MEMS

1. Biomaterials 23 (2002) 2737–2750
Evaluation of MEMS materials of construction for implantable
medical devices
Geoffrey Kotzara
, Mark Freasa,
*, Phillip Abelb
, Aaron Fleischmanc
, Shuvo Royc
,
Christian Zormand
, James M. Morane
, Jeff Melzakd
a
BIOMEC, Inc., 1771 East 30th Street, Cleveland, OH 44114, USA
b
NASA Glenn Research Center, 21000 Brookpark Road, Cleveland, OH 44135, USA
c
9500 Euclid Avenue, Lerner Research Institute, Cleveland Clinic Foundation, Cleveland, OH 44195, USA
d
Case Western Reserve University, 10900 Euclid Avenue, Cleveland, OH 44106, USA
e
8340 Hunting Dr., North Royalton, OH 44133, USA
Received 8 November 2001; accepted 14 December 2001
Abstract
Medical devices based on microelectro-mechanical systems (MEMS) platforms are currently being proposed for a wide variety of
implantable applications. However, biocompatibility data for typical MEMS materials of construction and processing, obtained
from standard tests currently recognized by regulatory agencies, has not been published. Likewise, the effects of common
sterilization techniques on MEMS material properties have not been reported. Medical device regulatory requirements dictate that
materials that are biocompatibility tested be processed and sterilized in a manner equivalent to the final production device. Material,
processing, and sterilization method can impact the final result.
Six candidate materials for implantable MEMS devices, and one encapsulating material, were fabricated using typical MEMS
processing techniques and sterilized. All seven materials were evaluated using a baseline battery of ISO 10993 physicochemical and
biocompatibility tests. In addition, samples of these materials were evaluated using a scanning electron microscope (SEM) pre- and
post-sterilization. While not addressing all facets of ISO 10993 testing, the biocompatibility and SEM data indicate few concerns
about use of these materials in implant applications. r 2002 Elsevier Science Ltd. All rights reserved.
Keywords: BioMEMS; Biocompatibility; Implant; Sterilization; ISO 10993
1. Introduction
To date, the majority of the development effort in the
microelectro-mechanical systems (MEMS) field has
focused on sophisticated devices to meet the require-
ments of industrial applications. However, MEMS
devices for medical applications (BioMEMS) represent
a potential multi-billion dollar market, primarily con-
sisting of microminiature devices with high functionality
that are suitable for implantation. These implanted
systems could revolutionize medical diagnostics and
treatment modalities. Implantable muscle microstimu-
lators for disabled individuals have already been
developed [1]. Precision sensors combined with inte-
grated processing and telemetry circuitry can remotely
monitor any number of physical or chemical parameters
within the human body and thereby allow evaluation of
an individual’s medical condition. Ultimately, we expect
that the same device will be able to administer a
therapeutic treatment as needed or as instructed
remotely [2]. At the other end of the spectrum, MEMS
processing technology is also being used to fabricate
micropatterned molds to process biocompatible poly-
mers for cell culturing [3] and to fashion functionally
simple passive microdevices like retinal implants [4,5],
neural electrodes [6–13], needles, microblades, and bio-
capsules [14,15].
Preliminary effort has focused on BioMEMS design
and fabrication with the intent of achieving certain
functionality. However, there are other hurdles that
must be overcome in order to realize the commercial
*Corresponding author.
E-mail address: mfreas@biomec.com (M. Freas).
0142-9612/02/$ - see front matter r 2002 Elsevier Science Ltd. All rights reserved.
PII: S 0 1 4 2 - 9 6 1 2 ( 0 2 ) 0 0 0 0 7 - 8

2. promise of these devices, and these hurdles are often
overlooked until much too late in the design process.
For example, the devices must be packaged and
sterilized, and the resulting device must be compatible
with the host into which it is implanted. A pioneer in
this field has stated that ‘‘Biocompatibility is the single
most complex issue facing in vivo sensor development
and it needs addressing up front in the sensor design’’
[16]. Data on the biocompatibility and sterilizability of
MEMS materials, however, are surprisingly limited and
data from currently accepted standard tests are almost
non-existent. This is in part due to the practice of
encapsulating these materials to isolate them from
biological fluids. For example, pacemakers typically
enclose all of the electronics with the exception of the
lead wires in an hermetic welded titanium canister,
blood pressure sensors use a gel to isolate the sensor
element from the fluid [17], and electronic transponders
[18] and implantable muscle microstimulators [19] are
glass encapsulated. In the future, in order to improve
functionality and reduce size, ever increasing numbers of
MEMS devices will have direct patient contact thus
requiring that biocompatibility testing be performed on
MEMS materials of construction.
Previous biocompatibility testing has been application
specific and thus limited in scope. Furthermore, the pre-
testing procedures employed in many of the studies did
not correspond to acceptable sterilization protocols for
a clinical device [20–22] or were not specified [1,17,
23–27,28]. While this is suitable for screening purposes
in a research context, it is not sufficient for devices
subject to regulatory scrutiny. In the medical arena, 20
issued, or soon-to-be issued, standards on pre-clinical
and clinical evaluation of medical devices exist. There is
agreement, however, that the ISO 10993 battery of tests
represent the minimum requirements that must be met
by all of the participating nations. The Food and Drug
Administration has also adopted the ISO 10993
standards per blue book memorandum in 1995: #G95-
1, entitled ‘‘Use of International Standard ISO 10993,
‘Biological Evaluation of Medical Devices’FPart 1:
Evaluation and Testing’’ [29].
The results of biocompatibility assessments must be
included in submissions to the FDA and other medical
regulatory bodies around the world before the devices
can be marketed. This paper is the first to report the
results of basic biocompatibility testing on MEMS
materials using an internationally recognized test matrix
(ISO 10993) and to evaluate the effects of sterilization
on material properties. The resulting data can be used
by BioMEMS developers to guide material selection and
choose sterilization methods compatible with materials
and device function. In a broader sense, it is hoped that
dissemination of such data will also improve acceptance
and understanding of BioMEMS in the scientific,
medical, and regulatory communities.
1.1. Previous work
For device commercialization, the published data on
the biocompatibility of MEMS materials are substan-
tially limited. Test protocols have varied from researcher
to researcher, and testing in accordance with ISO 10993,
‘‘Biological Evaluation of Medical Devices’’, is almost
non-existent. Some samples have been sterilized, but the
effects of the sterilization method have not been
explored. Control materials, both positive and negative,
have rarely been used making interpretation of the data
difficult. Additionally, many of the biocompatibility
tests that have been performed to date have looked at
the MEMS materials as adjuncts to other devices fur-
ther complicating the interpretation of the test results
[30–41]. An example is the use of silicon carbide as a
coating to improve the wear resistance of certain
orthopedic implants or as a coating on tantalum stents
to reduce thrombogenicity. In these and other studies,
the mechanical loading is highly non-representative of
the end use of MEMS devices [30–32]. The coating
materials are also very thin and can unbond or be
abraded from the substrates resulting in the creation of
multiple types of wear debris that can affect the
biocompatibility testing outcomes. At best, these types
of tests can serve as general indicators with respect to
MEMS material biocompatibility.
Silicon-based devices have been implanted in various
shapes and locations and examined for acceptable
biologic response. Silicon sieve electrodes have been
fabricated using boron etch step and silicon micro-
machining techniques. When implanted in rat peripheral
taste nerve fibers for 91–118 days, 21 of 28 implants
successfully demonstrated nerve regeneration [5]. Silicon
plates with a pre-determined pore distribution have been
implanted in the rat pancreas, liver, kidney, and spleen
for 7 months. They were cleaned ultrasonically and
soaked in hydrogen peroxide for 30 min prior to
implantation. Tissue reaction assessed using light and
scanning electron microscopy indicated minimal to
moderate response [15]. However, no effort was made
to determine if the response was due to the material or
the processing. Arrays of silicon microshafts have been
implanted in rabbit cortices for 6 months and neuron
density measured as a function of distance from the
microshafts. Material effects were found to be small
while the effects of geometry were more pronounced. It
was concluded that silicon shafts with ultrasharp, chisel
tips and smooth sides could be inserted with o10 mm kill
zones [11]. Phosphorous-doped monocrystalline silicon
10 Â 10 arrays of 80 mm diameter, 1.5 mm long electro-
des (Fig. 1) have been implanted in feline cortical tissue
[42]. Fifteen arrays were implanted for 24 h, and 12
additional arrays were implanted for 6 months. Leuko-
cytes were uncommon and macrophages were found
one-third of the time in the chronic implants. It was
G. Kotzar et al. / Biomaterials 23 (2002) 2737–27502738

3. noted that arrays insulated with polyimide had a greater
involvement of macrophages [42]. In subsequent in-
carnations, the design was revised to incorporate thin
silicon nitride films deposited by low-pressure chemical
vapor deposition (LPCVD) as the insulator. Chronic
implantations of these devices demonstrated a fibrotic
tissue response in the meningies in as many as 50% of
cases.
Recently, the biocompatibility of silicon has been
examined in a number of studies by Bayliss and
Buckberry [43,44], Bayliss et al. [44–47], Kubo et al.
[48], and Mayne et al. [49]. They have demonstrated that
nanocrystalline silicon does ‘‘not exhibit significant
cytotoxicity’’ [45]. However, one in vitro study that
examined a silicon-bearing bioglass demonstrated the
formation of nodules on periodontal ligament fibro-
blasts which has been attributed to silicon release from
the glass [48]. When nanostructured surfaces on silicon
were examined as substrates for neuron (B50) cells in
culture, these researchers found that, compared to
polished bulk silicon and plasma-enhanced chemical
vapor deposition (PECVD) polycrystalline silicon, the
most successful surface was the mesoporous silicon
(approximately 10 nm pore size) [15]. An earlier study by
these researchers demonstrated the dependence of the
results on the cell type. When culturing CHO cells on
nanostructured silicon, nanocrystalline PECVD silicon
performed substantially better than the other materials
[43]. As a part of their investigations, these same
researchers reported that only autoclaving was suitable
for ensuring a sterile environment for all substrate types
[46].
Several papers have discussed testing silicon carbide
(SiC) in vitro. In one study using macrophages,
fibroblasts, and osteoblast-like cells, the a and b forms
of SiC particles were dry heat sterilized at 1801C for 4 h.
The researchers reported that the cytoxicity results for
the two forms of SiC showed clear trends for all cell
lines. Both forms of SiC were highly toxic at concentra-
tions >0.1 mg/l, and the a form of the material was
more highly cytotoxic than the b form. However, the
researchers reported that this difference was not
statistically significant [50]. Another study tested SiC
deposited from radiofrequency (RF) sputtering using
alveolar bone osteoblasts and gingival fibroblasts for 27
days. The investigators reported that ‘‘Silicon carbide
looks cytocompatible both on basal and specific
cytocompatibility levels. However, fibroblast and osteo-
blast attachment is not highly satisfactory, and during
the second phase of osteoblast growth, osteoblast
proliferation is very significantly reduced by 30%’’
[26]. According to another paper, in a 48 h study using
human monocytes, SiC had a stimulatory effect
comparable to polymethacrylate [27]. Cytotoxicity and
mutagenicity has been performed on SiC-coated
tantalum stents. Amorphous SiC did not show any
cytotoxic reaction using mice fibroblasts L929 cell
cultures when incubated for 24 h or mutagenic potential
when investigated using Salmonella typhimurium mu-
tants TA98, TA100, TA1535, and TA1537 [30]. An
earlier study by the same authors of a SiC-coated
tantalum stent reported similar results [31]. This second
to last listed study was the only one of the above that
employed an ISO protocol, ISO 10993-5 for cytotoxi-
city. Not coincidentally, this study examined a device
intended for use as a clinical implant subject to FDA
scrutiny.
Silicon dioxide (SiO2) has been tested in vivo. A
peripheral nerve electrode with a grid of ten silicon bars
coated with SiO2 each 40 mm wide and 160 mm apart and
silastic cuff was transected onto a rabbit nerve. By 32
days post-operatively, the EMG of the affected muscles
had partially recovered. The EMG of the affected
muscles was indistinguishable from the contralateral
control muscles after 150 days. At 332 days, the
conduction properties of the implanted nerve confirmed
that the nerve was capable of conduction through the
silicon grid [33].
Research on the biocompatibility of silicon nitride
(Si3N4) has approached the issue from a number of
different directions. One study examined the interaction
between Si3N4 nanopowders with biochemical media
[51] while others have examined its potential as a
material for total joint replacements [22,52]. Early
studies examined its effects on rabbit stromal cell
proliferation with conflicting results. One of these
studies examined the difference between in vitro and in
vivo testing using rabbit skeletal cells and tissue. In vitro
testing showed that marrow stromal cells (MSC)
attached initially to the upper portions of porous
Si3N4 ceramic test disks. However, after 4 weeks the
cells were only attached to the disk edges. When fresh
marrow, or first passage MSC, had been inoculated into
diffusion chambers with and without Si3N4 and
implanted intraperitoneally for 5 weeks, they formed
cartilage, bone, and fibrous tissue. There was tissue
differentiation adjacent to Si3N4 but not within the
pores. In contrast, when the Si3N4 implants were
Fig. 1. Micromachined neural electrode.
G. Kotzar et al. / Biomaterials 23 (2002) 2737–2750 2739

4. inserted into femoral marrow cavities, they were
surrounded initially by woven bone. After 12 weeks of
implantation, mature bone permeated those implants
with pore sizes of 255764 mm [25]. This latter result was
supported by an earlier in vivo study using rabbits that
also demonstrated that MSCs will proliferate and
produce bone matrix, indicative of tissue ingrowth, in
porous Si3N4 intramedullary implants [22].
Several in vitro studies on Si3N4 using the human
osteosarcoma MG-63 cell line have been performed.
These studies examined the effect of the Si3N4 on the
production of IL-1b and TNF-a as indicators of
inflammatory responses. In one study, Si3N4 disks and
particulates were tested with the human osteoblast-like
MG-63 cell line in vitro for 48 h. Materials were steam
autoclaved at 2701C for 20 min. The researchers
reported that the incubation of MG-63 cells with 1,
10, or 100 mg/ml of Si3N4 particles did not decrease
DNA synthesis compared to the cells in the polystyrene
control media. Furthermore, cells grown on the surfaces
of reaction-bonded silicon nitride disks (RBSN) resulted
in an increased expression of cytokines IL-1b and
TNF-a compared to cells propagated on the control
surfaces. In contrast to those results, the expression of
IL-1b and TNF-a of cells propagated on the surfaces
of sintered-reaction-bonded Si3N4 disks (SRBSN) ap-
peared to be the same as that of control cells on the
polystyrene surfaces [20]. A second study produced
similar results. The researchers reported that cells
propagated on RBSN fared poorly compared to those
propagated on SRBSM which suggested that the
process of sintering in the manufacturing of Si3N4 was
critical in maintaining the proliferation as well as
promoting the metabolism of the MG-63 cell line [52].
A third study was performed using the same protocol as
the earlier studies: solid and particulate (1, 10, or 100 mg/
ml) Si3N4 and the MG-63 cell line. This latter study also
examined the effect of the fabrication process (RBSN vs.
SRBSN) on the generation of IL-1b and TNF-a. The
results of this study paralleled those of the previous
study with the exception of the response to the 100 mg/ml
concentration. For 100 mg/ml, the TNF-a expression
was greater than that for the controls [53]. These results
indicate that material processing may impact biocompat-
ibility results.
Researchers at the Ecole Polytechnique Federale de
Lausanne, reported using the photoresist epoxy SU-8,
commonly used in high aspect ratio MEMS fabrication
procedures, as an insulating agent on their neural
electrodes. They also have used SU-8 as a base for cell
cultures. These represent chronic applications since the
cultures persisted for more than 3 months. The
researchers concluded that SU-8 ‘‘should be’’ biocom-
patible, but performed no specific tests to determine this
beyond their own application. (Marc Heuschkel, private
communication).
2. Materials, processing and test methods
2.1. Materials
The following materials were selected for inclusion in
this study: (1) single crystal silicon (Si), (2) polycrystal-
line silicon (polysilicon), (3) silicon oxide (SiO2), (4)
silicon nitride (Si3N4), (5) single crystal cubic silicon
carbide (3C-SiC or b-SiC), (6) titanium (Ti), and (7) SU-
8 epoxy photoresist. Of these, polysilicon, Si3N4, and
3C-SiC were deposited by chemical vapor deposition
(CVD), SiO2 by thermal oxidation of Si, Ti by physical
vapor deposition (PVD), and SU-8 by spin coating.
These materials were chosen because they represent
several major categories of MEMS materials, ranging
from Si and Si-derivatives (SiO2 and Si3N4) commonly
used in conventional MEMS, to inert materials, such as
3C-SiC, for chemically and biologically harsh environ-
ments. Titanium was selected because it is a metal
currently used in many biomedical applications and SU-
8 was included to represent the class of polymer thin
films that show promise in microfabricated BioMEMS
devices. Typical MEMS processing methods were used
in all cases. These materials also represent the major
functional classes of materials found in MEMS devices.
Single crystal silicon, for example, is the most commonly
used substrate in bulk and surface micromachining.
Polysilicon is, without question, the most commonly
used structural material in surface micromachined
devices, and 3C-SiC is receiving attention both as a
structural layer material and as a protective coating
material for harsh chemical and high wear environ-
ments. Silicon dioxide is widely used as a sacrificial and
electrical isolation material, especially in polysilicon
surface micromachining. Silicon nitride is used in
conventional MEMS as an electrical insulator and as a
non-conducting structural material. Titanium can be
used for electrical contacts and SU-8, an EPON epoxy-
based resin photoresist, can be used as a polymeric
structural material, an optical waveguide, encapsula-
tion, or as an insulation layer.
2.2. Sample preparation
The thin film samples were deposited by various
standard methods on prime grade, 100 mm-diameter,
(1 0 0) silicon wafers that are commonly used as
substrates in MEMS devices. The wafers were polished
on both sides, and had a thickness ranging from 450 to
480 mm. The wafers were boron-doped (p-type conduc-
tivity) with a resistivity ranging from 1 to 50 O cm. The
wafers were acquired from the manufacturer in a
‘‘process ready’’ condition, and therefore were not
exposed to ambient laboratory conditions until the film
deposition processes were to be initiated. Although
many of the processing steps have batch capability and
G. Kotzar et al. / Biomaterials 23 (2002) 2737–27502740

5. thus could accommodate wafers from other processes,
the wafers for this study were intentionally segregated
and processed separately to eliminate the risk of cross-
contamination.
The film deposition processes could be classified into
three categories: (1) CVD, (2) PVD, and (3) spin
coating. Prior to the deposition of the CVD films
(polysilicon, Si3N4, and 3C-SiC) and the thermally
grown SiO2, the designated wafers were cleaned using
RCA Laboratories cleaning procedure, as per standard
laboratory protocols. The main purpose of the RCA
clean was to insure the purity of the CVD films by
removing contaminants in an ex situ manner. The RCA
process is divided into two major steps, commonly
known as SC-1 and SC-2. SC-1 was performed in a
heated (801C) mixture of H2O, NH4OH, and H2O2, and
was used to remove organic contaminants, while SC-2
was performed in a heated (801C) mixture of H2O, HCl,
and H2O2, and was sufficient to remove ionic con-
taminants. A short immersion in HF was performed
between the two steps to remove any SiO2 film that may
have formed on the surface of the silicon wafers during
the SC-1 step. Between each step, the wafers were rinsed
in deionized (DI) water, and after the SC-2 step, the
wafers were dried in an N2-fed spin rinse dryer. For the
polysilicon, Si3N4, and SiO2 depositions, the wafers were
immediately loaded into the appropriate furnace and the
CVD process was initiated.
For the wafers used as substrates for the 3C-SiC films,
a second in situ cleaning procedure was performed
inside the SiC reactor, detailed below. The wafers used
as substrates for the PVD and spin-coated thin film
samples were not RCA cleaned, since such processing is
not required and is not commonly performed as part of
normal practice. Instead, the surfaces of these wafers
were cleaned using an in situ RF-sputter cleaning
procedure also detailed below.
2.2.1. SiO2
Silicon dioxide films were grown on the aforemen-
tioned silicon wafers using a high-temperature, thermal
oxidation process. Immediately following the RCA
clean, the wafers were loaded into a batch process,
horizontally oriented, atmospheric pressure oxidation
furnace (MRL Industries), which idles at a temperature
of 8001C in a N2 ambient. The furnace is equipped with
automatic loading and unloading capabilities in order to
minimize thermal shock of the wafers. Once loaded, the
furnace temperature was ramped up to 10001C at a rate
of 51C/min in N2. Once stabilizing at 10001C, the N2
ambient is substituted with a mixture of O2 and H2,
initiating a process commonly known as wet oxidation.
The flow rates of O2 and H2 were 6 standard liters per
minute (slm) and 9 slm, respectively. Under these
conditions, an SiO2 film with a mean thickness of
4887 (A, as measured by a Rudolph EL ellipsometer and
verified by a Nanospec 4000 optical reflectometer, was
grown in approximately 88 min. As expected, the
thickness uniformity was extremely high, with a varia-
tion of only 0.8% across the 100 mm-diameter sub-
strates. Both sides of each wafer were coated in a single
oxidation run. The as-deposited films had a specular
appearance when observed optically.
2.2.2. Si3N4
Silicon nitride films were deposited on RCA-cleaned
silicon wafers by a LPCVD process in a horizontal,
autoloading tube furnace (MRL Industries) designed to
deposit films on both sides of each wafer. The process is
performed at a temperature of 8201C and a pressure of
280 mTorr. Dichlorosilane (Si2H2Cl2) and ammonia
(NH3) are used as silicon- and nitrogen-containing
precursor gases, respectively. Prior to film deposition,
the loaded reaction furnace was evacuated to a base
pressure of 20 mTorr. Using a Si2H2Cl2 flow rate of 18
standard cubic centimeters per minute (sccm) and a NH3
flow rate of 108 sccm, stoichiometric Si3N4 films were
deposited at a deposition rate of about 30 (A/min. The
mean film thickness was 2061 (A, with a variation of
2.25%. Like the SiO2 films, the Si3N4 films had a
specular optical appearance.
2.2.3. Polysilicon
Doped polysilicon films were deposited by LPCVD
using an in situ doping process in a batch furnace tube
similar in general construction to the unit used to
deposit the Si3N4 films (MRL Industries). The process
uses silane (SiH4) as the silicon source and phosphine
(PH3) as a phosphorus-doping source. The phosphine
concentration from the source bottle was 5% in H2. The
deposition temperature was 6151C and the deposition
pressure was 300 mTorr. The SiH4 flow rate was
105 sccm and the PH3 flow rate was 5 sccm, which
resulted in a 5000 (A-thick film being deposited in
185 min. The mean thickness variation was not recorded
for these samples. Both sides of each wafer were coated
during a single deposition run.
2.2.4. 3C-SiC
3C-SiC films were grown on (1 0 0) silicon wafers in a
custom-built, RF-induction heated, atmospheric reactor
detailed elsewhere [1]. The reactor uses a SiC-coated
graphite susceptor that can only hold two wafers per
run, therefore, a second cleaning procedure was
performed on each wafer to augment the RCA clean.
This procedure was performed at the onset of the
growth process and involved exposing heated wafers to
a H2 ambient at a flow rate of 25 slm. This in situ
cleaning step was performed at 10001C for 5 min, after
which the susceptor is cooled to below 5001C and
propane (C3H8) is added to the H2 flow at a rate of
84 sccm. The propane concentration at the source bottle
G. Kotzar et al. / Biomaterials 23 (2002) 2737–2750 2741

6. is 15% in H2. After injection of C3H8, the susceptor
temperature is increased to 12801C and held at this
temperature for 90 s. A thin epitaxial 3C-SiC layer
(o100 (A) is formed on the silicon surface during this
step. After 90 s, the C3H8 flow rate is dropped to 26 sccm
and SiH4 is added to the mixture at a flow rate of
102 sccm. The concentration of SiH4 in the source bottle
is 5% in H2. Under these conditions, single crystalline
3C-SiC films are grown at a rate of 167 (A/min. For this
study, 5000 (A-thick films, with a thickness variation of
about 30%, were grown in 30 min. Unlike the other
CVD processes, this process deposits films on a single
side of each wafer, requiring two growth runs to
completely coat each wafer with 3C-SiC.
2.2.5. Ti
Titanium thin films were deposited by PVD using a
Denton Vacuum, Inc. Discovery 24 high vacuum
magnitron sputtering system. The system is capable of
accommodating two, 100 mm-diameter substrates, but it
can only deposit films on a single side, requiring two
deposition runs for each wafer. After loading the
deposition chamber, the system was evacuated to a base
pressure of 7.6 Â 10À7
mTorr. Once the base pressure
was reached, the wafer surfaces were cleaned in situ by
argon ion bombardment. The RF power and chamber
pressure for this process were 125 W and 5 mTorr,
respectively. After the in situ cleaning step, the deposi-
tion chamber was evacuated to a pressure near
6.0 Â 10À7
Torr, after which the sputter deposition
process was initiated. Titanium films were deposited
onto the silicon wafers by bombarding an elemental
titanium target with energetic argon ions. The argon
flow rate was 10 sccm, resulting in a background
chamber pressure of 5 mTorr. The deposition time was
42 min, yielding Ti films with a mean thickness of
5040 (A, based on a deposition rate of 120 (A/min. The
thickness and the deposition rate were determined by
measuring the step height of a titanium film deposited
on a masked glass slide. A Dektak 3030 stylus
profilometer was used for the step height measurements.
Optical techniques could not be used to map the
thickness variation across each wafer, so data regarding
thickness uniformity were not available.
2.2.6. SU-8
The SU-8 25 was purchased from MicroChem
Corporation (Newton, MA). Prior to application of
the SU-8 25, the wafers were cleaned in a piranha
solution. In piranha cleaning, the wafers are immersed
in a bath of a 3:1 mixture of concentrated 95% H2SO4
and 30% H2O2 for 20 min, followed by a 5 min rinse in
DI H2O, and finally dried in nitrogen ambient. The SU-
8 25 was hand applied to the wafers, and the wafers were
then accelerated to 1315 rpm over 5 s and were spun at
that speed for an additional 13 s. This produced a
coating of SU-8 B55 m thick. After which, the wafers
were subjected to a two-stage soft bake on a program-
mable hot plate. The first stage was for 5 min at 651C
followed by a second stage of 931C for 15 min. Ramp
rates for both stages were 51C/min. The wafers were
then blanket exposed for 4 s at 50 mW/cm2
. The post-
exposure bake was performed in two stages consisting of
1 min at 501C and 5 min at 951C with ramp rates of 51C/
min. The second side was then coated in the same
manner.
2.3. Packaging and sterilization
The wafers were packaged in Case Western Reserve
University’s Microfabrication Laboratory (MFL) clean-
room. Individual wafers were placed in 5.5 in  8.5 in
pouches and sealed using a Packworld Medical H400/
AT Heatsealer. Pouches were purchased from Tolas
Health Care Packaging and were made from TPP-0060
and TPF-0563. These pouch materials are suitable for
both autoclaving and gamma sterilization. After the
wafers were sealed in pouches, half were sent to STERIS
Isomedix Services in Morton Grove, IL for gamma
irradiation that varied between 51.2 and 53.6 kGy. This
was intended to simulate a 2 Â gamma sterilization of
2.5 Mrad per exposure. In commercial sterilization
practices, it is not possible to achieve a uniform dose
of radiation. Some parts of the load may be exposed to
twice the required dose so that other areas receive the
minimum required for the desired sterility assurance
level. The remaining half of the samples were autoclaved
at the Cleveland Clinic Foundation using an AMSCO
Scientific SI-120 Eagle Century Series. (Serial # AMS-
CO 10L5FC) 2 cycles for 25 min/10 min dry at 1211C.
2.4. Test methods
This paper takes the first step in performing basic
biocompatibility tests on MEMS materials using an
internationally recognized test matrix (ISO 10993).
NAMSA (Northwood, OH) performed all biocompat-
ibility and material characterization testing. We have
chosen the following basic tests as a starting point:
2.4.1. Materials characterization tests
Subtle differences in processing and constituents can
change the outcome of biocompatibility tests. For this
reason, it is important to unambiguously characterize
the material tested so that others can duplicate favorable
outcomes. This characterization can be accomplished by
clearly describing the fabrication and handling of the
test materials and by characterization of the specimens
produced. The latter tests are part of the ISO 10993-14
protocol. The 10993-14 series of tests is composed of a
variety of techniques to provide a ‘‘fingerprint’’ of the
materials. These include infrared analysis, various
G. Kotzar et al. / Biomaterials 23 (2002) 2737–27502742

7. thermal analysesFthermal gravimetric analysis, differ-
ential thermal analysis, and differential scanning calor-
imetryFvarious mechanical tests, and potential
extractables. Infrared and thermal analyses are primar-
ily used with organic materials and as such were not
employed to characterize the MEMS materials reported
here.
The materials we fabricated were extracted under
various conditions. Materials are characterized accord-
ing to these protocols for two basic reasons. First, the
tests establish a baseline that allows the results of later
biologic testing to be linked to specific extractables, and,
second, they establish the presence and nature of any
extractables that have the potential to find their way into
the human body. From the perspective of biocompat-
ibility, the components that can be extracted are the
most significant. ‘‘The United States Pharmacopoeia
includes physicochemical tests based on water and
isopropanol extracts that are particularly useful in
defining materials as rich or poor in extractables and
in categorizing a specific material’s extractables in
general terms, such as non-volatile residue, residue on
ignition, buffering capacity, heavy-metals content,
ultraviolet absorption, and turbidity’’ [54]. In addition,
scanning electron microscopy was performed on the
samples before and after sterilization to detect any
obvious visual changes.
2.4.2. Aqueous physiochemical tests
For these tests, 19.3–23.7 g portion of the test material
was extracted at 701C for 24 h using a ratio of material
to USP purified water of about 1:5. Non-volatile
residue, residue on ignition, heavy metals and buffering
capacity were determined. For the non-aqueous physio-
chemical tests, a 16.3–20 g portion of test material was
extracted at 701C for 24 h using a ratio of material to
isopropyl alcohol of about 1:5. Non-volatile residue,
residue on ignition, turbidity, and UV absorption were
determined. Isopropyl alcohol is used to evaluate
lipophilic extractables.
2.4.3. Biocompatibility tests
As previously mentioned, biocompatibility tests are
specified by ISO 10993 based on mode of use, bodily
tissue or fluid contacted, and duration of contact. The
tests include cytotoxicity, sensitization, irritation, acute
systemic toxicity, subchronic toxicity, genotoxicity,
implantation, hemocompatibility, chronic toxicity, car-
cinogenicity, and reproductive. However, certain tests
are common to all implant applications. Cytotoxicity,
short-term muscle implant, and long-term muscle
implant tests are such test regimens and yield a good
understanding of acute and chronic response to
biomaterials.
2.4.4. Cytotoxicity
The cytotoxicity tests we performed followed the ISO
10993-5 standard: ‘‘Test for CytotoxicityFIn Vitro
Methods’’. A single extract of the test article was
prepared using single strength minimum essential
medium (1 Â MEM) supplemented with 5% serum
and 2% antibiotics. Extracts for six of the seven test
materials were based on the USP ratio of 4g:20 ml.
Portions of the test articles were extracted in the 1 Â
MEM at 371C for 24 h. For SiC, the extract was based
on the alternative USP ratio of 60 cm2
:20 ml and was
extracted under the same conditions. Each test extract
was then placed onto three separate confluent mono-
layers of L-929 mouse fibroblast cells which had been
propagated in 5% CO2. Two milliliters of the test
extract, the reagent control, the negative control (high-
density polyethylene), and the positive control (tin
stabilized polyvinylchloride) were placed into each of
the three 10 cm2
test wells and were incubated at 371C in
the presence of 5% CO2 for 48 h. The test well contents
were then examined microscopically (100 Â ) to deter-
mine any change in cell morphology and the percent
lysis. Test well contents were also examined for
confluency of the monolayer, and color as an indicator
of resulting pH.
2.4.5. Implantation
The ISO standard governing implant effects is
10993-6: ‘‘Tests for Local Effect after Implantation’’.
Sterile implant samples, approximately 1 mm  10 mm,
were prepared aseptically. USP negative controls (poly-
ethylene) of the same dimensions were sterilized by
steam. Rabbits were implanted with the eight samples,
four test and four negative, for periods of 1 and 12
weeks, and were euthanized prior to excising muscle
tissue and examining the implant site macroscopically.
A microscopic evaluation was conducted to further
define any tissue response. Implant sites were examined
microscopically for signs of inflammation such as the
presence of polymorphonuclear cells, lymphocytes,
plasma cells, macrophages, giant cells, and gross
necrosis. They were also examined for fibroplasia,
fibrosis, and fatty infiltrates.
2.4.6. Scanning electron microscopy
Scanning electron microscopy (SEM) was used to
inspect the surfaces of the all samples to look for
evidence of sterilization-induced damage that might
affect the biocompatibility of the materials. A Hitachi
S4500 scanning electron microscope was used for the
SEM analysis. Surfaces were examined both at low
magnification and high magnification to locate and
identify a wide range of treatment-induced surface
defects. For insulating films such as SU-8, a thin
(o100 A-thick) Pd film was deposited on the surface
G. Kotzar et al. / Biomaterials 23 (2002) 2737–2750 2743

8. to eliminate charging effects that might be misinter-
preted as damage.
3. Results
3.1. Aqueous physiochemical tests
Samples for this series of tests were extracted based
upon weight rather than USP specified surface area.
Area limits are therefore provided for comparison to
determine the significance of the reported test parameter
values. Based on the results summarized in Table 1, all
materials except the silicon nitride and the SU-8
photoresist fall below detectable levels for all four
categories for materials that can be extracted in water.
The silicon nitride samples showed only non-volatile
residues that were above detectable levels, while the SU-
8 photoresist produced non-volatile residues and buffer-
ing capacity that were above detectable limits.
3.2. Non-aqueous physiochemical tests
As shown in Table 2, all of the test materials except
for SU-8 produced residues that were at or below
detectable limits. The SU-8 samples produced non-
volatile extracts exceeding the detection limits and
ignition residues that were right at detection limits.
However, the non-volatile residues were less for the non-
aqueous extraction than they were for the aqueous
extraction test. While the limits for these tests have not
yet been established, samples that are low in isopropyl
alcohol extractables are preferred for use devices that
contact blood and other tissues.
3.3. Cytotoxicity
The 1 Â MEM test extract was less than a Grade 2,
mild reactivity, in all cases (Table 3). For all test
material materials, confluent monolayers were present in
all test wells and the pH was similar to the negative
control at 48 h.
3.4. Implantation with histopathology
All seven materials were classified as non-irritants.
The detailed results of the 1- and 12-week rabbit muscle
implantation tests are shown in Table 4.
3.5. Scanning electron microscopy
For all materials included in this study, SEM analysis
revealed no discernable damage associated with the
sterilization treatments. Representative SEM micro-
graphs are shown in Figs. 2–7 taken for the Ti and
SU-8 samples. Figs. 2–4 are SEM micrographs from the
Ti specimens. Figs. 5–7 are SEM micrographs from the
SU-8 specimens.
4. Discussion
The potentially multi-billion BioMEMS industry is in
its’ infancy. Based on the Nexus study published in 1998
[55], the next phase of MEMS commercialization is
expected to be Biomedical MEMS. Currently, MEMS-
based pressure sensors are small enough to fit through
1 mm catheters [56], but are priced to be disposable, and
pacemakers have incorporated accelerometers to pace
the heart in proportion to patient activity [17,57]. In the
future, it is expected that as miniaturization continues,
more hybrid devices will be created, and that new types
of sensors will be developed. Already, researchers are
developing a pressure sensor that is as small as a grain of
rice, requires no battery, external catheter leads or
antenna and can be read using an external RF source
from 6 in away [58]. Glucose sensors, implantable
microstimulators, and MEMS infusion pumps are also
being developed. The Defense Department’s Advanced
Research Projects Agency (DARPA) is developing new
MEMS sensors that can monitor all major body systems
such as temperature, oxygenation, and respiration [17].
The growth rate of emerging product types such as drug
delivery systems and the ‘‘lab on a chip’’ are forecast to
exceed 65%/year. The growth rate for existing medical
products of in vitro diagnostics, heart pacemakers, and
hearing-aid mechanisms is forecast to be 21%/year [55].
While these microminiature devices are technically
feasible and there is a substantial commercial impetus,
a critical issue that remains is the biocompatibility
testing of the basic MEMS materials and processing
according to international standards.
Few of the implant materials developed over the last
four decades for orthopedic and cardiac applications
and more recently for dental applications are suitable
Table 1
Results of aqueous extraction tests
MEMS material Test extract
Non-
volatile
residue
Residue
on
ignition
Heavy
metals
Buffering
capacity
Silicon o1 mg o1 mg p1ppm o1 ml
Thermal oxide o1 mg o1 mg p1 ppm o1 ml
N-doped poly o1 mg o1 mg p1 ppm o1 ml
Silicon nitride 8 mg o1 mg p1 ppm o1 ml
Titanium o1 mg o1 mg p1 ppm o1 ml
SU-8 9 mg 1 mg p1 ppm 4.2 ml
Silicon carbide o1 mg o1 mg p1 ppm o1 ml
Limits based on
area
15 mg 5 mg 1 ppm 10 ml
G. Kotzar et al. / Biomaterials 23 (2002) 2737–27502744

9. for use in MEMS devices. The most notable of those
from the orthopedics and cardiology field are Ti and
certain medical grades of epoxy, while the research into
dental materials has supplemented Ti with Si3N4 and
3C-SiC. Even in the case of ‘‘biocompatible’’ materials
such as titanium, the processing techniques for MEMS
are typically different than a traditional implant, thereby
adding a new variable in the biocompatibility equation.
In addition, it was questioned if MEMS materials would
tolerate steam or gamma sterilization.
We suspected that Ti and SU-8 would likely suffer
most from the sterilization treatments and elected to
present these material SEM micrographs here. However,
no steam or gamma sterilization-induced damage was
observed in either of these materials. Other than the
observably smaller grain size for the autoclaved sample
as compared with the as-deposited and gamma sterilized
Ti samples, no differences in the Ti samples could be
observed. We believe that the difference in grain size is
not attributable to the sterilization procedure, but rather
small variations in the Ti sputtering process, such as
temperature, deposition rate or ambient pressure, which
is to be expected for samples prepared sequentially
rather than in a batch process. Perhaps somewhat
surprisingly, no differences were observed in the SU-8
samples either. However, SEM analysis was rather
difficult on the SU-8 samples, since the featureless and
extremely smooth surfaces made it extremely difficult to
image the samples. In most cases, we were required to
locate a dust particle (Figs. 4 and 5) or position the
electron beam very near the edge of the sample (Fig. 6)
in order to acquire a proper image. The roughened
surface shown in Fig. 6 is not due to gamma irradiation,
but rather is mechanical damage to the SU-8 film caused
during SEM sample preparation.
Cytotoxicity test results routinely correlate well with
short-term implant studies. They are so useful for
screening materials that may be used in medical devices
that they are required for every type of medical device.
‘‘Testing for cytotoxicity is a good first step toward
ensuring the biocompatibility of a medical device. A
negative result indicates that a material is free of
harmful extractables or has an insufficient quantity of
them to cause acute effects under exaggerated conditions
with isolated cells. However, it is certainly not, on its
own merit, evidence that a material can be considered
biocompatibleFit is simply a first step. On the other
hand, a positive cytotoxicity test result can be taken as
an early warning sign that a material contains one or
more extractable substances that could be of clinical
importance. In such cases, further investigation is
required to determine the utility of the material’’ [59].
The requirements of an ISO-protocol cytotoxicity test
are rather strict. For a test to be valid, the reagent and
negative controls must demonstrate no reactivity (Grade
0) and the positive control must demonstrate a reactivity
of Grade 3 or 4. The significance of the cytotoxicity
protocol is that if an implanted material is shown to
have low volumes of soluble components it will have a
limited ability to affect a living organism unless the
extracts are highly toxic. The materials characterization
tests we preformed demonstrated that all of the
materials we examined generated very low volumes of
extractables, and the cytotoxic response showed that the
Table 2
Result of non-aqueous extraction tests
MEMS material Results
Non-volatile residue Residue on ignition Turbiditya
UV absorption
Optical density
Silicon 0.002 l max ¼ 345 nm
Thermal oxide o1 mg p1 mg 1.8 NTU 0.006 l max ¼ 315 nm
N-doped poly o1 mg p1 mg 1.1 NTU 0.030 l max ¼ 190:1 nm
Silicon nitride o1 mg p1 mg 1.1 NTU 0.003 l max ¼ 335 nm
Titanium o1 mg p1 mg 1.3 NTU 0.227 l max ¼ 190:1 nm
SU-8 4 mg 1 mg 0.2 NTU 3.298 l max ¼ 210 nm
Silicon carbide o1 mg p1 mg 0.5 NTU 0.01 l range=190–370 nm
a
NTU=Nephelometric turbidity units.
Table 3
Cytotoxicity test data
MEMS material Reactivity grade
Test
sample
Negative
control
Positive
control
Less than
Grade 2
(mild
reactivity)?
Silicon 0.0 0.0 4.0 Yes
Thermal oxide 0.0 0.0 4.0 Yes
N-doped poly 0.0 0.0 4.0 Yes
Silicon nitride 0.0 0.0 4.0 Yes
Titanium 0.0 0.0 4.0 Yes
SU-8 0.0 0.0 4.0 Yes
Silicon carbide 0.0 0.0 3.3 Yes
G. Kotzar et al. / Biomaterials 23 (2002) 2737–2750 2745

10. extractable components were of low toxicity. While
these tests provide a good qualitative correlation with
the results from studies performed in vivo, exceptions
can arise in cases where the tissue damage in vivo by the
extractables operates through more complex mechan-
isms.
Intramuscular implant tests can provide an indication
of the more complex mechanisms of cytotoxic response.
‘‘Implanting a test article inside the body of a laboratory
animal is the most direct means of evaluating a medical
device material’s potential effects on the surrounding
living tissue. Samples are cut to size, if necessary;
sterilized; and implanted aseptically. Then, after a
period of time ranging from weeks to months, the
implant sites are examined. Attention is focused entirely
on local effects that occur in response to the presence of
Table 4
One- and 12-week rabbit muscle implantation results
MEMS material Irritant ranking score D between test
sample and controla
Non-irritant (0–2.9)
Test sample Control Slight irritant (3.0–8.9)
Moderate irritant (9.0–15.0)
Severe irritant (>15.0)
1-week implantation
Silicon 6.0 6.7 0.0 Non-irritant
Thermal oxide 6.0 5.0 1.0 Non-irritant
N-doped poly 7.3 6.0 1.3 Non-irritant
Silicon nitride 6.3 6.7 0.0 Non-irritant
Titanium 7.3 8.0 0.0 Non-irritant
SU-8 8.3 7.3 1.0 Non-irritant
Silicon carbide 10.0 8.0 2.0 Non-irritant
12-week implantation
Silicon 2.0 4.7 0.0 Non-irritant
Thermal oxide 2.0 4.7 0.0 Non-irritant
N-doped poly 2.3 4.0 0.0 Non-irritant
Silicon nitride 3.0 2.7 0.3 Non-irritant
Titanium 2.0 3.3 0.0 Non-irritant
SU-8 4.3 2.0 2.3 Non-irritant
Silicon carbide 1.7 2.0 0.0 Non-irritant
a
Negative difference is recorded as zero.
Fig. 2. As-deposited surface finish of sputter deposition elemental
titanium. Fig. 3. Surface finish of autoclaved sputter deposition elemental
titanium.
G. Kotzar et al. / Biomaterials 23 (2002) 2737–27502746

11. the test material that has been in intimate contact with
living tissue’’ [60]. This protocol is useful for evaluating
the local tissue response in the neighborhood of the
implanted material; it is not intended to evaluate the
potential systemic effects of a medical implant, which
requires an entirely different protocol.
These are in vivo tests and thus can determine any
localized effects on the tissues by a device or material.
The primary evaluation method is the macroscopic
evaluation of the size of the capsule surrounding the
implant. If a material is highly reactive, capsules with a
thickness of 2–4 mm can be found, while the capsule
surrounding the negative control may be so small that it
is not visible. The tissue surrounding the implant is
examined histopathologically, and, at the microscopic
level, the nature and extent of the cellular reaction
is evaluated and scored. If the reaction is severe,
inflammatory cells are present in greater numbers and
dead muscles cells will be found surrounding the
implant. The results of the implantation tests for the
MEMS materials we examined demonstrated, with few
exceptions, performance indistinguishable from the
negative controls. The noted exceptions were N-doped
polycrystalline silicon, SU-8 and 3C-SiC at 1 week and
SU-8 and Si3N4 at 12 weeks. In all cases, however, the
differences were not significant.
The biocompatibility of Ti has been well established
for commercially pure grades as well as several alloys.
ASTM standards (F67, F136, F620, F1108, F1285,
F1472, F1580, F1713, and F1813) exist for cast,
wrought, and forged Ti and Ti-alloys that are intended
for implantation. According to ASTM F67, unalloyed
Ti may be furnished in the hot-rolled, cold-worked,
forged or annealed conditions with descaled, sand-
blasted or ground surface finishes permitted. However,
established standards do not exist for alternative
Fig. 6. Surface finish of autoclaved SU-8.
Fig. 7. Surface finish of gamma sterilized SU-8.
Fig. 4. Surface finish of gamma sterilized sputter deposition elemental
titanium.
Fig. 5. As-deposited SU-8.
G. Kotzar et al. / Biomaterials 23 (2002) 2737–2750 2747

12. fabrication forms of Ti, such as sputtering, for surgical
implant purposes. The surface finish for sputtered Ti is a
function of the sputtering process and therefore only
limited variations in the finish are possible without
additional post-application processing. The surface
finish of the samples we tested is shown in Figs. 2–4.
During our testing, there were no characteristics of the
biologic response that were attributable to the granular
surface structure of the sputtered Ti and the results of
the biocompatibility testing described here are consis-
tent with those reported elsewhere for sputtered Ti
[61,62].
Established standards for implantable grades of Si3N4
and SiC do not currently exist. The forms and
compositions of these materials are still evolving as
potential applications are developed. Many of the earlier
studies examining the biocompatibility of SiC and Si3N4
employed materials generated by fabrication methods
suited to other implantable applications, such as RF
sputtering for SiC and reaction-bonding and sintered
reaction-bonding for Si3N4. The results of the biocom-
patibility tests for these fabrication methods may not
apply to materials fabricated using MEMS technology.
Our results for Si3N4 and SiC show that when the
materials are generated using MEMS fabrication
techniques, they elicited no significant non-biocompa-
tible responses to the test battery employed in this series.
This closely correspond with the results reported by
Orth et al. [63] who compared the biocompatible
response of SiC and Si3N4 with that of alumina when
implanted in rat femora. Those researchers demon-
strated that while alumina was more bio-inert than
either of the two silicon-based ceramics, neither SiC nor
Si3N4 produced an unacceptable result.
The biocompatibility testing discussed in this paper
did not uncover any MEMS material that was not
biocompatible when subject to the processing, packa-
ging, and sterilization methods described herein. How-
ever, due to the complex nature of biocompatibility, one
cannot conclude that these same materials will pass the
full battery of ISO 10993 testing, nor if minor changes in
processing will affect the results. In summary, this
testing is but the first step in determining if Si,
polysilicon, SiO2, Si3N4, 3C-SiC, Ti, and SU-8 epoxy
photoresist and their respective MEMS processing
methods are suitable for implantable medical devices.
These seven materials were tested using a baseline
battery of ISO 10993 physicochemical and biocompat-
ibility tests required by the Food and Drug Adminis-
tration and other international regulatory agencies. The
results are consistent with previous studies where
samples were carefully prepared, cleaned, and sterilized
and did not uncover any MEMS material that was not
biocompatible. Although these data can serve as a
reference document for BioMEMS device designers, it
should be noted that additional biocompatibility testing
will be required to meet the full ISO 10993 requirements
for implants.
Acknowledgements
This work was supported by a grant from the
Glennan Microsystems Initiative.
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