APPLIED PHYSICS LETTERS 102, 183511 (2013)

High strain biocompatible polydimethylsiloxane-based conductive graphene
and multiwalled carbon nanotube nanocomposite strain sensors
Curtis Lee, Louis Jug, and Ellis Menga)
Department of Biomedical Engineering, University of Southern California, Los Angeles,
California 90089-1111, USA(Received 22 January 2013; accepted 24 April 2013; published online 9 May 2013)
High performance strain sensors were achieved featuring simple, low-cost construction involving
the screen printing of combinations of multi-walled carbon nanotube and graphene nano-platelet
nanocomposites on biocompatible and flexible polymer substrates. Conductivity and thermal
coefficients of resistance of different conductive nanocomposite sensor materials were measured.
The zero current resistance and gauge factor of printed sensors was characterized. The
combination of high strain operation (up to 40%), high gauge factor (GF > 100), and biocompatible
construction pave the way for applications such as minimally invasive in vivo strain measurements.
C 2013 AIP Publishing LLC. [http://dx.doi.org/10.1063/1.4804580]
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Advances in materials and processing methods, including
microcontact printing,1 oriented nano-wires,2 “wavy” silicon,3and large area reduced graphene oxide,4 have allowed electronic components (e.g., transistors) to be built on flexible
substrates such as polydimethylsiloxane (PDMS), which is abiocompatible, flexible, and transparent polymer that has beenshown to be useful for flexible electronics.5,6 Electronics embedded in or on soft PDMS substrates are now possible bymaking them thin enough to be flexible,7 using geometries
that allow bending,6 or incorporating conductive filler material into the PDMS itself.8
Carbon nanotubes (CNTs) possess favorable mechanical,
thermal, and electrical properties,9 including intrinsic piezoresistivity resulting in their popularity as a conductive filler material in PDMS-based electronics such as a transistors8 andstrain sensors.7,10–12 Vacuum filtration7 and chemical assisted
vapor deposition10 can create thin films of CNTs for transfer
onto PDMS. CNTs were also incorporated into PDMS using
sonication,13 shear mixing,14 magnetic stirring,15 or by hand16
to create a conductive PDMS (CPDMS) composite patternedusing molds,11,14 micromolds,16 or microcontact printing.12
CNTs are preferred over traditional conductive filler materials
such as carbon black and Ag powder because of their low percolation threshold in PDMS (0.2 wt. % to 5 wt. % for carbon
nanotubes8,11 versus 10 wt. % for carbon black and 83 wt. %for 1–2 lm Ag particles17).
Graphene is an alternate low cost18 filler material thathas been used with PDMS to form flexible electronics.
Graphene shares similar electronic properties as carbon
nanotubes19,20 but is more amenable to patterning and bulk
manufacturing through solution based exfoliation,21 epitaxial
growth on metals,22 or epitaxial decomposition of SiC.20
Graphene-based flexible strain sensors were realized using a
number of techniques to create thin piezoresistive layers.
One approach is to deposit graphene on thin layer23 or 3D
matrix24 of metal (either copper or nikel), encapsulate thelayer in PDMS, and then etch away the metal. A layer of graphene platelets was achieved by filtration followed bya)Electronic mail: ellis.meng@usc.edu

0003-6951/2013/102(18)/183511/5/$30.00

transfer of the resulting layer onto a polymer substrate.25
Similarly, a dispersion of graphene oxide platelets was combined with a polymer precursor and reduced in situ.26
Hybrid CNT þ graphene strain sensors were previously
reported using vacuum filtration of a dispersion of
CNT þ graphene to form a conductive layer that was subsequently transferred onto a polymethylmethacrylate (PMMA)
substrate.25 Here, we present hybrid strain sensors featuring a
piezoresistive composite made from multiwalled carbon
nanotubes (MWNT), graphene nanoplatelets (GNPs), or a
combination of the two in a PDMS prepolymer to form ascreen printable CPDMS. Piezoresistive CPDMS was sandwiched between PDMS layers to realize flexible strain sensors with high gauge factor and capable of high strain
operation. Although piezoresistive composites consisting of
CNTs and PDMS were previously reported,11,15 biologicalapplications were not considered and thus a stiffer PDMS
(Sylgard 184) was used. Here, medical grade PDMS (MDX4-4210, a US Pharmacoepia class VI material) was used toform a nanocomposite sensor material.
The piezoresistivity of CPDMS, useful for strain andother sensing applications, arises primarily from (1) straininduced break up of conductive networks, (2) the intrinsic
piezoresistivity of the filler, and (3) the change in resistancedue to electron tunneling, which changes dramatically as afunction of filler separation. Simulations indicate thatapplied strain does not sufficiently disrupt conductive nanonetworks to significantly alter bulk resistance of material,27suggesting that the other two mechanisms are responsible.
Piezoresistivity contributed by CNTs in conductive composites is dependent on the diameter of tube used. For thicker
MWNT tubes having a diameter "65 nm, there is poor loadtransfer, and piezoresistivity is dominated by inter tube distance changes during strain.28 However, for thinner tubes
with diameters around 10 nm, the intrinsic piezoresistivityof the tubes become significant.29 The nanotubes used in
the current study were 20–30 nm in diameter, and therefore
the intrinsic piezoresistivity of the MWNTs are expected tocontribute, but not dominate the overall piezoresistivity ofthe device. The contribution of electron tunneling is102, 183511-1

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Lee, Jug, and Meng

dependent on the distribution and geometry of the conductive networks formed by the filler material, which is in largepart determined by the matrix material and processing parameters.30 The electron tunneling effects are expected todominate at low filler concentrations and under high strain
when low density of conductive networks is present.
CPDMS in the present work was created by measuringwith a precision scale a predetermined amount of carbon filler material, either GNP and/or MWNT (cheaptubes.com,
Brattleboro, VT), and PDMS (MDX4-4210, Factor II, Inc.,
Lakeside, AZ) base into a glass beaker to create a CPDMS
prepolymer. To this was added 10–15 mL of a solvent mixture consisting of a 30:70 ratio of IPA and Stoddard solvent(by volume). The mixture was then placed in an ultrasonicbath (Bransonic 3510, Branson Ultrasonic Corp., Danbury,CT) for 15–18 h to achieve thorough mixing of the filler material with the PDMS base and evaporate the solvent. Using
this method, CPDMS base with a range of concentrations of
MWNT and GNP were prepared (Fig. 1). Crosslinker (in aratio of 1:10 to the base) was added to the nanocomposite
conductive filler-PDMS base mixture using a planetary
mixer (Thinky Corp., Laguna Hills, CA), and the obtained
conductive prepolymer was screen printed onto a 500 lm
thick layer of non-conductive medical grade PDMS using alithographically defined brass screen prepared using a previously developed method.31 Briefly, CPDMS prepolymer was
spread across a custom stencil using a plastic squeegee. The
stencil was made by etching away a lithographically definedpattern in a 76 lm thick brass shim (4.5 # 6 in.) sandwiched
between two sheets of a negative, dry-film photoresist. The
photoresist was exposed using a high resolution transparency
mask (Mikacolor, Los Angeles, CA) and UV light source(45 mJ/cm2) and developed in a dilute sodium hydroxidebath. The exposed brass regions were etched away using ferric chloride. The parts for this etching kit were purchasedfrom MicroMark (Berkeley Heights, NJ).

Appl. Phys. Lett. 102, 183511 (2013)

The screen-printed CPDMS was then placed under avacuum for 3–4 h to remove any remaining solvent and thencured at 80 $ C. A second layer of non-conductive medical
grade PDMS was added to fully encapsulate the CPDMS.
Robust electrical connections were made by threading finewires through the contact pads several times or making aconnection using conductive epoxy (Microcircuit Silver
Type O, Transene Company, Inc., Danvers, MA).
Conductivity is achieved when the concentration of filler
is sufficient to form contiguous networks within the nonconductive matrix either by direct contact of filler particles(network formation) and/or by electron tunneling throughthe polymer between closely spaced particles. This percolation threshold concentration is a function of the uniformityof distribution (which is in turn related to the viscosity of thepolymer and the preponderance of aggregation of the filler
material),32 aspect ratio, and degree of curvature of the filler
material. Single-walled carbon nanotubes (SWNT) arecurved,28 however, MWNT are relatively straight, and therefore the degree of curvature can be ignored.
Conductivity was calculated from sheet resistance measurements using a four point probe (Rs ¼ 4.53(V/I) whenprobes have equal spacing, films are <40% of probe spacing,and if the edges of the film are more than four times the spacing of the probes8,16). Initially, 35 # 45 mm rectangular patterns of CPDMS were printed onto a non-conductive PDMS
substrate to ensure that the edges effects were minimized.
However, subsequent four point probe measurements performed directly on un-encapsulated sensors (1.8 # 30 mm)
yielded identical resistance values; therefore, four point
measurements were later made directly on exposed, unencapsulated sensors.
The percolation threshold was 5 wt. % for MWNT-only
and MWNT-GNP composites. The percolation threshold for
GNP composites was 12.5 wt. % which can be explained by
the smaller aspect ratio of the platelets compared to tubes
and is similar to values reported by other groups working
with graphene polymer composites.18 Once the percolationthreshold is crossed, traditional percolation theory33 statesthat conductivity is related to the concentration of the filler
material by the following equation:rcom ¼ r0 ð/ ' /c Þt

FIG. 1. (a) Photograph of screen-printed MWNT/GNP strain sensor (scalebar is 5 mm). (b) Cross sectional SEM image of strain sensor thick film prepared by freeze fracture (scale bar is 1 lm). GNPs appear as flakes (denotedby orange, thick arrows). MWNTs appear as solid white dots or filaments(denoted by thin blue, arrows). For comparison, (c) GNPs (scale bar is 1 lm)
and (d) MWNTs (scale bar is 500 nm) are shown in their pre-processedstate.

for/ > /c ;(1)

where rcom is the conductivity of the composite, r0 is the
conductivity of the filler material, / is the concentration of
filler material, /c is the concentration of the filler material at
the percolation threshold, and t is an empirically determined
exponent for each system that is related to the aspect ratio,
curvature, and degree of aggregation of the filler material.
This exponential trend has been confirmed in literature andagrees with experimental results (Fig. 2). Above concentrations of 13% filler material, the MWNT composites were too
viscous to screen print and the GNP composites did not cure.
To evaluate the utility of these nanocomposites forstrain sensing applications, the zero current resistance(ZCR), thermal coefficient of resistance, soak test performance, and gauge factor were characterized. Zero current resistance was calculated by recording the resistances of thedevices from 1–10 lA using a precision multimeter

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Appl. Phys. Lett. 102, 183511 (2013)

FIG. 2. Conductivity data comparing experimental results against those forsimilar composites found in literature. Lu et al.11 used 40–60 nm diameter
MWNTs and Liu and Choi12 used 20–40 nm diameter MWNTs.

(Keithley 2400; Keithley Instruments, Cleveland, OH) andextrapolating the resistances back to zero current (Fig. 3(a)).
Measuring ZCR in addition to conductivity revealed the current dependency of conductivity which was higher at lower
conductivities. The dependency of resistance with current
has been suggested in literature to be associated with theincreased number of shells contributing to conductance at
higher currents in MWNTs34 and the linear increase in prevalence of electron tunneling with applied voltage.28,35 The
screen printing process allowed three devices to printed at
one time and among each batch, the standard error of the
ZCR was as low as 3% (data not shown). However, between
batches and across different filler concentrations, the standard error of the ZCR was "20%, with the exception of 69%
standard error for the 5% MWNT devices. These data suggest improvements in controlling the thickness of screen
printed conductive layers (e.g., automating the process)could reduce the standard error.
The temperature sensitivity of the composite, or temperature coefficient of resistance (TCR), was determined bymeasuring the device resistance under a constant current bias
(10 lA) at different temperatures from 30 to 80 $ C (10 $ C
increments) and taking the slope of the resulting curve (Fig.
3(b)). The TCR for composites made from MWNTs wereshown to have a weak dependence on composition and be
negative, whereas the TCR for GNP devices were positive.
Both of these results are consistent with literature,36,37 where
it is suggested that filler materials with lower aspect ratios(GNPs) will have a positive TCR due to polymeric expansion with temperature that causes increased tunneling resistance which do not affect fillers with high aspect ratios wherecontact resistance dominates.37
Soak testing of strain sensors was performed in 1X PBS(37 $ C for 21 days) to simulate in vivo conditions and isongoing (data not shown). Although a decrease in resistanceof was observed, likely attributed to penetration of conductive ions into the PDMS, sensors were still operational. It is
expected that a thin Parylene coating (few lm) will prevent
migration without significantly increasing the stiffness of thesensors.

FIG. 3. (a) Zero current resistance (mean 6 SE, n ¼ 6–12) and (b) TCR(mean 6 SE, n ¼ 3–5).

Gauge factor (GF) is a useful measure of merit for piezoresistive strain sensors and is equal to the normalizedchange in resistance divided by the strain
DR DR
GF ¼ R ¼ R ;
DL
e
L

(2)

where R is the nominal as-fabricated, undeformed strain
gauge resistance, DR is the total change in resistance, L isthe original unstrained strain gauge length, DL is the change
in length following application of strain, and e is the applied
strain. Metal and semiconductors are typical materials used
for strain gauges. Metal strain gauges typically have relatively low gauge factors (GF " 2). Although semiconductors
have improved gauge factors (GF " 100), mechanical properties limit their use to low strain (<0.1%) applications.
Thick film materials for strain sensing have reported GFs in
the range of 2–12.11,12,38
Gauge factors of the strain devices were obtained using a
motorized stage (Z812; Thorlabs, Newton, NJ) that allowed
for user defined strain patterns and a precision multimeter(Keithley 2400; Keithley Instruments, Cleveland, OH). The
sensors were mounted to the motorized stage using custom
acrylic clamps with a known separation distance and subjected to uniaxial strain using LabVIEW (National Instruments,

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Lee, Jug, and Meng

Appl. Phys. Lett. 102, 183511 (2013)

FIG. 4. (a) Strain pattern applied to device. (b) Raw
data of resistance versus time. (c) Resistance versus
strain (with first 2 cycles removed). Gauge factor
obtained by averaging slope of linear region.
(d) Gauge factors for various compositions including
MWNT and/or GNP fillers (mean 6 SE, n ¼ 4–6).

Austin, TX) to control the motor (Figs. 4(a) and 4(b)). Gauge
factor was calculated by averaging the linear fit to relative resistance versus uniaxial strain data during the stretchingphase of 3–5 strain cycles (Fig. 4(c)) (the R2 value of the linear regressions ranged from 0.93 to 0.99). Sensors exhibited a“breaking in” cycle, in which the first 1 or 2 cycles had either
a higher or lower gauge factor compared to subsequent cycles(Fig. 4(b)), and these cycles were not used when calculating
gauge factor. This phenomena is likely due to the differing
stiffness between the filler materials and PDMS; the initialstrain cycles likely displace some filler relative to the PDMS
Two operation regimes were observed: a smaller GF at low
strain and a larger value at higher strains (Fig. 4(c)). Thisphenomenon was previously observed in other CNT-polymer
composites in which it was hypothesized that at lower strains
the change in resistance is dominated by breaking electricalcontacts between overlapping CNTs, but at high strain thechange in resistance is dominated by increasing the tunneling
distance between tubes that are already separated.39 The present study focused on the device’s maximum GF in preparation for high strain biomedical applications, and therefore
only the GF at higher strain was calculated.
Gauge factor for our composites ranged from 2 to 22 forhigh concentrations of filler material (10 wt. % MWNT and
12.5 wt. % GNP) and up to 100 for lower filler concentrations
closer to the percolation threshold (5 wt. % MWNT and
3.5 wt. % MWNT þ 1.5 wt. % GNP mixture) (Fig. 4(d)). Thisphenomenon is attributed to the dominance of electron tunneling at filler concentrations close to the percolation threshold; the distance between filler particles increases when thesensors are stretched while still allowing conduction throughtunneling. For higher concentrations of filler material, theconnections are made mostly by direct contact which are lesseffected by strain than tunneling electrons. This result is similar to other reported strain sensors made from graphene where
it was found that the GF was dependent on initial resistance40
and suggests that high sensitive strain sensors can be madefrom CPDMS near the percolation threshold of the filler material. The repeatability of GF varied based on concentration.
Lower filler concentrations were associated with much higher
variabilities (standard error values approaching "20% of theGF) whereas the standard error of the GF of the higher concentrations were within "10% of the GF, suggesting that
CPDMS at lower concentrations were much more sensitive toprocessing than higher concentrations.
In summary, strain sensors featuring simple, low-cost
construction involving the screen printing of MWNT and
GNP nanocomposites on biocompatible medical grade polymer substrates was achieved. The biocompatibility of thesubstrate, low cost of the GNPs, and strain sensitivity and performance, including high strain operation and gauge factors,of these sensors uniquely enable applications in wearable andimplantable sensors in health diagnostics and monitoring.
This work was funded in part by an NSF CAREER
Award (ECS-0547544) and the University of Southern
California Viterbi Fellowship (CL). The authors would liketo thank Dr. Donghai Zhu, Mr. Eric Welder, and the members of the USC Biomedical Microsystems Laboratory for
their assistance.

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