Advances in Bioelectromagnetics for Implantable Systems

Carlos J. Cela 1, Anil K. RamRakhyani 1, Sundar Srinivas 1,Gerard Hayes 2, Michael Dickey 2, Gianluca Lazzi 1

1University of Utah, Salt Lake City, UT, 84112, USA. 2North Carolina State University,Raleigh, NC, 27695, USA.

Abstract— Implantable medical devices are used to sup-plement, monitor, or replace physiological function for avariety of medical conditions. In this article we highlight aset of topics that we believe could eventually be instrumentalin bringing the next generation of implantable devices toexistence. These include multi-scale numerical modeling andsimulation, magnetic neural interfaces, high-performancetelemetry, and the use and application of flexible antennas.

Index Terms— Numerical simulation, Magnetic stimula-tion,Biomedical telemetry, Microelectronic implants, Anten-nas

I. INTRODUCTION

Some common requirements for the design and opera-tion of biomedical implantable devices include a thoroughunderstanding of the bioelectromagnetic interactions andneural response, interfacing the electronics with the humanbody for purposes of stimulation or recording, provideelectrical power to the implant, and the need to transmitinformation to and from the implant. The sections belowdescribe some of the current collaborative efforts by theauthors in these key areas.

II. BIOELECTROMAGNETIC MULTI-SCALE MODELINGFOR PREDICTION OF NEURAL RESPONSE

Numerical methods have been used for decades to solvebioelectromagnetic problems at different temporal andspatial scales. Most work done to date focus on the solutionof bioelectromagnetic problem as a particular scale [1],[2]. At the cellular spatial scale, different electric circuitanalogs to model the bioelectric dynamics of cellularmembrane and single cells have been described [3], [4],while at the bulk tissue scale (i.e. without cellular structuredifferentiation), numerical formulations have been used todetermine electric and magnetic field components insidetissue caused by external or internal sources [5], [6], [7].Implantable neural interfaces have components belongingto different spatial and temporal scales: the dynamics ofa single synapse, the behavior of a neural cell axon, thegeometry of the stimulating and recording electrodes differorders of magnitude in size and time granularity with eachother. It is interesting to make the analysis of these differ-ent scales converge to a single method, because this allowsfor a seamless model of the entire bioelectrical behavior.We argue that a suitable domain for this convergence to

take place is circuit network theory; this is so because theunderlying theory is well understood, differential equationscan be represented using equivalent circuits, and there exista wealth of tools and methods available for numericaltreatment of circuit networks. Of the many numericalformulations available for bioelectromagnetic problems atthe tissue scale, there are two methods, the admittancemethod [8] and the impedance method [2], that representthe problem domain in terms of an equivalent electricalcircuit. The conversion of the model to a network circuitequivalent is different depending on the spatial scale,but in all cases involves the characterization of localelectrical, chemical, and/or thermal behavior by replacing aportion of the model with lumped circuital elements. In thecase of cellular modeling, the electrochemical behavior ofthe citomembrane is modeled using a repeating circuitalpattern representing a membrane patch [9]. In the caseof bulk tissue, the mass to be modeled is subdivided insmall volume cells (voxels) considered homogeneous, andthe equivalent electric network is derived from the knowndielectric properties [10] of the constitutive tissue of eachvoxel [7]. Adaptive meshing techniques can be used toreduce the size of the model while keeping a relativelysmall error compared to homogeneous meshing [2], [7].Once the model is described in terms of an equivalentcircuit network, the external stimulation sources are added,and tools from circuit theory are used to form a linearsystem and to solve it. The result is usually expressedin terms of the electric field and current density at everypoint inside the model. Once the system is solved, neuralresponse prediction can be estimated using a variety ofcriteria, including the calculating the activation function[1], or using known threshold values [11].

III. MAGNETIC NEURAL INTERFACES

Neural stimulation for medical implants is commonlyrealized by means of localized charge injection usingmetallic electrodes or electrode arrays. This presents sev-eral drawbacks, the potential need for excessive chargedensity to achieve stimulation with contact electrode arrayswhen electrode size is small, and the lack of tolerance withrespect to imperfect contact between the electrode contactsand the neural tissue. In addition, the dynamics of theelectrochemical interaction between electrode and tissue

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is complex, and depends on frequency, current density,temperature, electrode composition, etc [12], [13].

Our research group is investigating an alternative ap-proach, by developing a new class of microcoils and micro-coil arrays to be used as neural microstimulators in neuralinterfases. These coils generate time-varying magneticfields, and the mechanisms of neural stimulation are basedon eddy currents and their gradients of the magneticallyinduced electric fields. There are several advantages inusing a magnetic field for interfacing: coils do not needdirect contact with the tissue to be stimulated, so theycan be completely insulated, removing the requirementof biocompatibility for the metal in the windings, andpreventing the possibility of electrochemical reactions inthe vecinity of the implant. Further, coils can potentiallyoffer a larger number of options to control the shape ofthe induced magnetic fields (and therefore eddy currents)compared to traditional stimulators, and their operationwill not be affected by contact capacitances.

Methods to increase the induced electric fields andother electromagnetic quantities such as using ferrite cores,ferrite slabs, figure of eight coils etc. are being investi-gated. The theoretical and numerical results predicted thepossibility of stimulating neural tissues using smaller coilsembedded with ferrite cores and low voltages to drive thecoils. The coil sizes were reduced to 1 cm in diameterand the drive voltage was reduced to 40 V, compared to1kV for transcranial magnetic stimulation [14], [15]. Thenumerical predictions have been experimentally verified ona rat sciatic nerve in-vivo.

IV. HIGH-PERFORMANCE TELEMETRY

In the context of this article, we use the term telemetry inthe general sense of wireless interfacing technologies thatenable bidirectional data transfer between the implanteddevice and external electronics while simultaneously pow-ering the implanted device. Based on the direction of thedata transfer, wireless links are referred to as uplinks (fromthe driver to the load circuit) and downlinks (from the loadto the driver circuit). In general, the available data rate onthe link strongly depends on the carrier frequency, systembandwidth and data modulation scheme. Due to simplicityof the design, ASK (Amplitude Shift Keying) and OOK(ON-OFF Keying) based modulation schemes are popularfor low data rate communication. To achieve higher datarates, BPSK (Binary Phase Shift Keying) is used [16].Downlink communication can be implemented using LSK(Load Shift Keying) [16] or similar schemes, in whichthe effective load of the link (i.e. power consumed by thedevice) is modulated with the data signal and detectedat the driver side. For purposes of power transfer, thepower transfer efficiency depends on the size, structure,relative location, and environment surrounding of the coils.Conventional two-coil based inductive links are extensively

Fig. 1. Experimental prototypes of coils used for magneticneural interfases.

studied in the literature [17], [18], and power transferefficiency is formulated as in Equation 1. In general, thestrategy to achieve an efficient wireless power transferis to design the coils so they resonate at the inductivelink operating frequency. In these systems, power transferefficiency is a strong function of the quality factor (Q-factor) of these coils and the coupling between the transmitand receive coils. Due to the finite nature of the source andload resistance, Q-factors of the coils are limited to smallvalues (10 to 30) [19], [17], which restricts the maximumachievable power transfer efficiency of any given two-coilsystem.

η =k2QdQl

1 + k2QdQl= ηdl (1)

A. Multi-coil Approach

Multiple-coil based power delivery [20] is an alterna-tivetechnique used to decouple adverse effects of sourceand load resistance from the coils, and in this way achievehigh Q-factor on coils. Multi-coil system typically usesmore than two coils for power transfer. Figure 2 showsa magnetically coupled four-coil system. The externalpower source is connected to the driver coil and implantedelectronics is connected to the load coil. Transmitter andreceiver coils are resonating high-Q passive coils. At the

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operating frequency of the system, and considering highvalues for the magnetic couplings k1 and k3, and a high Q-factor for transmitter and receiver coils, the power transferefficiency of a four-coil power transfer system can beapproximated as Equation 2 [20].

Driver coil

Transmittercoil

Receivercoil

Loadcoil

12

3

Fig. 2. Block diagram of four-coil based telemetry system. Loadand receiver coils are implanted. Driver and transmitter coil arepart of the external device.

η ' k21QdQt

1 + k21QdQt

k22QtQr

1 + k22QtQr

k23QrQl

1 + k23QrQl

= ηdtηtrηrl (2)

From Equation2, it can be seen that the effect of lowvalue of k2 (coupling between external and implantedcoils) have been compensated by the high Q-factor trans-mitter and/or receiver coils.

V. FLEXIBLE ANTENNAS AND COILS

There is a growing interest in the development offlexible antenna technologies to satisfy diverse applicationsincluding implantable medical devices [21], [22]. Formedical implants, in addition to the typical antenna perfor-mance parameters (e.g. efficiency, bandwidth), the physicalparameters (e.g. size, weight, dynamic reliability). Thematerials of construction typically define the mechanicalperformance of the antenna and provide a critical designconsideration. We present a brief overview of the currenttrends of flexible antenna systems. The systems are dividedinto three categories based on their material compositions:

1) Thin conductors and substrates2) Flexible conductors and substrates3) Stretchable conductors and substrates

A. Thin conductors and substrates

Rigid materials can be rendered flexible by making themthin in at least one dimension; an every day example isaluminum foil. Most of the initial flexible antenna systemsconsisted of thin sheets of etched copper on thin substrates,such as polyimide and PCB materials (such as 0.008 G10or FR-4). Mechanically, these antennas can be curved inone dimension or folded in multiple dimensions. Elec-trically, these antennas perform well. The conductor and

substrate losses are minimized through the use of copperand low-loss dielectric substrates. However, the use of rigidmetals (such as etched copper or stamped stainless steel)limits the dynamic flexing capabilities of the antennas,and subjecting the antennas to repetitive flexure couldfatigue and crack the conductive elements. Furthermore,this construction does not allow for the antenna system tostretch.

B. Flexible conductors and substrates

The flexibility of rigid conductors (e.g., copper) can beimproved by making them thin in two dimensions; that is,using wires, threads, and fabrics. Relative to thin films,these geometries have an improved degree of freedomfor flexibility and can be arranged to facilitate stretching(e.g. using a weave pattern or a coil geometry). In arepresentative application, a gold wire could be embeddedin a PDMS substrate to form the antenna. Conductivethread assemblies consist of dielectric fibers (such asyarns and glass fibers) with small interwoven metallicwires. Electrically, these antennas perform well relativeto conventional microstrip elements. The substrate lossesare minimized through the use of low-loss dielectric sub-strates. Mechanically, these antenna systems can conformto complex, irregular surfaces (such as clothing). Althoughthese antennas are more flexible than a rigid conductorconstruction, the metallic components (thin wires or coatedplastics) are still vulnerable to fatigue and wear over time.

C. Stretchable conductors and substrates

To overcome limitations with respect to stretchability,researchers have identified two types of structures: thin-film metallization of the antenna conductor on elastomericsubstrates [23] and the formation of liquid metal radiatingelements in micro-fluidic channels within an elastomericsubstrate [24], [25]. It has been shown [23] that thinmetal films (such as <100nm thick gold films) depositedon elastomeric substrates (such as PDMS) can stretchreversibly without losing electrical conduction. The pres-ence of dense arrays of micro-cracks in the substrateallow for the thin film to twist and flex when strainedup to 100% strain without losing conductivity. Up to 20%elongation has been reported with acceptable electrical per-formance after hundreds of thousands of cycles. However,when subjected to excessive stress, micro-cracks can formwithin the gold film that result in higher resistance andlower antenna efficiency. In the absence of micro-cracks,these antennas perform well electrically. The conductorand substrate losses are minimized through the use ofgold and low-loss dielectric substrates. Fluidic antennasrepresent an emerging group of antenna systems thatare extremely flexible, highly-stretchable, and maintainacceptable electrical performance. These antennas fullyadopt the mechanical properties of the encasing substratematerial (such as PDMS) and are therefore flexible and

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mechanically durable. Previously reported liquid metal an-tenna configurations include single-layer dipoles and loopsand multi-layered patches and coils. They are well suitedfor a broad range of applications, including telemetryand biomedical applications. Unlike solid metal structuresor metallized surfaces, liquid metal conductors do notcrack or fatigue during stretching. The metal is typicallycomposed of gallium or gallium alloys. Gallium has low-toxicity, negligible vapor pressure, and a conductivity thatis approximately an order of magnitude lower than copper.It is a low viscosity fluid, which allows it to flow inresponse to stress to maintain electrical continuity [26].Importantly, its surface forms a thin oxide film rapidlyin the presence of air that helps stabilize mechanically themetal. The metal can be shaped into the form of an antennaby injecting it into channels with the desired shape. Theresulting antennas adopt the mechanical properties of theencasing substrate since the conductive elements are liquid.Therefore, mechanically, these antenna systems conform tothe widest array of complex shapes and irregular surfaces.The ability to stretch an antenna offers a simple route tochanging its shape, and therefore, function. Reconfigurablefluidic antennas (through the dynamic control of the flowof the liquid metal through various microchannel struc-tures) have been demonstrated recently [27].

VI. CONCLUSION

Advances in bioelectromagnetics for implantable sys-tems are currently driven by the availability of new mate-rials and technologies, including commonplace computingprocessing power, which help to bring to life ideas hatshort time ago were not realisable. The topics presentedare currently focus of extensive research efforts by ourgroup, and represent the state of the art in the field.

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