FEATURE ARTICLE
Neural Interfaces
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A Materials Roadmap to Functional Neural Interface Design
Steven M. Wellman, James R. Eles, Kip A. Ludwig, John P. Seymour, Nicholas J. Michelson,
William E. McFadden, Alberto L. Vazquez, and Takashi D. Y. Kozai*
Clinical scientists have also applied knowledge gained from basic neuroscience
studies to develop interfaces with the
nervous system for therapeutic or assistive purposes in patients with injury and
disease.[2–4] For example, electrical stimulation using implantable neural interfaces–or neuromodulation devices—have
received Food and Drug Administration
(FDA) market approval for the treatment
of a diverse set of conditions including
epilepsy, depression, Parkinson’s Disease,
sleep apnea, blindness, deafness, obesity,
urinary and fecal incontinence, and hypertension. Patients have also used similar
devices to detect and direct brain signals
to control robotic limbs, bypassing their
injured or degenerated spinal cord.[3,4]
While these clinical successes highlight
the potential of these technologies, concerns over large variability in therapeutic/
assistive efficacy, long-term reliability, and
health risks[5–8] prevent these devices from
reaching their full potential. This article
reviews the recent advances and continued
materials based challenges of this rapidly
expanding field.
Advancements in neurotechnologies for electrophysiology, neurochemical
sensing, neuromodulation, and optogenetics are revolutionizing scientific
understanding of the brain while enabling treatments and preventative
measures for a variety of neurological disorders. The grand challenge in
neural interface engineering is to seamlessly integrate the interface between
neurobiology and engineered technology to record from and modulate neurons over chronic timescales. However, the biological inflammatory response
to implants, neural degeneration, and long-term material stability diminishes
the quality of the interface overtime. Recent advances in functional materials are aimed at engineering solutions for chronic neural interfaces, yet,
the development and deployment of neural interfaces designed from novel
materials have introduced new challenges that have been largely unaddressed. Many engineering efforts that solely focus on optimizing individual
probe design parameters, such as softness or flexibility, downplay critical
multidimensional interactions between different physical properties of the
device that contribute to overall performance and biocompatibility. Moreover,
the use of these new materials present substantial new difficulties that must
be addressed before regulatory approval for use in human patients is achievable. In this review, the interdependence of different electrode components is
highlighted to demonstrate the current material-based challenges facing the
field of neural interface engineering.
1. Introduction
1.1. Neural Signals
Implantable neural interfaces are important tools for capturing and modulating the sophisticated computations of the
nervous system. This technology has seen an explosion in
research, innovation, and potential applications. In order to
better understand plastic changes in neural networks, these
interface components must maintain stability over the long
time periods associated with memory formation and learning.[1]
S. M. Wellman, J. R. Eles, N. J. Michelson, W. E. McFadden,
Prof. A. L. Vazquez, Prof. T. D. Y. Kozai
Department of Bioengineering
Center for the Basis of Neural Cognition
McGowan Institute of Regenerative Medicine
NeuroTech Center
University of Pittsburgh Brain Institute
Center for Neuroscience at the University of Pittsburgh
University of Pittsburgh
208 Center for Biotechnology
300 Technology Drive, Pittsburgh, PA 15219, USA
E-mail: tdk18@pitt.edu
The success of an implanted neural interface is contingent
upon the quality of the signals that the interface can detect from
the nervous system, as well as the reliability and precision with
which the interface can modulate the nervous system. Extracellular neural recordings can be used to isolate individual neuronal action potentials (also known as “single-units” or spikes),
Dr. K. A. Ludwig
Department of Neurologic Surgery
200 First St. SW, Rochester, MN 55905, USA
Dr. J. P. Seymour
Electrical and Computer Engineering
1301 Beal Ave., 2227 EECS, Ann Arbor, MI 48109, USA
The ORCID identification number(s) for the author(s) of this article
can be found under https://doi.org/10.1002/adfm.201701269.
DOI: 10.1002/adfm.201701269
Adv. Funct. Mater. 2017, 1701269
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which represent the most basic code in neuron communication. Alternatively, extracellular recordings that do not isolate
individual action potentials are called multiunit recordings,
which arise when two or more single-units from different
nearby neurons are recorded simultaneously. Finally, local field
potentials (LFPs) are generated by the summation of neural
activity from a large population of neurons within a spatial
region. LFPs are often comprised of low frequency oscillations
(0.1–120 Hz) that are less sensitive to spatial interference.
For neuroprosthetics applications, single-unit and multiunit
recordings are often preferred over LFPs, due to the specificity
of the information conveyed, but all have demonstrated useful
read-out information.
Although neuronal signal transduction is electrical in nature,
it is propagated at each axon/dendrite terminal, called a synapse, by the release of neurotransmitters. Unlike action potentials and LFPs, neural chemical sensing electrodes are specific
to certain electroactive neurotransmitters such as dopamine or
adenosine. These biosensors rely on a technique called fastscan cyclic voltammetry (FSCV) to measure the concentration
of molecules based on their oxidation and reduction potentials. While these devices do not operate in the same temporal
domain as traditional electrodes, their sensitivity to neurotransmitter-specific signaling is critical for assessing neuronal-type
specific activity. Neurochemical recordings utilizing FSCV
electrodes have been obtained in human patients by multiple
groups,[9,10] and have been proposed as a fundamental human
research tool as well as a critical sensing component for closedloop neuromodulation systems.[11,12] When combined with traditional electrophysiological recording modalities, FSCV neurochemical measurements can be instrumental in understanding
the complex multimodal therapeutic effects and limiting side
effects from neuromodulation therapies to treat psychiatric
disorders.
In addition to sensing electrical current or neurochemicals,
electrode-based neural interfaces can also be used to introduce
signal into the nervous system by injecting current into local
tissue (0.1 to >10 000 Hz). Electrode designs can vary dramatically depending upon the anatomical location and therapeutic
application for stimulation. For example, simple bipolar or
tripolar electrode configurations are typically used to stimulate
the vagus nerve for epilepsy and depression, while indwelling
microcortical arrays with hundreds of electrode sites are used
to provide sensory feedback for a prosthetic.
Recent advances in genetics have expanded the neural interface toolset beyond electrode-based technologies. Optogenetics
combines genetics with light-sensitive proteins to instill membrane channels or activity reporters in selective cell populations. Genetic targeting provides a level of control not possible
with traditional electrode technology and is attained using
several approaches. A common approach is to rely on specific
promoters that regulate the expression of genes in cells targeted by the promoter. Major constraints for optogenetics are
the requirement for genetic manipulation, the limited availability of promoters, and the potential variability of genetic
transcription. Many of the biological challenges for optogenetics are due to poor control over the number of transgenes
inserted into the genome, gene insertion loci, promoter
enhancers/inhibitors/silencers activity based on insertion
Adv. Funct. Mater. 2017, 1701269
Kip A. Ludwig is the associate director of the Mayo
Clinic Neural Engineering
Laboratories in Rochester,
Minnesota. He received
his Ph.D. in Biomedical
Engineering from the
University of Michigan. Prior
to Mayo he led translational
device programs at the
National Institutes of Health
under the SPARC, BRAIN,
and CREATE Devices Initiatives. He also developed a minimally invasive neuromodulation therapy that is for sale in
seven countries and in a U.S. Pivotal Trial.
John P. Seymour’s research
is in the area of advanced
neurotechnology with a focus
on circuit mapping tools.
His research has addressed
reduced tissue reactivity and
improved optical, electrical,
and mechanical performance. He earned his B.S.
with Honors in Engineering
Physics from Ohio State
University and his M.S. and
Ph.D. in Biomedical Engineering from the University of
Michigan. His industry experience includes working at
NeuroNexus as a principal scientist. He now holds a faculty research position of assistant research scientist in the
Department of Electrical Engineering at the University of
Michigan developing novel neural interface systems.
Takashi D. Y. Kozai is an
assistant professor in the
Bioengineering Department at
the University of Pittsburgh.
He received his Ph.D. in
Biomedical Engineering
from the University of
Michigan, and his BAs from
the University of Colorado at
Boulder in molecular, cellular,
and developmental biology
and biochemistry. His primary
research interest is in seamlessly interfacing neurobiology
and neurotechnology through multimodal analysis, including
in vivo multiphoton imaging, functional in vivo electrophysiology, and biomaterials engineering. He is also interested in
elucidating molecular and cellular pathways of brain injuries
and diseases as well as developing intervention strategies.
loci, and for virus optogenetics, trade-offs between diffusion
radius and transduction rates.[13] In addition, due to biophysics
and protein kinetic limitations, high sensitivity opsins will
have slower channel kinetics while fast opsins will have lower
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sensitivity.[14] For the purpose of this work, the focus will be on
the challenges in delivering and extracting light over the desired
spatial resolution with implantable devices (Section 2.15). However, it is critical to understand that optogenetic control will be
an ongoing challenge for both biological engineering and materials engineering of neural interfaces.
In general, neural interfaces intended to stimulate nervous
tissue or record neural activity from the nervous system face
a similar material based challenge. The fundamental goal is
to enable a higher number of input/output channels from the
nervous system, while only creating a minimal disruption of
tissue. Avoiding tissue damage is important both for the longterm patency of the interface as well as for minimizing safety
risks to the patient. As will be explained in subsequent sections, this fundamental trade-off motivates a number of interdependent design choices with interactions that are often not
considered.
1.2. Current Challenges
Current challenges with chronically implantable neural interfaces can largely be categorized into performance reliability
and variability issues. Invasive surgical procedures to implant
chronic neural interfaces carry inherent risks, such as brain
infection, surgery related hemorrhaging, blood clots, edema,
etc. As such these interfaces would ideally be designed to
minimize tissue trauma, remain both functional and biocompatible for the life-time of the patient to avoid corrective
surgery or reimplantation, and be removable in the event of
unforeseen complications. Across all neural implant classes,
there is an average decline of the signal-to-noise ratio (SNR)
in the communication between the nervous system and neural
interfaces over time. The resulting decline in neural interface
reliability is believed to be, in part, due to the inflammatory
tissue response triggered by the surgical implantation procedure. This initial reaction can progress into a chronic immune
response, ultimately turning into a foreign body response and
even leading to migration of the tissue or implant.[7,15] While
the details of the biological challenges with chronic neural
interfaces have been discussed elsewhere,[8] this reactive tissue
response degrades the electrical characteristics of the interface
by forming a high impedance encapsulation sheath in conjunction with neural degeneration. This response decreases
the amplitude of the detected neural signals by increasing the
minimum distance from the electrode to the nearest neuron
firing action potentials, thereby increasing the applied current
necessary for exciting nearby neurons, and limiting the precision of activation via electrical stimulation.[16,17] A number of
acute and delayed stab wound studies show that the tissue
can recover to some extent when the device is removed,[17–20]
suggesting that certain physical properties of chronically
implanted neural interfaces are responsible for this tissue
reaction.[20]
The focus on minimizing the foreign body response has
increased the emphasis on studying advanced materials to overcome specific design parameters hypothesized to attenuate the
inflammatory tissue response (see reviews[8,21–28]). However,
large variability in histology and recording performance[8,29–34]
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have made it difficult to disentangle contributions from the
many physical properties of these devices to the myriad of
observed multimodal failure modes.[5,8,35–37] These complex
failure modes produce histological outcomes that, in isolation, are poor predictors of actual performance outcomes as
measured by brain-machine neural communication with high
degrees of freedom and SNR.[34,35] Similarly small amounts of
platinum particulates have been shown to break off from the
electrode during chronic stimulation resulting in a foreign
body response, but multiple studies have demonstrated the
presences of these particulates post-mortem without causing
clinically significant safety concerns or deterioration of performance.[38–41] Current literature often highlights the advances of
individual material properties and therefore has a limited discussion of the broader obstacles limiting the advancement of
the field. While excellent reviews can be found on advancing
individual design properties[21–28,42] or expanding upon the biological tissue reaction,[8] this review aims to highlight the interdependence of multiple design and material properties as well
as to guide the successful of balancing and titrating of these
properties. Through this multidimensional cost-benefit analysis, engineers can design functional neural interfaces capable
of obtaining selective, high-fidelity, long-lasting readouts of
brain activity. Material selection, the design and assembly of the
materials, and strategies for surgical delivery of the implantable components will likely provide critical breakthroughs in
achieving improved device performance.
2. Material Considerations for Engineering
The simplest electrode design consists of a conductive component to transfer electrical charge, insulation to spatially constrain neuronal signals, and a connector that provides external
interfacing with signal processing devices. However, there are
numerous intrinsic material properties (conductivity, flexibility,
material strength, and chemical stability) as well as extrinsic
design parameters (geometry, density, insertion, and packaging) that must be considered to achieve long-term biocompatibility and functionality. These individual intrinsic and extrinsic
properties are interdependent on one another; optimizing one
can negatively impact performance or lead to loss of functionality. The goal of this review is to introduce a number of the key
material and design properties as well as their complex interdependence on one another. Understanding these engineering
challenges are critical for the development of devices with longterm functionality.
2.1. Electrical Property/Electrode Traces
One of the most important components of any electrode technology is the conductor, which forms the electrical connection,
also known as electrical trace, that carries current (Figure 1a).
Every electrode must include a conductive element which transfers detected signal from the tissue to an integrated wireless
transducer or external electrical device, such as a headcap for
cortical recording electrodes, or transmits current to the electrode site from a percutaneous connector or implanted pulse
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Figure 1. Basic Electrode Anatomy. a) Randles circuit representation of the electrode, which includes the electrode trace (red dash), insulation (green
dash), and electrode site (blue dash: Relectrode and Celectrode) where neural potentials (Vneuron) are recorded from. Note: Postelectrode equipment is
not shown. b) Cross-section of a microthread electrode[43] showing the carbon fiber core and Parylene insulation. c) Layer-by-layer (LbL) assembled
composite electrodes using d) Au nanoparticles or e) carbon nanotubes. Reproduced with permission.[44] Copyright 2012, American Chemical Society.
generator. The conductance (G) is the material property that
describes the ease with which charges can move within it
G=
1
A
=σ
R
l
R=ρ
(1)
l
A
(2)
where R is the resistance of the conductor, A is the cross-sectional area of the conductor, l is the length, σ is the electrical
conductivity in siemens per meter (S·m−1), and ρ is the electrical resistivity or specific electrical resistance (Ω·m).
As shown in Equation (1) and Equation (2), the dimensions
of the material can play a crucial role in determining electrical
properties. With the emergence of nanoscale and angstrom
diameter wires, it is important to consider how the size of the
electrical trace can constrain the maximum signal or stimulation
current that can safely be passed through the device. The minimum cross-sectional area (A) is limited by the heating power (P,
Joule heating, also known as Ohmic or resistive heating), especially for electrical stimulation and chemical sensing electrodes
P = I 2R = I 2 ρ
l
A
(3)
where I is electrical current. At some finite A the electric current will produce enough heat that it might burn/oxidize or
Adv. Funct. Mater. 2017, 1701269
melt the conductor, reducing its conductivity. Given all of these
considerations, specific materials and the method of deposition have unique advantages and disadvantages that should be
considered when choosing an electrical trace conductor for particular applications.
Metals, particularly noble metals, are perhaps the oldest
and most common electrical conductors used for neural electrodes.[45,46] Their high electrical conductivity originates from
the overlapping valence and conductance bands of metal atoms,
allowing for relatively free movement of electrons through the
solid “electron sea” lattice.[47] Selection of metal conductors
requires careful consideration of conductivity, mechanical stability for handling, chemical stability (namely, corrosion resistance), and biocompatibility. In general, heavy transition metals
such as gold (Z = 79), platinum (Z = 78), and iridium (Z = 77),
as well as tungsten (Z = 74) or tantalum (Z = 73; for short-term
applications) display a number of these desirable properties.
The noble metals and alloys can generally remain relatively
chemically inert over the chronic timescales that would affect
other conductor classes.
Advances in microfabrication and photolithography techniques enabling well controlled and reproducible batch fabrication has led to the use of semiconductors and polycrystalline
materials which are more compatible with these microfabrication processes. Semiconductors are materials with small bandgaps (between the valence band and conductance band) that can
be made conductive by adding electron acceptors (p-dopants) or
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electron donors (n-dopants) impurities. Polycrystalline materials are composed of small crystalline regions separated by
grain boundaries.[48] These grain sizes dramatically impact the
conductivity, making them more conductive than their amorphous material counterparts.[49–51] Polycrystalline silicon (polysilicon) is a highly pure, polycrystalline form of silicon that is
easily fabricated into microelectrodes and is compatible with
complementary metal-oxide-semiconductor (CMOS)[52] and
micro-electromechanical systems (MEMS)[53] technology. Additionally, the conductive properties of polysilicon can be modulated by varying processing conditions or dopant densities.[51]
For these reasons, polysilicon is a popular material used in
microfabricated silicon arrays. While silicon technology can be
more brittle than metal, it offers a greater design space than
traditional microwires.
Unlike inorganic semiconductors, conductive polymers
(CPs) are organic polymers that conduct current.[54–59] CPs
are composed of monomeric compounds linked in chains
of alternating single and double bonds and doped with a stabilizing counter-ion.[60] Conjugated polymers have narrower
band gaps, allowing electrons to move easily between the conducting band and valence band. Dopants increase the conductivity of polymers by either removing electrons from the
valence band (p-doping) or adding electrons to the conduction
band (n-doping). The most common biocompatible conductive
polymers used for neural interfaces are polypyrrole (PPy),[61]
poly(ethylene dioxythiophene) (PEDOT),[61] and polyaniline,[62]
with PEDOT being the most electrochemically stable.[28] Most
conductive polymers, however, do not contain intrinsic charge
carriers and must be doped with a counter ion to reduce electrical resistance and heating, and improve conductance. Therefore, they require counter ions that are partially oxidizing
(p-doping) electron acceptors (such as LiClO4, I2, or AsF3) or
alternatively partially reducing (n-doping) electron donors
such as Na or K. While CPs may be either chemically or electrochemically synthesized,[63,64] careful consideration of the
counter-ion’s charge or polarity (zwitterions), and concentration
is necessary for electrochemical deposition.[28] Conductive polymers generally have lower conductivity than most metals, but
other advantageous properties make them popular for coating
electrode sites (see Section 2.4).[65] However, recent efforts are
aimed at developing conductive polymer composite wires with
improved mechanical properties.[66–68] Currently, considerable
challenges remain due the brittleness of deposited CPs which
complicate the fabrication of electrodes comprised entirely
of CPs.
Carbon-based materials (graphene, carbon fiber, carbon
nanotubes) are also being developed as the conductive components of electrodes. Some carbon allotropes are highly conductive due to the carbon–carbon pi bonds. Carbon-based
materials also have useful mechanical properties such as high
tensile strength and modulus. Single carbon fibers (CF) are
implemented similarly to metal microwires[69–73] (Figure 1b).
Graphene, a hexagonal lattice monolayer of carbon,[74] has
exceptional flexibility, surface area, and conductivity.[75] Graphene and graphene oxide (GO) have been implemented as
bulk conductors,[75] coatings,[76] and dopants.[77,78] The flat
structure of graphene makes it ideal for patterning planar
implants.[79–81] Recently developed high aspect ratio all-carbon
Adv. Funct. Mater. 2017, 1701269
crumpled graphene transistors demonstrate high sensitivity
and improved spatial resolution, with no observed increase in
source–drain resistance at up to 56% compressive strain.[82] In
contrast, carbon nanotubes (CNTs) are cylindrically wrapped
graphene sheets. Single-walled carbon nanotubes are comprised of a single layer and multiwalled carbon nanotubes have
multiple layers. These CNTs are emerging as novel substrates
for a variety of neural engineering applications due to their high
mechanical strength and electrical properties.[83–85] CNTs also
can have a range of electrical properties based on the orientation in which the graphene is wrapped in a tube. This property
is represented by a pair of indices (n,m), and it modulates the
CNTs’ conductive properties: from metallic, to quasi-metallic,
or semiconductive[86,87] (see review[28]). Due to challenges of
fabricating long CNTs, they are generally employed for the
electrode site materials (see Section 2.4),[85] however, there
has been some effort in developing and patterning CNT containing composite conductors, such as through layer-by-layer
deposition methods for neural interfaces (Figure 1c–e).[28,88–91]
For applications where stretchability rather than flexibility is
required, CNT-doped copolymer matrices can achieve up to
130% strain.[92] These composite conductors are engineered
by bringing conductive particles close enough for electrons to
jump from one particle to another, known as percolation networks. The small spaces between each particle can absorb more
strain before breakage, improving the mechanical properties of
the device. However, the tradeoff is that these small spaces lead
to increased resistivity compared to solid metals.
To summarize, ideal conductors should achieve exceptional
conductivity, long-term functional durability, electrochemical
stability, low cross-sectional area, and the ability to withstand
stress and strain forces. As devices become smaller and vary in
their geometries, materials that can be handled easily during
microfabrication are preferred since their conductive properties can be altered with processing. Metals wires are traditional
materials with good conductivity with limited stretchability and
the ability to assemble into large arrays. Some other conductors
trade compatibility with microfabrication techniques and
increased design space for assembly into arrays with increased
brittleness. Composite materials generally trade decreased
conductivity with increased stretchability and/or durability.
Often, the decreased conductivity needs to be compensated
by increasing the cross-sectional surface area. As such, for
developing functional implants, it is critically important to
balance all of these material properties as well as those in the
following sections.
2.2. Insulation Materials
The other critical component of chronically implantable microelectrode technology is the dielectric layer, or insulation layer.
Some commonly employed insulating materials include varnish, epoxy, parylene, glass, Teflon, polyimide, fused silica,
silicon oxide, silicon nitride, or silicone elastomers.[28,93] Dielectric materials generally have a large bandgap between the conductive and valence band, which blocks electrical conduction
at low voltages. While the functional component of any electrode is the conductive element, insulating dielectric materials
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are equally critical to preventing parasitic capacitance from
allowing the signal to leak into the tissue or into adjacent conductors. The insulating material property is defined primarily
by its resistivity (ρ) and the dielectric constant (εr) or the relative
static permittivity. For simplicity one may consider the effect of
a dielectric material in a parallel plate capacitor (C)
C = ε rε 0
A
d
(4)
where ε0 is the electrical constant (ε0 ≈ 8.854 × 10−12 F⋅m−1),
d is the thickness of the insulation layer, and A is the crosssectional area of the insulation layer along the electrical trace.
Insulating materials are characterized as having a relatively low
εr, that is, the ratio of the capacitance of a dielectric material
to that of a vacuum which tends to 1. The shunt capacitance
(Cshunt, Figure 1a) is the capacitance across the insulation layer
from the center of the electrode conductor to the conductive
electrolyte bath conductor, which results in an inversely frequency dependent impedance across the electrode, behaving as
a low pass RC filter with frequency fc[94–96]
Vout-shunt = Vin
fc =
1
1 + ω 2R 2C 2
1
ωRC
(5)
(6)
where ω is the angular frequency of the waveform (rad s−1).
With respect to Equations (3) and (4), materials with high
dielectric constants or low dielectric thickness lead to high
capacitance, coupling of electrical signal to adjacent traces, and
attenuating the recorded signal (Vout-shunt). At very thin dielectric thicknesses, the insulation becomes less effective due to
electron tunneling.
For high density arrays, current density accumulation in a wire
or electrode can cause a change in current density in adjacent
wires through capacitive charging with the two wires acting as
the capacitor (Ccross-talk: aka, cross-talk capacitance, capacitive
coupling or mutual capacitance, Figure 1a). In multichannel
electrode arrays, cross-talk is caused by capacitive coupling of
one circuit (the “aggressor”) to another (the “victim”)
∆Vvictim =
Cadj
1
⋅
× ∆Vaggressor
C grd + Cadj Ragressor (C grd−a + Cadj )
1 +
R victim (C grd−v + Cadj )
(7)
where ∆Vvictim is the potential change induced by the “aggressor”
circuit onto the “victim” circuit, ∆Vaggressor is the potential
change in the “aggressor circuit,” Cgrd is the total capacitance
relative to ground, Cadj describes the capacitor formed between
the circuits, Ragressor and Rvictim are the resistive loads of the
coupled circuit on the conductors respectively, and Cgrd−a and
Cgrd−v are the capacitances between the aggressor and ground
and the victim to ground respectively. Equivalent circuit models
developed for electrodes implanted into tissue show that crosstalk increases the effective amplifier impedance (comprised of
the head-stage impedance and shunt capacitance) above the
Adv. Funct. Mater. 2017, 1701269
effective electrode impedance (comprised of the resistance of
the electrolyte solution as well as resistance and capacitance of
the electrode) with increasing signal frequency.[95,96] As a result,
crosstalk can attenuate recorded signals, cause false signal in
adjacent electrodes during recording, or result in unwanted
accumulation of charge density between adjacent electrodes
during stimulation.[97,98] This effect is especially prominent
in small multielectrode arrays in which resolution is important, as signal in one wire can affect signals in nearby wires.
Inert materials or materials with low dielectric constants, such
as ceramics or polyimide, reduce crosstalk compared to other
poor insulating materials such as silicon, a more conductive
substrate.[99] While the source of cross-talk is traditionally in
the interconnects and packaging, it becomes an increasing
important issue as engineers attempt to design higher density
arrays.[100]
A number of polymer insulating materials have become popular due to the ability to uniformly and reproducibly coat complex microscale shapes, lower elastic modulus, attain chronic
stability in physiological environments, and achieve compatibility with other microfabrication techniques for selectively controlled de-insulation of active sites.[101] These include parylene
and polyimide variants where the side chains (or R groups)
alter the material’s dielectric, mechanical, and sometimes even
bioactive properties.[101–103] Careful selection or combination
of functional R groups can also be used to functionalize bioactive surfaces (see Section 2.14). Polydimethylsiloxane (PDMS)
is another insulator for neural tissue interfaces due to its biocompatibility.[104] However, effective PDMS-based dielectrics
for implants require substantial thickness due to its high water
absorption rate, which increases its dielectric constant and
shunt capacitance, and decreases the cutoff frequency. Other
polymers with very high water absorption cannot be used as
insulation materials for chronic implants. For example, many
epoxies are unsuitable for use as insulators for chronic electrodes because they absorb water, negatively impacting dielectric properties as well as allowing for corrosion conditions in
devices utilizing dissimilar metal connections within the insulation, especially under stimulation.[105,106]
Another property that requires consideration, especially
for electrical stimulation, is the material’s dielectric strength,
which is a measure of the field strength a material can withstand before electric breakdown occurs. At high voltages (on the
order of 106 V cm−1, the limit increases with decreasing relative
permittivity and thickness) dielectric materials “break-down”
and become conductive.[107] Dielectric strength, in contrast
to the dielectric constant, decreases with surface defects (pinhole defects; Figure 2a), increases with temperature, increases
with signal frequency, and can decrease with water absorption.
Careful consideration of the dielectric property, water absorption, and insulation thickness are important for chronic neural
interfaces. Exceeding these voltages will usually also lead to
other failure modes such as gross degradation or melting.[108]
However, for chronic electrical stimulation, lower DC voltages
that are either deliberately induced to power active components
on the electrode, or naturally arise during stimulation can exacerbate failure of the insulation, especially at points of defects.
For example, the increase in time to failure for an Al2O3 and
parylene bilayer is, in some situations, 13 times faster when
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Figure 2. Insulation Failures. a) Strong stimulation currents lead to insulation failure that start as pin hole defects (red arrow) in the insulation (cyan).
Complete Insulation failure (green arrow) can be identified over regions of PEDOT/PSS electrochemical deposition through the insulation (cyan).
b) Many dielectric polymers such as parylene are susceptible to cracking (magenta arrow) and delamination (cyan arrow). Reproduced with permission.[111] Copyright 2012, IOP Publishing. c) Planar arrays made from crystalline polymers form triple junctions (red dashed circle) when implanted.
The two crystalline layers cannot be perfectly aligned due to entropy (blue/white angles). d) Ions and water molecules in the tissue thermodynamically
degrade the interface of the two crystalline layers a crack forms. e) Cracks thermodynamically propagate until insulation fails. Note: this is not an
issue if the triple junction crack propagation failure rate is slower than the lifetime of the patient. f) Self-healing smart polymers with self-annealing
functional groups and reseal delamination. g) Insulation adhesion layers (e.g., Au) are sometimes used to reduce insulation delamination failures.
h) In the tissue, the wire (W), exposed adhesion layers (Au), and the tissue environment creates a galvanic cell, and accelerates corrosion. Reproduced
with permission.[112] Copyright 2011, Elsevier.
a 5 V bias is applied.[109] Also, thin dielectric films risk the possibility of forming a simple low-pass RC circuit. An RC low-pass
filter allows frequencies from 0 Hz (DC) to its cutoff frequency
to pass while attenuating higher frequencies, which can contain
critical single-unit information. In this case, a negative capacitance amplifier can help correct for the various capacitive impedances introduced by the conductor/insulator capacitors.[94,110]
Dielectrics are essential in ensuring electrical current travels
unimpeded through the conductor preventing any loss in electrical quality through shunting or attenuation. As a result, an
ideal insulator would be thin and have a low dielectric constant,
high dielectric strength, achieve virtually no capacitance, have
minimal potential for water absorption, and simultaneously
thick enough to occlude fluid ingress. These optimal properties
would allow small electrical traces to be packed close together
while limiting parasitic capacitance and cross-talk, leading
to limited signal loss or attenuation. In addition, the material
must be flexible (low Young’s modulus) and remain durable
under chronic cycling. Since a wide range of materials fit into
the insulator class, other mechanical and adhesion properties
must also be considered for applicability to specific electrode
designs and functionality requirements.
2.3. Adhesion of the Insulator
During microelectrode production, the insulator is typically
either deposited uniformly over the entire conductive material,
Adv. Funct. Mater. 2017, 1701269
as in the case of fabricated microelectrodes, or hand-made
using backfill in the case of neuromodulatory devices. Active
electrode sites are formed by removing insulation in a controlled fashion or by controlling the wire length that protrudes
from the insulation sleeve. Exposure methods for microwires,
such as mechanical impact, cutting with a razor, pyrolysis
(flame), electric arc, plasma arc, and laser can lead to large
variability in site size, impedance, and recording quality.[113–117]
These methods can also create collateral damage to the surrounding insulation layer, such as cracking and delamination
(Figure 2b). Cracking and delamination lead to electrode failure
as the geometric area of the exposed conductor becomes too
large.[118] In turn, this averages the recorded potentials across
a larger electrode surface area and attenuates the neural signal
(see Section 2.4). This underscores the importance of carefully
consider the biocompatibility of the underlying materials. For
example, many semiconductors (i.e., gallium or arsenide) are
toxic in vivo.[119] While adhesion layer coatings such as Au and
Ti have been employed to limit delamination, this can lead to
rapid corrosion of the electrode site material (Figure 2g,h; see
Section 2.5).[120]
A much more controlled approach for exposing the electrode site or depositing the same electrode site is through
lithographic techniques, such as through the use of photoresist
masks.[121] However, unlike microwires that have conformal
coatings of insulation, planar microelectrodes are generally
created by sandwiching conductors between two insulation
materials. With silicon technology, low pressure chemical vapor
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deposition of silicon oxide is used to conformally insulate the
polycrystalline silicon traces,[122] and provides one of the highest
quality SiO2 films available. However, thin films of silicon oxide
have a tendency to bow from compressive internal stresses.[123]
For this reason, an additional silicon nitride dielectric layer or
silicon nitride followed by another silicon oxide dielectric layer
are used to balance the bowing and provide additional biostability compared to a single SiO2 layer.[124,125]
When two crystalline materials are used to sandwich electrical traces and then are implanted into tissue, a triple junction
forms (Figure 2c) often leading to delamination (Figure 2d,e).
The thermodynamics of crack propagation at triple-junction
sites between crystalline materials has been well studied,
especially in electrolytic solutions such as the cerebral spinal
fluid.[35,126–128] One approach to reducing delamination is
through thermal annealing, a method that utilizes heat and
pressure to improve the seal between two polymer layers.[129]
However, this has been shown to impact device performance
and sensitivity by influencing bulk polymer properties.[129,130]
Polymers that can withstand temperatures near their melting
points for longer periods of time such as amorphous (noncrystalline) polymers may be another alternative to forming better
thermally annealed insulators for avoiding triple-junctions.
Another approach to prevent crack propagation is through the
use of “smart” self-healing materials that can respond to incidences of damage.[131] For example, polymers can be functionalized with reactive chemical species,[132–144] two polymer layers
with appropriate functional groups (CONH2 and COOH)
have been shown to self-anneal into imine bonds in vitro that
chemically seal two layers together over time without the need
for thermal annealing (Figure 2f).[145] In the event of material
degradation or delamination, the exposed functional groups
may reform covalently through thermodynamic reactions.
However, further investigations are required to determine the
self-healing potential of these functionalized polymers in vivo.
Unlike planer probes that involve polymer–polymer delamination failure, designs such as microwires and bed of needle
arrays benefit from being coated with a single conformal layer
of insulation. However, these devices still suffer from delamination and cracking at the conductor–insulator interface near their
recording sites, especially if the conductor microactuates during
stimulation.[36,118,146,147] Careful material selection and deposition methods are critical for long-term functional performance.
2.4. Electrode Size and Materials
Electrode site materials and site sizes are important aspects that
impact the sensitivity and performance capabilities of neural
microelectrodes. Electrode site designs need to carefully consider the electric field strength to selectively attenuate signals
resulting from more distant neurons
Vmonopole =
Vdipole =
1 q
⋅
4 πε 0 r
1 qdcos (θ )
⋅
4 πε 0
r2
Adv. Funct. Mater. 2017, 1701269
(8)
(9)
where q is the magnitude of the action potential, r is the distance from the source, d is the thickness of the cell membrane,
ε0 is the permittivity constant of the tissue, and θ, the angle
between the current source and the electrode, is equal to 0. In
reality, extracellular action potential characterization depends
on the distance between the electrode and the cell, and the
radius and dendritic density of the cell.[148] In general, empirical measurements indicate that the voltage drop-off is between
that of Vmonopole and Vdipole.[149–152] This voltage can be expressed
simply by
n
1
I
⋅ xi
4
r
πσ
i
e
i =1
V =∑
( where 1 < x < 2)
(10)
where σe is the conductivity of the extracellular space, n is the
number of current point sources which the electrode is exposed
to, and x is 1 < x < 2 depending on the size, geometry and type
of neuron. Small contact sites (<<1000 µm2) have the advantage
of greater selectivity, and can potentially distinguish between
multiple single-units with higher signal amplitudes. This is
because only a small number of neurons will be within close
proximity to small recording sites and detected voltage of distant neurons decay rapidly with r (Figure 3a). However, small
electrode sites are impaired by increased impedances and
thermal noise.[150,153] On the other hand, lowering the impedance by increasing the recording site surface area leads to attenuation of the signal amplitude (Figure 3a,b).[150,154] Large sites
can be modeled as small sites in parallel with increasing r with
distant parallel elements detecting very low V. Voltage elements
in parallel are calculated as the average of all parallel elements.
As a result, the fall-off of the electric field generated by a nearby
neuron is steep; larger electrodes tend to average strong signal
from the nearby neuron with weak signals that are more distant, effectively decreasing the signal recorded. Therefore, it is
important to select electrode site materials that lower impedance without increasing the geometric site size. For single-unit
recordings, the ideal electrode has the lowest possible surface
area and impedance.
Stimulation electrodes require additional material considerations. The electrode site material and stimulation amplitude
on stimulating electrodes determine whether the electrode
is capacitive (non-Faradaic), involving electrostatic or electrolytic interactions between the electrode–electrolyte interface,
or Faradaic, which involves redox reactions between the electrode and surrounding medium.[94,158] Non-Faradic electrodes
stimulate neurons through the diffusion of ions by charging
and discharging the capacitive layer that forms at the electrode
surface, which is considerably safer than Faradic charge injection methods since the electrode material is chemically stable.
Non-noble metals, such as titanium, tantalum, and stainless
steel metals inject charge mainly through this double layer
capacitance.[159] However, there are charge transfer limitations compared to Faradaic electrode materials.[160,161] Irreversible Faradic interactions occur when oxidized ions diffuse away from the electrode before they can be reduced and
redeposited onto the electrode site, which is called electrode
dissolution and can have harmful effects on the surrounding
tissue.[162] This is particularly true with metals that generate
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Figure 3. Electrode site size and material properties influence recording and stimulating performance. a) A small contact site surface area b) compared
to a larger contact site surface area records larger voltage amplitudes (VT) from a single point source (a neuron) provided by the inverse square law of
electric field strength. Note: I2 and I3 are negative. c–f) Charge transfer during electrical stimulation with varying geometric surface areas at the electrode
interface. Increasing surface area from (c) to (d) increases the capacitance (Ce2 > Ce1) and halves the resistance (Re1 > Re2). e) Creating a pore introduces
a resistance (Rp’) with a cross-sectional area of A and length l. Increasing pore length from (e) to (f) begins to impact ion movement within the electrodetissue interface and attenuates electrode performance. g–j) SEM images of electrode site materials: g) PEDOT/CNT, scale bar 1 µm, h) PEDOT/GO, scale
bar 1 µm, i) PEDOT/PSS, scale bar 1 µm, and j) IrOx, scale bar 500 nm. g,i Reproduced with permission.[155] Copyright, 2016, IEEE. h Reproduced with
permission.[77] Copyright 2013, the Royal Society of Chemistry. j Reproduced with permission.[156] Copyright 2010, Elsevier. Distribution of charge density
representing the edge effect for k) a planar high-surface area electrode and l) circular electrode. Adapted with permission from.[157] Copyright 2009,
Frontiers Media SA. m) Recording site electrochemically deposited with polypyrrole (PPy). The octagonal recording site geometry allowed the polymer
to grow into eight teardrop segments with an absence of coating from the center. c–f,m) Reproduced with permission.[28] Copyright 2014, Springer.
Adv. Funct. Mater. 2017, 1701269
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low molecular weight ions during stimulation. Noble metals
such as platinum (Pt), iridium (Ir), palladium (Pd), and rhodium (Rh) are used during neural stimulation due to their corrosion resistance.[163,164] While platinum and iridium have limitations on charge injection capacity,[165,166] platinum–iridium
alloys improve mechanical and electrochemical properties,
making them suitable as implantable interfaces. Another concern during tissue stimulation is the formation of hydrogen
and oxygen gas through the electrolysis of water. The potential
range that occurs before the reduction of water (H2 formation)
or oxidation of water (O2 formation) is defined as the water
window. Platinum electrodes have reported charge injection
limits between 50 and 150 µC cm−2 using biphasic 0.2 ms
pulses, safely within limits of the water window.[167] However,
platinum dissolution has been shown to occur at as little as
20 µC cm−2 so alternative parameters during stimulation must
be carefully considered.[168]
On the other hand, pseudocapacitive electrodes are made
from electrode materials that transfer charge through Faradaic modes, but some of these metal ions can be reversed
back onto the electrode before the ion diffuses away by rapidly applying the opposite charge. These electrode materials are sometimes considered ideal stimulating electrode
materials because their method of charge injection involves
reversible redox reactions that avoid the harmful dissolution
of metal ions but have high charge transfer capabilities. For
example, iridium oxide (IrO2) electrodes are used for their
mixed conductivity and reversible Ir3+/Ir4+ redox reactions
that permit charge injection capacities of about 0.5–8 mC
cm−2.[169,170] Sputtered iridium oxide films, electrodeposited
iridium oxide films, and activated iridium oxide films generate the same iridium oxide as an electrode material, but
with widely different surface morphologies.[171–173] Other
common materials for pseudocapacitors include transition metal oxides, nitrides, and sulfides such as manganese
(MnO2), iron oxide (Fe3O4), titanium nitride (TiN), and titanium sulfide (TiS2).[174–180]
While many safe stimulation electrodes exist, they are limited in their ability to spatially and selectively activate small
populations of neurons.[181] Ideally, a stimulation electrode
will have a very small surface area such that it minimizes
activation of nearby axons and neurons to a locally confined
area.[28,182–184] However, attempting to increase specificity by
lowering the geometric area results in higher charge densities
(Q/A) that can damage tissue.[185–187] To address tissue safety
issues, most stimulating electrodes are designed to have large
surface areas. One way to safely alter the electrical properties
of both recording and stimulating electrodes is to increase
the electrochemical surface area (ESA) relative to the geometric surface area (GSA) (Figure 3b–d). The ESA of a probe
involves introducing micro and nanoscaled features that alter
the roughness and porosity of the surface of the electrode,
improving charge transfer with the surrounding electrolyte
solution. Likewise, increasing ESA lowers impedance,[188] so
charges of equivalent amplitude can be injected with lower
electrode potentials, decreasing the risk of tissue damage.
However, there is an upper limit to increasing the ESA, since
at some point ions diffusion through the pore become negatively impacted (Figure 3e). For a sinusoidal AC waveform, the
Adv. Funct. Mater. 2017, 1701269
penetration depth λ into the pore of a rough surface can be
modeled as
λ=
1 κr
2 c dω
(11)
where r is the pore radius (cm), κ is the electrolyte conductivity (Ω−1 cm−1), cd is the surface capacitance (F cm−2), and ω
is the angular frequency of the waveform (rad s−1). For a given
frequency ω, there will be a maximum penetration depth λ at
which increasing the surface area beyond this distance will not
influence charge transfer, due to resistance to ionic current as
a result of increasing pore length (Figure 3e).[189] Likewise in
chronic implants, resistance to ion diffusion across the coating
can be exacerbated in vivo as reactive tissue and cells block
pores across the surface of the electrode, likely overriding any
beneficial enhancements over time from increasing the electrochemical surface area.[190] Lastly, it may be worth considering
how the geometry of the electrode site leads to an “edge effect,”
where charges are more concentrated on the edges of an electrode during stimulation due to repelling like charges,[157,191]
leading to preferential corrosion at the edges of the electrode
site (Figure 3j,k).[192] Therefore, even when relatively safe stimulation parameters are followed, irreversible redox reaction may
still occur at the edges which may lead to corrosion of the electrode, damage of the tissue, as well as facilitate the delamination of the electrode sites from around its edges.
One method of increasing the ESA:GSA ratio to reduce
impedance is with CNT, graphene, CP coatings, or other nanostructured surfaces (Figure 3f–j).[28,83,85,121,193] While these materials generally have lower impedances per electrochemical surface area compared to many metals, they are easier to deposit
with high ESA. For stimulation, CPs require more careful consideration of the counter ion.[194,195] Among the CPs, PEDOT
is considered to be the most stable electrically conducting polymer coating that has demonstrated decreased impedance, similar charging injecting capabilities to iridium oxide electrodes,
biocompatibility with tissue, and long-term mechanical stability
in vivo.[28,196,197] Doped CPs have the potential for high storage
capacities. PEDOT doped with paratoluene sulfonate (pTS) can
produce charge storage capacities of about 130 mC cm−2.[198]
The specific dopant and deposition parameters can dramatically
alter the surface morphology of the electrode thus influencing
charge capacities. For example, poly(styrene sulfonate) (PSS)doped PEDOT electrodes display a reduced storage capacity of
about 100 mC cm−2.[198] These polymers can also be chemically
modified to form a plethora of derivatives, each with unique
electrochemical and biological properties(see Section 2.14).[199]
In contrast to stimulating electrodes or electrodes to record
electrophysiological signals, electrodes used to measure electroactive neurochemical concentrations using FSCV are critically dependent on the interaction between the electrode surface and adsorbed electroactive neurochemicals of interest.[200]
FSCV measurements typically consists of two steps: 1) a small
DC holding potential is applied to the electrode to promote
adsorption of neurochemicals of interest to the electrode surface while repelling confounding interferents and 2) the application of an approximately 10 ms triangular voltage pulse that
sweeps from −0.4 to 1.5 V back to −0.4 V (the cathodic/anodic
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limits and sweep rate may vary based off of the neurochemical
of interest and electrode material). The triangle voltage pulse
causes electroactive neurochemicals adsorbed to the surface to
oxidize and reduce at specific potentials, creating brief peaks
in Faradaic current that can be used to identify specific neurochemicals and changes in their concentration. Consequently,
materials used for FSCV materials must (1) enable small electrode designs that minimize biofouling/scarring that would
prevent the adsorption of the neurochemical to the surface,
(2) minimize the number of other electrochemical reactions
that occur at the surface that could confound measurement of
the neurochemical of interest, (3) have minimal drift in open
circuit potential that would generate variability in the true
voltage applied during the FSCV pulse, and (4) be able to both
adsorb the neurochemical of interest while allowing the neurochemical to desorb once oxidized/reduced so that changes in
local concentrations of the neurochemical can be monitored
over time.[201] The ultimate goal is for the FSCV electrode to
have both high sensitivity (detect low concentrations of a
neurochemical) and high selectivity (able to differentiate the
neurochemical of interest from interferents).
The traditional electrode material for FSCV recordings are
small carbon fiber electrodes, which can also be cycled at high
levels to deliberately cause dissolution of carbon at the surface.
This removes biofouling and uncovers new unoccupied sites
for adsorption to enhance signal in a chronic environment.[202]
Numerous alternatives to traditional carbon have been studied
including noble metals like gold, platinum/iridium, iridium
oxide, conductive polymers, boron-doped diamond, graphene,
and pyrolyzed photoresist.[203–211] Boron-doped diamond is a
promising material for chemical sensing due to its good conductivity, chemical stability, and biocompatibility, but it has
a low double-layer capacitance that is not ideal for electrical
stimulation.[212,213] Some of these newer materials like borondoped diamond have yet to demonstrate sensitivity comparable
to carbon fibers. For other new materials the inherent drift in
electrochemical reactions at the surface of the material over
time ultimately limit sensitivity, or the inability of the neurochemicals of interest to desorb after oxidation/reduction once
adsorbed dramatically decrease sensitivity over time. FSCV
waveforms and subsequent analysis to maximize sensitivity
and selectivity utilizing a small carbon fiber electrode was
optimized over the course of decades, while the same degree
of optimization for other material candidates has yet to be
performed. Of the recently explored new materials, pyrolyzed
photoresist, which behaves the most like traditional carbon but
is more amenable to modern multielectrode fabrication processes, is arguably the most promising.
2.5. Chemical Stability of Materials
Developing chronically implantable devices require careful
selection or development of materials and assemblies that outlast the lifetime of the subject. When the adhesion layer (e.g., Ti
underneath Au or Ir) or the interface of the conductor and electrode site material becomes exposed to the tissue (due to dielectric cracking, delamination, electrode site dissolution,[214,215]
or general preparation), the tissue fluid and the two separate
Adv. Funct. Mater. 2017, 1701269
metals form a short circuit galvanic cell that continuously corrodes the electrode material and degrades the electrical characteristics of the microelectrode. Gold-plated tungsten wires,
for instance, have been shown to corrode readily in phosphatebuffered saline (Figure 2g,h).[112,111] Not only does this negatively impact the electrical properties of the electrode, but it
can also lead to toxic metal ion generation.[216] Silver electrodes,
for example, have been shown to rapidly oxidize in the body,
leading to large implant mass loss and rampant tissue inflammation.[217–219] Corrosion in tissue commonly occurs between
two metals, metals of varying purity, or a heterogeneous alloy
surface, where one metal dissolves and deposits onto the surface of the other metal through ion migration facilitated by the
electrolyte solution.[220] This can be avoided through homogenously alloying multiple metal elements together. Corrosion
resistant, higher atomic weight transition metals with good biocompatibility, such as platinum or iridium, are chosen as preferred electrode materials.[21] In addition to improving corrosion
resistance, alloys combine the mechanical and electrochemical
properties of two metals, allowing for tunable mechanical and
electrical properties.[221–224] For example, pure platinum is too
soft for use as a cortical penetrating microwire, so it is commonly implemented in an alloy with the stiffer metal iridium,
thus resulting in a stiffer, lower impedance electrode.[223,224]
Even without metal adhesion layers, some metals can be
prone to delamination and corrosion in physiological environments. Platinum black,[225] iridium oxide,[226] and nanoporous
materials[227] can all be used to increase the ESA of the probe.
However, careful selection of material selection, site size, and
deposition parameters are necessary as these coatings are often
prone to delamination or electrode dissolution.[214,215,228] For
example, platinum black suffers from low mechanical stability
and poor adhesion to the electrode substrate,[229] while iridium
oxide films become more prone to degradation and instability
during electrical stimulation as electrode site sizes becomes
smaller and charge densities increase.[214]
Electrodeposition of CPs can also control corrosion. PEDOT
is especially promising due to its stability in oxygen-rich,
hydrated environments,[61] since oxidation further polymerizes
the CP rather than corroding the metal.[28] Likewise, undoped
PPy coatings have been shown to impart corrosion resistance
on steel.[230] However, while CP coatings offer electrochemical
benefits and a wide array of options for functionalization, they
have difficulty maintaining lamination[231] and are brittle.[232]
Therefore, when using CP coatings, it is also important to consider the underlying conductor material. Care must be taken to
ensure that the CP adheres well to the electrode site material to
prevent delamination. Adhesion can also be improved by appropriate dopant selection or by surface premodification such as
roughening.[233] For example, addition of self-assembled alkylsilane monolayers improved PPy adhesion.[234] The size of the
dopant molecule is an important factor to consider. For example,
PEDOT doped with large PSS (M.W. 100,000) had an increased
tendency to delaminate due to smoother electrode morphology
when compared to doping with smaller pTS (M.W. 171.19)
and perchlorate (ClO4; M.W. 99.45).[235] However, small dopants
can leach out of the CP matrix eventually leading to a decrease
in conductivity and long-term instability in vivo.[236] On the
other hand, larger dopants such as PSS (M.W. 70,000)[212]
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typically remain trapped within the CP matrix, and therefore
do not contribute to a loss in conductivity.[237] Unsurprisingly,
appropriate selection of the conductive polymers (monomers
and/or oligomers) used for electrochemical deposition is also
critically important for the success of the electrode. Although,
PPy is commonly used, due to its ease of growth and low toxicity, PEDOT exhibits superior chemical and electrochemical
stability. Coating thickness is also an important consideration, with thicker PEDOT coatings displaying lower impedance measurements on average compared to thinner coatings,
but are more prone to delamination.[196] It has been suggested
that rough-patterned surfaces offer better adhesion of polymers
to the substrate than smooth surfaces.[233,235,238] Furthermore,
the applied potentials required for electrochemical deposition can cause material dependent corrosion which can disturb the integrity of CP coating, where oxidation of the metal
can block electron transfer resulting in a nonuniform, patchy
coverage.[239]
The risk of corrosion can also vary depending on whether
the electrode is used for stimulation or recording: excessive
electrode polarization during stimulation has been shown to
produce irreversible electrochemical reactions that can degrade
metals (such as stainless steel[240]) which are biologically stable
otherwise. This has led to the exploration of conductive polymers and CNT composite materials for recording and electrical stimulation electrode materials.[84,155,241–245] CNT doping
improves the charge injection capacity through several mechanisms, including increased coating surface area, CNT electrical
conductivity, and the additional charge transfer mechanism
of small cationic molecules entering the coating to serve as
counter-ions against the trapped anionic CNTs when the CP
is reduced. These attributes provide the coating with substantial non-Faradaic (capacitive) and Faradaic character in both
oxidized and reduced states.[246] Mechanical stability tests following both acute and chronic stimulation, as well as chronic
soaking, revealed PEDOT doped with multiwalled nanotubes
(PEDOT/MWNT) and layer-by-layer CNT composites exhibited
none of the delamination or cracking that are typical of PEDOT
coatings or IrOx sites.[84,241,247] Interestingly, PEDOT doped
with carboxylic acid functionalized CNT maintained recording
quality despite increases in impedance over 3–8 weeks post
implant.[248] However, careful consideration of size and concentration are necessary since ejected nanoparticles can cause frustrated phagocytosis, cell cycle arrest, apoptosis, and membrane
lysis.[249,250]
Lastly, the biological response to the implant contributes to
electrode failure in a variety of ways. Initially, biofouling, or protein adsorption on the electrode surface, leads to encapsulation,
which increases the resistance and capacitance of the electrode,
while reducing the electrochemical surface area and attenuating the high frequency content of the electrical stimulus.[237]
This is particularly true for high ESA electrode materials where
the protein and glial cells restrict ionic diffusion.[251] A study for
cardiac pacing electrodes showed that the advantages for having
a high ESA electrode were lost 3–8 weeks after implant unless
antiinflammatory steroids were released from the electrode
site.[252] Biofouling has also been shown to be exacerbated with
stimulating electrodes. For example, IrOx electrode sites used
for chronic in vivo stimulation or CV testing demonstrated
Adv. Funct. Mater. 2017, 1701269
more biological adhesion compared to nonstimulating electrode sites or sites used only for electrochemical impedance
spectroscopy.[226] Besides decreasing electrode performance,
the tissue response also continuously facilitates the degradation of the insulation and electrode site materials. Cracked or
delaminated insulation is further exacerbated by the body heat
and infiltrating immune cells that produce reactive oxygen
species.[5,253,254] Activated macrophages have been speculated
to break down cracked parylene coatings.[101] Efforts toward
reducing the biological factors’ contribution to material degradation and recording failure are aimed at reducing biofouling,
and improving the stability of the underlying materials. For
intracortical electrodes, biofouling can be minimized by modifying the electrode surface properties and minimizing damage
to the blood–brain barrier (BBB) upon electrode insertion.
Various hydrogel and polymer coatings can be applied to electrode surfaces to increase hydrophilicity or incorporate anti-biofouling molecules or drugs. Common anti-biofouling polymers
used in surface functionalization include polyethylene glycol
(PEG) and PEG methacrylate (PEGMA). Boron-doped-diamond
electrodes have also been shown to exhibit reduced biofouling
and more stable electrochemical properties than conventional
TiN electrodes.[43,255] Combining anti-biofouling surface chemistries with image guidance technology to avoid tearing major
blood vessels during insertion may minimize overall inflammation and glial encapsulation.[33] A combination of materials and
design selection will help improve the long-term performance
of implantable interfaces.
2.6. Size and Geometry: Surface Area and Diffusion
Initially, the size and geometry of implants were believed to
play a small role in recording performance and tissue integration when comparing three microelectrode designs with crosssectional areas ranging from 1450–16,900 µm2.[256] However, in
vitro and in vivo studies with polymer fibers show glial attachment significantly decreased when the fiber diameters were
less than 12–5.9 µm, respectively.[257,258] In the brain, it was
shown that 20, 50, and 150 µm2 cross-sectional areas significantly improved neural health and reduced glial encapsulation
compared to 3264 µm2 (Figure 4a–c).[259] Lastly, microelectrodes
with 55.4 µm2 cross-sectional area demonstrated significant
improvement in chronic single-unit recording performance
when compared to traditional devices with a 1875 µm2 crosssectional footprint.[132] This has led to a number of hypotheses
on how size and geometry may be designed to enhance chronic
recording performance.[8,248] For example, devices with holes
or lattice structures might improve recording performance by
allowing trophic factors to diffuse from one side of the probe
to another, which can potentially improve the health of the
tissue around the probe.[8,248] In addition, tissue can regenerate
through the probe holes, which both improve electrical and
chemical signaling between the tissues on either side of the
probe as well as prevent probe migration by providing a soft
anchor once the tissue heals after the initial insertion injury.[260]
These, in turn, have been demonstrated to impact chronic
neural recordings and the efficacy of stimulation therapies.[261]
Another hypothesis is that these lattice structures reduces the
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Figure 4. Implant size impacts acute and chronic tissue integration. a) Tissue around the probe’s (b) thin polymer lattice structure showed significant
reduction in encapsulating cells and improved neural density (c). Reproduced with permission.[259] Copyright 2007, Elsevier. Copyright 2017, Nature
Publishing Group.[267] d,e) Two photon imaging of tissue strain in vivo from a Michigan Electrode Array (d) and carbon fiber microthread electrode (e).
Neurons are green, while recording sites and astrocytes are red. Cyan outline highlights microthread electrode. Neurons in panel (d) are much more compressed and oval/elliptical than neurons in panel (e), indicating increased mechanical strain from the embedded electrode volume. Reproduced with permission.[34] Copyright 2014, Elsevier. f,g) Chronic histology shows carbon fiber microthreads with an 8 µm diameter reduced tissue reactivity and improved
neuron density of microthread (g) compared to silicon electrodes (f). Reproduced with permission.[132] Copyright 2012, Nature Publishing Group.
surface area for biofouling,[132] as well as the diffusion barrier
for proinflammatory cytokines, which cause an accumulation
of these cytokines around the probe.[262] Similarly, electrocorticography (ECoG) arrays also benefit from a more open architecture, due to tissue growth through the holes, which minimizes
the amount of scarring formed between the electrode sites
and the brain.[263,264] Reducing the overall cross-sectional area
of the implant reduces the overall volume of the implant. For
implantable probes, lower implant volumes reduces the tissue
displacement volume that occurs during device implantation,
which in turn decrease the mechanical strain experienced by
neighboring neurons (Figure 4d,e).[8,34] This strain has been
shown to inhibit activity of nearby neurons and cause additional inflammation that negatively impacts the nearby tissue
health.[265,266] Finally, the location of the electrode sites on the
implantable probe is thought to influence recording characteristics. For instance, electrode sites located at the tip of the
probe are hypothesized to have a larger viewing radius (see
Section 2.7), or volume of tissue in which extracellular action
potentials can be detected, compared to sites located along the
shank.[121] In contrast to tip electrodes, sites along the shank
have less visibility to tissue located directly behind or to the side
of the electrode that is partially obstructed by the insulating
shank.
The benefits of minimizing the volume and cross-sectional
area of implants has motivated advances in microscale and
Adv. Funct. Mater. 2017, 1701269
nanoscale electrode technology. Wire shank devices with
1–10 µm diameters have been fabricated with silicon, gold,
carbon fiber, and conductive polymer composite substrates
with excellent electrical properties and minimal host-tissue
response.[268–272] On the nanoscale, etching and chemical
vapor deposition of silicon has enabled the creation of devices
<150 nm diameter for in vitro and in vivo applications while
using tobacco mosaic virus as a biotemplate has enabled
metal wires to be as thin as <20 nm in diameter.[273–279] However, these introduce new challenges,[28] including handling,
packaging, resistive heating, and in vivo chronic implantation requirements.[176,177] In addition, decreasing the size of
the implant corresponds with a smaller electrode site, which
increases impedance. This can be addressed by increasing the
ESA:GSA ratio(see Section 2.4).[43,72,73,193]
Given that FSCV measurements require that neurochemicals of interest either directly adsorb to the electrode surface
or diffuse into the vicinity directly adjacent to the electrode,
there has been a fundamental push toward the use of FSCV
electrodes with diameters of 10 µm or less.[280] The diameter of
the carbon fiber electrode is important to decrease local damage
to the synaptic end-terminals which are the neurochemical
sources, as well as decrease biofouling and/or reactive gliosis
which may create a barrier preventing the adsorption/diffusion of neurochemicals to the surface of the electrode.[280] Prior
studies have demonstrated in vitro that biofouling related to
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protein adsorption to the surface of the electrode can dramatically reduce sensitivity to dopamine and other neurochemicals
of interest.[281] As carbon fibers are mechanically brittle, a stringent requirement of the use of carbon fiber FSCV electrodes
with diameters 10 µm or less may be problematic in terms of
handling and long-term implantation.[282]
There is no published work directly comparing different
diameter sizes for FSCV electrodes and neurochemical
recording quality in a chronic in vivo experimental preparation.
Nevertheless, difficulties have been observed in obtaining acute
dopamine measurements when implanting a 50 µm diameter
boron-doped diamond FSCV electrode, likely due to trauma
to the adjacent tissue preventing dopamine release locally.[283]
However, using a dopamine reuptake inhibitor to promote
dopamine diffusion from intact tissue into traumatized tissue
did enable dopamine detection with these larger devices.[283]
These data support a substantial body of work demonstrating
that (1) implantation of a 220 µm diameter dialysis probe can
dramatically impact the release and reuptake of neurochemicals at least 220–250 µm from the microdialysis probe,[284,285]
(2) tissue near probe tracks created by 280 µm diameter microdialysis probes exhibit ischemia and endothelial cell debris
which is most prominent at 4 h post implant,[286] (3) at 24 h the
microdialysis probe tracks are surrounded by hyperplasic and
hypertrophic glial and glia processes,[286] and (4) 7 µm diameter
carbon fiber probes produce a diffuse disruption of nanobead
labeling, but no focal disruption of blood vessels, evidence of
endothelial cell debris, or glial activation.[286] Across all interface
modalities, there is increasing evidence that subcellular geometries improve long-term tissue integration.
is possible to avoid penetrating surface pial vessels and large
penetrating blood vessels,[289] but with multiple, densely packed
shanks, there is a higher probability for vascular injury, microhemorrhage, and exacerbated scarring and cell death, which
further increases with greater shank density.[290–293] Despite
this, there are few studies that define a “critical” shank spacing
to avoid compounding damage. Both electrophysiological performance and gliosis are progressively worse for Michigan style
devices and microwire devices at tip-to-tip spacings <600 µm,
while wire shanks that are 1 mm apart are indistinguishable
from solitary electrodes.[294–296] For tapered devices, the base-tobase spacing of shanks should also be considered. The blood
supply for the underlying cortex enters through the surface of
the brain, and it is possible that the high-strain of narrow baseto-base separation could disrupt that blood flow.[297] Evaluating
the effect of shank density on the hemodynamics of the underlying cortex should be examined in future studies.
Common types of multishank arrays (silicon planar
Michigan arrays, silicon bed-of-needles Utah arrays, and
metal microwire arrays) have been compared in 1–4 monthlong rodent studies.[298–300] These studies show that metal
wire shanks tend to outperform planar or bed-of-needles style
silicon devices in chronic electrophysiology with higher SNRs
and lower impedances, as well as with less long-term BBB dysfunction. In addition to different materials, these devices vary
in shank density, shank shape, and platform area, complicating
the interpretation of these results. Future studies should systematically vary these parameters to determine their significance for long-term probe viability.
2.8. Flexibility: Softness/Geometry
2.7. Volumetric Density and Channel Count
For recording applications, the demand for a greater number
of channels has increased. More recording channels allows for
a greater volume of tissue to be sampled, and further, multiple
channels located closely together allows for the sampling of overlapping volumes of tissue.[152,287,288] This makes triangulation
of the signal across multiple electrodes possible. An important
consideration for multishank devices is the impact of the overall
footprint of the device on the tissue health. More specifically,
the volumetric density that the device occupies in the brain over
a given amount of space can impact performance. Even if the
same volume is implanted into the brain, the maximum strain
on the tissue can be reduced by distributing the implant volume
over a larger tissue volume as demonstrated by finite element
modeling (Figure 5). For example, a single shank tetrode made
from four 50 µm diameter electrodes may generate a tissue
response similar to a single 100 µm diameter wire, but would
generate a very different tissue reaction when compared to four
50 µm diameter wires that are implanted with 1 mm spacing
over a large region of the cortex. However, additional considerations may be required based on how the shank spacing impacts
injury (tearing and compressing) to the underlying vasculature.
When considering configurations of multiple penetrating
shanks, the most important factors are the density of the shanks
and the configuration and footprint of the platform piece where
the individual shanks come together. With a single shank, it
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Once in the brain, a neural probe spends the rest of its life battling a wet, chemically dynamic, and mechanically tumultuous
environment. This mechanical environment is dominated by
“micromotions,” which are the movements of the brain in the
cavity of the skull in response to respiration, pulsation, and everyday movement.[301] This effect is likely to be greater in larger
primates than rodents due to larger and more intricate brains
and interstitial space.[302] While the material composition of the
probe will largely determine how it fares in the mechanical and
chemical milieu of the brain, many groups have also shown
that the flexibility of the probe can have an impact on how the
host-tissue response and micromotion-related strain is generated by the probe.
Flexible implantable devices are believed to reduce the
mechanical strain experienced by tissue due to motion especially in the brain when devices are anchored to the skull. The
brain is normally floating inside a cerebrospinal fluid filled
dural sac which in turn floats inside of the skull. In humans,
the brain and skull are usually separated by about 2–4 mm
while this number is closer to 100 µm in mice.[35] As such,
large forceful movements or impacts to the skull can lead to
substantial brain tissue displacement and mechanical strain to
the tissue surrounding rigid implants. In addition, there are
smaller micromotions and physiological motions in the brain
due to respiration and heart-related blood flow pulsation.[303,304]
These small movements can increase when craniotomies and
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Figure 5. Probe insertion results in high strain between shanks of multishank devices, with more strain generated between densely packed shanks.
A finite element simulation of Michigan-style planar silicon probe insertion into a viscoelastic brain phantom is shown. The maximum strain value of
each element following insertion is displayed. a) When the probe shanks are spread far apart (500 µm tip-to-tip spacing), there is little overlapping
strain between shanks. b,c) As tip-to-tip spacing gets tighter, there is increased tissue strain between shanks (300 and 150 µm). Scale bar = 100 µms.
duratomies are left open and unsealed.[305] Finite element modeling studies have shown that stiff or rigid devices that are
anchored to the skull resist complying with these brain movements.[306,307] In turn, this can lead to a perpetual generation
of mechanical strain and proinflammatory cytokines or even
reinjure the tissue resulting in device failure.[5,8,34,308] As such,
there has been increased focus on developing implantable technologies that are much more compliant, which can be achieved
in several ways
kc _ Normal =
l
E⋅A
(12)
kc _ planar =
4l 3
Ewt 3
(13)
kc _ lateral =
4l 3
Etw 3
(14)
kc _ radial =
64l 3
3π E (do4 − di4 )
(15)
k=
1
kc
(16)
where kc is compliance, l is the device length, A is the crosssectional area, E is the elastic modulus (or Young’s modulus),
w is the width of a planar devices, t is the thickness of a planar
devices, do is the outer diameter of a cylindrical devices, di is the
inner diameter of a cylindrical devices, and k is stiffness. The
compliance of a device can be improved by employing materials
with a low elastic modulus, reducing the cross-sectional area of
the device, or increasing the length of the device. In order to
reduce the rigidity of neural implants from silicon backbone
substrates and large metal backbones (100’s GPa), engineers
have turned to polymers and thin-film technology.[309,310] Polymers can have softer material properties (Young’s modulus),
low thermal expansion, exceptional solvent tolerance, and
good dielectric constants. However, polymers can also have
high water absorption, poor solvent tolerance, and changing
Adv. Funct. Mater. 2017, 1701269
dielectric properties that increase when it has absorbed water
or ions, which negatively impact the chronic performance of
implanted devices. Therefore, polymers selected for device
design must satisfy these broad range of criteria.
Hydrogels are among the few materials that can be engineered to have similar mechanical properties to brain tissue
due to their high water content.[311–313] Microglia have been
shown to be responsive to the stiffness of a material, migrating
from softer to stiffer surfaces.[314] However, hydrogels lack
many of the other material properties necessary for functional
performance, such as electrical, strength, and durability properties. Furthermore, the swelling of hydrogel materials can lead to
additional neural cell death if the hydrophilicity, crosslink density, and weight of the polymer is not properly optimized.[315,316]
This greatly limits the ability to construct functional devices
from low elastic modulus hydrogels.
Reducing the cross-sectional area of the probe track has
been shown to yield significant improvement in chronic neural
recording performance, but as the dimensions decrease, other
challenges emerge such as the mechanical handing requirement related to insertion and durability.[73] Similarly, overall
increases in device length lead to new challenges. For specific anatomical targets, length can be increased by entering
the tissue at a very distant region, though this substantially
increases tissue injury. Alternatively, a meandering design,
such as a sinusoid, could be employed to increase the length
of the device without dramatically increasing the implantation
related injury.[317–320] The placement of trifilar coils along the
lead combined with strategic anchor points to prevent migration, such as with suture pads, may improve device performance, similar to pacemakers.
The last decade has seen intense excitement in the development of wearable and soft electronics[321–325] but rarely are these
new materials used in vivo. Little is known about the performance of wearable systems for long-term in vivo use and the
number of papers showing in vitro monitoring or actuation
is quite limited. Another major challenge is that all of these
devices currently require much larger substrate thickness
and feature sizes relative to thin flexible arrays patterned with
photolithographic approaches such as polyimide, parylene, or
SU-8 substrates.[129,130,326–340] Most applications currently favor
thin substrates and small feature sizes because the tradeoffs for
biomimetic elasticity at the expense of size are still unproven.
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In addition, as implantable probes become more flexible,
motion related electromagnetic artifacts are likely to manifest
within the probe. Motion artifacts are caused by the movement or bending of a current carrying conductor, resulting in
a change in the induced electromagnetic field. Although these
artifacts are typically observed in the external connector cable
(see Section 2.12), increasing implant flexibility may increase
the likelihood of motion artifact occurrences due to perturbances from physiological micromotion. These motion artifacts
are further exacerbated by changes in the internal resistance
of soft conductors or in the capacitance of the insulation due
to mechanical strain, which alters the alignment of the material’s conductive molecules and the cross-sectional area of the
material. This, in turn, can dramatically impact the amplitude
and waveform characteristics of the electrical signal. While
improving the softness and flexibility of the device is hypothesized to improve chronic tissue interfaces, it can negatively
impact conductivity, dielectric properties, impedances, handling
requirements for insertion, material strength, packaging, and
scaling.[66,341] Balancing these properties with advanced composite materials or a blend of multiple materials has proven to
be the critical challenge for developing functional implantable
devices.[66,68] Ideally these devices will be anchored to the region
which they are intended to sample and deliver their signal reliably amid all the stresses and strains produced by normal brain
motion.
2.9. Implantation Technology
Well controlled probe insertion techniques play a critical role
in optimizing recording performance and minimizing injury
during surgery.[342] This is particularly true for thin, flexible
polymer probes that experience buckling when penetrating
tough pia mater and, once successful, deflect from original site
of implantation.[343,344] Insertion must generate enough strain
to puncture the leptomeninges and brain tissue (as little as 0.05
shear strain or 1 mN penetration force), but without causing
damage to the surrounding neural tissue, which occurs with
as little as 0.1 strain at a 10 s−1 strain-rate or 3% cyclic strain
at 0.2 Hz.[344–349] The pia mater is relatively heterogeneous
with differing regions of thickness and surface vasculature,
so the location of insertion has implications on insertion success. Implanting through thicker pial blood vessels increases
compression on the cortex before penetration occurs, while
insertions through smaller vessels demonstrate reduced compression on the brain.[350]
In pursuing designs that minimize insertion injury, in vivo
and in vitro probe insertion experiments consistently show that
rigid devices with sharper-tipped and lower cross-sectional area
devices have lower penetration forces, less tissue dimpling,
and less tissue strain during insertion.[346,351–355] Considering a
single penetrating shank, the major implicated parameters are
tip shape, presence of edges, and cross-sectional area and profile. Devices with smaller profiles display reduced cortical insertion injury in postmortem histological analysis <1 week postimplant.[356] However, sharper tips have other trade-offs such
as leading to more vascular damage, especially at lower speeds
in in vitro models.[346] In addition to device geometry, implant
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surface chemistry has been shown to play an important role in
insertion. Notably, both hydrophobic (silane and Parylene C)
or hydrophilic (acid functionalization) surface treatments were
able to reduce penetration force (silane and acid treatment) or
friction force during insertion (Parylene C).[352,354]
There is less consensus about the effect of insertion speed
on implantation damage. The brain is thought to be a hyperviscoelastic material, which means the faster the tissue deforms,
the higher the Young’s modulus (stiffer).[352,355,357–361] Thus, at
higher insertion speeds it may require more stress to generate
enough strain to puncture. In vivo puncture studies of single
shank devices have corroborated this, showing more dimpling, higher penetration force, greater penetration damage,
and ultimately more long-term gliosis at higher insertion
speeds.[352,355,362,363] Interestingly, lower insertion speeds could
lead to more vascular damage, which may be due to differing
mechanical properties of brain and blood vessels. While strategies to map and avoid larger blood vessels with two-photon
microscopy can minimize most of this vascular damage, this
may be impossible for devices with multiple shanks.[289] For
these multishank devices, high-speed insertion velocities
>8.3 m s−1 have demonstrated improved insertion success,
reduced hemorrhaging, and improved gliosis.[290,291,360]
So far insertion has been reviewed with respect to rigid,
brittle substrates that can be machined into pointed shanks for
facile penetration. However, stiff devices result in mechanical
mismatch with the surrounding tissue (discussed previously in
Section 2.8).[306,364] In contrast, softer, more flexible substrates
are more prone to buckling upon insertion.[344,365] To address
this, groups have proposed using delivery shuttles to deliver
devices into the brain. Some simple designs use electrostatic
interactions to adhere flexible probes to a silicon substrate.[366]
Following implantation, moisture in the brain can disrupt
the interactions between the silicon substrate and the flexible
probe to allow for shuttle retrieval. This design can be scaled
by using a “sewing machine” method, which uses an insertion
needle that can mechanically attach and detach from a flexible
probe for more rapid implantation.[367] However, these require
additional design considerations for the coupling mechanism.
Alternative approaches involve fully or partially coating a flexible device with a sacrificial, resorbable polymer.[368–374] In vivo
studies have shown that insertion wounds around resorbable
insertion shuttles can fully close with limited evidence of an
inflammatory lesion at 4 weeks postimplant.[375] While this is a
promising strategy, current fabrication methods are costly and
labor intensive, and finished products are often very sensitive
to moisture and handling.[318,371] These materials and design
issues of scalability and ease-of-use must be addressed before
these next generation techniques can be clinically realized.
2.10. Mechanical Strength
Materials must maintain robust mechanical strength despite
implementation in smaller and more flexible devices,
demanding careful material selection and geometry design. As
devices become smaller and more flexible with material selection and design geometry, the strength of the material becomes
an increasing problem. Mechanical strength varies with
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material, size, orientation, and geometry. Therefore, mechanical
properties are often quantified in different ways: yield strength,
compressive strength, shear strength, tensile strength, fatigue
strength, and impact strength. Decreasing cross-sectional area
(A) leads to increased stress on the implant
σ stress =
F
A
(17)
where the stress on the device (σstress) is a function of the force
(F) applied on the device area (A). The stress on the implant
causes strain (ε) or deformation of the material, while the
stress–strain relationship is linear
σ stress = Eε
(18)
As the relationship becomes nonlinear, elastic deformation
of the implant becomes plastic, resulting in permanent deformation. Brittle materials such as varnish, glass, or even diamond undergo limited plastic deformation before failing.[376,377]
In contrast, tough metals can undergo extensive plastic deformation before reaching failure, making them more resistant
to mechanical stresses.[378] However, over time, small plastic
deformations can cause failure through fatigue. Elastomers are
a class of polymers that are capable of enduring large amounts
of elastic deformations, and are the preferred choices for nextgeneration devices.[22,27]
While strong polymers such as aramid or polybenzoxazole
exist, their high resistivities, high dielectric properties, and
sometimes poor chemical stabilities as thin films in physiological solutions make them difficult to implement as building
blocks for neural interface technologies.[379,380] As such, a major
challenge becomes balancing size with multiple mechanical,
electrical, and chemical properties of both conductors and
insulators. One approach to balance mechanical strength with
conductivity is to incorporate carbon-based materials, such as
graphene or carbon fibers into composite materials.[28,132,381]
These composite conductors represent flexible alternatives to
brittle thin-film metal traces.[28,84,89,90,382–385] For example, layerby-layer assembled polyelectrolytes in particular have demonstrated some limited self-healing capacity.[386] The noncovalently bonded individual particles display some limited mobility
within the composite material resulting in the material’s flexibility and capacity to heal defects. Alternatively, controlled
addition of defects (adding tears or holes in strategic locations)
in thin-film metal traces allow greater distribution of stress,
allowing the device to undergo greater amounts of strain before
failure.[387]
One material with exceptional mechanical strength and stiffness is silicon carbide (SiC). SiC is a wide band gap semiconductor with highly saturated electron drift velocity (leading to
high electron mobility and good conductivity at high current
frequencies) and a high thermal conductivity.[388] While SiC has
more limitations in lithography and microfabrication compared
to other forms of silicon, it can be grown readily and doped
with p-type or n-type dopants.[388] It is also extremely resistant
to physical wear and corrosion[388,389] and has improved biocompatibility compared to silicon[48] and surgical steel.[390] This has
renewed interest in applying silicon carbide for chronic neural
Adv. Funct. Mater. 2017, 1701269
interfaces. The amorphous phase of SiC is used as an insulating coating, while the polycrystalline phase is conductive.[48]
SiC has also been incorporated into nanocomposites with conductive polymers, augmenting the conductivity of SiC and the
mechanical properties of the employed conductive polymer.[391]
Additionally, the electrical conductivity of this composite can be
“tuned” by varying the particle loading.[391]
2.11. Mechanical Mismatch of Subcomponents and Geometry
While the mechanical mismatch between the implant and tissue
has been well characterized,[306,307] mechanical mismatch of
implant subcomponents has recently been emphasized as critical points of failure.[35] For example, iridium and bulk silicon
have elastic modulus’ of 528 and 200 GPa, respectively.[35] As a
result of this mismatch, mechanical strain occurs, especially on
the lower modulus material during implant movement, such as
from physiological motion. Finite element modeling revealed
that these strains are further focused on protruding geometries,
leading to material failure at these mechanical mismatched
interfaces as demonstrated by in vivo experiments (Figure 6).[35]
These finite element models also demonstrated that geometry
and placement of these subcomponents impact the durability
and mechanical strain experienced. Specifically, the center
of the substrates experience the least amount of strain, while
outer surfaces, especially on protrusions and edges, experience
the greatest amount of strain. In turn, these impact the integrity of the electrical trace and insulation material. Furthermore,
these models emphasized that with simple and uniform geometries, these strain profiles are more uniformly distributed and
shared more evenly across the entire device rather than along
sharp edges.[35] In some situations, simple and uniform geometries may have improved strength and durability compared to
complex lithographic designs (Figure 6f).[35]
2.12. Tethering and Packaging
The penetrating shaft of an intracortical electrode brings electrode sites close to neurons of interest, but in many applications, the neural signal must be transferred to an internal pulse
generator or an external system. To accomplish this, the shanks
and electrode traces of the device are joined at a platform at
the surface of the brain (Figure 7). The platform interfaces with
the skull-anchored connector, which interfaces with external
devices. The design of the platform and connector (jointly
referred to as the device’s packaging) has significant ramifications on the long-term viability of the device. Extracorporeal
connectors are at risk for traumatic damage and material failure
due to environmental exposure.[253,392] Metals, ceramics, and
polymers have all been used for the packaging of subcutaneous
implants, but they vary in their benefits: ceramics are known
for biocompatibility and corrosion resistance yet lack proper
mechanical strength and reliability, metals have high ductility
and durability but are usually bulky and difficult to scale down
in size, and polymers allow for tunable mechanical properties
but suffer from absorption of fluids and swelling.[393] The current gold standard is to use a titanium casing with insulated
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Figure 6. Mechanical failure from nonhomogenous components and geometries. a) A finite element model (FEM) of mechanical strain on a planar
silicon array with iridium electrode sites under a 1 µm micromotion in tissue. b) A strain profile along the protruding electrical trace (red dashed line),
substrate surface (black solid line), and at the center of the substrate (blue dash dot line) as indicated in (a). c) Strain profile of (a) from the top view.
Lowest strain (black) and highest strain (purple) regions are indicated with arrows. d) SEM of a planar silicon array explanted 189 d post implant in
mice V1 cortex. e) 2× zoom in of (d). f) FEM showing an order of magnitude less, evenly distributed strain profiles along simple elegant geometries.
Reproduced with permission.[35] Copyright 2015, Elsevier.
wiring to hermetically seal electrical components from exposure to biological fluids,[394] and help protect device circuitry
from electromagnetic interference.[395] Packaging choices for
Adv. Funct. Mater. 2017, 1701269
size and form-factor of the platform and connector can also
enable on-board electronics such as preamplifiers and filters or
incorporate wireless chips as modules.[267,396] This is especially
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Figure 7. Schematic of different neural probe packaging strategies.
Most commercially available technologies are anchored to the skull by
cement or bone screws in order to achieve mechanical stability (top).
The platforms and shanks of a device can either be directly attached to
the connector (rigidly anchored) or attached to a wire bundle intermediary (tethered). Next generation devices could be unanchored (Bottom).
These devices would use telemetry to send neural data to external neural
processors. Inset: meningeal fibrosis of the wire bundle of a Utah array
that was implanted in a nonhuman primate for ≈4 years (inset, top).
Second-harmonic generation imaging of fibrous tissue around wire
bundle shows thick collagenous tissue (≈500 µm thick) surrounding
wire bundle (inset, bottom). Nonhuman primate data is courtesy of the
Batista Lab, University of Pittsburgh.
true of fully subcutaneous deep brain stimulating (DBS) systems, which house their power source and electronics in an
Implanted Pulse Generator (IPG). The IPG is typically secured
in a distant tissue compartment, necessitating durable material
choices to withstand movements. One of the more common
issues for patients using DBS[397] and pacemaker[398] devices is
lead breakage and infection. The mechanical mismatch between
the hermetically sealed casing and the insulated wires make
this joint a preferential failure point (see Section 2.11). Efforts
to minimize potential failure modes involve wireless devices
that power IPGs through radiofrequency (RF) telemetry. Still,
this strategy requires significantly more power to operate, bulky
device components, and is complicated by tissue that impedes
wireless coupling between subcutaneous implants and external
devices.[399] Transparent ceramics can avoid this issue as well
as enable optical surveying of neural activity and optogenetics
(Section 2.15).[400]
Packaging choices also affect the biological response to
implants. Devices with larger platforms such as 10 × 10 or 4 ×
4 shank Utah arrays are prone to fibrous encapsulation by the
surrounding meninges. While the burden of penetrating shafts
likely plays a role in this outcome, a similar encapsulation
Adv. Funct. Mater. 2017, 1701269
process occurs for platforms without penetrating shafts, such
as silicone ECoG implants, which sit sub or epidurally on the
surface of the brain.[264,401,402] The platform can become enveloped in a fibrous capsule as early as 7 d postimplant, with
encapsulation persisting for the lifetime of the device.[263,264,401]
For these implants, fibrosis can be mitigated by leaving fenestrations throughout the substrate or using a resorbable substrate that leaves only electrode sites and traces on the surface
of the brain.[263] Additionally, one study[403] examining nonfenestrated ECoG grids, found that hematomas formed in 14% of
cases after subdural implantation of electrocorticography grids
in human children. Subdural hematomas located beneath an
ECoG array can interfere with recording. These grids can also
form depressions on the surface of the brain, indicating that
the weight of an ECoG array is an important consideration.
However, mesh-based micro-ECoG devices with a more open
architecture suggest that they allow tissue vasculature to wrap
up and around electrode sites, reducing the formation of collagen scar tissue between electrodes and the brain in at least
one study.[263]
In devices with low-profile or 2D platforms that can be fully
embedded in dental acrylic, there are no reports of fibrous
encapsulation, but possible fibroblast and circulating white
blood cell influx into the parenchyma from the meningeal
space.[263,403–405] Cell influx can be mitigated by implanting
devices fully in the cortex, eliminating any device-meninges
interface.[404,406] While the significance of this cell population
is unknown, it presents as an important consideration for
designing the meninges-device interface in future probes.
Platforms are also important fixation points. Platforms that
are not fixed can migrate during the brain’s recovery from
implantation surgery. Anchoring can either be accomplished
by directly attaching the platform to the connector (rigidly
anchored) or held in place by a wire bundle that attaches to the
connector (tethered) (Figure 7). In the latter scenario, the connector is cemented and/or screwed into nearby skull bone. The
method of tethering and packaging of the wire bundle cables,
however, can affect the long-term viability of an implant. For
one, due to their proximity and thin insulation layer, wires in
the bundle are the primary source of parasitic capacitance (Section 2.2, Equation (6)). Additionally, several studies in rodents
have shown that tethering the device leads to increased reactive microglial and astrocytic gliosis and neuronal cell death
compared to untethered devices in both planar, Michigan
style arrays[404,406] and microwire style arrays.[300] Untethered
microwire arrays also outperform tethered microwire arrays in
long-term electrophysiology SNRs.[300] This is likely due to slack
in the wire bundle, which can allow for more gross movement
of the implanted device during micromotion and impacts. Analysis of the implant tracks for tethered devices corroborate this,
which appear to be more elongated than tracks around untethered devices.[407] Movement of the wire bundle can also be a
source of motion artifact and lead fracture.[408–412] This could be
somewhat reduced by placing the bundle in a trifilar coil, which
can allow expansion and contraction of the bundle. It is also
possible that the wire bundle generates its own tissue response,
which can further affect the relative motion of the brain and
the implant. This is seen in the wire bundles of Utah arrays in
nonhuman primates, which can be become fully encapsulated
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in fibrous tissue after 4 years of implantation (Figure 7). The
potential for additional fibrous tissue buildup at the interface
of the wire bundle and platform may explain why Utah arrays
in rats are frequently slanted and partially explanted.[411] The
packaging material and design are critically important components that impact overall functional performance that are often
underappreciated.
2.13. Wireless Transducers
While untethered devices have better outcomes than tethered devices, they still must be anchored to the skull, which
causes incongruous motion of the brain and implanted device
(Figure 7). It is no surprise that devices that are not anchored to
the skull or to a gel sealant have reduced astrocytic and microglial gliosis than their anchored counterparts.[404,406] The percutaneous nature of the connector also presents an infection risk
that is much lower for fully subcutaneous deep brain stimulators.[406,413,414] These issues press the need for fully implantable, wireless recording systems that are not anchored to the
skull. While completely wireless, closed-loop systems for motor
cortex recording and spinal stimulation have been achieved and
able to treat gait dysfunction, the brain–computer interface is
still bulky and transcranial.[415,416] Smaller, fully implantable
wireless modules have been fabricated that can directly interface with commercially available neural probes.[417–420] These
devices still require in vivo validation, but are a promising
strategy to adapt mature technology to the next generation. Less
invasive options include fully injectable millimeter scale electrodes which can interface with neural signal processors via RF
or ultrasound.[25,421] Similar devices could also be incorporated
on an intravascular stent-electrode device to minimize invasiveness even further.[422] It is unclear, however, if these devices can
record single neuronal unit activity.
There are also challenges involved with wireless stimulation
of neurons. Clinically available deep brain stimulators require
subcutaneous stimulation generators and batteries, but these
are bulky and batteries must be periodically replaced. One
possible solution for this is to use an external energy source
to stimulate neurons. Toward this end, RF or ultrasound stimulation can be targeted to electrode devices to stimulate neurons, though in vivo validation is required.[423–425] Infrared
light can be used to thermally stimulate neurons, but it must
be delivered through an optically transparent “optrode” device
to achieve localized delivery and avoid light scattering.[424,426,427]
Visible light can also be used to stimulate neurons by using a
photovoltaic substrate, which can convert light into electric current, though they often require toxic materials and have poor
stimulation efficiency.[408,428–431] A somewhat less invasive
strategy is to deliver nanomaterials such as gold or magnetic
iron oxide nanoparticles to a neuronal population and stimulate through IR and near-IR light, ultrasound, or magnetic
fields.[432–435] This also offers advantages over conventional
stimulation through infrared or transcranial magnetic stimulation in that nanoparticles can be targeted to specific populations
of neurons through localized injection, magnetic targeting, or
modified surface chemistry.[427,434,436,437] Many nanomaterials
that can be incorporated into wireless stimulation can also be
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used as stimulated drug-release carriers to further increase the
therapeutic flexibility of wireless stimulation.[434,437–441] However, nanoparticles can undergo photothermal reshaping at
relatively low energies (0.6–1.5 eV) driven by curvature-induced
surface diffusion at temperatures well below the melting point
threshold.[442] This is particularly important for nanoparticles
that have specific shape and aspect ratio that allow tuning to
specific ranges of wavelengths or frequencies.[443] For chronic
applications, the long-term stability of nanorods and nanoparticles on stimulation performance and biocompatibility need
to be considered at temperatures well below the melting point
of these particles. More recently, using carbon fibers as wireless
antennas, carbon fiber photoconverters have been stimulated
using photovoltaic and photothermal principles and with better
spatially localized neural activation pattern than achievable by
optogenetics or electrical stimulation.[182] Emerging materials
and new wireless transduction technology together promise
novel modalities for interfacing with the nervous system and
mitigation of the risks of percutaneous devices. However, these
new systems must be thoroughly characterized to identify
unique challenges and limitations.
2.14. Molecular and Biochemical (Coatings) Engineering
The chemical and mechanical properties of the probe’s surface
play a large role in determining how the brain and device will
interact with one another. Researchers exploit surface chemistry
to mitigate molecular and cellular fouling of the probe, modulate the neurodegenerative and corrosive behavior of encapsulating cells, and deliver therapeutics. As protein adsorption and
cellular attachment to the neural probe’s surface are the first
steps in forming scar tissue, many groups are pursuing surface
chemistry approaches to prevent molecular and cellular adhesion to the probe.[182,356,444–446] The most basic approach to this
is to modulate the hydrophobicity of the probe’s surface.[447–449]
While hydrophobic surfaces prevent hydrophilic domains of
proteins from binding to the surface of the probe, hydrophobic
domains can still attach.[450] Hydrophilic and zwitter-ionic coatings can prevent both hydrophobic and hydrophilic proteins
from attaching by using hydrogen bonds and electrostatic
interactions, respectively, to hold a water layer at the interface of the probe and brain. This displaces proteins and cell
binding domains, preventing attachment.[446,451] Various PEG,
poly(OEGMA), and hyaluronic acid hydrophilic coatings have
been incorporated onto neural probes to prevent fouling and
gliosis in vitro and in vivo.[271,452,453] As an added benefit, hydrophilic or zwitterionic polymer chains can also sterically prevent
adsorption and adhesion.[271,454,455] Unsurprisingly, coatings are
more effective in preventing gliosis in vivo if they are applied
to the whole penetrating shank as opposed to just the electrode
sites, however electrode site coatings can suppress increases in
impedance after initial implantation.[453,455]
Despite promising results, many current antifouling coatings fail over time as coating materials oxidize and molecules
and cells find vulnerabilities within the coating.[446,450,455] Alternative coating strategies seek to modulate the phenotype of contacting cells instead of preventing attachment. These strategies
seek to improve neuronal attachment and regeneration while
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preventing microglial and astrocyte adhesion and activation. By
reducing inflammatory cell activation, reactive oxidative species (ROS) output can also be reduced, which has been shown
to cause material degradation of most commercially available
probe types.[254,456–458] Cells can be modulated by modifying the
surface topography of the probe, immobilizing bioactive molecules to the surface, and/or applying drug-releasing coatings
to the probe.
Coating stiff devices with a low density or soft material with a
Young’s modulus closer to that of brain tissue can reduce gliosis
and improve neuronal survivability.[314,459,460] Because hydrogel
materials lack the electrical, strength, and durability properties
necessary for functional devices, they are commonly used as
surface coatings with and without bioactive molecules which
have been shown to improve the tissue response.[314,316,461–463]
One disadvantage of hydrogel coatings is that they can push
neurons away increasing the distance to the nearest neuron,
especially when it swells, which decreases recorded amplitude.
Furthermore, depending on the hydrogel crosslink density, it
can provide a scaffold for infiltrating macrophages and neutrophils. Hydrophobic, silicone based gels can avoid swelling issue
and still yield lower glial attachment.[464] Macroscopic gel coatings that cover electrode sites may additionally increase device
impedance, though the effect is minimal for swollen hydrogels.
Furthermore, conductive polymers can be electropolymerized
within the hydrogel to improve the electrical properties.[465] Like
soft coatings, patterned surfaces can also have an impact on
adherent cell behavior. Etched silicon substrates with 3D grid
patterns or pore features reduced microglial attachment while
improving neuronal attachment in vitro and in vivo, respectively.[466–468] Improved and directed neuronal cell attachment
have also been demonstrated in vitro for synthetic polymer,
gold, and extracellular-matrix based micro and nanopatterned
surfaces.[469,470] Conductive-polymer and carbon nanotube or
graphene oxide composite coatings have also been shown to
improve cell attachment and neural stem cell differentiation,
though it is unclear if this is due to the topographical features
of the coating, or if there are other chemical or mechanical
pathways in play.[470,471]
Attaching bioactive molecules to the surfaces of neural
probes can also modulate cell behavior. Molecules can either
be attached directly to the probe’s surface, or can be chemically tethered[472] or mechanically interlocked[236] to polymer
coatings. Since immobilized molecules cannot directly act on
intracellular targets, most coating strategies use molecules that
target extracellular targets. Attaching extracellular matrix molecules such as laminin, peptide sequences from laminin, or the
proteoglycan hyaluronic acid to the surface can “camouflage”
probes such that host cells do not recognize the devices as foreign. These coatings have been shown to reduce gliosis and
prevent neuronal cell death.[177,236,472,473] Immobilized cell adhesion proteins like L1, which brokers neuron–neuron binding,
have been shown to prevent gliosis and neuron death as well as
to encourage neurite regeneration in vivo (Figure 8).[452,474–477]
Other immobilization strategies seek to actively suppress
inflammation. Attaching antagonists of proinflammatory
surface receptors or ROS scavengers to remove extracellular
inflammatory species can reduce gliosis and spare neurons
in vivo.[478,479] Small bioactive molecules, such as superoxide
dismutase mimic ROS scavengers, have the added benefit of
lower immunogenicity and better in vivo stability compared
with fragile peptides and proteins.[479] On the other hand,
ligands that mimic the physiological signals for transcytotic
transport can be attached to nanoparticles to enable transport
across the blood–brain barrier.[480]
Drug eluting coatings have the advantage of modulating
cell behavior at greater distances and acting on intracellular
targets, whereas surface-bound bioactive molecules can only
directly act upon membrane targets of adherent cells. Perhaps
the most explored drug release target is the antiinflammatory
Figure 8. Microglial encapsulation of neural probes is attenuated by L1 cell adhesion molecule coating. Both ultrasmall carbon fiber microthread
electrodes (left, blue outline) and silicon Michigan microelectrodes (middle, blue outline) are encapsulated by microglial processes (green) after 6 h
of implantation in mice. Electrode devices can be “camouflaged” by covalently attaching bioactive molecules to their surfaces, as shown by an L1 cell
adhesion molecule coated Michigan electrode (left, blue outline), which is relatively unburdened by microglial encapsulation. Reproduced with permission.[8] Copyright 2015, American Chemical Society. Reproduced with permission.[475] Copyright 2017, Elsevier.
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drug dexamethasone. Dexamethasone, which has been effective in reducing gliosis, sparing gliosis, and reducing electrode
impedance, can be administered passively from polymer coatings or actively administered from conductive polymer coatings
or porous metal electrode sites by electrical stimulation, systemic injection, and microdialysis infusion.[480–489] It is unclear,
however, if the effect extends beyond the release period, or if
the standard host–tissue response is simply delayed by drug
release. Similar dexamethasone-eluting pacemaker leads can
have long-periods of drug release (>10 years) with reduction in
fibrosis and lower pacing thresholds.[490,491] Other small molecules are being explored as well. Antiinflammatory cucurmin
and resveratrol delivered through systemic injection or resorbable polymeric coatings on devices are effective in preserving
neurons and repelling glial scarring, but only for as long as
the drug is being released.[491–493] Protein release from resorbable polymers, conductive polymers, and hydrogel coatings has
enabled the release of alpha-melanocyte stimulating hormone
to reduce inflammation as well as neural growth factor, neurotropin-3, and brain-derived growth factor to stimulate neurite extension toward devices.[494–498] This was demonstrated
in glass cone electrodes loaded with growth factor have shown
stable neural recordings for at least 5 years, though the glass is
brittle and susceptible to shattering; it remains unclear if the
stability of the recordings can be attributed to the growth factors or other design features.[499,500] Therefore, glass devices
should be pursued with extreme caution due to their fragility.
In addition to drug delivery, stem cell delivery has also been
achieved by seeding stem cells onto laminin or alginate coated
probes, though once in the brain most cells differentiate along
a non-neuronal fate.[499,501]
Lastly, conductive polymers have long been used as a method
to attach bioactive molecules to the electrode site and for electrically stimulated controlled drug release. The electrical advantages of CPs[501,502] can be further enhanced by co-depositing
conductive macromolecules into the coating such as graphene oxide or carbon nanotubes.[502,503] In vivo, these loaded
conductive polymer coatings have proven to be safe, chemically and mechanically durable, and beneficial for long-term
neural recordings, with larger spike amplitudes, lower noise
floors,[155,271,504–506] and are stable over long periods of stimulation.[155,271] They have also been implicated in encouraging neuronal attachment and neural stem cell differentiation.[471,507,508]
Combined, these features make conductive polymers a versatile tool that can be incorporated into any electrode device,
including those with soft substrates.[509] There are molecular
and biochemical coating strategies available for devices of all
materials, conformations, and dimensions. Each device should
consider how to leverage surface chemistry to improve the
host–tissue response while maintaining electrical properties.
2.15. Optogenetics and Optical Engineering
The development of protein based light-sensitive ion channels (opsins) and fluorescent sensors (voltage sensors, calcium
sensors, sensing fluorescent reporters) have become a critical
tool for achieving cell-specific interrogation. Light delivery and
virally targeted opsins have gained increasing scientific and
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commercial interest as a potential technology for neuromodulation. Optogenetic tools have encountered material challenges
related to achieving high-density and efficient light control (see
ref. [267] for review), but these technologies are only reliable
for months and intended for animal use. This section discusses
the material challenges of long-term optical stimulation that
includes the packaging of light sources. It is important to note
that optogenetics also faces important molecular biology challenges[14,510,511] including transduction reliability and regulatory
concerns.[512]
The central component of an optical stimulation device
is typically a light emitting diode (LED) or laser diode, and
depending on the application, the use of materials such as
quantum dots or photoluminescence elements might be desirable. One key advantage of LEDs (including µLEDs and organic
LEDs [OLEDs]) is this element can directly illuminate the target
tissue and thereby simplify the optical system. LEDs provide
a wide illumination profile (typically a cosine function, i.e.,
Lambertian) and thus are an efficient and effective way to illuminate reasonable volumes of tissue close to the component
surface. However, if the presence of the LED(s) interferes with
nearby sensors, generates too much heat, or if the required
LED lifetime is long (e.g., ≥1 year), then the use of waveguides
to deliver the light from an external light source such as a laser
diode might be a better long-term strategy. Coupling light into a
waveguide can be very efficient when using a laser diode since
those have pinpoint directed emission, while the coupling efficiency of an LED is exceptionally poor by comparison unless
the waveguide is larger than the surface area of the LED.[513]
A common challenge for all light sources is their energy
requirement to generate sufficient irradiance, while minimizing
heat generation due to inefficiencies in the current driver,
quantum efficiency, and light outcoupling. It has been estimated that optogenetic stimulation requires as much as 1000fold more power than electrical stimulation to excite matching
tissue volumes.[512] Therefore, the electrical insulation must
have high dielectric strength to compensate for the additional
electrical power necessary to drive LEDs and be robust against
insulation failure. This is especially true in locations outside
of the brain that are prone to mechanical forces or movement
(e.g., muscle). However, power requirements also depend on
the distance to the target, opsin sensitivity, and the irradiance
required to cause a therapeutic effect. As the efficiency of LEDs
and opsins improve, currently available low power µLEDs and
OLEDs become practical for in vivo implantation (Figure 9).
One inorganic LED system having a pixel size of 10 × 15 µm
has been demonstrated to control neural activity with as little as
60 nW (optical power), so even though the power efficiency is
only ≈1%, the heat generation does not exceed safety limits.[514]
Organic LEDs having a pixel size of 6 × 9 µm have also been
demonstrated to drive cellular activity.[515,516] Many challenges
still remain. Solid state light emission has yet to achieve longevity for at least a year when being used in vivo. Devices also
need to be tested for biocompatibility given the potentially reactive materials such as Ag, Al, ITO, GaN, AlGaNP and many
other possible light emitting diode layer materials.
Incoherent µLEDs and coherent vertical cavity surface
emitting lasers (VCSELs) are rigid inorganic devices but still
have the potential to achieve a flexible form factor suitable
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Figure 9. Sample light delivery systems for optogenetic stimulation. a) Optroelectrode design using gallium nitrate (GaN) microlight emitting diodes
(µLEDs) and electrode recording sites stacked together using heterogenous packaging. Reproduced with permission[322] Copyright 2013, American
Association for the Advancement of Science. b) Cross-section along a monolithically fabricated µLED and microelectrode (dash-line on top inset,
bottom inset shows the layers along the cross-section). The electrode typically consists of low impedance materials like iridium deposed on layers of
SiO2 (green). Gallium nitride (GaN) multiquantum wells (MQW) are connected to an anode and cathode to form the LED and silicone shank is the
base layer. Reproduced with permission.[514] Copyright 2015, Elsevier. c) Organic LED (OLED) layers have also been recently demonstrated in vitro
using these dimensions. Potential failure of LED-based illuminating systems is the possibility of failure of the deposited materials and water ingress to
leak the voltage to power the LED into the surrounding tissue. This will electrically stimulate the tissue instead of optically. Reproduced with permission.[516] Copyright 2016, the American Association for the Advancement of Science. d,e) Another prototype of optrode where a waveguide is used to
deliver light. e) Waveguide cross-section showing the transmission, cladding, and silicon substrate layers. Reproduced with permission.[517] Copyright
2015, Nature Publishing Group.
for the surface of the brain or conformal wrapping around
a nerve.[518,519] To do so, one must thin the light source and
heterogeneously package each component onto a substrate to
provide flexibility and insulation.[322,520] The challenge here is
to electrically bond the light source to a substrate, eliminate
any voids at the interface, and apply a pin-hole-free conformal
coating.[521] In practice, this approach usually fails because of
defects at sharp edges and voids at interfaces.[522] The usual
failure modes of all implantable devices are applicable here
including water and ion ingress, hydrolysis, and delamination. Since these are power consuming elements, the failure
is accelerated from the electric fields generated with every
pulse of current used to generate the light (Figure 9).[109] To
be practical in human use, one must demonstrate stability
after millions or billions of pulses and many years of soaking,
which to date has not be done in any optogenetic stimulation
system.
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OLEDs have the potential to be highly flexible, conformal,
and multicolor. However, they are also the most challenging
since they are additionally prone to fail in the presence water
vapor or oxygen, which have high transmission rates in polymers.[523] Still, consumer demand for flexible displays and
wearable electronics has spurred large R&D investments in this
technology. The most common approach to reduce the permeability of water and oxygen without compromising flexibility
is the deposition of a hybrid stack of alternating organic/inorganic layers (Figure 9c).[523] Atomic layer deposition of Al2O3
and ZnO with low water absorption polymers (e.g., Barix) is
one example of the hybrid approach.[516,524]
If the light source and electronics must be implanted and
designed for long-term use, then hermetic packaging is currently the most feasible solution to prevent in vivo failure.[522,525]
A hermetic feedthrough for light is also required, which is
another material challenge and unlike electrical feedthroughs is
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not a mature medical technology. Usually hermetic enclosures
are large and must be positioned many millimeters or centimeters from the therapeutic target, and a waveguide must be engineered with high transmission and biostability.
The core and cladding of a waveguide are equally important
material choices and both have an optical requirement and lifetime requirements. The index of refraction is the most critical
optical parameter and should ideally maximize the transmission efficiency (and reduce Fresnel losses if covering either end
of the waveguide). Some common waveguide core and cladding
combinations that may be lithographically patterned include
SU8/PMMA,[526] SU8/tissue[527] polysiloxane/polysiloxane,[528]
and SiON/SiO2 (Figure 9d,e).[529,530] A simpler manufacturing
technique is fiber pulling using polymer/polymer and many
forms of fused silica. Unlike the microfabricated waveguides,
these can be combined as bundles and that may increase flexibility. Conversely, the microfabricated systems are particularly
useful for multimodal device design like optoelectrodes and
drug delivery.[517] One group has also developed a technique
for co-polymer-metal extrusion so that the tip becomes a moderate-density optoelectrode tool.[269] Many questions remain
about all of these optical systems. A few studies have tested
SU8 and EpoCore and found optical losses become stable
out to 3–6 months of soaking[326,527] but longer testing is still
required and for each optical system. Polymer fibers (extruded
cylinders) have the most potential as a flexible waveguide core,
which would enable a similar form factor as used in conventional neuromodulation leads. When combined with a hermetically packaged laser diode and accompanying electronics, it is
similar to an IPG and flexible lead, e.g., spinal cord stimulator
and DBS stimulators. A stylet could provide the stiffness to
reach the spinal cord or deep brain structure, then is removed
to allow the waveguide to move more naturally with the tissue.
An optical waveguide, especially one designed for peripheral
nervous system, should be flexible in all directions so as to
minimize tissue reactivity. Polymer fibers are typically made of
polycarbonate, PMMA, polystyrene,[531] and Cytop (fluorinated
polymer) but only Cytop has good transmission between 400
and 550 nm.[532] Regardless, termination to a hermetically packaged laser diode remains a major challenge.
Another constraint of optical stimulation is heat generation.
Light requirements usually depend on the tissue volume and
distance from the light source. A Monte Carlo simulation tool
available online provides an excellent resource for predicting
heat generation in tissue due to light absorption and irradiance
profiles as a function of the waveguide diameter, optical power,
and wavelength.[533] Fortunately, the duty cycle and pulse width
for ChR2 stimulation is typically very low (e.g., 5%–15% duty
cycle, 2–5 ms duration) and can become even lower at shorter
pulse durations. Recent work demonstrated that even when
using pulse-width modulation within this small time window
the electrophysiological effect is equivalent and therefore power
consumption can be significantly reduced.[534] When optical
pulses are in the range of 1 s however, then the rate of heat
dissipation is no longer high enough to prevent local tissue
heating. Additionally, heat from the light source and any nearby
electronics inside the implanted electronics enclosure will have
to be added to the radiation heat. Although the heat deposition
at these levels might not adversely impact tissue, it is sufficient
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to elicit changes in tissue perfusion to scavenge increases in
heat.[535,536] One recent report stated a maximum energy of
0.5 mJ to avoid these physiological responses.[535]
Lastly, there is increased desire for electrode materials that
are transparent in tissue for optical stimulation through an
electrode, read-out of fluorescent neural activity sensors, and
compatibility with optical imaging modalities. In order to image
through a probe, the material needs to be both transparent, and
have the same refractive index as tissue. This is particularly
true with multiphoton microscopy and holography, where nonuniform refraction of light can alter the convergence of the twophoton light path. This can be addressed with uniform “sheet”
or “slab” arrays such as an ECoG array, though these devices
are prone to dense fibrous encapsulation quickly after implantation (Section 2.12). Transparent electrode site and electrical
trace materials have also become desirable. Thin-film indium
tin oxide (ITO) is a mostly (≈80%) transparent conductor that
has been gaining attention for optogenetic applications.[537,538]
More recently, graphene conductors with ≈90% transparency
have also become a popular electrode material for optogenetic
applications due to improved strength and lower resistivity.[80,81]
However, even with one to four layer thick graphene, these
transparent materials generate photoelectric artifacts[80] (see[182]
review). Even when light source placement is carefully considered, tissue is a very turbid medium that readily scatters photons. The resulting photoelectrical activation can impact the
detected signal[182] and can even cause photoelectrochemical
Faradaic redox reactions if proper material selections are not
made.[539] A long-term challenge for ITO with in vivo optogenetic application is that ITO oxidizes overtime, which results
in darkening of the film and a dramatic rise in impedance.[540]
Development of new materials for optogenic stimulation and
detection need to consider these additional challenges beyond
the electrode consideration discussed in Sections 2.1–2.14.
Although numerous issues involving optogenetics are
highlighted, its future is bright especially for basic-science
researchers mapping neural circuits. Solutions for long-term
use of optogenetic neuromodulation will require new advances
in materials to provide robust electrical insulation, chemically
stable, flexible, and safe light delivery via LEDs or waveguides,
and for some applications, near-transparent materials. The
opsin transduction challenges will also need to be resolved
along with the regulatory challenges, and the therapy will have
to be highly efficacious relative to electrical stimulation or pharmaceutical interventions for its long-term translational viability.
3. Challenges and Opportunities for NeuralInterfaces Located Outside the Brain
Although implants in the brain often capture the academic and
public imagination, neural interfaces for use outside the brain
are by far the largest current therapeutic market. The cochlear
prosthetic for deafness is one of the earliest neural interfaces
FDA approved in the United States. Vagal nerve stimulators
for epilepsy, spinal cord stimulators for pain, and Finetech–
Brindley stimulator bladder control stimulators were all FDA
market approved well before DBS for Parkinson’s disease. Currently, the single largest implantable neural interface market is
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the epidural spinal cord stimulators for pain. However, recent
FDA approvals include: the Argus II Retinal Prosthetic from
Second Sight,[541] the CVRx Rheos Barostim system for hypertension,[542] the Inspire hypoglossal nerve stim system for sleep
apnea,[543] the Enteromedics abdominal vagus block for obesity,[544] the Medtronic InterStim sacral nerve stimulator for urinary/fecal incontinence,[545] and the StimWave percutaneously
deliverable and wirelessly powered spinal cord stimulation
system.[546]
In general, many of the same design constraints and material trade-offs for neural interfaces in the brain apply to neural
interfaces placed outside the brain (Figure 10), but the risks/
benefits of each trade-off changes for the specific therapeutic
application and implant location. More severe side effects such
as intracranial hemorrhage leading to paralysis, stroke, and
seizures that are risks in brain implants are dramatically less
prevalent in the periphery, and replaced with other therapy
limiting side-effects such as nerve damage leading to motor
dysfunction, unintended muscle activation, bradycardia/tachycardia, dyspnea, pain and paresthesia. Whereas devices such as
the CVRx baroreflex activation system or the Finetech–Brindley
bladder stimulator work very well as a single minimally invasive
stimulating electrode or bipolar/tripolar epineural cuff, systems
intended to generate multifaceted motor signal decoding and
tactile/proprioceptive feedback for control of an artificial arm
may require dozens if not hundreds of precisely located stimulating and recording electrodes. Similarly, an indwelling longitudinal intrafascicular electrode that violates the epineurium,
or the flat interface electrode (FINE) which redefines a nerve’s
geometry by flattening the nerve to allow for closer contact with
central nerves may be completely suitable for use in the axon
fibers.[547,548] However, penetrating electrodes increase the risk
of disrupting the blood–nerve barrier (BNB), which is essential
in circulating oxygen and nutrients to individual nerve fascicles.[549–551] Likewise, FINE or book-cuff electrodes have been
shown to increase BNB permeability,[552] which allows for the
leakage of inflammatory plasma proteins into the extracellular
space.[545] Additional design considerations are necessary to
minimize device related trauma. More recently, regenerative
peripheral nerve interfaces (RPNIs) are currently being developed to provide seamless integration of transected peripheral
nerve with electronic interfaces by minimizing fibrotic encapsulation, promoting revascularization, and improving signal
transduction for neuroprosthetic control.[553,554] Similar to
Figure 10. Device considerations and challenges encountered during electrode design in the central nervous system and peripheral nervous system.
Axes rank design components from lowest to highest priority for both PNS and CNS devices. Note that device challenges between the two systems are
not mutually exclusive. For example, foreign body response is not as disruptive in the PNS (moderate priority) than the CNS (high priority).
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targeted reinnervation, when the nerve terminates on a muscle
graft then the action potential is transduced into a large amplitude coordinated muscle response that is read out as an electromyogram. However, unlike targeted reinnervation, RPNIs treat
even small diameter nerves in situ using small muscle grafts
and so can potentially achieve a large number of afferent and
efferent interfaces.
Stresses on the neural interface and leads connecting to the
implantable pulse generator (IPG), as well as relative movement
of the targeted tissue with respect to the implanted electrode,
also vary greatly based on location of the implant. Epidurally
placed spinal cord stimulation electrodes for pain are known to
move relative to the underlying spinal cord during changes in
postures, motivating strategies to anchor and stimulate on the
dorsal root ganglia to provide more consistent therapy. Spinal
cord stimulation systems, cochlear implants, and occipital
nerve stimulators, suffer most frequently from electrode displacement, extrusion, or breakage.[555–559] Relative movement of
nearby tissue inducing migration is less of an issue for vagal
nerve stimulators or carotid sinus nerve stimulators, but insulation to direct current preferentially to the underlying nerve of
epineural cuffs is more important. Spiral cuff electrodes, such
as those used to stimulate the vagus nerve in epilepsy patients,
allow for multiple contact points with a flexible geometry that
wraps around the nerve and tethers it in place.[560–563] Efforts
to reduce migration of PNIs involve device redesign such as
suture pads along the lead or mesh electrodes which provide
additional points of suturing to the fascia, connective tissue
that encloses nerve and muscle tissue.[564] Modifications to surgical techniques have also seen success, such as the use of a
silicone elastomer adhesive to cement the electrode to the nerve
cable.[565] Electrodes placed on or in peripheral nerves in the
arm to record/stimulate for prosthetic control may have leads
that are tunneled back over the shoulder joint back to an IPG
or wireless base station for control, and relative movement
across that joint is particularly problematic for wear and tear on
tissue and lead insulation. Moreover, electrodes placed outside
the brain and tunneled back to an IPG tend to be much more
consciously perceptible to patients, leading to a host of “irritation” related patient manipulation of the electrode, tunneled
lead or pulse generator. For example, there is a common condition known as “IPG Twiddler’s Syndrome” where the patient
attempts to spin the IPG on a 2-suture axis within the subcutaneous pocket, which has later led to 3- and 4-suture fixation of
the IPG.[566–568] Similarly, one of most common failure points in
the cardiac pacemaker or DBS system is at the ISO-1 connector
(International Organization for Standardization) which connects the electrode lead to the IPG. The mechanical mismatch
between the hermetically sealed titanium housing of the IPG
placed subcutaneously or submuscular in the chest, the bore
in the RF-transparent header section, and the lead insulation
wears the insulation over time, slowly exposing the underlying
trifilar coil leading to extraneous stimulation of the muscles/
nerves in the chest.
Although more clinically available, PNS devices face challenges similar to those faced by CNS systems. Flexible materials are required to bend with macromotion and retain fit to
nearby anatomy. Greater stresses on implants also demand
more mechanically robust materials, conductors that will not
Adv. Funct. Mater. 2017, 1701269
fail due to fracture, insulation that will maintain closure, and
traces and connectors that can withstand stresses induced
by mechanical mismatch. More creativity is required in
defining device geometry, in order to complement the specific
strengths of the materials involved. All of the considerations
regarding chemical stability of exposed materials also apply as
modified for a peripheral environment. Ultimately, materials
for PNS devices must work synergistically to maintain functionality within some of the most active regions of the body.
4. Regulatory, Reimbursement, and Other
Practical Considerations
Despite numerous advances in electrode fabrication processes,
electrode and insulation materials, and bioactive coatings within
the academic realm, existing clinical deep brain stimulators,
spinal cord stimulators, and vagal nerve stimulators still draw
heavily from the designs, materials, and underlying concepts
from the cardiac pacemaker/defibrillator. There are two primary
drivers for the slow speed of adoption of new fabrication processes, passive materials, and biologically active materials. The
first is the large jump from low volume animal studies across
a few academic labs, to widespread clinical use in hundreds of
thousands of patients at facilities across the world. Even the
cardiac pacemaker/defibrillator, which is implanted in 900,000
patients a year worldwide and relies on a simple surgical procedure and well understood materials that have been in practice
for 50 years, has a significant rate of device failure. The comprehensive FOLLOWPACE study, which followed over 1500 pacemaker patients over an average of 5.8 years, found that 12.4% of
patients developed pacemaker complications within the first two
months after surgery, while 9.2% of patients experienced complications after two months.[569] Many of the novel materials discussed in this review such as conductive polymers and carbon
fibers are manufactured using processes where the intermediate
stages of the material may be toxic, or where potentially toxic
substances are used to treat the material to achieve the final
result. In an industry supply chain designed to manufacture a
product at scale, many of the manufacturing steps or constituent
materials are provided by an outside supplier. When manufacturing at the scale of 10,000 implants per year, it can be difficult
for the end-system manufacturer to guarantee that the supplier
has not changed their process in some small way that might
accidentally introduce a leachable toxic substance into material
in contact with tissue. Moreover, simple unknowns like shipping
the electrode coated with IrOx overseas, or understanding the
potential consequences if a surgeon drops the sterile package
with the electrode and leads before implantation have to be carefully understood for any change in design. Finally, the introduction of biologically active components can greatly complicate the
regulatory pathway. The “combined” devices often have to go
through the FDA Office of Combination Products and review
from both the FDA Center for Devices and Radiological Health
and the FDA Center for Biologics Evaluation and Research, dramatically increasing the regulatory burden and bar for safety.
Given the biological component, there is understandably a concern about variability in manufacturer process/supply chain/
handling leading to unforeseen consequences in the patient.
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Table 1. Constraints on electrode device design and ideal solutions for recording and stimulating applications. Constraint columns represent
opposing extremes for different engineering parameters.
Constraint
Ideal
Constraint
Conductor
Low conductivity
High electrical resistance
Ohmic Heating
Infinite conductivity
Zero cross-sectional area
Electrochemically stable materials
Flexibility
Mechanical strength
High cross-sectional area
Increased stiffness
Low stretchability
Increased device footprint
Dielectric thickness
Ultrathin
Low dielectric strength
Dielectric constant of vacuum (1)
High shunt capacitance
Greater impact from water absorption
Electron tunneling
Parasitic capacitance
No capacitance
No water absorption
Low dielectric constant
Self-healing
Thick Insulation
Large device size & footprint
Increased susceptibility for delamination
Increased stiffness
Site size
Very small
High impedance
Charge density safety limit (stimulation)
Zero surface area electrode site
Zero impedance
Very large
Signal is distributed and attenuated
Large activation area (stimulation)
Site coating
Resistive materials
Corrosion
Material dissolution
High ESA:GSA ratio
High charge injection capability
Non-Faradaic charge transfer
Good biocompatibility
Capacitive materials
Low charge injection capabilities
Size and Geometry
Small cross-sectional area
Very flexible
Poor Strength
Low device volume
High device strength
Large cross-sectional area
Increased stiffness
Increased tissue strain
Increased glial encapsulation
Volumetric density
Low shank density
Discontinuous sampling volume
Spatially continuous recording/stimulating radius
Large spacing between electrode
shanks
Low tissue strain
High shank density
Increased tissue strain BBB disruption
Meningeal gliosis
Flexibility
Very stiff
Mechanical mismatch with soft tissue
Stiff enough to insert/handle
Flexible enough to not activate inflammatory mechanosensitive receptors
Rigid enough to minimize motionrelated electromagnetic artifact No
migration
Very flexible
Difficult to penetrate pia mater
Greater motion artifact
Device migration
Insertion
Slow insertion
BBB Compression and Disruption
Tissue Strain and Dimpling
No vascular disruption
No tissue strain
Controlled insertion speed and angle
Fast insertion
Ballistic trauma
Increased gliosis
BBB Severing
Very rigid
Mechanical mismatch with soft tissue
Activates inflammatory mechanosensitive
receptors
Elastic Modulus of tissue
Good conductivity at low crosssectional area
Good dielectric at thin insulation layer
High yield strength
Very soft
Low Conductivity
Low Dielectric Constant
Difficult to insert
Packaging
Rigidly anchored device
Increased tissue strain
Impact related trauma
Infection route
Unanchored to the skull
Wireless device
Flexibly Tethered device
Increased fibrotic encapsulation
Implant migration
Infection route
Motion artifacts
Molecular and Cellular
Pharmaceutical
Drug release side effects
Toxic polymer degradation
Hydrogels swell
No biofouling
No inflammation or scarring
No neurodegeneration
Neurotrophic, Neuroregeneration,
Neural Growth Factors
Neurocamouflaging/Neural Adhering
Surface
Biological
Hydrogels swell
Stem cell rejection by immune system
Proteins degrade quickly
Elastic modulus (softness)
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The second driver of the slow speed of adoption of new
materials/concepts/process into clinical use is the high cost in
both time and money to obtain market approval from the FDA,
and the realities of insurance reimbursement. A pivotal trial in
the United States for an implantable neuromodulation/neuroprosthetic device to obtain FDA premarket approval often takes
more than 5 years to complete and can cost well in excess of
100 million dollars. Moreover, obtaining FDA market approval
does not guarantee that insurance payers will reimburse for
the therapy, or reimburse at a price necessary to sustain a commercial manufacturer. As new materials/processes introduce
significant unknowns into the regulatory and reimbursement
pathways, the prospective reward for adopting these new concepts must be both highly likely and offset the large costs and
risks inherent in regulatory approval given the price for reimbursement set by insurance payers.
Despite these drivers slowing the adoption of new materials/concepts/processes, in 2015 RetroSense Therapeutics LLC
received an investigational new drug exemption from the FDA to
transduce a very limited number of human subjects in order to
restore light sensitive channels in the retina of patients blinded
by retinitis pigmentosis.[570] Although this is a Phase 0–1 study
to determine safety, and only ambient light is used for illumination, this is a critical step in moving toward optogenetic therapies.[571] The jump from animal models is quite large, however,
and demonstrating transduction, specificity of transduction, and
maintenance of transduction utilizing an adeno-associated virus
gene vector to enable optogenetic therapies in humans has yet to
be demonstrated. To date optogenetics has primarily been used
in genetically homogenous murine models, with no studies as
of yet demonstrating consistent specificity and maintenance of
transduction over a long period of time across a large cohort in a
large animal model. A study using adeno-associated virus ocular
subretinal injection to address c-MER proto-oncogene tyrosine
kinase (MERTK) mutations in six human patients blinded by
retinitis pigmentosis, for example, failed to demonstrate any
transduction in three of the patients, and only one patient maintained expression out to 2 years.[571]
Figure 11. Complexity of Functional Neural Interface Engineering. There is a complex network of interdependence between design properties and
material constraints, which must be carefully considered when engineering functional neural interface devices. Optimizing a single material property
can often negatively impact another material characteristics directly or indirectly through downstream relationships. This can push the overall design
outside functional performance constraint windows, ultimately resulting in a nonfunctional device. Therefore, neural interface engineering requires
precise balancing of between intrinsic and extrinsic material properties, and their impact on overall performance. Bolded words represent tunable
design parameters, and italics denote influencing factors or constraints. Functional performance characteristics are listed in the center, where the
biological, electrical and mechanical domains overlap. Blue and red arrows represent positive or negative correlations between terms. Abbreviations
include blood–brain barrier (BBB), PC (pseudocapacitive), µ/n (micro/nano surface), electrochemical to geometric surface area ratio (ESA:GSA),
volumetric density of tissue recording (VD). Note: Impedance increase and decrease can both impact performance positively or negatively, depending
on the cause. See Table 1 for simplified form.
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www.afm-journal.de
Although the retina is attractive for initial demonstrations
of optogenetics (as photoreceptors can be illuminated without
an implantable device), more diverse therapeutic applications
will require implantation of yet to be developed human-ready
fiber optics or µLEDs for illumination. This will require a host
of developments to enable a safe, reliable, and cost-efficient
system at scale, as outlined in the excellent review “From
optogenetic technologies to neuromodulation therapies.”[512]
Even given a safe and reliable implantable optogenetic transduction and illumination system, clinical adoption will be key
for a successful optogenetics business case. At present only
5%–20% of the patients indicated for standard DBS for Parkinson’s elect to receive an implant; optogenetics will have to
demonstrate a large improvement over electrode technologies
to overcome hesitance to undergo both manipulation of the
genome in addition to surgical implantation of a fiber optic or
µLED system.
5. Conclusion
The complexity of neural interface design, coupled with regulatory and reimbursement hurdles, forecasts a challenging
road forward for medical device innovators. Understanding
how individual material properties contribute to the success
or failure of a neural interface has allowed researchers to
experiment with the optimization of specific design variables.
Although this has resulted in advancements in materials science and the current understanding of ideal design (Table 1),
major obstacles remain for translation to functional implantable devices. This is because single parameter optimization
alone is insufficient in producing long-term functional neural
interfaces, due to unavoidable tradeoffs that result from the
interdependence of design variables. For example, hydrogels can be engineered to have a similar elastic modulus as
the brain tissue, however they lack many other important
characteristics, including electrical and insulating properties, chemical stability, or material strength and durability.
Conductive polymers provide benefits of corrosion resistance, high ESA, and favorable electrical properties, but they
are lacking in mechanical stability, biostability, and softness.
And subcellular devices have favorable softness and flexibility,
and induce minimal strain on the surrounding tissue, yet they
suffer from high resistance and low mechanical strength. It
is clear that there is a complex network of interplay between
intrinsic and extrinsic material and design properties, and
the resulting biological, mechanical and electrical responses
(Figure 11). Therefore, the ultimate challenge in developing
functional devices is understanding interdependence of these
variables, and then carefully balancing and titrating these
variables with existing or emerging biomaterials. Alternatively, new modes of technology (e.g., wireless transducers)
could eliminate the need for some of these variables. These
innovations, however, require careful perspective and consideration to critically analyze what new limiting variables are
introduced with these new signal transduction modes. Only
through honest evaluation of limitations and failure modes,
as well as outliers, are we able to engineer transformative next
generation technologies.
Adv. Funct. Mater. 2017, 1701269
Acknowledgements
T.D.Y.K., S.M.W., A.L.V., and N.J.M. were supported by NIH
NINDS 1R01NS094396. K.A.L. was supported by The Grainger
Foundation. J.P.S. was supported by NIH (U01-NS090526-01) and the
NSF (1545858). A.L.V. was supported by NIH (R01-NS094404). The
authors would like to thank Dr. Aaron Batista Lab for providing the
histology samples used in Figure 7.
Conflict of Interest
The authors declare no conflict of interest.
Keywords
bioelectronics, electrodes, microelectromechanical systems, photonics,
sensors
Received: March 8, 2017
Revised: April 10, 2017
Published online:
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