IOP PUBLISHING

JOURNAL OF MICROMECHANICS AND MICROENGINEERING

doi:10.1088/0960-1317/20/9/095028

J. Micromech. Microeng. 20 (2010) 095028 (12pp)

Low-cost carbon thick-film strain sensors
for implantable applications
Christian A Gutierrez and Ellis Meng
Department of Biomedical Engineering, University of Southern California, 1042 Downey Way,
DRB-140, Los Angeles, CA 90089-1111, USA
E-mail: ellis.meng@usc.edu

Received 17 March 2010, in final form 23 July 2010
Published 27 August 2010
Online at stacks.iop.org/JMM/20/095028
Abstract
The suitability of low-cost carbon thick-film strain sensors embedded within a biomedical
R
MDX4-4210) for implantable applications is investigated.
grade silicone rubber (Silastic!
These sensors address the need for robust cost-effective implantable strain sensing technology
for the closed loop operation of function-restoring neural prosthetic systems. Design,
fabrication and characterization of the sensors are discussed in the context of the application tostrain/fullness measurements of the urinary bladder as part of the neuroprosthetic treatment oflower urinary tract dysfunction. The fabrication process, utilizing off-the-shelf screen-printing
materials, is convenient and cost effective while achieving resolutions down to 75 µm. Thismethod can also be extended to produce multilayer embedded devices by superposition ofdifferent screen-printable materials. Uniaxial loading performance, temperature dependence
and long-term soak testing are used to validate suitability for implantation while
proof-of-concept operation (up to 40% strain) is demonstrated on a bench-top latex balloon
bladder model.
(Some figures in this article are in colour only in the electronic version)

1. Introduction(figure 1). It is known that the internal bladder pressure
undergoes only small changes during filling (∼400 Pa) due
to creep and stress relaxation in the bladder wall; muchlarger changes are encountered during voiding (∼5000 Pa)[3]. Furthermore, pressure measurements can be confoundedby body-movement-induced artifacts [4]. Thus, interestingly,
pressure data alone do not provide an adequate or reliablemeasure of fullness. In fact, bladder wall strain correlates
more effectively with bladder volume and is a direct and easily
observable measure of bladder fullness [5].
Despite the biomechanical evidence, the majority ofefforts to determine bladder fullness use pressure as theprimary indicator [6–8].
This may be due, in part,to the ability to readily catheterize pressure sensors forinternal measurements, but these approaches require repeat
procedures, are generally uncomfortable and can be painfulfor patients. Clearly, there exists a need for the developmentof more aptly suited technologies to address the uniquechallenges presented by physiological mechanical systems. A
simple yet robust approach is adopted here to the measurementof strain for in vivo applications. The focus of this paper is thedesign, fabrication and application of carbon thick-film strain

In 2008, it was estimated that approximately 259 000
Americans suffer from chronic spinal cord injury (SCI),
and each year this number increases by about 12 000 [1].
SCI patients with lower urinary tract dysfunction face many
challenging medical problems that have both an adverse and a
profound impact on daily life [2]. In normal patients, stretchsensitive mechanoreceptors located in the bladder wall respondto filling and transmit the sense of fullness to the autonomic
spinal cord centers. In SCI patients, however, the sensory input
to drive micturition (bladder voiding) is lost due to damage tothe signal conducting spinal cord.
Restoration of bladder function through the use of
neuroprosthetic systems has been an active area of researchfor some time. However, a significant and unmet requirement
for these systems is the need for feedback information
regarding the fullness/state of the bladder. Closed-loop,
tether-free operation necessitates the integration of a sensor
system that provides this information to the function-restoring
neuroprosthetic system capable of triggering bladder voiding
in response to the detection of maximal bladder filling
0960-1317/10/095028+12$30.00

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J. Micromech. Microeng. 20 (2010) 095028

C A Gutierrez and E Meng

A commonly cited figure-of-merit for the strain sensitivityof strain sensors is known as the gauge factor (GF) and isdependent on the choice of piezoresistive material. It is defined
as
!R/R
!R/R
GF =
=
,
!L/L
ε
where R is the nominal undeformed strain gauge resistance,!R is the total change in resistance, L is the original unstrained
length of the strain gauge, !L is the total change in length
under strain and ε is the applied strain across the gaugearea. Thick-film materials have been shown to exhibit GFs
in the range of 2–10 [14–16], which is nominally better than
traditional thin-film metal strain gauges (GF ∼ 2). However,the benefits of high-strain operation (>0.1%), compatibility
with implantable substrates (such as silicone rubbers), andcost-effective fabrication and packaging techniques make thickfilms considerably more attractive for the application discussed
here.

Figure 1. Urinary bladder anatomy showing the muscles andefferents involved in micturition. A separate function-restoring
neuroprosthetic system uses existing efferent circuitry via spinal
cord stimulation to control bladder voiding. The system utilizes
strain sensor feedback for the purpose of detecting bladder fullness
via bladder wall strain measurements to activate bladder voiding as
required.

2.2. Fabrication of the screen-printing apparatus
sensors for implantable sensor systems in the context of theneuroprosthetic treatment of urinary tract dysfunction due to
SCI. Due to the ease of use and ability to create precise patterns
on a variety of substrates, thick-film-based technologies are
regarded as being compact, robust and relatively inexpensive[9]. These features make thick-film-based sensor technologies
an attractive option for implantable sensor systems.

Thick films can easily be patterned by screen printing. Thismethod of pattern transfer involves the use of a physical mask
acting as a stencil or screen through which the thick-film paste
is applied. The predefined open areas in the screen thus transfera thick-film image onto the underlying substrate. Despite
the numerous technical advances and high costs associatedwith production-grade screen-printing equipment, the ultimate
printing resolution of these systems still lies around 50 µm.
The approach adopted here utilizes low-cost, off-the-shelfscreen-printing materials employing a lithographically definedbrass screen and an acrylic frame, with reproducible printing
resolution down to ∼75 µm.
A brass shim stock sheet assortment was purchased from
Precision Brand (Downers Grove, IL) along with a metal
etching kit from MicroMark (Berkeley Heights, NJ) containing
all necessary lithographic and etching tools and materials.
Brass sheets of 25–75 µm thickness were used. Brass sheets
were cut to size (11.5 × 15.2 mm) and cleaned using a wet
abrasive pad to remove the native oxide layer. A negative dryfilm photoresist was then cut to size and applied to both sidesof the wet metal sheet. The photoresist-metal sandwich wasthen placed in between two protective laminator sheets andpassed through a standard business card laminator provided aspart of the kit. The resist was then patterned lithographicallyunder ultraviolet light (45 mJ cm−2 ), utilizing a standardtransparency mask (Mikacolor, Los Angeles, CA). Photoresist
development was then performed in a dilute sodium hydroxide
solution, and the brass sheet was placed in a ferric chloride
bath at ∼80 ◦ C under constant agitation to etch away areasof exposed metal. After etching, a metal polish was applied
to the clean metal surface to protect the metal screen fromcorrosion and oxidation. A laser-machined acrylic frame wasconstructed to hold and stretch the metal screen above thesubstrate (figure 2).
R,
A precision metal-blade handheld squeegee (Permalex!
Transition Automation, Tyngsboro, MA) was used to press
the carbon paste through the lithographically defined patterns

2. Materials, design and fabrication
2.1. Material selection
Class III medical devices designed to be implanted into thebody must satisfy strict FDA regulations for biocompatibility
and bio-inertness. Therefore, material selection and theuse of proven biomedical grade materials is an importantdesign consideration at the outset of any implantable device
design. For these reasons, biomedical grade silicone rubber
R, Dow Corning, Midland, MI) was(MDX4–4210, Silastic!
selected as the strain sensor substrate due to its proven in vivo
performance and compliance with United States Pharmacopeia(USP) class VI and ISO 10993-1 material standards. Complete
encapsulation of thick-film devices within this substrate
maximizes material safety and eliminates the potential for an
immune response to the embedded thick-film material.
For the application discussed here, high-resistance carbon
thick-film paste (7082, DuPont, Wilmington, DE) was selected
as a suitable thick-film material because of its high resistance,
suitability for screen-printing [10] and strain-responsive
characteristics. Carbon-based pastes were previously used
as force sensitive resistors (FSRs) in artificial hair cell devices[11] and in tactile sensors, and they are known to exhibit robust
piezoresistive properties [12, 13]. The high-resistance natureof carbon thick films readily enables resistance measurements
directly with a multimeter without the need for bridge or signal
amplification circuitry, thereby further simplifying the sensor
interface.
2

J. Micromech. Microeng. 20 (2010) 095028

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(a)

(b)

Figure 2. (a) Custom made screen printing apparatus with brass screen, acrylic frame and metal blade squeegee. (b) Screen-printed carbon
coupons on a silicon wafer.

in the metal screen onto the underlying substrate. The precured as-printed carbon film thickness was dictated by the
thickness of the screen used. The final post-curing carbon film
thickness was typically 50–60% of the pre-cured thickness
due to shrinkage and evaporation of solvents during the curing
process.
2.3. Alternative deposition and patterning methods
2.3.1. Spinning. Spin coating was investigated as analternative method to producing uniform thick films (figure 3).
This was attempted on both silicon and glass substrates
at various speeds. Due to the highly viscous nature of
the standard paste, a thinner (3610—Dipropylene glycol
monomethyl ether, TCI, Portland, OR) was used to lower
the paste viscosity. It was found that although films weresuccessfully spun onto the substrate, their uniformity waspoor when compared to uniformities achieved with standard
photoresists. Significant thickness variation and striation
were clearly visible, especially at speeds below 2 krpm,
and these were likely due to the lack of any leveling agentsor striation-reducing additives that are commonly employed
in standard photoresist formulations. Additionally, the grain
size of the carbon particles (graphite) was of the order of
1–2 µm, so any aggregation or clumping of these grains also
contributed to the observed striation and surface roughness.

Figure 3. Film thickness as a function of spin speed measured onstandard 75 × 50 mm laboratory glass slides. Carbon paste wasthinned with dipropylene glycol monomethyl ether (mean ± SD,
n = 4).

prepared for relative etch rate comparisons (table 2). A
physical contact mask was utilized to shield portions of the
coupon while exposing the remaining areas. Photoresist
and Parylene C films exhibited greater etch rates whencompared with the carbon film under nearly all conditions,with maximum thick-film etch rates observed at 200 W, 100
mtorr etching conditions (RIE 80, Oxford Plasma Technology,
Oxfordshire, UK).

2.3.2. Plasma etching. Because of the organic natureof the carbon paste, it was hypothesized that reactive ion
etching (RIE) in oxygen plasma could provide a means ofpattern transfer at resolutions below screen-printable limits.
Compatibility testing with standard MEMS processes andmaterials was conducted to assess the potential of using
standard physical masking techniques (table 1). To test
etch performance, carbon test coupons were fabricated, as
previously described. Additional test coupons with photoresist(AZ 4903, Mays Chemical, Indianapolis, IN) and Parylene
C (Specialty Coating Systems, Indianapolis, IN) were also

Due to the chemical incompatibility of the carbon thickfilms with the photoresist (table 1) and the relatively poor
selectivity (table 2), chromium was selected as the etch mask
material to micropattern an array of carbon microelectrodes
serving as test features. A 2000 Å layer of chromium
was deposited by e-beam evaporation (Temescal, Edwards,
Livermore, CA) and patterned in a chrome etchant (CR-7,
Cyantek, Fremont, CA). The carbon film was then etched inoxygen plasma (150 W, 100 mtorr) for approximately 15 min.
3

J. Micromech. Microeng. 20 (2010) 095028

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(a)

(b)

Figure 4. Alternate method of pattern transfer: RIE plasma etching with the chromium etch mask. (a) Optical micrograph of a patterned
chromium mask over carbon film (carbon film roughness ±1 µm, 1–2 µm granule size). A uniform residue remains after etching. (b) Afterremoval of the chromium mask, residue can clearly be seen in the optical micrograph. Non-volatile components of the carbon paste
composition remain on the substrate.
Table 1. Carbon paste (Dupont 7082) MEMS process compatibility.
Process

ChemicalResultCleaning/rinsing
Cleaning/rinsing
Photolithography
Photolithography
Photoresist stripping and lift-off
Metal etching

DI water
Isopropyl alcohol
Photoresist
Photoresist developer
Acetone
Chrome etchantCompatible—no dissolution
Compatible—no dissolution
Not compatible—paste solvents attack photoresist
Compatible—no dissolution
Not compatible—dissolves carbon film
Compatible—no dissolution

Table 3. Strain sensor design parameters and measured unstrained
resistances.

Table 2. RIE etch rate comparison (µm min−1 ).

Power (W)

Pressure(mtorr)

100
100
100
100
150
150
150
150
200
200
200
200

100
200
300
400
100
200
300
400
100
200
300
400

AZ 4903

Etch rate(µm min−1 )Parylene C

CarbonDesignLength
(mm)

Width
(µm)

Calculated
resistance (k#)

Measuredresistance (k#)

0.495
0.317
0.148
0.01
0.569
0.493
0.341
0.444
0.907
0.689
0.794
0.505

0.427
0.344
0.117
0.234
0.43
0.582
0.298
0.414
0.641
0.606
0.422
0.472

0.213
0.139
0.118
0.057
0.419
0.375
0.305
0.258
0.629
0.435
0.414
0.284

1
2
3
4
5

1
1.5
2.5
4.75
5.5

500
100
100
100
75

3
12
36
72
112

3–6
10–16
24–35
35–55
48–70

Mixer, Thinky Corp., Tokyo, Japan) with a 10:1 base-tocuring-agent ratio. The prepolymer mix was poured onto
a 75 × 50 mm standard laboratory glass slide (∼500 µm
thick), degassed in a vacuum and cured in a convection oven at∼80 ◦ C for 30 min. The thick-film strain sensors were thenscreen printed onto the silicone sheet. Thin flexible 34AWG
wire was then adhered to the carbon contact pads using a small
drop of carbon paste at the contact pads. The films were cured
in a convection oven at 120 ◦ C for 30 min. The resultant
thick films had a thickness of ∼15 µm. A second layer ofsilicone (∼500 µm) was then poured, degassed and cured tocomplete the encapsulation of the thick-film sensors (figure 5).
A range of device designs of varying lengths and widths
were fabricated, yielding unstrained resistances in the range of
3–70 k# (table 3).

The result after the removal of the chrome mask is a clearlypatterned film with residue evident in the regions exposedto the plasma (figure 4). The result indicates that although
the carbon-particle filler appears to have been successfully
etched away, there appears to be a non-organic or nonvolatile component to the paste composition which is noteasily removed in oxygen plasma. Due to incomplete material
removal by RIE, this method of pattern transfer was notselected as the method of choice, although further work isnecessary to fully characterize the exact nature of the postetch residue.

2.4.1. Multilayer thick-film sensor array fabrication. With
the addition of an extra screen-printing step, the singlelayer fabrication process can be extended to a multilayered
process enabling low-resistance routing connections (figure 6).
Due to the inherently high resistance of the carbon thick film

2.4. Thick-film strain sensor fabrication
R,
Biomedical grade silicone rubber (MDX4-4210, Silastic!
Dow Corning, Midland, MI) was prepared (AR-250 Hybrid

4

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(b)

(a)

(c)

Figure 5. (a) Fabrication process for embedded carbon thick-film strain sensors within biomedical grade silicone. (b) Illustration of asingle-layer carbon strain sensor embedded between PDMS layers. (c) Photograph of a carbon thick-film strain sensor with attached wires
on a glass microscope slide (slide: 25 × 75 mm). Figure 5(b) reprinted from [17] with permission of the Transducer Research Foundation.

(a)

(b)

Figure 6. (a) Fabrication process for a flexible multilayered thick-film device with high-resistance sensors and low-resistance connective
routing. (b) A fabricated 3 × 3 array of sensors in a floret configuration. Figure 6(b) reprinted from [17] with permission of the Transducer
Research Foundation.

connections within the same substrate enables additional
sensor configurations. A proof-of-concept 3 × 3 array of
2-axis strain sensors in a floret configuration was fabricatedusing this method. High-conductivity silver paste was utilized
as the routing material (Silver epoxy paste type 1, Transene
Inc., Danvers, MA, USA). It was patterned onto the silicone
substrate and then cured in an oven. The silver pattern was thenaligned manually under a microscope to the carbon screen prior
to printing of the carbon features. The carbon thick film wasutilized here, it would be impractical to utilize the samematerial for both the sensing element and the connective
wire routing within the substrate. This would yield anexcessively high resistance contribution from the routing
connection and obscure strain-responsive variations in theregion of interest. Furthermore, strain response artifacts
from the routing connections themselves could lead toconfounding measurements. Therefore, the ability to pattern
both high-resistance features with low-resistance routing5

J. Micromech. Microeng. 20 (2010) 095028

C A Gutierrez and E Mengdeposited directly on top of the cured silver pattern followed by
one more curing step to set the carbon. The silver was printed
first because of the higher curing temperature required, thereby
preventing excessive baking of the carbon material.

3. Experimental methods
3.1. Preparation of carbon films for sheet resistancemeasurements
Rectangular carbon films measuring 35 × 45 mm and withthicknesses ranging from 8.6 to 22.5 µm were screen printed
onto standard 75 × 50 mm laboratory glass slides for sheet
resistance measurements. The films were cured in a convection
oven at 120 ◦ C for 30 min. Sheet resistance is a useful measureof resistance for devices utilizing a single layer of resistive
material with uniform thickness. Sensor resistance can easily
be determined in the design stage if the sheet resistance is
known. Furthermore, thick-film sheet resistance, in particular,
is coupled to the GF, which is an important figure-of-merit forstrain sensing technology [14]. Units of sheet resistance areexpressed as #/! or ‘ohms-per-square’. The resistance scales
linearly with the number of squares (of any side length) thatcan be drawn in the footprint of the resistive film.
A four-point probe measurement technique was utilized
to assess the sheet resistance Rs given by
! "
V
π
,
(1)
Rs =
ln(2) I

Figure 7. Schematic of the uniaxial strain measurement setup.

device designs while high-strain loading tests the ultimate
yield strength and fracture limits of the thick-film material in
addition to assessing linearity and performance for high-strain
applications. Both of these loading conditions are importantin determining the most appropriate sensor design for bladderwall strain measurement applications.
3.3. Temperature dependence characterization and soak
testing
Temperature dependence characterization was conducted on
three devices which were placed in an environmental chamberoven and connected to a precision multimeter (Keithley
2400, Keithley, Cleveland, OH) for resistance measurements.
The chamber temperature was increased from ambient room
temperature (23 ◦ C) to a maximum of 80 ◦ C by incrementsof 10 ◦ C. The resistance of each device was recorded at eachtemperature point after allowing the system to reach thermal
equilibrium.
Long-term soak testing was also conducted on threedevices completely immersed in a 1× phosphate-bufferedsaline (PBS) solution. PBS provides an ionic environment
similar to that found in the body. The solution and devices
were maintained in a sealed container at core body temperature(37 ◦ C) for the duration of the experiment. Resistancemeasurements were taken initially once every day with amaximum interval of three days towards the end of the 50 daysof testing, at which point the experiment was concluded.

where π/ ln(2) = 4.53 for an infinite sheet. This expression is
only valid for equal probe spacing between all four probes.
For film thicknesses <40% of probe spacing, and if theedges of film are more than about four times the spacing
distance from the measurement point, no correction factor
is necessary [18–20]. Standard probe spacing is usually about
1 mm for most four-probe measurement systems, making it
relatively easy to fabricate large area films by screen printing
for accurate sheet resistance characterization. A four-pointprobe resistivity stand (302, Lucas Labs, Gilroy, CA) was used
to measure sheet resistance. A sourcemeter (Keithley 2400,
Keithley, Cleveland, OH) was used in 4-wire sense mode tosimultaneously supply current and measure voltage.
3.2. Uniaxial loading apparatus

A dynamic mechanical testing system (ELF3100, Bose Corp.,
Eden Prairie, MN) was used to apply controlled strain to
the microsensors while the resistance was simultaneously
monitored with a precision multimeter (Keithley 2400,
Keithley Instruments, Cleveland, OH).
The sensors were mounted to the test apparatus via
specimen clamps with a known separation distance (figure 7).
The clamp separation distance was computer controlled andwas increased by discrete steps (minimum size 0.4 µm)
thereby applying known levels of uniaxial strain. Sensor
performance in low (<0.1%) and high (0.1–2%) strain regimes
was evaluated. Various sensor designs were tested (table3) under these two loading conditions. Low-strain loading
allows for comparison of GF performance across several

3.4. Bladder model
To replicate strain sensor monitoring of bladder fullness, acustom-engineered bladder model consisting of a latex sheet(500 µm thick) mounted in a laser-machined fixture (10 cm
diameter opening) was constructed. Precise control of
inflation/deflation of the bladder, and thus the applied strain,
was achieved using a calibrated air pressure system. Sensors
were mounted to the latex surface with a cyanoacrylate-based
USP Class VI biomedical adhesive (MG 100, Adhesive System
Inc, Frankfort, IL) (figure 8).
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(a)

(b)

Figure 8. (a) Schematic of the bench-top bladder model setup. (b) Photograph of a custom-engineered latex bladder setup. Strain sensors
attached to the latex wall responded to applied strain arising from controlled inflation of the model bladder. Figure 8(b) reprinted from [17]with permission of the Transducer Research Foundation.

Sensor resistance was monitored while internal bladderpressure was varied and recorded.
3.5. Directional discrimination capability of multilayered
sensor arrays
The directional discrimination capability of dual-axis floretconfigured multilayered strain sensors was evaluated. Several
floret elements within the array were selected and monitored.
The real time resistance response of a single floret sensor
element was measured under single-axis loading during
application of parallel and orthogonal strain to test the onaxis/off-axis resistance response. Strain was applied manuallyin lateral directions with approximately equal magnitude.
Figure 9. Resistance–temperature dependence.

4. Results and discussion

using thinner films or by alternative paste recipes with lower
carbon content. The drawback of using a thinner film is thatthe film becomes more fragile and film fracture may occur at
lower strain levels. Additionally, due to the fairly large grain
size (1–2 µm) found in the paste composition, a homogenous
film requires thicknesses of ∼10 µm to avoid cracking during
the curing process. A modest increase in sheet resistancewould therefore be possible at the cost of film durability. The
measured sheet resistance is, however, still sufficient for theintended application. Characterization of the GF for the thickfilm ‘strip’ sensors fabricated here is discussed in later sections.

4.1. Sheet resistance measurements
Carbon films were fabricated for resistivity and sheet
resistance characterization, as described previously. These
measurements yielded an average resistivity of 1.50 ±0.045 # cm (mean ± SD, n = 4), yielding an average
sheet resistance of 1 k# !−1 for 15 µm thick films. When
compared to sheet resistance values typically attained formetals, the values achievable with the carbon thick film wereseveral orders of magnitude greater. The high resistance
simplifies the layout of carbon sensing elements, which
obviates the need for serpentine structures usually required bythin-film metal sensors to achieve appreciable resistances, thus
minimizing overall sensor footprint. Based on these results,
a straightforward ‘strip’ design was selected, providing direct
resistance measurement capability, small footprint and strain
measurement along one axis.
Generally, there exists a relationship between the sheet
resistance and the GF for thick films. For films with sheet
resistances less than about 1 k# !−1 , GFs of about 2 are
typically observed, while for films with sheet resistances
greater than about 100 k# !−1 , GFs up to 12 are typicallyobserved [14, 15, 21–24]. Thus increasing the sheet resistancegenerally leads to increased GF and therefore sensitivity.
Increasing sheet resistance could potentially be achieved by

4.2. Resistance–temperature dependence
The measured temperature dependence was parabolic acrossall devices measured (figure 9). This parabolic resistance–
temperature characteristic is a well-documented phenomenon
for thick-film materials. Several theories have been put
forward to explain this behavior, including electron percolation
theory [25] (commonly employed to explain the chargetransport mechanism in semiconductors [26, 27]).
However, the most widely accepted theory proposed bySion et al [28] argues that a combination of thermal coefficientof expansion (TCE) and thick-film thermal coefficient ofresistance (TCR) produces the parabolic dependence. It was
7

J. Micromech. Microeng. 20 (2010) 095028

C A Gutierrez and E Meng

(a)

(b)

Figure 10. Data from nine consecutive strain cycles were plotted for two devices: designs 1 and 5 (as per table 3). (a) Strain sensitivity for a
5 k# device. (b) Strain sensitivity for a 48 k# device. The maximum GF achieved was ∼2.5 with design 5.

shown conclusively that the combination of the exponential
temperature characteristics of resistivity for thick-film resistors(TFRs) and the effects of thermal strain arising from TCE
mismatch between the TFR and substrate produced theparabolic characteristic.
For in vivo applications, sensor performance over a wide
temperature range is not as important as understanding thesensitivity to temperature variations near body temperature.
The in vivo environment is considered to be highly temperaturestable, with thermal fluctuations not exceeding ±1 ◦ C fromthe average core body temperature (37 ◦ C). The result forthree carbon sensors tested indicates that the inflection pointof parabolic curve lies close to 37 ◦ C. This is beneficial
because it implies that small temperature variations aroundcore body temperature will have a minimal effect on resistancemeasurements. Maximum error introduced by temperaturefluctuations in an in vivo setting is estimated to be within
0.1% (figure 9). This makes it possible to attain reliablestrain measurement data without the need for additionaltemperature compensation circuitry, further simplifying sensor
implementation.

cycles per device. Generally, all sensor designs exhibiteda positive linear response with strain. Design 1 (adj.
R2 = 0.980) exhibited the lowest sensitivity (GF ∼0.85), while
design 5 (adj. R2 = 0.974) exhibited the highest sensitivity(GF ∼2.5). High-resistance devices (design 5) provided amore uniform and repeatable response that was, however,somewhat less linear over this loading range. Althoughgood linearity was observed with all designs, hysteresis
during loading/unloading cycles contributed to overall strain
measurement errors and limited measurement accuracy for
low-resistance designs (figure 10(a)). A strain measurement
accuracy of ±14.3% of full scale (0.07% strain) was achieved
with design 1 providing very limited performance for lowstrain applications. An absolute strain measurement resolutionof ±0.005% was achieved with high resistance sensors (design5) and this includes all measurement and hysteresis errors
over the nine loading cycles. Over the range of operation (0–0.07% strain), this yields a measurement accuracy of ±7.14%
of full-scale, suggesting that under low-strain conditions,high-resistance sensors provide an optimum combination ofsensitivity, accuracy and repeatability. Long-term loading andfatigue testing are still required to assess strain measurement
performance over prolonged periods of cyclic stress.
High-strain loading was conducted on five samples(designs 1–5). Each sensor was tested once until fracture
was observed or until approximately 2% applied strain wasachieved. For sensors in which fracture occurred, no further
data were collected. For sensors that did not fracture, a single
loading/unloading cycle was recorded (figure 11). Under
high-strain loading conditions, high zero-strain resistancedevices (>20 k#, designs 3–5) exhibited a nonlinear response
followed by fracture of the carbon pattern (indicated by anopen circuit reading). The points of fracture were clearlyvisible under an optical microscope, but the silicone rubber
substrate remained intact. Fracture generally occurred in
the range of 1–1.5% strain. As strain was reduced during
the unloading cycle, the separated carbon sensor segments
were brought back into contact as indicated by a finite
resistance reading. These results further suggest that highzero-strain resistance designs are most suitable for low-strain
applications.

4.3. Soak testing
To further verify the suitability for long-term implantation andperformance, three sensors were passively and continuously
soaked in PBS at 37 ◦ C to simulate in vivo conditions. No
degradation in sensor resistance was observed after 50 days ofpassive soaking, at which point the experiment was concluded.
Furthermore, there was no visible degradation of the silicone
substrate nor migration of any carbon materials through thesubstrate. The positive results validate the fabrication process
and verify the encapsulation capability of MDX-4 for thickfilm sensors in vivo.
4.4. Uniaxial strain measurement
Sensor performance in low (<0.1%) and high (0.1–2%) strain
regimes were evaluated. Low-strain testing was conductedon three representative sample devices: designs 1, 3 (datanot shown) and 5 as per table 3. The data in figure 10
were generated by nine consecutive loading and unloading8

J. Micromech. Microeng. 20 (2010) 095028

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Figure 12. Representative data demonstrating real-time pressure
tracking by a strain sensor (5 k#) attached to the latex bladder wall.
Reprinted from [17] with permission of the Transducer Research
Foundation.

Figure 11. High-strain loading performance for five different device
designs. Reprinted from [17] with permission of the Transducer
Research Foundation.

Low zero-strain resistance devices (<10 k#) exhibiteda very linear response up to 2% strain with no breakage
or cracking of the carbon film observed.
Over thisrange, no noticeable hysteresis was observed but multiple
loading/unloading cycles were not performed.
Themeasurement performance of low-resistance devices (designs
1 and 2) in this range of applied strain is significantly better
than under low-strain loading conditions. Measurement errors
over this range were estimated to be within ±0.75% of fullscale. In the context of bladder strain applications wherelarge strain variations are expected, the performance of lowresistance devices appears to be more suitable. We therefore
found low zero-strain resistance designs both more robust andlinear over a wider range of strain, and these would be preferred
for bladder strain measurement applications.

consistent measurements. A ‘break-in’ period was requiredfor both the latex sheet and the attached sensor to achieve
steady measurements. However, after repeated cycling, sheet
relaxation resulted in a decreased sensor response for the same
applied bladder pressure during later cycles. Steady-statestrain measurements were also conducted at intermediatepressures (0.4, 0.6, 0.8 psi) and confirmed that the sensors
were capable of tracking sustained levels of steady-statestrains over a period of 10 min.
4.6. Multilayer sensor array dual-axis capability
A clear difference in the response between the on-axis(parallel) and off-axis (orthogonal) strain was consistently
observed across several devices (figure 13). Elongation of
a sensor element along a particular axis may also provide
increased sensitivity along that axis while simultaneously
reducing directional cross-talk. The floret configuration isvery compact (1.5 mm2 ) and is well suited to applications
requiring an array of multi-axis sensors.
Directional
discrimination within a monolithic film enables potentiallyinteresting applications to complex systems with anisotropic
strain profiles. For example, the electrically evoked contractile
response of muscle tissues could be monitored by a multi-axis
strain sensor array affixed to the tissue surface. Multi-axis
strain information from such a system could provide feedback
regarding optimum electrical stimulation parameters, such as
timing, frequency and amplitude, and help in the design andplacement of stimulation electrodes for advanced prosthetic
systems.

4.5. Bladder performance
Pressures up to 1 psi (6.89 kPa) were applied corresponding tomaximum strains of 30–40% on the latex bladder surface.
The pressure and thus strain was successfully tracked bymonitoring the resistance of a sensor (5 k#—design 1)
mounted to the latex surface (figure 12). The bladderunderwent multiple inflation/deflation cycles over a period of
83 min under continuous monitoring. Although the measured
strains are greater than previously tested limits of the sensors,
there are at least two factors that contribute to the observed
behavior. The first is that the relatively stiff mounting adhesive
between the sensor and bladder dampens and limits the straintransmitted to the sensor. This highlights the need for an
appropriate fixation method. The second factor is that the
bladder model applies a combination of tangential and radialstrains rather than simple uniaxial strain. Due to sensor
placement near the center of the bladder (where tangential andradial strains are equal) and the presence of the adhesive, thetotal strain is distributed and dampened, allowing the sensors
to capture bladder inflation dynamics without fracture failure[29].
Relaxation in the latex sheet after repeatedinflation/deflation cycles was a significant factor in

4.7. Suitability for in vivo applications
Commercially available strain sensors are commonly
packaged in a flexible plastic substrate, such as polyimide,
while solid-state semiconductive sensors utilize silicon as
the substrate material. These sensors typically employ
piezoresistive materials, such as metals (i.e. gold, platinum
and other alloys) and semiconductors. Both have a long,9J. Micromech. Microeng. 20 (2010) 095028

C A Gutierrez and E Meng

Table 4. Young’s modulus of common strain sensor substratematerials and relative processing costs for small-medium production
volume. For comparison, the modulus of soft tissue is included.
Material

Young’s modulus (MPa)Processing cost

Soft tissue
Silicone rubber
Silicone
Polyimide
Parylene-C∼ 0.1 [33–37]
0.2–0.5 [38, 39]
130 000 [31, 40, 41]∼ 7500 [41]∼ 4000 [42, 43]

n/a
Low
High
Low-medium
Low-medium

performance of biomedical grade silicone rubber and therelatively flat temperature dependence (at physiologically
relevant conditions) of the thick films utilized here, makethis combination of materials well suited to implantableapplications.
The feasibility of measuring bladder strain with thick-film
sensors has been demonstrated here, but consideration of the
final attachment mechanism and location is still required forin vivo use. Various attachment parameters, including suturematerial and sensor location, were investigated previously ingoats [4]. It was found that devices sutured superficially on
the bladder wall (peritoneum) detached within 12 months.
Furthermore, devices sutured within the wall (between the
bladder muscle and internal bladder epithelium) failed due toerosion through the epithelium into the bladder lumen. The
most promising attachment location was found to be between
the bladder muscle and peritoneum with non-absorbablepolypropylene sutures. A simple modification to the current
sensor design can include reinforced suture holes throughwhich the device can be easily attached to this locationin vivo.
Our present study has focused on characterizing single‘strip’ devices of varying lengths and widths, 1–5.5 mm and
75–500 µm, respectively, with a total device thickness ofapproximately 1–2 mm. Further refinement of sensor design
would allow for more precise targeting of the physiologically
relevant strains encountered during bladder filling and voiding.
The bladder possesses a remarkable wall structure that canaccommodate both large distensions (up to 11 times its resting
volume [35]) during normal urine filling as well as rapid releaseback to a relaxed resting state. Bladder wall strain ratios up to
70% are possible [5]. Therefore, designing a sensor capable ofoperation within an acceptable range of strain, for example inachieving >80% voiding efficiency [44, 45], is required. By
controlling the device thickness, the strain response to appliedstress can be modified. Thicker devices will experience lessdeformation under the same loads when compared with thinner
devices. In this manner, an optimum combination of device
robustness and response can be achieved. Similarly, theelasticity of the silicone substrate material can be tailored byadjusting the prepolymer to curing agent ratio. Additional
sensor layouts can also be considered that extend the strain
measurement capability of this technology. For instance, across-hatched or interleaved sensor pattern can be created.
Additionally, the silicone substrate itself can also be patternedin a similar manner by cutting or coring out unused andunpatterned regions. The resulting strain-sensitive ‘mesh’ will

Figure 13. Resistance measurements of single floret sensor element.
Strain was applied manually in orthogonal lateral directions to test
the on-axis/off-axis resistance response. Some cross-talk was
observed, but a larger on-axis response is clearly obtained. The
figure is representative of a typically observed response. Reprinted
from [17] with permission of the Transducer Research Foundation.

extensive history in strain sensor design [30, 31]. However,these traditional approaches are most suitable for low-strain
conditions (<0.1%). Measurement of bone strain has been
demonstrated in vivo [32] with traditional thin-film sensors.
However, the low-strain, high-stiffness conditions specificto bone are not applicable to soft tissues found throughout
the body. Furthermore, fabrication of sensors with thesematerials generally requires clean room facilities to deposit andpattern thin films. Following fabrication, extensive packaging
and assembly steps are typically required for in vivo use,
adding further complexity and cost. Although commercial
sensors are relatively inexpensive (due to economies of scale)and simple to use, there exist several factors limiting theirapplication in vivo. A major drawback of solid-state siliconsensors is their inability to provide cost-effective solutions toniche problems with low to medium (1–10 thousand per year)
production volume requirements [9]. Additionally, there exists
a large mechanical mismatch between soft tissues and common
substrates, such as polyimide and silicon. Such substratematerials can be up to five orders of magnitude stiffer than soft
tissue (table 4). This mismatch makes it extremely difficult
to employ such sensors in vivo. Silicone rubbers possessmechanical properties that more closely match those of soft
tissue, like the bladder. Furthermore, the final mechanicalproperties of the rubber can easily be adjusted by changing theratio of elastomer to curing agent in the prepolymer mixture.
For these reasons, thick-film-based strain sensors offer
several advantages over traditional approaches for in vivo
applications. Thick films are straightforward to implement,
compatible with established microfabrication techniques andinexpensive to process when compared with traditional
metal thin-film strain sensors.
Furthermore, thick-film
patterns are readily microfabricated on flexible polymer
substrates that provide superior mechanical matching tosoft tissues over traditional rigid substrates [38, 39, 41].
These advantages, coupled with the superior encapsulation
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J. Micromech. Microeng. 20 (2010) 095028

C A Gutierrez and E Meng

be much more responsive to imposed stresses than a solid slabof the same footprint. This technique can improve sensor
robustness and strain measurement range when piezoresistive
materials of low yield strength, such as thin-film metals, areutilized [46]. The ability to easily modify sensor design
parameters allows the fabrication of a range of device andmaterial properties that can be engineered to closely match
those of the bladder and to operate within its physiologicalrange.

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5. Conclusion
Simple, robust, low-cost carbon thick-film-based sensors were
designed, fabricated, tested and shown to be effective inmeasuring bladder fullness and strain on a bench-top model. A
straightforward cost-effective fabrication process was utilized
without the need for large, expensive equipment or facilities.
The use of USP class VI and ISO 10993-1 approved medicalgrade encapsulation and adhesive materials enables the use of
these sensors for various implantable applications. The strain
sensitivity and performance of these sensors were investigatedusing uniaxial strain measurements and a bench-top bladdermodel, with positive results demonstrating the ability tomeasure large strains at physiologically relevant strain levels.
Temperature dependence and soak testing performance werecharacterized and validated for the in vivo environment. The
fabrication process can also be extended to include multilayer
devices for arrayed applications. Future work can include
wireless telemetry for external measurements as well as
integration with preclinical neuroprosthetic systems currently
under development in animal models. This work provides astarting point for the fabrication of various other thick-filmbased technologies intended for use with implantable medical
devices where closed-loop operation and sensory feedback
information are required.

Acknowledgments
This work was supported in part by the Engineering ResearchCenters Program of the NSF under Award Number EEC0310723, an NSF CAREER Award (ECS-0547544) and theBill and Melinda Gates Foundation (CG). We would like tothank Ms Alice Cho for assistance in performing uniaxial
strain studies and Mr Jason Geathers and Mr Lawrence Yu
for assistance in fabrication. We would also like to thankProfessor James Weiland for access to the Bose ELF3100
mechanical testing system.

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