IEEE SENSORS JOURNAL, VOL. 7, NO. 12, DECEMBER 2007
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Parylene-Based Strain Sensors for Bone
Gloria Y. Yang, Student Member, IEEE, Garrett Johnson, William C. Tang, Senior Member, IEEE, and
Joyce H. Keyak
Abstract—This paper presents a novel flexible implantable device to provide high-resolution mechanical strain data from a bone
surface in real time. The design of the device has been verified with
finite-element analysis, and a prototype has been successfully fabricated which consists of a thin-film metal gauge encapsulated between two layers of parylene-C. The prototype has been characterized simultaneously with a commercial strain gauge using tensile
testing. The results indicated that the strain sensitivities of the prototype were approximately 2.5 times greater than those of commercial gauges. In addition, real-time strain data collection has been
successfully demonstrated on bone surfaces with the novel devices
using mechanical testing of chicken tibiae in three-point bending.
Index Terms—Bone, parylene, piezoresistive devices, strain
measurement.
I. INTRODUCTION
ONITORING strain on the surface of bones in real time
would allow better understanding of the biomechanical behavior of the musculoskeletal system. Knowledge of
bone strains would also facilitate advances in musculoskeletal
diagnostics, rehabilitation monitoring, and feedback, as well
as improved data collection and clinical studies to develop
advanced orthopaedic implants. However, currently available
devices for measuring strain are too large (typically 2 5 mm)
to provide measurements with suitable resolution. These
gauges are also difficult to mount on bones because of their
large size and the bone’s irregular surface topology. If the
strain sensors were formed using a thin and flexible membrane
as the carrier material, the gauges would conform to the bone
surface better, and thereby provide more accurate strain data. In
addition, gauges with greater strain sensitivity would provide
more detailed measurements with better precision. Previously,
flexible strain gauges with polydimethylsiloxane (PDMS) as
the carrier material were developed [1]–[3]. However, it was
found that only silicone-based adhesives would form strong
adhesion between the PDMS and the bone surface. This type
M
Manuscript received January 1, 2007; revised August 29, 2007; accepted August 31, 2007. The associate editor coordinating the review of this paper and
approving it for publication was Prof. Fabien Josse.
G. Y. Yang is with the Department of Electrical Engineering and Computer
Science, University of California, Irvine, CA 92697 USA (e-mail: gyyang@uci.
edu).
G. Johnson is with the Department of Biomedical Engineering Department,
University of California, Irvine, CA 92697 USA (e-mail: [email protected]).
W. C. Tang is with the Departments of Biomedical Engineering and Electrical
Engineering and Computer Science, University of California, Irvine, CA 92697
USA (e-mail: [email protected]).
J. H. Keyak is with the Departments of Orthopaedic Surgery and Biomedical Engineering, University of California, Irvine, CA 92697 USA (e-mail:
[email protected]).
Color versions of one or more of the figures in this paper are available online
at http://ieeexplore.ieee.org.
Digital Object Identifier 10.1109/JSEN.2007.909923
Fig. 1. Drawing of the gauge design. L is the gauge length, w is the trace width,
p is the pitch between turns, and n is the number of turns.
of adhesives usually requires a day to be fully cured at room
temperature, making PDMS-based sensors less desirable for
measuring bone surface deformations in vivo. Hence, this paper
presents a new type of carrier material for strain sensors for this
specific application.
II. DESIGN AND SIMULATION
A. Design
Chlorine-substituted poly-para-xylylene (Parylene-C) has
been widely used as a material to encapsulate medical devices
such as cardiac pacemakers [4], probes for neural prostheses
[5], neurocages [6], and cell manipulation platforms [7]. Its
wide popularity is due to its biocompatibility, low moisture
vapor transmission rate compared with epoxy, silicones, and
urethanes [8], and ease with which a conformal layer can be
deposited at room temperature. Therefore, Parylene-C was
chosen as the carrier material for strain sensors reported in this
paper.
The operating principle of the current strain gauge is based on
piezoresistivity as the transduction mechanism. The fractional
change in resistance is proportional to the applied strain in direct
response to induced stress. The gauge piezoresistive metal was
designed with a serpentine pattern with a target resistance of
180 . Previously, it was found that metal trace failures often
occurred at the corners of the turns during fabrication. In order
to increase the fabrication yield and mechanical robustness of
the gauges, an end loop at each turn and tapered traces from
the bond pads were incorporated in the design, as illustrated in
Fig. 1.
To ensure that attachment of the strain gauge to the bone surface would produce a negligible effect on the bone behavior,
the stiffness of the device should be much less than that of the
target bone. The Young’s modulus of Parylene-C is between
2 and 5 Ga [5], which is an order of magnitude lower than that
of human cortical bone (20.5 GPa) and trabecular bone tissue
(18 GPa) [9], [10]. According to [11], the cortical thickness
1530-437X/$25.00 © 2007 IEEE
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IEEE SENSORS JOURNAL, VOL. 7, NO. 12, DECEMBER 2007
varied from 0.3 mm at the superior half of the human femoral
neck to 6 mm thick at the inferior half of the neck. The matrix
length and width of the sensor of 6.1 and 4.1 mm, respectively,
and the cortical thickness of 1 mm were used in the analysis of
sensor influence on the bone biomechanics. From (1) and (2), a
total Parylene-C membrane thickness of 12 m, which was two
orders of magnitude less than the calculated requirement, was
chosen for the device reported in this paper
(1)
(2)
where is the stiffness, is the cross-sectional area, is the
Young’s modulus, is the length of the senor or bone, and is
the thickness.
A Wheatstone bridge circuit configuration was usually
adopted to measure the change in the resistances of the strain
gauges. According to [12], the power consumption of the
Wheatstone bridge was inversely proportional to the gauge
resistance, as shown in (3). In this design, the overall power
budget was estimated to be 100 W with gauge resistor of
100 . As our device resistance is higher than 100 , hence
less power would be consumed. The power could be supplied to
the sensor through inductive coupling once it is fully implanted
Fig. 2. Example of simulated plots of strain gauges by ANSYS with tensile
deformation of 20 m in uniaxial direction. These showed the stress distribution (left plot) and strain distribution (right plot) in the gauge region. The color
scheme indicates different stress and strain levels. In the stress distribution plot,
navy blue
–53 MPa and green
–241 MPa. In the strain distribution
–.025% of strain, and teal
–0.27% of strain.
plot, light blue
:
:
=5
= 0 18%
= 194
= 0 25%
(3)
where
is the power dissipated through the Wheatstone
is the initial rebridge, is the gauge factor of the sensor,
is the minimum measurement
sistance of the strain gauge,
is the input noise of the first amplifier, and
resolution,
is the bridge bias voltage.
Fig. 3. Process flow for the microstrain sensor. (a) Deposit a stress-free
flexible membrane of Parylene-C (6-m thick) as the bottom structural layer.
(b) E-beam evaporate 10-nm Cr and 120-nm Au, followed by wet etch to
pattern the metal traces. (c) Deposit the top structural layer of Parylene-C of
6-m thickness. (d) Etch Parylene-C with oxygen plasma to expose the wire
bonding pads. (e) Peel off the device from silicon wafer, and bond wires to
electrode pads.
B. Simulations
To verify the designs, ANSYS finite-element modeling tool
was used. The strain and stress distributions within the gauges
were investigated when they were under tensile deformations
in the and directions. The model included a thin-film gold
m,
m,
m
gauge with dimensions of
, embedded in the center of a 1 cm 1 cm
and
Paylene-C membrane. A 2D-model was adopted, since thickness of the gauges was approximately two orders of magnitude
less than the width of the traces. In all simulations, deformations
were applied to the edges of the Parylene-C membrane.
Fig. 2 shows a uniform stress distribution along the gauge
lengths and high stress concentrations at the corners of the turns.
The gauge was uniformly elongated with average strain close
to the applied strain of 0.2%, validating the feasibility of the
Parylene-based strain gauge design.
III. FABRICATION
The thin metal film gauges sandwiched by membranes of
Paylene-C were fabricated using surface micromachining technology. The process only requires two masks and the steps are
described in details below and shown in Fig. 3.
1) First layer of Parylene-C: A 6- m Parylene-C layer was
deposited using the Parylene deposition equipment (Specialty Coating Systems, Model PDS-2010 Labcoter© 2).
The Parylene thickness depended on the amount of evaporated polymer. It was 1 m of Parylene-C deposited per
1 g of polymer.
2) Patterning metal traces: A 120-nm gold film with a 10-nm
chrome adhesion layer was deposited onto the Parylene
layer with e-beam evaporation. Photolithography was used
followed by wet etching to pattern the metal gauges.
3) Second layer of Parylene-C: Another 6- m Parylene-C
layer was deposited to encapsulate the metal traces.
4) Exposure of bond pads: Since the etch rate of AZ-4620
photoresist is faster than Parylene-C [13], a thicker layer
m was patterned on the
of AZ-4620 photoresist
Parylene-C. The Parylene-C was etched using reactive ion
etching (RIE) (Plasma-Therm 790 Series), with oxygen
flows set at 50 sccm, pressure at 200 mT, and power at
350 W. The etching rate was found to be approximately
0.5 m min.
5) Device separation: Each individual device was peeled off
from the silicon wafer.
YANG et al.: PARYLENE-BASED STRAIN SENSORS FOR BONE
Fig. 4. Example of fabricated device. The matrix length and width of the device
were 6.1 and 4.1 mm, respectively. The gauge length and grid width of the sensor
were 350 and 390 m, respectively. Inset shows the enlarged picture of the
gauge region.
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Fig. 5. Sample under test in the Instron machine. Inset shows the extensometer
in use to measure the reference strain in the region where our device and the
commercial gauge were attached.
Gauges with two different lengths, 350 and 500 m, width of
10 m and pitch of 10 m between each turn were fabricated,
as shown in Fig. 4.
IV. CHARACTERIZATIONS AND RESULTS
Two types of experiments were conducted to characterize the
performance of the devices.
A. Electromechanical Testing
Aluminum rods were machined according to ASTM E8-01
standard [14], with a diameter of 6.35 mm (0.25 in.) at the reduced section and gauge length of 2.54 cm (1 in.). The surface of the rods was pretreated with ethanol to remove oil and
grease, then abraded with silicon-carbide paper to remove thin
oxides, and finally reapplied ethanol repeatedly till no discoloring occurred. All devices were installed within 30 min on the
aluminum rods after the surface preparation.
The strain gauges were affixed onto the aluminum rods with
certified M-Bond 200 (Vishay Micro-Measurements) adhesive.
Commercial strain gauges (Omega Engineering, Inc.) were also
installed on the rods next to our devices using the same procedures. After wires were soldered on the bonding pads, epoxy
adhesive (Hysol M-31Cl, Loctite) was used to encapsulate the
bonding regions to provide robust wire bonds, as well as a protective coating. The specimens were left at room temperature
for more than 72 h before tensile testing was performed.
In order to obtain repeatability of the results, aluminum rods
would need to be maintained within the elastic region when the
devices were under test. To do so, an Instron testing machine
(model 3367) with an extensometer was used to characterize the
mechanical properties of the aluminum rod by applying longitudinal tensile forces. From the result, the yield strain of the rod
was found to be approximately 0.2%. The Young’s modulus,
yield strength, and ultimate tensile strength were also found to
be 58 GPa, 128 MPa, and 173 MPa, respectively. These values
were close to those provided by the manufacture, hence validating the testing method and setup.
The Instron machine was again used to apply tensile forces
to the aluminum rods with strain gauges on them. The extensometer was placed in the reduced section of the rod across
the gauges for measuring the reference strain (as illustrated in
Fig. 5).
The gauges under test were prestrained three times to 10%
, to ensure
more than the intended test strain, i.e., 1000
Fig. 6. Strain sensitivities of the measured Parylene-based gauge and commercial gauge. The Parylene-based gauge demonstrated repeatable data when it was
tested from 0% to 1% of strain for six test runs.
that the yield strain of the aluminum rod was not reached. The
strain loading cycles were
Each strain gauge was tested in at least four runs. The
resistance changes of our device and the commercial strain
gauge were measured with digital multimeters (HP34401A)
and recorded in real-time simultaneously by computers, while
the devices were tested.
The relative change in resistances was calculated in order
to find the gauge sensitivity. Examples of measured results are
plotted in Fig. 6. The gauge factor , can be obtained from the
slope of the curves with the following equation:
(4)
where
is the percent resistance change, and is the strain
experienced by the strain gauge.
The average gauge factor of the Parylene-based gauges was
2.56, with standard deviation of 0.06, and standard uncertainty
of 0.015 over 16-test runs with three different gauges from the
same batch. The average gauge sensitivity of commercial strain
gauges was calculated to be 1.03 with standard deviation of
0.097 and standard uncertainty of 0.024 over 15- test runs with
three gauges, although the gauge factor provided by the manufacturer was 2. An example of measured strain sensitivity and
the calculated correlation coefficients from a tested Parylenebased and a commercial gauges over six test runs were listed
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IEEE SENSORS JOURNAL, VOL. 7, NO. 12, DECEMBER 2007
TABLE I
COMPARISON OF THE SENSITIVITY, LINEARITY AND REPEATABILITY OF
A PARYLENE-BASED AND A COMMERCIAL GAUGE
in Table I. The measured gauge sensitivities of the Parylenebased gauges were found to be approximately 2.5 times higher
than those of the commercial gauges. This could be due to different materials used, i.e., ours was gold and the commercial
gauge was made of constantan. Different sensor dimensions
could also play a role, especially when the thicknesses differ
by an order of magnitude. Moreover, different manufacturing
process (e-beam evaporation against electroplating) could cause
the grain size and grain boundary thickness of the metal materials to differ and, hence, contribute to higher gauge factor.
The Parylene-based gauges also demonstrated a linear relationship with changes in resistance versus strain for a large range
of strain. The data showed a high degree of linearity with correlation coefficients, , higher than 0.9991 and 0.9929 for the
Parylene-based and commercial gauges, respectively.
Furthermore, it was observed that the metal traces of the Parylene-based gauge yielded, causing open circuit, when it was
elongated to 2.2% of strain. According to [15], the tensile yield
strain for human cortical bone and trabecular tissue were found
to be 0.73 0.05% and 0.62 0.04%, respectively. The measured results demonstrated that the device was capable of monitoring strain in the required strain range below the yield strain,
i.e., 0 to 0.5%, for nondestructive bone testing in vivo. The Parylene-based gauges are also suitable for measuring surface strain
in the osteoporotic femoral head, where the strain levels (av) were found to be 70% higher on average
erage of 520
) [16].
than those in a healthy one (average of 304
Fig. 7. (a) Sample under test in the MTS hydraulic machine. (b) Close-up
picture of the bone specimen after it had been fractured into two pieces. The
Parylene-based gauge also broke into two pieces, indicating its ability to conform to the bone surface deformation well and to measure the maximum strain
when the bone fractured.
B. Three-Point Bending on Chicken Tibiae
Chicken tibiae were chosen as test specimens for the threepoint bending for convenience and cost considerations. To expose the bone surfaces, the soft tissues on the chicken tibiae
were removed with a scalpel. The ends of the bones were cut, so
that they could fit into a metal mold. Polymethyl-methacrylate
(PMMA) was mixed in 1:5 ratio by volume. The mixed solution was poured into the mold where the bone had been fixed
in the desired location. Once it was cured, 2.54 cm (1 in.)-cubic
PMMA was formed at each end of the tibiae. The tibiae were
wrapped with wet cloths to maintain the moisture on bone surface during the molding.
Ethanol was used to remove the grease on the bone surfaces
before strain gauges were installed with M-Bond 200 adhesive.
Parylene-based gauges and commercial gauges were placed at
equidistance (5 mm) from the midpoint of the tibiae. Epoxy was
Fig. 8. Strain measurements of the three chicken tibiae. The top graph shows
the force that was applied to the tibiae versus time. The bottom graph shows the
strain measurements of Parylene-based and commercial gauges versus time.
used to secure the bonding sites after wires were soldered onto
the bond pads.
A materials testing system (MTS) servo-hydraulic testing machine was set up to apply three-point bending loading with free
ends to the chicken tibiae (as shown in Fig. 7).
A force was applied to the center of the bone, on the side opposite the gauges. In such cases, the gauges experienced tension.
Both gauges monitored the bone surface strain simultaneously
YANG et al.: PARYLENE-BASED STRAIN SENSORS FOR BONE
as the force was applied, and the resistances of the sensors were
measured in real-time by digital multimeters and recorded by
computers. Since strain gauges could not be reused, new strain
gauges were installed on each chicken tibia for testing.
Bone surface strains were calculated from the measured
changes in resistances recorded from both Parylene-based
gauges and commercial gauges. As bone surfaces are irregular,
the strain distributions would be different as the positions
varied. This caused the measured strain levels from both gauges
to differ slightly, as shown in the plots for specimen 1 and 3
in Fig. 7(b). However, the curves from both gauges in the plot
exhibited similar profiles for specimen 2. The Parylene-based
gauges demonstrated the ability to measure strain up to 2.6%,
at which point the bone fractured into two as illustrated in
Figs. 7(b) and 8.
V. CONCLUSION
This paper has presented the design and fabrication of a
microstrain gauge encapsulated in a biocompatible Parylene-C
membrane. It has been shown that these gauges exhibit greater
strain sensitivities, hence providing better strain data resolution
than that of the commercial gauge tested in this work. The
capability of these gauges to monitor strain on chicken tibiae
has also been demonstrated. Ultimately, this Parylene-based
strain gauge combined with a wireless telemetry circuit can be
used to monitor the real-time surface deformation of a bone
in vivo.
ACKNOWLEDGMENT
The authors would like to thank T. Kaneko, W. Dang,
Y.-H. Hsu, and W.-M. Wong for their technical assistances.
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Gloria Y. Yang (S’04) received the B.E. degree in electrical engineering and
the M.S. degree in biomedical engineering from the University of New South
Wales, Sydney, Australia, in 2000. She is currently working towards the Ph.D.
degree in the Department of Electrical Engineering and Computer Sciences at
University of California, Irvine.
Garrett Johnson was born in Riverside, CA, and now lives in Irvine, CA. He
is a fourth year biomedical engineering major and plans to work in the biotechnology field after graduation. He also has plans to attend graduate school in the
future.
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William C. Tang (S’86–M’90–SM’02) received the B.S., M.S., and Ph.D. degrees in electrical engineering and computer sciences from the University of
California at Berkeley, Berkeley, in 1980, 1982, and 1990, respectively. His invention and seminal thesis work on the electrostatic comb drive has become one
of the crucial building blocks for many sensors and actuators in the MEMS field.
His career includes developing MEMS for the automotive industry, while
working at the Ford Research Laboratory, Dearborn, MI, and Ford Microelectronics, Inc., Colorado Springs, CO. In 1996, he joined the Jet Propulsion
Laboratory, Pasadena, CA, and pursued MEMS applications for space explorations. In 1999, he assumed the responsibility as the Program Manager for
the MEMS programs at the Defense Advanced Research Projects Agency. In
2002, he joined the faculty as a Professor with the Department of Biomedical
Engineering, University of California, Irvine, with a joint appointment with
the Department of Electrical Engineering and Computer Science. His current
research interests are in micro and nanoscale technologies for wireless medical
implants and microbiomechanics.
Dr. Tang is a member of the Biomedical Engineering Society, a Fellow and
Chartered Physicist with the Institute of Physics, and a Fellow of the American
Institute for Medical and Biological Engineering.
Joyce H. Keyak did her undergraduate work at the University of California,
Berkeley, in mechanical engineering and received the Ph.D. degree from
the University of California, San Francisco/Berkeley (Joint Program in
Bioengineering).
She specializes in the biomechanics of bone and in evaluating the effects of
osteoporosis and tumors on bone strength.