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# Multisite bio-stimulating implants magnetoelectrically powered and individually programmed by a single transmitter

Authors:
Multisite bio-stimulating implants magnetoelectrically
powered and individually programmed by a single
transmitter
Zhanghao Yu, Joshua C. Chen, Yan He, Fatima T. Alrashdan,
Benjamin W. Avants, Amanda Singer, Jacob T. Robinson,
Kaiyuan Yang
Rice University, Houston, TX
Implantable bioelectronics for electrically modulating activities of
specific cells have shown great success and exciting potential in
treating a wide range of diseases. Some of the most representative
therapies are cardiac pacemakers and neuromodulators for motor
function restoration, pain relief and neural disorder treatment [1,2].
While several wireless miniaturized bio-stimulators have been
demonstrated [3-6], most of them lack the capability of coordinated
multisite stimulation, which is shown to be more effective in many
scenarios [1,2]. Equipping an implant with electrode/LED arrays is a
straightforward approach to add extra stimulation channels [7-9], but
the deployment flexibility of stimulating spots is limited due to leads.
[10] shows a wired retinal stimulator array to scale up the driving
capability and ensure synchronization, but the heavy use of leads
severely limits its applications. A two-site heart pacing system [2] is
proposed with two independently powered and controlled implants
for flexible leadless deployment. Because the implants are
multiplexing, they face stricter EM exposure constrains for power
transmission, more challenging device synchronization, and limited
scalability to more implants. To circumvent these problems, this
paper presents a hardware platform for coordinated and miniaturized
multisite stimulating implants, wirelessly powered and controlled by
a single TX. Magnetoelectric (ME) wireless power transfer with high
power and efficiency, low body absorption, and less sensitivity to
misalignment [4,5], is co-designed with a robust SoC to enable
reliable operation and individual programmability of the implants. The
presented system features: (1) robust operation with 2V source
amplitude variations, covering up to 40mm distance between TX and
implants; (2) individual addressability and programmability of each
implant, leveraging PUF IDs; (3) >90% chip efficiency for 1.5-to-3.5V
stimulation with fully programmable parameters; (4) no extra TX
output power required for additional implants; (5) miniaturized
implants with 6.2mm3 volume and 30mg mass.
The implant integrates a ME film, a capacitor, on-board electrodes,
and a SoC. The SoC interfaces with the ME film to receive power
and data and drives programmable stimulation. ME induced voltage
is rectified to Vrect and then converted by an adaptive switched-
capacitor power converter (SCPC), which provides proper voltage
and buffers energy on the off-chip capacitor for stimulation, and
provides VDD_H as a high-voltage supply for SoC. A 1V supply VDD_L
is generated by LDO (Fig.1 bottom). Each implant cycles through
charging, data transfer and stimulation phases. To maintain reliable
synchronized operation of multiple implants under different ME
voltages caused by different implantation depth and body movement,
the phase transitions are solely controlled by the TX with a short
notch of magnetic field (Fig.2 top). Comparator outputs in the active
rectifier are reused as watchdog signals to detect the notches.
Meanwhile, a global clock is extracted from the source by sensitive
clock recovery circuit, ensuring synchronization among all implants.
Individually programing every implant by a shared TX is critical for
effective and flexible stimulating therapies. Downlink data transferred
by ASK modulation contains a preamble for real-time demodulation
threshold calibration, an 8-bit ID for addressing, and a 19-bit data
payload for calibration and stimulation settings. The data update
controller checks the ID in packet against the on-chip ID to decide
whether to accept the new data. The on-chip 8-bit ID is realized with
CMOS physical unclonable functions (PUF) leveraging transistor
intrinsic variations to cheaply generate and store device-specific IDs
(Fig.2 bottom). A inverter chain based PUF design with native NMOS
regulation [11] is employed. Because of the narrow operating
temperature range and the native voltage regulation, 15-cycle
temporal majority voting (TMV) is sufficient to filter out thermal noise
and ensure PUF’s reliability. The ID generation is triggered by power-
on reset signal and clock gated after the ID is loaded to registers.
Variations of input voltage and
power of implants, caused by their
distance and misalignment with the
TX, are unavoidable in practice,
especially for multisite implants.
Thus, robust power recovery to
support stimulation across a wide
source conditions is highly desired.
Simply generating a high enough
voltage for stimulation driver
(VDD_stim) may ensure robustness
but will suffer from high power loss
and thus heat dissipation [3]. Alternatively, unregulated voltage
stimulation by directly driving electrodes with charged capacitors has
high efficiency but sacrifices precise charge deposition control [6]. To
achieve the desired robustness and efficiency without a complicated
feedback and reconfiguration loop as in [12], the proposed SCPC
directly generate a VDD_stim that is 10% higher than the desired
stimulus amplitude, and relies on the off-chip capacitor and a
regulator-style stimulation driver to support regulated mono- and bi-
phasic stimulus. Regulation of VDD_stim is realized by disconnecting
the capacitor from SCPC, the core of which is a charge pump and a
charging controller, once it reaches the desired level. High-speed
amplifiers inside the stimulation driver regulates the stimulus. To
save power, the amplifier will only be turned on in the stimulation
phase. The SCPC also includes an always-on high voltage selector
to generate VDD_H, which connects VDD_H to the higher one between
Vrect and VDD_stim and guarantees cold startup using Vrect (Fig. 3).
Fig.4 captures the operation waveforms of the implant. VDD_stim is
charged up and regulated to 2.75V, then drops to 2.15V after the
2.5V, 1.2ms bi-phasic stimulation. It is verified that the implant
maintains its operation with maximum stimulation amplitude (3.5V)
under large ME source variations (1.52.7V). 90% stimulating
efficiency is achieved as long as the amplitude is larger than 1.5V.
Power transfer at various distances are measured, which shows a
maximum TX-RX distance of 40mm and a highest power transfer
efficiency (PTE) of 1.03%. Individual programming of two implants
by a single TX is illustrated in Fig.5 (top left).
An in-vitro test with the a 2cm thick porcine tissue as a medium is
conducted (Fig.5, top right), which demonstrates flexible implant
deployment covering a space with 35mm radius, and synchronized
stimulations by two implants with programmed 0.01-to-0.8ms delays.
Based on simulation of the specific absorption rate (SAR) and the
electric field induction in a coil-generated 330kHz magnetic field, a
magnetic strength of 0.1mT, which is enough to sustain implant’s
functionality, can be delivered to a depth of 60mm without violating
the IEEE C95.1-2019 standards (unrestricted environment).
The proposed system is further validated in-vivo using a transgenic
line of Hydra vulgaris as a model for muscle stimulation and a rat
model for neural stimulation. Hydra naturally express a calcium
sensitive fluorescent protein, GCaMP7b, as well as voltage-gated ion
channels. To model synchronous stimulation of muscle tissue, two
hydra are used. In order to synchronize the muscle contractions, we
provide 3.5V, 20Hz, 1.2ms pulse width, biphasic stimulation pulse
trains. This results in >200% GCaMP7b fluorescence increases
which demonstrates activation of ion channels resulting in stimulus
aligned muscle contractions in both organisms (Fig.6 top left). We
also stimulate the sciatic nerve of the rat with varying amplitudes. A
graded response in the intensity of the rat leg kick is measured with
EMG of the plantar muscles (Fig.6 top right). The comparison table
with other bio-stimulating systems is given in Fig. 6 (bottom).
References:
[1] S. Harkema et al., Lancet, 2011, DOI: 10.1016/S0140-6736(11)60547-3.
[2] H. Lyu et al., Sci. Rep., 2020, DOI: 10.1038/s41598-020-59017-z.
[3] D. K. Piech et al., Nat. BME., 2020, DOI: 10.1038/s41551-020-0518-9.
[4] A. Singer et al., Neuron, 2020, DOI: 10.1016/j.neuron.2020.05.019.
[5] Z. Yu et al., TBioCAS, 2020, DOI: 10.1109/TBCAS.2020.3037862.
[6] H. M. Lee et al., JSSC, 2015, DOI: 10.1109/JSSC.2014.2355814.
[7] Y. K. Lo et al., ISSCC, 2016, DOI: 10.1109/ISSCC.2016.7418067.
[8] Y. Jia et al., ISSCC, 2018, DOI: 10.1109/ISSCC.2018.8310387.
[9] Y. Jia et al., ISSCC, 2020, DOI: 10.1109/ISSCC19947.2020.9063065.
[10] T. Tokuda et al., TBioCAS, 2010, DOI: 10.1109/TBCAS.2010.2078508.
[11] D. Li et al., ISSCC, 2019, DOI: 10.1109/ISSCC.2019.8662537.
[12] H. M. Lee et al., JSSC, 2013, DOI: 10.1109/JSSC.2013.2266862.
Active
Area
Capacitor for Rectifier
Capacitor
for
SCPC
Capacitor
for
LDO
Capacitor for
Data and
Control
1mm
0.8mm
Die micrograph.
© 2021 IEEE. Personal use is permitted, but republication/redistribution requires IEEE permission.
Implants
Control Stimulators
B: Wired Array
of Stimulators [10]
C: Multiple TXs +
Stim. Implants Pairs [2]
TXs
Stim. Implants
TX
Stim.
Implant
Biventricular
Cardiac Pacing
+ High Channel Density
+ Compatibility
Fixed Channel Quantity
Limited Deployment
+ Flexible Deployment
Higher Body Absorption
+ Flexible Deployment
+ Higher Efficiency
+ Good Synchronization
+ Good Scalability
+ High Channel Density
High Risks of Infection
Fixed Channel Quantity
Limited Deployment
A: Single Implant with
Electrode/LED Array [7-9]
TX
Stim. Implants
AB
Multisite Spinal
Cord Stimulation
2mm
6.2mm3
30mg
Proposed Implants Power Managem ent
Data Recovery
Electrodes
ME Film
SoC
Central
Controller
PUF Register
Data
Data
Write
CLKdata
Control
WDa,b
Data
Active
Rectifier
VDD_H
Vrect
Stimulation
Controller
VDD_stim
CLK
Recovery
ID
Check
Vref
Cstore
Voltage
Ref.
SCPC
Phase
Switch
Amp. & Timing
Vref
CLKG & CLKstim
Demodulator
VDD_L
Stimulation
Driver
VDD_stim
Phase
Switch
Stim. Generation
POR
LDO
+ POR
Vrect
VDD_H
This Work: Single TX +
Multiple Stim. Implants
Applied
Magnetic
Field
+ -
+ -
+ -
Coil + -
PZT Metglas
V
ME
Films
Magnetoelectric (ME) Effects
Electrodes
Fig. 1. Concepts of multisite bio-stimulation and various multisite
stimulating system structures; illustration of the implant, architecture
of its SoC, and principles of ME power transfer for multiple RXs.
ID Check
Vrect
+
-
ENtrack
Data Recovery
Controller
Data
PUF
with TMV
POR
POR
ENPUF
Data
Register
ENtrack
Pass?Y N
Data Update
D[18:0]
32 bits
Vth Tracking
2 bit8 bits
ID
Check Gap Data
Update
19 bits
...
...
Footer
6 bits 4 bits 5 bits 3 bits
Stim.
Amp. Stim.
Delay
Stim.
Pulse Width
Vref
Trimming
...
...
Data Transmission Protocol and Individually Programm ing Scheme
Vac2
Vac2
Vrect
+
-
Vac1
D Q
CK
RN
+
-
POR
ENdata
WDa
WDb
CLKG
Vac1
Vac1
Operating Scheme an d Synchronized Timing Reference s Generation
fsfs/n
ENdata&ENstim
CLKdata
CLKstim
VSS
ME
Envelope
Vrect
CLKG
ENdata
ENstim Vref(0.2V)
D Q
CK
RN
S Q
RQN
CNT[3]
Register
RN
CK
D
ENPUF
CLKG
PUF Cell
Selector
RN
CK
S[3:0]
S[3]
X8
MUX
PUF Cell
Frequency
Divider
R
Rw
Rw
TMV
Counter
RN
CK
POR
POR
PUFout[7:0]
PUF with Tempor al Majority Voting (TMV)
R
POR
VDD_L
VSS
0.7V
PUF[i]
Native
NMOS
PUF
Cell
...
...
...
...
...
...
...
ENPUF
PUF[i] [0] [1] [7]
Errors due to Noise
Y
RW
PUFOUT[7]
PUFOUT[1]
PUFOUT[0]
1
0
1
Count
0 1
0
20
40
60
Percentage (%)
PUF Output (V)
CLKdata
Data Transfer Stimulation
Envelope
Extractor
Fig. 2. Diagram of SoC operating phase transition and clock
recovery; diagram of data recovery circuitry and schematics of PUF.
Fig. 3. Principles of the proposed highly efficient voltage stimulation;
schematics of adaptive power converter and stimulation driver.
VDD_stim
Stimulation at 20Hz with Rload = 1kΩ
Data
2.5V
-2.37V
2.75V
2.15V
Data
Update
1.2ms
Bi-phasic
Mode
Demodulation
Vth Tracking
1V 40ms
ID
Check
Stimulation Phase
50ms
Charging
Measured Stimulating EfficiencyMeasured Stimulation with Source Variation s
0 5 10 15
0
1
2
3
4 Stim. Amplitude
VDD_stim
Stimulation Amplitude Setting
Voltage (V)
0
20
40
60
80
100
Stim. Efficiency (%)
Stim. Efficiency
> 90%
0.05 0.15
0
1.75
3.5
5.25
1.2
1.8
2.4
3.2
Voltage (V)
Time (sec)
Vrect
0.1
VME =2.7V VME =2.2V VME =1.5V
Notches for
phase transi tion
Drops due to
Cstore charging
Stim.
3.5V
2.4V
1.2V
Measured Power Delivery vs. Distances
0 5 10 15 20 25 30 35 40 45
0
1
2
3
4
TX-RX Distance (mm)
Vrect (V)
0.01
0.1
1
Peak PTE (%)
Operating Distance > 40mm
(Vrect > 1.4V)
0.6mT
0.09mT
Fig. 4. Measurements of operating waveforms of the proposed
implant; 3.5V stimulation with varying ME voltage; stimulating
efficiency with various amplitudes; power delivery versus distances.
Progmaming of Multiple Implants with a Si ngle TX
Synchronized Stimulations with Programma ble
Amplitude, Pulse Width and Delay (In-Vitro)
Implant A
ID: 11010111
Implant B
ID: 01111000
Vrect:A
Vrect:B
Stimulation:A Amp. Change:
1V 2V
0.5V 20ms
Fixed Amp.:2V
Stimulation:B
Same Data
Different Data
Stim.:A
Stim.:B
Stim.:A1.2ms
1.2ms
0.6ms
3.25V
1.5V
2.5V
1.5V
Stim.:A
1.2ms
Delay
0.8ms
1V
3.25V
Delay
0.4ms
1.2ms
3.25V
1V
Stim.:B
0.6ms
Bone
3.16
Local Safety Analysis of the 330-kHz
Magnetic Field in COMSOL
In-Vitro Test with Porcine Tissue
Implants with
Testing Pins
Porcine Tissue
Magnetic
Driver
Diameter: 6cm
Turns: 15
Porcine Tissue
Thickness: 2cm 30
X (mm)
10 20 30
10
20
Y (mm)
TX Coil
Lateral & Longitudinal
Operating Distances
40
40
TX Coil
Osciloscope
Power Supply AFG
6cm
Muscle
Skin+Fat
Air
Spinal Cord:
40mm
Cardiac:
60mm
1.00
0.32
0.10
0.032
mT
0.1mT
Simulated
IEEE Standard
(Unrestriced)
Simulated
(Localized)
IEEE Standard
(Unrestricted)
Skin 67.1 0.00913
Fat 56.2 0.15
Muscle 36.6 0.51
Bone 10.6 0.00819
Max. Electric Field (V/m)
69.05
(IEEESTD.
2019.8859679)
Max. SAR (W/kg)
2
(IEEESTD.
2019.8859679)
Stim.:B
Stim.:A
Stim.:B
Fig. 5. Measurement of individual programmability of implants with
a single TX; safety analysis; in-vitro test with porcine tissue and
shmoo plot of operating space; synchronized stimulations and
stimulations with programmable delays.
Hydra Contractions with & without Stimulation Rat EMG Response vs. Stimulation Amplitude
0
200
400
600
Stim.
On
10sec10sec 10sec
0100 200
Hydra A
Hydra B
Spontaneous
Contractions
0100 200
ΔF/F (%)
Time (s)
Hydra
GCaMP7b
Fluorescence
Permanent Magnet (Bias)
Implant
Microscope
Objective
Stereotrode
TX
Coil
0
-1
0
1
2
3
Time (ms)
3V
2V
1V
0.25V
Stimulation
Amplitude:
Averaged EMG
Response (mV)
Bi-phasic Stimulation
(1.2mS pulse width,
3Hz frequency)
EMG of
Plantar Muscles
Sciatic Nerve
Standard Deviations)
In-Vivo Validations
This Work ISSCC'16 [7] ISSCC'18 [8] Sci. Rep.'20 [2] ISSCC'20 [9] Nat. BME.'20 [3]
Neural / Cardiac Neural Neural Cardiac Neural Neural
180 180 350 180 350 65
Magnetoelectric Inductive Ind uctive Inductive Inductive
Ultrasonic
Single TX +
Multi. Implants
Electrode
Array
μLED
Array
Two TXs +
Two Implants
Electrode &
μLED Array
N/A
Voltage Current Optical Voltage Optical, Current Current
Amplitude 3.5V, 4b 0.5mA, 7b N/A 3V, Fixed 0.77mA, 5b 0.4mA, 3b
Pulse Width 1.2ms, 4b 8ms 2ms, 2b 0.3ms, 3 levels 0.64ms, 4b Continuous
Frequency Continuous Continuous 10Hz, 2b Continuous 400Hz, 4b Continuous
Delay 0.8ms, 5b No No No No No
PUF ID Multiplexer Multiplexer Differe nt fcarrier Multiplexer N/A
9864 300 320700 4
1 x 0.8 5.7 x 4.4 1 x 1 0.85 x 0.45 5 x 3 1 x 1
6.2 500 12.2 10.1 N/A 1.7
40 N/A 7 60 N/A 55
Implant Volume (mm3)
Max. RX-TX Distance (mm)
Stimulation Mode
Stimulation
Parameters
(Max., Resolution)
SoC Power (μW)
Chip Size (mm2)
Target Application
Process (nm)
Multisite Stimulation
Strategy
5
Fig. 6. Synchronous muscle activations of Hydra in response to
electrical stimulations; EMG response of rat with various stimulation
amplitudes; comparisons with state-of-the-art stimulating systems.
... When we apply a magnetic field to the material, the magnetostrictive material generates a strain that is coupled to the piezoelectric layer that, in turn, generates an electric field [35]. Thus, by applying an alternating magnetic field at the acoustic resonant frequency of the film we can efficiently deliver power to our implant [35,36,38,39]. In addition to delivering power, we can also transmit data to our implant by modulating the frequency of the applied magnetic field. ...
... The frequency shift results in a change in the amplitude of the received voltage, which can be interpreted as a digital bit sequence that specifies the stimulation parameters for the implant [38,39]. Taken together, the complete wireless EVNS system consists of an external magnetic field transmitter, a ME film that harvests power and data from the magnetic field, and a custom integrated circuit that interprets the digital data and generates the electrical stimulus delivered by the electrodes (Fig. 1a). Figure 1b demonstrates a conceptual overview of the system implemented in a large animal model where a surface coil can be used to wirelessly transmit a magnetic field to power and program the implant for endovascular stimulation. ...
... While the implant itself might only move a few millimeters once fixed within the tissue, it is easy to imagine misaligning a wearable transmitter by a centimeter or more, which remains within our alignment tolerances. Furthermore, the use of a wearable transmitter is also possible due to the low magnetic field strengths that are required to activate high voltages in these ME thin films where only ~1 mT field strengths are required for the power densities required to activate neurons through stimulation by the material itself [35] or by powering custom integrated circuits [38,39]. This will allow the technology to be readily translated into the clinic and even permit patients to use implants at a home-setting. ...
Preprint
Full-text available
Implanted bioelectronic devices have the potential to treat disorders that are resistant to traditional pharmacological therapies; however, reaching many therapeutic nerve targets requires invasive surgeries and implantation of centimeter-sized devices. Here we show that it is possible to stimulate peripheral nerves from within blood vessels using a millimeter-sized wireless implant. By directing the stimulating leads through the blood vessels we can target specific nerves that are difficult to reach with traditional surgeries. Furthermore, we demonstrate this endovascular nerve stimulation (EVNS) with a millimeter sized wireless stimulator that can be delivered minimally invasively through a percutaneous catheter which would significantly lower the barrier to entry for neuromodulatory treatment approaches because of the reduced risk. This miniaturization is achieved by using magnetoelectric materials to efficiently deliver data and power through tissue to a digitally-programmable 0.8 mm2 CMOS system-on-a-chip. As a proof-of-principle we show wireless stimulation of peripheral nerve targets both directly and from within the blood vessels in rodent and porcine models. The wireless EVNS concept described here provides a path toward minimally invasive bioelectronics where mm-sized implants combined with endovascular stimulation enable access to a number of nerve targets without open surgery or implantation of battery-powered pulse generators.
... When we apply a magnetic field to the material, the magnetostrictive material generates a strain that is coupled to the piezoelectric layer that, in turn, generates an electric field 37 . Thus, by applying an alternating magnetic field at the acoustic resonant frequency of the film, we can efficiently deliver power to our implant [37][38][39]41,42 . In addition to delivering power, we can also transmit data to our implant by modulating the frequency of the applied magnetic field. ...
... In addition to delivering power, we can also transmit data to our implant by modulating the frequency of the applied magnetic field. The frequency shift results in a change in the amplitude of the received voltage, which can be interpreted as a digital bit sequence that specifies the stimulation parameters for the implant 41,42 . Taken together, the complete wireless EVNS system consists of an external magnetic field transmitter, an ME film that harvests power and data from the magnetic field, and a custom integrated circuit (IC) that interprets the digital data and generates the electrical stimulus delivered by the electrodes (Fig. 1a). Figure 1b shows a conceptual overview of the system implemented in a large animal model where a surface coil can be used to wirelessly transmit a magnetic field to power and programme the implant for endovascular stimulation. ...
... Furthermore, the use of a wearable transmitter is also possible due to the low magnetic field strengths that are required to activate high voltages in these ME thin films. These thin films require only <1 mT field strengths for the power densities needed to activate neurons through stimulation by the material itself 37 or by powering custom integrated circuits 41,42 . This will allow the technology to be readily translated into the clinic and may even permit patients to use implants at a home setting. ...
Article
Full-text available
Implantable bioelectronic devices for the simulation of peripheral nerves could be used to treat disorders that are resistant to traditional pharmacological therapies. However, for many nerve targets, this requires invasive surgeries and the implantation of bulky devices (about a few centimetres in at least one dimension). Here we report the design and in vivo proof-of-concept testing of an endovascular wireless and battery-free millimetric implant for the stimulation of specific peripheral nerves that are difficult to reach via traditional surgeries. The device can be delivered through a percutaneous catheter and leverages magnetoelectric materials to receive data and power through tissue via a digitally programmable 1 mm × 0.8 mm system-on-a-chip. Implantation of the device directly on top of the sciatic nerve in rats and near a femoral artery in pigs (with a stimulation lead introduced into a blood vessel through a catheter) allowed for wireless stimulation of the animals’ sciatic and femoral nerves. Minimally invasive magnetoelectric implants may allow for the stimulation of nerves without the need for open surgery or the implantation of battery-powered pulse generators.
... This paper is an extended version of [13], with more comprehensive qualitative and quantitative analysis of ME power transfer to multiple devices, detailed discussions on circuit implementations, and additional experimental results on the system robustness. The rest of the paper is organized as follows: Section II presents the design of the proposed system; Section III gives detailed implementations of the implant SoC; Section IV shows the experimental results, including functional validations, in-vitro tests and in-vivo experiments [16]; measured efficiency of the ME samples are also added to the curve. ...
Preprint
This article presents a hardware platform including stimulating implants wirelessly powered and controlled by a shared transmitter (TX) for coordinated leadless multisite stimulation. The adopted novel single-TX, multiple-implant structure can flexibly deploy stimuli, improve system efficiency, easily scale stimulating channel quantity, and relieve efforts in device synchronization. In the proposed system, a wireless link leveraging magnetoelectric (ME) effect is co-designed with a robust and efficient system-on-chip (SoC) to enable reliable operation and individual programming of every implant. Each implant integrates a 0.8-mm2 chip, a 6-mm2 ME film, and an energy storage capacitor within a 6.2-mm3 size. ME power transfer is capable of safely transmitting milliwatt power to devices placed several centimeters away from the TX coil, maintaining good efficiency with size constraints, and tolerating 60 degree, 1.5-cm misalignment in angular and lateral movement. The SoC robustly operates with 2-V source amplitude variations that spans a 40-mm TX-implant distance change, realizes individual addressability through physical unclonable function (PUF) IDs, and achieves 90% efficiency for 1.5-3.5-V stimulation with fully programmable stimulation parameters.
... For example, a mmsized ME film activated using a low amplitude and lowfrequency magnetic field has been used to successfully provide therapeutic deep brain stimulation in a freely moving rodent model for Parkinson's disease [19]. Exploiting the CMOS technology, the first programmable neural implant leveraging the ME effect (MagNI) has been proposed in [20], [21], [25]. 1 2 3 4 5 6 7 8 9 10 11 12 13 14 15 16 17 18 19 20 21 22 23 24 25 26 27 28 29 30 31 32 33 34 35 36 37 38 39 40 41 42 43 44 45 46 47 48 49 50 51 52 53 54 55 56 57 58 59 60 A c c e p t e d M a n u s c r i p t ...
Article
Full-text available
Abstract-Objective.Compared to biomedical devices with implanted batteries, wirelessly powered technologies can be longer-lasting, less invasive, safer, and can be miniaturized to access difficult-to-reach areas of the body. Magnetic fields are an attractive wireless power transfer (WPT) modality for such bioelectronic applications because they suffer negligible absorption and reflection in biological tissues. However, current solutions using magnetic fields for mm-sized implants either operate at high frequencies (>500 kHz) or require high magnetic field strengths (>10 mT), which restricts the amount of power that can be transferred safely through tissue and limits the development of wearable power transmitter systems. Magnetoelectric (ME) materials have recently been shown to provide a wireless power solution for mm-sized neural stimulators. These ME transducers convert low magnitude (<1 mT) and low-frequency (∼300 kH) magnetic fields into electric fields that can power custom integrated circuits or stimulate nearby tissue. Approach: Here we demonstrate a battery-powered wearable magnetic field generator that can power a miniaturized MagnetoElectric-powered Bio ImplanT "ME-BIT" that functions as a neural stimulator. The wearable transmitter weighs less than 0.5 lbs and has an approximate battery life of 37 hr. Main results: We demonstrate the ability to power a millimeter-sized prototype "ME-BIT" at a distance of 4 cm with enough energy to electrically stimulate a rat sciatic nerve. We also find that the system performs well under translational misalignment and identify safe operating ranges according to the specific absorption rate limits set by the IEEE Std 95.1-2019. Significance: These results validate the feasibility of a wearable system that can power miniaturized magnetoelectric implants that can be used for different neuromodulation applications.
Article
This article presents a hardware platform including stimulating implants wirelessly powered and controlled by a shared transmitter (TX) for coordinated leadless multisite stimulation. The adopted novel single-TX, multiple-implant structure can flexibly deploy stimuli, improve system efficiency, easily scale stimulating channel quantity, and relieve efforts in device synchronization. In the proposed system, a wireless link leveraging magnetoelectric (ME) effect is co-designed with a robust and efficient system-on-chip (SoC) to enable reliable operation and individual programming of every implant. Each implant integrates a 0.8-mm² chip, a 6-mm² ME film, and an energy storage capacitor within a 6.2-mm³ size. ME power transfer is capable of safely transmitting milliwatt power to devices placed several centimeters away from the TX coil, maintaining good efficiency with size constraints, and tolerating 60°, 1.5-cm misalignment in angular and lateral movement. The SoC robustly operates with 2-V source amplitude variations that spans a 40-mm TX-implant distance change, realizes individual addressability through physical unclonable function (PUF) IDs, and achieves 90% efficiency for 1.5-3.5-V stimulation with fully programmable stimulation parameters.
Article
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Clinically approved neural stimulators are limited by battery requirements, as well as by their large size compared with the stimulation targets. Here, we describe a wireless, leadless and battery-free implantable neural stimulator that is 1.7 mm3 and that incorporates a piezoceramic transducer, an energy-storage capacitor and an integrated circuit. An ultrasonic link and a hand-held external transceiver provide the stimulator with power and bidirectional communication. The stimulation protocols were wirelessly encoded on the fly, reducing power consumption and on-chip memory, and enabling protocol complexity with a high temporal resolution and low-latency feedback. Uplink data indicating whether stimulation occurs are encoded by the stimulator through backscatter modulation and are demodulated at the external transceiver. When embedded in ex vivo porcine tissue, the integrated circuit efficiently harvested ultrasonic power, decoded downlink data for the stimulation parameters and generated current-controlled stimulation pulses. When cuff-mounted and acutely implanted onto the sciatic nerve of anaesthetized rats, the device conferred repeatable stimulation across a range of physiological responses. The miniaturized neural stimulator may facilitate closed-loop neurostimulation for therapeutic interventions. A wireless and battery-free 1.7-mm3 neural stimulator implanted onto the sciatic nerve of rats allows for repeatable stimulation across a range of physiological responses.
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About 30% of patients with impaired cardiac function have ventricular dyssynchrony and seek cardiac resynchronization therapy (CRT). In this study, we demonstrate synchronized biventricular (BiV) pacing in a leadless fashion by implementing miniaturized and wirelessly powered pacemakers. With their flexible form factors, two pacemakers were implanted epicardially on the right and left ventricles of a porcine model and were inductively powered at 13.56 MHz and 40.68 MHz industrial, scientific, and medical (ISM) bands, respectively. The power consumption of these pacemakers is reduced to µW-level by a novel integrated circuit design, which considerably extends the maximum operating distance. Leadless BiV pacing is demonstrated for the first time in both open-chest and closed-chest porcine settings. The clinical outcomes associated with different interventricular delays are verified through electrophysiologic and hemodynamic responses. The closed-chest pacing only requires the external source power of 0.3 W and 0.8 W at 13.56 MHz and 40.68 MHz, respectively, which leads to specific absorption rates (SARs) 2–3 orders of magnitude lower than the safety regulation limit. This work serves as a basis for future wirelessly powered leadless pacemakers that address various cardiac resynchronization challenges.
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This paper presents the first wireless and programmable neural stimulator leveraging magnetoelectric (ME) effects for power and data transfer. Thanks to low tissue absorption, low misalignment sensitivity and high power transfer efficiency, the ME effect enables safe delivery of high power levels (a few milliwatts) at low resonant frequencies ( $\sim$ 250 kHz) to mm-sized implants deep inside the body (30-mm depth). The presented MagNI (Magnetoelectric Neural Implant) consists of a 1.5-mm $^2$ 180-nm CMOS chip, an in-house built 4 × 2 mm ME film, an energy storage capacitor, and on-board electrodes on a flexible polyimide substrate with a total volume of 8.2 mm $^3$ . The chip with a power consumption of 23.7 $\mu$ W includes robust system control and data recovery mechanisms under source amplitude variations (1-V variation tolerance). The system delivers fully-programmable bi-phasic current-controlled stimulation with patterns covering 0.05-to-1.5-mA amplitude, 64-to-512- $\mu$ s pulse width, and 0-to-200-Hz repetition frequency for neurostimulation.
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A major challenge for miniature bioelectronics is wireless power delivery deep inside the body. Electromagnetic or ultrasound waves suffer from absorption and impedance mismatches at biological interfaces. On the other hand, magnetic fields do not suffer these losses, which has led to magnetically powered bioelectronic implants based on induction or magnetothermal effects. However, these approaches have yet to produce a miniature stimulator that operates at clinically relevant high frequencies. Here, we show that an alternative wireless power method based on magnetoelectric (ME) materials enables miniature magnetically powered neural stimulators that operate up to clinically relevant frequencies in excess of 100 Hz. We demonstrate that wireless ME stimulators provide therapeutic deep brain stimulation in a freely moving rodent model for Parkinson's disease and that these devices can be miniaturized to millimeter-scale and fully implanted. These results suggest that ME materials are an excellent candidate to enable miniature bioelectronics for clinical and research applications.
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A power-efficient wireless stimulating system for a head-mounted deep brain stimulator (DBS) is presented. A new adaptive rectifier generates a variable DC supply voltage from a constant AC power carrier utilizing phase control feedback, while achieving high AC-DC power conversion efficiency (PCE) through active synchronous switching. A current-controlled stimulator adopts closed-loop supply control to automatically adjust the stimulation compliance voltage by detecting stimulation site potentials through a voltage readout channel, and improve the stimulation efficiency. The stimulator also utilizes closed-loop active charge balancing to maintain the residual charge at each site within a safe limit, while receiving the stimulation parameters wirelessly from the amplitude-shift-keyed power carrier. A 4-ch wireless stimulating system prototype was fabricated in a 0.5-μm 3M2P standard CMOS process, occupying 2.25 mm2. With 5 V peak AC input at 2 MHz, the adaptive rectifier provides an adjustable DC output between 2.5 V and 4.6 V at 2.8 mA loading, resulting in measured PCE of 72 ~ 87%. The adaptive supply control increases the stimulation efficiency up to 30% higher than a fixed supply voltage to 58 ~ 68%. The prototype wireless stimulating system was verified in vitro.