patients’ opportunities to practice and apply skills in their everyday environments—a notion referred to as ecologically valid treatment. The aim of this chapter is to provide an overview of mobile technology in the context of cognitive rehabilitation, with a special focus on applications and interventions targeted towards patients with ABI. We provide a general background on cognitive rehabilitation as a treatment model, discuss opportunities and considerations for developing mobile rehabilitation

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solutions for users with cognitive impairment, and provide a general overview of the state of the research on mobile cognitive rehab treatments. Finally, we provide recommendations for future development and research in this area.

24.2

Cognitive Rehabilitation

Cognitive rehabilitation has been defined as “a systematic, functionally oriented service of therapeutic cognitive activities, based on an assessment and understanding of the person’s brainbehavioral deficits. Services are provided by qualified practitioners and are directed to achieve functional changes by [1] reinforcing, strengthening, or reestablishing previously learned patterns of behavior, or [2] establishing new patterns of cognitive activity or compensatory mechanisms for impaired neurological systems” (p. 62) [8]. The process of cognitive rehabilitation can be broadly categorized into interventions which are either restorative or compensatory. Restorative interventions, also referred to as direct training or process-specific, are aimed at the impairment level and involve repetitive stimulation of a domain, typically in a systematic, methodical manner. Compensatory approaches are targeted at the functional level and are based on the assumption that although damaged neurological functions cannot be restored, one’s intact strengths and abilities can be used to circumvent impairments. To this end, compensations can reduce the extent of limitations on activities and participation and promote successful community reintegration in spite of residual physical and cognitive impairments. Cognitive rehabilitation initially focused on the use of restorative approaches. Within the past few decades, the focus has been more on compensation for areas of impairment [9]. Historically, the goal of the restorative approaches has been to reduce or eliminate underlying cognitive impairment [10]. Restorative interventions are focused on exercises targeting specific cognitive impairment after brain injury or disease. Although the notion of directly diminishing cognitive impairment is appealing, many have

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noted that a program of rehabilitation will only have relevance to patients and their families to the extent that it improves daily functioning. In other words, successful cognitive rehabilitation should produce functional improvements in patients’ daily lives, and not just higher scores on laboratory tests of cognition. The International Classification of Functioning, Disabilities, and Health (ICF) model was developed by the World Health Organization (WHO) to operationalize the functional changes associated with known medical conditions [11]. This model classifies changes associated with brain injury into changes in (1) body functions and structures, (2) activity, and (3) participation. Body functions and structures are measured by assessment of physical or mental functions, and impairment is measured by deviation from expected levels of performance. Activity limitations are an individual’s inability to complete basic or instrumental activity of daily living (ADL/IADL) (e.g., inability to recall appointments, follow a recipe while cooking, and follow a medication regimen or balance a checkbook). Participation restrictions are a loss or change in social roles (e.g., loss of a job or inability to parent children). Participation is highly individualized and needs to be assessed by self or family report regarding the degree to which an individual is [1] a productive member of society and/or [2] integrated into family and community life. Participation restrictions reflect, for example, the degree to which individuals are limited in their life roles (e.g., ability to run a household or maintain a network of friends and family, employment, education, and volunteer activities). In this model, there is a dynamic interaction between impairments (both physical and cognitive), activity limitations, and participation restrictions which collectively determine an individual’s reintegration into the community. The ICF proposes that ADLs and participation in life roles are the two domains which measure the impact of an injury or illness on daily functioning. It then follows that these same domains best measure treatment effectiveness. Not surprisingly, recent studies of the effectiveness of

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cognitive rehabilitation largely focus on ADLs and participation as their main measures of treatment outcomes [12]. Cognitive rehabilitation is often time- and labor-intensive and can require multiple sessions to help teach and apply new skills to improve daily functioning. While the restorative approach alone may not be adequate for significant gains in functioning, these approaches are used within the framework of the ICF to rebuild discrete skills in coordination with other approaches tied to functional outcomes. Restorative approaches have a clear and important role in the rehabilitation process by being paired with compensatory approaches towards the goal of improving functioning. Distinct compensation skills, such as those used to increase visual scanning, decrease impulsivity, and/or decrease episodes of forgetfulness will only be effective to the extent that they are reliably and consistently used. While it may seem that once a skill or strategy is learned, it will then be readily applied, that in fact is not necessarily the case [13]. No matter how effective a strategy may be in practice, it will not be effective for the patient until they have the appropriate skills in place to use that strategy in their day-to-day life. This process requires repeated practice, particularly within the patients’ usual environments (i.e., their homes and communities). A compensation strategy designed to enhance functioning can be beneficial only to the extent that its purpose is understood and it is reliably and consistently implemented. Implementation can be accomplished in one of two ways: external cueing or skills training. External cueing can be helpful for giving both reminders to complete a task as well as instructions on how to complete it. This may involve, for instance, setting a timer to remind the patient to use the appropriate strategy, possibly including information on how to use that strategy. A caregiver may also provide a direct cue to the patient to use the correct strategy when needed. While such a reminder from a caregiver may be effective, it also means remaining dependent upon some external agent which reduces independence. It also places an

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increased burden on caregivers which can add to caregiver burnout. Skills training, in contrast, works to ensure that strategies are used correctly by teaching the patient-specific skills which will help them to directly apply the strategy. In this case, a targeted intervention is used to (re)learn some discrete skills. Training is delivered using trials of mass learning until skill mastery is consistently demonstrated. The patient is then able to independently execute the specific strategy, although some level of cueing may still be necessary for this to happen. Given that this approach involves improving performance on a specific task or skill, it does not always have direct relevance to ADLs. However, these discrete skills are crucial, if not necessary, to using strategies that ultimately lead to improvement in overall functioning. This approach, though often more time-consuming, ultimately results in the individual acting with a greater level of agency. An example of evidence-based cognitive rehabilitation approach which combines compensatory and restorative approaches is planner acquisition method described by [14]. This has long been a centerpiece of cognitive rehabilitation, into which other compensation strategies are woven. While this strategy is focused on improving memory functioning, the calendar/planner has also been demonstrated as a strategy to address problems including organizational skills, behavioral control, and other areas of cognitive impairment which impact daily functioning. This approach has been supported in multiple subsequent studies as a powerful method to improve functioning in persons with severe memory impairment [12]. The approach works by using both restorative and compensatory approaches in concert, using a three-step process of acquisition application, and adaptation. The acquisition phase focuses on learning the names, location, purpose, and use (i.e., acquiring semantic knowledge) for the components of the calendar/planner. Restorative approaches (e.g., errorless learning, spaced retrieval) are used to learn the components of the

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calendar/planner. Errorless learning is a restorative approach designed to teach specific information to patients by leveraging preserved procedural memory abilities. Learning of the target behavior occurs through the active involvement of the therapist and occurs without their conscious control over its use. In the application phase, the patient practices using the memory notebook in various real-life and/or role-play tasks in the clinic. Working together the patient and therapist choose tasks that are relevant to the patient’s life and develop strategies which help to compensate for impairments which interfere with daily functioning. In the final adaptation phase, skills learned in the first two stages (acquisition and application phases) are applied within community and naturalistic settings. A task that was previously performed in the clinic can now be performed in the patient’s own environment (e.g., remembering to take medications, returning phone calls, and remembering scheduled appointments). The impact on daily functioning in this approach is tied to the use of a specific compensation strategy (in this case, the planner/calendar). However, the use of this strategy is dependent upon the use of restorative approaches that help ensure the strategy is used effectively. Thus, while both the restorative and compensatory components are necessary, neither used in isolation is sufficient to reach goals of greater independence. Success is achieved by using both approaches in concert. The three-stage approach using a combination of both restorative and compensation methods has also been effective to address problems in other cognitive domains, including impairments in attention, executive skills, visual-spatial skills, and communication abilities [15]. Thus, while both restorative and compensatory approaches are important, describing cognitive rehabilitation as making a choice between the two is a false dichotomy and can even be counterproductive. Rather, it is useful to understand the unique contributions of both approaches and how they are used together to achieve goals of greater independence.

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24.3

Mobile Technology for Rehabilitation

The WHO defines mobile health (mHealth) technology as “medical and public health practice supported by mobile devices”, including (but not limited to) smartphones and tablets and biometric/activity monitors (e.g., wearables). Mobile health (mHealth) has the potential to enhance rehabilitation care in multiple ways, including by providing evidence-based education, supporting compliance to treatment plans, supporting monitoring and management of biometrics and chronic symptoms, facilitating patient–provider communication, and promoting effective long-term self-management [16]. Through mHealth technology, healthcare providers can both collect and share health-related information with their patients. When communication via mHealth occurs in only one direction, as in the case of devices that remotely track biometrics (e.g., heart rate monitors) or apps that collected self-reported data (e.g., symptoms trackers), this is considered a one-way system. Another example of a one-way system would be an app that provided information or guided intervention without collecting any data from the user (e.g., a daily meditation app). A two-way system involves a back-and-forth exchange of information between the user (patient) and healthcare provider. Evidence suggests that twoway systems may be preferred by patients, family members, and clinicians [17]. Mobile technology provides new opportunities for delivering interventions in the “here and now” as patients go about their usual activities— a model referred to as ecological momentary intervention (EMI). EMI has recently garnered widespread interest in healthcare technology development for two principal reasons—its ability to extend an intervention beyond the standard treatment context (e.g., clinician’s office) and into individuals’ daily lives, and the advantage of providing an intervention in individuals’ natural environments, encouraging them to apply new skills and behaviors to their actual experiences [18]. Evidence suggests that EMIs

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encourage the practice of new behaviors and skills [19–21], which may improve generalization and increase the impact of interventions (e.g., [22, 23]) when combined with face-to-face clinical contacts. For cognitive rehabilitation, this means interventions could be delivered at the very moment that a person with ABI is performing a cognitively demanding task in their home or community, such as planning a shopping list or preparing a meal. Most adults across all age groups, including adults with disabilities, persons with TBI, use smartphones [24–26]. Beyond being beneficial for using mHealth technology, smartphones are the central hub around which most mHealth technology functions, addressing the frequently voiced need from persons with cognitive disabilities for a single interface to manage a variety of cognitive and health needs [16, 17, 27]. It is therefore not surprising that applications of mHealth technology to health management have grown exponentially in the last decade. In 2017, there were more than 325,000 mHealth apps available, representing a 25% increase over the prior year [28] Apps addressing cognitive issues represent a relatively small share of the market. Most mHealth apps focus on fitness (36%), stress management (17%), and diet (12%), with the remaining 35% focused on disease management [29]. Despite this tremendous growth in the general market, evidence points to a dearth of products that address the particular needs and concerns of users with disability. Disabilityfocused apps represent less than 2% of all mHealth apps examined [30]. A recently published study evaluated the current state of mobile health care for people with disabilities based on reviews of the scientific literature and web-based resources for locating apps, as well as a survey of 377 users with disabilities [31]. They concluded that mHealth development and application for people with disabilities is in its early stages, with only a handful of mHealth tools targeted towards persons with disability and virtually no scientific evidence of their effectiveness. Furthermore, despite reporting a relatively high rate of

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adoption of mHealth apps (40%), survey respondents pointed to problems with accessibility and concerns about the accuracy or relevance of the content in mainstream mHealth apps for disabled users. A recent systematic review of mHealth apps specifically within the domain of rehabilitation found only seven apps focused on TBI rehabilitation [32]. These apps addressed a variety of issues such as symptom monitoring, wayfinding, memory problems, and emotional and behavioral concerns. Only three apps were evaluated with a quantitative outcome assessment, and only two of these were in randomized controlled trials (RCTs) [32]. The authors concluded that early app design and evaluation efforts show promising results, but clearly, more research is clearly needed.

24.4

Ethical Considerations

The commercially, rather than scientifically, driven proliferation of mHealth apps presents additional challenges to applying mHealth technology in practice. Even as research on effective mHealth technology grows, clinicians may struggle to find evidence-based mHealth apps, particularly those addressing the specific needs of individuals with cognitive disabilities, amid the sea of apps available (less than 1% of currently available mHealth apps are evidence-based) [33, 34]. A survey of over 500 rehabilitation clinicians reported that only 23% felt knowledgeable about available rehabilitation technology and only 51% felt comfortable using this technology in practice [35]. Similarly, in another study of healthcare professionals working with youth with brain injuries, despite 75% of patients using mHealth technology, only 42% of providers discussed or facilitated this mHealth use in their patients [36]. Mirroring this, one study found that only 10% of participants with brain injury reported that a clinician discussed the use of mHealth services with them following their injuries [36, 37]. This highlights a critical issue —individuals want to use mHealth technologies, but they are not receiving informed recommendations from their healthcare providers. This

creates a high risk that these individuals may use mHealth technology that is inappropriate for (or even detrimental to) their health, unique needs, and rehabilitation goals. This concern is heightened for those individuals with cognitive impairments that limit insight and decisionmaking. A 2018 systematic review of studies on evaluating mHealth apps concluded that clinicians assess the following about mHealth apps to determine if they are appropriate for patients: (1) design, (2) information/content, (3) usability, (4) functionality, (5) ethical issues, (6) security and privacy, and (7) user-perceived value [38]. Inclusion of the patient in this process is essential, to ensure a match between the patient’s needs, values, and preferences and the features, potential risks and benefits, and costs of the mHealth technology. Individuals with cognitive disabilities do report that smartphones generally provide accessible, convenient, effective, and acceptable support. However, they also desire simplicity, better training and support for learning and continuing to use new technology, and better accessibility features (e.g., larger font) [17, 37, 39–42]. Poor accessibility may explain the disparity in mHealth use between those with and without a disability [16]. Inclusion of people with disabilities in the development and evaluation of mHealth technology is essential, though remains a critical gap in current mHealth development that contributes to this ongoing disparity [16]. Even in the growing literature supporting mHealth use for rehabilitation in populations with cognitive impairment, very little work has been done to identify the best approaches to training individuals to use the technology in their daily lives, leading to frustration and technology abandonment. Broadly, practitioners could support the use and adoption of mHealth by their patients by helping them identify appropriate and evidence-based mHealth apps specific to their needs, values, and preferences and providing hands-on opportunities for their patients to practice with the technology itself before moving on to use it independently [27]. The same approaches used in cognitive rehabilitation (e.g., errorless learning, vanishing cues) could also be

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used to acquire the skills needed to use mHealth technology [43], though instruction should be tailored to the individual learner rather than adopting a one-size-all approach. Further, several strategies should be designed within mHealth platforms themselves, to support ongoing use and reduce abandonment, as we discuss later in this chapter. Despite its potential, mHealth is not a magic bullet and should not replace—nor be implemented in the absence of—health providers. Careful consideration of privacy, safety, and emotional well-being when using mHealth technology is paramount. Mobile technology is a primary part of the future of health care, presenting opportunities to provide quality care to more individuals, but we need to address the challenges inherent in mHealth—especially for those with cognitive impairment—to maintain an effective and ethical practice.

24.5

Mobile Technology for Cognitive Assessment

Mobile technology has a number of capabilities that hold promise for enriching the assessment of cognitive function. Built-in sensors—such as positional sensors (e.g., accelerometer, gyroscope, and GPS), media sensors (e.g., microphone and camera), inherent sensors (e.g., device timer), and participatory user–device interactions (eg, screen interactions, metadata input, app usage, and device lock and unlock)—could provide helpful information for assessing individuals’ behavior for the purposes of training, monitoring, diagnosis, or rehabilitation [44]. For example, sensor-enriched mobile assessment of memory could leverage information from human–device interactions—such as the number of times an individual interacts with the screen to gather necessary information to complete a memory task. Furthermore, mobile technology could serve as a platform for administration, scoring, interpretation, and storage of adapted versions of traditional “paper and pencil” neuropsychological tests. A scoping review published in 2021 found encouraging evidence for

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the use of mHealth technology to collect patientreported outcomes after acquired brain injury, including those capturing neurocognitive functions in day-to-day life [27]. However, despite this great promise, the application of mobile technology for neurocognitive assessment is still in its infancy, and to date, there are no validated mobile assessment tools for neurocognitive assessment in individuals with ABI.

24.6

Mobile Technology for Cognitive Rehabilitative Treatments

There are clear opportunities within the traditional framework of cognitive rehabilitation for mobile technology to enhance treatment. Mobile devices can provide external cueing via alarms and notifications. Using mobile technology in the home and community can support the repeated practice of skills and habits without relying on limited therapist availability for direct supervision in the clinic. Furthermore, mobile technology’s ability to deliver instruction and support practice within the home and community allows skill acquisition and practice to take place in the very contexts in which they will be applied. These features have been leveraged by targeted cognitive rehabilitation applications. In addition to targeted applications, many standard smartphone functions and apps geared towards the general population are now regularly incorporated into cognitive rehabilitation. As indicated earlier, training in the use of external aids—planners, calendars, memory notebooks— has always been a critical aspect of compensatory treatments. Traditional paper and pencil aids however can be bulky and easily misplaced, while smartphones allow users with cognitive impairment to easily and discretely take advantage of electronic alarms, calendars, to-do lists, GPS, and other cognitive aids as they go about their daily lives. A recent study from the TBI Model Systems National Database study on internet use (N = 337) of individuals with TBI found that 95% owned a mobile phone and over 75% used their smartphone to access the internet

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[24]. A smaller survey of individuals with TBI (N = 29) found that use as a memory and organizational aid was a commonly cited benefit of smartphones for this population [37]. A 2007 review of the literature on efficacy of using external memory aids for managing memory disorders identified critical research gaps in the study of electronic external aids for users with TBI, including, a poor understanding of factors that lead to long-term adoption of external aids [45]. The authors also note that further research is needed on the design and selection of aids and the evaluation, instruction, and ongoing monitoring of people using the devices [45]. The vast majority of mHealth apps are not evidence-based at present [16]. There is however a small but growing literature speaking to the use of smart technologies to both help improve functioning and monitor functioning for persons with ABI. In a systematic review looking at literature through May 2019, Kettlewell and coauthors (2019) were only able to identify four studies that employed a randomized controlled trial methodology and focused on functioning. They concluded that there was insufficient evidence available to support the benefit of personal smart technologies to improve outcomes in this population [46]. In a scoping review, published the same year, Juengst and co-authors (2019), found that mHealth showed some potential for persons with TBI to use as a compensation strategy for cognitive impairment, as well as a method for monitoring reducing symptoms and a way of addressing social educational goals. Looking at literature published between 2012 in 2019, the authors identified 16 papers which met their criteria. These studies used multiple approaches, with most studies focusing on everyday memory functioning and using simple prompts to help complete a specific task at a given time. The studies also collectively show the importance of the involvement of a human component in the intervention and the need for the contact to be meaningful to the person served. Also evident was the importance of understanding the cognitive difficulties experienced by persons with TBI that may cause greater difficulty in learning or

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using the technology itself. The qualitative evaluation of these interventions indicated encouraging outcomes in terms of better symptom management, decreasing memory problems and mood difficulties, and improving overall well-being than more formal outcome measures showed alone [47]. More recently, Juengst and colleagues (2021) conducted a scoping review of literature from 2015 to 2019 of studies which used mHealth technology for assessment of patient-reported outcomes in community-dwelling individuals with ABI which included ecological momentary assessment for data collection. 12 manuscripts were identified which met inclusion criteria. The interventions were at various stages of development and testing, supporting a range of cognitive skills from everyday memory to executive function (planning, organization, goal setting) to community participation (social, academic/vocational) [27]. Though the evidence remained insufficient to inform clinical guidelines, it did show promise for the future, and overall users with TBI were enthusiastic about using mHealth technology as a compensatory strategy [17]. Given the great variability with respect to what constructs were measured as well as the frequency duration and timing of the exact data collection, the authors were unable to make recommendations regarding either the optimal content or timing of specific mHealth platforms for capturing information. The authors did note that, as with previous reviews, personalization of the technology to account for the unique needs and abilities of each person was crucial to success. This idea has long been one of the main tenants of formulating rehabilitation treatment: to gather data and ask questions to determine which approach has the best ‘‘goodness of fit’’ between technology and a given individual [48].

24.7

Future Directions

Though the extant evidence for use of mHealth technology for cognitive rehabilitation is certainly promising, there are several areas for

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future development so that we can move towards developing clinical practice guidelines. As discussed in this chapter, mobile technology has the potential to extend the duration and intensity of cognitive rehabilitation (including the delivery of ‘booster’ sessions or maintenance therapies), provide assessment and treatment in ecologically valid contexts, and permit time-sensitive interventions. Long-term use of mobile technology for cognitive rehabilitation should consider patients’ evolving needs as they progress through their individual recovery trajectory, as well as their meaningful life roles and personal treatment goals, which are also subject to change over time [49]. Future design efforts should capitalize on the complimentary expertise of various stakeholder groups from both the “technology domain” (technology developers) and the “healthcare domain” (clinicians, consumers with brain injury, and their caregivers). This co-design approach engages stakeholders representing patients, clinicians, and healthcare administration at the frontend of design to identify clinically meaningful problems and concerns, while still allowing for the opportunistic exploration of “technologyinspired” approaches that envision new healthcare technologies that would be overlooked by strictly “problem-driven” design approaches. That said, we should not forget to pause and ask not just what we can do but also what we should (or should not) do. Inclusion of relevant stakeholders—from patients to their family members to clinicians to payors—is crucial to ensuring meaningful, feasible, scalable, and ethical development of mHealth to support cognitive rehabilitation [49]. Ongoing work is needed not only in the development and scientific evaluation of specific mHealth technologies and platforms but also in how we operationalize, implement, and reimburse for cognitive rehabilitation services provided via mHealth. For the reasons discussed above, the goal of cognitive rehabilitation is predominantly on compensation and adaptation —that is, long-term behavioral changes to support function—rather than restoration, hence, we also need to carefully consider long-term

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consumer engagement and technical support for using these technologies independently and indefinitely. Lastly, the rapid evolution of mHealth often leaves us scrambling to keep up, both with high-quality evidence (especially as technologies rapidly develop, rendering earlier technologies obsolete) and with important ethical considerations. Below we offer suggestions of some specific promising directions for future development. First, we need to employ a framework for thinking about how to operationalize the specific therapeutic components to include in mHealth technology for different cognitive rehabilitation targets. This is an important, but often overlooked, step for both research and clinical implementation. For research, we cannot test what components are most effective if we do not have a concrete way to measure those components directly. For clinical implementation, we need standardized definitions for documentation and reimbursement. The Rehabilitation Treatment Specification System (RTSS) provides this kind of framework for defining specific treatment components in rehabilitation interventions. The RTSS characterizes rehabilitative treatments according to the targets, ingredients, and mechanisms of action. Thus, RTSS provides a coherent theory-based framework that researchers can use to systematically test the effects of specific ingredients on specific targets [50]. Classifying mHealth treatments for cognitive rehabilitation according to this approach could greatly accelerate development, as it would enable researchers to test the effectiveness of mobile delivery of specific treatment ingredients, and these evidence-based treatment ingredients could then be implemented across various platforms as technology continues to advance. Second, we need to develop processes for including mHealth technology in the ongoing treatment process and in clinician workflow. This includes processes for adapting mHealth technology use in response to patient progress or decline [49]. Adaptations may be required for the content, timing, and/or frequency of assessments or EMI in response to these changes. Processes may be external to the technology—such as plans

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for ongoing re-evaluation to ensure the technology is still meeting the patient’s and family’s needs, and process internal to the technology— such as built-in flexibility and customizability within the apps themselves. Individuals will need feedback from providers as well, so the use of mHealth should continue to be paired with, rather than replace, contact with healthcare providers. Input from clinicians and developing training for clinicians to enable them to identify evidencebased mHealth apps and to provide effective training in the use of mHealth technology to their patients will be critical. Further, clinicians may have different preferences and needs for how they access the data collected from their patients via mHealth, how the data are presented to them (e.g., design of a “clinician portal” in a two-way mHealth system), and how mHealth use would impact, and be incorporated into, their clinical workflow. Third, we need to develop strategies for engaging individuals with disabilities and their caregivers in the use of mHealth technology, both in the short and long term. This can be done in several ways. One critical approach is to include end-users in the development of mHealth technology. Doing so will promote accessibility, design consistent with patient needs, preferences, and values, and long-term adoption. It will also allow us to identify—and hopefully, circumvent or prevent—barriers to mHealth technology use that result in frustration with technology abandonment. In addition to addressing external factors, like appropriate and accessible tech support, strategies need to be embedded into mHealth technology design to promote engagement and long-term use. Such strategies that have been identified in several past studies include iterative feedback about progress (e.g., not only having individuals track symptoms but also providing feedback and what these symptoms mean), ability to customize mHealth content and notification based on personal needs and preferences (e.g., limiting notifications only to the most salient components for that individuals), and including “gamification” (e.g., use of in-app challenges and rewards based on engagement). Such “gamification” has been successfully employed in other

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mHealth technologies, such as those targeting exercise, weight loss, or smoking cessation [51– 53], and in one study targeting symptom management after youth concussion [54]. Last, we need to develop better training and early support for learning how to use mHealth technology, to ensure that individuals are able to continue using mHealth technology independently. Training should be individualized based on users’ familiarity and comfort with mHealth technology and on their cognitive strengths and limitations. As technology rapidly develops, so do the future directions that can be explored for using mHealth in cognitive rehabilitation. Technology developers refer to a new era of personal technology—the “post-app era”—wherein the way that users interact with their smartphones is becoming more streamlined. Increasingly users will rely on mobile technology to carry out dayto-day tasks without ever needing to open a traditional mobile application. Cloud computing, advancing in machine learning and artificial intelligence (for example, conversational user interfaces know as virtual assistants or “chatbots”), and the seemingly inevitable transition into the “post-app era” are currently paving the way for previously unthought-of mHealth technology. Exciting as this may be, we cannot rush ahead without considering the ethical implications. These include, but certainly are not limited to the following. First, we must consider what we are tracking or providing to individuals with cognitive disabilities in the absence (or synchronous presence) of a human clinician. Second, we must carefully determine what mHealth technology can replace in current practice versus how it would be best used to complement or supplement current practice. Third, advances in mHealth technology should be made to close rather than widen health disparities. While highincome countries remain at the forefront of developing the latest mobile technologies used in health care, the rate of penetration of such technologies in low- and middle-income countries has recently exceeded that of their wealthier neighbors [26]. Together, there is much potential for mHealth technology to improve the

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effectiveness and reach of cognitive rehabilitation interventions, but we must ensure it is done with the same consideration afforded to cliniciandelivered care.

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559 13. Bergquist TF, Jackets MP. Awareness and goal setting with the traumatically brain injured. Brain Inj. 1993;7(3):275–82. 14. Sohlberg MM, Mateer CA. Introduction to cognitive rehabilitation: theory and practice. Guilford Press; 1989. 15. Eberle R, Rosenbaum A. Cognitive rehabilitation manual and textbook: translating evidence-based recommendations into practice. 2nd ed. In ACRM Publishing; In Press. 16. Ramey L, Osborne C, Kasitinon D, Juengst S. Apps and mobile health technology in rehabilitation: the good, the bad, and the unknown. Phys Med Rehabil Clin. 2019;30(2):485–97. 17. Osborne CL, Juengst SB, Smith EE. Identifying usercentered content, design, and features for mobile health apps to support long-term assessment, behavioral intervention, and transitions of care in neurological rehabilitation: an exploratory study. Br J Occup Ther. 2021;84(2):101–10. 18. Heron KE, Smyth JM. Ecological momentary interventions: incorporating mobile technology into psychosocial and health behaviour treatments. Br J Health Psychol. 2010;15(1):1–39. 19. Hay CE, Kinnier RT. Homework in counseling. J Ment Health Couns. 1998;20(2):122. 20. Kazantzis N, L’Abate L. Handbook of homework assignments in psychotherapy. Springer; 2007. 21. Newman MG, Consoli A, Taylor CB. Computers in assessment and cognitive behavioral treatment of clinical disorders: anxiety as a case in point. Behav Ther. 1997;28(2):211–35. 22. Kalichman SC, Kalichman MO, Cherry C, Swetzes C, Amaral CM, White D, et al. Brief behavioral self-regulation counseling for HIV treatment adherence delivered by cell phone: an initial test of concept trial. AIDS Patient Care STDS. 2011;25(5):303–10. 23. Malow RM, West JA, Corrigan SA, Pena JM, Cunningham SC. Outcome of psychoeducation for HIV risk reduction. AIDS Education and Prevention. 1994 24. Baker-Sparr C, Hart T, Bergquist T, Bogner J, Dreer L, Juengst S, et al. Internet and social media use after traumatic brain injury: a traumatic brain injury model systems study. J Head Trauma Rehabil. 2018;33(1):E9-17. 25. Morris JT, Sweatman M, Jones ML. Smartphone use and activities by people with disabilities: user survey 2016. J Technol Pers Disabil. 2017;5:50–66. 26. Pew Research Center. Mobile fact sheet. Pew Research Center. 2018; 27. Juengst SB, Terhorst L, Nabasny A, Wallace T, Weaver JA, Osborne CL, et al. Use of mHealth technology for patient-reported outcomes in community-dwelling adults with acquired brain injuries: a scoping review. Int J Environ Res Public Health. 2021;18(4):2173.

560 28. mHealth Economics 2017 Report: Status and trends in digital health | R2G [Internet]. research2guidance. [cited 2022 Jan 5]. https://research2guidance.com/ product/mhealth-economics-2017-current-status-andfuture-trends-in-mobile-health/. 29. Aitken M. Patient adoption of mHealth: use, evidence and remaining barriers to mainstream acceptance. IMS Institute of Healthcare Informatics; 2015 Sep. 30. Constantino T. IMS Health Study: patient options expand as mobile healthcare apps address wellness and chronic disease treatment needs [Internet]. IMS Institute for Healthcare Informatics; 2015 Sep [cited 2022 Jan 5]. https://www.businesswire.com/news/ home/20150917005044/en/IMS-Health-StudyPatient-Options-Expand-Mobile. 31. Jones M, Morris J, Deruyter F. Mobile healthcare and people with disabilities: current state and future needs. Int J Environ Res Public Health. 2018;15 (3):515. 32. Nussbaum R, Kelly C, Quinby E, Mac A, Parmanto B, Dicianno BE. Systematic review of mobile health applications in rehabilitation. Arch Phys Med Rehabil. 2019;100(1):115–27. 33. Anthes E. Mental health: there’s an app for that. Nature News. 2016;532(7597):20. 34. Free C, Phillips G, Galli L, Watson L, Felix L, Edwards P, et al. The effectiveness of mobile-health technology-based health behaviour change or disease management interventions for health care consumers: a systematic review. PLoS Med. 2013;10(1): e1001362. 35. Morris J, Jones M, Thompson N, Wallace T, DeRuyter F. Clinician perspectives on mRehab interventions and technologies for people with disabilities in the United States: a national survey. Int J Environ Res Public Health. 2019;16(21):4220. 36. Plackett R, Thomas S, Thomas S. Professionals’ views on the use of smartphone technology to support children and adolescents with memory impairment due to acquired brain injury. Disabil Rehabil Assist Technol. 2017;12(3):236–43. 37. Wong D, Sinclair K, Seabrook E, McKay A, Ponsford J. Smartphones as assistive technology following traumatic brain injury: a preliminary study of what helps and what hinders. Disabil Rehabil. 2017;39(23):2387–94. 38. Nouri R, R Niakan Kalhori S, Ghazisaeedi M, Marchand G, Yasini M. Criteria for assessing the quality of mHealth apps: a systematic review. J Am Med Inform Assoc. 2018;25(8):1089–98. 39. Bedell GM, Wade SL, Turkstra LS, HaarbauerKrupa J, King JA. Informing design of an app-based coaching intervention to promote social participation of teenagers with traumatic brain injury. Dev Neurorehabil. 2017;20(7):408–17.

A. R. Rabinowitz et al. 40. Lannin N, Carr B, Allaous J, Mackenzie B, Falcon A, Tate R. A randomized controlled trial of the effectiveness of handheld computers for improving everyday memory functioning in patients with memory impairments after acquired brain injury. Clin Rehabil. 2014;28(5):470–81. 41. Narad ME, Bedell G, King JA, Johnson J, Turkstra LS, Haarbauer-Krupa J, et al. Social Participation and Navigation (SPAN): description and usability of app-based coaching intervention for adolescents with TBI. Dev Neurorehabil. 2018;21(7):439–48. 42. Powell LE, Wild MR, Glang A, Ibarra S, Gau JM, Perez A, et al. The development and evaluation of a web-based programme to support problem-solving skills following brain injury. Disabil Rehabil Assist Technol. 2019;14(1):21–32. 43. Ehlhardt LA, Sohlberg MM, Kennedy M, Coelho C, Ylvisaker M, Turkstra L, et al. Evidence-based practice guidelines for instructing individuals with neurogenic memory impairments: What have we learned in the past 20 years? Neuropsychol Rehabil. 2008;18(3):300–42. 44. Templeton JM, Poellabauer C, Schneider S. Enhancement of neurocognitive assessments using smartphone capabilities: systematic review. JMIR Mhealth Uhealth. 2020;8(6): e15517. 45. Sohlberg MM, Kennedy M, Avery J, Coelho C, Turkstra L, Ylvisaker M, et al. Evidence-based practice for the use of external aids as a memory compensation technique. J Med Speech-Lang Pathol. 2007;15(1):xv–xv. 46. Kettlewell J, das Nair R, Radford K. A systematic review of personal smart technologies used to improve outcomes in adults with acquired brain injuries. Clin Rehabil. 2019;33(11):1705–12. 47. Juengst SB, Hart T, Sander AM, Nalder EJ, Pappadis MR. Mobile health interventions for traumatic brain injuries. Curr Phys Med Rehabil Rep. 2019;7 (4):341–56. 48. Wright BA. Motivating children in the rehabilitation program. 1983; 49. How T-V, Hwang AS, Green REA, Mihailidis A. Envisioning future cognitive telerehabilitation technologies: a co-design process with clinicians. Disabil Rehabil Assist Technol. 2017;12(3):244–61. 50. Hart T, Dijkers MP, Whyte J, Turkstra LS, Zanca JM, Packel A, et al. A theory-driven system for the specification of rehabilitation treatments. Arch Phys Med Rehabil. 2019;100(1):172–80. 51. Burkow TM, Vognild LK, Johnsen E, Bratvold A, Risberg MJ. Promoting exercise training and physical activity in daily life: a feasibility study of a virtual group intervention for behaviour change in COPD. BMC Med Inform Decis Mak. 2018;18 (1):136.

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52. Rajani NB, Weth D, Mastellos N, Filippidis FT. Use of gamification strategies and tactics in mobile applications for smoking cessation: a review of the UK mobile app market. BMJ Open. 2019;9(6): e027883. 53. Timpel P, Cesena FHY, da Silva CC, Soldatelli MD, Gois E, Castrillon E, et al. Efficacy of gamificationbased smartphone application for weight loss in

561 overweight and obese adolescents: study protocol for a phase II randomized controlled trial. Ther Adv Endocrinol Metab. 2018;9(6):167–76. 54. Worthen-Chaudhari L, McGonigal J, Logan K, Bockbrader MA, Yeates KO, Mysiw WJ. Reducing concussion symptoms among teenage youth: Evaluation of a mobile health app. Brain Inj. 2017;31 (10):1279–86.

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Telerehabilitation Technology Verena Klamroth-Marganska, Sandra Giovanoli, Chris Awai Easthope, and Josef G. Schönhammer

Abstract

Due to physical and cognitive deficits, it is often difficult and costly for individuals who have suffered a stroke to access on-site neurorehabilitation. Telerehabilitation offers the opportunity to improve the rehabilitation process as it can provide intensive supervised rehabilitation in the home environment. The term telerehabilitation refers to the provision of therapeutic services at a distance, enabled by electronic telecommunication and information technologies. Services are provided through a variety of technical systems with different purposes and capabilities. This chapter provides an overview of technical solutions for providing telerehabilitation services to treat the main consequences of stroke, namely paresis of the upper and lower extremities, and communication difficulties. We describe the

V. Klamroth-Marganska (&) Research Unit for Occupational Therapy, School of Health Professions, Zurich University of Applied Sciences ZHAW, Zurich, Switzerland e-mail: [email protected] S. Giovanoli . C. A. Easthope . J. G. Schönhammer cereneo Foundation, Center for Interdisciplinary Research CEFIR, Vitznau, Switzerland e-mail: [email protected] C. A. Easthope e-mail: [email protected] J. G. Schönhammer e-mail: [email protected]

communication tools, sensor technologies, virtual reality systems, and robots for service delivery and explore the facilitators and barriers to successful implementation. Evidence is summarized in the context of teleassessment, telemonitoring, and teletherapy. Keywords

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Telehealth Teletherapy Teleassessment Stroke

25.1

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. Telemonitoring .

Introduction

In neurorehabilitation, clients face numerous barriers to accessing usual on-site appointments including geographic isolation, limited resources, and shortage of time; all these may lead to the lack of compliance with rehabilitation regimens [1]. Telerehabilitation may help to overcome these barriers and is a suitable supplement or even substitute to usual rehabilitation. The term telerehabilitation describes the use of information and communication technologies (ICT) for the delivery of rehabilitation services to people at a distance, for example, in the home environment, in the out-client area, in the in-client sector, or at school [2, 3]. Early attempts to use phones for teletherapy of people with aphasia were made as early as the 1970s. Vaughn et al. designed a device that combined phones “with a

© The Author(s), under exclusive license to Springer Nature Switzerland AG 2022 D. J. Reinkensmeyer et al. (eds.), Neurorehabilitation Technology, https://doi.org/10.1007/978-3-031-08995-4_25

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variety of terminal devices, such a speaker phone for speech, handset for other auditory signals, a ‘‘Touchtone’’ keypad for pointing responses, a teletypewriter for typing, a ‘‘Telenote’’ device for handwriting, and computer terminals for more generalized applications” [4, 5]. Standardized assessments such as the modified Barthel Index were also successfully carried out on the phone in the 1980s [6]. In the age of digitalization and the Internet, technologies are advancing, becoming cheaper, more accessible and ubiquitous, and available to a greater number of people. As ICT continues to develop, telerehabilitation is becoming a global treatment option.

25.1.1 Benefits Telerehabilitation offers many advantages as compared to conventional in-person therapy: By eliminating the need to travel long distances, telerehabilitation helps clients comply with treatment protocols. The remote delivery of service is especially important for those who have difficulties traveling due to physical impairments or who live in rural or remote settings. Telerehabilitation can increase the frequency of services and enhance continuity of care [2]. Perhaps even more importantly, telerehabilitation promotes clients’ involvement and empowers them to care for and manage their medical needs and therapeutic interventions [2, 7]. Clients’ empowerment and engagement are key elements to enhance neuroplasticity and facilitate functional recovery in neurological disorders [8, 9]. Most health care, perhaps as much as 85%, is self-care [10, 11]. The home environment as an authentic environment for the experiences of functioning encourages clients to develop problem-solving skills and actively engages them in the rehabilitation process [7]. Through the use of new technologies, the client is not alone in his home rehabilitation. As therapists have access to data from home, clients can self-train knowing that their therapist is tracking their progress. Automatized real-time feedback from technologies motivates and engages the client, thus, further promoting neuroplasticity [12, 13].

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Educational materials that directly address the current state (as assessed during therapy) can be delivered promptly when needed and improve client knowledge. “The greater the understanding and comprehension that clients have in terms of the rehabilitation process, goals, diagnosis, and healing process, the more invested they will be in their own recovery” [8]. Clients with neurological disorders have often complex healthcare needs that require a multidisciplinary coordination of their rehabilitation. Services are provided by many professionals including physicians, psychologists, rehabilitation engineers, audiologists, nurses, and educators, out-client rehabilitation, and, above all, by occupational and physical therapists as well as speech–language pathologists [3]. Telerehabilitation can be integrated into telemedicine platforms that allow to connect all stakeholders including the client and share all relevant information in real-time.

25.1.2 Interaction from a Distance In telerehabilitation, therapists and clients communicate from a distance and exchange health data with the use of technologies. These include video conferencing systems, instant messaging platforms, mobile health (mhealth) applications, electronic client portals, and digital client platforms. By incorporating additional technologies, telerehabilitation enables not only audiovisual interaction but also real-time exchange of ``handson'' information. Wearables, for example, monitor specific types of physiological parameters that are readily accessible from outside the human body. Robotic devices can give assistance and transmit information about forces and movements to give therapists a deeper understanding of the client’s sensorimotor performance [14, 15]. Environmental sensors provide information about how clients interact with their environment. Taken together, the potential of interaction between therapists and clients is not restricted to audiovisual communication. Innovations in medical devices such as nanosensor technologies and

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virtual reality may expand the possible information exchanged in the future [16].

25.1.3 Synchronous and Asynchronous Therapy Telerehabilitation may be delivered synchronously when therapist and client communicate in real-time [17]. Usually, they speak via video conferencing and telemedicine systems. Telerehabilitation may also be performed asynchronously when information for therapy or education is delivered to the client so that he/she can train alone, or when the client trains alone and recorded data is transmitted to the therapist [17]. After a familiarization phase under therapist supervision, clients can continue therapy independently. Thus, telerehabilitation shifts from remote one-to-one therapy to (partially) supervised self-training. Self-training may involve medical applications (apps) or gaming software run on flat screens, head-mounted devices, or projection systems, that provide clients with visual and auditory feedbacks (but may also include other sensory inputs such as touch, movement, balance, and smell) [18]. The user interacts with the virtual environment by a mouse or joystick, cameras, sensors, or haptic (touch) feedback devices. The data may be further analyzed or processed (e.g., to recognize trends of recovery over days or weeks). Robotics for self-therapy enable clients to practice independently with mechanical assistance [19]. Self-training with integrated feedback may empower and motivate clients to engage in therapy.

25.1.4 Acceptance of Telerehabilitation Although telerehabilitation is applied regularly in rural regions, its general implementation and acceptance was rather limited among therapists and clients until the Coronavirus disease 2019 (COVID-19) pandemic forced a rapid adoption of

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telerehabilitation. Because severe acute respiratory syndrome coronavirus 2 (SARS-CoV-2) is highly contagious, vulnerable persons were isolated starting in 2020. Unfortunately, those who require rehabilitation are often vulnerable (e.g., 90% of stroke clients in Switzerland are 65 or above) [3]. Furthermore, the pandemic went along with very early in-client discharge and a suspension of rehabilitation services in out-client settings in 2020, which decreased the access to rehabilitation services and their availability [7, 8]. To aggravate the situation, SARS-CoV-2 infection is not only associated with pneumonia, but also with neurological complications: The hypercoagulable states may lead to stroke and other neurological manifestations. Thus, COVID19 did not only decrease the supply of rehabilitative services but also additionally increased the number of persons who required these. Telerehabilitation was a promising way to fill the supply gap of rehabilitative measures during the pandemic. As telerehabilitation reduces the contact between vulnerable persons and the environment, it enables the therapy of persons in strict isolation [15]. Thus, traditional in-person visits were often replaced by tele-services. In Switzerland, for example, about two-thirds of occupational therapists (OT) provided telerehabilitation during the COVID-19 pandemic lockdown although they were often not reimbursed and the pandemic did hardly allow to implement measures associated with the successful implementation of telerehabilitation such as education and training, as well as administrative and technical support [20]. The service relied mainly on audiovisual interaction. The media used most was phone followed by chat services, email, video conferencing systems, and short messages services (SMS). One reason for this rather low use of modern technologies may be clients’ preferences. 95% of people aged 65–69 already use the Internet, but only 35% of people aged 85 and over do (numbers from Switzerland 2020 [21]). Three-quarters of the OTs rated their clients’ experience of telerehabilitation as positive or rather positive during the lockdown [20]. About two-thirds of OTs described their own

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experience as broadly positive or rather positive, and about one-quarter as negative or rather negative [22]. Usually, it needs two for out-client telerehabilitation: clinicians and clients. What does it take to improve the telerehabilitation experience for both user groups? The uptake of tele-services is influenced by facilitators and barriers. Almathamy et al. [23] reviewed scientific literature regarding factors that influence clients’ adoption of telemedicine. The telemedical counseling they focused on was delivered with synchronous video conferencing systems or software, most studies reported on specially developed systems. The authors distinguished internal and external factors that influence the adoption. While the former describes the users’ behaviors and motivations, the latter refers to the system and the surrounding environment. External barriers described were mainly technical (e.g., low internet speed or difficulty to use the system). Internal barriers included resistance to technology, the lack of eye as well as physical and social contacts with telemedicine, and clients’ security concerns. Accordingly, understanding of the technology and high internet speed were important facilitators. Family involvement during treatment facilitated the use of systems. For the service to be effective, clinicians do not only need to find it useful for clients, but they themselves require to be ready for system uptake [24]. To achieve successful implementation and sustainable use of telerehabilitation, it needs preparation, teaching, and support from an organization. Jafni et al. adopted findings from a critical care information system and identified seven significant factors that influence the adoption of telerehabilitation by therapists [25, 26]: . The system’s functions and features should be more useful, faster, and easier compared to existing systems. . The delivered information must be accurate and understandable for every stakeholder including the client.

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. Service quality requires follow-up services and responsiveness to support requests. . Users should receive regular training that targets technology skills, safety, and security. Implementation plans help avoid problems. . System performance must be reliable. . Factors that influence users’ decisions on system use include professional hierarchy, age, education, system ease of use, and cutting-edge technology [26]. The system must be effective and useful to meet the expectations even of stakeholders with limited IT skills. . A strategy is needed to get stakeholders to adopt the new system. This includes stakeholder involvement, interactive communication (e.g., meetings), and a “champion” who can influence and encourage others to use the system.

25.1.5 Effectiveness of Telerehabilitation The effectiveness of telerehabilitation has not yet been sufficiently proven. There are several studies showing that telerehabilitation interventions have either better or equal salutary effects on motor, higher cortical, and mood disorders as well as quality of life compared with conventional in-person interventions. However, evidence is based mostly on pilot projects that are small in sample size and proof-of-concept in nature, and randomized controlled trials (RCT) are scarce. A meta-analysis in 2020 by Laver et al., involving 22 RCTs, examined the efficacy of telerehabilitation after stroke [12]. The authors concluded that data are not sufficient to draw definitive conclusions as there was variance in interventions and comparators among studies, few adequately powered studies, and several studies with risk of bias. However, at this point telerehabilitation seems not inferior to inperson services or usual care regarding improvement in upper limb function, improvement of health-related quality of life,

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independence in ADL, or reduction of depressive symptoms [12]. Therapy is an important part, but it is not the only service that is delivered in rehabilitation: Rehabilitation involves also monitoring, assessment, prevention, assistance, supervision, education, consultation, and coaching [3]. We focus on telemonitoring, teleassessments, and teletherapy because current research, especially in these areas, employs new technologies that go beyond audiovisual interaction. Telemonitoring is based on the observation of behavioral, biological, or psychological signals. Standardized teleassessments complement this continuous realworld data in providing metrics that mirror clinical assessments [27]. The term teletherapy describes the delivery of sensory motor and cognitive rehabilitative treatment.

25.1.6 Stroke We focus on stroke as the manifestations and the recovery process after stroke are good examples of how different technological advances may cover the entire neurorehabilitation spectrum of a client. Stroke is one of the leading causes of longterm disability in the world and survivors often need long-term neurorehabilitation treatment. About 26% remain disabled in basic ADL and 50% suffer from reduced mobility due to hemiparesis [28]. Aphasia and depression are other frequent causes of disability. Thus, stroke is a disease of immense public health importance with growing economic and social consequences [29]. We focus on telerehabilitation for the main complications of stroke, namely paralysis or loss of muscle movement of the upper and lower extremities, and communication difficulties.

25.2

Upper Extremity

A frequent consequence of stroke is a decline in functioning of the upper extremities, like the shoulders, arms, hands, and fingers [30–32]

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which, in turn, is often associated with a deterioration in life quality [33, 34]. After a stroke, stroke survivors enter a recovery process. Recovery refers to the improvement in one or more components of an individual’s functioning over time [35]. Importantly, at the level of motor behavior, recovery must be distinguished from compensation [36–38]. Behavioral recovery refers to actions that reflect restitution of pre-morbid movements, that is, actions using the same anatomical body parts (i.e., muscles, joints, effectors) for task accomplishment as before the neurological disease, whereas compensation refers to performing a task with different/additional body parts [35]. For example, when an individual reaches for a cup on a table, only shoulder, elbow, wrist, and finger movements may be employed before a stroke. Conversely, after a stroke, a person might make additional trunk movements in direction of reach to support arm and hand movements.

25.2.1 Assessment of Motor Functioning 25.2.1.1 Clinical Assessment Observation-based Clinical Assessment Neurorehabilitation research examines the factors of interventions that optimize recovery, like the type of therapy approach and the time of delivery in clients’ rehabilitation journey [35]. Evidence for the effectiveness of an intervention is key for generation and accumulation of knowledge, which is fundamental for decision-making in the treatment of individuals [39]. However, the effectiveness of an intervention study is often not comparable to the effectiveness of other studies, due to differences in measured constructs and assessment tools [39]. Therefore, many efforts were made, like systematic reviews and Delphi studies, to determine a set of outcome measures for motor functioning that should be used universally [39–45].

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Outcome measures are typically sorted according to several categories [39, 44]. First, they are categorized by the International Classification of Functioning, Disability, and Health (ICF) domains. The ICF uses the term functioning to refer to an individual’s health and disability status. Functioning is regarded as a general term and is defined by three sub-domains [46]: 1. Body structure and body functions are anatomical parts of the body and physiological and psychological functions. Problems in body structure or function are defined as impairments. 2. Activity is the execution of tasks by an individual. Problems with task accomplishment are referred to as activity limitations. 3. Participation is an individual’s involvement in life situations. Problems during participation in real-life situations are called participation restrictions. For example, when an individual experiences a stroke, the incident might result in impairments to the motor cortex (body structure) and the execution of shoulder movements (body function), which might limit reaching movements for objects (activity), which might be a restriction when shopping for groceries (participation). Clinical outcome measures are often based on the observation of individuals’ behavior by a clinical evaluator [44]. Technology-based outcome measures are obtained from sensor data that are usually transformed (e.g., the spatial position of the wrist over time is transformed to movement speed or the number of movement units [45]. Moreover, technology-based outcome measures at the ICF activity level can be classified as capacity (i.e., maximal activity, performed in a controlled environment) and performance measures (i.e., activity in a real-life setting [39, 47, 48]). Outcome measures must fulfill strict criteria regarding validity, reliability, responsiveness, clinical utility, and expert consensus to be recommended [39, 44]. Widely recommended examples of clinical outcome measures of the upper extremity are the Fugl–Meyer Assessment for Upper Extremities

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(FMA-UE [49, 50]; function level of ICF) and the Action Research Arm Task (ARAT [51, 52], activity level of ICF). The ARAT is an example of a capacity measure. The FMA-UE covers 33 items that examine reflexes, voluntary movements, and movement synergies of shoulder, elbow, wrist, and fingers. Required materials are a reflex hammer and objects for movement task, like a sheet of paper or a tennis ball. Evaluators rate the degree of function for each item on a three-point scale. Administration time is about 30 minutes [53]. To provide a specific example, in one item of FMAUE individuals need to abduct their arm. The behavior of interest is the shoulder abduction function (which, after stroke, may be accompanied with undesired elbow movement). The ARAT comprises of 19 movement tasks, in which clients make reaching, grasp, grip or pinch movements, using standardized objects, like wooden blocks, marbles, pieces of metallic tubes, etc. Evaluators rate clients’ ability to perform each task and quality of movement on a four-point scale. Testing time is about 5–20 minutes [52, 54]. One task, for example, is to pick up a wooden ball from a table and put it onto a shelf that is placed within reach. Of interest is the entire sequence of movements. Clinical outcome measures at the ICF participation level are usually questionnaires that assess participation in daily life as perceived by clients [39, 43]. An example of a recommended questionnaire is the Stroke Impact Scale [43]. Clinical outcome measures, such as the FMAUE and the ARAT, are well established in clinical practice and have excellent psychometric properties [43]. However, they rely on direct observation, with client and evaluator being present at the same location. Hence, the question arises of how the assessment of motor functioning can be realized in remote settings. One solution is to video-record clinical assessments remotely and to evaluate the videos at a different location [55]. This method has excellent psychometric properties but implies additional time and material costs for cameras and the operation of standardized video-shooting procedures. Thus,

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for remote assessment, technology-generated outcome measures might be a more actionable solution. Technology-Based Clinical Assessment Technology-generated outcome measures are considered as an important complement to clinical outcome measures because of several advantages [39]: They are objective [45] (as they do not depend on the interpretation of an observer) and more sensitive [56] (as they yield continuous scale data). Thus, they can gather more detailed information about recovery which provides better information for the personalization of treatment regimina [56, 57]. Another advantage is that technology-generated outcome measures can distinguish between behavioral recovery and compensation, as they can measure both types of behavior separately, which extends information yielded by the ARAT [57, 58]. For the ICF function level, an expert committee recommended to employ a reaching task in the horizontal plane and capture kinematic outcome measures with an optoelectrical camera system [45, 59]. Optoelectrical systems use multiple cameras with infrared illuminators and triangulation algorithms to reconstruct the threedimensional (3D) position of reflective markers [60]. Markers are placed at defined anatomical landmarks of individuals’ body segments to measure their position and to derive kinematic measures. Position measurement of a marker has less than 1 mm error and measurement of joint angles has an error range of 1–degrees [61]. Alternatives are networks of IMUs and markerless camera systems (multiple high-quality video cameras and computer vision-based analysis). Compared to marker-based systems, these are slightly less [62] or similarly accurate [63], respectively. IMU networks are mobile, whereas markerless camera systems leave individuals’ bodies free of markers and sensor straps. However, both systems are similar to optoelectrical systems in price range and complexity of analysis. Hence, these systems imply similar barriers

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to the widespread use in clinical practice and at home as marker-based systems. Moreover, various force measurement devices are recommended to measure grip strength (entire hand), precision grip [64] strength (individual fingers), and assess finger individuation [65] (control of one finger independently of the others). An advantage of the assessments for hand and finger strength is that they are quick and easy to administer as compared to other clinical or technology-based assessments, and that muscle strength early after stroke is a good predictor in statistical models of motor improvement capacity [66, 67]. For the ICF activity level, an expert committee recommended a 3D reaching task (“drinking task”) and optoelectrical cameras for kinematic measurement [57, 68–70]. For this task, recommendations extend to the analysis of kinematic outcome measures [45], such as the number of movement units of the arm endpoint and angular velocity of the elbow, and trunk displacement and arm abduction angle [58]. Recovery, as opposed to compensation, corresponds to a smaller number of movement units of the arm endpoint, higher angular velocity of the elbow, less trunk inclination, and less arm abduction [57]. While the 3D reaching task assesses capacity at the ICF activity level, experts recommended to assess performance by monitoring arm use in real life, using measures like threshold activity counting [71], which can be obtained with accelerometers or inertial measurement units (IMUs) [39, 72]. However, a standardized way for application and analysis of this assessment is still missing and requires further research [39]. The described technology-based outcome measures are rather established in research than in clinical application. The equipment is expensive (about 10′000–20′000 USD), application is time-intense, and analysis is rather complex, especially in the case of optoelectrical camera systems [39]. Hence, their immediate application in remote settings appears to be impracticable.

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25.2.1.2 Tele-Assessment Low-cost sensor technologies, such as IMUs and depth-sensing cameras, and advances in videobased pose tracking algorithms are expected to enable technology-based assessment in clinical practice and real-life settings, once usability and standardization application and analysis have matured [42, 45]. Accelerometers and IMU sensors are the most frequently used wearable sensors in neurorehabilitation research [73, 74] (see Wang et al. for more types of wearable sensors). Accelerometers measure linear acceleration, gyroscopes measure angular velocity and magnetometers measure magnetic fields; an IMU is an assembly of these three components and captures the mentioned quantities in three orthogonal dimensions, which are usually aligned with the device’s housing [75]. Usually, 2–4 sensor units are used and placed at wrists or lower arms, upper arms, or the sternum [74]. Data are frequently used to estimate the orientation of the device in order to estimate orientation of body segments or joint angles [74], with errors lying in the three to eight-degree range [60]. Another technology for clinical practice and home-use are standard digital RGB (Red–Green– Blue) video cameras (e.g., like those of a webcam) combined with infrared depth sensors [76] (RGB-D). Algorithms use these video and depth data to estimate and track 3D joint positions. The position data can be used to estimate joint angles, which have a similar error to IMUs depending on the viewing angle [60, 77]. Pose estimation algorithms [78] are based on computer vision and deep learning, and only use videos of standard digital RGB cameras [79]. The algorithms detect and track joint positions and yield two-dimensional (2D) or 3D position data, depending on the capabilities of the algorithms. Validation studies showed that accuracy can be similar to that of depth-sensing cameras when the plane of body motion and camera viewing angle are aligned (e.g., trunk movement in the frontal plane is captured by a frontal camera [80]. Stereo cameras might be the key technology to provide accurate but low-cost motion sensing

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[81]. Stereo camera systems consist of two or more RGB video cameras. Depth information is inferred by the difference in an object’s positions in the two camera images. These low-cost technologies can be used to capture motion data and measure motor functioning at each ICF level [73]. Motion data are commonly used to classify the degree of motor function impairment or activity limitation [73]. Observation-based assessments such as the FMU-UE or the ARAT yield ordinal values. Supervised machine learning algorithms can estimate the observation-based scores from the motion data, captured during the clinical assessment tasks [82, 83] or during different functional tasks [84, 85]. For example, FMA-UE scores were estimated from sensor data that were recorded during eight functional motor tasks [85]. Hence, this approach may require fewer movement samples than FMA-UE or ARAT assessments [84]. A drawback of wearable sensors is that they require donning, while camera systems do not have this disadvantage [73]. Estimating clinical scores at the ordinal level has the advantage of linking motion data to validated assessments of motor functioning, but it does not take advantage of the continuous scale data that is provided by the sensors [73]. However, the use of continuous data requires the identification and psychometric validation of the kinematic metrics, as it is the case with the recommended outcome measures obtained with optoelectrical cameras [57]. As mentioned above, motor functioning at the ICF activity performance level has already been quantified by assessing arm motion in real-life, for several hours to several days [73]. Accelerometer data were used, with sensors being placed at one or both wrists [73], and measures as activity counts were employed [71] Activity counts measured whether acceleration magnitude was above a particular threshold [71]. In neurological clients, arm use of the more affected upper limb was compared with the less affected or with data from healthy individuals [71, 73, 86, 87]. Function can be defined as the functionality of a body structure as is the case in the ICF framework. Functional motion can be also defined as

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the goal-directed and volitional motion of the upper limbs in relation to an object or target during ADL [88]. One approach aims to detect this functional motion and identify its nature during ADL in real-life settings [73]. In these studies, functional motion is defined as the goaldirected and volitional motion of the upper limbs with regard to an object or target in ADL [88]. Functional motion covers functional primitives (e.g., reaching, transporting, repositioning) that can be combined into functional movements (e.g., drinking from a cup) and functional activities (e.g., having dinner). For this approach, usually, IMUs are used and placed at individuals’ wrists and/or various other positions [73]. First, several repetitions of each functional task are executed in a standardized setting and categorized by human observers. Then, supervised machine learning is used to estimate the category using the motion data [89–91]. This approach ultimately allows quantification of individuals’ functioning on all ICF levels, for example by counting the number of functional activity executions (activity performance) or by measuring movement quality within functional activities [89].

25.2.2 Teletherapy Neurorehabilitation research and guidelines suggest that recovery of motor functioning is positively influenced by interventions that follow principles of motor learning, such as high intensity and repetitive practice [92–97]. Recovery is positively influenced by the use of technology, as it facilitates the provision of therapy that follows these suggestions [39, 98]. In clinical application, rehabilitation technology may comprise of end-effector and exoskeleton rehabilitation robots linked to multimodal virtual or augmented reality environments [99]. Moreover, low-cost rehabilitation systems may be used that require less direct supervision by health professionals to deliver additional therapy in clinical and home-based settings [12, 98].

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Rehabilitation systems for home use can be broadly categorized as feedback systems, virtual reality (VR) systems, and robotic systems [100]. Feedback systems use human motion data to provide feedback about motor performance, primarily in real-life environments [74, 101], whereas VR systems use individuals’ motion data to provide interaction with artificially generated environments [102]. The systems may complement each other, as feedback systems are specialized in monitoring and feedback, whereas VR systems focus on training [103]. Robotic systems are usually seen as a separate class [100], even when combined with VR displays, because they comprise of extensive robotic apparatus [104].

25.2.2.1 Feedback Systems Feedback systems usually consist of wearable sensors (accelerometers or IMUs) on the wrist of the impaired upper limb and sometimes on other body segments (for reviews see [73, 74]). Motion data are used for the estimation of arm use intensity, recognition of specific arm movements, recognition of particular exercises, and movement quality [73]. These analyses yield feedback such as information about upper limb use, the number of completed arm exercises, or numeric values for the degree of compensation [74]. In several studies, feedback was provided in form of vibrotactile signals, as these do not require visual attention. For example, vibrotactile feedback was applied to the affected arm as a reminder to use the arm more frequently [105, 106]. In another study, visual feedback was provided on conventional laptops or tablet screens in form of numbers or graphs which showed summary information of relevant metrics [107–109]. Several studies provided evidence that feedback systems improve motor performance and suggest that feedback systems have the potential to motivate individuals to move, increase adherence, and support motor learning [105, 106, 109]. However, evidence is sparse, especially in form of RCT [74, 110]. Further research is required to explore the full potential of feedback

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systems, notably with respect to systems that provide feedback during motor performance [73].

25.2.2.2 VR Systems VR systems use data from motion sensing technologies to allow interaction between individuals and virtual environments [102]. The rationale behind VR systems is that immersion and use of games motivate individuals to practice [111]. VR systems can be classified by the type of motion sensing system and display technology [111]. Motion sensing systems (for lists of systems see [104, 112]) can be grouped into desktop (e.g., joysticks), body-mounted (e.g., IMU systems, sensor gloves, head-mounted displays), and contact-free systems (e.g., camera-based systems). Individual motion sensing systems may contain various motion sensing technologies (e.g., force sensors, pressure sensors, IMUs, depth sensors). Display technologies of VR systems can be categorized into visual (e.g., 2D flat panel displays), auditory (e.g., two speakers for stereo sound), and haptic displays (e.g., vibrotactile actuators, force feedback [102]). VR systems differ in the degree of immersion and, hence, can further be classified as nonimmersive systems (e.g., systems with conventional tablets) and immersive systems (e.g., 3D stereoscopic vision head-mounted displays with surround sound [104]). Immersion in this definition is a technical quality and refers to the capability of systems to simulate the real world and generate authentic virtual experiences [113, 114]. Individuals may experience the virtual environment from a first- or third-person view. Motion data of individuals’ body segments or joints are used to enable synchronous action in the virtual environment. Examples of commercialized VR rehabilitation of the upper limbs can be found in a recent review [104]. An umbrella review of the meta-analysis concluded that there is evidence of a benefit of VR on motor function, but that evidence is weak as study quality is low or very low [114]. According to the review, important moderating factors for a benefit

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of VR are the quality of immersion, interactivity, and customization of VR environments. Immersion and interactivity are important because differences in perceived distance between VR and real life, as well as system delays, might hinder the transference of skills from VR environments to real life [115]. This is critical as clinical evidence for the transfer of skills from VR training to real life is inconclusive [116]. Customization of VR environments is an important capability of VR systems as individuals with motor deficits might also have deficits in perception or cognition (e.g., individuals with attention deficits might require environments with reduced distraction [103]). Thus, VR systems engineered for neurorehabilitation are preferable to commercial VR gaming solutions that rather are designed for able-bodied individuals [114]. Several reviews examined the efficacy of VRbased training in home-based settings [103, 117]. Across reviews, only a few randomized controlled trials were identified. These generally showed a benefit of VR telerehabilitation that is comparable to conventional therapy, but the numbers of participants per study were low, VR Systems and intervention protocols differed strongly, and there was a wide variety in outcome measures. One RCT [118] that observed a benefit of VR-based telerehabilitation (and demonstrated noninferiority to conventional face-to-face therapy) applied several techniques to influence therapy outcomes such as behavioral contracts, stroke education, high ease of use, many input devices to cover a large range arm movement functions, frequent interaction with clients, multiple means of providing client feedback, solutions for creating appointment, and reminders. In several clients, this led to more than 1000 arm movement repetitions per client and day. This points to the conclusions that many aspects of therapy require optimization to obtain the desired doses of practice and beneficial effects on arm movement recovery.

25.2.2.3 Robotic Systems There is a large and quickly increasing number of rehabilitation robots and robotic devices for the upper extremities (a recent review lists

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commercial devices for home use [119]). Robotic devices can be classified into grounded endeffectors, grounded exoskeletons, and wearable exoskeletons [120]. End-effector devices are usually only attached to individuals’ distal arm parts, such as the hands or fingers. Exoskeletons target one or more joints of a paretic limb and are attached to adjacent segments. Systems can be further divided into passive, active, and interactive systems, depending on whether they merely stabilize, actively control, or react to individuals’ movement input (for further classification see [121]). Systems are often equipped with sensors and, hence, can measure kinetics and kinematics that can be used to control VR-based exercises. Robotic devices for home use mainly target wrist, hand, and fingers [120]. Figure 25.1 shows an example of an interactive end-effector robot, the ReHandyBot, a portable robot for hand rehabilitation [122]. In clinical settings, robot-assisted therapy resulted in similar or larger improvements in upper extremity motor function as conventional interventions [123–125], regardless of the type of robotic arm training device [126]. Furthermore, for home-based settings, studies also suggest a benefit of robot-assisted therapy that is similar to conventional therapy, but the number of studies is rather small [127]. Drawbacks of robotic systems are their obtrusiveness, low comfort, and high costs [128, 129]. A recent trend are soft robotic gloves [130]. Soft robotic gloves are wearable and can facilitate hand and finger movements during activities of daily living or while playing VR-based serious games. Many studies report beneficial effects on motor recovery in clinical and home-based settings [130].

25.3

Lower Extremity

While upper limb impairments are common after stroke, the same is true for gait impairments. Over 80% of stroke survivors demonstrate a gait impairment [131] that recovers to some extent in the first two months after stroke [132]. Despite this, community ambulation often remains compromised in most survivors [133–135]. Gait

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ability specifically has major implications for health; it is an essential predictor for functional independence [136, 137] and long-term survival [137, 138] after stroke, along with participation and quality of life [139, 140]. Regaining gait ability is hence one of the most frequently prioritized goals of stroke survivors [131, 141].

25.3.1 Assessment of Motor Function 25.3.1.1 Clinical Assessment Observation-Based Clinical Assessment Gait ability can be separated into two major domains: Functional gait ability and gait quality. Assessments commonly employed in an in-clinic environment typically target the first domain. Best practice guidelines [43, 44] include the 10m walk test (10MWT) [142], 6-min walk test (6MWT) [143], and timed-up-and-go test (TUG) [144], as they provide a complementary representation of a client’s functional gait status and are feasible to implement clinically. These tests are commonly performed throughout stroke rehabilitation, albeit with a variability in terms of protocol and also time/distance criteria [145]. The 10MWT aims to measure maximal gait velocity and reflects the maximal capacity that a client has at a given point in time. For stroke survivors throughout all phases of recovery, the test demonstrates excellent sensitivity (MDC90 = 0.34–0.1 m/s from acute to chronic, respectively), and reliability (ICC = 0.83–0.95 from acute to chronic, respectively [145–147]). The MCID for this test in subacute stroke is estimated between 0.16 and 0.22 m/s [148]. The 6MWT aims to measure gait endurance. Equally, the clinometric properties of this assessment are well understood, and it demonstrates excellent reliability (ICC = 0.97–0.99 from acute to chronic, respectively) and sensitivity (MDC90 = 52 - 28 m from acute to chronic, respectively) [149] in stroke survivors throughout their recovery [150]. In terms of construct validity, the 6MWT correlates significantly with aerobic capacity, mobility, walking speed, strength, balance, and participation [150].

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MCID for slow walkers in the subacute phase is estimated at 44m [148]. The TUG represents a compound test that measures the ability to perform sequential motor tasks relative to walking and turning [144, 151]. In this capacity, it can be considered a representation of complex movement [147]—one of the four pillars considered necessary for successful community ambulation [139]. In terms of clinometric properties, the TUG is considered reliable (ICC = 085 - 0.96) [152] and measures with an MDC95 of 2.9–3.5 s [147, 153]. Technology-Based Clinical Assessment The second domain of gait assessment—gait quality—is not typically breached in clinical routine [154]. Procedures informing on this domain, in order of complexity, include observational gait assessments, pressure mat recordings, monocular video recordings (RGB/RGBD), inertial-measurement-unit-based recordings, insole-based recordings, multi-camera video recordings, and 3D motion capture with force measurements. The objective and time-efficient quantification of gait quality has long carried the promise of identifying key deficits and guiding therapeutic intervention, however, this has yet to manifest on a large scale in clinical settings [155]. The barriers to large-scale adoption lie mainly in the cost, complexity, training requirements, and unwieldiness of current technology [156]. Advances especially in two domains carry the potential to overcome these limitations and have been frequently evaluated: Improved spatial reconstruction algorithms for IMUs, and improved computer vision pose estimation [157]. Typically, gait quality is described in three domains: Spatio-temporal metrics that describe the spatial and temporal characteristics of the endpoint within the resolution of a single gait cycle (e.g., step length, stance time, double support phase), joint kinematics (time-series of angles between segments that span from one gait event to the next), and kinetics (time-series of the ground reaction force for each limb) [158]. Gait events are considered when the foot strikes the ground (Foot-strike) and when the foot leaves the ground (Foot-off) and are used as anchors to

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normalize all cycles to a common timescale [159]. For brevity, this chapter will forgo technologies that are not translatable to the home environment, e.g., force plates. The clinical utility of metrics derived within these domains is not yet well understood and pathology- and deficitdependent [160, 161]. Camera-based systems are typically best suited for pre-defined walkways or treadmill situations [161], where a large amount of gait cycles can be collected within a small movement volume. For gait, the same optical technologies introduced above are applicable for the calculation of joint angles and segment positions during walking: Multi-camera RGB and RGB-D setups, and single-camera RGB and RGB-D. The benefits and drawbacks are equally largely comparable to the application for the upper extremities, however for walking, the capture volume and movement amplitude are larger, and occlusions are less frequent. Compared to optoelectrical 3D motion capture, the accuracy for multi-camera RGB systems using contemporary pose estimation networks is in the range of 2–3° joint angle deviation [162] and 1–15 mm deviation in 3D joint center position estimation [80]. Multicamera systems remain complex, as the camera locations must be calibrated relative to each other before measurement. Moving to a single RGB-D camera with an added depth channel results in a reduced accuracy between 5 and 6° for joint angles [163], however, this is at least partially dependent on the camera placement [77]. This is equally applicable when using 2D monocular RGB video to estimate 3D pose [164–166], where out-of-plane errors in the major movement directions can be reduced by strategic camera placement [167]. In synthesis, these values compare favorably to estimated MCIDs for lower limb angles in stroke survivors, which range from between 3 and 9° for the major lower extremity joints and planes [168, 169]. Associated, nascent technologies that may enrich the industry are event-based cameras [170], 2D LiDAR sensors [171–173], or moving camera setups [174–176]. These technologies offer various benefits (high framerate, reduced setup complexity, and increased capture volume,

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respectively), however, have not yet reached a level of maturity and accuracy for gait analysis that is sufficient for clinical application. Inertial measurement technology provides a series of additional benefits compared to camera solutions, that are especially relevant to gait. The measurement becomes location-independent and can be performed for long distances overground. Individual measurement units however need to be attached to the subject and calibrated, leading to additional setup time and potential inaccuracies. In a recent review, Poitras et al. report accuracies for joint angle estimation in contemporary commercial systems that are slightly lower than camera-based systems (Hip: < 9.3°; Knee < 11.5°; and Ankle < 18.8°) [177]. However, concerning spatio-temporal parameters [178], IMUs demonstrate accuracy sufficient to describe clinically relevant metrics such as stance duration and asymmetry within the limits of meaningful change after stroke [179–183].

25.3.1.2 Tele-Assessment Although the advances in technology-based assessments bring the flexibility to perform gait assessments outside clinical environments and without supervision, the link between in-clinic and assessments performed in clients’ domestic environments is not yet well understood. The International Classification of Functioning Disability and Health (ICF) suggests that there is a difference between these two settings [46]. The measurements that are carried out in standardized environments such as the clinic reflect the best performance of the clients or their capacity and the assessments that are carried out during daily activities are more representative of the clients’ actual performance. Within this thought framework, it can be imagined that a 1:1 translation of clinical fixed-time/distance tests to home rehabilitation may not be trivial, especially concerning functional gait ability. Concerning gait quality, comparisons between clinic and home have been performed using both IMU and camera-based systems [184].

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In these comparisons, the accuracy of the technology remains comparable with the clinical setting, however usability and feasibility challenges must be considered. For cameras, an important aspect is the capture volume: Treadmills are rare in client’s typical homes. Furthermore, the calibration of multiple cameras is a time-consuming and painstaking process that is not feasible for clients to perform autonomously. Fixed-camera installations that require minimal calibration are costly to install and require building modifications. Home-recording technologies are thus practically limited to monocular RGB/RGB-D video and stereoscopic cameras that house the sensors in a single discrete unit. In either case, due to the limited field of view camera-based gait analyses are limited to very short walkways or discrete lower limb movements performed on the spot [185]. As a complicating factor, the nonstandardized scenes as backgrounds which are common in home applications can confound the tracking algorithms, leading to decreased accuracy [186, 187]. The practical use of cameras for gait analysis in home settings is hence quite limited and confined to a dedicated space. For IMUs, the limitations concerning the volume of capture do not exist. As these systems are fully portable, they cannot limit the assessment space. However, IMU-based systems also bring challenges that are not yet fully addressed [188]. These are mainly in the usability spectrum, for example, the client-friendly attachment of IMUs to skin/clothing while retaining sufficient data quality, or zero-interaction data transfer and charging solutions [189]. Novel solutions, such as seamlessly including sensing units into shoes with wireless chargers [190], can be expected to provide solace to these usability barriers as costs to components fall and new form factors become available. Finally, many IMU systems require a calibration step in which predefined movements are necessary which may not be possible for a given client or which may lead to recording errors if performed too rapidly or erroneously.

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Having scoped the limitations of each technology, there remains the question of which activity to record in a home environment, especially in the context of gait quality. Optimally, the same time/distance tests as performed in the clinic would translate seamlessly to the home environment for both functional gait ability and gait quality. Practical considerations, however, put the use of these tests into question. Specifically, few home environments provide the space necessary for a 6MWT (per clinical definition a 40 m straight walkway), nor even for a 10MWT with acceleration and deceleration phases. These tests are then performed outdoors, which complicates the question of a standardized setting with even surfaces [191]. Furthermore, there remains the question of how to initiate and end recordings for time/distance-limited tests. User interaction at the extrema of the tests invariably leads to a dual-task situation, therefore recording should best be triggered automatically. In the case of time, this is relatively trivial relying on a simple countdown timer. In the case of distance, however, the cumulative error in real-time spatial reconstruction of an IMU can lead to a roughly 5% error in the distance estimation [192]. A 10MWT for instance could be 9.5 m or 10.5 m long, hardly sufficiently reproducible for a reliable clinical outcome. Augmenting IMUs with other technology such as GPS can be helpful in improving the distance estimation [193], however, this is limited to longer tests such as the 6MWT and is strongly sensitive to satellite coverage which can limit its applicability in large cities [194]. The question here arises, whether there is a need for fixed-time/distance tests, or whether the continuous data collection during a portion of a day provides sufficient gait data to either extract representative test segments [195] or predict test performance [196–198]. Continuous data collection or “gait monitoring”, holds the promise to also generate novel digital mobility outcomes (DMOs) [199–201] that can enable greater insight into a client’s status. A plethora of algorithms are available to extract sequences of interest, such as gait, from the continuous data stream [202, 203]. From the current technology situation and a clinical utility perspective, it is unsurprising that the

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state of the art is gyrating towards continuous monitoring of gait using IMUs [204–207]. Open questions that remain are the location of the sensors (common are shoe [180], ankles [208], pelvis or pocket, and wrist [209] or a combination of these) and the form factor (dedicated gait monitoring hardware, smartphone sensors, or smartwatch sensors) [210].

25.3.2 Teletherapy Concerning teletherapy, applications focus either on gait volume and intensity or on gait quality. Volume and intensity are collected throughout the day and presented as summative feedback on demand via web or tablet applications [211], similar to what is present in the lifestyle and wellness segment. Indeed, significant effort has been expended in evaluating lifestyle and wellness wearables for gait monitoring in elderly and clinical populations [212–216]. Applications for gait quality are typically proposed as gait training movements in a stationary setting [13, 217]. This can be performed asynchronously [218] or synchronously [219] with a health professional. There are countless exergaming-like applications that use IMU and/or camera systems coupled to visual displays that provide training programs and movement feedback in various levels of detail [220–224]. This segment is evolving rapidly, accelerated by readily available open-source and lightweight pose estimation algorithms for 2D RGB systems. An overview of commercial offerings would exceed the scope of a single chapter and probably be outdated before the book is printed. A few research studies [225, 226] have engaged in the challenge of providing continuous real-time feedback on gait quality during ecological activity [227, 228]. While biofeedback is a promising approach in clinical settings [229, 230], challenges remain that limit the transfer to real-world scenarios. Outside of technical and usability limitations such as mounting points and battery capacity, the challenges here entail the appropriate detection of movement context [231] and the careful selection of relevant movement features and feedback

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modalities and mapping [101, 232, 233]. This area is expected to demonstrate significant innovation in the near future [234–239]. Mixed reality (MR) is a promising pathway for gait training scenarios (Fig. 25.2) [240, 241]. Compared to the previously described VR systems, MR enables the client to mix digital and real worlds through different technologies [242]. One approach for MR environments uses a video stream of RGB video cameras mounted to a headset to display a video of the environment onto a closed head-mounted display (HMD). The environment can be augmented with digital creations that are anchored in physical space, appearing in the HMD along with the video feed. These systems can be highly immersive; however, lag of the video feed and a corresponding mismatch of vestibular and visual information is still a challenge [243, 244]. To compensate for this, video resolution can be scaled down, resulting in a visually jarring imperfect representation of the surroundings. A further issue is the question of video white balance, which is adapted automatically to the lighting of scene and can lead to jumps in brightness of the video stream [245]. Equally open is the question of the matching of light source positioning and environmental situations between virtual and real objects [245]. As technology progresses on the fronts of mobile processing and graphics performance, camera, and display technologies, these issues may be resolved. A second approach uses translucent displays to superimpose a hologram image into the real world [246]. This elegantly circumvents the questions of lag and environment resolution as only the digital portion of the mixed reality needs to be rendered. There is rapid development in this sector, however, the field of view remains limited leading to a low level of immersion. In both cases, the usability of the device for remote deployment to clients [247, 248] and realistic multimodal feedback [249, 250] are remaining challenges that a new generation of rehabilitation-friendly devices should address. The perspective is that MR provides a pathway for highly promising VR-based training interventions from the clinic to extend to a client’s home [16, 251, 252].

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In synthesis, the technology situation is highly favorable to expand on the solutions currently available for gait in home environments. Improvements in device form factors, usability, measurement accuracy, presentation quality, and visualization of gait data synergize to unlock powerful new insights and provide training over a much longer rehabilitation period than previously feasible.

25.4

Communication

Effective communication in neurological conditions requires intact language, motor speech, and voice production. Aphasia is a language disorder where the use and the comprehension of meaningful speech are affected while motor speech disorder refers to a condition where a person has problems creating or forming speech sounds needed to communicate [253, 254]. In neurological voice disorder, the coordination or strengths of muscles that are needed to produce phonation are affected. All three conditions can occur after stroke, both, independently of each other or in combination, leaving roughly a third of all stroke survivors with some sort of communication impairment [255–257]. These impairments have a great impact on those affected as the majority of labor force depends nowadays on communication skills [258] but also social participation and independent living are dramatically affected by impaired communication. Clients with aphasia after stroke “described intense feelings of frustration, hopelessness, isolation, and depression at not being able to talk” [259]. Telerehabilitation is nowadays attracting broad interest in the field of speech and language therapy. For successful telerehabilitation, there is a need for optimal technology use in both, synchronous and asynchronous therapy delivery. To guide therapists along the recovery path, reliable and meaningful remote measures of voice, speech, and language functions are essential. Besides the direct need of valid measures for therapy delivery, communication measures might also serve as a general indicator of well-being

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Fig. 25.1 The ReHandyBot (picture: Rehabilitation Engineering Laboratory, ETH Zurich/Stefan Schneller). Fingers and thumb are fixed to separate grippers. Grippers’ movement is coupled which allows training hand opening and closing. Moreover, assistive or resistive forces allow haptic feedback and interaction with virtual objects

including emotional load and measure of participation, or early sign of cognitive decline [260, 261].

25.4.1 Assessment of Speech and Language Functions 25.4.1.1 Clinical Assessment Observation-Based Clinical Assessment Clinical assessment of functional voice, speech and language impairments consists of a wide range of measurement instruments. They largely depend on the perceptual evaluation by professional speech and language pathologists or task-

specific performance measures (dysarthria [262], aphasia [263]). Many of these measurement instruments may not be specific to the stroke population but are used in the respective conditions (i.e., voice, speech, or language impairments) independent of the underlying disease. Most of the functional assessments related to communication are obviously highly languagespecific. In the following paragraph, we will give only a few examples of voice, speech,and language scales in English or German to point out the limitations of remote assessments. The German Bogenhausener Dysarthrieskalen (BoDyS) is a tool to assess voice and speech impairments in different tasks such as picture description [264]. For each task, the speech and

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Fig. 25.2 A client after stroke performing a mixed reality gait training parkour with a holographic head-mounted display. Obstacles that require specific gait adaptations are placed throughout a given physical environment, transforming the space into a virtual rehabilitation gym (from [240])

language pathologist rates different symptoms perceived, such as hoarseness or slow talking speed. Even though audio recordings for potential algorithmic analysis are created during the assessment, the analysis requires manual scoring of voice and speech pathologies. An expert panel has recently defined a core outcome set to assess language impairments in aphasia research [263]. It includes the Western Aphasia Battery-Revised (WAB-R), a diagnostic tool widely used by clinicians to assess language skills [265]. In German-speaking countries, the Aachener Aphasie Test (AAT) is a widely used diagnostic tool for differential diagnosis and follow-up assessment of language impairment [266]. Scales were also developed to assess communication in a more general term. No consensus, however, was reached by Wallace and

colleagues on which measures to use in a core outcome set for research, and the discussion on how to best measure communication is ongoing [263]. Measurement instruments related to participation—especially important once a client is at home—are usually complemented by selfreported outcome measures and questionnaires related to emotional well-being and quality of life —measures which seem more meaningful to assess the impact of a disorder on a client’s daily living, but may not give much information on the underlying impairments [263, 267]. Taken together, many of the clinically used assessment tools strongly rely on subjective observational methods, i.e., speech and motor behavior are observed by a professional speech and language pathologist and require (rather timeconsuming) manual scoring. There is a call for

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more technology-dependent outcome measures which may help to understand underlying impairment and treatment effects as stated by Elisabeth Armstrong in her commentary article: Let’s utilize technology to the maximum and contribute our expertise at the highest level, while at the same time not simply counting what can be counted and neglecting “what counts” [268]. Technology-Based Clinical Assessments Voice Analysis

Instrumented acoustic voice analysis is a quantitative noninvasive method to complement a multidimensional approach for evaluation of initial impairment and the recovery/therapy progress [269]. Over the past decades, several acoustic measures have been proposed to be sensitive to (overall) voice quality or specific perceptual dimensions such as hoarseness or breathiness ([270] and other reviews). In standard acoustic voice assessments, a voice profile (vocal intensity and pitch measures), frequency and amplitude perturbation (jitter, shimmer), and harmonicity measurements (harmonic to noise ratio) are usually determined that provide information about the underlying functional impairment and recovery progress [270, 271]. The feasibility and validity of such metrics highly depend on the client's task (e.g., sustained vowels vs. reading) and instructions (e.g., comfortable loudness vs. maximal loudness) as well as room setup and technical specifications (e.g., microphone quality, setup, and calibration). An expert panel of the American Speech-LanguageHearing Association has recently approached this issue and published a standardization protocol defining instructions and specifications for qualitative acoustic voice analysis in great detail [269]. Even in a very standardized setup, acoustic parameters are highly variable due to dependence on the vocal intensity used [272]. Most of the traditional perturbation measures depend on the computation of the fundamental frequency which may not be reliable in severe dysphonia. For this reason, Patel et al. further

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suggested the calculation of cepstral peak prominence, a measure based on spectral analyses for estimating general dysphonia severity and overall breathiness from connected speech, i.e., during a reading task or a conversation [269, 273–275]. Several computer programs exist which enable instrumented voice analysis also in a clinical setting. One example used in research and clinic is the comprehensive open-source program PRAAT, which was initiated and continuously developed by Jan Boersma and David Weenink from the University of Amsterdam (http://www.PRAAT.org/). Many commercially available voice analysis software exist that come with technical equipment that allows for rapid voice assessment without requiring complex calibration and in-depth technical knowledge. Speech and Language Analysis

Similar to acoustic voice analysis, certain speech and language parameters can be automatically calculated from audio files using algorithms integrated into voice analysis software such as PRAAT. Such parameters may include speech rate and overall talking time, but also more complex estimates of fluency extracted from connected speech [276, 277]. Fully automized analysis of audio samples to assess language functions is often limited to macroparameters. More relevant [278, 279] linguistic parameters such as lexical diversity, mean utterance lengths, repetitions, and selfcorrections can be extracted from transcripts with the aid of specialized speech analysis systems [279]. Examples include the “Aachener Spontansprachanalyse” [280, 281] for German or the “Computerized language analysis” (CLAN) for English [282] which, once the audio sample is manually transcribed, assists with word tagging and linguistic analysis. More contentdependent discourse measures can be assessed using natural language processing by computerized systems [283, 284]. As part of the Aphasia Bank Project, large amounts of standardized speech samples (e.g., the English story-retelling task “Cinderella”)

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from aphasic clients were—and still are—collected, transcribed, and made available to researchers for in-depth discourse analysis [285, 286]. This Big Data approach contributes to the understanding of discourse differences in mildly impaired stroke survivors that may not with scales such as the WAB-R due to ceiling effects [287]. Even though computer-assisted methods facilitate manual transcription, the process is still time-consuming and thus often not feasible in daily clinical practice. Advances in Automated Speech Recognition (ASR)—the automatic conversion of audio files into written text [288]— could solve this problem in the future. A recent study suggests that ASR could be used in a clinic setting for aphasic clients. However, the quality of transcripts needs to be investigated separately for each communication disorder [289].

25.4.2 Tele-Assessments In a remote setting, one way to assess speech and language functions may be to apply standardized assessment batteries or scales via video call in a synchronous session with a speech and language pathologist. This way of applying assessments is limited by the availability of analog equipment on the client’s side, and for many tools, a fully computerized adaptation is still lacking. Furthermore, the validity of such teleconferencebased application of standardized test batteries needs to be investigated—as has recently been done for the widely used WAB-R [290]—and standards need to be developed [291]. Overall, remote assessments via teleconference systems are feasible, but a revised computerized version rather than ad hoc digital adaptation of valid analog tools would be highly desirable. Recording audio samples during a teleassessment via video call for automated analysis is possible, however, signal processing circuits such as automatic gain control or noise cancellation may modify the original microphone signal [269, 271]. Setup and implementation at a

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client’s home may not (easily) be controllable. Audio samples may still be used for synchronous or asynchronous manual scoring of voice, speech, or language impairment or transcription. Some automated tele-diagnostic instruments are also available for web-based administration or use with smartphone applications. Smartphone applications have been developed for selfadministration of acoustic voice assessments in clinical telepractice settings which can produce valid calculations of acoustic parameters [292]. As described above, a highly standardized setup in a clinical setting is crucial for valid acoustic voice assessments. It is thus questionable whether such smartphone-based audio recordings achieve the quality to successfully guide teletherapeutic interventions. For speech and language, there are applications mainly used for research purposes which allow remote application of certain standardized speech and language assessments (e.g., [293]).Tele-monitoring. Using microphones for monitoring voice, speech, or language functions in daily living is feasible. However, it faces several limitations: Privacy, environmental noise, recording quality, and interference with speech from communication partners. One solution for the privacy issue is to only store information that does not need content recording, such as overall talking time, as done in people with aphasia after stroke [294]. However, this metric does not seem to provide much information about functional recovery as it lacks correlation to impairment level but might be an indicator of participation [294]. Accuracy of measuring talking time is highly dependent on environmental noise. Researchers have started to address this limitation by measuring voice functions via neck surface vibration measures with a single accelerometer placed on the subglottal neck surface [295]. Relevant metrics such as fundamental frequency and sound pressure level as well as spectral parameters can be captured through skin acceleration and correlate for the most part with the parameters received from acoustic measures through a microphone [275, 296, 297].

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25.4.3 Teletherapy 25.4.3.1 Synchronous Therapy Many therapists, clinics or specialized telerehabilitation providers offer speech and language therapy via video conferencing [298], also called telepractice [299, 300] (ASHA 2018). Some use available video call platforms such as Webex or Zoom with standard video call features [301, 302]. Others may use specialized speech and language tele-therapy platforms with integrated digital exercises or social functionality [303]. Such synchronous speech and language therapies were found to have similar effects as in-person therapy (reviewed by [304]). With increasing functionality such as joint editing of digital material and integrated audio recordings, an increased range of therapy content may be provided in the future. A high-quality microphone and high-speed internet connection are crucial, especially for vocal tele-therapy, and technical hurdles might limit use in the elderly. There are, however, also advantages of video-based synchronous over inperson therapy. Group therapy and group chats for people with communication problems are easily implementable. Group therapy was found to be a valuable and cost-efficient way of delivering therapy. Group chats that socially connect people with similar impairments, medical history, and interests have a positive effect on participation and well-being which in turn can improve specific linguistic effects [305–307]. Another advantage is the high context specificity in which the therapy takes place, allowing speech and language therapy to be tailored to a client’s functional environment [308]. 25.4.3.2 Asynchronous Therapy Many commercially available computer programs exist which have digitalized logopedic exercise material and can be used by clients for independent functional training at home. Commercially available programs, such as Tactus Therapy Solutions, Lingraphica, Constant Therapy, or

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Neolexon are built to supplement cliniciandelivered therapy; the therapist creates a personalized training program and has access to training results or audio recordings via a specialized therapist interface. In many exercises, difficulty levels are automatically adapted based on errors and response time, allowing training without constant evaluation and exercise adaptation by the therapist. Automated detection of errors is mainly available for written input. Thus, any form of automated result or performance feedback is mostly restricted to receptive exercises (i.e., the correct answer can be chosen among suggestions) or to written content. Advances in automatic speech recognition will open the path to more independent verbal training in the future [309]. Feedback during voice training does not rely on speech recognition and helps to adapt vocal behavior—namely pitch and/or loudness—during training [310]. The use of neck vibration monitoring system enables biofeedback during daily living and promotes changes in vocal behavior [311, 312]. Retention of the adapted behavior, however, seems to be strongly dependent on the type and frequency of feedback used [312]. Taken together, such computerized asynchronous therapies are usually well received by stroke survivors and seem a valid option to increase training time, even though efficacy may not exceed in-person therapy effects [313] (see [314] for a review). The field of telerehabilitation is still in its infancy and technologies are developing rapidly. Comprehensive combinations of synchronous and asynchronous options might be more successful in the future. The Australian Tele-CHAT (i.e., Comprehensive High-dose Aphasia Treatment) program is an example of a combination of synchronous, asynchronous, and group therapy and it also supports education and social exchange among stroke survivors. Feasibility and acceptability of this program is still under scientific investigation, but it seems clear that for an efficient and effective therapy delivery, the use of different therapy modes needs to be optimally coordinated.

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25.5

Final Remarks

Technological solutions are expected to become an integral part of telerehabilitation. These should be complemented by data platforms that consolidate the diverse streams of information into a coherent and actionable whole. These data can be leveraged by healthcare professionals to gain unprecedented insights into recovery profiles, individual therapeutic needs, and requirements for sustainable technological solutions. AI is already demonstrating some promising performance in specific medical tasks. However, the details of decision-making of the algorithms are often poorly understood, and there is still no robust interpretability and explainability, thus hindering the widespread use of AI in medicine. Ongoing developments of methods for visualization, explanation, and interpretation of AI models aim at making AI transparent and explainable (XAI), and thus robust for application in the medical field [315]. In the future, XAI may allow to target and scale interventions based on a large amount of data precisely and reliably to the needs of the individual. Clinical trials show promising results with telerehabilitation. The Corona pandemic has promoted the acceptance of telerehabilitation among therapists and clients. To promote widespread and long-term implementation in routine clinical practice, it is up to stakeholders to increase adoption through measures on telecommunications infrastructure, legal certainty around digital health policies and legislation, data security, technical support, teaching, and reimbursement. The combination of increasing digital literacy, scalability requirements of the neurorehabilitation industry, regulatory changes, and an increased enthusiasm for new technologies are creating a breeding ground for tele-neurorehabilitation to gain wider acceptance and fully deliver on its promise. In the long term, telerehabilitation could become a ubiquitous tool that enables access to health care and intense, personally optimized neurorehabilitation regardless of geographical location.

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Note that this book chapter primarily focused on technology and services for assessment and treatment at the levels of ICF function and activity. Future work will certainly also cover clients’ functioning at the ICF participation level, especially because improvements in quality of life and participation in social activities are the higher-order goals of neurorehabilitation. Thus, technology and service design will increasingly need to incorporate clients’ social and environmental contexts. Barriers in social contexts can be reduced, for example, through simulation of clients’ motor limitations to healthy individuals, using VR and robotic systems [14]. Healthy individuals can thus experience and gain an intuitive understanding of clients’ motor limitations [14]. Moreover, caretakers can aid clients to leverage smart homes and smart cities to reduce barriers to social context and increase engagement in these. For example, smart door locks might reduce the stress of clients when entering new social interactions [316] and social games in smart cities can engage to move and interact with others [317]. Hence, as we press on into the future, new technologies and services will become increasingly pervasive in our everyday lives. These technologies can be leveraged for rehabilitation at a distance and might enable a more direct impact on participation than previous approaches.

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Part VI Robotic Technologies for Neurorehabilitation: Upper Extremity

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Forging Mens et Manus: The MIT Experience in Upper Extremity Robotic Therapy Hermano Igo Krebs, Dylan J. Edwards, and Bruce T. Volpe

Abstract

MIT’s motto is “Mens et Manus” (Mind and Hand) and we have adopted it as the guiding rule (principle) for our line of research: using robotics and information technology to re-connect the brain to the hand. Training and treatment protocols enhance this re-connection phenomenon, reduce impairment, increase function, and improve the quality of life beyond natural recovery. This chapter describes our efforts towards attaining this goal since the initial development of the MIT-MANUS in 1989. Numerous clinical trials involving thousands of participants working with (receiving therapy using) different versions of the MIT-Manus have been conducted since then and we have created a complete robotic gym for the upper extremity. In fact, for over 10 years, the American Heart

H. I. Krebs (&) Massachusetts Institute of Technology, 77 Massachusetts Avenue, Office 3-155, Cambridge, MA 02139, USA e-mail: [email protected] D. J. Edwards Moss Rehabilitation Research Institute, Elkins Park, USA e-mail: [email protected] B. T. Volpe Feinstein Institutes for Medical Research, Manhasset, NY, USA e-mail: [email protected]

Association and the Veterans Affairs/Department of Defense endorsed the use of robot-assisted therapy in stroke rehabilitation for upper extremities, and we have been focusing on how to tailor and augment therapy to a particular patient’s need and in determining who is a responder (and non-responder) to this kind of intervention. Keywords

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Rehabilitation robotics Robot-assisted therapy Robotic therapy MIT-Manus Upper extremity Neuro-modulation Stroke

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Introduction

The use of robotic technology to assist recovery after a neurological injury has proven to be safe, feasible and effective—at least to reduce impairment—for the upper extremity of stroke participants. Nevertheless, there is vast room for improvement. But what is the best way to pursue further improvement? Ultimately, we would like to prescribe customized therapy to optimize and augment a patient’s recovery. In this chapter, we review our experience developing upper extremity robotic therapy and applying it in clinical practice. Based on that experience, we propose the most productive way to refine and optimize this technology and its application. Needless to say, this personal viewpoint will almost certainly neglect or under-emphasize important developments; however, that should not

© The Author(s), under exclusive license to Springer Nature Switzerland AG 2022 D. J. Reinkensmeyer et al. (eds.), Neurorehabilitation Technology, https://doi.org/10.1007/978-3-031-08995-4_26

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be construed as a dismissal of other work but more as a symptom of the explosive growth of research in this field. Despite its inevitable limitations, we trust our perspective may have value. This chapter is organized as follows. We first discuss the expert consensus expressed in the American Heart Association guidelines (AHA) and introduce the MIT-Manus robotic gym (Sects. 26.1.1 and 26.1.2). We then discuss some of our evidence with over 1,000 stroke participants in both sub-acute and chronic phases of recovery that are in line with AHA’s guidelines (Sects. 26.2.1 and 26.2) and what we believe to be the best way to employ robotic technology (Sect. 26.2.3). We also discuss a few misconceptions that have been experimentally debunked including some approaches that failed to augment outcomes on robot-mediated therapy (Sects. 26.2.4. thru 26.2.6). We review some of the robot control schemes employed in our studies and which one promotes the best outcomes (Sect. 26.2.7), as well as the potential of robot-mediated assay to evaluate recovery (Sect. 26.2.8). We wrap-up this chapter with a short discussion and some closing remarks (Sect. 26.3).

26.1.1 The State of the Art The 2010 American Heart Association (AHA) guidelines for stroke care recommended that: “Robot-assisted therapy offers the amount of motor practice needed to relearn motor skills with less therapist assistance. Most robots for motor rehabilitation not only allow for robot assistance in movement initiation and guidance but also provide accurate feedback; some robots additionally provide movement resistance. Most trials of robot-assisted motor rehabilitation concern the upper extremity (UE), with robotics for the lower extremity (LE) still in its infancy... Robot-assisted UE therapy, however, can improve motor function during the inpatient period after stroke.” In 2010 AHA suggested that robot-assisted therapy for the UE has already achieved Class I, Level of Evidence A for Stroke Care in the Outpatient Setting and Care in Chronic Care Settings. It suggested that robot-

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assisted therapy for UE has achieved Class IIa, Level of Evidence A for stroke care in the inpatient setting. Class I is defined as: “Benefit >>> Risk. Procedure/Treatment SHOULD be performed/administered;” Class IIa is defined as: “Benefit >> Risk, IT IS REASONABLE to perform procedure/administer treatment;” Level A is defined as “Multiple populations evaluated: Data derived from multiple randomized clinical trials or meta-analysis” [1]. In 2016 AHA did not differentiate between inpatient and outpatient stroke populations and it suggested Class IIa across different time points following a stroke. This is not an isolated opinion. The 2010 Veterans Administration/Department of Defense guidelines for stroke care came to the same conclusion endorsing the use of rehabilitation robots for the upper extremity. More specifically, the VA/DOD 2010 guidelines for stroke care “Recommend robot-assisted movement therapy as an adjunct to conventional therapy in patients with deficits in arm function to improve motor skill at the joints trained.” The VA/DOD suggested that robot-assisted therapy for the UE has already achieved rating level B, “A recommendation that clinicians provide (the service) to eligible patients. At least fair evidence was found that the intervention improves health outcomes and concludes that benefits outweigh harm” [2]. These endorsements came on the 21st anniversary of our initial efforts begun in 1989 that led to what became known as “MIT-Manus.” It would be difficult to deny the impact of this work on neuro-rehabilitation, described by a senior clinical colleague as “perhaps one of the most important developments in neuro-recovery in the last 75 years” (the late Dr. Fletcher H. McDowell, Director, The Burke Rehabilitation Center, Cornell-NY Hospital, White Plains, NY). Creating this level of trust required decades of perseverance. The enormity of the challenge cannot be understated. This type of research is the antithesis of the rapid-fire breakthroughs expected in, say, information technologies. It requires slow and painstaking experimental trials and the creation of a large body of experimental evidence to demonstrate progress, but that is essential. Neuro-rehabilitation depends on neural

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plasticity and its potential to augment recovery. The central challenge of rehabilitation robotics is to provide tools to manage and harness positive plastic changes. It is not simply to automate conventional practices. Primarily due to a lack of tools for measurement and experimental control, many conventional practices lack the support of scientific evidence. As a result, there is no clear design target for the technology nor any reliable “gold standard” against which to gauge its effectiveness. The message seems clear: We must study the process of neuro-recovery as well as the technologies that might augment this process.

26.1.2 An Upper Extremity Gym of Robots To begin with, we had to invent the technology since the available technologies were inadequate. We developed interactive robots to work with the shoulder-and-elbow (with and without gravity compensation), the wrist and the hand, as well as combinations of these modules. We further developed exoskeletal robots for neuroscience research (see Fig. 26.1).

26.1.2.1 Modularity We chose to pursue a modular approach for several reasons. The foremost was entirely pragmatic: As we intended to introduce new technology to a clinical environment, it needed to be minimally disruptive—i.e., not too big, complex or intimidating. A secondary reason was our recognition that engineers were unlikely to create optimal technology on the first pass. Though a design to address over 200 DOFs of the human skeleton was technically feasible, it would have been large, complex and—most important—difficult to revise or modify. With a modular approach, individual modules could be refined and optimized without re-design of other modules. 26.1.2.2 Gravity-Compensated Shoulder-And-Elbow Robot The centerpiece of our effort for the upper extremity became known as MIT-Manus, from

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MIT’s Motto “Mens et Manus” (Mind and Hand). Unlike most industrial robots, MITManus was configured for safe, stable, and highly compliant operation in close physical contact with humans. This was achieved using impedance control, a key feature of the robot control system. Its computer control system modulated the way the robot reacted to mechanical perturbation from a patient or clinician and ensured a gentle compliant behavior. The machine was designed to have a low intrinsic end-point impedance (i.e., be backdrivable) to allow weak patients to express movements without constraint and to offer minimal resistance at speeds up to 2 m/s (the approximate upper limit of unimpaired human performance, hence the target of therapy, and the maximum speed observed in some pathologies, e.g., the shock-like movements of myoclonus). MITManus had 2 active degrees of freedom (DOF) and one passive DOF. It consisted of a semi-direct-drive, five-bar-linkage SCARA mechanism (Selective Compliance Assembly Robot Arm) driven by brushless motors [3, 4]. Since then, several variants were deployed commercially.

26.1.2.3 Gravity-Compensated Shoulder-Elbow-AndWrist Exoskeletal Robot Based on their mechanical interface with a human, robots can be classified as end-effector or exoskeletal designs. End-effector robots interact with the human via a handshake, i.e., the interaction takes place through a single point of contact. In other words, there is power exchange only at the tip of the robot. Exoskeletal robots are mounted on distinct human limb segments with more than one point of contact. End-effector robot designs like the MIT-Manus are simpler, and afford significantly faster “don” and “doff” (set-up time much smaller) than exoskeleton designs, but typically occupy a larger volume. We employ a “rule of thumb” to guide us in the selection of configuration based on the target range of motion. For limb segment movements requiring joint angles to change by 45 degrees or less, end-effector designs appear to offer better

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Fig. 26.1 A Gym of Upper Extremity Robots. The top row, left shows a person with chronic stroke working with the anti-gravity shoulder-and-elbow robot. The top row, middle panel shows a person working with the planar shoulder-andelbow robot. The top row, right panel shows the wrist robot during therapy at the Burke Rehabilitation Hospital. The middle row, left panel shows the hand module for grasp and release. The middle row, middle panel shows reconfigurable robots. The robotic therapy shoulder-and-elbow and wrist modules can operate in stand-alone mode or be integrated into a coordinated functional unit. The middle row, right panel shows the shoulder-and-elbow and hand module integrated into a coordinated functional unit. The bottom row shows an exoskeletal robot with 3 active DOFs designed for psychophysical studies of the shoulder, elbow and wrist. For this exoskeletal robot, the links must be adjusted to the person’s limb segments (using laser pointers). Once the arm, forearm, and wrist are properly adjusted, psychophysical experiments can assist or selectively apply perturbation force fields to the shoulder, elbow, and wrist (either flexion/extension or abduction/adduction)

compromises. Conversely, exoskeletal designs appear to offer better choices for larger ranges of motion. That said, in some circumstances, the application dictates the configuration. One such

case occurs during psychophysical experiments in which we may want to carefully apply and control perturbations to one, but not another, joint. Hence we designed a highly-backdrivable,

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3 active DOF, gravity-compensated shoulderelbow-and-wrist exoskeletal robot as shown in Fig. 26.1. Two Exos can be configured for bimanual use (see other examples in [5–7]).

26.1.2.4 Gravity Non-compensated Shoulder-And-Elbow Robot A 1-DOF module was conceived to extend the benefits of planar robotic therapy to spatial arm movements, including movements against gravity. Incorporated in the design are therapists’ suggestions that functional reaching movements often occur in a range of motion close to shoulder scaption. That is, this robotic module was designed for therapy to focus on movements within the 45° to 65° range of shoulder abduction and from 30° to 90° of shoulder elevation or flexion [8]. The module can permit free motion of the patient’s arm, or can provide partial or full assistance or resistance as the patient moves against gravity. As with MIT-Manus, the system is highly-backdrivable. 26.1.2.5 Wrist Robot To extend treatment beyond the shoulder-andelbow, we designed and built a wrist module for robotic therapy [9]. The device accommodates the range of motion of a normal wrist in everyday tasks, i.e., flexion/extension 60°/60°, abduction/ adduction 30°/45°, and pronation/supination 70°/70°. The torque output from the device is capable of lifting the patient’s hand against gravity, accelerating the inertia, and overcoming most forms of hypertonicity. As with all of our exoskeletal designs, we purposely under-actuated the wrist robot with fewer degrees of freedom than are anatomically present. Not only does this simplify the mechanical design, but it also allows the device to be installed quickly without problems of misalignment with the patient’s joint axes. In this case, the axes of the wrist’s ulnar-radial and flexion-extension joints do not intersect and the degree of non-intersection varies between individuals [10]. If robots and humans had the same number of degrees of freedom but these were not

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co-aligned, the motion might evoke excessive forces or torques. By allowing the human joint more degrees of freedom than the robot, excessive loads are avoided. Ease of use is another critical consideration in all our designs. We consider it a major determinant of success or failure in the clinical rehabilitation environment. The wrist robot must be attached to or removed from the patient (donned or doffed) within 2 min. Finally, the wrist robot module can be operated in isolation or mounted at the tip of the shoulder-and-elbow, gravity-compensated robot. Hence it enables a combination of translating the hand (with the shoulder-and-elbow robot) to a location in space and orienting the hand (with the wrist robot) to facilitate object manipulation.

26.1.2.6 Hand Robot Moving a patient’s hand (for purposes of poststroke training focusing on the fingers) is not a simple task since the human hand has 15 joints with a total of 22 DOF (27 including the wrist); therefore, it was prudent to determine how many DOF are necessary for a patient to perform the majority of everyday functional tasks. Here our clinical experience with over 2,000 stroke patients was invaluable in that it allowed us to identify what was most likely to work in the clinic (and what probably would not). Though individual digit opposition (e.g., thumb to pinkie) may be important for the unimpaired human hand, it is clearly beyond the realistic expectations of most of our patients whose impairment level falls between severe and moderate; a device to manipulate 22 DOF is unnecessary (or at least premature). Our hand therapy module is a novel design that converts rotary into linear movement using a single brushless DC electrical motor as a free-base mechanism with what is traditionally called the stator being allowed to rotate freely [11]. The stator (strictly, the “second rotor”) is connected to a set of arms, while the rotor is connected to another set of arms. When commanded to rotate, the rotor and stator work like a double crank and slider mechanism, in opposing configuration, where the crank is represented by

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a single arm and the slider is the shell or panel that interacts with the hand of the patient (see Fig. 26.1). The hand robot is used to simulate grasp and release with its impedance determined by the torque evoked by relative movement between stator and rotor. A torsional spring (connected in geometric parallel) is available to compensate for a patient’s hypertonicity (inability to relax). The hand robot is capable of providing continuous passive motion, strength, sensory, and sensorimotor training for grasp and release; it can be employed in stand-alone operation or mounted at the tip of the planar robot (see Fig. 26.1).

26.2

Harnessing Plasticity to Augment Recovery

26.2.1 Clinical Evidence for Inpatient Care The earliest pilot experiments using the MITManus device to alter the outcome of post-stroke upper extremity impairment demonstrated remarkable safety and significant effectiveness whether the patients were treated within a month, at the inpatient rehabilitation setting, or 6 months after the acute stroke, in an outpatient setting [12, 13]. Volpe et al. reported composite results of the controlled tests of robotic therapy compared to appropriate sham conditions in 96 stroke inpatients admitted to Burke Rehabilitation Hospital in White Plains, NY [14]. All participants received conventional neurological rehabilitation during their participation in the study. The goal of these trials was to build on the positive pilot data and test whether movement therapy had a measurable impact on recovery. Consequently, we provided one group of patients with as much movement therapy as possible to address a fundamental question: Does goal-oriented movement therapy have a positive effect on neuromotor recovery after stroke? In retrospect, at the time of these studies, the answer to this question was far from clear.

Placement of subjects in an experimental (robot-trained) or control (robot-exposure) group was done in a random fashion. Individuals in the experimental group received no less than 25 sessions of sensorimotor robotic therapy for the paretic arm (one-hour session daily every weekday). Patients were asked to perform goaldirected, visually-guided and visually-evoked reaching movements with their paretic arm. MIT-Manus’ low impedance and low-friction assured that the robot would not suppress the patient’s attempts to move. The robot afforded gentle guidance and assistance only when a patient could not move or deviated from the desired path [15]. We named this intervention “sensorimotor” therapy and it was similar to the “hand-over-hand” assistance that a therapist often provides during usual care. It is interesting to note that this form of “assistance as needed,” which has been a central feature of our approach from the outset, has been adopted by other groups [12, 16, 17]. Individuals assigned to the robot-exposure (control) group were asked to perform the same planar reaching tasks as the robot therapy group. However, the robot did not actively assist the patient’s movement attempts. When the subject was unable to reach toward a target, he or she could assist with the unimpaired arm or the technician in attendance could help to complete the movement. The robot supported the weight of the limb while offering negligible impedance to motion. For this control group, the task, the visual display, the audio environment (e.g., noise from the motor amplifiers) and the therapy context (e.g., the novelty of a technology-based treatment) were all the same as for the experimental group, so this served as a form of “placebo” of robotic movement therapy. Patients in this group were seen for only 1 h per week during their inpatient hospitalization. The study was “double blinded” in that patients were not informed of their group assignment and therapists who evaluated their motor status did not know to which group patients belonged. Standard clinical evaluations

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included the upper extremity sub-test of the FuglMeyer Assessment (FM, maximum score = 66); the MRC Motor Power score for four shoulderand-elbow movements (MP, maximum score = 20); and the Motor Status Score (MSS, maximum score = 82) [18–20]. The Fugl-Meyer test is a widely accepted measure of impairment in sensorimotor and functional grasp abilities. To complement the Fugl-Meyer scale, Burke Rehabilitation Hospital developed the Motor Status Scale to further quantify discrete and functional movements in the upper limb. The MSS scale expands the FM and has met standards for interrater reliability, significant intra-class correlation coefficients and internal item consistency for inpatients [21]. Although the robot-exposure (control) and robot-treated (experimental) groups were comparable on admission, based on sensory and motor evaluation and on clinical and demographic scales, and both groups were in-patients in the same stroke recovery unit and received the same standard care and therapy for comparable lengths of stay, the robot-trained group demonstrated significantly greater motor improvement (higher mean interval change ± sem) than the control group on the MSS/E and MP scores (see Table 26.1). In fact, the robot-trained group improved twice as much as the control group in these measures. Though this was a modest beginning, it provided unequivocal evidence that movement therapy of the kind that might be delivered by a robot had a significant positive impact on recovery.

Table 26.1 Burke rehabilitation hospital inpatient studies

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26.2.2 Clinical Evidence for Chronic Care The natural history of motor recovery of the paretic upper limb after stroke reveals a dynamic process that has traditionally been described by a period of flaccidity that is followed by changes in tone and reflex, as well as the frequent development of synkinesis or associated movement disorders. This synkinesis is characterized by involuntary, composite movement patterns that accompany an intended motor act [22]. Complete motor recovery, when it occurs, will unfold rapidly in hours or days. The more commonly observed partial recovery, with broad variability in final motor outcomes, unfolds over longer periods [23, 24]. At the time of our initial studies, the state of knowledge regarding motor recovery post-stroke indicated that the majority of gains in motor abilities occurred within the first three months after stroke onset, and that over 90% of motor recovery was complete within the first five months [25]. However, we were able to recall one third of the 96 stroke inpatients mentioned earlier three years after discharge. We observed that both groups continued to improve after discharge from the hospital and after five months post-stroke. Our data suggested that previous results limiting the potential of chronic patients’ recovery were based on the effects of general rather than task-specific treatments during the recovery period post-stroke [23, 24]. In 2010 the Veterans Affairs completed the VA-ROBOTICS study (CSP-558), a landmark multi-site, randomized clinical trial in a

Between group comparisons: Final evaluation minus initial evaluation

Robot trained (N = 55)

Control (N = 41)

PValue

Impairment measures (±sem) Fugl-Meyer shoulder/elbow (FM-se)

6.7 ± 1.0

4.5 ± 0.7

NS

Motor power (MP)

4.1 ± 0.4

2.2 ± 0.3

6 months) stroke was recruited in this study.

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Each subject participated in two test conditions, i.e., with constraint force and treadmill only, with a 10-min sitting break in between. The order of the two test conditions was randomized across subjects. Each subject walked on a treadmill without force for 1-min (i.e., baseline), followed by the application of the constraint force for an additional 7 min (adaptation). Then, the constraint force was removed without warning and the subjects continued walking on the treadmill for another 1-min (post-adaptation). Afterwards, subjects were allowed to take a 1-min standing break and walked on the treadmill for another 5 min with the application of constraint force (readaptation). Overground walking gait speed was tested before treadmill walking, immediately post treadmill walking, and 10-min after the end of treadmill walking. For the treadmill only condition, a protocol that was similar to the constraint force condition was used but no force was applied. Averages of EMG activity (during stance phase) of the paretic leg for the constraint force and treadmill only conditions are shown in see Fig. 32.2 [105]. For the constraint force condition, the EMG integral of TA, MG, SOL, VM, RF, MH, ADD, and ABD increased during the early adaptation period, and still remained higher during the late adaptation period. Further, the enhanced muscle activity of TA, MG, RF and MH was partially retained following the release of the constraint force during the post-adaptation period. In contrast, for the treadmill only condition, the EMG integral of all muscles showed modest change during the course of the adaptation period and post-adaptation period, Fig. 32.2. For the constraint condition, the symmetry of step length during overground walking significantly improved after treadmill walking (F2,13, = 3.56, p = 0.043, n = 14, data from two subjects were excluded due to outlier, ANOVA). The post-hoc test indicated that the symmetry index of step length significantly decreased from 0.13 ± 0.06 at baseline to 0.09 ± 0.07 at 10-min after the end of treadmill walking, p = 0.03, (a smaller number means a more symmetrical step length). Overground walking speed had no

M. Wu

significant changes (p > 0.05). For the treadmill only condition, there were no significant effects on step length symmetry, and overground walking speed (all p > 0.05). Between conditions, the comparison showed more improvement in the symmetry of step length after treadmill walking with constraint force than after treadmill only (p < 0.04) suggesting that subjects walked with a more symmetrical gait pattern after treadmill walking with the application of the leg constraint force than after treadmill only condition.

32.2.2 Locomotor Training in Human with Incomplete SCI 32.2.2.1 Introduction Recent reviews of clinical studies on the effectiveness of current robotic training in humans with SCI suggest that robot-assisted gait training is not superior to other gait training modalities [74, 106]. One possibility is that these robotic training modalities do not provide adequate challenges to drive motor learning in humans with SCI during locomotor training [107]. For instance, muscle activities are significantly lower during passive guided, robotic locomotor training than with physical therapist-assisted treadmill training in humans with SCI [108]. Previous studies have shown that an error-augmentation training paradigm may enhance arm recovery in individuals post-stroke [85]. Thus, we postulated that error-augmentation also would facilitate motor learning during locomotor training in humans with SCI. By applying a resistance load to the leg during treadmill walking, which may increase leg kinematic errors [109], recent studies have indicated that humans with SCI adapt to the resistance load applied and demonstrate an aftereffect consisting of an increase in step length following load release [110]. The presence of aftereffects suggests the formation of anticipatory locomotor commands in response to the resistance load. In particular, a previous study indicated that this aftereffect could be transferred from treadmill training to overground walking in humans with

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A Flexible Cable-Driven Robotic System: Design and Its Clinical …

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Fig. 32.2 Step-by-step integrated EMG of the paretic leg during the course of treadmill walking for the conditions with a constraint force applied to the non-paretic leg and treadmill only (control). Data shown are average across 15 subjects post-stroke (data from one subject were excluded due to large artifacts). The step numbers were different across subjects during the adaptation period because the treadmill speeds were different. Thus, the data from the first 160 steps and the last 30 steps during the adaptation period were used to calculate the average of integrated EMG. (Adopted from Wu et al. [105]). Abbreviation: tibialis anterior (TA), medial gastrocnemius (MG), soleus (SOL), vastus medialis (VM), rectus femoris (RF), medial hamstrings (MH), adductors (ADD) and abductors (ABD)

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SCI [111]. However, locomotor adaptation and the aftereffects are generally short-lived, i.e., the increase in step length returns back to baseline within tens of steps during the post-adaptation period, after one session of force perturbation training, which may have a limited clinical impact on walking function in humans with SCI. A previous study using a split-belt treadmill paradigm indicated that prolonged repeated exposure to split-belt perturbation induces a long-term retention of improved step length symmetry in individuals post-stroke [112]. Thus, we postulated that a prolonged repeated exposure to swing resistance perturbations during treadmill training might also induce long-term retention of improved step length, resulting in improvements in the walking function of humans with SCI. The purpose of this study was to determine whether robotic resistance or assistance treadmill training by using a cable-driven robotic system would be effective in improving locomotor function in humans with chronic incomplete SCI [113]. Ten individuals with chronic incomplete SCI (i.e., >12 months post injury) with an injury level from C2 to T10 were recruited to participate in this study. All subjects were classified by the American Spinal Injury Association (ASIA) as ASIA grade D. In order to test the locomotor training effect of the cable-driven robot in the SCI population, an 8 weeks training trial was conducted using a randomized crossover schedule. Specifically, subjects were blocked by gait speed into slow (0.5 m/s) subgroups and then randomized to either the assistance or resistance training first. After the first 4 weeks of training, subjects from both groups were switched from assistance to resistance or from resistance to assistance training, and completed another 4 weeks of training. Three assessments of gait were used to determine the training effects.

32.2.2.2 Results In this pilot study, 8 out of 10 subjects finished 8 weeks of robotic treadmill training, with 2 subjects dropping out of the study. For the 8

M. Wu

patients that finished 8 weeks of robotic gait training, we found a significant improvement in self-selected overground walking speed (p = 0.03, one-way repeated measures ANOVA), i.e., the gait speed improved from 0.67 ± 0.20 m/s to 0.76 ± 0.23 m/s (see Fig. 32.3a). Fast walking speed also improved from 0.96 ± 0.31 m/s to 1.06 ± 0.32 m/s, although no significant difference was obtained due to the small sample size (p = 0.19, one-way repeated measures ANOVA). In addition, scores on the Berg Balance Scale significantly improved from 42 ± 12 at pre-training to 45 ± 12 post 8 weeks of robotic gait training (see Fig. 32.4). There were no significant changes in walking distance at the pre and post robotic training evaluation sessions (p = 0.12), although averaged 6-min walk distance increased from 223 ± 81 m at pre-training to 247 ± 88 m at post training (see Fig. 32.3b). In order to compare the effect of 3DCaLT robotic treadmill training versus conventional treadmill training on walking function in people with SCI, we conducted a randomized controlled study. In particular, one limitation of previous robotic gait training systems is the constraint of the pelvis movement in the mediolateral direction [3], which may limit the weight shifting that occurs during the training. The ability to initiate and control weight shifting is a prerequisite for independent walking [63]. The weight shifting ability of many patients with SCI is impaired. Insufficient weight shifting to the ipsilateral leg may limit the unloading of the contralateral leg, which may affect the leg swing of the contralateral leg because the load afferent is a key input that modulates the translation from stance to swing during locomotion [114]. Thus, improving the weight shifting ability of the patients with SCI may improve their walking function. The goal of this study was to determine whether 3DCaLT robotic treadmill training would improve walking function in people with SCI. We hypothesized that the integration of the weight shifting component into the treadmill training paradigm would improve the weight

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A Flexible Cable-Driven Robotic System: Design and Its Clinical …

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Fig. 32.3 Self-selected overground gait speed (a) and 6 min walk distance (b) at pre and post 8 weeks of robotic treadmill training in human SCI. The bar and error bar indicates the mean and standard deviation of the gait speed and walking distance across 8 subjects for pre and post training. Asterisk (*) indicates a significant effect of treatment

were recruited in this study and were randomly assigned to receive robotic or treadmill only training, and underwent 6 weeks of locomotor training. The 3DCaLT cable-driven robotic system was used to provide controlled forces to the pelvis and legs during treadmill walking. Outcome measures consisted of overground walking speed, 6-min walking distance, and other clinical measures, and were assessed pre and post 6 weeks of training, and 8 weeks after the end of training (Fig. 32.4).

Fig. 32.4 Berg balance scale at pre and post 8 weeks robotic treadmill training in humans with SCI. The bar and error bar indicates the mean and standard deviation across 8 subjects for pre and post training. Asterisk (*) indicates a significant effect of treatment of robotic gait training

shifting ability and/or balance of humans with SCI, resulting in improvements in walking speed and endurance after robotic treadmill training. Sixteen humans with motor incomplete SCI (i.e., ASIA Impairment Scale Level of C or D)

32.2.2.3 Results Fourteen participants completed all the training and evaluation sessions and 2 participants dropped out of the study. Robotic treadmill training with weight shift induced significant improvement in the endurance of patients with SCI. Specifically, 6-min walking distance significantly increased after robotic treadmill training (p = 0.02) [115]. Post-hoc tests indicated that 6min walking distance significantly increased from 120 ± 37 m to 157 ± 59 m (29% increase) after robotic training (p = 0.04), and remained to be significantly greater than baseline at the follow up test, i.e., 151 ± 60 m (24% increase, p = 0.04). In addition, self-selected overground walking speed was intended to

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M. Wu

increase after robotic treadmill training, i.e., from 0.33 ± 0.15 m/s to 0.39 ± 0.20 m/s (15% increase), although this was not significant (p = 0.07), and was 0.38 ± 0.19 m/s (13% increase) at the follow up test. There were no significant changes in fast walking speed (p = 0.16) after robotic treadmill training. In contrast, treadmill only training induced no significant improvement in the endurance of patients with SCI. Specifically, the 6-min walking distance had no significant change after treadmill only training (p = 0.65). In addition, self-selected walking speed (p = 0.89) and fast walking speed (p = 0.43) showed no significant change pre and post treadmill only training. Between group comparisons indicated that gains in 6-min walking distance were significantly greater for the robotic training group than that for treadmill only training group (p = 0.03), but gains in self-selected and fast walking speeds were not significantly different between the two groups (p = 0.06 and p = 0.12 for self-selected walking speed and fast walking speed, respectively), Fig. 32.5.

32.2.3 Locomotor Training in Children With CP 32.2.3.1 Introduction As mentioned previously, weight shifting in the mediolateral direction is one of the key components during human locomotion [116]. However, the weight shifting ability is often impaired in children with CP compared to children who are typically developed [117]. For instance, children with CP performed weight shifting less efficiently as demonstrated by a shorter range of motion of the center of pressure (COP) and slower velocity of COP displacement during standing compared to children who are typically developed. It was suggested that weight shift training might improve dynamic balance during walking in children with CP [118]. Thus, we postulated that applying a mediolateral assistance force at the pelvis during the stance phase of gait might facilitate weight shifting in children with CP during treadmill walking, and repeat practice

Fig. 32.5 Changes in 6 min walking, a, self-selected overground walking speed, b, fast walking speed, c, pre and post 6 weeks of robotic treadmill training or treadmill only training, and 8 weeks after the end of training. Data were averaged across subjects in each group. * indicates a significant difference, p < 0.05. (Adopted from Wu et al. [115])

of weight shifting during treadmill training may improve dynamic balance and improve walking speed in children with CP. Evidence from spinalized mice indicates that motor learning is more effective with assistance as needed than with a fixed trajectory paradigm [81]. Thus, a robotic system that allows for variability in the stepping pattern during treadmill training will be effective in improving walking speed in children with CP. In addition, results from motor learning studies indicate that when there are more similarities between learning tasks and the application of those tasks, a greater transfer of motor skills will take place

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A Flexible Cable-Driven Robotic System: Design and Its Clinical …

[118]. Thus, a robotic BWSTT technique that provides less constraints and allows for a natural dynamic gait pattern during treadmill walking will be an effective method for transferring motor skills from treadmill training to overground walking in children with CP as measured by increased overground walking speed after robotic BWSTT. The purpose of this study is to assess improvements in the locomotor function of children with CP after robotic treadmill training with the application of applying controlled forces to both the pelvis, for facilitating weight shift, and the leg at the ankle, for facilitating leg swing, through the 3DCaLT robotic gait training system [83]. We hypothesized that robotic treadmill training that applies assistance at the pelvis for facilitating weight shifting would be more effective than treadmill only training in improving walking function in children with CP. Twenty three children with CP were recruited (14 boys and 9 girls, the average age were 10.9 ± 3.2 years old, Gross Motor Function Classification System (GMFCS) levels [119] were I–IV). Each subject was randomly assigned to either a robotic training group (n = 11) or a treadmill only training group (n = 12). Treadmill training was performed 3 times per week for 6 weeks. Gait assessment was made pre, post 6 weeks of robotic treadmill training, and at 8 weeks after the end of the training, using gait speed, endurance, and clinical measures of motor function (the dimensions D (standing) and E (walking, running, jumping) of the Gross Motor Function Measure (GMFM-66), [120]). Results. Eleven participants from the robotic training group and 10 participants from the treadmill only group completed all the training and evaluation sessions (the dropout rate was 8.7%). The walking function of children with CP improved after 3DCaLT robotic treadmill training. Specifically, self-selected walking speed significantly increased after robotic training (p = 0.03), see Fig. 32.6 [121]. The post-hoc test indicated a significant difference between pre versus post training tests (15.4% increase, p = 0.04), although there were no significant differences

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between pre versus follow up tests (p = 0.08). Six-minute walking distance significantly increased after robotic training (p = 0.048), Fig. 32.6. Post-hoc test indicated a significant difference between pre versus post training tests (12.8% increase, p = 0.04), although there were no significant differences between pre versus follow up tests (p = 0.28). The GMFM scores had no significant changes after robotic training (p > 0.06). Walking speed and endurance had no significant change after treadmill only training (p > 0.05), see Fig. 32.6. In addition, treadmill only training induced no significant change in dimension E of GMFM (p = 0.34), but induced a significant increase in the dimension D of GMFM (p = 0.01). Post-hoc tests indicated a significant difference between pre versus post training tests (p = 0.03), and pre versus follow up test (p = 0.02). A greater gain in 6-min walking distance was obtained for the participants from the robotic group than that from the treadmill only group (p = 0.01), but the gain in self-selected walking speed had no significant difference between the two groups (p = 0.12). Specifically, changes in 6-min walking distance were 42.2 ± 57.4 m (post training) and 25.1 ± 52.0 m (follow up) after robotic training, and were −3.8 ± 35.9 m (post training) and −8.2 ± 46.8 m (follow up) after treadmill only training. Changes in selfselected walking speed were 0.10 ± 0.15 m/s (post training) and 0.09 ± 0.09 m/s (follow up) after robotic training, and were 0.04 ± 0.11 m/s (post training) and 0.04 ± 0.11 m/s (follow up) after treadmill only training. In order to determine whether robotic resistance versus assistance training was more effective in improving walking function in children with CP, we conducted another randomized controlled study. Specifically, results from motor learning studies indicated that active training is more effective than passive training in improving the efficacy of motor training [40]. Thus, active engagement from children with CP might improve the efficacy of locomotor training. We proposed that applying a resistance force to leg

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Fig. 32.6 Average of self-selected, fast walking velocities pre, post 6 weeks of robotic treadmill training, a, or treadmill only training, b, and 8 weeks after the end of treadmill training, i.e., follow up. Three trials were tested and averaged across each test session and averaged across participants for each group. c. average of 6-min walking distance pre, post 6 weeks of robotic treadmill training or treadmill only training, and 8 weeks after the end of training. Error bars indicate the standard deviation of each gait parameter (n = 10 for the robotic treadmill training group, data from one subject was excluded for the 6-min walking distance test because this subject was sick immediately before post test, which significantly impacted his endurance performance based on subject’s self-report). Error bars indicate the standard deviation of each gait parameters. SSV, self-selected velocity; FV, fast velocity. * indicates a significant difference, p < 0.05. (Adopted from Wu et al. [121])

swing may force them to be more actively involved. In addition, applying a resistance force to leg swing may produce a deviation in step kinematics, i.e., an increase in the kinematic errors, which is supported by previous studies in individuals post-stroke [122, 123] and spinal cord injury [109]. Error augmentation may

M. Wu

accelerate motor learning during treadmill training [87, 124], resulting in an improvement in the efficacy of locomotor training in children with CP. On the other hand, providing leg assistance force may facilitate leg swing, which imitates the way that physical therapist provides leg assistance during treadmill training, and improve locomotor function in children with CP through use-dependent motor learning mechanisms [86]. Thus, the purpose of this study was to assess functional changes after robotic resistance versus assistance treadmill training in children with CP. We hypothesized that children with CP from both groups would show improvements in locomotor function, although greater improvements were expected after resistance than that after assistance training. In this study, 23 children with CP (11 males and 12 females) were recruited. The average age of participants was 10.6 years old (ranged from 6 to 14). Their Gross Motor Function Classification System level ranged from I–IV (I (1), II (11), III (8), IV (3)). Participants were randomly assigned to receive assistance or resistance force applied to both legs at the ankle during treadmill walking. Training sessions took place 3 times a week for 6 weeks.

32.2.3.2 Results Twenty participants completed all the training and assessment sessions with 3 participants dropping out. Robotic resistance treadmill training improved walking function in children with CP. Specifically, fast walking speed significantly increased after robotic resistance treadmill training (F (2, 9) = 4.12, p = 0.03), [125]. Post-hoc tests indicated a significant increase from baseline to post testing (18% increase, from 0.98 ± 0.39 m/s to 1.13 ± 0.38 m/s, p = 0.01), although there was no significant difference between baseline and follow up testing (17% increase, p = 0.13). Self-selected walking speed tended to increase (32% increase), but this was not significant (F (2, 9) = 1.66, from 0.63 ± 0.30 m/s to 0.72 ± 0.24 m/s, p = 0.21).

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A Flexible Cable-Driven Robotic System: Design and Its Clinical …

In addition, 6 min walking distance significantly increased after resistance training (F (2, 9) = 10.04, p = 0.001). Post-hoc tests indicated a significant increase from baseline to post testing (30% increase, from 272.7 ± 113.0 m to 336.3 ± 104.9 m, p = 0.01), and from baseline to follow up testing (35% increase, from 272.7 ± 113.0 m to 353.9 ± 125.8 m, p = 0.001). In contrast, robotic assistance treadmill training induced modest changes in walking function in children with CP. Specifically, both fast and self-selected walking speeds had no significant changes after assistance training (F (2, 9) = 1.85, from 0.84 ± 0.34 m/s to 0.84 ± 0.37 m/s, p = 0.19, and F (2, 9) = 0.51, from 0.54 ± 0.22 m/s to 0.52 ± 0.18 m/s, p = 0.6, for fast and self-selected walking speeds, respectively). In addition, 6 min walking distance had no significant change after assistance training (F (2, 9) = 0.48, from 216.3 ± 116.8 m to 230.1 ± 119.2 m, p = 0.63). Greater functional gains were observed for participants who underwent robotic resistance training than for those who underwent robotic assistance training. Specifically, changes in selfselected walking speed, fast walking speed, and 6 min walking distance after treadmill training were significantly greater for participants from the resistance group than that from the assistance group (F (1,1) = 5.36, p = 0.03, F (1,1) = 10, p = 0.003, and F (1,1) = 12.23, p = 0.001, for self-selected walking speed, fast walking speed, and 6 min walking distance, respectively), Fig. 32.7. Functional gains were partially retained or even slightly increased during the follow up period. GMFM scores significantly changed after robotic resistance training (F (2, 9) = 5.21, p = 0.02). Post-hoc test indicated significant differences between post testing and follow up testing scores (p = 0.02), although there was no significant difference between pre testing versus post testing scores (p = 0.89), and pre testing versus follow up testing scores (p = 0.05). In contrast, GMFM scores had no significant changes after robotic assistance training (F (2, 9) = 0.49, p = 0.69).

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32.2.3.3 Discussion The purpose of these studies was to determine the effect of robotic treadmill training using the 3DCaLT on walking function in adults with chronic stroke and motor incomplete SCI, and children with CP. We found that locomotor training using the 3D cable-driven robotic system may induce improvements in locomotor function in adults with stroke and SCI, and children with CP. In particular, results show that the functional gains after robotic treadmill training were greater than that after treadmill only training in adults

Fig. 32.7 Changes in walking self-selected walking speeds, (a), fast walking speed, (b), and 6-min walking distance, c, after robotic resistance/assistance treadmill training, and 8 weeks after the end of the training, i.e., follow up test. Data shown in the figure are the mean and standard deviation of walking speeds and distance across participants. * indicates a significant difference, p < 0.05. (Adopted from Wu et al. [126])

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with SCI and children with CP. Further, the improvements in walking function were partially retained at 8 weeks after the end of the training, indicating the clinical significance of such robotic treadmill training.

32.2.4 Improved Walking Function in Individuals PostStroke Applying a controlled resistance or assistance load to the paretic leg during treadmill training using the cable-driven robotic system significantly improved walking function in individuals post-stroke. However, robotic resistance training, which was applied to the paretic leg, was not superior to assistance training in improving endurance, balance in individuals post-stroke [87]. The increase in kinematic errors produced by the resistance load may elicit an error correction process that accelerates motor learning during locomotor training in individuals post-stroke. For instance, for the subjects who were assigned to the resistance training group, the resistance applied to the paretic leg produced a deviation in leg kinematics, that is, increased kinematic errors. An enhanced error has been shown to be more effective than passive guidance in improving arm performance in individuals post-stroke [85]. For the lower limb, previous studies indicated that individuals post-stroke adapted to the resistance load applied to the paretic leg and showed an aftereffect consisting of increased step length of the paretic leg after load release [122, 123]. Further, repeated exposure to resistance load during treadmill training may induce a prolonged retention of aftereffects of the paretic leg in individuals post-stroke. In this study, repeated exposure to a resistance load was applied to the paretic leg during 6 weeks of treadmill training. As a result, the step length of the paretic leg during overground walking increased after resistance training, suggesting that the aftereffect of an increased step length may be accumulated

M. Wu

and transferred from one context (i.e., treadmill walking) to another context (i.e., overground walking) in individuals post-stroke. In particular, we observed a partial retention of the increased step length of the paretic leg at follow up. On the other hand, for subjects who were assigned to the assistance training group, an assistance force provided to the paretic leg may facilitate the leg swing to induce a longer step length on the paretic side during treadmill training. The increased step length of the paretic leg may be accumulated and transferred to overground walking through 6 weeks of locomotor training, resulting in an improvement in walking function after assistance treadmill training in individuals post-stroke. However, because the assistance force was applied at the paretic leg facilitating the leg to swing forward, instead of resisting the leg to induce kinematic deviation, we postulated that the motor learning mechanisms involved in robotic assistance training would be different from those involved in resistance training. A use-dependent motor learning mechanism may be involved during robotic assistance treadmill training [86]. The synaptic efficacy of sensorimotor pathways involved in the leg swing of the paretic leg may be enhanced by repetitive stepping assisted by the cabledriven robot [127]. Maintaining variation in kinematics during BWSTT is considered to be critical in improving the locomotor function in individuals post-stroke. For instance, results from animal experiments show that motor learning is more effective with a robotic algorithm that allows variability in the stepping pattern than with a fixed trajectory paradigm [81]. In addition, results from the human study have shown that intralimb coordination after stroke was improved by physical therapist-assisted BWSTT, which allowed for kinematic variability, but not robotic gait training with fixed trajectory, which reduces kinematic variability [128]. In the current study, the 3DCaLT robotic system, which is highly back drivable, has limited constraints on the leg kinematics during treadmill training [82]. The cable- driven system

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A Flexible Cable-Driven Robotic System: Design and Its Clinical …

can be moved by the patient with the smallest possible resistance as opposed by the robot. Thus, the cable system allows the patients greater flexibility in controlling their gait patterns. The cable-driven robotic did not significantly affect the variability in ankle trajectory while the controlled load was applied to the legs during treadmill walking, see Fig. 32.8a. For instance, a previous study indicated that there were no significant changes in the variability of ankle trajectory of humans with SCI for different loading conditions (i.e., at baseline, with cable attached, and with assistance load applied (ANOVA, p = 0.6), see Fig. 32.8b [82]. This type of training seems more effective than fixed trajectory training in improving locomotor function in individuals post-stroke. Results from this study also indicated that resistance training was not superior to assistance training in improving speed, endurance, balance in individuals post-stroke. One possible reason may be due to the compensatory movement of the non-paretic leg that often occurs during treadmill training when the resistance force was applied to the paretic leg during the swing phase [88]. For instance, in this study, we found enhanced muscle activities in hip extensors, and ankle plantarflexors when the constraint (i.e., resistance) force was applied to the non-paretic leg during the swing phase (which is corresponding to the time of period of the stance phase of the paretic leg). The enhanced muscle activities in hip extensors and ankle plantarflexors of the paretic leg may be because individuals post-stroke need to counteract the backward resistance force and move the body forward over the standing leg (i.e., the paretic leg) by extending the hip and ankle joints. These results suggest that the application of targeted constraint force to the non-paretic leg during the swing phase may induce forced use of the paretic leg. This is also consistent with previous CIMT, a strategy that has been extensively used in arms training in individuals post-stroke to improve motor function and daily use of the paretic arm [129]. Further, repetitive forced use of the paretic leg may result in an enhanced muscle activity of the

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Fig. 32.8 a Ankle trajectories in the sagittal plane are shown from one patient with incomplete SCI during treadmill walking without the attachment of a cable robot. The solid thick line shown is the ensemble average trajectory across 7 step cycles. b Variability of ankle trajectory for 3 different loading conditions, i.e., baseline without attachment of the cable-driven robot; cable robot attached with 4 N pretension load, and cable robot attached with controlled assistance load. Path deviation of ankle trajectory in the sagittal plane for each condition was used to quantify the variability of ankle movement of each subject during treadmill walking. The bar and error bar indicate the mean and SD of the RMS error of ankle trajectory across subjects. (Modified from Fig. 3 of Wu et al. [82])

paretic leg even when the constraint force was removed during the post-adaptation period, which may be achieved through a use-dependent motor learning mechanism [104]. The repetitive movement or neural activity in a particular pattern over time, such as enhanced muscle activity of the paretic leg, in this case, may influence future movement or neural activity patterns even when the external perturbation is removed [130].

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32.2.5 Improved Walking Function in Humans with SCI The locomotor functional gains obtained using the 3DCaLT robotic gait training system are comparable to or even greater than that of using treadmill only training or robotic systems with a fixed trajectory control strategy. For instance, in a randomized trial involving 27 participants with SCI, the use of robotic-assisted BWSTT with a fixed trajectory did not significantly increase walking velocity (mean difference was −0.05 m/s) [54, 106, 131]. However, in another study with 20 subjects with chronic SCI, results indicated that the use of robotic-assisted treadmill training with a fixed trajectory may significantly improve walking speed in the SCI population [2]. The functional gains were 0.11 ± 0.11 m/s following robotic gait training, which is comparable to the results obtained in the current study. In addition, results from the current study indicate an improvement in balance control in human SCI following cable-driven robotic gait training, i.e., Berg Balance Scale scores increased 3.3 ± 2.3. This is a functional gain not previously seen in studies with the Lokomat. The unnecessary medial-lateral support may reduce the potential functional gains in balance control following robotic gait training using the Lokomat. Recent studies indicate that there is a strong relationship between balance and walking capacity in patients with SCI [132, 133]. Thus, training stereotypical gait patterns in human SCI without challenging balance control may squander training time by focusing the training on the impairment that is not the bottleneck for achieving a greater walking speed [134]. In addition, the intensive task-specific walking practice may be delivered through a cabledriven robotic-assisted BWSTT system with the help of only 1 therapist, and can be performed for a longer duration (dependent upon the tolerance of the patient) thereby increasing the amount of practice of stepping behaviors. While the sample size is small, our results indicate that the improvements in locomotor function in our ambulatory subject population were statistically

M. Wu

significant, with self-selected gait speed and Berg Balance Scale scores increasing by 0.09 ± 0.10 m/s (13%) and 3.3 ± 2.3 (8%), respectively, post robotic training. These improvements were qualitatively similar to those achieved by people with a similar diagnosis and chronicity of injury who performed therapistassisted BWSTT [54]. Thus, 3DCaLT robotic BWSTT may achieve comparable functional gains when compared to therapist-assisted BWSTT, but can substantially reduce the labor effort and personnel cost of physical therapists. Applying pelvis assistance could facilitate weight shift and induce additional challenges in maintaining lateral balance control and stability of the hip of the standing leg of patients with SCI. Thus, we speculate that patients with SCI may have to recruit additional muscle activation of the hip abductors/adductors to maintain lateral balance during walking. Repeated exposure to a pelvis assistance force during long-term treadmill training may improve motor control of hip abductors/adductors through use-dependent motor learning mechanisms [36], resulting in an improvement in lateral balance control, particularly of the standing leg of patients with SCI after robotic training. The improvement in the lateral balance of the standing leg may allow more time for patients with SCI to take a longer step length with the swing leg, which is supported by the trend of improvements in single leg support time after robotic training. In addition, given the importance of lateral balance control during walking, the improvements in lateral balance may reduce the energy cost [135], resulting in a more efficient gait pattern. This may be one reason why we observed a significant improvement in the endurance of patients with SCI after robotic treadmill training. In contrast, results from this study indicated that treadmill only training did not induce significant improvements in the walking function of patients with SCI, which is consistent with several previous systematic reviews [106, 136]. For instance, the average gain in 6-min walking distance obtained after treadmill only training was 11.0 ± 24.4 m, which is comparable to gains obtained from previous randomized

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A Flexible Cable-Driven Robotic System: Design and Its Clinical …

controlled studies using treadmill training [136], but is only approximately equal to 30% of the gain obtained after robotic treadmill training, i.e., 36.8 ± 30.5 m (>minimal clinical important difference, [137], and > clinical meaningful difference, which was suggested to be 31 m) [136]. One possible reason why treadmill only training seems less effective than robotic-assisted treadmill training may be that the challenge inherent in the task of treadmill only training for patients with SCI was not strong enough to induce improvement.

32.2.6 Improved Walking Function in Children with CP We observed significant improvements in selfselected walking speed and endurance after robotic training, but not after treadmill only training. The participants from the 3DCaLT robotic training group showed a greater increase in 6-min walking distance that those from the treadmill only training group. In addition, greater functional gains in walking were obtained for participants from the resistance group than those from the assistance group, and the functional gains obtained after resistance training were partially retained during the follow up period. Results from this study suggest that treadmill training in conjunction with the application of controlled forces to the pelvis and/or resistance force to the legs while allowing for a natural stepping pattern is effective in improving walking function in children with CP. Applying assistance load to the pelvis during treadmill training may improve weight shifting ability in children with CP. In this study, the 3DCaLT robotic system provided assistance force for facilitating weight shift at the targeted phase of gait while children with CP walk on a treadmill. The repeat practice of weight shifting during treadmill training may improve the weight shifting ability of children with CP through a use-dependent motor learning mechanism [86]. In addition, applying a mediolateral assistance force at the pelvis may enhance muscle activation of hip abductors/adductors, key muscles for

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balance control in the frontal plane during walking [138]. Further, repeat activation of these sensorimotor pathways induced by repeat pelvis assistance load during treadmill training may reinforce circuits and synapses used for lateral balance control during walking through usedependent neuronal plasticity mechanisms [36], leading to long-term improvements in lateral balance control. The improvement in balance control may lead to improved lateral stability on the stance leg, allowing for the contralateral leg to move forward, resulting in improvements in walking speed and endurance in children with CP after training. In addition, greater improvements in walking function were observed for children who underwent resistance training than those who underwent assistance training. One possible reason for the differences in functional gains may be due to children undergoing resistance training were more actively involved in the locomotor training session than those who underwent assistance training. Specifically, for children who were assigned the resistance training group, the resistance load applied to both legs during the swing phase may force participants to generate additional joint torque to counteract the load and move the leg forward during treadmill training, which may require participants to increase voluntary activation through enhanced supraspinal input to the motoneuron pool and/or increase motoneuron excitability [139]. In contrast, for children who underwent assistance training, the central nervous system may adapt to the assistance force applied to the leg(s) during the swing phase of gait by reducing the motor output of the leg muscles [140], probably due to optimization of the energy cost [141]. As a consequence, the training effect could be suboptimal. This is also consistent with a previous study, which indicated that only a modest increase in the gait speed of children with CP was observed after robotic training in which a passive guidance force was applied to both legs [77]. The functional gains obtained after robotic resistance treadmill training are comparable to or even greater than that from previous studies

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using conventional treadmill training. For instance, the functional gain in walking speed (i.e., 0.15 m/s, although 61.5 m, MCID) was greater than that after treadmill training with manual assistance (i.e., −25.0 − 19.8 m) [7, 55], which is comparable to functional gains obtained after assistance training (i.e., 13.9 m), and after robotic training using the GaitTrainer (i.e., 69 m) [143]. This study has several limitations. For instance, the sample size is small, which warrants further studies involving a larger cohort. In addition, while participants were randomly assigned into two different groups (i.e., 3D robotic and treadmill only training, or robotic resistance and assistance), physical therapists who conducted intervention and outcome assessments were not blinded, which might potentially bias the results. Further studies are needed to optimize this resistance training paradigm.

32.2.7 Other Advantages of the Cable-Driven Robotic System The cable-driven robot system can apply for compliant assistance as needed or even resistance as tolerated to the paretic leg (s) during treadmill training. The cable-driven robot system is easy to setup compared to an exoskeleton robot system, such as the Lokomat, which requires the rotation center of robotic arms to be aligned with the patient's hip and knee joints [62]. The setup time

M. Wu

of the cable-driven system is shorter than that of the exoskeleton systems, which is critical for the long-term treadmill training. In addition, the cost of the cable-driven robot system is less expensive than the current robotic systems, such as the Lokomat or AutoAmbulator. Thus, the cabledriven robotic system has multiple potential advantages to allow for the delivery of this type of therapy to a larger patient population. Robotic-assisted treadmill systems provide for the training of a repetitive walking pattern that is critical for locomotor recovery in individuals post-stroke or with SCI. However, the sensory feedback provided to the patients who are trained on the treadmill is distinct from overground walking. For instance, the optical flows are different for these two walking conditions. Visual cues are in conflict with proprioceptive signals from the legs during treadmill walking, which is not experienced during overground walking [144]. Such factors may limit the transfer of the motor skill learned on the treadmill to overground walking. For instance, a previous study showed a partial transfer of motor adaptation obtained from split-belt treadmill training to overground walking [145].

32.3

Conclusion

The cable-driven locomotor training system proposed in this study provides a promising adjunct for the treatment of patients post-stroke, patients with incomplete SCI, and children with CP through robotic-assisted treadmill training. This system is highly back drivable, complaint, and allows patients to voluntarily move their legs during BWSTT. The 3DCaLT can apply controlled assistance/resistance forces to the pelvis (in the frontal plane) and legs (in the sagittal plane) at the targeted phase of gait while subjects walk on a treadmill. Results from these randomized controlled studies suggest that applying targeted lateral assistance force to the pelvis during treadmill training seems more effective than treadmill only training in improving endurance in adults with SCI and in children with

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A Flexible Cable-Driven Robotic System: Design and Its Clinical …

CP. In addition, applying a targeted resistance force to both legs during treadmill training to induce active involvement of children with CP is crucial for improving the effect of treadmill training. However, applying a resistance force to the paretic leg during treadmill training was not greater than applying assistance force for improving walking function in individuals poststroke, which may be due to the compensatory movement from the non-paretic leg. In addition, the cable-driven robot is easy to set up and costeffective to allow for delivery of this type of therapeutic intervention to a larger patient population. Acknowledgements These studies were supported by NIH/NICHD, R01HD083314, R01HD082216, R21HD058267, R21HD066261, and PVA, #2552, NIDRR/RERC, H133E100007. We thank Dr. Schmit BD, Dr. Hornby TG, Dr. Gaebler-Spira DJ, Dr. Rymer WZ, Dr. Roth E, Dr. Yen SC, Dr. Wei F, Dr. Hsu CJ, Ms. Kim J, Mrs. Dee W, Ms. Arora P, Mrs. Rafferty M, Ms. Landry JM, and Mrs. Zhang YH for their assistance.

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treadmill walking in individuals with incomplete spinal cord injury. Phys Ther. 2006;86(11): 1466–78. Yen SC, Landry JM, Wu M. Size of kinematic error affects retention of locomotor adaptation in human spinal cord injury. J Rehabil Res Dev. 2013;50 (9):1187–200. Houldin A, Luttin K, Lam T. Locomotor adaptations and aftereffects to resistance during walking in individuals with spinal cord injury. J Neurophysiol. 2011;106(1):247–58. Yen SC, et al. Locomotor adaptation to resistance during treadmill training transfers to overground walking in human SCI. Exp Brain Res. 2012;216 (3):473–82. Reisman DS, et al. Repeated split-belt treadmill training improves poststroke step length asymmetry. Neurorehabil Neural Repair. 2013;27(5):460–8. Wu M, et al. Robotic resistance treadmill training improves locomotor function in human spinal cord injury: a pilot study. Arch Phys Med Rehabil. 2012;93(5):782–9. Pearson KG. Proprioceptive regulation of locomotion. Curr Opin Neurobiol. 1995;5(6):786–91. Wu M, Kim J, Wei F. Facilitating weight shifting during treadmill training improves walking function in humans with spinal cord injury: a randomized controlled pilot study. Am J Phys Med Rehabil. 2018;97(8):585–92. Inman V, Ralston HJ, Todd F. Human walking. Baltimore: Williams and Wilkins; 1981. Ballaz L, et al. Impaired visually guided weightshifting ability in children with cerebral palsy. Res Dev Disabil. 2014;35(9):1970–7. Liao HF, et al. The relation between standing balance and walking function in children with spastic diplegic cerebral palsy. Dev Med Child Neurol. 1997;39(2):106–12. Palisano R, et al. Development and reliability of a system to classify gross motor function in children with cerebral palsy. Dev Med Child Neurol. 1997;39(4):214–23. Russell DJ, et al. The gross motor function measure: a means to evaluate the effects of physical therapy. Dev Med Child Neurol. 1989;31(3):341–52. Wu M, et al. Effects of the integration of dynamic weight shifting training into treadmill training on walking function of children with cerebral palsy: a randomized controlled study. Am J Phys Med Rehabil. 2017;96(11):765–72. Savin DN, et al. Poststroke hemiparesis impairs the rate but not magnitude of adaptation of spatial and temporal locomotor features. Neurorehabil Neural Repair. 2013;27(1):24–34. Yen SC, Schmit BD, Wu M. Using swing resistance and assistance to improve gait symmetry in individuals post-stroke. Hum Mov Sci. 2015;42: 212–24. Wu M, et al. Repeat exposure to leg swing perturbations during treadmill training induces

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A Flexible Cable-Driven Robotic System: Design and Its Clinical … long-term retention of increased step length in human SCI: a pilot randomized controlled study. Am J Phys Med Rehabil. 2016;95(12):911–20. Wu M, et al. Robotic resistance treadmill training improves locomotor function in children with cerebral palsy: a randomized controlled pilot study. Arch Phys Med Rehabil. 2017;98(11):2126–33. Wu M et al. Robotic resistance treadmill training improves locomotor function in children with Cerebral Palsy: a randomized controlled pilot study. Arch Phys Med Rehabil. 2017. Edgerton VR, et al. Training locomotor networks. Brain Res Rev. 2008;57(1):241–54. Lewek MD, et al. Allowing intralimb kinematic variability during locomotor training poststroke improves kinematic consistency: a subgroup analysis from a randomized clinical trial. Phys Ther. 2009;89(8):829–39. Wolf SL, et al. Retention of upper limb function in stroke survivors who have received constraintinduced movement therapy: the EXCITE randomised trial. Lancet Neurol. 2008;7(1):33–40. Classen J, et al. Rapid plasticity of human cortical movement representation induced by practice. J Neurophysiol. 1998;79(2):1117–23. Mehrholz J, Kugler J, Pohl M. Locomotor training for walking after spinal cord injury. Cochrane Database Syst Rev. 2008(2):CD006676. Scivoletto G et al. Clinical factors that affect walking level and performance in chronic spinal cord lesion patients. Spine (Phila Pa 1976). 2008;33 (3):259–64. Lemay JF, Nadeau S. Standing balance assessment in ASIA D paraplegic and tetraplegic participants: concurrent validity of the Berg Balance Scale. Spinal Cord. 2010;48(3):245–50. Hornby TG, Reinkensmeyer DJ, Chen D. Manually-assisted versus robotic-assisted body weight-supported treadmill training in spinal cord injury: what is the role of each? PM R. 2010;2 (3):214–21.

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135. Donelan JM, et al. Mechanical and metabolic requirements for active lateral stabilization in human walking. J Biomech. 2004;37(6):827–35. 136. Mehrholz J et al. Is body-weight-supported treadmill training or robotic-assisted gait training superior to overground gait training and other forms of physiotherapy in people with spinal cord injury? A systematic review. Spinal Cord. 2017. 137. Norman GR, Sloan JA, Wyrwich KW. Interpretation of changes in health-related quality of life: the remarkable universality of half a standard deviation. Med Care. 2003;41(5):582–92. 138. Winter DA, et al. An integrated EMG/biomechanical model of upper body balance and posture during human gait. Prog Brain Res. 1993;97:359–67. 139. Ekblom MM. Improvements in dynamic plantar flexor strength after resistance training are associated with increased voluntary activation and V-to-M ratio. J Appl Physiol (1985). 2010;109(1):19–26. 140. Wu M, et al. Kinematic and EMG responses to pelvis and leg assistance force during treadmill walking in children with Cerebral Palsy. Neural Plast. 2016;2016:5020348. 141. Reinkensmeyer DJ, et al. Slacking by the human motor system: computational models and implications for robotic orthoses. Conf Proc IEEE Eng Med Biol Soc. 2009;2009:2129–32. 142. Norman G. The effectiveness and effects of effect sizes. Adv Health Sci Educ Theory Pract. 2003;8 (3):183–7. 143. Smania N, et al. Improved gait after repetitive locomotor training in children with cerebral palsy. Am J Phys Med Rehabil. 2011;90(2):137–49. 144. Torres-Oviedo G, Bastian AJ. Seeing is believing: effects of visual contextual cues on learning and transfer of locomotor adaptation. J Neurosci. 2010;30(50):17015–22. 145. Reisman DS, et al. Split-belt treadmill adaptation transfers to overground walking in persons poststroke. Neurorehabil Neural Repair. 2009;23 (7):735–44.

Body Weight Support Devices for Overground Gait and Balance Training

33

Andrew Pennycott and Heike Vallery

Abstract

Regaining the ability to walk overground, to climb stairs and to perform other functional tasks such as standing up and sitting down are important rehabilitation goals following neurological injury or disease. However, these activities are often difficult to practice safely for patients with severe impairments due to the risk of injury, not only to the patient but also to therapists. The emergence of various technologies that provide a degree of body weight support can play a role in rehabilitation focused on recovering overground gait and balance functions. These can greatly reduce the risk of falls and thus allow more intense and longer training sessions. Therefore, the systems empower individuals with the ability to practice the types of activities and functions they need in order to return home and to be A. Pennycott · H. Vallery(B ) Faculty of Mechanical, Maritime and Materials Engineering, Delft University of Technology, Delft, The Netherlands e-mail: [email protected] A. Pennycott e-mail: [email protected] H. Vallery Department for Rehabilitation Medicine, Erasmus MC, Rotterdam, The Netherlands

reintegrated into the community as much as possible. This chapter explores the origin of body weight supported devices and considers which groups could derive benefit from the training. An overview of the main training platforms available today—which comprise both robotic and non-robotic technologies—is then provided, followed by a discussion regarding outcomes of the devices thus far and possible future directions of the technology. Keywords

Robotics · Rehabilitation · Body weight support · Gait · Walking · Stroke · Spinal cord injury

33.1

Clinical Rationale for Body Weight Supported Training

33.1.1 Origins and Evolution In body weight supported (BWS) gait training, a harness is placed around a person’s torso and / or pelvis and connected to an unloading system in order to provide a variable degree of unloading as the subject walks. Originally motivated by a rich literature of studies with felines [6] and rodents [18] which demonstrated that stepping patterns could be restored through treadmill-based

© The Author(s), under exclusive license to Springer Nature Switzerland AG 2022 D. J. Reinkensmeyer et al. (eds.), Neurorehabilitation Technology, https://doi.org/10.1007/978-3-031-08995-4_33

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training with BWS, this training method enables patients with a high degree of weakness and/or poor coordination to undergo gait training following neurological and musculoskeletal injuries. One major advantage of body weight support devices is that they reduce the fear of falling and thereby allow individuals to concentrate more freely on the main training tasks. It has been shown that fear itself can have effects on gait characteristics; for example, older adults who reported being more fearful during walking tended to walk more slowly with a greater step width and shorter step length in one study [14]. There are low-cost solutions available for providing assistance from therapists during transfers and gait training such as the gait and walking belt devices [46], which allow therapists to apply supportive forces close to a patient’s centre of gravity and provide handles for assistance. It has been shown that the risk of falling was lower in rehabilitation programs incorporating gait belts and similar devices, and that using a gait belt during assisted falling reduced the risk of injury [50]. However, due to the exertion levels required and sometimes awkward postures required for performing assistance, manual gait training has been associated with increased risk of injury to therapists [12]. Moreover, research has shown that increased fall efficacy—a term representing an individual’s degree of confidence in conducting activities of daily living without falling—is a predictor of gait and indeed other functional rehabilitation outcomes [8]. Fear of falling may have contributed to the findings of some earlier research, for instance with incomplete spinal cord injured subjects [17], which pointed to greater gains being achieved through treadmill-based as opposed to overground walking, since in the latter studies, the subjects were supported only by simple ambulatory assistive devices such as crutches and canes and hence likely had a more pronounced fear of falling during the training. Another possible contributor may be that walking with body weight support could reduce fatigue and hence increase the feasible training duration. Indeed, it has previously been shown that body weight support reduces the net metabolic rate

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during gait [16]. Furthermore, ‘verticalisation’ of individuals, which can be promoted through BWS, is an important goal in rehabilitation of different groups such as stroke patients, since it can have benefits regarding circulation, prevention of pneumonia and clots, while also providing stimulation for the autonomic and sensory nervous systems [26]. The first generation of body weight support systems were mounted over a treadmill with therapists providing assistance to the patient. The observation that the training duration and consistency were frequently limited by therapist fatigue led to the development of robot-assisted gait training platforms such as the Lokomat [11]. Beneficial effects on kinematic, kinetic and spatiotemporal parameters of robot-assisted training have been noted, for instance, for chronic stroke patients [7]. However, other studies have suggested that for certain groups such as subacute stroke participants with moderate to severe walking impairments, the variability and diversity afforded by overground walking may be more beneficial than treadmillbased training with robotic assistance [19]. This could be partly due to the many changes in gait biomechanics that occur when walking on a treadmill [30]. It has also been argued that the treadmill belt could make the training less intense and challenging, and that the treadmill-based systems cannot provide task-specific overground training [25]. Several new body weight support systems now offer a safe environment that is not dependent on treadmills, which help individuals to overcome their fear of falling and thereby to practice overground gait. This technology could lead to improved rehabilitation outcomes compared with previous treadmill-centred methods. Researchers have argued that overground training could prove beneficial due to greater kinematic variability being permitted in addition to offering gait training in a more real-world environment [56].

33.1.2 Target Groups There are various groups of people who can benefit from body weight supported gait and balance training such as individuals who have suffered

Body Weight Support Devices for Overground Gait and Balance Training

from an injury or neurological pathology. Though the various patient groups are affected in different ways and have different patterns of gait and balance impairment, the main task of the training devices remains the same: to provide a degree of body weight support so that the people can perform gait and balance training with a reduced risk of falling. The incidence of spinal cord injury (SCI) has been estimated to lie between 10.4 and 83 per million people per year [57]. In addition to the detrimental effects on sensory and motor function, which can manifest in gait as reductions in gait speed and alterations in walking pattern [28], SCI can lead to secondary health conditions such as ulcers, amputations and major depressive disorder [27]. In addition to potential benefits with regard to ambulation, robot-based rehabilitation has shown positive results concerning posture, intestinal, cardio-respiratory and metabolic function [20]. Over 50 million people per year worldwide have a traumatic brain injury (TBI) [32], which is characterised as a change in brain function and other pathology due to an external force [36]. As well as the impacts on cognitive, behavioral, and emotional functioning, the effects on motor skills can lead to considerable alterations in gait, including decreased walking speed and step length, with exaggerated knee flexion on initial ground contact with the foot being evident [54]. The impact on walking also leads to increased incidence of falls in this group [35]. Independent walking is a common discharge rehabilitation goal for TBI patients [24]. Stroke is the second leading cause of death globally and a major cause of long-term disability in adults [45]. Furthermore, as stroke is associated with increased age, the prevalence of stroke could well be increasing due to the ageing of the population [15]. It can lead to impaired ambulatory function, which is in turn associated with a decreased quality of life [39]. Multiple sclerosis (MS) is a disease that causes the myelin sheath of nerve cells in both the brain and spinal cord to be damaged, leading to sensory and motor impairment and physical and mental issues. The incidence of multiple sclerosis is

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over 35 per 100,000 people and appears to be increasing with time [52]. MS manifests in gait by decreased gait speed, endurance, step length, cadence and joint kinematics [9]. For multiple sclerosis patients, in common with the other patient groups, gait is strongly associated with wider participation in society [23]. Cerebral palsy (CP) represents the most prevalent physical disability in children, affecting from 2 to 2.5 per 1,000 children in the United States [29]. CP can affect gait development in various ways, and CP patients typically have stiff knee action in the swing phase, crouched gait, excessive hip flexion, intoeing, and ankle equinus (limited dorsiflexion of the foot due to a lack of flexibility in the ankle joint) [55]. As noted for the other groups in this section, lower limb and gait dysfunction can have wider impacts on the overall quality of life for CP patients [22].

33.2

Overground Training Devices

33.2.1 Robotic Devices Motivation Robotic body weight support devices use actuation to control the forces acting on an individual, possibly in multiple directions and varying over the gait cycle. They can actively track the movement of a subject in both the vertical and horizontal directions during gait, thus avoiding undesired interference between the individual and the device itself. Adjusting the level of unloading is usually conveniently done through user interfaces that are operated by therapists. Moreover, the devices, which are summarised in the following sections, can be broadly categorised as being ceiling-mounted or mobile frames. Ceiling-Mounted Systems The ZeroG® gait and balance training system (Fig. 33.1a) is commercially available through Aretech, LLC (Ashburn, Virginia, US). The system can provide around 200 kg and 90 kg of static and dynamic of body-weight support, respectively. Mounted to an overhead track, a small motor propels a trolley to track a patient’s movements. More than one ZeroG trolley can be placed on the

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same track, thereby providing the opportunity for multiple patients to train simultaneously. Patients can practice walking overground, walk up and down steps, perform sit-to-stand movements and practise other balance-centred tasks. These activities are important since the patients will encounter such challenges in their everyday lives. The SafeGait 360◦ (Gorbel Medical, Victor, NY, U.S.) and the Vector Gait & Safety System® (Bioness Inc., Valencia, CA, U.S.) are further examples of ceiling-mounted body weight support systems that actively track a subject’s movement during gait in the longitudinal direction. A drawback of systems based on single rails is that they restrict the user to a specific path. Another limitation is that the cable will unavoidably transmit lateral force components whenever the individual moves laterally. Even if these lateral movements are small during gait, the ‘pendulum effect’ due to the mounting below a pivot could potentially disturb balance or support it more than intended or needed [13, 41]. This stabilising effect makes it more difficult to purposefully train lateral balance. However, there can be benefits from stabilising a patient in the lateral plane in some cases, particularly those with lateral propulsion syndrome [38]. Therefore, multiple systems have been proposed that enable training in a 3D workspace. The Free Levitation for Overground Active Training (FLOAT) system, commercially available through Reha-Stim (Schlieren, Switzerland) and shown in Fig. 33.1b), is a 3D body weight support system that capitalises on cable robot technology developed at ETH Zürich [48]. The system transmits forces to a person via wires that are actuated by motorised winches positioned at the ceiling in the four corners of the desired workspace. Conventional cable robots tend to have a limited workspace because the angle of the cables with respect to the horizontal plane determines how much force needs to be transmitted for a given vertical unloading level. To enable a large workspace without incurring excessive cable forces, the FLOAT uses a mechanical configuration of moving cable deflection units (pulleys) [51]. The deflection units are not actu-

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ated but rather are moved by the tension in the cables they deflect. The design reduces moving masses to a minimum and it enables control of a three-dimensional (3D) force vector, including relieving patients of a percentage of their body weight and providing longitudinal assistance or resistance. Nevertheless, a design drawback of the FLOAT is that, as in conventional cable robots, all the winches must act in combination to actuate the different degrees of freedom, rather than in a decoupled fashion. This requires high-power actuators because they need to serve both the lowspeed, high-force vertical degree of freedom and the high-speed, low-force horizontal degrees of freedom. The RYSENTM (Motek Medical B.V., Houten, the Netherlands) is also a cable robot but has lower-power motors. This 3D body weight support system (Fig. 33.1c) mechanically decouples the degrees of freedom across the motors such that these can make different speed-torque tradeoffs [42]. The RYSEN uses moving cable deflection units on rails and posterior-anterior motion is actuated by motors. Springs are applied in series with two main motors being used for predominantly vertical actuation and a double-sided variable-radius winch for lateral actuation. The low-power motors limit the bandwidth of closedloop control in the vertical direction but the design does achieve very good force tracking in the horizontal direction [42, 43]. Besides cable robot technology, another option to provide 3D body weight support is to use gantry systems. The Active Response Gravity Offload System (ARGOS) is used by NASA for astronaut training in simulated reduced gravity environments and is based on an active overhead gantry crane system. Motion in the horizontal plane is controlled by electric motors, while the degree of body weight support is controlled by a crane connected to the user via a steel cable and shock absorber. The NaviGAITor is a similar example of a multidirectional body weight support system [44]. A movable bridge is mounted on a pair of rails to enable movement in the longitudinal direction, while lateral motion is permitted by movement

Body Weight Support Devices for Overground Gait and Balance Training

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Fig. 33.1 Robotic body-weight support devices. a Zero-G (courtesy of Aretech, US), b THE FLOAT (courtesy of REHA-STIM MEDTEC AG, CH), c Motek Medical’s RYSENTM at the Rehabilitation Center Heliomare, Wijk aan Zee, NL (courtesy of Motek Medical), d ANDAGO (courtesy of Hocoma AG, CH)

of a trolley along the bridge. Actuation is realised in these two directions via electric motors, and a further motor drives a hoist to provide vertical forces in order to realise different degrees of body weight support. A major advantage of 3D systems over systems based on single rails is that, in principle, they do not restrict the user to a specific path, and, if they are sufficiently transparent, do not cause restoring horizontal forces, thereby preventing the aforementioned pendulum effect. However, a 3D setup may impose different limitations on the practice space; for example, it fits less easily into narrow or curved rooms or corridors. Furthermore, such a device can typically only support one patient at at time, the practice walking length is limited, and the design may require high ceilings for installation. An inherent limitation of the gantry-based systems is that large masses need to be moved when tracking a walking user. Even if closed-loop force control is applied, the user will feel some remaining inertia because the reduction in apparent inertia is limited in causal control schemes [10]. This

limits the devices’ ability to render purely vertical forces on a user, potentially disturbing their gait by imposing undesired horizontal force components. Mobile Frames The Andago® (Fig. 33.1d), developed by Hocoma AG (Volketswil Switzerland), comprises a mobile frame mounted on wheels and a BWS [34]. Patient trunk movements are tracked, and hence he or she can practice walking without being confined to one specific training room. The training platform can be used in patient-following mode in which the person’s movements are followed not only in the forward and backward directions but also in turning movements, straight-line mode in which turning inputs are not followed, and finally in manual mode in which the device is controlled by a therapist via a joystick. A further system which was originally mobile is the KineAssist [40], which interacts with users through a pelvis and torso harness. The device senses interaction forces at the pelvis and thereby controls the movement of a robotic platform. Today, however, the KineAssist is only available as a treadmill-mounted system from Woodway (Waukesha, WI, U.S.).

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Fig. 33.2 Passive body weight support devices. a Zero-G passive (Courtesy of Aretech, US), b LiteGait (courtesy of Mobility Research, US)

33.2.2 Non-Robotic Devices Motivation Various non-robotic systems are available on the market that provide static body weight support during overground gait and balance training. The principal advantage of these systems is clearly their cost: they do not require the expensive sensing and actuation hardware of their robotic counterparts. The disadvantage is that the vertical support cannot be precisely controlled and the options for applying forces in the horizontal plane—for instance perturbation forces for balance training—are more limited. Nevertheless, the devices do allow people who cannot support their entire body weight to practice overground walking. Ceiling-Mounted Systems A large number of commercial solutions are available that are based on passive trolleys mounted on rails in order to provide support during overground ambulation. For example, the design of the ZeroG-Passive system (Aretech, LLC, Ashburn, VA, US) closely follows the robotic ZeroG platform, but rather than actuating the trolley position via a motor, it is simply pulled along by the patients as they ambulate (Fig. 33.2a). The FreeStep SAS (Biodex, Inc, Shirley, New York, US) operates similarly to the ZeroG-Passive and also uses small and lightweight trolleys.

In light of the findings concerning how BWS reduces the risk and fear of falling, a particularly innovative building design is the Shirley Ryan AbilityLab in Chicago (formerly known as the Rehabilitation Institute of Chicago, RIC). Many areas of this hospital are equipped with rails mounted to the ceiling such that training with a safety harness and passive trolleys is not limited to dedicated gym areas. Instead, users can walk with a safety harness along ‘patient highways’ in many locations in and close by their own rooms. This reduces the risk of falls and reduces the transfer time between therapy sessions. If these aspects are considered when planning the construction of a hospital or rehabilitation center, the installation of such rails is much less expensive than retrofitting. However, the noise levels of passive carts on the rails should also be taken into consideration. Mobile Frames The LiteGait system by Mobility Research (Tempe, AZ, U.S.) comprises a mobile cart mounted on castors and an overhead bracket for attachment to a harness (Fig. 33.2b). Owing to a locking system, the device can be used for either treadmill or overground training. Other castor-mounted systems include the NxStepTM Unweighing system by Biodex (Shirley, New York, US) and the PhysioGait

Body Weight Support Devices for Overground Gait and Balance Training

(HealthCare International, Langley, US). Like their mobile robotic support counterparts, mobile frames are not tied to being used in a specific training room. However, a disadvantage is the relatively large mass of the device that must be moved by the patient, which potentially affects the gait biomechanics.

33.3

Device Characteristics

33.3.1 Transparency To maximise motor learning outcomes and functional gains, training devices should only provide just enough support to create a safe environment; the devices should be able to ‘hide’ their presence in terms of interaction forces between the device itself and the user. The degree to which this is possible is often referred to as transparency [5]. This can be promoted through hardware design that emphasises low inertia and friction as these factors can lead to higher interaction forces and distort the mechanics of gait. In robotic body weight support devices, the ability of a system to track the movement of an individualwillgoverntherealisabledegreeoftransparency to a large extent. As mentioned in Sect. 33.2.1, even with closed-loop force control, the device dynamics cannot be ‘hidden’ completely, and the physical mass of a device is a governing factor of transparency, with friction between static and moving parts also playing a role. Again, systems mounted on mobile platforms and 3D gantry-based systems tend to have limited transparency since relatively large masses must be moved by the user.

33.3.2 Vertical Support Forces For body weight support systems, levels of body weight support above 30% appear to significantly alter some gait characteristics such as the kinematics and kinetics of the hip and knee joints [3]. Indeed, modelling has shown that there could be changes in the sign of joint moments above this level of unloading [43]. Therefore, an excessive

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level of body weight unloading potentially causes large kinematic and kinetic changes from physiological gait and therapists should aim to apply levels of vertical support below this apparent threshold during training sessions. Besides the mean value of support over the gait cycle, variations in vertical unloading should also be considered. Some systems intrinsically exhibit variable forces, especially when using counterweights or springs. A prevalent paradigm is to keep vertical unloading forces as constant as possible, with the argument in favour of this approach being that gravity itself is constant. This presumed requirement increases device complexity to keep force levels constant despite the oscillatory vertical movement of the user during gait. This goal is often pursued by closed-loop force-controlled actuation. However, although a constant unloading vertical force can relieve the body of a portion of its weight, its inertia in the vertical direction remains unchanged, which disturbs the relationship between gravitational and inertial forces and may thus alter the natural frequency of gait. In fact, simulation results suggest that a simple passive spring suspension—in which a spring stiffness is chosen to favour natural gait cadence— distorts gait mechanics to a lesser extent than more complex approaches that actively control the vertical force to a constant level [4]. During gait, acceleration-dependent inertial forces can largely be compensated by position-dependent spring forces and the relationship between weight and inertia is thereby maintained constant despite the unloading, and therefore, the gait frequency remains closer to its physiological value (withoutunloading). Hence, a passive elastic suspension is not only easier and less expensive to realise, but also may disturb gait to a lesser degree than an actively controlled constant force suspension. This insight can, furthermore, influence the design of robotic systems, namely the inclusion of series-elastic elements in combination with lowbandwidth actuators. For example, the RYSEN employs series springs that cover oscillations in the vertical direction during gait where the lowpower motors lack bandwidth. Only for move-

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ments with lower frequency need the motors move to change the setpoint of the springs (mainly in order to manoeuvre over stairs or to sit down).

33.3.3 Longitudinal Forces The notion that individuals would prefer zero or positive forces in the longitudinal direction (forces acting in the forward direction of ambulation) due to energy efficiency considerations has been challenged by research using the RYSEN [43]. Though resistive longitudinal forces applied by the devices would require greater energy consumption by the subject during gait, healthy subjects in this study tended to favour small negative forces when walking in the robotic platform. The reason suggested was that unlike a backward fall, a forward fall can be recovered through a swift movement of the swing leg and, therefore, a small negative force encouraging a slight forward lean during gait could allow subjects to feel safer and have less fear of dangerous falls. Overground gait training devices can also include perturbation forces in various directions, including along the longitudinal axis. For example, an additional module has been integrated into the FLOAT platform to apply perturbation forces in different directions [37]. It has been argued that applying perturbations in this way, as opposed to a treadmill, presents less risk of falling and hence less risk of injury and a reduced sense of fear during the training.

33.3.4 Lateral Forces As mentioned in Sect. 33.2.1, research has pointed to a pendulum effect from the body weight support mechanism itself: the systems can produce lateral restoring forces which decrease the challenge of maintaining balance in the frontal plane. Therefore, systems that reduce these restoring forces could be useful for gait training that incorporates active balance control. Systems that do not permit lateral movement are likely to lead to greater restoring forces and hence may limit the ability

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to train balance in the frontal plane, which is an important element of gait [33]. However, systems that, in principle, do allow lateral movement may still generate undesired lateral forces, depending on their degree of transparency. Moreover, the point of application of the resultant unloading force needs to be considered to know which stabilising or destabilising moments are applied about the center of mass as discussed in the next section.

33.3.5 Harness and Attachment Ease of attachment to the harness is important as long set-up times will detract from the time available for actual training. Wheeled mobile platforms should offer sufficient space so as to allow wheelchair access and to avoid necessitating additional transfers. Comfort while exercising in the harness should also be considered in the design as discomfort may limit the training duration tolerable by the subjects and, similarly to the fear of falling, may detract from their concentration on the active gait task. This has been considered in some designs for which different harness versions are available for male and female users. The attachment of the harness also determines the line of action of the resultant force acting on the user and how the unloading forces are distributed along the body. The location of the line of action with respect to the body’s center of mass defines the resultant moments of the unloading force. Inmoments caused by ground reaction forces, this governs the rate of change of centroidal angular momentum, which is a key variable for bipedal stability and balance control. Results with able-bodied subjects indeed indicate that the self-selected walking speed is affected by the attachment mechanism [43]. Furthermore, the optimal type of attachment may vary according to the degree of impairment of the subject. For less impaired subjects, attachment near the pelvis could be better since this will not produce stabilising moments about the hip joint, while for individuals who are less able to maintain their balance independently, attachment

Body Weight Support Devices for Overground Gait and Balance Training

higher up on the torso would be better due to the stabilising moments this affords. For treadmill walking, attachment via a harness has been shown to reduce vertical acceleration. This is due to restrictions in both linear and rotational movements imposed by the harness that lead to the trunk being less able to absorb shocks than in normal gait [1].

33.4

Outcomes of Overground Gait Training

There have been a limited number of studies thus far which have investigated the outcomes of applying the various devices described here to overground gait and balance for different patient groups. In contrast, there has been a greater volume of research focused on the outcomes of treadmill-based training and rehabilitation as summarised, for instance, in the review by Wessels et al. [53]. Although research has suggested that conventional therapy incorporating overground walking with manual assistance seems to yield broadly comparable outcomes as compared to treadmill training with BWS [31], there have not been comprehensive comparisons between overground training using BWS and treadmill-centred training. Nevertheless, the results from studies using BWS devices in rehabilitation have generally been encouraging so far; some of the main findings from the studies with overground devices are summarised below. Huber and Sawaki compared overground gait training for non-traumatic SCI using the Zero-G system, finding better outcomes in terms of sphincter control for the robot-assigned group compared to standard-of-care therapy [21]. Both groups achieved significant gains in Functional Independence Measure (FIM) scores but no significant cross-group differences were apparent. Anggelis et al. compared the outcomes of training with and without dynamic body weight support using the Zero-G platform in traumatic brain injury patients [2]. The Zero-G group demonstrated significantly greater improvements in functional independence measures (FIM) as

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well as cognitive improvements than the control group. The authors suggest that this is due to the greater intensity facilitated by dynamic body weight support systems. Furthermore, the importance of reducing the fear of falling and thereby allowing the individuals to concentrate more fully on the training tasks was highlighted. Brunelli et al. conducted a study with subacute stroke patients and compared body weight supported overground training with the LiteGait to conventional physiotherapy [8]. While both groups showed improvements in all the outcome measures—which included the Rivermead Mobility Index, Barthel Index, and the Six Minute Walk Test—greater gains were shown concerning the Functional Ambulation Classification for the individuals who participated in the BWS training. Tay et al. compared outcomes for patients undergoing training with the Andago system in addition to conventional therapy to other individuals who only had conventional therapy [47]. Statistically significant FIM gains were made in both groups; though the robotic training group showed greater improvement in functional ambulation scores, there were no statistically significant differences in the other outcomes. Van Hedel et al. assessed the use of the Andago platform for children and youths with gait impairments [49]. They observed that the device prevented several falls and also that variability in stride duration and the degree of inter-joint coordination were higher with the overground training than treadmill walking.

33.5

Future Directions

The cost of the various overground robotic devices remains an obstacle to their widespread adoption in rehabilitation programs. Notably, the actuators used to track gait movement lead to additional safety requirements and hence higher certification costs. Non-robotic systems are, therefore, much less expensive than robotic body weight support platforms. Since the former category may provide most of the potential benefits of the body weight support systems at much lower cost, research comparing usability and clinical outcomes—for exam-

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ple in terms of functional gait—between passive and robotic systems is needed in order to evaluate whether the additional costs of actuation are actually justified. More generally, there remains a paucity of results concerning clinical outcomes from using the various devices for different patient groups. Additional evidence for the efficacy of the devices from studies with larger subject groups is needed to justify funding of the devices by health care systems. Though there are promising results showing improvements achieved through overground robotic training, whether these lead to improved ambulation in the long term remains uncertain. Therefore, studies including longer-term followups for the various different types of training are needed, along with investigations as to whether or not training in the clinical setting can really translate into increased community participation and integration. Acknowledgements This chapter is an update of a previous edition that was co-authored by Joe Hidler (Aretech, US) and Arno Stienen (Motek Medical B.V., Houten, NL). Disclosure of Financial Interests H. Vallery is an inventor on multiple patents and patent applications related to the FLOAT and to the RYSEN and may financially benefit from sales of either device.

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Epilogue: Robots for Neurorehabilitation—The Debate

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John W. Krakauer and David J. Reinkensmeyer

Abstract

There is an ongoing debate over the value of robotic movement training devices for clinical rehabilitation practice and their promise for the future. In this Epilogue, we break down the debate into a series of specific propositions and then provide commentary from two perspectives. JK is a neurologist and neuroscientist who has developed novel location-based immersive training environments combined with proprietary gaming software that encourages active exploration

J. W. Krakauer Departments of Neurology, Neuroscience and Physical Medicine & Rehabilitation, Baltimore, MD, USA e-mail: [email protected] The John Hopkins University School of Medicine, Baltimore, MD, USA The Santa Fe Institute, Santa Fe, NM, USA D. J. Reinkensmeyer (&) Department of Mechanical and Aerospace Engineering, University of California at Irvine, 4200 Engineering Gateway, 92617 Irvine, CA, USA e-mail: [email protected] Department of Anatomy and Neurobiology, University of California at Irvine, 4200 Engineering Gateway, 92617 Irvine, CA, USA Department of Biomedical Engineering, University of California at Irvine, 4200 Engineering Gateway, 92617 Irvine, CA, USA

to reduce upper limb impairment after stroke. DR is an engineer and neuroscientist who has developed robotic and sensor-based rehabilitation training systems and computational models of motor recovery. Keywords

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Robots Rehabilitation Stroke Movement training Sensory-motor recovery Plasticity Motor learning Somatosensation

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Proposition 1: Robotic movement training devices are useful scientific tools. DR: Agree. Robotic movement training devices such as those shown in Fig. 34.1 have facilitated key contributions to rehabilitation science. In terms of scientific method, they have made scientists and clinicians define the content of rehabilitation movement training because they must be specifically designed and precisely programmed before they can be used to do anything. And then, when they are used to assist in movement training, they maintain a record of every force and movement the patient experiences, generating an abundance of information that has only just begun to be tapped for scientific purposes. These twin “superpowers” of codification and quantification have led to more precise hypotheses—about the effects of different types of training paradigms and about how we might detect and predict movement recovery.

© The Author(s), under exclusive license to Springer Nature Switzerland AG 2022 D. J. Reinkensmeyer et al. (eds.), Neurorehabilitation Technology, https://doi.org/10.1007/978-3-031-08995-4_34

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J. W. Krakauer and D. J. Reinkensmeyer

Fig. 34.1 The authors in robotic movement training devices

Some key scientific findings enabled by robotic devices include the importance of efference (i.e., active patient effort) in movement training [1–3], and relatedly, the algorithmic tendency of the motor system to “slack” when physical assistance is provided [4–7]; the interplay of error, challenge, and motivation in stimulating motor learning [8– 12]; the presence of multiple learning/adaptation process during recovery [13, 14]; the importance of proprioception for the effectiveness of motor training [15]; and the diagnostic power of movement kinematics such as movement smoothness [16–18]. There have also been many equivocal therapeutic findings in studies conducted with robotic therapy devices. This is likely due to underpowering related to the high variability and modest size of treatment response, even with the improved quantification ability. But at least robots have forced scientists to ask well-defined questions, which is an improvement over the field from when I entered it 30 years ago. JK. Agree. I think that robots broadly defined can be thought about and used in three ways. The first is as scientific tool to study motor control and motor learning. The seminal work on forcefield adaptation by Reza Shadmehr and

colleagues is a key example. Second, robots can be used to quantify motor deficits in patients, the studies by Stephen Scott and colleagues, and Amy Bastian and colleagues serve as good examples of this. Third, and most questionably in my view, is as therapeutic devices. Proposition 2: The treatment effect due to robotic movement training is so modest that the technology is not a clinically useful tool. DR: Agree with caveats. Yes, the mean treatment effect due to robotic therapy is modest, below what can be considered clinically important. But note that this isn’t a result unique to robotic therapy—it’s a problem with all neurorehabilitation movement training and probably reflects fundamental limits imposed by damaged neural tissue. Also note that when you consider the statistical distribution of patient responses, some percentage (I estimate 10–30%) experience treatment effects due to robotic therapy that are above the minimally clinically important difference (MCID). So, while the statement is true on average, it’s not accurate for all individuals. Much of the challenge of robotic therapy (and neurorehabilitation in general) now lies in

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Epilogue: Robots for Neurorehabilitation—The Debate

figuring out who responds, to what protocol, and why, in order to better shape patient and treatment selection. Finally, robotic movement training devices typically remain relatively difficult to use and thus are not clinically viable because of uptake issues, not effectiveness issues (they mostly match and sometimes exceed the results achievable with conventional training—see below). Thus, the dream of providing therapists and patients with a tool that allows them to automate key aspects of movement training with less cost has only been achieved sporadically by certain therapists and facilities. JK: Agree with caveats. The reason why the treatment effect of robots has been modest (to say the least) is because of conceptual confusion rather than a failure of engineering. This confusion is a sad echo of a similar conceptual confusion that lies at the heart of rehabilitation in general. It is a strange move to defend robotics by saying that the whole field is doing poorly! What is the confusion? It takes two forms—the first is the assumption that assisted repetition of movements is an effective form of training—it isn’t. The second is the assumption that motor learning is the way to fix a motor control deficit—it isn’t. With regard to the first false assumption—completing the proprioceptive feeling of the correct kinematics of a movement is no more likely to help you find the motor commands required than watching someone serve will turn you into a tennis player. Unfortunately, this haptic guidance fallacy will not die and it is not rescued by the fact that patients make some initial contribution to the movement. With regard to the second assumption, motor learning capacity is normal in patients with hemiparesis from corticospinal lesions—if learning is all that was required for repair then patients who have had years to learn after stroke should all have recovered! Proposition 3: Continuing to refine the device design and training algorithms won’t improve on the treatment effect size enough to matter. DR: Agree with caveats. I must admit that I am skeptical that continuing to tweak device designs or movement training algorithms can have a major effect unless coupled with plasticityenhancing interventions. However, can science

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really ever confidently assert an absolute negative? Recall dogmas such as “There are no physical laws other than Newton’s” or “The adult brain can’t create new neurons”? Somewhere out there, there is an ingenious Ph.D. student who just might knock it out of the park with a new approach. And, in the meanwhile, I like to keep in mind that science is often Edisonian, in the sense that it relies on extensive trial-and-error, but also, and perhaps more relevantly, in the sense that it requires protracted effort and the careful assembling of incremental improvements to finally generate meaningful effects. Take as one inspiring example the evolution of the treatment of childhood leukemia [19]. JK: Agree with caveats. Again what is needed is clearer biological thinking not better engineering. As is so often the case these days, the methodological tail wags the conceptual dog. It is an unfortunate historical fact that rehabilitation of late has been driven by an unholy alliance between engineers and clinicians with a relative lack of neuroscience in between. This technoutopianism badly needs tempering. What it has led to is a fuzzy thinking and confusion. For example, is a robot meant to mimic, assist, or ultimately substitute for a therapist? Or is a robot meant to be a complementary approach? The lack of clear answers to these questions is exasperating. That said, I do think that a next generation of robots could be used to enhance recovery of the upper limb after stroke and other forms of neurological injury. For example, to provide varying degrees of weight support, allow bilateral training, and restrict compensatory strategies. What we must stop doing is seeing robots primarily as devices to provide triggered assistance along the direction of movement. Indeed, a recent study showed that robotic weight support was sufficient to improve 3D movements in a circle-drawing task with no additional improvements obtained when path assistance was added [20]. Thus, to the degree that some therapeutic gains have been seen for assistive modes, they are small and probably attributable to a second-order effect acting through motivation and effort, rather than any direct effect on either learning or repair.

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Repetition of this kind is not synonymous with deliberative practice. As to motivation, there are other ways to promote it that do not require expensive actuated robots [21]. Proposition 4: Refining patient selection criteria won’t improve on the treatment effect size enough to matter. DR: Disagree. I think the evidence is already pretty clear for the upper extremity after stroke that people with some residual corticospinal tract will benefit more than people with totally disrupted CST, and the size of the additional benefit is above the MCID for a fraction of patients [22, 23]. Timing matters enough to be meaningful as well, as the recent study by Dromerick et al. convincingly showed (training in the subacute phase was significantly more beneficial than training in the chronic phase) [24]. Baseline proprioception is a powerful predictor of benefit as well [15, 25]. The challenge comes in the logistical and ethical implementations of these criteria. But I do believe that, if they could be implemented appropriately, the result would be to bump up the mean treatment effect of clinical rehabilitation practice. JK. Agree. It is always a last resort to say that there is a subset of patients that will respond to a treatment that is not showing impressive results overall. I of course agree that we should stratify intelligently, especially when investigating new approaches. It would have been odd if the first heart transplants had been tried on octogenarians with severe medical comorbidities. That said you can’t rescue a new approach that was illconceived conceptually from the get-go by holding out hope that there are some patients out there that will show large effect sizes. It is indeed the case that there are some chronic patients that respond well to larger doses and intensities of rehabilitation of any kind. Whether this is true recovery/restoration versus peripheral effects, strengthening or compensation remains unclear. Again, what is the specific biological hypothesis about robotics over other forms of approach that make it the ideal choice for a subset of latent recoverers? I have yet to hear a clear answer.

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Proposition 5: Robotic therapy treatment effects would not be significantly amplified even if patients were given orders of magnitude more training time with the devices. DR: Disagree with caveats. The amount of movement practice typically achieved by neurologic patients is woefully small—in-clinic therapy is limited and home adherence is low, even if the prescribed exercise program is unambitious. Historically, even studies that have set out to study “high dose” have probably undershot the target as well. Jeffers et al.’s meta-study in a rodent model of stroke suggested the dose–response curve is a nonlinear, “hockey stick” function—flat at first, up to a relatively high threshold of training activity, and then correlated thereafter [26]. For walking after stroke, a recent study convincingly showed the benefit of a large dose of walking compared to a small dose [27]. But in that study, which had three dose levels, the therapeutic effects of the two higher dose levels were about the same, and those effects, while clinically significant, were just clinically significant. So, maybe it’s a “broken hockey stick”. Variations in low doses (which clinicians might sometimes mistakenly think of as relatively “high” because of historic practice habits) don’t matter (the blade); variations that get past the threshold matter (the shaft), but there is no zone of “hyperdose” that can ultimately overcome the effects of severe neuroanatomical damage (the broken, top shaft). See Chap. 3 for an extended discussion of these issues. JK Agree. We can all agree that patients need more behavioral intervention at higher dose and intensity focused on impairment reduction. The question is what should this behavioral intervention be? At first blush, robots would seem to be ideal—high-dose and high-intensity movement repetition machines! Alas, this superficial impression is wrong but still, as I have said above, holds powerful sway over the minds of many in the field. Motor learning is never about simple repetition and recovery is not about motor learning. What we need to do is take a step back and think. The success of treadmill training coupled with epidural stimulation for spinal cord injury is

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a reason for optimism and points to a potential analogous use of upper limb robots as “treadmills” for the arm. As I said above, for a robot to be effective it must help patients find their own commands not just provide them with kinematic sensory experiences. I am just not convinced that assistive robots are the best way to promote exploration of residual command space. Proposition 6: There are no mechanistic reasons to believe that providing active assistance with a robotic therapy device should be more effective than conventional rehabilitation exercise that does not provide physical assistance. DR: Disagree with caveats. In the case of gait training, if some physical assistance isn’t provided, the walking task can’t be safely practiced. In that case, it is difficult for me to believe that providing no assistance can achieve similar results to providing some assistance (robotic or not). But even when there aren’t safety issues, providing active assistance with a robot can provide benefits in at least three ways. First, providing active assistance improves motivation to engage in training [10]. Improved motivation can in turn improve not only dose achieved (by encouraging a greater willingness to engage in practice) but also the amount of motor learning achieved (possibly through dopaminergic mechanisms associated with experiences of success) [28, 29]. Yet, it may also be possible to generate similar motivational effects without active assistance provided by robots, for example, through sensor-based systems with clever feedback about success, or by sociopsychological effects—such as collaboration, cooperation, or simply practicing in a group [30]—it remains to be seen. Second, active assistance promotes neural strengthening. Weakness is a strong predictor of limb function after stroke but cannot be attributed to muscle atrophy. Repeated stimulation of descending pathways improves force output by multiple mechanisms. For example, repeated efferent stimulation that is accompanied by movement performance feedback likely helps the motor system sort out which residual CST

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neurons are most useful to activate desired muscles with adequate force. We have previously laid out a model of this concept of neural strengthening with mathematical detail [31–33]. The point for this discussion is that a robotic device that appropriately titrates mechanical assistance will help a person repetitively engage damaged descending pathways (in a motivated way!—see above). This will cause improvements in the neurally mediated component of strength, and, indeed, this is a well-documented outcome of robotic movement training [34, 35]. Improvements in neural strength should be expected to translate, in turn, to some amount of impairment reduction on a variety of scales (and indeed, neural strengthening is impairment reduction by definition!). For some subsets of activities, this will translate to modest improvements in function. Third, movement with active assistance promotes somatosensory stimulation compared to what happens during self-driven movement that is very impaired. And by stimulation I don’t mean kinematic demonstration, which I agree has limited benefit not only in movement rehabilitation but also in movement training with unimpaired participants [36]. Rather, I mean generation of a greater diversity of somatosensory input. Such somatosensory stimulation may then work through use-dependent, Hebbian, or reward-based learning mechanisms to drive beneficial cortical plasticity. If the somatosensory story turns out to be true, then the role of robots may increase, as stimulating somatosensation seems like it will require active technologies that impose movements. In summary, then, what do assistive robots bring that can serve as a useful tool for therapists and patients? Imagine you are a therapist who would like your very weak patient to practice a large dose of functionally relevant movement in a motivating way. You also want to make sure that the patient is motivated and continuously puts out the effort needed to rebuild neural strength. Then, as a bonus, you would like them to experience a diversified portfolio of somatosensory input. Ideally, you’d like for them to do this without you being constantly present, as you are incredibly busy and the amount of reimbursed time you have with them is limited.

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JK. Agree with caveats. At the risk of repeating myself: What is the extra thing that a robot does compared to conventional therapy? Is there a qualitative difference or a quantitative one? “Proprioceptive stimulation” is a slippery notion. As David agrees, it can’t just be proprioceptive “demonstration”, analogous to watching the correct movement. So what is it? Is the proprioceptive stimulation instructing cortex or is it increasing the receptivity of segmental circuits to residual descending commands? Not yet knowing the answer would be ok if the effect sizes of robotic intervention were large, but sadly they are not. In the recent RATULS trial [37], the largest robotic rehabilitation trial to date, robotassisted training of the upper limb was no better than usual care on the primary outcome measure, the ARAT at 3 months. Additionally, robotics was no better than enhanced upper limb training (EULT) on the Fugl–Meyer score and was inferior to UELT on ADLs at 3 months. These results are, to say the least, hardly a ringing endorsement for robotic rehabilitation no matter how badly David wants to put a positive spin on RATULS (see below). My belief is that assistive robotics falls between two chairs—it is neither optimal for impairment reduction nor functional training. Robots have just not worked out. If they are to have a future in rehabilitation then movement-path assistance needs to go to the back of the line. There are other ways to “help a person repetitively engage damaged descending pathways (in a motivated way!)” as David states, without path assistance. Proposition 7: Robotic training holds no advantages compared to other non-robotic technologies for training, such as sensor-based training. DR: Agree with caveats. In RATULS, as in the vast majority of previous trials of robotic movement training, robot-assisted training was no worse than some form of enhanced upper limb training for the primary outcome (ARAT success) and it was better than usual care for

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impairment reduction (measured by UEFM change, which, by the way, is meaningful to patients at around 4 points [38, 39]—RATULS found the mean benefit over usual care was 2.8 points, indicating that a substantial fraction of robot participants exceeded by 4 points what usual care participants achieved). But, surprisingly to me at this mature stage of clinical testing, the robotic movement training in RATULS was delivered “face-to-face” “with a one-to-one patient-to-therapist ratio” (quotes from the study). The study thus did not test what is perhaps the most obvious potential advantage of robotic movement training: motivating realworld patients to engage in further beneficial training outside of formal therapy times. That said, I have to acknowledge that my ideas about the uniquely, positive package of influence that active assistance might exert over and above sensor-only approaches still needs to be further experimentally verified. I could fall back on the robot superpowers of codification and quantification mentioned above, but sensors have these superpowers to some degree as well and are missing only the power of being able to precisely control patterns of force application. So, pragmatically, yes, I agree—use of sensor-based systems for routine clinical practice seems warranted for the most part at this time. By way of full disclosure, I’d like to note that the two systems that I have worked on that have been commercialized with some success—as ArmeoSpring and MusicGlove—are essentially sensorbased systems with some functionally targeted. non-motorized, mechanical design. ArmeoSpring foregoes path assistance, as John suggests, providing gradable weight support for 3D hand movement. JK: Agree with caveats. Actuated robots are not going to be required in general—can use far cheaper devices for weight support and sensing. That said, I do think that a next generation of robots could occupy a niche for specific forms of intervention and measurement, especially of forces, that address the problem of 3D movements

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Epilogue: Robots for Neurorehabilitation—The Debate

in general, and the complexities of the paretic shoulder in particular. Acknowledgements DJR supported by NIH grant R01HD062744 and NIDILRR grant 90REGE0005-01. Disclosure David Reinkensmeyer has a financial interest in Hocoma AG and Flint Rehabilitation devices, makers of rehabilitation equipment. The terms of these arrangements have been reviewed and approved by the University of California, Irvine in accordance with its conflict of interest policies.

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Correction to: Spinal Cord Stimulation to Enable Leg Motor Control and Walking in People with Spinal Cord Injury Ismael Seáñez, Marco Capogrosso, Karen Minassian, and Fabien B. Wagner

Correction to: Chapter 18 in: D. J. Reinkensmeyer et al. (eds.), Neurorehabilitation Technology, https://doi.org/10.1007/978-3-031-08995-4_18 Chapter 18 (Spinal Cord Stimulation to Enable Leg Motor Control and Walking in People with Spinal Cord Injury) was previously published non-open access. It has now been changed to open access under a CC BY 4.0 license and the copyright holder updated to ‘The Author(s)’. The book has also been updated with this change.

The updated version of this chapter can be found at https://doi.org/10.1007/978-3-031-08995-4_18 © The Author(s) 2023 D. J. Reinkensmeyer et al. (eds.), Neurorehabilitation Technology, https://doi.org/10.1007/978-3-031-08995-4_35

C1

Index

A Aachener Aphasie Test (AAT), 579 Aachener Spontansprachanalyse, 580 AAT. See Aachener Aphasie Test (AAT) ABI. See Acquired brain injury (ABI) ABLE exoskeleton, 275 AbleX rehabilitation system, 534 Accelerometers, 71, 331, 473 Acquired brain injury (ABI), 549, 550, 687 ACT3D arm robot, 92, 326, 626 ActiGait® stimulator, 408 Actigraphy, 331 Action potentials, 402 Action Research Arm Test (ARAT), 72, 473, 535, 606–607, 638, 643, 653 Active leg exoskeleton (ALEX), 666 Active Response Gravity Offload System (ARGOS), 748 Active video games, 445 Activities of daily living (ADL), 44, 58, 154–155, 330, 331 actuated joints, functional training, 626 ARMin robot system, 627, 631–632 upper limb training, 76 Activity, 568 Activity-driven recovery, 347 AD. See Autonomic dysreflexia (AD) Adaptation, 224–225 Adaptive haptic training paradigms, 253–254 ADL. See Activities of daily living (ADL) Adverse events (AE), of robotic locomotor training, 153 Affective computing, 213 Affective pathway, 196 AFO. See Ankle-foot orthosis (AFO) After-effects, 224 AHA. See American Heart Association (AHA) AI. See Artificial intelligence (AI) Air Hockey game, 643 ALEX. See Active leg exoskeleton (ALEX) ALS. See Amyotrophic lateral sclerosis (ALS) Amadeo, 308, 310 Amadeo System, 121, 140 American Heart Association (AHA), 598, 712 American Spinal Injury Association (ASIA), 63, 728 Impairment Scale (AIS), 63, 379 Amyotrophic lateral sclerosis (ALS), 512

ANDAGO device, 276, 300, 301, 305, 749 Animal studies different reflex circuits, recruitment of, 377 locomotor circuits, recruitment of, 377 spinal circuitry in, 148–149 Anklebot description, 703 design features and measurement, 703 drop foot, 703 pediatric version, 703, 704 torque profile, 703 Ankle-foot orthosis (AFO), 708 Anterior circulation infarcts (PACI), 154 Anthropometrics, 278 robotic training, setup time, 278–279 task specificity, 279 Aphasia, 567, 577 Aphasia Bank Project, 580 ApoE. See ApolipoproteinE (ApoE) ApolipoproteinE (ApoE), 358–359 Apps, mobile health, 553–554 ARAT. See Action Research Arm Test (ARAT) ARGOS. See Active Response Gravity Offload System (ARGOS) Arm and hand movements, assessing clinical scores, estimating, 471 Action Research Arm Test (ARAT), 473 Box and Block Test (BBT), 473 estimating more than one clinical scale, 473–474 Fugl-Meyer Assessment, Upper-Extremity subscale (FMA-UE), 472–473 Functional Ability Scale (FAS), 473 movement kinematics, estimating, 470–471 upper limb training, wearable sensors to facilitate, 474–475 Arm Coordination Training Robot, 625 Armeo Booom system, 531, 627, 654 Armeo Power robot, 74, 529, 636, 637 Armeo robot, 140 Armeo Senso system, 309, 311, 474–475 Armeo Spring robot, 66, 273, 310, 530–531, 625, 637, 652–653, 762 ARM-Guide, 650 ARMin robot system, 67, 275, 326 actions and mechanism

© The Editor(s) (if applicable) and The Author(s), under exclusive license to Springer Nature Switzerland AG D. J. Reinkensmeyer et al. (eds.), Neurorehabilitation Technology, https://doi.org/10.1007/978-3-031-08995-4

765

766 actuation principle, 626–627 mechanical structure, 624–625 number and type of actuated joints, 625–626 robot-supported arm therapy system, 624 actuated joints, 625–626 applications, 623–624 Armeo Power®, commercial version, 636, 637 Armeo products, 637 ARMin I robot, 627 ARMin II robot, 627 ARMin III robot, 627, 628, 636 ChARMin robot, 633–636 end-effector-based robots, 624–625 evaluation clinical pilot studies, 638–639 clinical trials, 639–641 with healthy subjects, 636–638 perspectives for future clinical testing, 641 with stroke patients, 638 exoskeleton robots, 624–625 goals of, 638 measurement functionality of, 632–633 motorized robots, 626–627 nonmotorized robots, 626–627 ongoing testing, 641–644 innovative approach, 642–643 multiplayer games, 643 novel neuro-animation experience, 643 technical development, 641–642 ROM measure, 632 technical data for, 637 technical evaluation, 627 therapy modes ADL training, 631–632 game therapy, 629–631 passive and active mobilization, 629 upper-limb rehabilitation, 644 ARMON orthosis, 272–273 Arm rehabilitation, 214–215 Artificial intelligence (AI), 253 Ashworth Scale/Test, 324, 672 ASIA. See American Spinal Injury Association (ASIA) ASR. See Automated Speech Recognition (ASR) “Assist-as-needed” (AAN) control scheme, 276, 630 Assisting Hand Assessment, 293 Assistive robots adaptive tools, 124 BAXTER robot, 121, 122 BMI-controlled robotic limbs, 127, 128 Care-O-bot 4, 121, 122 Cyber 310 robotic arm, 120 DeVAR, 120 HERB, 121 iARM, 121 JAC02 robotic arm, 121 KARES II system, 120 MoVAR device, 121 Raptor Wheelchair Robot System, 120 from rehabilitation therapy to, 312

Index SEM Glove, 126 Associative learning, 4, 5 Asynchronous therapy, 565, 582 Atalante, 685, 689 ATLAS Pediatric Exoskeleton, 299, 304–305 Attention pathway, 196 Auditory stimuli, 439 Augmentation error augmentation combined effects of forms of, 233, 234 for enhanced training, 231–232 functional, 238–241 guidance vs., 228–231 statistical approaches to personalize, 236–238 of feedback to leverage, 227–228 movement, 228 Augmented (external) feedback, 195, 308–309, 311, 433 Augmented reality, 305–306 AutoAmbulator® device, 666, 706, 721, 738 Automated adaptation scheme, 280 Automated Speech Recognition (ASR), 581 Autonomic dysreflexia (AD), 266 Autonomic nervous system, 266 Autonomy, 194

B BAC Glove. See Bidirectionally Actuated Cable (BAC) Glove Balance and mobility of stroke survivors, 476 clinical scores, estimating 10-m walk test (10 MWT), 476 Timed Up-and-Go (TUG) test, 476 Trunk Impairment Scale (TIS), 476–477 to facilitate gait training in the clinic, 477 Barthel Index, 158, 325, 638 BATRAC device, 532, 654 BAXTER robot, 121, 122 BBS. See Berg Balance Scale (BBS) BBT. See Box and Block test (BBT) BCIs. See Brain-computer interfaces (BCIs) BDNF. See Brain derived neurotrophic factor (BDNF) Belgrade Grasping-Reaching System, 411, 412 Berg Balance Scale (BBS), 449, 704, 724, 725, 728, 729, 736 Berkeley Bionics. See EksoNR™ Berkeley, George, 226 Berkley Lower Extremity Exoskeleton (BLEEX), 684 Bernstein’s ‘repetition without repetition’ principle, 304 Bidirectional brain–computer interface (BD-BCI), 511 Bidirectionally Actuated Cable (BAC) Glove, 125, 127 Bilateral hand movements, by neural coupling, 136–138 Bilateral vs. unilateral motor learning, 610–611 Bi-Manu-Track device, 140, 625, 640, 650 Bi-Many-Track system, 415 “Bio-cooperative” control system, 214–215, 217–218 Biological models, 349–352 BiOMOTUM Spark, 685, 693–694 Bionic Glove, 76, 411, 412, 413

Index BLEEX. See Berkley Lower Extremity Exoskeleton (BLEEX) Blocked trials, 235 BMI. See Brain-machine interface (BMI) Body structure and body functions, 568 Body weight support (BWS) devices, 156, 160–161, 275–276, 670, 745, 747, 749 characteristics harness and attachment, 752–753 lateral forces, 752 longitudinal forces, 752 transparency, 751 vertical support forces, 751–752 clinical rationale for training origins and evolution, 745–746 target groups, 746–747 future directions, 753–754 gait training, 745, 753 overground gait training, outcomes of, 753 overground training devices non-robotic devices, 750–751 robotic devices, 747–749 Body-weight-supported treadmill training (BWSTT), 45, 146, 148, 305, 665–666, 695, 703, 719–722, 731, 734, 736 Body-worn sensors, 65–66 BONES robot, 103 Boost, 657, 658–659 Box and Block test (BBT), 43, 73, 448, 473, 484 Brain-computer interfaces (BCIs), 277 BCI controlled FET definition, 415–416 potential mechanisms, 418–421 SCI/stroke, for upper and lower limb function restoration, 417–418 technology, 416 stroke motor rehabilitation gait impairment, 509–510 neuroprosthetics, 512–514 operation, 510 physiotherapy. See Physiotherapy, BCI post-stroke motor impairment, 509 working process, 511 and virtual reality, 440–441 Brain derived neurotrophic factor (BDNF), 358 Brain injury, movement impairments treatment. See Impairment-based neurorehabilitation devices Brain lesions cerebellar lesions, 10–11 system plasticity, 6 Brain-machine interfaces (BMI), 127, 128 Brain plasticity, VR, 442–443 Brain-spine interface, 382 Brain structure and function, 359–360 Burke rehabilitation hospital inpatient studies, 603 BWS devices. See Body weight support (BWS) devices

767 BWSTT. See Body-weight-supported treadmill training (BWSTT)

C Cable Actuated Dexterous (CADEX) glove, 123 Cable-driven locomotor training system (3DCaLT) developments and ongoing testing cable-driven robot system, 738 children with CP, 730–734 human, incomplete SCI, 726–730 individuals post stroke, 724–726 walking function in children with CP, 737–738 walking function in humans with SCI, 736–737 walking function in individuals post-stroke, 734–735 neural plasticity, 719 pathophysiology cerebral palsy, children with, 718–719 SCI, 718 stroke, 718 robotic BWSTT, children with CP, 722–724 robotic gait training humans with SCI, 722 individuals post-stroke, 721 therapeutic action/mechanisms and efficacy children with CP, 720–721 humans with SCI, 720 individuals post-stroke, 719–720 Caden-7, 625 Cadense shoe, 711–712 CADEX glove. See Cable Actuated Dexterous (CADEX) glove Camera-based systems, 574 Canadian Occupational Performance Measure (COPM), 72 Capabilities of Upper Extremity Instrument (CUE-T), 69, 77 Capacity for activity, 469 Capacity for recovery of function, 41, 47, 49 Cardiovascular training, challenges of, 209 Care-O-bot 4, 121, 122 Carry-over effect, of FES, 406–407 Case Western Reserve University (CWRU)/VA neuroprosthesis, 410 Catch mechanism, 197 Catch trials, 233 Ceiling effect. See Diminishing returns Ceiling-mounted BWS systems, 750 Center of pressure (COP), 730 Central nervous system (CNS), 264, 280 Central pattern generators (CPGs), 7, 146–148, 151, 371, 376, 377, 378, 379 Cerebellar lesions, 10–11 Cerebral palsy (CP), 290, 296, 297, 302, 303, 305, 310, 633, 747 See also Upper extremity robotic therapy

768 children with, 718–719 locomotor training, 730–734 neuroplasticity, 719 pathophysiology, 718–719 robotic BWSTT, 722–724 task-oriented practice, 720–721 impaired balance, 702 Lokomat technology, 673 Cervical spinal cord injury measures of activity after, 66–73 measures of body functions and structures after, 63–66 measures of participation after, 74 Challenge point framework, 251, 253 ChARMin robot, 307, 308, 310, 633–636 CHEP. See Contact heat evoked potential (CHEP) Children, rehabilitation therapy in, 289, 290–291, 311, 313 applying, 293 congenital vs. acquired neurological lesions, 294–296 consultation and test training, 296 intensity of training, 296 patient selection, 294 assistive technology and, 312 cerebral palsy, 290 concerns about using, 311–312 implementation of technologies assessments, 293 environment, 293 scheduling robotic therapies, 292–293 technologies, 291–292 therapists, 292 technology supported lower extremity, 296–297 clinical evidence, 302–306 pediatric lower extremity systems, 297–302 technology supported upper extremity, 306 clinical evidence, 310–311 pediatric lower extremity systems, 306–309 Chronic pain management, spinal cord stimulation for, 370 CIMT. See Constraint-induced movement therapy (CIMT) Circulation infarcts (POCI), 154 CLAN. See Computerized language analysis (CLAN) Cleveland FES Center, 405 Clinical assessment, 323 definition, 322 drawbacks of, 325 observed-based assessments, 324–325 time-based assessments, 324, 325 upper/lower limb function, 324 Clinical Outcomes Assessment Compendium (COAC), 61 Clinical therapy program, implementation of robots into, 268 CNS. See Central nervous system (CNS) COAC. See Clinical Outcomes Assessment Compendium (COAC) Cognition, 267

Index Cognitive assessment, mobile technology for, 555 Cognitive rehabilitation acquisition phase, 552 adaptation phase, 552 application phase, 552 mobile technology for, 550–552, 553–554 treatments, 555–556 Combinatory robotic training, 276–277 Communication, 577 speech and language functions observation-based clinical assessment, 578–580 technology-based clinical assessments, 580–581 tele-assessments, 581 teletherapy asynchronous therapy, 582 synchronous therapy, 582 Compensatory approaches/strategy, 43, 550 Compensatory movement control, 44, 48 Competent model of human walking, 702–711 Anklebot, 703–704 ankle mechanical impedance, 706 MIT-Skywalker, 706–708 restoration of walking, 708–709 Robotic Exosuit Augmented Locomotion (REAL), 709–711 seated position training, 704–706 Soft Exosuit, 708–709 soft robotic exosuit technology, 709 Compex motion-based neuroprosthesis, 76 Compex motion multichannel neuroprosthesis, 410 Computational neurorehabilitation models, 334–335 as dynamical system models, 346–347 essential characteristics of, 346 multiple time scales in, 347 of sensory cortex, 350 theoretical learning rules in, 347 errors, learning from, 348 feedback, learning without, 348 rewards, learning from, 348–349 types, 349 qualitative biological models, 349–352 quantitative predictive models, 352–353 Computer gaming, 535, 536, 629–631 See also Upper-extremity movement training Computerized language analysis (CLAN), 580 Conditional task difficulty. See Functional task difficulty Congenital vs. acquired neurological lesions, 294–296 Constraint-induced movement therapy (CIMT), 46, 139–140, 453, 526, 528, 529, 725, 735 Contact heat evoked potential (CHEP), 63, 64 Control algorithm, for robotic training devices, 276 Control and feedback interfaces, 280 Conventional gaming consoles, 283 Conventional therapy, 414, 639 Cooperative hand movements contralateral reflex responses to ipsilateral arm nerve stimulation, 139 control of, 140 neural coordination of, 135–141

Index neural coupling bilateral hand movements by, 136–138 in post-stroke subjects, 138–139 reflex response pattern during, 137 robotic assisted therapy, 139–141 COP. See Center of pressure (COP) COPM. See Canadian Occupational Performance Measure (COPM) Corticospinal tract (CST), 43, 359 COVID-19 pandemic, 217, 537, 565 CPGs. See Central pattern generators (CPGs) Crank-and-rocker mechanism, 721, 723 CST. See Corticospinal tract (CST) Customized therapy, 236, 237 Cyber 310 robotic arm, 120 Cybersickness, 432 Cyclic body unloading and loading, role of, 149

D Daily life mobility, robotic devices for, 151 Dampace device, 625, 626 DARPA. See Defense Advanced Research Projects Agency (DARPA) Defense Advanced Research Projects Agency (DARPA), 684 DEMMI. See De Morton Mobility Index (DEMMI) De Morton Mobility Index (DEMMI), 158 Depression, 567 long-term, 4, 5, 8 Desktop Vocational Assistant Robot (DeVAR), 120 DeVAR. See Desktop Vocational Assistant Robot (DeVAR) Diego device, 308, 310 Different reflex circuits, recruitment of, 377 Digital health metrics, 322, 323–324, 337 computational models, for neurorehabilitation, 334–335 influencing therapy decisions with, 335 interpretation of, 333 robotic and sensor-based technologies, 335, 336–337 robotic assessments and, 325–331 advanced digital health metrics, 327–330 non-robotic technology-based assessments, 330–331 platforms, 331–332 raw sensor data, 325–327 selection and validation of, 332–333 Digital mobility outcomes (DMOs), 576 Digital-palmar prehensile strength, 65 Diminishing returns, 172 Directional error vectors, 348 Disconnection syndrome, 27 Discrete movement, 702 DLR Light-Weight Robot III arm, 127 DMOs. See Digital mobility outcomes (DMOs) Dolphin scenario, 643 Dopamine, 7, 9, 250, 386 Dose–response relationship, 42, 47, 303

769 Drop foot, 407–410, 703 Dynamical system models, 346–347 Dynamic dominance hypothesis, 27–28 Dynamic tilt table, 302

E EA. See Error augmentation (EA) ECoG. See Electrocorticogram (ECoG) Ecologically valid treatment, 550 Ecologically valid walking and balance tasks, 198 EDC. See Extensor digitorum communis (EDC) Edisonian, 759 EEG. See Electroencephalogram/electroencephalography (EEG) Egocentric video, 67, 71 EHPA project. See Exoskeleton for Human Performance Augmentation (EHPA) project Eight-order model, 233 EksoNR™, 157, 160, 274, 667, 685, 687–688 Electrocorticogram (ECoG), 511, 512, 513 Electroencephalogram/electroencephalography (EEG), 127, 359, 442, 510–512, 514–516, 517 Electromyography (EMG), 24, 25, 45, 536, 658, 726, 727 impairment-based neurorehabilitation devices, 91 upper limb rehabilitation, 64–65 VAEDA Glove, 125 Electrophysiology CHEP, 64 EMG, 64–65 MEP, 64 NCS, 64 SSEP, 64 EMG. See Electromyography (EMG) End-effector-based systems, 273–275, 291, 299, 624–625 Enhanced OPTIMAL Theory, 196 Enhanced upper limb therapy (EULT), 75, 605, 762 Epidural electrode, 370, 373 ERD. See Event-related desynchronization (ERD) Erigo exoskeleton, 298, 302 Error amplification, 255 Error augmentation (EA) algorithms, 48 application for stroke rehabilitation, 239 combined effects of forms of, 233, 234 concept, 271 for enhanced training, 231–232 functional, 238–241 guidance vs., 228–231 statistical approaches to personalize, 236–238 Error fields construction, 237–238 Error-free learning, 226 Error perception, 232 Errors, learning from, 348 Error space, 237 ERS. See Event-related synchronization (ERS) E-skin sensors, 491–492 E-textiles, 489–491

770 ETHZ-Paracare device, 76, 411, 413 EULT. See Enhanced upper limb therapy (EULT) European-Union-funded MIMICS project psychophysiological integration in arm rehabilitation, 214–215 in leg rehabilitation, 215–216 Event-related desynchronization (ERD), 416, 511, 515 Event-related synchronization (ERS), 511 Evidence of effectiveness, 566, 567 Excitatory spinal circuits, recruitment of, 376 EXCITE clinical trial, 656 Exercise robotic devices, 528–529 therapy, 526, 527 Exergames, 291, 303, 308–309 Exo-Glove Poly II, 125, 127 Exoskeleton for Human Performance Augmentation (EHPA) project, 684 Exoskeleton robots, 624–625, 326–327, 274–275, 291 See also specific entries Atalante, 685, 689 available devices, 684 BiOMOTUM Spark, 685, 693–694 BWSTT, 695 Class II devices, 696 considerations for clinical use documentation, 695 general assessment, 694 limitations, 696 safety considerations, 695–696 training strategies, 695 EksoNR™, 685, 687–688 GEMS-H, 685, 690–691 HandSOME, 124 history of, 683–684 Honda Walking Assist Device, 685, 689–690 Indego, 685, 688–689 Keeogo, 685, 691–692 modular exoskeletons, 689–692 MyoSuit, 685, 693 regulatory status and future expectations, 696 ReWalk™, 685–687 ReWalk Restore, 685, 692–693 rigid lower-body exoskeletons, 684–689 soft exoskeletons, 692–694 Exosuits, 708 Expectancies, 194 EXTEND orthosis, 532, 533 Extensor digitorum communis (EDC), 114, 115 External focus of attention, 194 Extrinsic feedback, 433 Extrinsic motivation, 194 EyeToy® system, 451

F FAC. See Functional Ambulation Categories (FAC) Falls, safety from, 197 FAS. See Functional Ability Scale (FAS)

Index FCR. See Flexor carpi radialis (FCR) FDA. See Food and Drug Administration (FDA) FDP. See Flexor digitorum profundus (FDP) FDS. See Flexor digitorum superficialis (FDS) Fear of falling, 746, 750, 753 Feedback learning without, 348 parameters, 280 systems, 571–572 Feedforward control, 233 FET. See Functional electrical stimulation (FES) therapy (FET) Fictive locomotion, 147 FIM. See Functional Independence Measure (FIM) FINGER robot, 123, 655, 656 FitMI system, 534 Fixed guidance/disturbance, 249 Flaccid muscle tone, 42 Flexor carpi radialis (FCR), 612 Flexor digitorum profundus (FDP), 114, 116 Flexor digitorum superficialis (FDS), 114, 116 Flexor hypertonicity, 116, 129 FLOAT system. See Free Levitation for Overground Active Training (FLOAT) system Flow theory, 436 FMA. See Fugl-Meyer Assessment (FMA) fMRI. See Functional magnetic resonance imaging (fMRI) FoG. See Freezing of Gait (FoG) Food and Drug Administration (FDA), 61, 62, 408, 684, 696 FreeBal device, 530, 654 FreeD module, 299, 304, 668–669 Free exploration therapy, 228, 229 Freehand system, 76, 411, 412 Free Levitation for Overground Active Training (FLOAT) system, 276, 300–301, 305, 748, 752 FreeStep SAS, 750 Freezing of Gait (FoG), 174 Fugl-Meyer Assessment (FMA), 91, 155, 158, 325, 352, 363, 418, 530, 603, 606, 633, 638, 639, 640, 653, 720 Fugl–Meyer Assessment for Upper Extremities (FMA-UE), 239–240, 472–473, 568, 643 Functional Ability Scale (FAS), 473 Functional Ambulation Categories (FAC), 152, 156, 158, 293 Functional electrical stimulation (FES) therapy (FET), 76–77, 440, 510, 516–517, 276, 277, 390. See also Brain-computer interfaces (BCIs) carry-over effect, 406–407 definition, 402, 406 neuroplasticity effect, 406–407 orthoses/robotic devices, 414–415 physiology, 402–404 potential mechanisms, 418–421 vs. robotic therapies, 415 SCI lower limb function restoration, 409–411

Index upper limb function restoration, 413–414 and spinal cord stimulation (SCS), 387 conceptual differences, 387–388 practical differences, 388 stroke lower limb function restoration, 407–409 upper limb function restoration, 411–413 technology, 404–406 upper limb training, 536 Functional error augmentation, 238–241 Functional Independence Measure (FIM), 23, 418, 720, 753 Functional magnetic resonance imaging (fMRI), 6, 10, 137–138, 359, 442–443 Functional Mobility Scale, 293 Functional Reaches Test, 449 Functional task difficulty, 251, 252

G Gait Enhancing and Motivating System (GEMS-H), 685, 690–691 GAITRite™ system, 330 Gait robots, 209 Gait Trainer (GT), 156, 273, 300, 667–668, 676, 721, 723 Gait training, wearable sensors to facilitate, 477 Game-based VR applications in rehabilitation. See VR exergames Game therapy ARMin robot system, 629–631 labyrinth game, 630 ping-pong game, 630 virtual reality scenarios, 630 for neurorehabilitation, 283 GCS. See Glasgow Coma Scale (GCS) Gene expression, neuroplasticity, 4–5 Generalization, 235, 250 Genetic markers, 358–359 Genomic analysis, 363 GENTLE/s robot, 625, 626 GEO system, 273, 274, 299, 668 German Bogenhausener Dysarthrieskalen (BoDyS), 578 GestureTek® IREX® video capture system, 449 GEXP therapy. See Graded exposure (GEXP) therapy Gillette Functional Assessment Questionnaire walking scale, 293 Glasgow Coma Scale (GCS), 158–159 GloreHa system, 123 GMFCS. See Gross Motor Function Classification System (GMFCS) GMFM. See Gross Motor Function Measure (GMFM) Goal Attainment Scale, 293 Graded and Redefined Assessment of Strength, Sensibility and Prehension (GRASSP), 68, 76, 77, 641 Graded difficulty, 198 Graded exposure (GEXP) therapy, 199, 200 Grail system, 300, 305 Graphical user interface (GUI), 629 Grasp and Release Test (GRT), 69

771 GRASSP. See Graded and Redefined Assessment of Strength, Sensibility and Prehension (GRASSP) Gravity-compensated shoulder-and-elbow robot, 599 Gravity-compensated shoulder-elbow-and-wrist exoskeletal robot, 599–601 Gravity non-compensated shoulder-and-elbow robot, 601 Gross Motor Function Classification System (GMFCS), 303, 731, 732 Gross Motor Function Measure (GMFM), 293, 302 GRT. See Grasp and Release Test (GRT) GT. See Gait Trainer (GT) GUI. See Graphical user interface (GUI) Guidance augmentation, 228–231

H Haken-Kelso-Bunz model, 610–611 HAL. See Hybrid assistive leg (HAL) Hand function restoration assistive devices adaptive tools, 124 BAXTER robot, 121, 122 BMI-controlled robotic limbs, 127, 128 Care-O-bot 4, 121, 122 Cyber 310 robotic arm, 120 DeVAR, 120 HERB, 121 iARM, 121 JAC02 robotic arm, 121 KARES II system, 120 MoVAR device, 121 Raptor Wheelchair Robot, 120 SEM Glove, 126 brain plasticity, 119 current state of technology, 128–129 electromyography (EMG), 127 neuromechanics, 114–115 non-wearable devices fixed base, 123–124 hand-end-effector coupling, 121–123 independent robots, 120–121 pathophysiology SCI, 118–119 stroke, 115–118 synaptogenesis, 119 therapeutic devices Amadeo system, 121 FINGER robot, 123 HandSOME, 124 PneuGlove, 124 VAEDA Glove, 125 X-Glove, 125 wearable hand robotics active extension assistance, 124–125, 126–127 active flexion assistance, 125–127 control, 127 Hand movements. See Arm and hand movements, assessing; Cooperative hand movements

772 Hand robot, 601–602 HandSOME orthosis. See Hand Spring Operated Movement Enhancer (HandSOME) orthosis Hand Spring Operated Movement Enhancer (HandSOME) orthosis, 124, 532, 533 Hand-worn sensors, 71 Handy 1 workstation, 120 “Haptically demonstrate”, 248 Haptic and tactile stimuli, 439–440 Haptic assistance, 248 Haptic cueing, 248 Haptic disturbance methods, 249 Haptic error augmentation, 253 Haptic feedback, 248, 254, 516 Haptic Kitchen game, 643 HapticMASTER robot, 91 Haptic noise, 249 Haptic nudges, 486 Haptic rendering, 254 Haptic training aim of, 250 effectiveness of, 248, 252 forces, 254–255 functional task difficulty, 251, 252 learning benefits of, 252 methods, 248–249 on motor (re) learning, 251–253 nominal task difficulty, 251–252 paradigms, 255 strategy, 255 terminal feedback, 252 visual feedback, 252 HapticWalker device, 668 Harness, 752–753 HAS. See Hybrid assistive systems (HAS) Head-mounted display (HMD), 577, 676 Health care clients, 202 crisis, 537 mobile technology and, 555, 558 professionals, 167, 202, 203 systems, 754 Health Technology Assessment (HTA) Program, 605 Heart rate control model-based, 209–211 using treadmill speed and visual stimuli, 211 Hebbian learning, 226 Hebbian plasticity, 348, 350, 511 Hebb’s rule, 419, 719 Hemiparetic stroke, 23, 91, 93, 103 Hemorrhage, ischemia vs., 359 HERB. See Home Exploring Robot Butler (HERB) HERO Glove, 125, 128 Hirob, 312 HMD. See Head-mounted display (HMD) Hodgkin-Huxley formalism, 373 Hoehn, Margaret, 172 Home-based robotic systems, 282 Home Exploring Robot Butler (HERB), 121

Index Homeoplastic processes, 348 Homeostatic plasticity, 348, 349, 350 Honda Walking Assist Device, 685, 689–690 HOS. See Hybrid orthotic systems (HOS) Hotspots, 381 HTA Program. See Health Technology Assessment (HTA) Program Human “in the loop,” 207–208 Human–machine interface, 279 automated adaptation scheme, 280 control and feedback interfaces, 280 conventional gaming consoles, 283 feedback parameters selection, 280 home-based robotic systems, 282 mechanical interfaces, 279 robotic assessment and therapy documentation, 281 robotic therapy at home, 281–282 Humans, spinal circuitry in, 148–149 Human walking, competent model of. See Competent model of human walking Hybrid assistive leg (HAL), 667 Hybrid-assistive limb systems, 274 Hybrid assistive systems (HAS), 410, 414 Hybrid orthotic systems (HOS), 414

I iARM robot, 121 ICAE. See Intensive conventional arm exercise (ICAE) ICARE study. See Interdisciplinary Comprehensive Arm Rehabilitation Evaluation (ICARE) study ICF model. See International Classification of Functioning, Disabilities, and Health (ICF) model ICMS. See Intracortical microstimulation (ICMS) ICT. See Information and communication technologies (ICT) IISART. See International Industry Society in Advanced Rehabilitation Technology (IISART) Immersive VR modalities, 432 Impairment-based neurorehabilitation devices abnormal synergies and weakness quantification, 90–96 ACT-3D robot, 91–92, 94 hand finger orthosis, 94 HapticMASTER robot, 91 muscle coactivation, 91 paretic limb movements, 91, 93 shoulder/elbow flexion/extension, 93 stereotypic movement, 90–91 wrist and fingers force, 93 device design, 102–104 independent joint control loss, 100–102 patient motivation, 106–107 practical implications, and translation, 106–107 rehabilitation robotic therapies, 90 rehabilitation specialist acceptance, 104–106 robotic devices, 97–98 scientifically underpinned concept, 100 SENSING technologies, 98–100

Index sensorimotor deficits, 90 spasticity quantification, 96–97 upper limb synergies, in hemiparetic stroke, 108 Impedance control, 35 latency response, 24 limb impedance, 25–26 reflexive resistance, 26 rehabilitation, implications, 27 selective compensation, 24 Impedance controller, gait orthosis technology architecture, 671 disadvantage, 671 Implantable pulse generator (IPG), 382 Implanted electrodes, 405 IMUs. See Inertial measurement units (IMUs) Indego, 274, 685, 688–689 Industry partners, 202 Inertial measurement units (IMUs), 65–66, 71, 330, 331, 332, 471, 693 Informational pathway, 196 Information and communication technologies (ICT), 563 Inhibitory spinal circuits, recruitment of, 376–377 Initial values, 172 Injuries. See also Cervical spinal cord injury; Impairment-based neurorehabilitation devices; Spinal cord injury (SCI); Traumatic brain injury (TBI) acquired brain injury (ABI), 549, 550, 687 cervical spinal cord injury, 63–74 SCI, hand function restoration, 118–119 InMotion® Arm/Hand, 121 InMotion2, 307–308 Innovative movement therapy, in children, 304 Innowalk/Innowalk Pro devices , 298, 302 Instrumented hand finger orthosis, 94 Instrumented TUG (iTUG), 476 Intelligent Robotic Arm, 625 Intensive conventional arm exercise (ICAE), 609–610 Interdisciplinary Comprehensive Arm Rehabilitation Evaluation (ICARE) study, 359 Internal models, 224 International Classification of Functioning, Disabilities, and Health (ICF) model, 60–61, 323, 527, 551, 568, 569, 575 International Classification of Functioning, Disability and Health, Child & Youth version (ICF-CY), 290 International Industry Society in Advanced Rehabilitation Technology (IISART), 267 International Spinal Cord Injury Society (ISCoS), 63 International Standards for the Neurological Classification of Spinal Cord Injury (ISNCSCI), 63 Inter-trial noise, 233 Intracortical microstimulation (ICMS), 513 Intrinsic feedback, 195, 433 Intrinsic motivation, 194 Intrinsic Motivation Inventory, 213, 655 IPG. See Implantable pulse generator (IPG) IREX® system, 451 Ischemia vs. hemorrhage, 359

773 Ischemic stroke, 154 ISCoS. See International Spinal Cord Injury Society (ISCoS) ISNCSCI. See International Standards for the Neurological Classification of Spinal Cord Injury (ISNCSCI) ISO 9241–210 standard, 279 iTUG. See Instrumented TUG (iTUG)

J J JAC02 robotic arm, 121 Jamar dynamometer, 65 Jebsen–Taylor Hand Function Test (JTHFT), 28, 30–31, 128 Jintronix software, 449 Joint position reproduction (JPR), 674 JPR. See Joint position reproduction (JPR) JTHFT. See Jebsen–Taylor Hand Function Test (JTHFT)

K KAIST. See Korea Advanced Institute of Science and Technology (KAIST) KARES II system, 120 Keeogo device, 685, 691–692 Key purchase decision-makers, 202 KINARM Exoskeleton Robot, 328, 529 KineAssist device, 749 Kinect™, 65, 330, 445, 448, 452–454, 533 KITE BCI-FET system, 417 Knowledge of performance, 195 of results, 195 Korea Advanced Institute of Science and Technology (KAIST), 120

L LabVIEW program, 723 LACI. See Lacunar anterior circulation infarcts (LACI) Lacunar anterior circulation infarcts (LACI), 154 Language function, assessment of observation-based clinical assessment, 578–580 technology-based clinical assessments speech and language analysis, 580–581 voice analysis, 580 LARA. See Lever-Assisted Rehabilitation for the Arm (LARA) LE. See Lower extremity (LE) LEAPS study, 666, 703, 712 Learning damaged brain/spinal cord. See Neuroplasticity from errors, 348 without feedback, 348 rates, 347 from rewards, 348–349 Leg rehabilitation, 215–216

774 Lever-Assisted Rehabilitation for the Arm (LARA), 657–658 LEXO system, 273, 625 Linear-time invariant (LTI) difference equations, 347 Linear track system, 532 LiteGait device, 750, 753 LOC. See Loss of consciousness (LOC) Locomotion, 376, 377, 379, 382 neurophysiology underlying, 147–148 robots, 274 Locomotor circuits, recruitment of, 377 Locomotor training cable-driven robot system, other advantages of, 738 children, CP, 730–734 walking function, 737–738 human with incomplete SCI, 726–730 walking function, 736–737 individuals post-stroke, 724–726 walking function, 734–735 robotic devices for, 150–151 Lokohelp system, 668 Lokolift system, 670 LokomatPro Sensation, 672 Lokomat systems, 157, 158, 160, 209–211, 215, 269, 274–276, 292, 298–299, 302–303, 304, 311, 326, 667, 721, 723 assessment tools mechanical stiffness, 672–673 range of motion, 673–674 voluntary force, 673 clinical measures, 672–674 clinical outcomes, 676 current version, 667 description, 666 feature of, 674 mechanical resistance measurement, 673 orthosis design, 668 in position control mode, 670 spinal cord-injured patient, 667 VR-enhanced, 675 Long-term depression (LTD), 4, 5, 8 Long-term potentiation (LTP), 4, 5, 8 LOPES. See Lower-extremity powered exoskeleton (LOPES) Loss of consciousness (LOC), 158–159 Lower extremity (LE), 573 motor function, assessment of clinical assessment, 573–575 tele-assessment, 575–576 rehabilitation in children, 296, 312 clinical evidence, 302–306 pediatric lower extremity systems, 297–302 rehabilitation robotics, 327 Anklebot, 703–704 MIT-Skywalker, 706–708 Soft Exosuit, 708–709 Variable-Friction Cadense shoes, 711–712 teletherapy, 576–577

Index Lower-extremity powered exoskeleton (LOPES), 274, 276, 327, 666 Lower limb (LL) function, 42, 43, 44, 47 Lower limb movement amount, providing feedback on, 487–488 Lower limb movements, monitoring, 481–482 Lower limb rehabilitation gait technologies, 297 LTD. See Long-term depression (LTD) LTP. See Long-term potentiation (LTP) LYRA system, 273

M Machine learning algorithms (MLA), 213, 472 Maestro hand exoskeleton, 123 Magnetometry, 71 MAL test. See Motor Activity Log (MAL) test Manual muscle testing (MMT), 63 Manumeter, 483, 484 MAS. See Modified Ashworth Scale (MAS); Motor Assessment Scale (MAS) Maximum voluntary contraction (MVC), 117 MCID. See Minimally clinically important difference (MCID) MCS. See Motor Capacity Scale (MCS) MDD. See Minimal detectable difference (MDD) MDR. See Medical Device Regulation (MDR) Mechanical impedances, 671, 702 Mechanical interfaces, 279 Mechanomyography (MMG), 472 Mechatronic system, 665 Medical Device Regulation (MDR), 277, 278 Medical Devices Directive 93/42/EEC (MDD), 278 Memory dynamics model, 235 MEPs. See Motor evoked potentials (MEPs) MGA-Exoskeleton, 625 MI. See Motor imagery (MI); Motricity Index (MI) MIME robot, 140, 611, 650 Minimal detectable difference (MDD), 63 Minimally clinically important difference (MCID), 63, 605, 758, 760 Mirror Image Motion Enabler, 625 Mirror neuron system (MNS), 443 MIT-Manus robot, 75, 140, 273, 308, 598, 599, 625, 650 MIT-Skywalker asymmetric speed programs, 706 balance training, 707 and Cadense shoe, 711 conventional gait physiotherapy, 706 discrete training mode, 707 kinematically based robot-assisted gait therapy, 706 removing the floor constraint, 711 rhythmic training mode, 706 speed-enhancing programs, 706 treadmill training, 706–708 vision distortion programs, 706 Mixed reality (MR), 577 MLA. See Machine learning algorithms (MLA) MMG. See Mechanomyography (MMG)

Index MMT. See Manual muscle testing (MMT) MNS. See Mirror neuron system (MNS) Mobile BWS frames, 749–751 Mobile exoskeleton, 157, 159 Mobile health (mHealth), 553–554 Mobile technology, 549 for cognitive assessment, 555 for cognitive rehabilitation, 550–552, 553–554 for cognitive rehabilitative treatments, 555–556 ethical considerations, 554–555 future directions, 556–559 Model-based heart rate control, 209–211 Modern video analysis techniques, 495–496 Modified Ashworth Scale (MAS), 281, 324, 327, 617, 640, 672 Modular exoskeletons, 689–692 Monosynaptic reflex, recruitment of, 375–376 MoreGait device, 282 MossRehab, 164 Motion data, 570 Motivation, 250 gaming elements challenge, 436 goal setting, 435 rewards, 435–436 sense of progress, 436 socialization, 436–437 pathway, 196 Motivational effect, 654–655 Motor Activity Log (MAL) test, 240, 448, 639 Motor adaptation, 224–225 Motor Assessment Scale (MAS), 152 Motor Capacity Scale (MCS), 68 Motor control principles, 35, 224 impedance control latency response, 24 limb impedance, 25–26 reflexive resistance, 26 rehabilitation, implications, 27 selective compensation, 24 motor lateralization dynamic dominance hypothesis, 27–28 left hemisphere damage, 28, 30 neural optimization process, 27 rehabilitation, implications, 31–32 right hemisphere damage, 28, 30 motor learning aftereffects, 32 forward model, 32 population-coding model, 33 rehabilitation, implications, 34 skill learning, 33 visuomotor adaptation and generalization, 32–33 optimal control feedback control, 23 invariant characteristics, 20–21, 22 movement patterns, 20, 21 optimization principles, 21–22 rehabilitation, implications, 23–24

775 Motor control, spinal cord stimulation for animal studies, lessons from different reflex circuits, recruitment of, 377 locomotor circuits, recruitment of, 377 electrically activated neural structures, 373–374 general principles, 373 post-synaptic activation of neural circuits, 374 excitatory spinal circuits, recruitment of, 376 inhibitory spinal circuits, recruitment of, 376–377 monosynaptic reflex, recruitment of, 375–376 Motor-driven cable apparatus, 723 Motor-enabling effect of SCS, 379 Motor evoked potentials (MEPs), 64, 359, 612 Motor function observation-based clinical assessment, 567–569, 573–574 technology-based clinical assessment, 569, 574–575 tele-assessment, 570–571, 575–576 Motor imagery (MI), 440–441 Motor impairment, 306 Motorized robots, 626–627 Motor lateralization, 35 dynamic dominance hypothesis, 27–28 left hemisphere damage, 28, 30 neural optimization process, 27 rehabilitation, implications, 31–32 right hemisphere damage, 28, 30 Motor learning, 4, 35, 235–236, 247–248, 251, 253, 256, 345, 347 aftereffects, 32 assessing, 250 forward model, 32 movement variability and, 304 performance metrics, 250 population-coding model, 33 principles, 269–271 adaptability, 434–435 dosing, 434 enriched environments, 432–433 intrinsic and extrinsic feedback, 433 motivation, 435 task specificity, 433–434 rate, 255 rehabilitation, implications, 34 research in, 248 skill learning, 33 task performance variability, 250 visuomotor adaptation and generalization, 32–33 Motor Power score (MRC), 603 Motor rehabilitation, VR systems, 437 Motor skills, 247 Motor Status Score (MSS), 603 Motor task, 251–252 Motor variability, 249, 255–256 Motricity Index (MI), 155, 158 MoVAR device, 121 Movement augmentation, 228 Movement kinematics, estimating, 470–471 Movement smoothness, 327–328

776 Movement variability, 253 MRC. See Motor Power score (MRC) MS. See Multiple sclerosis (MS) MSS. See Motor Status Score (MSS) Multi-directional over-ground devices, 305 Multimodal human–machine interaction in rehabilitation, 207–208 Multimodal Rehabilitation Program, 487 Multiple sclerosis (MS), 370, 747 robotic gait training evidence in, 171 persons with, 170–171 practical application of, 171, 172 Multiple time scales, 347 Musculoskeletal system and skin, 266–267 MusicGlove system, 475, 534, 762 MVC. See Maximum voluntary contraction (MVC) MyndMove® stimulator, 67, 76, 404, 405, 411, 412 MyoPro, 126 Myosuit device, 269, 270, 685, 693 Myro, 295, 309, 311 My Spoon robot, 120

N NASA. See National Aeronautics and Space Administration (NASA) National Aeronautics and Space Administration (NASA), 264 National Health Service (NHS), 605 National Institute of Disability and Rehabilitation Research (NIDRR) support, 650 National Institute of Health Stroke Scale (NIHSS) score, 334, 361 NaviGAITor, 748 Negative damping (negative viscosity), 228–229 Negative error, 233 Neofect smart glove, 534 Nerve conduction study (NCS), upper limb rehabilitation, 64 NESS H200, for hand paralysis, 411, 412 Ness Handmaster, 76 NESS L300 Go, for foot drop, 408 Neural circuits, post-synaptic activation of, 374 recruitment excitatory spinal circuits, 376 inhibitory spinal circuits, 376–377 monosynaptic reflex, 375–376 Neural coupling afferent pathway of, 137–138 bilateral hand movements by, 136–138 efferent pathway of, 138 fMRI recordings, 137–138 ipsi- /contralateral M1 cortical areas, 138 in post-stroke subjects, 138–139 reflex responses, 137 unilateral leg nerve stimulation, 137 Neural networks, 231

Index Neural process, VR, 444 Neurapraxia, 42 NeuroFenix gameball, 534 Neurological conditions, 265–266 Neurological disorders, 324, 525–527, 564, 719 Neurological gait impairments, 306 Neuromodulation, 370, 378, 386 therapies, 276–277 Neuromodulatory networks, 536 Neuromuscular electrical stimulation (NMES), 402 Neuronal plasticity, 527 Neurophysiology, 42, 48 Neuroplasticity adults, 719 learning in CNS, 3–4 mechanism cellular plasticity, 5 gene expression, 4–5 plasticity, spinal cord, 6–7 subcortical contributions, 7–9 system plasticity, brain, 6 multiple functional forms of, 226–227 and recovery, 224–226 rehabilitation therapy cerebellar lesions, 10–11 lesions of cortex, 9–10 spinal lesions, 11–13 training-induced, 224 Neuroprosthesis, 379, 388, 409 See also Functional electrical stimulation (FES) therapy (FET) Neurorehabilitation, 224–225, 235, 237, 247, 256, 322 See also Robots, in neurorehabilitation computational models for, 334–335 robotics to stimulate, 248 NHPT. See Nine Hole Peg Test (NHPT) NHS. See National Health Service (NHS) NIDRR support. See National Institute of Disability and Rehabilitation Research (NIDRR) support NIHSS score. See National Institute of Health Stroke Scale (NIHSS) score Nine Hole Peg Test (NHPT), 324, 328 Nintendo® Wii™ , 445, 451, 533 NJIT-RAVR system, 440 NJIT-TrackGlove system, 439 NMES. See Neuromuscular electrical stimulation (NMES) Nominal task difficulty, 251–252 Non-invasive (transcutaneous) SCS, 389–390 Nonlinear “loss function”, 232 Non-motor data, collecting, 496 Nonmotorized robots, 626–627 Non-robotic devices ceiling-mounted BWS systems, 750 mobile frames, 750–751 motivation, 750 Non-robotic technology-based assessments, 330–331 NxStep™ system, 750

Index O Observation-based clinical assessment, 573–574, 578–580 Occupational therapists (OT), 421, 565 OCSP. See Oxford Community Stroke Project (OCSP) Oculus Quest, 306 Odstock® Dropped Foot Stimulator (ODFS® Pace), 408 Oncology, 358 OpenPose algorithm, 496 Operative control, 214–215, 217–218 Optical motion capture, 65 Optimal control, 34 feedback control, 23 invariant characteristics, 20–21, 22 movement patterns, 20, 21 optimization principles, 21–22 rehabilitation, implications, 23–24 Optimal rehabilitation robotics strategy, 90 Optimizing Performance through Intrinsic Motivation and Attention for Learning (OPTIMAL Theory), 194–197, 250 OT. See Occupational therapists (OT) Oura Health, 497 Overcompensation, 233 Overground gait training, outcomes of, 753 Overground training BWS devices non-robotic devices ceiling-mounted systems, 750 mobile frames, 750–751 motivation, 750 overground walking BWS systems, 670 robotic devices, 747 ceiling-mounted systems, 747–749 mobile frames, 749 motivation, 747 Overload, 172 Oxford Community Stroke Project (OCSP), 154

P Pablo system, 534 PAM. See Pelvic assist manipulator (PAM) Parastep system, 409–410 Parker Hannifin exoskeleton. See Indego Parkinson’s disease (PD), 388 robotic gait training, 171–173 evidence in, 175–177 persons with, 173–175 practical application of, 177–178 Participation, 568 Participation and Quality of life (PAR-QoL) Tool-Kit, 74 Participation restrictions, 551 Passive and active mobilization, 629 Passive gravity support systems, 529–531 Passive range of motion (ROM), 278 Passive trolleys, 750

777 Passive wearable exoskeleton, 124 Path controllers, 249, 672 Patient selection criteria, 760 Payment agencies, 202 PCI. See Physiological cost index (PCI) PD. See Parkinson’s disease (PD) PD controller. See Position-Derivative (PD) controller Pediatric Armeo Spring, 308, 309 Pediatric Lokomat, 722 Pediatric lower extremity systems, 297–302 Pediatric neurorehabilitation, 290, 296, 297 Pedometer, 488 Pelvic assist manipulator (PAM), 667 Percutaneous electrodes, 405 Performance of activity, 469 feedback, 195 Performance-based adaptive haptic training, 249 Performance-based progressive robotic therapy, 615 Performance-degrading (haptic disturbance), 248, 249, 252–253 Performance-enhancing (haptic guidance), 248, 249, 252 Personalized training, challenges and opportunities for, 235–236 PEXO exoskeleton, 312 PHANToM haptic device, 328 PHANToM Omni device, 328 PhysioGait system, 750 Physiological cost index (PCI), 408 Physiological integration of humans and rehabilitation technologies examples of, 211–212 heart rate control using treadmill speed and visual stimuli, 211 implementing into daily clinical routine, 212 model-based heart rate control, 209–211 rationale of, 209 Physiotherapy, BCI clinical applications, 517–519 post-stroke motor recovery, hypothesized mechanisms of, 515 stroke rehabilitation and mechanisms, 514–517 PI controller. See Proportional-integral (PI) controller Ping-pong game, 630 Plasticity, 348, 349, 350, 526 See also Neuroplasticity Play™ games, 451 PneuGlove, 124 Pneumatic actuators, 125 Pneumatically operated gait orthosis (POGO), 667 POGO. See Pneumatically operated gait orthosis (POGO) Pose estimation algorithms, 570 Position-Derivative (PD) controller, 670 Potential for biological recovery, 42–44 Powered devices, 310 PRAAT program, 580 Praxis neuroprosthesis, 410

778 Precision measurement, 357, 363 Precision rehabilitation, 357, 470 initial attempts at, 358 behavioral approaches and first clinical precision rehabilitation tools, 360–362 brain structure and function, 359–360 genetic markers, 358–359 stroke etiology, ischemia vs. hemorrhage, 359 oncology, 358 Predictive models, 352–353 Predict Recovery Potential (PREP and PREP2) algorithm, 334, 359–360, 362 Prevention-focused approach, 192–193, 197, 199 Probability density function, 228 Progression, 172 Proportional and/or derivative controllers, 249 Proportional-integral (PI) controller, 381 Propositional recovery model, 334 Proprioception, 46, 48, 123, 324, 328, 360, 656, 760 Proprioceptive afferents, 372, 374, 375 Proprioceptive effect, 656 Proprioceptive feedback, 6, 12, 149, 277, 379, 390, 471, 516 Psychophysiological integration in arm rehabilitation in MIMICS project, 214–215 examples of, 216–217 implementing into daily clinical routine, 217–218 in leg rehabilitation in MIMICS project, 215–216 rationale of, 212–213 PUMA-260 robot, 120 Pyramidal tract, 42

Q Qualisys, 65 Qualitative “biological” models, 349–352 Quantitative “predictive” models, 352–353 QuickDASH. See Quick Disabilities of the Arm, Shoulder, and Hand (QuickDASH) Quick Disabilities of the Arm, Shoulder, and Hand (QuickDASH), 72

R Radar-like systems, 494, 495 Radio tags and radar-like technologies, 494–495 RAE. See Resonating Arm Exerciser (RAE) RAGT. See Robotic-assisted gait training (RAGT) RAHFT. See ReJoyce Arm and Hand Function Test (RAHFT) Ranchos Los Amigos Hospital Functional Activities (RLAH), 70 Randomized controlled trials (RCTs), 105, 163, 171, 175, 176, 448, 517, 527, 528, 534, 536, 554, 566, 572, 650, 654 Randomized training, 235 Range of motion (ROM), 65, 326 ARMin robot system, measurement, 632 robotic gait orthosis technology, 673–674

Index Rapael Smart Board, 532 Rapael smart glove, 309 Raptor Wheelchair Robot System, 120 RATULS trial. See Robot Assisted Training for the Upper Limb after Stroke (RATULS) trial Raw sensor data, robotic assessments and, 325–327 RCTs. See Randomized controlled trials (RCTs) REAL. See Robotic Exosuit Augmented Locomotion (REAL) REAPlan device, 308, 310, 633 Recovery, 346 of function, 49 Regulatory Fit Theory, 193, 199 Regulatory Focus Theory, 192–194 Rehabilitation, 385–386, 746, 747 See also Children, rehabilitation therapy in motor abilities during, 152 robotics, 265, 267, 270, 281 technology-aided automated recognition of human psychological states in, 213 patient psychological states in, 213 psychological state recognition in, 213 Rehabilitation Joystick for Computerized Exercise (ReJoyce) system, 67, 273, 534–536 Rehabilitation Technologies Division of Applied Resources Corp. (RTD-ARC), 120 Rehabilitation technology design, 253–256 See also Upper limb rehabilitation, SCI; Walking and balance training performance adaptive haptic training paradigms, 253–254 haptic training paradigms, 255 implications for, 48–49 motor variability, 255–256 neurophysiological basis of, 41–49 nonlinear relationships in neurorehabilitation recovery, 43 principles for, 42–47 task-relevant information, 254–255 Rehabilitation Treatment Specification System (RTSS), 557 RehabNet system, 441 ReHapticKnob, 123 Reha-Slide system, 532 Reinforcement learning (RL), 226, 348–349 ReJoyce Arm and Hand Function Test (RAHFT), 66 ReJoyce system. See Rehabilitation Joystick for Computerized Exercise (ReJoyce) system Reliability, 62 REM-AVC trial, 653, 654 ReoAmbulator device, 666 ‘Repetition without repetition’ principle, Bernstein’s, 304 ReSense inertial measurement units, 67 Residual neural circuits, 42 Resonating Arm Exerciser (RAE), 657 Responsiveness, 62–63 Restoration of walking robotic devices for daily life mobility, 151

Index for locomotor training, 150–151 Restorative approaches, 550, 551, 552 Restorative interventions, 550 Restorative therapies robots technical aspects for, 272 body weight support devices, 275–276 combinatory robotic training approaches, 276–277 control algorithm for robotic training device, 276 end-effector devices vs. exoskeletons, 273–275 individually tailored training, 278 training devices complexity, 272–273 user-centered design process, 277–278 Reversibility, 172 ReWalk™, 274, 667 clinical applications, 686 features, 685 mobility training, 686 ReWalk Personal, 685 ReWalk Rehabilitation, 685 ReWalk Restore, 685, 692–693 Rewards, learning from, 348–349 RGO system, 410 Rhythmic movements, 702 Rigid lower-body exoskeletons, 684–689 RL. See Reinforcement learning (RL) RLAH. See Ranchos Los Amigos Hospital Functional Activities (RLAH) Robot-assisted haptic training, 247 haptic methods, on motor (re) learning, 251–253 haptic training methods, 248–249 motor learning assessment, 250 rehabilitation technology design, 253–256 Robot Assisted Training for the Upper Limb after Stroke (RATULS) trial, 75, 605–608, 762 Robotic assessments, 322, 323, 335 See also Digital health metrics and digital health metrics on advanced digital health metrics, 327–330 based on raw sensor data, 325–327 non-robotic technology-based assessments, 330–331 exemplary, 329, 330 future directions, 331–335 robotic and sensor-based technologies, 335, 336–337 usability and influence of, 331–332 Robotic-assisted gait training (RAGT), 145–147, 298 locomotion, neurophysiology underlying, 147–148 motor abilities during rehabilitation, 152 multiple sclerosis evidence in, 171 persons with, 170–171 practical application of, 171, 172 Parkinson’s disease, 171–173 evidence in, 175–177 persons with, 173–175 practical application of, 177–178 reasons for, 152 robotic devices for daily life mobility, 151 for locomotor training, 150–151

779 specific effects of, 152 specific neurological conditions. See also specific conditions evidence for, 153 introduction related to, 153 practical application of, 153 walking-related, 153 spinal circuitry in animals and humans, 148 cyclic body unloading and loading, role of, 149 spastic muscle tone, 149–150 spinal cord injury evidence in, 168 persons with, 167–168 practical application of, 168–170 stroke, 154 evidence in, 156–157 persons with, 155–156 practical application of, 157–158 traumatic brain injury, 158–159, 162 evidence, 163–164 persons with, 162–163 practical application of, 164–167 unspecific effects of, 152–153 using end-effector devices, 305 using exoskeleton devices, 302–303 Robotic-assisted training, 46, 267 Robotic devices, 269, 283–284, 564, 747 ACT-3D robot, 91–92, 94 ceiling-mounted BWS systems, 747–749 exercise, upper limb training, 528–529 mobile frames, 749 motivation, 747 therapeutic approach, 264 Robotic Exosuit Augmented Locomotion (REAL), 709–711 Robotic gait orthosis technology Ashworth Test, 672 assessment tools lower limb proprioception, 674 mechanical stiffness, 672–673 range of motion, 673–674 voluntary force, 673 biofeedback, 674–676 goals of, 674 manual treadmill therapy comparison, 674 muscular efforts estimation, 674 stance and swing phase, 674 VR and computer game techniques, 675 body weight support system, 670 BWSTT, 665–666 clinical outcomes, 676 control strategies, 670–672 disadvantage, 671 feedback, 671 impedance controller, 671 patient-cooperative, 671 position control mode, 670 Gait Trainer, 667, 676 Lokolift system, 670

780 orthosis design drives, 669 mechanical aspects, 668–669 safety, 669–670 spasticity, 672–673 Robotic movement training, 757–758 Robotic orthosis, 157, 158, 160 Robotic/technology-based assessments, 323 Robotic therapy. See also Upper extremity robotic therapy devices, 267 goal of, 272 at home, 281–282 scheduling, 292–293 Robotic training principles, 268 motor learning principles, 269–271 training parameters, 269 Robots, in neurorehabilitation, 263 See also Assistive robots human–machine interface, 279 automated adaptation of training, 280 control and feedback interfaces, 280 conventional gaming consoles, 283 feedback parameters selection, 280 home-based robotic systems, 282 mechanical interfaces, 279 robotic assessment and therapy documentation, 281 robotic therapy at home, 281–282 patients requirements, 265 autonomic nervous system, 266 cognition, 267 musculoskeletal system and skin, 266–267 neurological condition, 265–266 restorative therapies, robots for anthropometrics, 278–279 body weight support devices, 275–276 combinatory robotic training approaches, 276–277 control algorithm for robotic training device, 276 end-effector devices vs. exoskeletons, 273–275 individually tailored training, 278 legal challenges, 277–278 training devices complexity, 272–273 user-centered design process, 277–278 robotic devices, 264 robotic training principles, 268–269 feedback and virtual reality, 271–272 motor learning principles, 269–271 training parameters, 269 technology readiness level, 264, 265 therapists’ requirements, 267 clinical therapy program, robots in, 268 therapists instruction, 268 user requirements, 264 Robot studios, 644 Robot-supported arm therapy system, 624 Rocking therapy, 657 ROM. See Range of motion (ROM) RT300 FES, 406

Index RTSS. See Rehabilitation Treatment Specification System (RTSS) RUPERT devices, 275 Russian current stimulation, 391, 392 Rutgers Arm II, 532 RYSEN system, 276, 300, 301, 305, 748, 750, 751

S SaeboFlex, 124, 533 SaeboGlove, 124 SafeGait, 748 Safety, from falls, 197 Samsung Gait Enhancing and Motivating System—Hip (GEMS), 685, 690–691 SARAH system. See Semi-Automated Rehabilitation at the Home (SARAH) system Scalar rewards, 348 SCARA. See Selective Compliance Assembly Robot Arm (SCARA) SCIM. See Spinal Cord Independence Measure (SCIM) SCRIPT orthosis, 532 SCs. See Sensorimotor contingencies (SCs); Spinal cord stimulation (SCS) SDSS. See Spatially sequentially distributed electrical stimulation (SDSS) Secondary somatosensory (S2) cortical areas, 137–138 Second-order model, 233 SeeMe system, 450 Selective Compliance Assembly Robot Arm (SCARA) , 599 Selective Control Assessment of the Lower Extremity Scale, 293 Self-paced treadmill with VR environment, 305 Self-training, 565 SEM Glove. See Soft Extra Muscle (SEM) Glove Semi-Automated Rehabilitation at the Home (SARAH) system, 496 SENSING technologies, for home-based quantification of arm motor deficits, 98–100 Sensorimotor contingencies (SCs), 430 Sensorimotor impairment, 322, 324, 333 Sensorimotor rhythms (SMRs), 511, 515 Sensorimotor therapy, 602, 616 Sensory cortex, computational model of, 350 Sensory-motor relationships, distortion of, 224–225 Serious games, 445, 533–536 Serotonin, 386 Severe acute respiratory syndrome coronavirus 2 (SARS-CoV-2), 565 Short-interval cortical inhibition (SICI), 612 SICI. See Short-interval cortical inhibition (SICI) 6-m walk test (6MWT), 152, 155, 156, 158, 167, 168, 170, 302, 487, 573, 576, 704, 705, 724–725, 729, 730–733, 737, 738 Slacking, 46 Sliding spring actuators, 125 SMA. See Supplementary motor area (SMA)

Index SMART Arm, upper limb training, 532 SMRs. See Sensorimotor rhythms (SMRs) Soft exoskeletons, 692–694 Soft Exosuit, 708–709 Soft Extra Muscle (SEM) Glove, 126 Somatosensory evoked potentials (SSEP), upper limb rehabilitation, 64 Sony® Eye-Toy® system, 445 Spasticity, 672–673 quantification, 96–97 robot-assisted assessments of, 327 in walking impairment, 324 Spastic muscle tone, 43, 44, 45, 149–150 Spatially sequentially distributed electrical stimulation (SDSS), 405 Spatiotemporal SCS, 380, 383–385, 388 controlled by brain signals, 382 controlled by residual kinematics, 381–382 Specificity, 172 Speech and language analysis, 580–581 Speech and language functions observation-based clinical assessment, 578–580 technology-based clinical assessments speech and language analysis, 580–581 voice analysis, 580 Spinal circuitry in animals and humans, 148 cyclic body unloading and loading, role of, 149 spastic muscle tone, 149–150 Spinal Cord Independence Measure (SCIM), 67, 411 Spinal cord injury (SCI), 264, 265, 266, 376, 388, 718, 747 See also Upper limb rehabilitation, SCI age at incidence, 118 ARMin IV robot for rehabilitation of, 634 BCI-FET for upper and lower limb function restoration, 417–418 cervical SCI measures of activity after, 66–73 measures of body functions and structures after, 63–66 measures of participation after, 74 exoskeleton EksoNR™, 687–688 Indego, 688 ReWalk™, 685–686 falls, 118 FES therapy lower limb function restoration, 409–411 upper limb function restoration, 413–414 functional training, in persons, 11–12 hand function restoration, 118–119 limitations of locomotor rehabilitation facilitated by SCS, 385–386 plasticity spinal reflex, 7 task-specific, 7 prerequisite, training, 12–13 recovery of sensorimotor functions after stroke and, 41–49

781 rehabilitation, 13 robotic gait training evidence in, 168 persons with, 167–168 practical application of, 168–170 spatiotemporal SCS for neuroprosthetics and neurorehabilitation in animal models of, 380, 383–385 controlled by brain signals, 382 controlled by residual kinematics, 381–382 spinal neuronal circuits, plasticity, 11 studies of SCS for improving motor and autonomic functions after, 386–387 tonic SCS for recovery of voluntary motor control, 378–380, 383 voluntary movement after, 372 Spinal Cord Injury Center of the University Hospital Balgrist, 666 Spinal cord stimulation (SCS), 369 for chronic pain management, 370 and Functional Electrical Stimulation (FES) , 387–388 for improving motor and autonomic functions, 370–371, 386–387 for leg motor control following motor disorders, 371–373 limitations of locomotor rehabilitation facilitated by, 385–386 for motor control different reflex circuits, recruitment of, 377 electrically activated neural structures, 373–374 general principles, 373 locomotor circuits, recruitment of, 377 post-synaptic activation of neural circuits, 374–377 motor-enabling effect of, 379 as promising neuroprosthetic technology and neurorehabilitation therapy after SCI, 388 spatiotemporal SCS, 380–385 tonic SCS, 378–380, 383 transcutaneous SCS, 389–394 Spontaneous recovery, 347 SpringWear system, 531 State-space representation, 346 Stationary sensors, 330 Statistical error augmentation, 236–238 Stereo cameras, 570 Stimulette edition5, 402 STIMuSTEP®, 408 Strengthening effect, 655–656 Stroke, 266, 550, 567, 747 See also Balance and mobility of stroke survivors; Brain-computer interfaces (BCIs); Motor control principles; Wearable sensors for stroke rehabilitation ARMin IV robot for rehabilitation of, 634 BCI-FET for upper and lower limb function restoration, 417–418 deficits, 19 etiology, 359 FES therapy

782 lower limb function restoration, 407–409 upper limb function restoration, 411–413 pathophysiology, 718 recovery of sensorimotor functions after, 41–49 robotic gait training, 154 evidence in, 156–157 persons with, 155–156 practical application of, 157–158, 159, 160–161 Stroke motor rehabilitation BCI system neuroprosthetics, 512–514 physiotherapy, 514–519 working process, 511 gait impairment, 509–510 post-stroke motor impairment, 509 Subcutaneous electrodes, 405 Supervised learning, 226, 348 Supplementary motor area (SMA), 136 Swedish Help Arm, 272 Synchronous therapy, 565, 582 System identification, 235, 238

T Tabletop therapy systems, 531–532 TACI. See Total anterior circulation infarcts (TACI) Tailwind device, 532, 654 Task-oriented neurorehabilitative training, 272 Task-Oriented Physiotherapy (TOP), 303 Task specificity, 433–434 TBI. See Traumatic brain injury (TBI) tDCS. See Transcranial direct current stimulation (tDCS) Technology-based clinical assessments, 574–575 speech and language analysis, 580–581 voice analysis, 580 Technology readiness level (TRL), 264, 265 Teleassessment, 567, 581 Telemonitoring, 567 Telerehabilitation technology, 537, 563 acceptance of, 565–566 benefits, 564 communication, 577 speech and language functions, assessment of, 578–581 teleassessments, 581 teletherapy, 582 effectiveness of, 566–567 interaction from distance, 564–565 lower extremity, 573 motor function, assessment of, 573–575 teletherapy, 576–577 stroke, 567 synchronous and asynchronous therapy, 565 upper extremity, 567 motor functioning, assessment of, 567–571 teletherapy, 571–573 Teletherapy, 571, 576–577 asynchronous therapy, 582

Index feedback systems, 571–572 robotic systems, 572–573 synchronous therapy, 582 VR systems, 572 10-m walk test (10 MWT), 152, 156, 170, 171, 175, 177, 302, 324, 451, 476, 573, 704, 711 Tenodesis grasp, 44 TENS. See Transcutaneous electrical nerve stimulation (TENS) TES. See Transcranial electrical stimulation (TES) Tetraplegia, 118–119, 121, 126, 128 Tetraplegia Hand Activity Questionnaire (THAQ), 68 THAQ. See Tetraplegia Hand Activity Questionnaire (THAQ) Theory of Work Gamification, 195–196 Therapeutic and functional electrical stimulation, 536 Therapists rehabilitation technologies, 292–293 robotic therapy devices, 267–268 Therapy-Wilmington Robotic Exoskeleton (T-WREX) Armeo Spring®, 75–76, 530, 626, 636, 652–654 for arm training, 649, 651 conventional therapy with, 74 rehabilitation therapy with, 76 upper-extremity movement training, 650–652 THERA-Trainer LYRA, 299, 300, 668 Thorsen’s Functional Test, 69 3DCaLT. See Cable-driven locomotor training system (3DCaLT) 3D Guidance Trak-STAR system, 707 3D systems, 749 Threshold Hypothesis, 43, 483, 484 Threshold to detection of passive motion (TTDPM), 674 Time constant of error decay, 231, 232 Timed Up-and-Go (TUG) test, 476 TIS. See Trunk Impairment Scale (TIS) TMS. See Transcranial magnetic stimulation (TMS) Tone compensating orthoses, 532–533 TOP. See Task-Oriented Physiotherapy (TOP) Toronto Rehab Institute—Hand Function Test, 70 Total anterior circulation infarcts (TACI), 154 T-/Pneu-WREX device, 276 Training schedules, 235 Transcranial direct current stimulation (tDCS), 277, 611–612 Transcranial electrical stimulation (TES), 64 Transcranial magnetic stimulation (TMS), 5, 64, 77, 612 Transcutaneous electrical nerve stimulation (TENS), 390 Transcutaneous spinal cord stimulation, 390–391, 393–394 Transparency, 280, 751 Traumatic brain injury (TBI), 266, 550, 747 robotic gait training, 158–159, 162 evidence, 163–164 persons with, 162–163 practical application of, 164–167 Treadmill -based methods, 199–201

Index challenges of, 209 speed and visual stimuli, heart rate control using, 211 system, 300 training, 160–161 Anklebot, 703–704 MIT-Skywalker, 706–708 task-specific motor practice, 719 Trexo system, 299 TRL. See Technology readiness level (TRL) Trunk Control Test, 155, 158 Trunk Impairment Scale (TIS), 476–477 TTDPM. See Threshold to detection of passive motion (TTDPM) TUG test. See Timed Up-and-Go (TUG) test

U UC. See Usual care group (UC) UEFMA. See Upper Extremity Fugl-Meyer Assessment (UEFMA) UEMS. See Upper Extremity Motor Score (UEMS) Ultra-wideband (UWB) radio systems, 494 Unsupervised learning, 226, 348 Upper extremity, 567 motor functioning clinical assessment, 567–569 tele-assessment, 570–571 rehabilitation, 327–328 in children, 306–312 teletherapy, 571 feedback systems, 571–572 robotic systems, 572–573 VR systems, 572 Upper Extremity Fugl-Meyer Assessment (UEFMA), 359, 415, 447, 535, 655 Upper Extremity Motor Score (UEMS), 640–641 Upper-extremity movement training with mechanically assistive devices, 649–659 Boost, 658–659 case study, 649–651 clinical evidence, 651–654 clinical validation with ArmeoSpring, 652–654 democratizing, 656–659 Lever-Assisted Rehabilitation for the Arm (LARA), 657–658 motivational effect, 654–655 motor recovery, 654–656 other approaches, 654 proprioceptive effect, 656 Resonating Arm Exerciser (RAE), 657 robotic rehabilitation, 650 strengthening effect, 655–656 T-WREX, 651–652 Upper extremity robotic therapy American Heart Association (AHA) , 598 gravity-compensated shoulder-and-elbow robot, 599 gravity-compensated shoulder-elbow-and-wrist exoskeletal robot, 599–601

783 gravity non-compensated shoulder-and-elbow robot, 601 gym of upper extremity robots, 600 hand robot, 601–602 MIT-Manus, 598, 599 modularity, 599 plasticity, augment recovery bilateral vs. unilateral motor learning, 610–611 chronic care, 603–608 clinical perceptions, 608–610 inpatient care, 602–603 neuro-modulation, 611–612 neuro-recovery, 613–616 robot-mediated assay, 616–617 transition-to-task, 608 VA/DOD guidelines, 598 wrist robot, 601 Upper limb function, 42, 44, 47, 58, 59, 67–75, 78, 271 Upper limb (UL) movement, promoting/monitoring, 480–481, 485 providing feedback on exercise activities, 486–487 providing feedback on UL movement amount, 485 reminders to move, 486 Upper limb rehabilitation, SCI ARMin robot system, 640, 644 biomechanical measures, 65–66 angles, ROM and trajectories, 65–66 assessment of muscle strength, 65 body-worn sensors, 65–66 optical motion capture, 65 robotic systems, 66 clinical outcome assessments, 68–70, 72–73 distinctions of tetraplegic upper limb, 59–60 electrophysiology, 63–65 CHEP, 64 EMG, 64–65 MEP, 64 NCS, 64 SSEP, 64 importance of upper limb, 59 measurement properties, 62–63 measures of activity, 66–73 measures of body functions and structures, 63–66 measures of participation, 74 outcome measure frameworks, 60–62 technology in assessment and rehabilitation of, 57–78 therapeutic approaches, 74–77 therapy technologies, 306–307 upper limb training neuromodulation systems for, 76–77 passive workstations and interactive platforms for, 75–76 robotic systems for, 74–75 wearable technology role, 71 Upper limb training clinical adoption of passive devices, 537–538 compensatory strategies, 526–527 functional recovery mechanisms, 526–527

784 linear track system, 532 neuromodulation systems for, 76–77 passive exercise therapy devices, summary of, 539–540 passive gravity support systems, 529–531 passive workstations and interactive platforms for, 75–76 ReJoyce system, 534, 536 restoring hand function, 528 robotic exercise devices, 528–529 robotic systems for, 74–75 serious games, 533–536 SMART Arm, 532 tabletop therapy devices, 531–532 telerehabilitation, 537 therapeutic and functional electrical stimulation, 536 tone compensating orthoses, 532–533 wearable sensors to facilitate, 474–475 Usability tests, 277 User-centered design process, 277–278 User-centered technologies, 191 considerations adding resistance to increase workload, 199–201 ecologically valid tasks with graded difficulty, 198 motivational virtual experiences with sustained attention, 198–199 providing safety from falls, 197–198 frameworks intrinsic motivation and attention for learning theory, 194–197 Regulatory Focus Theory, 192–194 future development, 201–203 stakeholders, 202–203 US Model Spinal Cord Injury System, 281 Usual care group (UC), 605 UWB radio systems. See Ultra-wideband (UWB) radio systems

V VAEDA Glove. See Voice and Electromyography-Driven Actuated (VAEDA) Glove Vagus Nerve Stimulation (VNS), 536 Val66met polymorphism, 358 Validity, 62 Vanden Berghe Hand Function Test, 70 Vanderbilt exoskeleton. See Indego Van Lieshout Test (VLT), 641 Variable-Friction Cadense shoes, 711–712 VA-ROBOTICS study, 603–606, 608, 609 Vector Gait & Safety System, 275, 748 Ventral tegmental area (VTA), 8–9 Verticalisation, 746 Vertical support forces, 751–752 Veterans Administration/Department of Defense (VA/DOD) guidelines, for stroke care, 598 Vicious cycle, 485 Vicon system, 65

Index Video analysis techniques, 495–496 Virtual Peg Insertion Test (VPIT), 328, 329, 330, 333, 334 Virtual reality (VR), 195, 199, 200, 271–272, 291, 302, 305 balance and gait activity promotion, 452–454 custom systems, 448–451 non-custom systems, 451 neuroscience brain plasticity, 442–443 visuomotor representations, 443–444 robotic gait orthosis technology, 675 scenarios for arm training, 630 stroke sensorimotor neurorehabilitation auditory stimuli, 439 brain-computer interfaces, 440–441 gaming elements, 435–437 haptic and tactile stimuli, 439–440 immersion and cybersickness, 432 motor learning principles, 432–435 point of view, 438–439 presence and embodiment, 430–432 visual presentation, 438 upper extremities custom systems, 445–447 non-custom systems, 448 upper limb training, 529 VR-coupled locomotor system, 434 Virtuous cycle, 485 Visual feedback for heart rate control, 211 Visuomotor distortion, 231, 233 Visuomotor representations, 443–444 VLT. See Van Lieshout Test (VLT) VNS. See Vagus Nerve Stimulation (VNS) Voice analysis, 580 Voice and Electromyography-Driven Actuated (VAEDA) Glove, 125 Voluntary force, 673 VPIT. See Virtual Peg Insertion Test (VPIT) VR. See Virtual reality (VR) VRACK system, 453 VR exergames, 195–196 VR systems, 572 VTA. See Ventral tegmental area (VTA)

W Walk-Aide®, 408 Walking and balance training performance user-centered technologies for considerations, 197–201 frameworks, 192–197 future development, 201–203 Wearable cameras, 492–494 Wearable feedback, 488 Wearable sensors for stroke rehabilitation, 470 arm and hand movements, assessing

Index Action Research Arm Test (ARAT), 473 Box and Block Test (BBT), 473 clinical scores, estimating, 471–474 facilitating upper limb training, 474–475 Fugl-Meyer Assessment, Upper-Extremity subscale (FMA-UE), 472–473 Functional Ability Scale (FAS), 473 movement kinematics, estimating, 470–471 balance and mobility of stroke survivors, 476 10-m walk test (10 MWT), 476 clinical scores, estimating, 476–477 to facilitate gait training in the clinic, 477 Timed Up-and-Go (TUG) test, 476 Trunk Impairment Scale (TIS), 476–477 emerging technologies and applications, 488 collecting non-motor data, 496 e-skin sensors, 491–492 e-textiles, 489–491 modern video analysis techniques, 495–496 radio tags and radar-like technologies, 494–495 and traditional technologies, 496–497 wearable cameras, 492–494 future scenario, 477–478 to measure movement in the field, 478–480 critical information learned from wearable sensing, 482–484 monitoring lower limb movements, 481–482 monitoring upper limb movements, 480–481 to motivate movement and exercise, 484 lower limb movement amount, providing feedback on, 487–488 upper limb movement, promoting, 485–487 Weight-supporting technologies, 310

785 Western Aphasia Battery-Revised (WAB-R), 579 Wheelchair-mounted assistive robots iARM, 121 JAC02 robotic arm, 121 Wii™ Balance Board, 445 Wii™ Fit game, 445 Wii™ Resort games, 445 Wii™ system, 445, 448, 452 WMFT. See Wolf Motor Function Test (WMFT) Wolf Motor Functional Ability Scale, 240 Wolf Motor Function Test (WMFT), 72, 240, 359, 448, 472, 473, 534, 638, 639 Workload, adding resistance to increase, 199–201 WREX, 650 Wrist forces, cortical control of, 350 Wrist robot, 601 Wrist-worn sensors, 71

X Xbox 360, 445 Xbox Kinect, 533 X-Glove, 125

Y Yahr, Melvin, 172 YouGrabber system, 309, 311

Z ZeroG-Passive system, 750 ZeroG system, 275, 747