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IEEE JOURNAL OF BIOMEDICAL AND HEALTH INFORMATICS, VOL. 19, NO. 3, MAY 2015
Enabling Wireless Powering and Telemetry for Peripheral Nerve Implants Rangarajan Jegadeesan, Member, IEEE, Sudip Nag, Member, IEEE, Kush Agarwal, Student Member, IEEE, Nitish V. Thakor, Fellow, IEEE, and Yong-Xin Guo, Senior Member, IEEE
Abstract—Wireless power delivery and telemetry have enabled completely implantable neural devices. Current day implants are controlled, monitored, and powered wirelessly, eliminating the need for batteries and prolonging the lifetime. A brief overview of wireless platforms for such implantable devices is presented in this paper alongside an in-depth discussion of wireless platform for peripheral nerve implants covering design requirements, link design, and safety. Initial acute studies on the performance of the wireless power and data links in rodents are also presented. Index Terms—Inductive coupling, neural implants, peripheral nerve implants, telemetry, wireless power delivery.
I. INTRODUCTION EURAL implants [1]–[6] have become ubiquitous in recent times with some of them [3], [7]–[9] receiving FDA approvals for commercial use. Some of these neural implants are powered using nonrechargeable batteries, some using rechargeable batteries, and others sustainably powered by a wireless power transfer scheme. Implants powered using batteries have limited lifetime, and a surgical intervention is required to replace them. As batteries have finite recharge cycles, wirelessly charging the batteries only provides an incremental lifetime. Battery-free operation using sustainable wireless powering can extend the lifetime of implants as long as needed. For example, this approach is being adopted in current day retinal implants [9]. In this study, we first present the different approaches proposed to wirelessly power neural implants and compare their performance metrics against the implant requirements. We then discuss the implantable peripheral nerve prosthesis and its wireless design requirements. We propose a wireless platform
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Manuscript received November 3, 2014; revised February 23, 2015; accepted April 15, 2015. Date of publication April 21, 2015; date of current version May 7, 2015. This work was supported by the National Research Foundation, Singapore, under the Grant Award NRF-CRP10-2012-01. R. Jegadeesan and S. Nag are with the Singapore Institute for Neurotechnology Singapore, 117456 (e-mail: [email protected]; sudipnag1@ gmail.com). K. Agarwal is with the Department of Electrical and Computer Engineering, National University of Singapore, Singapore 117576, and also with Singapore Institute for Neurotechnology, Singapore 117456 (e-mail: agarwal.kush@ nus.edu.sg). N. V. Thakor is with the Singapore Institute for Neurotechnology, Singapore 117456, and also with the Department of Electrical and Computer Engineering, Department of Bioengineering, and the Department of Medicine, National University of Singapore, Singapore 117576 (e-mail: [email protected]). Y.-X. Guo is with the Department of Electrical and Computer Engineering, National University of Singapore, Singapore 117576, and also with the National University of Singapore Suzhou Research Institute, Suzhou 215123, China (e-mail: [email protected]). Color versions of one or more of the figures in this paper are available online at http://ieeexplore.ieee.org. Digital Object Identifier 10.1109/JBHI.2015.2424985
Fig. 1. Powering an implant device using the near-field inductive power transfer scheme. Vector B represents the magnetic field created by the transmitting coil that induces an EMF (V + , V − ) in the receiving coil.
(power delivery and data telemetry) for the peripheral nerve prosthesis (recording and stimulation implants) starting from the architecture and design to implementation and experimental results in rodents. The entire wireless platform was designed with an intent to meet the safety standards. To the authors’ knowledge, no work exists that has showcased a wireless platform with large power handling capability (>100 mW) and functional telemetry, while still adhering to safety standards (shown using both experiments and simulation). II. BACKGROUND Near-field inductive coupling [10]–[12], midfield resonant coupling [13], [14], near-field capacitive coupling [15], [16], far-field electromagnetic coupling [17]–[19], and acoustics [20], [21] are the various methods proposed to power implantable devices. Each method has its own merits and demerits as discussed in the following sections. A. Near-Field Inductive-Coupling Scheme The near-field inductive-coupling scheme works on the principle of electromagnetic induction. A transmitting coil placed close to the skin produces time-varying magnetic field, which induces an electromotive force (EMF) in the receiving coil placed inside the body as shown in Fig. 1.
2168-2194 © 2015 IEEE. Personal use is permitted, but republication/redistribution requires IEEE permission. See http://www.ieee.org/publications standards/publications/rights/index.html for more information.
JEGADEESAN et al.: ENABLING WIRELESS POWERING AND TELEMETRY FOR PERIPHERAL NERVE IMPLANTS
The induced EMF is tapped by the implanted device for its operation. There are numerous ways in which the induced EMF can be tapped by the implant device, the most efficient way being resonant coupling [22], [23]. Resonance improves power transfer efficiency, and in most cases, maximizes efficiency. Series and parallel resonant topologies equally benefit the power transfer link; however, the link characteristics vary with the type of resonance it uses. Various resonant network topologies that can be used with inductive coupling are analyzed, and the optimal design methodology to choose the link parameters (topology selection, optimal frequency of operation and load optimization) is presented in [12]. Inductive power link optimization based on the geometry [10], [11], [24], source-load decoupling [25], and multicoil linkage [26] have been proposed for powering generic implantable devices. With a complete understanding of the inductive power transfer scheme and its efficient design techniques, the focus of the current day research has shifted toward overcoming practical problems in implantable applications— namely coil misalignment, flexibility, and biocompatibility of the implant coil. Coil misalignment is common in inductively coupled links as the implant and external coils are in a moving frame. Coil alignment using permanent magnets is one of the earliest proposed solutions to overcome this issue and is widely used in current day implants [7], [8]. The alignment magnets generate heat during MRI (strengths of 1.5, 2 T), requiring them to be surgically removed before scanning. The problem has now been claimed to be addressed using suspended magnets, which align with the incident field allowing compatibility with MRI [6]. However, alternative approaches such as feedback coil [27] and back-telemetry [28], [29] based closed-loop power control have also been proposed to maintain sufficient power at implant. Flexibility of the implant coil allows hassle-free placement inside the body and efficiently uses the space allowed for the implant. The major shortcoming of the inductive powering link is that coil flexion detunes the resonance and the received power at the implant drops significantly. Approaches using flexible coils with serpentine structures of gold [30], [31] and liquid metals [32] have been proposed, although they report a marked drop in power delivered. The inductive-coupling scheme to transfer power to implants was the earliest proposed method [33], and many works have demonstrated system-level implementations of the scheme in retinal [34], [35], cortical [29], [36], cochlear [7], [8], cardiac [37], [38], and neural stimulator [3], [39] implants. Of all the proposed wireless power transfer schemes, only the near-field inductive power transfer scheme has been shown to be functional in actual implants and almost all of the FDA approved, wirelessly powered implants use it. B. Near-Field Capacitive-Coupling Scheme The near-field capacitive-coupling scheme works on the principle of charge displacement affected by capacitive coupling. Metal patches, one above the skin and one below the skin are used to create a capacitor supporting the displacement currents to transfer power wirelessly. Two such capacitances, one for the
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Fig. 2. Powering an implant device using the near-field capacitive-coupling scheme. Vector E denotes the electric field between the capacitor plates formed by the transmitting and receiving patches.
forward current and the other for the return path are used to build a capacitive coupled link as shown in Fig. 2. The modeling of the near-field capacitive power transfer scheme and the optimization for efficient operation is shown in [15] and [16]. Unlike the inductive power transfer scheme, the capacitive power transfer scheme supports flexible patch realizations, and the performance does not degrade much with flexion. It was shown that the capacitive scheme can deliver much larger power to the implant than the inductive scheme [40], while still adhering to the specific absorption rate (SAR) limits of 2 W/kg [41]. The main drawback of the capacitive power transfer scheme is its short range of operation. The power transfer efficiency drops significantly when the separation between the patches is more than 5 mm [15] and, hence, this technique is recommended only for subcutaneous applications. So far, no system-level functionality has been reported with a capacitive-coupled power transfer scheme. C. Far-Field Electromagnetic-Coupling Scheme The far-field electromagnetic-coupling scheme works on the principle of mutual impedance between a transmitting and receiving antenna. An external transmitting antenna radiates electromagnetic waves, which when incident on tuned receiver antenna placed inside the body generates a current across its terminals. The generated current is rectified and used by the implant device as shown in Fig. 3. The design of a far-field-based wireless power transfer scheme for biomedical implants has been presented in [17] and [19]. System-level design has been presented for cardiovascular pressure measurement implant in [42]. The far-field method can support a large range (few tens of millimeter) and can power deep seated implants using antennas much smaller than the ones used in near-field implants. However, since the link operates at gigahertz frequencies, power rectification losses are higher. This
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IEEE JOURNAL OF BIOMEDICAL AND HEALTH INFORMATICS, VOL. 19, NO. 3, MAY 2015
Fig. 3. Powering an implant device using a far-field electromagnetic-coupling scheme.
Fig. 5. Wireless power delivery using an acoustic scheme showing a piezo electric ultrasound generator producing oscillations and exciting a piezoelectric implanted receiver.
Fig. 4. Wireless power transfer using the midfield resonant-coupling scheme showing a transmitting antenna producing focused beams of electromagnetic fields (H t , H r ) and a miniaturized antenna (dimension a) receiving the power from both the far-field and near-field components.
coupled with the inherent low power transfer efficiency of the scheme makes it unattractive. D. Midfield Resonant-Coupling Scheme Midfield resonant power transfer scheme uses both near-field inductive-coupling and far-field electromagnetic coupling to transmit power from the external transmitter to receiver placed deep inside the body as shown in Fig. 4. The design of the midfield resonant-coupling scheme is presented in [14]. The optimal frequency of operation for various implant locations in the body is derived and reported in [13]. Transmitter design and optimization for the midfield resonantcoupling scheme is presented in [43]. The system-level design of powering scheme for cardiac implant is presented without experimental results in [44], and a miniaturized pacemaker implant much smaller than the traditional one was built and powered using this scheme and tested in a rabbit model [14]. Midfield resonant coupling is an emerging scheme proposed for use in deep seated ultralow power implants. This scheme boasts tiny implant receivers and functionality up to few tens of millimeter. However, the power handling capability is quite low (less than 5 mW).
Fig. 6. Comparison of performance metrics of wireless power delivery schemes.
E. Acoustic Scheme The acoustic power transfer scheme uses ultrasound to vibrate a piezoelectric energy harvester implanted inside the body as shown in Fig. 5. One design of the acoustic scheme is presented in [21]. A system-level design of the acoustic scheme for wireless powering of microoxygen generator implant is presented in [20], and it is claimed that the ultrasound-based power transfer scheme has lower power fluctuations caused by variations in alignment when compared with the near-field power transfer schemes. However, long-term effects of tissue vibration caused by this power transfer scheme are yet to be reported. F. Comparison of Power Delivery Schemes The power transfer capability and the range of power transfer offered by the aforementioned powering schemes have been compared against the neural implant requirements as shown below in Fig. 6. It can be seen that the near-field inductive-coupling
JEGADEESAN et al.: ENABLING WIRELESS POWERING AND TELEMETRY FOR PERIPHERAL NERVE IMPLANTS
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scheme can most comprehensively address the requirements of most of the neural implants. As the power requirements of the next-generation implants are expected to reduce, and the demand for smaller size and form factor increases, the benefits of emerging techniques, such as the midfield resonant coupling and acoustic power transfer can be reaped. Data telemetry to and from the implant is vital for controlling and monitoring its functionality. Out of all the proposed power transfer schemes, only the near-field methods and ultrasound are capable of coupling data on to the power link. More often than not, the control signal to the implant has low bandwidth and is transmitted over the power link using amplitude modulation or on–off keying. One exception is the retinal implant, where large bandwidth data corresponding to image acquisition need to be transmitted. A separate link for data transmission based on inductive coupling is used in this application [45], [46]. Similarly, small bandwidth telemetry from implants can be transmitted over the same power link using load modulation techniques as presented in [47]. However, for large bandwidth telemetry from implants, separate data link using near-field coupling [48] or traditional radiation-based far-field transceivers [49], [50] are used. III. PERIPHERAL NERVE PROSTHESES Peripheral nerve injuries result in dysfunction of the motor and sensory functions in limbs. Most of the times, surgical repair techniques proposed for such nerve injuries do not provide complete functional recovery. The best approach is to directly convey the nerve signals to the denervated muscles and restore functionality immediately and not wait for nerve regeneration. Such an approach needs a neural recording implant that can record and classify the nerve signals and convey them to the stimulator implant that provides patterned stimulation to achieve functional actions like hand grasp etc. The system-level diagram of the nerve prosthesis is shown below in Fig. 7. In this study, we present the wireless platform for the recording and stimulator implants aimed at building peripheral nerve prosthesis. A. Stimulator Implant Functional muscle stimulation is achieved by injecting current into corresponding muscles using electrodes in a particular sequence at a particular strength [51]. The stimulator implant has controllable multichannel accurate current drive capability to inject charge to stimulate the target muscles. The external control unit powers the implanted stimulator and sends control data defining the strength, duration, and type of stimulation along with the stimulation pattern to the implanted stimulator as depicted in Fig. 8(a). The stimulator unit acknowledges receipt of the stimulator control data using backchannel impedance modulation. B. Neural Recording The neural recording implant picks up neural signals from the nerve using multichannel electrodes, which are then amplified, digitized, and sent wirelessly to an external control unit as shown in Fig. 8(b). The external control unit wirelessly powers
Fig. 7.
Proposed peripheral nerve prosthesis in hand.
the implant unit and collects nerve signals from the implant, classifies them, and sends corresponding control signals to the stimulator unit to create a patterned stimulation. C. Wireless Power Delivery and Telemetry Requirements Functional muscle stimulation requires up to 10 mA of current with electrode impedances varying from 500 Ω to 1 kΩ. Hence, a peak power of 100 mW needs to be delivered to the stimulator implant for a sustainable operation. The stimulator implant also needs 4.8 kb/s (16 bytes of data to be sent within a 50 ms listening window of the stimulator to provide information about pulse duration, pulse strength, phase and packet frame bytes) bi-directional data telemetry for receiving the control signals and providing acknowledgement. The neural recording implant needs 35 mW of power for amplification, analog to digital conversion and digital logic functionality. A 1.3 Mb/s data link is required to transfer the neural signals (8 channel, 20 kHz data) from the implant to the external control unit. The complete requirements for the wireless platform are shown in Table I below. IV. WIRELESS PLATFORM FOR PERIPHERAL NERVE IMPLANTS A. Wireless Platform for Stimulator Implant Choosing the right wireless platform for the implant begins with the power requirements for the implant. With large power requirements (exceeding 10 mW) for both the stimulation and recording implants, the only proven method to sustainably transfer power to the implants is the near-field inductive-coupling scheme and is chosen for these implants. The data requirements
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Fig. 9. Schematic of the implant section of wireless platform for muscle stimulator.
Fig. 8. (a) Wireless scheme for the muscle stimulator implant shown using a cross-sectional view of the forearm. The power delivery and bidirectional data telemetry is done using the inductors L 1 and L 2 . (b) Wireless scheme for the neural recording implant shown using a cross-sectional view of the upper arm. The power is delivered using coils L 1 and L 2 , whereas data telemetry is performed using coils L D 1 and L D 2 . TABLE I SUMMARY OF REQUIREMENTS FOR THE WIRELESS PLATFORM USED IN PERIPHERAL NERVE PROSTHESES Fig. 10. Schematic of the external section of wireless platform for muscle stimulator. Implant type
Power Requirement
Data to Implant
Data from Implant
Stimulation Recording
100 mW 35 mW
4.8 kb/s NA
4.8 kb/s 1.3 Mb/s
are quite different with a stimulator implant needing bidirectional telemetry at low bandwidth (4.8 kb/s) and the recording unit requiring large bandwidth telemetry (1.3 Mb/s for four channel neural signal acquisition at 20 KHz) from the implant. Bidirectional telemetry for the stimulator is easily achieved by modulating the data on to the power link. However, for the recording implant, a separate data link is needed to meet the bandwidth requirements. The wireless platform to power the peripheral nerve implants and perform telemetry is discussed in the following sections.
Power to the implanted stimulator is delivered using an inductive power link comprising of a transmitting coil L1 placed outside the body in the forearm and a receiving coil L2 , implanted just beneath the skin in the forearm as shown in Fig. 8(a). The parallel resonant topology is chosen to meet the 10-V rectified voltage requirement for the stimulator. The power received at the implanted coil L2 is utilized by the stimulator implant at two voltage levels. While the current stimulator circuitry directly taps part of the power from the unregulated voltage (which needs to be at least 10 V under loading conditions for proper operation), the digital logic circuitry draws its power from the 3.3 V regulated supply. Control data are sent from the external stimulator unit to the implant using an on–off keying technique (see M2 in Fig. 10).
JEGADEESAN et al.: ENABLING WIRELESS POWERING AND TELEMETRY FOR PERIPHERAL NERVE IMPLANTS
This keying disturbance is reflected at the implant side as intermittent absence of power and is recovered as control data using a diode demodulator. The capacitor Co (see Fig. 9) momentarily powers the digital circuitry during power fluctuations caused by modulation. The received data are decoded by the digital logic, and the decoded command is used to set the stimulation parameters. The data are looped back to the transmitter as an acknowledgement (ACK) using impedance modulation effected by the detuning capacitor Cm and switch M4 . This impedance modulation is reflected at the transmitting coil L1 as voltage fluctuations. Since the transmit power is quite large, the voltage fluctuations cannot be directly demodulated without using extra circuitry for high-voltage protection of demodulator. The class E amplifier is also detuned by the demodulator input increasing power losses. Hence, a small reader coil LD 1 is used to pick up the modulated data from the transmit coil which is then demodulated and recovered using a logic detector as acknowledgement. It is worth mentioning that the coupling between the transmitting coil L1 and the data reader coil LD 1 needs to be kept to a minimal value (coupling coefficient
IEEE JOURNAL OF BIOMEDICAL AND HEALTH INFORMATICS, VOL. 19, NO. 3, MAY 2015
Enabling Wireless Powering and Telemetry for Peripheral Nerve Implants Rangarajan Jegadeesan, Member, IEEE, Sudip Nag, Member, IEEE, Kush Agarwal, Student Member, IEEE, Nitish V. Thakor, Fellow, IEEE, and Yong-Xin Guo, Senior Member, IEEE
Abstract—Wireless power delivery and telemetry have enabled completely implantable neural devices. Current day implants are controlled, monitored, and powered wirelessly, eliminating the need for batteries and prolonging the lifetime. A brief overview of wireless platforms for such implantable devices is presented in this paper alongside an in-depth discussion of wireless platform for peripheral nerve implants covering design requirements, link design, and safety. Initial acute studies on the performance of the wireless power and data links in rodents are also presented. Index Terms—Inductive coupling, neural implants, peripheral nerve implants, telemetry, wireless power delivery.
I. INTRODUCTION EURAL implants [1]–[6] have become ubiquitous in recent times with some of them [3], [7]–[9] receiving FDA approvals for commercial use. Some of these neural implants are powered using nonrechargeable batteries, some using rechargeable batteries, and others sustainably powered by a wireless power transfer scheme. Implants powered using batteries have limited lifetime, and a surgical intervention is required to replace them. As batteries have finite recharge cycles, wirelessly charging the batteries only provides an incremental lifetime. Battery-free operation using sustainable wireless powering can extend the lifetime of implants as long as needed. For example, this approach is being adopted in current day retinal implants [9]. In this study, we first present the different approaches proposed to wirelessly power neural implants and compare their performance metrics against the implant requirements. We then discuss the implantable peripheral nerve prosthesis and its wireless design requirements. We propose a wireless platform
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Manuscript received November 3, 2014; revised February 23, 2015; accepted April 15, 2015. Date of publication April 21, 2015; date of current version May 7, 2015. This work was supported by the National Research Foundation, Singapore, under the Grant Award NRF-CRP10-2012-01. R. Jegadeesan and S. Nag are with the Singapore Institute for Neurotechnology Singapore, 117456 (e-mail: [email protected]; sudipnag1@ gmail.com). K. Agarwal is with the Department of Electrical and Computer Engineering, National University of Singapore, Singapore 117576, and also with Singapore Institute for Neurotechnology, Singapore 117456 (e-mail: agarwal.kush@ nus.edu.sg). N. V. Thakor is with the Singapore Institute for Neurotechnology, Singapore 117456, and also with the Department of Electrical and Computer Engineering, Department of Bioengineering, and the Department of Medicine, National University of Singapore, Singapore 117576 (e-mail: [email protected]). Y.-X. Guo is with the Department of Electrical and Computer Engineering, National University of Singapore, Singapore 117576, and also with the National University of Singapore Suzhou Research Institute, Suzhou 215123, China (e-mail: [email protected]). Color versions of one or more of the figures in this paper are available online at http://ieeexplore.ieee.org. Digital Object Identifier 10.1109/JBHI.2015.2424985
Fig. 1. Powering an implant device using the near-field inductive power transfer scheme. Vector B represents the magnetic field created by the transmitting coil that induces an EMF (V + , V − ) in the receiving coil.
(power delivery and data telemetry) for the peripheral nerve prosthesis (recording and stimulation implants) starting from the architecture and design to implementation and experimental results in rodents. The entire wireless platform was designed with an intent to meet the safety standards. To the authors’ knowledge, no work exists that has showcased a wireless platform with large power handling capability (>100 mW) and functional telemetry, while still adhering to safety standards (shown using both experiments and simulation). II. BACKGROUND Near-field inductive coupling [10]–[12], midfield resonant coupling [13], [14], near-field capacitive coupling [15], [16], far-field electromagnetic coupling [17]–[19], and acoustics [20], [21] are the various methods proposed to power implantable devices. Each method has its own merits and demerits as discussed in the following sections. A. Near-Field Inductive-Coupling Scheme The near-field inductive-coupling scheme works on the principle of electromagnetic induction. A transmitting coil placed close to the skin produces time-varying magnetic field, which induces an electromotive force (EMF) in the receiving coil placed inside the body as shown in Fig. 1.
2168-2194 © 2015 IEEE. Personal use is permitted, but republication/redistribution requires IEEE permission. See http://www.ieee.org/publications standards/publications/rights/index.html for more information.
JEGADEESAN et al.: ENABLING WIRELESS POWERING AND TELEMETRY FOR PERIPHERAL NERVE IMPLANTS
The induced EMF is tapped by the implanted device for its operation. There are numerous ways in which the induced EMF can be tapped by the implant device, the most efficient way being resonant coupling [22], [23]. Resonance improves power transfer efficiency, and in most cases, maximizes efficiency. Series and parallel resonant topologies equally benefit the power transfer link; however, the link characteristics vary with the type of resonance it uses. Various resonant network topologies that can be used with inductive coupling are analyzed, and the optimal design methodology to choose the link parameters (topology selection, optimal frequency of operation and load optimization) is presented in [12]. Inductive power link optimization based on the geometry [10], [11], [24], source-load decoupling [25], and multicoil linkage [26] have been proposed for powering generic implantable devices. With a complete understanding of the inductive power transfer scheme and its efficient design techniques, the focus of the current day research has shifted toward overcoming practical problems in implantable applications— namely coil misalignment, flexibility, and biocompatibility of the implant coil. Coil misalignment is common in inductively coupled links as the implant and external coils are in a moving frame. Coil alignment using permanent magnets is one of the earliest proposed solutions to overcome this issue and is widely used in current day implants [7], [8]. The alignment magnets generate heat during MRI (strengths of 1.5, 2 T), requiring them to be surgically removed before scanning. The problem has now been claimed to be addressed using suspended magnets, which align with the incident field allowing compatibility with MRI [6]. However, alternative approaches such as feedback coil [27] and back-telemetry [28], [29] based closed-loop power control have also been proposed to maintain sufficient power at implant. Flexibility of the implant coil allows hassle-free placement inside the body and efficiently uses the space allowed for the implant. The major shortcoming of the inductive powering link is that coil flexion detunes the resonance and the received power at the implant drops significantly. Approaches using flexible coils with serpentine structures of gold [30], [31] and liquid metals [32] have been proposed, although they report a marked drop in power delivered. The inductive-coupling scheme to transfer power to implants was the earliest proposed method [33], and many works have demonstrated system-level implementations of the scheme in retinal [34], [35], cortical [29], [36], cochlear [7], [8], cardiac [37], [38], and neural stimulator [3], [39] implants. Of all the proposed wireless power transfer schemes, only the near-field inductive power transfer scheme has been shown to be functional in actual implants and almost all of the FDA approved, wirelessly powered implants use it. B. Near-Field Capacitive-Coupling Scheme The near-field capacitive-coupling scheme works on the principle of charge displacement affected by capacitive coupling. Metal patches, one above the skin and one below the skin are used to create a capacitor supporting the displacement currents to transfer power wirelessly. Two such capacitances, one for the
959
Fig. 2. Powering an implant device using the near-field capacitive-coupling scheme. Vector E denotes the electric field between the capacitor plates formed by the transmitting and receiving patches.
forward current and the other for the return path are used to build a capacitive coupled link as shown in Fig. 2. The modeling of the near-field capacitive power transfer scheme and the optimization for efficient operation is shown in [15] and [16]. Unlike the inductive power transfer scheme, the capacitive power transfer scheme supports flexible patch realizations, and the performance does not degrade much with flexion. It was shown that the capacitive scheme can deliver much larger power to the implant than the inductive scheme [40], while still adhering to the specific absorption rate (SAR) limits of 2 W/kg [41]. The main drawback of the capacitive power transfer scheme is its short range of operation. The power transfer efficiency drops significantly when the separation between the patches is more than 5 mm [15] and, hence, this technique is recommended only for subcutaneous applications. So far, no system-level functionality has been reported with a capacitive-coupled power transfer scheme. C. Far-Field Electromagnetic-Coupling Scheme The far-field electromagnetic-coupling scheme works on the principle of mutual impedance between a transmitting and receiving antenna. An external transmitting antenna radiates electromagnetic waves, which when incident on tuned receiver antenna placed inside the body generates a current across its terminals. The generated current is rectified and used by the implant device as shown in Fig. 3. The design of a far-field-based wireless power transfer scheme for biomedical implants has been presented in [17] and [19]. System-level design has been presented for cardiovascular pressure measurement implant in [42]. The far-field method can support a large range (few tens of millimeter) and can power deep seated implants using antennas much smaller than the ones used in near-field implants. However, since the link operates at gigahertz frequencies, power rectification losses are higher. This
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IEEE JOURNAL OF BIOMEDICAL AND HEALTH INFORMATICS, VOL. 19, NO. 3, MAY 2015
Fig. 3. Powering an implant device using a far-field electromagnetic-coupling scheme.
Fig. 5. Wireless power delivery using an acoustic scheme showing a piezo electric ultrasound generator producing oscillations and exciting a piezoelectric implanted receiver.
Fig. 4. Wireless power transfer using the midfield resonant-coupling scheme showing a transmitting antenna producing focused beams of electromagnetic fields (H t , H r ) and a miniaturized antenna (dimension a) receiving the power from both the far-field and near-field components.
coupled with the inherent low power transfer efficiency of the scheme makes it unattractive. D. Midfield Resonant-Coupling Scheme Midfield resonant power transfer scheme uses both near-field inductive-coupling and far-field electromagnetic coupling to transmit power from the external transmitter to receiver placed deep inside the body as shown in Fig. 4. The design of the midfield resonant-coupling scheme is presented in [14]. The optimal frequency of operation for various implant locations in the body is derived and reported in [13]. Transmitter design and optimization for the midfield resonantcoupling scheme is presented in [43]. The system-level design of powering scheme for cardiac implant is presented without experimental results in [44], and a miniaturized pacemaker implant much smaller than the traditional one was built and powered using this scheme and tested in a rabbit model [14]. Midfield resonant coupling is an emerging scheme proposed for use in deep seated ultralow power implants. This scheme boasts tiny implant receivers and functionality up to few tens of millimeter. However, the power handling capability is quite low (less than 5 mW).
Fig. 6. Comparison of performance metrics of wireless power delivery schemes.
E. Acoustic Scheme The acoustic power transfer scheme uses ultrasound to vibrate a piezoelectric energy harvester implanted inside the body as shown in Fig. 5. One design of the acoustic scheme is presented in [21]. A system-level design of the acoustic scheme for wireless powering of microoxygen generator implant is presented in [20], and it is claimed that the ultrasound-based power transfer scheme has lower power fluctuations caused by variations in alignment when compared with the near-field power transfer schemes. However, long-term effects of tissue vibration caused by this power transfer scheme are yet to be reported. F. Comparison of Power Delivery Schemes The power transfer capability and the range of power transfer offered by the aforementioned powering schemes have been compared against the neural implant requirements as shown below in Fig. 6. It can be seen that the near-field inductive-coupling
JEGADEESAN et al.: ENABLING WIRELESS POWERING AND TELEMETRY FOR PERIPHERAL NERVE IMPLANTS
961
scheme can most comprehensively address the requirements of most of the neural implants. As the power requirements of the next-generation implants are expected to reduce, and the demand for smaller size and form factor increases, the benefits of emerging techniques, such as the midfield resonant coupling and acoustic power transfer can be reaped. Data telemetry to and from the implant is vital for controlling and monitoring its functionality. Out of all the proposed power transfer schemes, only the near-field methods and ultrasound are capable of coupling data on to the power link. More often than not, the control signal to the implant has low bandwidth and is transmitted over the power link using amplitude modulation or on–off keying. One exception is the retinal implant, where large bandwidth data corresponding to image acquisition need to be transmitted. A separate link for data transmission based on inductive coupling is used in this application [45], [46]. Similarly, small bandwidth telemetry from implants can be transmitted over the same power link using load modulation techniques as presented in [47]. However, for large bandwidth telemetry from implants, separate data link using near-field coupling [48] or traditional radiation-based far-field transceivers [49], [50] are used. III. PERIPHERAL NERVE PROSTHESES Peripheral nerve injuries result in dysfunction of the motor and sensory functions in limbs. Most of the times, surgical repair techniques proposed for such nerve injuries do not provide complete functional recovery. The best approach is to directly convey the nerve signals to the denervated muscles and restore functionality immediately and not wait for nerve regeneration. Such an approach needs a neural recording implant that can record and classify the nerve signals and convey them to the stimulator implant that provides patterned stimulation to achieve functional actions like hand grasp etc. The system-level diagram of the nerve prosthesis is shown below in Fig. 7. In this study, we present the wireless platform for the recording and stimulator implants aimed at building peripheral nerve prosthesis. A. Stimulator Implant Functional muscle stimulation is achieved by injecting current into corresponding muscles using electrodes in a particular sequence at a particular strength [51]. The stimulator implant has controllable multichannel accurate current drive capability to inject charge to stimulate the target muscles. The external control unit powers the implanted stimulator and sends control data defining the strength, duration, and type of stimulation along with the stimulation pattern to the implanted stimulator as depicted in Fig. 8(a). The stimulator unit acknowledges receipt of the stimulator control data using backchannel impedance modulation. B. Neural Recording The neural recording implant picks up neural signals from the nerve using multichannel electrodes, which are then amplified, digitized, and sent wirelessly to an external control unit as shown in Fig. 8(b). The external control unit wirelessly powers
Fig. 7.
Proposed peripheral nerve prosthesis in hand.
the implant unit and collects nerve signals from the implant, classifies them, and sends corresponding control signals to the stimulator unit to create a patterned stimulation. C. Wireless Power Delivery and Telemetry Requirements Functional muscle stimulation requires up to 10 mA of current with electrode impedances varying from 500 Ω to 1 kΩ. Hence, a peak power of 100 mW needs to be delivered to the stimulator implant for a sustainable operation. The stimulator implant also needs 4.8 kb/s (16 bytes of data to be sent within a 50 ms listening window of the stimulator to provide information about pulse duration, pulse strength, phase and packet frame bytes) bi-directional data telemetry for receiving the control signals and providing acknowledgement. The neural recording implant needs 35 mW of power for amplification, analog to digital conversion and digital logic functionality. A 1.3 Mb/s data link is required to transfer the neural signals (8 channel, 20 kHz data) from the implant to the external control unit. The complete requirements for the wireless platform are shown in Table I below. IV. WIRELESS PLATFORM FOR PERIPHERAL NERVE IMPLANTS A. Wireless Platform for Stimulator Implant Choosing the right wireless platform for the implant begins with the power requirements for the implant. With large power requirements (exceeding 10 mW) for both the stimulation and recording implants, the only proven method to sustainably transfer power to the implants is the near-field inductive-coupling scheme and is chosen for these implants. The data requirements
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Fig. 9. Schematic of the implant section of wireless platform for muscle stimulator.
Fig. 8. (a) Wireless scheme for the muscle stimulator implant shown using a cross-sectional view of the forearm. The power delivery and bidirectional data telemetry is done using the inductors L 1 and L 2 . (b) Wireless scheme for the neural recording implant shown using a cross-sectional view of the upper arm. The power is delivered using coils L 1 and L 2 , whereas data telemetry is performed using coils L D 1 and L D 2 . TABLE I SUMMARY OF REQUIREMENTS FOR THE WIRELESS PLATFORM USED IN PERIPHERAL NERVE PROSTHESES Fig. 10. Schematic of the external section of wireless platform for muscle stimulator. Implant type
Power Requirement
Data to Implant
Data from Implant
Stimulation Recording
100 mW 35 mW
4.8 kb/s NA
4.8 kb/s 1.3 Mb/s
are quite different with a stimulator implant needing bidirectional telemetry at low bandwidth (4.8 kb/s) and the recording unit requiring large bandwidth telemetry (1.3 Mb/s for four channel neural signal acquisition at 20 KHz) from the implant. Bidirectional telemetry for the stimulator is easily achieved by modulating the data on to the power link. However, for the recording implant, a separate data link is needed to meet the bandwidth requirements. The wireless platform to power the peripheral nerve implants and perform telemetry is discussed in the following sections.
Power to the implanted stimulator is delivered using an inductive power link comprising of a transmitting coil L1 placed outside the body in the forearm and a receiving coil L2 , implanted just beneath the skin in the forearm as shown in Fig. 8(a). The parallel resonant topology is chosen to meet the 10-V rectified voltage requirement for the stimulator. The power received at the implanted coil L2 is utilized by the stimulator implant at two voltage levels. While the current stimulator circuitry directly taps part of the power from the unregulated voltage (which needs to be at least 10 V under loading conditions for proper operation), the digital logic circuitry draws its power from the 3.3 V regulated supply. Control data are sent from the external stimulator unit to the implant using an on–off keying technique (see M2 in Fig. 10).
JEGADEESAN et al.: ENABLING WIRELESS POWERING AND TELEMETRY FOR PERIPHERAL NERVE IMPLANTS
This keying disturbance is reflected at the implant side as intermittent absence of power and is recovered as control data using a diode demodulator. The capacitor Co (see Fig. 9) momentarily powers the digital circuitry during power fluctuations caused by modulation. The received data are decoded by the digital logic, and the decoded command is used to set the stimulation parameters. The data are looped back to the transmitter as an acknowledgement (ACK) using impedance modulation effected by the detuning capacitor Cm and switch M4 . This impedance modulation is reflected at the transmitting coil L1 as voltage fluctuations. Since the transmit power is quite large, the voltage fluctuations cannot be directly demodulated without using extra circuitry for high-voltage protection of demodulator. The class E amplifier is also detuned by the demodulator input increasing power losses. Hence, a small reader coil LD 1 is used to pick up the modulated data from the transmit coil which is then demodulated and recovered using a logic detector as acknowledgement. It is worth mentioning that the coupling between the transmitting coil L1 and the data reader coil LD 1 needs to be kept to a minimal value (coupling coefficient
