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CRITICAL REVIEW

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Materials for microfabricated implantable devices:a review
Cite this: Lab Chip, 2015, 15, 4256

Kee Scholten and Ellis Meng*
The application of microfabrication to the development of biomedical implants has produced a newgeneration of miniaturized technology for assisting treatment and research. Microfabricated implantabledevices (μID) are an increasingly important tool, and the development of new μIDs is a rapidly growing fieldReceived 10th July 2015,
Accepted 15th September 2015that requires new microtechnologies able to safely and accurately function in vivo. Here, we present a
review of μID research that examines the critical role of material choice in design and fabrication. Materials
commonly used for μID production are identified and presented along with their relevant physical

DOI: 10.1039/c5lc00809c

properties and a survey of the state-of-the-art in μID development. The consequence of material choice
as it pertains to microfabrication and biocompatibility is discussed in detail with a particular focus on the

www.rsc.org/loc

divide between hard, rigid materials and soft, pliable polymers.

1. Introduction
Department of Biomedical Engineering, Univ. of Southern California, Los Angeles,
CA 90089-1111, USA. E-mail: ellis.meng@usc.edu

Kee Scholten is a postdoctoral
scholar in the department of Biomedical Microsystems Laboratory at the University of Southern California. Kee graduatedfrom the California Institute ofTechnology (Caltech) in May
2009 with a BS in Applied Physics. He completed his PhD in
Applied Physics at the Universityof Michigan under Professor
Edward Zellers in 2014, with afocus on opto/microfluidic senKee Scholten
sors for gas-phase micro-devices.
His research interests are in microfluidic sensors, and the synthesis of microfabricated devices with nanoscale structures.

4256 | Lab Chip, 2015, 15, 4256–4272

Biomedical devices intended for in vivo implantation,
whether diagnostic or therapeutic, face familiar challenges ofEllis Meng is Professor of biomedical and electrical engineering in the Viterbi School of Engineering at the University ofSouthern California. She is also
Dwight C. and Hildagarde E.
Baum Chair of the Departmentof Biomedical Engineering. She
received the B.S. degree in engineering and applied science andthe M.S. and Ph.D. degrees inelectrical engineering from the
California Institute of TechnolEllis Mengogy (Caltech), Pasadena, in
1997, 1998, and 2003, respectively. She directs the Biomedical
Microsystems Laboratory founded in 2004 which conducts
research on biocompatible polymer technology and micromachining, sensors and actuators, microfluidics, and implantableand biomedical microelectromechanical systems (MEMS). Her
honors include the NSF CAREER award, Wallace H. Coulter Foundation Early Career Award, 2009 TR35 Young Innovator Under
35, Viterbi Early Career Chair, and ASEE Curtis W. McGraw
Research Award. She is on the editorial board of the Journal of
Micromechanics and Microengineering and is co-chair of the 2017
IEEE MEMS conference. She is also an active educator andauthored a textbook on bioMEMS.

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Lab on a Chipsize, power consumption, biocompatibility and efficacy. Engineering devices with micron-scale precision enables significant reduction in footprint, encumbrance and power
demand, and an accompanying increase in device complexity
and often capability. By adapting semiconductor micromachining processes to the field of biomedical implants,
researchers have opened broad new avenues of research andpromulgated a generation of new implantable technologies,
so-called microfabricated implantable devices (μID). Examples are numerous and include physiological sensors, biochemical sensors, neural prostheses, cochlear prostheses,
drug delivery devices, ocular implants, and vascular stents.
For the researcher, μIDs present an elegant solution for problems requiring targeted manipulation or observation of living
tissue; for the patient, μIDs present an option for minimallyinvasive point-of-care treatment. For the engineer, however,μIDs present a series of interconnected design challenges;
device structure and fabrication method are constrained bymaterial choice, which must satisfy desired physical properties as well as required hermeticity and biocompatibility.
As the body of research on μID development grows,
designs increasingly take advantage of a greater variety ofmaterials, and numerous reports describe implants fabricated from unconventional materials, including polymers
and biopolymers (e.g. polymethylmethacrylate, polyethylene
glycol, chitosan), synthetic materials and synthetic composites (e.g. hydrogels, polyĲN-isopropylacrylamide)), and nanostructured materials (e.g. nanoparticles, nanowires,graphene). The nature of an expanding field such as this precludes the compilation of an exhaustive list of all materials
under investigation. However, for implantable devices thepath from preliminary work to translation is long and perilous,and material composition is often a major determinant. Government regulations governing approval for medical devices
are stringent and often inconsistent between countries, andthe requisite testing to ensure compliance can be laborious
and costly.1 Many materials which are successful in in vitro
experiments may encounter unforeseen complications when
implanted, including unexpected immune responses or slow
but appreciable degradation. For μIDs, these issues can bemagnified; the extremely thin layers of materials common in
micro-devices may degrade at rates too slow to accurately measure, but that nonetheless lead to device failure within patient
lifetime, and even minor immune response can be sufficient toisolate a micron-sized device and prevent proper operation.
This review will discuss the role of material choice in thedesign and fabrication of μIDs, with a primary focus on asmall group of well-vetted materials that dominates the literature and has therefore played a critical role in establishing
the history of the field. For ease of analysis, we have broadly
characterized those surveyed as either ‘hard’ or ‘soft’, theformer comprising high Young's moduli materials such as
silicon, glass, metals and ceramics and the latter comprising
low Young's moduli polymers including polydimethylsiloxane(PDMS), polyimide, polyĲchloro-p-xylylene) (Parylene C), andbiodegradable polymers. For engineers embarking on theThis journal is © The Royal Society of Chemistry 2015

Critical review

design and development of a new device, choice of material
may be the first and most critical decision; material properties dictate available fabrication and packaging methods,
device performance and chemical compatibility, and, forμIDs specifically, material choice strongly influences interaction between the device and host. This review will explore
how material choice influences and, at times, limits design
and fabrication of μIDs, and will describe the advantages andchallenges of working with different materials, with a focus
on the differences in approaches for soft and hard devices.
We will begin with an overview of the physical and chemical
properties of commonly employed materials and the available
microfabrication methods and tools. A survey of μIDs
described in the established literature, categorized by material composition, follows, with a focus on how material selection influenced device design and operation. Finally, we present a discussion of current challenges affecting use ofdifferent materials in μID development.

2. Material considerations
2.1 Biocompatibility
The requirement of biocompatibility, that the material beboth safe to an implanted host and able to function in in vivo
conditions, represents perhaps the strictest prerequisite on
materials for μIDs.2 Williams formalized the definition as“the ability of a material to perform with an appropriate host
response in a specific situation”,3 acknowledging varying
requirements on materials in separate sites and applications.
Biocompatibility is a more significant concern for implants
intended for chronic use, but even for acute applications,
exposed material must satisfy demands of both patient safetyand device efficacy. Materials are tested for toxicity and carcinogenicity, as well as any leached or degradation products,
and to determine the degradation, corrosion and dissolution
profile at the ambient temperature, pressure and salinity typical of the in vivo environment. Host response is strongly
influenced not only by the chemical composition of theimplant material, but also surface morphology, crystallinity,
and surface energy, and bulk mechanical properties such aselastic constants and shape (see section 2.2).4 For a moredetailed discussion of material biocompatibility, the reader isreferred to reviews by Williams.4,5
Biocompatibility requirements exclude many materials
common in microfabrication and microelectromechanicalsystems (MEMS) for μIDs; many metals corrode quickly inin vivo environments6 while polymers with low melting orsoftening temperatures, or high solubility in water, failmechanically. For μIDs intended for chronic implantation,
there are considerations of water and gas intrusion,7 fatigue
failure,8 and failure due to immune response.9–11 Scar formation and other foreign body reactions present a significant
obstacle to long-term implantation,9,11 and are a common
failure mode for otherwise successful μIDs. Common strategies to countering (or limiting) the immune response include
careful design and implant placement, minimizing size to

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reduce tissue damage, biomolecular coatings,12,13 and pharmaceutical approaches;14–16 however, choice of material
remains critical.17
Common examples of materials considered biocompatible
and used in μIDs include: cobalt-chromium, iridium, titanium, platinum, nitinol, certain glasses, PDMS, Parylene C,polymethylmethacrylate (PMMA) and polyimide.1,7,18,19 Silicon, polysilicon, SU-8 polymer, SiO2, Si3N4, and SiC have beentested by Kotzar et al. under regulatory guidelines (FDA ISO
10-993), and minimal biocompatibility issues were reported
for all but SU-8.20 Subsequent tests of implanted devices likewise suggest silicon to be a viable material for μIDs,9,21 however such data falls short of what is needed for regulatory
approval. Extensive data on histopathology for corticalimplants specifically was compiled by Stensaas andStensaas,22 and organizes materials as non-reactive (e.g. aluminum, ceramic alumina, gold, platinum and certain productions of polyethylene and Teflon), reactive (e.g. SiO2,
nichrome, TiO2) and toxic (e.g. Silastic RTV silicone, silver,iron and copper).
Due to concerns of electrochemical corrosion, biofoulingand irritation, micromachined devices are frequently encapsulated in conformal polymers, bonded to glass, or enclosed
in biocompatible metal casings.23 Encapsulation is a common approach even for devices comprising exclusively biocompatible materials, as an insulating or hermetic seal isoften required to prevent water and soluble ions from damaging or shorting electrical connections. Material choice for
encapsulants must satisfy not only conditions of biocompatibility, but also desired low permeability, ease of deposition,and conformality.

Lab on a Chip

example, the elasticity or rigidity of a material, typically represented by the Young's modulus, must be considered toreduce mechanical mismatch between the device and surrounding tissue, an important criteria of minimizing tissue
damage and foreign body response.10,24,25 While mechanically compliant devices may better match the mechanicalproperties of soft tissue, high aspect ratio devices, such as
micro-needles or probes, require sufficient compressive andtensile strength to survive insertion without mechanicalbuckling.26 In general, common μID materials are considerably stiffer than tissue at most implantation sites, however
the elastic moduli of both biological tissues and μID materials vary over several orders of magnitude (Fig. 1). A material
selection that is ‘too soft’ for one site or application may beconsidered ‘too hard’ for another. Mechanical properties ofmaterials can also influence critical feature dimensions,which can subsequently determine device shape andresulting biocompatibility. For example, Edel et al. determined that the sharpness of the tip of a microprobe could
dramatically reduce the size of the neuron ‘kill zone’ surrounding implants.27
Thermal conductivity is of relevance for μIDs which may
produce heating of surrounding tissue. Electrical conductivity
and permeability to gas and water can determine whether amaterial requires encapsulation, or is suited as an
encapsulant, while thermal stability (e.g. softening, melting
and degradation temperatures) determines whether a material will survive in vivo. Other properties may only be relevant to particular devices, for example, optical transmissivity
and optogenetic implants. Table 1 provides values from literature of important material properties for commonly chosen
μID materials.

2.2 Physical properties
As with any MEMS device, material selection must accountfor any necessary constraints on mechanical, electrical, thermal, optical and chemical properties, but there are often
requirements of material properties unique to μIDs. For

2.3 Fabrication technologies
Material choice determines the microfabrication tools andtechniques available for μID production, and consequentiallythe minimum size and resolution and possible geometries

Fig. 1 Logarithmic plot of elastic (Young's) moduli for various biological tissues and common μID materials. Data points are representative values
available in literature, shaded areas span full range of values for selected materials in referenced literature.223–227,231,241–245

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Critical review

Table 1 Material properties of silicon, glass, Parylene C, PDMS and polyimide, adapted from literature

Silicona
Young's modulus (MPa)

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d

150 000−2

Hardness
Processing temperature
limits (°C)

Glass

850 kg mm
Tmelt

1414

Tglass—Tdegredation —Density (g cm−3)
Thermal conductivityc (W m−1 K−1)
Electrical resistivity (Ω cm)
Dielectric constantGas permeabilityc
N2(cm3 mm m−2 day−1 atm−1) CO2
O2
H262 750[ref. 223]
418d kg mm−2[ref. 228]821d
[ref. 223]–1700e[ref. 231]
—
—

2.23d
[ref. 223]–2.27e[ref. 231]
156 [ref. 235]
1.15d
[ref. 236]–1.4e[ref. 231]
2.3 × 105 (intrinsic) 1014d [ref. 238]
[ref. 237]
1016e[ref. 231]
11.9
4.6d [ref. 223]
3.9e [ref. 231]
—
—
—
—
—
—
—
—

2.329

Parylene Cb

PDMS

2760

0.360–0.870 [ref. 224]

90
[ref. 232]
125
[ref. 233]
1.289

—

0.084

0.168 (14.7 °C) 0.159 (50 °C)[ref. 226]

8.8 × 1016

2.4 × 1014–1.5 × 1015 [ref. 226] >1016 [ref. 227]

2.95–3.15

2.69–2.77 [ref. 226]

2.9–3.5 [ref. 227]

0.4
2.8
3.0
43.3

1.8–2.6 × 104 [ref. 226, 239]
2.1–2.5 × 105 [ref. 226, 239]
3.9–5.3 × 104 [ref. 226, 239]
4.3–5.8 × 104 [ref. 226, 239]

2.3 [ref. 226]
17.3 [ref. 226]
9.6 [ref. 226]
96.3 [ref. 226]

Polyimide

2300–8500[ref. 225–227]
80
30–80 (Shore A) [ref. 226, 229] 50–99 (Rockwell E)(Rockwell B)[ref. 230]290——

350 (oxidation) 750(decomposition) [ref. 234]
0.970–1.11 [ref. 226, 229]

325—410[ref. 225–227]
510–620[ref. 225–227]
1.39–1.42[ref. 226, 227]
0.12 [ref. 226]

a

Values adapted from ref. 231 unless otherwise stated. b Values adapted from ref. 240 unless otherwise stated. c Values for 25 °C unlessotherwise stated. d Values for Pyrex® borosilicate glass. e Values for SiO2.

that can be produced during fabrication. Different methods,
each with their limitations and idiosyncrasies, are available
for depositing, etching and patterning different materials,
therefore, material selection must be coincident with the initial design phase and accompany any discussions of criticaldimensions as well as outlines for process flow.
The majority of μIDs are fabricated partially, or entirely,
through use of standard cleanroom processing. Thisapproach offers several advantages including compatibility
with production of integrated circuitry, sub-micron precisionin machining, and often batch-scale fabrication. Thoughmany devices require manual packaging or final assembly
before deployment, the parallelization of cleanroom processing enables large variable experimentation for academic
research and low costs for commercial efforts. A key advantage of cleanroom processing is the large number of
established protocols and equipment for microfabrication,
developed over decades for the production of semiconductor
and MEMS devices. At the same time, adherence and compatibility with established protocols may impose unnecessary
limits on μID design. Common μID materials such as PDMS
and other polymers may not be ‘clean’ enough for use inmany facilities, while common cleanroom materials (e.g. silicon) have come to dominate the μID field because of theirubiquity among cleanroom microfabrication.
Alternate micromachining methods include direct laser
milling or writing, which has been utilized in metal andpolymers.28–32 Typical processes have resolutions between 10
and 100 μm for fabricating features, while surface

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morphology can be altered at the nanometer scale using laser
exposure.33 Direct writing requires serial processing, which
can increase cost and inter-device variation, but allows for
rapid prototyping and may remove the need for cleanroom
facilities. Of the solid freeform fabrication methods, only
stereolithography offers sufficient resolution to build microfabricated structures,34 and there is limited work detailingthe biocompatibility of the specialized resins required.35,36
Other research utilizes ad hoc fabrication techniques, forexample the Anikeeva group has produced optrode neural
implants using a repurposed optical fiber-pulling tower.37,38
Such approaches often yield excitingly novel devices, but the
fabrication methods are rarely generalizable beyond the original application. In any case, choice of material will largely
narrow the available microfabrication methods, as amenability to microfabrication is a key consideration when selecting
materials for production of μIDs. Below, we delineate someof the most common methods available for micromachiningof biocompatible μID materials.
Several methods for depositing thin films of metals,
including those common biocompatible choices listed above,
are ubiquitous in microfabrication facilities. Evaporation,using electron-beam or heated coils, and sputtering offer
batch-scale deposition of metal films up to hundreds ofnanometers thick, with single nanometer precision.2 Suchfilms can be patterned using lithographic shadow masks,
with resolution limited by the lithography procedure and surface roughness of the targeted substrate. Thicker metal structures (up to several mm) can be deposited throughLab Chip, 2015, 15, 4256–4272 | 4259

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electroplating, however this requires electrically conductive
substrates and can result in rough surfaces or other defects.2
Materials such as silicon, silicon nitride, silicon dioxide, andsome polymers (e.g. Parylene), can be deposited conformallyby way of chemical vapor deposition (CVD).2,39 Thoughnumerous different incarnations of the process exist, mostrequire high temperatures and low pressures and all require
an available chemical precursor. CVD processes offer a highdynamic range of thickness, from monolayers to tens of
microns, high uniformity and excellent conformality, but
may not be compatible with all substrates or materials, andare not amenable to shadow masking.39 Polymeric materials
such as PDMS and polyimide can be deposited in wet processes, or cast from solvents or uncured states.40 Spincoating offers highly uniform, but non-conformal, coverageof substrates at thicknesses ranging from single to hundredsof microns. Greater conformality can be afforded by casting
directly from an uncured state, but uniformity is sacrificed.40
For those materials not easily patterned during deposition, the availability of useful etching techniques is critical. A
large variety of tested techniques exist for silicon etching,including isotropic and anisotropic wet etchants, isotropic
gas etchants, isotropic plasma etchants, and deep reactive
ion etches (DRIE), many with high selectivity and etch rates,
facilitating surface and bulk micromachining for MEMS fabrication. However, for other materials common in μID development, options are often limited. Glass can be etched isotropically by HF and anisotropically with DRIE processes
using SF6 or C4F8,41–43 but etch rates for the latter are considerably slower than those achievable with silicon. Polyimide
and Parylene are typically patterned with reactive ion etches(RIE) and oxygen containing plasmas,44–49 as wet etching, inmost cases, is not feasible.46,50 Typical masking materials(i.e. photoresists) etch at similar rates, unfortunately withselectivity as low as 1 : 1 and thereby limiting etch depth.47
DRIE of Parylene can be achieved with cycles of SF6 and O2
plasma followed by C4F8, but reported aspect ratios are poor
compared to silicon DRIE.51 PDMS is notoriously difficult toetch, due to its high chemical inertness. Dry etching methods
exist that use reactive ion or microwave generated plasmas of
CF4 and O2 mixtures,52,53 however many of these methods
suffer from slow etch rates and poor aspect ratios, and often
yield rough surfaces.
Few MEMS or μIDs comprise a single homogeneous material, as such compatibility with other materials and material
processing, strength of inter-material adhesion and availability of bonding methods is a chief concern. For materials
requiring a biocompatible or hermetic seal for chronic
implantation, but are nonetheless commonly used in μIDs
(e.g. silicon), adhesion strength and durability withencapsulants is critical. For silicon μIDs, several reports havedescribed adhesion with conformal coatings of silicone rubbers,54,55 Parylene,55–59 and polyimide,60 as well as methods
for creating long-lasting hermetically sealed devices withanodic bonding to glass.59,61,62 An excellent comparative analysis by Treager of encapsulant efficacy against moisture4260 | Lab Chip, 2015, 15, 4256–4272

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intrusion highlights the superiority of metals and glasses
over organic adhesives and silicones.63 A thorough review
specific to silicon μID encapsulation has been compiled by
Wasikiewicz and Roohpour.55 For polymeric materials adhesion strength and surface area has been noted to improvewith plasma exposure or the use of adhesion promotors suchas silane A-174.56 Inter-material bonding and compatibility isalso important for hybrid devices, where multiple materials
serve structural purposes beyond encapsulation; these material interfaces include glass-silicon,64–67 glass-Parylene,68 silicon-polyimide,69 and silicon-Parylene.70 Reliability of intermaterial bonding is also critical for packaging and assemblyof microfabricated components. For a detailed description of
established methods for micropackaging and microassemblyof μIDs the reader is directed to texts by Schuettler and Stieglitz,71 and Inmann and Hodgins.72

3. ‘Hard’ materials for microfabricatedimplantable devices‘Hard’ materials, representing those non-polymeric materials
with high Young's moduli (>104 MPa), high hardness (>102
kg mm−2), high working temperatures (>500 °C), and effectively zero gas or liquid permeability, include silicon, glass,ceramics, and metals, the most common materials used in
MEMS production and historically the first materials used for
producing μIDs. As early as 1970, researchers were adapting
semiconductor manufacturing processes to create simple
implantable electrodes for recording neural activity.73 Following the onset of MEMS research and development, silicon
micromachining was used to produce the first μID pressure
sensors,74,75 drug delivery devices,76,77 and neural electrodearrays.78,79 Glass and metal structures became commonlyused in conjunction with silicon μIDs, providing biocompatible or hermetic encapsulation, or serving as mechanical orelectrical components. More recently microfabricated devices
comprising only glass or metal structures have grown in popularity, owing to new fabrication methods such as glass
reflow,80 and laser milling.29 Due to widespread familiarity
with silicon micromachining, and the importance of compatibility with CMOS processing and integrated circuitry (IC), silicon remains the most common choice of material for μID
fabrication. Hard materials, in general, are favored for theirmechanical strength, large thermal budget, and impermeability to liquid intrusion.
Hard μIDs are typically produced by way of bulk and/or
surface micromachining of preformed substrates. Bulk micromachining typically relies upon chemical etchants or physicalmachining methods to produce mechanically robust, highaspect-ratio three-dimensional structures for implantable
MEMS, while surface micromachining methods include UV
and electron-beam lithography, thin-film deposition (e.g.
CVD, sputtering and evaporation of metals, dielectrics andpolymers), oxide growth, and dopant implantation, withapplications for patterning planar devices such as electrodes
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methods, an impressive number of hard μIDs have beendeveloped and tested: electrochemical and optical probes forstimulating and recording neural activity, drug delivery
devices, physical, chemical and electrical sensors, and prostheses. This section will review microfabrication approaches
for a popular subset of hard substrates, namely silicon,
glasses, ceramics and metals, as they pertain to the development of μIDs, and examine the challenges of using suchmaterials.
3.1 SiliconThough best known for its semiconducting properties, silicon
is commonly used in μID production as a mechanical material or bulk substrate due to the ease with which it can be
micromachined with high resolution and repeatability. Several μIDs produced through silicon micromachining are

Fig. 2 Examples of microfabricated implantable devices machinedfrom silicon. (a) Intracortical recording electrode array: reprinted fromref. 91 © Elsevier 1998. (b) KOH etched silicon drug release microchip:
reprinted from ref. 9 © Elsevier 2003. (c) Cochlear prosthesis
stimulating electrode (inset) connected to application specificintegrated circuit: reprinted from ref. 94 © Elsevier 2008.

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Critical review

presented in Fig. 2. Wet etching of bulk silicon through lithographically patterned mask layers provides a low-cost, batch
scalable method for fabricating a variety of desired structures
on a micrometer size scale. Examples of such structures
include simple cavities. Several variations of wireless μID
pressure sensors comprise silicon cavities vacuum sealed bysilicon–silicon fusion bonding or silicon-glass anodic
bonding.65,66,75,81–85 The design, adapted for biomedical use
by Bäcklund et al. in 1990,75 consists of a vacuum sealed cavity
enclosed on at least one side by a thin membrane; deflectionsin the membrane due to changes in physiological pressure aremeasured capacitively65,66,75,81–84 or piezoelectrically.85 The
design is robust and versatile, facilitates simple integration ofplanar metal coils for wireless monitoring, and was successfully employed for monitoring intraocular75,83 and cardiovascular pressures.82,84,85
A similar fabrication approach was used to produce reservoir cavities for implantable drug delivery/releasedevices.9,77,86–89 Santini et al. used KOH, an anisotropic etchant, to create implantable arrays of square pyramidal microreservoirs sealed with individually addressed membranes that
could be electrochemically dissolved for controlled, in vivo
release of therapeutics such as polypetides.77,89 Elman et al.
used similar KOH etched reservoirs with release driven byelectro-thermal destruction of the sealing silicon nitride
membrane; the design has shown success in intracranialdelivery of chemotherapeutics in rodent models.86,87 Thesedesigns allow for large numbers of individually addressable,
low-volume reservoirs in a solid-state device, though the reliance on concentration based diffusion makes controllingdosing rate difficult.
Micromachining of silicon is also used to fabricate highaspect ratio shanks, as a structure for penetrating neuralprobes. μID neural probes comprise long and thin support
structures patterned with conductive electrodes for electrochemical recording and/or stimulation of brain tissue or
peripheral nerves. There has been considerable work on silicon neural probes and neural probe arrays, with the devices
known as the ‘Utah array’ and ‘Michigan array’ being themost successful. The ‘Utah array’, an intracortical microelectrode array composed of metal-tipped silicon needles, is fabricated by isotropically etching mechanically-cut silicon columns in a series of acid baths (5% HF, 95% HNO3).79,90,91
This process tapers the columns into thin needles with sharp
tips, enabling penetration into the neural cortex. The ‘Michigan array’ consists of electrode coated microprobes fabricated from stacked layers of boron doped silicon, SiO2, Si3N4,
and patterned metal traces, released from a silicon wafer by
etching away the undoped silicon with ethylenediamine-pyrocatechol.78,92 This is a versatile method that enables fabrication of very thin devices; in one report, the approach was
adapted to produce a flexible electrode array for a cochlear
prosthesis.93,94 These methods provide for batch-scale, parallel fabrication of high-aspect ratio structures, with precisely
controlled shank dimensions and electrode impedances andplacement. The selection of silicon enables the fabrication of

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sharp probes that easily penetrate brain tissue, and eases
connection to ancillary electronics.
In addition to wet etchants, dry silicon etchants, including
isotropic XeF2 gas, ionized plasma, and highly anisotropic
DRIE methods (e.g. Bosch processes), offer controlled micromachining with high selectivity for many masking materials.95,96 Dry micromachining of silicon avoids issues inherent to wet fabrication such as stiction, and DRIE enables
improved control of etch profile, however these processes frequently require serial processing and expensive instruments.
These techniques are increasingly used to produce highaspect ratio silicon structures with sharp features for MEMS
devices, which includes several examples of μIDs such as
microelectrode shanks for neural recording97,98 and implantable micromechanical pumps for drug delivery.99
In addition to its role as a support substrate or mechanical material in MEMS devices, silicon is the dominant choiceof semiconductor for fabrication of application specific integrated circuitry (ASIC) for μID controls. Use of silicon ICs inbiomedical implants predates the development of implantable MEMS by some years, and examples include conventionally fabricated devices such as pacemakers, neuro- and muscular-stimulators, and middle ear and cochlear auditory
prostheses.100 Inclusion of silicon ASICs in μIDs has becomea critical component of powered devices, particularly thosefor chronic implantation requiring circuitry for wireless communication and battery charging, in addition to control anddata logging. A more thorough discussion of IC technology inimplants, and biotelemetry in particular, can be found inreviews by Ko et al.100 and Receveur et al.101 ASICs and transducers are frequently fabricated separately then packaged
together, and as such choice of other materials for μID fabrication does not preclude use of silicon controls. However,direct integration of control circuitry is possible with silicon
μIDs, and offers reduced device profiles.102,103
3.2 Glass
Glasses, including borosilicate glasses such as Pyrex, quartz,
thermally grown or deposited SiO2, and silicon oxynitride, arefrequently used for sealing or encapsulating silicon μIDs dueto their high impermeability, chemical inertness, and biocompatibility. In addition, glasses and oxides are excellent
electrical insulators, which is desirable for wirelessly interrogated μIDs that require electrical insulation of microfabricated transmission coils or antennae. Biostability, however, can range greatly depending on specific compositionand even deposition method,104 and several glasses have highdissolution rates in saline. Dissolution rates in vivo have beenmeasured as low as 0.33 nm per day for CVD silicon nitride,and as high as 3.5 nm per day for CVD silicon oxide.105 In
tests of subretinal implants, a TEOS oxide (tetra-ethylorthosilicate) layer 0.5 μm thick dissolved entirely, leading to corrosion of the underlying silicon device, following one year of
implantation.106 Evidently the formation of pinhole-free, lowdefect layers is necessary for oxide insulated devices intended

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for chronic implantation. Alternatively, glass encapsulation
may make use of considerably thicker layers; accelerated age
testing on Pyrex-sealed silicon devices by the Najafi group
extrapolated lifetimes of greater than 100 years.107,108
Though glass is compatible with many lithographic andsurface patterning techniques, bulk micromachining of glass
is considerably more arduous compared with silicon. Etching
techniques for glasses include wet etching, dry etching, andDRIE as well as physical ablation methods such as
sandblasting, however, in general, these approaches lack therequisite combination of scalability, high resolution, and/or
high etch rates for ease of fabrication. As such, there are limited though notable examples of glass μIDs. A cooperative
effort by Allen and CardioMEMS™ produced a passive,
implantable pressure sensor for chronic arterial monitoring,
fabricated by sealing an LC resonator within a glass etchedcavity.109 The approach was similar to silicon designs
described above; HF was used to isotopically etch a recessedcavity in glass. Then after electrodepositing a planar metal
inductor coil, the cavity was hermetically sealed by a glass
plate supporting a corresponding coil. Haque et al. departed
from this approach and developed a novel glass reflow process for fabricating an intraocular pressure sensor.80 The process involves creating a silicon mold using a DRIE Bosch process, anodically bonding the etched silicon to a Pyrex wafer,and then heating the Pyrex such that the glass reflows into
the etched features. The silicon is then etched away, leaving
behind a smooth, homogeneous glass structure with highresolution features. Using this method, Haque et al. fabricated a glass μID housing an antennae, battery, ASIC, andthe capacitive diaphragm pressure sensor for chronic monitoring of intraocular pressure.80
A characteristic high optical transparency helped establish
glass as the material of choice for commercial fiber optics
and subsequently the material of choice for optical μID
research. The use of fiber optics in endoscopic surgery has
increased familiarity with the technology among medical professionals, motivating the development of all-glass pressure
and temperature sensors fabricated on fiber tips. Suchdevices rely on micromachined Fabry–Perot cavities or Bragg
gratings to create sensitive transducers for in vivo conditions.
Increasingly, the field of optogenetics utilizes μIDs comprising glass fibers or waveguides for exciting light-sensitive proteins or detecting fluorescence in neural research. Following
initial efforts relying on unmodified commercial fibers,110–112
the development of micromachined glass optical probes
offered spatially focused stimulation easily paired with silicon electrode probes. Zhang et al. used HF to taper an opticalfiber to a sharp point, and gold deposition to create an optical aperture of approximately 1 μm.113 The ‘optrode’ (Fig. 3a)
was manually integrated with a commercial silicon neuralprobe array to create a combined optical and electrical probe
for simultaneous excitation and monitoring of the light sensitive channelrhodopsin protein.113 Subsequent efforts adapted
this approach for other optogenetic circuits and animalmodels, as well as devices with multiple optrodes.114–116

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superior electrical insulation. Devices are fabricated by lithographically patterning metal electrodes and then individuallycarving each probe using a diamond blade.130 Devices are
finally released individually by mechanically dicing the substrate, underscoring the lack of methods for etching throughthick ceramic substrates. This approach requires serial processing and may not be amenable to the production of many
types of geometries; as such, there are still limited applications for ceramic μIDs.
3.4 Metal
Fig. 3 Examples of microfabricated implantable devices machinedfrom glass, ceramic, and metal. (a) Glass fiber ‘optrode’ for opticalneural stimulation: reprinted from ref. 113 © IOP Publishing 2009. (b)Ceramic shank electrode for electrochemical sensing of L-glutamate:reprinted from ref. 126 © Elsevier 2002. (c) Arterial stent fabricatedfrom stainless steel with electro-discharge machining: reprinted fromref. 133 © Springer 2011.

Abaya et al. recently presented an alternative fabrication
method that used a combination of wet etching and mechanical dicing to create an optrode array from a single quartz
substrate.117 Other glass μIDs for optogenetic research use
lithographic patterning to create embedded microfabricatedwaveguides; Wu et al. presented a method for creating opticalwaveguides comprising an oxynitride core clad in SiO2.118
Such designs offer greater integration with lithographicallypatterned electrodes, but require more complex fabrication.
3.3 Ceramic
Owing to a lack of suitable etching and micro-patterning
methods, there are few examples of ceramic μIDs despite thesuccessful use of the material in medical implants such as
hip prostheses.119 Ceramic packaging of microfabricateddevices has been adopted for implants, as new techniques
allow for a high density of electrical feed-throughs while
maintaining impressively low permeability to water.120–122
Similarly, low-temperature co-fired ceramics (LTCC) havefound applications as the support substrate for multilayer
implantable ASICs and antennae, particularly for devices
requiring wireless telemetry.123–125 The high dielectric constant of ceramic, and the ease with which low-profile, multilayer devices can be fabricated, have helped transition
ceramic devices to in vivo applications.
Work from the Gerhardt group has demonstrated the suitability of ceramics as a μID substrate for more varied devices.
Examples include electrochemical sensors for glutamate
detection126–128 and single neuron recording.129 Typicaldevices consist of ceramic shanks, terminating in tips <5 μmwide, with micro-patterned platinum electrodes on one orboth128 sides (Fig. 3b). Advantages include high mechanicaldurability and biocompatibility, and a decrease in shunt
capacitance compared to silicon electrodes as a result of theThis journal is © The Royal Society of Chemistry 2015

Several metals have well established biocompatibility and a history of successful use in implanted medical devices, notably
titanium and platinum, but also iridium, nitinol, some stainless steel variants, and other alloys. These metals exhibit
excellent stability in vivo as well as excellent impermeability,and have long been the material of choice of hermetic seals
around ‘large’ implants such as pacemakers. However, theneed for wirelessly powered or wirelessly charged μIDs often
precludes metal packaging, as opaqueness to electromagnetic
radiation renders it unsuitable. Metals offer high mechanicalstrength and, unlike silicon, glass, or ceramics, low risk ofbrittle fracture, and therefore a lower risk of catastrophic
mechanical failure. These features make metals an enticing
choice of material for μIDs, though the characteristic highconductivity of most metals complicates the design of any
device requiring surface patterned electronics. As such, useof metals for μID fabrication is typically reserved for applications requiring simple devices and robust structures, such as
coronary stents. Laser milling of stainless steel is a common
approach, but requires serial processing and provides limited
resolution.28,131 Electrical discharge machining (EDM) is analternative method that offers improved resolution and thepossibility of batch processing.132 The technique involves
generating electric spark discharge from microelectrodes tomill away material from a conductive substrates.132 EDM has
been successfully used to fabricate several stainless steel μIDs
including novel variations on coronary stents82,132,133(Fig. 3c) and also blood vessel cuffs for in vivo flow measurements.134,135 Fofonoff et al. demonstrated the use of EDM toproduce a high density array of titanium microprobes forneural recording.136 EDM fabricated titanium columns were
etched with HCl to create a tapered tip only a few microns
across, and then electroplated with platinum to provide a
conductive surface. The probes were insulated with a conformal layer of Parylene and the insulation at the tips removed
by laser ablation to create the recording electrode site. Arrays
were mechanically robust with precise spacing, but the single
electrode-site design limited applications.
Subsequent work has led to multi-site metal neural probes
by utilizing fabrication methods compatible with lithographic
patterning. Motta et al. used electroplating to create nickel
shanks with platinum recording electrodes insulated by silicon nitride.137 However, devices required an additional gold
encapsulation layer to prevent potentially dangerous leaching

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of nickel ions into the brain. An alternative approach by
McCarthy et al. takes advantage of recent advances in reactive
ion etching of titanium138 to bulk micromachine probes fromlithographically patterned titanium substrates.139,140

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3.5 Challenges of hard materials
The selection of hard materials incurs certain challenges,
specific to μID applications, which must be taken into
account during the design phase. Several such challenges
relate to issues of biocompatibility. For example, the majorityof the devices surveyed in this section require the addition ofan encapsulating barrier material to provide electrical insulation, to shield the body from potentially harmful materials or
to reduce protein adhesion and biofouling. Considerable
work and literature has been dedicated to ameliorating thisrequirement, and numerous encapsulating methods andmaterials are available, including oxides, polyurethane,
PDMS, Parylene and polyimide. Still, it remains a necessary
additional step, and researchers must be cognizant of concerns including adhesion and conformality of the barrier
layer, unwanted changes to electrical or electromagnetic
properties, and additional processing steps needed touncover critical features. Sharp edges or corners and hardsurfaces intrinsic to hard materials are also of concern, as
they can tear tissue during surgical placement or while
implanted. Additionally, severe mismatch between themechanical properties of such materials and the tissue at thetargeted site is believed to induce trauma, resulting frommicro-motion of the hard implant damaging surrounding tissue, and exacerbating the immune system's foreign-body
response. Numerous reports, compiled in reviews by Polikov
and Navarro,10,141 describe efforts to reduce such damage
through changes in design or electrode coating, but frequently, the stiff material composing the device is identified
as the problem.24,25,142 The high rigidity and stiffness of hardmaterials can also impede fabrication of devices requiring
flexible or deformable components. For μIDs such as nerve‘cuffs’, which require membranes with high mechanical compliance, hard materials are ill-suited.

4. ‘Soft’ materials for microfabricatedimplantable devices
Here we define ‘soft’ materials as those with low Young's
moduli (<104 MPa) and low hardness (<102 kg mm−2). Compared to silicon and glass these materials have lower thermal
conductivity, lower density, and much greater mechanicalcompliance. Soft materials commonly used in biomedical
implants include many plastics and rubbers, however thissection will focus on those materials with a demonstrated
amenability to microfabrication techniques and history ofuse in μID research, namely Parylene C, PDMS, and polyimide. Additional characteristics of such soft materials
include constrained working temperatures (<350 °C), highchemical inertness, and susceptibility to gas and vapor

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intrusion. Notably, these materials have well-documented
biocompatibility (many polymers are used in FDA approved
medical implants), with negligible issues of cytotoxicity, corrosion or instability, and as such have been used for decades
in conventional medical implants. Synthetic polymers and silicone rubbers, such as PDMS in particular, have long been afamiliar choice for construction or coating of implantable
catheters, shunts, fluidic valves, bladders, and cosmetic
implants among others, due to ease of fabrication, low cost,and minimal foreign body response incurred by implantation.143,144 Parylene and polyimide found early applications
as insulating coatings for implantable electrodes,145–148 andbecame increasingly common choices as biomedical encapsulation following studies confirming chronic biocompatibility.149,150 μIDs utilizing soft materials became a topic ofincreasing focus in the 1990s, following the proliferation ofpolymer MEMS and spurred in part by the development ofreplica molding and soft lithographic techniques. “Free-film”
devices, comprising soft material substrates, offer a combination of high biocompatibility, high flexibility, optical transparency, and electrical insulation that motivated continued
research. Following improvements in deposition and patterning methods, numerous soft μIDs were developed, which todate include flexible neural implants, retinal prosthetics,
pressure sensors, and drug pumps. More recent efforts havefocused on use of soft biodegradable materials, such as silk
fibroin, and the first examples of fully biodegradable and bioresorbable μIDs have been demonstrated.151
4.1 Surface micromachining of polymers: Parylene,
polyimide & PDMS
Fabrication of soft μIDs may entail surface micromachiningof planar films or soft-lithographic molding to create threedimensional or raised structures. The former entails depositing planar layers of polymers onto hard substrates (typicallysilicon wafers) using spin coating or CVD, and then using
photolithography to pattern additional layers of metal, insulation and other materials. Several examples of devices fabricated in this manner are displayed in Fig. 4. This approachallows for production of very thin, flexible devices, frequently
housing planar electrodes for electrochemical sensing or
stimulation. The technique is well suited for the developmentof flexible neural interfaces; the combination of thin profile,robust electrical insulation, and greater mechanical compliance addresses several problems facing hard μID neuralinterfaces and conventional microwire probes. Numerous
reports have described a range of devices fabricated using
Parylene152–159 or polyimide69,160–172 films, and in someinstances PDMS173–177 and liquid crystal polymer.178 The
most common design archetype is a thin penetrating shankcomposed of a polymer-metal-polymer sandwich intended for
implantation in the brain.69,153–158,161–164,166–168,171,179 Typicalfabrication relies on photolithography to pattern metal (frequently platinum) for electrodes and connective traces, andO2 RIE for exposing the electrodes from insulation andThis journal is © The Royal Society of Chemistry 2015

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Fig. 4 Examples of flexible microfabricated implantable devices
surface micromachined from soft polymers. (a) Neurotrophic sheathneural probe microfabricated from platinum and Parylene: reprinted
from ref. 157 © RSC Publishing. (b) Curved retinal micro-electrodearray with arrow denoting retinal tack hole: reprinted from ref. 184 ©
Elsevier 2008. (c) PDMS multi-electrode array for stimulation of spinal
cord, shown here wrapped around 2 mm diameter wire: reprinted
from ref. 177 © Springer 2007.

defining the shape of the probes. Several reports describe
hybrid devices incorporating polymer substrates with sections of silicon69,156,179 or silicon nitride,171 or thin sheets ofmetal,163 to increase mechanical strength without sacrificing
flexibility.
Alternative geometries have been developed for interfacingwith peripheral nerves, and include sieve electrodes170 andcurved devices for use as nerve ‘cuffs’.159,172,174–177 Nerve
cuffs consist of a micromachined polymer film patternedwith electrodes, which wraps around a peripheral nerve, providing contact for electrical recording or stimulation. These
devices exploit the flexibility of polymer μIDs to create geometries not feasible with rigid substrates. Surface micromachined polyimide172,180 and Parylene159 nerve cuffs havebeen demonstrated for sciatic nerves, but PDMS174–177 is amore common choice. Fabrication of PDMS devices often
involves use of photo-patternable PDMS to expose theelectrode sites.174,175 The use of RIE to etch away an insulating PDMS has also been demonstrated, but requires prohibitively long etch times.177 The ability to form micro-patterned
curved surfaces is a notable advantage of soft μIDs; examplesof other implantable technologies fabricated in this mannerinclude cochlear electrodes,181,182 retinal prostheses183–185
and catheter compatible flow sensors.186
4.2 PDMS soft-lithography
For μIDs requiring non-planar structures such as cavities,
actuators or fluidics, poor availability of chemical etchantswith high etch rates and high selectivity for soft polymers
preclude the use of bulk micromachining. Physical machining methods such as laser ablation are available, but, in general, these methods offer poor resolution and control, andThis journal is © The Royal Society of Chemistry 2015

Critical review

require serial processing. Instead, a mold of the structure canbe produced from thick photoresist or etched silicon, and thegeometry transferred to a soft polymer using soft-lithography,
micro-molding, or embossing. Typically, the desired material
is cast over the hard mold, in the form of an uncured elastomer precursor or deposited by CVD, and then either theunderlying structure is dissolved chemically or the device ispeeled off mechanically. This technique is a low cost
approach for creating devices requiring high-aspect, threedimensional geometries from soft materials.
Soft lithography with PDMS was first developed to patternmicrofluidic channels. Typical fabrication is conducted byfirst patterning a design in thick photoresist (frequentlySU-8), then casting and curing PDMS elastomer over the
mold. The cured PDMS containing a negative imprint of thepattern is physically peeled off the mold releasing the device.
In work by Lin et al., a PDMS microfluidic manometer was
fabricated in this manner for powerless sensing of intraocular pressure.187 The device consists of a PDMS microfluidic
channel connected to a reservoir of green dye sealed under adeformable PDMS membrane, and operates by transducingchanges in pressure into changes in the length of dye-filled
portion of the channel. Other PDMS μID pressure sensors
rely on strain gauges188 or contain LC resonators in hermetically sealed cavities, reminiscent of devices fabricated in silicon.189 In these efforts, the use of soft lithographic molding
replaced chemical bulk micromachining, simplifying fabrication and yielding pliable PDMS μIDs that did not require
additional encapsulation.
Similarly, the soft lithography method has been adapted
to fabricating implantable drug delivery devices relying on
PDMS fluidic reservoirs. In one design variation, a hollow
PDMS reservoir is integrated with a fluidic cannula(Fig. 5a);190–192 the reservoir can be filled and refilled by wayof a standard syringe owing to the self-resealing ability of theelastic PDMS. Electrolysis, driven by integrated electrodes,
can be used to electrochemically control drug delivery by generating bubbles that expel drug out of the cannula.190,191
Huang et al. developed an electrolysis-driven reservoir array,fabricated with PDMS soft lithography and developed for single use in vivo drug delivery.193 PDMS reservoirs, containing
drug aliquots, were integrated with electrodes and IC control,and sealed with a membrane that burst upon application ofelectrolysis. Other designs rely on passive drug release; Chenet al. sealed drug-filled PDMS microreservoirs with a nanohydrogel membrane that became permeable when the surrounding environment reached a certain level of acidity.194
Here, PDMS was chosen both for its biocompatibility and its
ability to bond with the polyĲN-isopropylacrylamide-co-MAA)
membrane.
4.3 Parylene micro-molding
Raised and 3D-structures can be fabricated in Parylene using
an analogous micro-molding technique; typically a mold ofpatterned sacrificial polymer (e.g. photoresist, polyethylene

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Lab on a Chipelectrodes202 that incorporate raised structures to delivery
neurotrophic agents and improve recording impedances.
Wang et al. developed Parylene based electrode arrays for retinal prosthesis, using micro-molding onto etched silicon
structures to create impressively sharp but flexible probes.203,204
The technique has also been used to create ancillary structures for silicon neural probes, such as an IC ‘pocket’ integrated onto a silicon probe array.205
4.4 Biodegradable materials

Fig. 5 Examples of microfabricated implantable devices fabricatedwith micro-molded polymers. (a) Refillable drug delivery device forchronic implantation, fabricated through PDMS soft-lithography:
reprinted from ref. 192 © RSC Publishing 2008. (b) Microfabricated bellows cast from Parylene C onto a polyethylene glycol mold: reprinted
from ref. 201 © Springer 2012.

glycol) or etched silicon is prepared, and then coatedconformally with Parylene by way of CVD. Following dissolution of the mold or removal of the Parylene by peeling, thegeometry is retained in the Parylene substrate. This technique allows for the fabrication of Parylene cavities, channels, reservoirs as well as more complex shapes with highaspect ratios, without relying on deep chemical or plasma
etching. Examples of μIDs fabricated in this approach include
several variations of intraocular pressure sensors,195–197 including a micro-molded Bourbon tube design197 and Parylene variants on LC resonator cavities,195,196 and electrochemical pressure sensors designed for monitoring hydrocephalus shunt
performance.198,199 A Parylene actuator for an implantable,
micro-scale drug pump using a combination of soft lithography and Parylene micro-molding was developed.68,200,201
The fabrication approach for the bellows requires a PDMS
mold that is used to cast stacked disks of polyethylene glycol (PEG), which are subsequently coated in a thin film ofParylene (Fig. 5b). Dissolving the PEG in water leaves
microfabricated Parylene bellows, which are installed in a
fluidic reservoir on-top of electrolysis electrodes. Generation of bubbles actuates the bellows, which displaces theliquid in the reservoir, allowing for precise, targeted delivery in vivo.
Several instances of Parylene based implantable electrodes
combine micro-molding with surface micromachining, to create implantable electrochemical probes with non-planar features. Examples include flexible electrode probes with microfluidic channels,158 ‘sheath’ electrodes153 and hemispherical

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Biomedical implants composed in part or entirely of
biodegradable materials confer several advantages, including mitigation of the host's foreign-body response, reduction
in size of implant following implantation, dissolution of temporary structures needed only during insertion, and often
obviation of need for explant surgery. The success of biodegradable polymers in time-release drug delivery capsules anddissolvable sutures has motivated recent work exploring biodegradable materials for μID fabrication. Several reports on
μID neural probes have described the use of biodegradable
polymers such as PEG,157,206 chitosan,207 and silk208,209 as
temporary stiffening agents to assist in implantation, however, there have been only a few reports describing the use of
biodegradable materials as the primary substrate in microfabricated implants (Fig. 6). Notable exceptions include workby Grayson et al. on the development of a biodegradable
drug-delivery array;210 the device consists of micro-reservoirs
embedded in comparatively slow degrading poly(L-lactic acid)(PLLA), preloaded with different compounds and sealed withthe fast degrading poly(lactic-co-glycolic acid) (PLGA). Fabrication entailed compression molding of a PLLA preform witha machined aluminum die. Fig. 6a shows an example of a
PLLA preform after the hot embossing process, showing theregularity, but limited resolution, of such fabrication. A similar fabrication method has been used to produce implantable
RLC resonators from biodegradable polymers (Fig. 6b),
embedded with conductive polypyrrole nanoparticles, for producing powered, biodegradable μIDs.211,212 Recently, Petersen
et al. demonstrated a hot embossing method using an etchedsilicon stamp to produce PLLA drug release devices with significantly improved resolution over traditional compression
molding techniques.213 Similar examples of biodegradable
devices include bio-resorbable coronary stents, fabricatedfrom preforms of PLLA and other polymers using direct laser
milling (Fig. 6d).214,215 However, as with compression molding, this approach is compatible with a limited subset of μID
designs.
Biodegradable silk membranes offer an enticing alternative. Recent efforts have developed a host of tools for microfabrication on biodegradable silk including methods for contact printing, soft-lithographic patterning and metal
deposition.216 Several reports outline adapting these methods
to produce biodegradable silk MEMS as optical217 andelectronic218 implantable devices. Fully functional, wirelessly
powered and successfully implanted devices have beenThis journal is © The Royal Society of Chemistry 2015

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these materials is not well understood. Due to the limited
thermal budgets of polymeric materials, many are not compatible with high temperature processing steps, including
some bake steps, bonding methods, and CVD steps. Low
working temperatures may also prevent use of solder connections, requiring the use of zero-insertion force connections or conductive epoxy for electrical connection. Thosematerials focused on in this section all suffer from somedegree of gas and vapor permeability. This can create problems such as bubbles of intruded gas forming during fabrication, and can also promote delamination of device structural layers after implantation. The latter presents a criticalissue for soft μIDs intended for chronic use, as the inevitable intrusion of water vapor can lead to electrical shorts
and catastrophic failures.
Low Young's moduli, though often desirable, can create
problems during implantation. Soft material neural shanks,
for example, frequently require a stiffening agent or tool toassist in insertion,221,222 which can complicate implantation
surgery. For some applications, such as inertial sensing,
stiffer materials may be preferred or even required, and many
applications ultimately require batteries and ICs built fromhard materials, undercutting the advantage of a soft μID
approach.

Fig. 6 Examples of microfabricated implantable devices built frombiodegradable materials. (a) Microfabricated holes in a PLLA preform
embedded using a hot embossing technique: reprinted from ref. 213 ©
Elsevier 2015. (b) RLC resonators fabricated from polyĲε-caprolactone)polypyrrole using a hot embossing technique: reprinted from ref. 211 ©
Elsevier 2013. (c) Dissolvable electronics microfabricated from Mg/
MgO and silk fibroin: reprinted from ref. 218 © AAAS. (d) Laser
machined stent fabricated from biodegradable PLLA: reprinted fromref. 215 © Elsevier 2014.

developed from silk fibroin substrates patterned with conductive magnesium coils and resistive heaters (Fig. 6c).219,220
Though use of biodegradable materials necessarily places significant restrictions on fabrication efforts, these nascentefforts show that many devices may still be realizable using
resorbable materials.
4.5 Challenges of soft materials
Challenges confronting the development of soft μIDs
include obstacles to both fabrication and operation. Thereare extensive, though often surmountable, limitations in thetools and methods available for microfabrication of thematerials discussed here. Slow etch rates and poor selectivity to photoresist masks restrict the types of structures thatcan be etched in soft materials, and make release of thickfilm devices difficult. There are few options for anisotropic
etching or DRIE. Photo-patternable variants of polyimide
and PDMS are often introduced as an alternative method
for surface micromachining, but the biocompatibility ofThis journal is © The Royal Society of Chemistry 2015

5. Conclusion and outlook
Given the diverse array of μID devices and applications,material choice will always be dependent on the requirements of any specific problem. Chief consideration must bethat of biocompatibility, ensuring both that implantedmaterials do not harm the body and that material resilienceis such that in vivo conditions do not prevent proper function of the device, and such issues will always be dependent
on questions of device placement, operation, and desired
duration. For devices intended for mass commercialization,
cost of material and ease of batch scale fabrication may
determine material selection, while for small-batch devices
intended as research material choice may be determined by
specific device design, or the availability of suitable microfabrication tools. The trend towards μIDs comprising soft
materials is likely to continue as on-going research yields
more microfabrication techniques developed specifically forsoft polymers. However, given the ubiquity of silicon fabrication facilities, the need for semiconductor based ICs, andapplications requiring mechanically hard or stiff substrates,
hard devices, and silicon in particular, are likely to remain
the most common choice for the near future. A large number of materials, new and old, with potential uses in μIDs
were not discussed here, and continuing research is likely
to produce even more in following years. Given suchincreasing options researchers should remain cognizant of
the available materials and their capabilities to best inform
design of new and improved μIDs.

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Conflict of interest disclosure
E. Meng has a significant financial interest in Fluid Synchrony LLC.

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Acknowledgements
The authors would like to acknowledge the Viterbi School ofEngineering for support of Dr. Kee Scholten, and the assistance of Hsiu (Mike) Yang in preparing the manuscript. This
work was partially funded by the NSF under award number
EFRI-1332394.

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