Micromachined Antenna Stents and Cuffs for
Monitoring Intraluminal Pressure and Flow
Kenichi Takahata, Member, IEEE, Yogesh B. Gianchandani, Senior Member, IEEE, and
Kensall D. Wise, Fellow, IEEE
Abstract—This paper describes two stainless steel microstructures that are microelectrodischarge machined from 50-m-thick
planar foil for intraluminal measurements of pressure and flow
(with potential for applications ranging from blood vessels to bile
ducts). The first structure is an inductive antenna stent (stentenna)
with 20-mm length and 3.5-mm expanded diameter. It is coupled
with capacitive elements to form resonant LC tanks that can be telemetrically queried. The resulting LC tanks are deployed inside silicone mock arteries using standard angioplasty balloons and used
in a passive telemetry scheme to sense changes in pressure and flow.
Using water as the test fluid, the resonant peaks shift from about
215 to 208 MHz as the flow is increased from 0 to 370 mL/min.
The second structure is a ring-shaped intraluminal cuff with two
400 2 750-m2 electrodes that are used to provide a direct transduction of flow velocity in the presence of a magnetic field. It is
fabricated in a manner similar to the stentenna, but with an insulating segment. The voltage has a linear dependence on flow rate,
changing by 3.1–4.3 V per cm/s of flow (of saline) over a 180 cm/s
dynamic range, with a magnetic field of about 0.25 T.
Index Terms—Bloodflow, cardiac, microelectrodischarge machining, sensor network, wireless.
N recent years, stents have come to play an essential role
in the treatment of cardiovascular and other diseases. Most
commonly implanted in the coronary artery, the cardiac stent
typically has mesh-like walls in a tubular shape, and once positioned by a catheter, is expanded radially by the inflation of
an angioplasty balloon. Its primary task is to physically expand
and scaffold blood vessels that have been narrowed by plaque
accumulation. However, renarrowing (restenosis) often occurs
due to blood clot (thrombus) formation, excess growth of intravascular tissues (proliferation), and further plaque deposition
[1]. While the availability of drug-eluting stents can suppress
these failures in certain cases, chronic monitoring of pressure
or flow is still needed in others. Another category of vascular
problems in which stents are used is aortic aneurysms, which
Manuscript received December 11, 2005; revised February 20, 2006. The exploration of microdischarge based process methods was supported in part by a
Grant from the National Science Foundation. The design and fabrication of the
pressure sensors was supported by the Engineering Research Centers Program
of the National Science Foundation under Award Number EEC-9986866. Subject Editor S. Shoji.
K. Takahata was with the Department of Electrical Engineering and Computer
Science, University of Michigan, Ann Arbor, MI 48109 USA. He is now with
the University of British Columbia, Vancouver, BC V6T 1Z4, Canada (e-mail:
[email protected]).
Y. B. Gianchandani and K. D. Wise are with the Department of Electrical
Engineering and Computer Science, University of Michigan, Ann Arbor, MI
48109 USA (e-mail: [email protected]; [email protected]).
Digital Object Identifier 10.1109/JMEMS.2006.880229
Fig. 1. A wireless link to an implanted sensor can be monitored by a portable
device placed in the residence or on the person of the patient. This can then
provide a link to a database through the internet or a broader wireless network,
permitting physicians to review the sensed parameters.
may be thoracic or abdominal. Such aneurysms have been surgically treated in the past, but the use of stents for this purpose
is increasing [2], [3]. While use of wireless pressure sensors
has been reported in association with these stents [4], the sensors are located in the aneurysm, not within the stent or path
of blood flow. For this reason, miniaturization is not essential,
and sensors of 1–2 cm length can be accommodated. Stents are
also used in the carotid artery [5], bile duct [6], pancreatic duct
[7], tracheobronchial airways [8], and in the esophagus [9] to
treat diseases ranging from atherosclerosis to carcinoma. Since
their intent is inevitably to facilitate the movement of some kind
of fluid, the innate ability to monitor intraluminal pressure or
flow can provide advance notice of the need for further tests or
For many applications, the absolute value of pressure and
flow is not required; a simple change in the average or instantaneous signal over time can provide a warning to be heeded by
a more direct and precise diagnosis. This eases the burden of
functional performance on the implanted device, and opens up
some options for engineering simpler solutions.
The use of wireless telemetry can alleviate a lot of the practical challenges associated with monitoring implanted sensors.
For example, pressure can be monitored using a microchip that
has a planar thin film inductor integrated with a micromachined
capacitive pressure sensor [10]–[14]. This LC tank circuit couples to a separate external transmitting coil via mutual inductance, responding to a change in pressure by a shift in the frequency at which the external coil shows a characteristic dip in
impedance or phase. If a stent is able to serve as an inductor/antenna and is integrated with microsensors, the planar inductor
can be eliminated from the LC tank (saving considerable space),
and the stent can become an inherent element in the wireless link
(Fig. 1).
1057-7157/$20.00 © 2006 IEEE
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While most commercially available vascular stents are made
by laser machining of stainless steel tubes, other constructions—such as threaded wire structures, or rings made from
shape-memory alloy and sutured to woven polyester—are
also used [3], [15]. It has been recently shown that stents cut
from planar metal foils by microelectrodischarge machining
offer appropriate mechanical properties [16]. These
stents use flexural designs and do not have any bonded or
welded seams. In assembling the device, a deflated angioplasty
balloon is threaded alternately above and below a series of
involute cross bands that lie between two longitudinal side
beams, and then expanded by a normal angioplasty procedure.
This design approach permits stents to be fabricated from steel
foil by using planar electrodes as “cookie cutters” that are lithographically fabricated on a silicon substrate [17]. It also offers
compatibility with other planar microfabrication processes
based on lithographic techniques. While the stents reported in
[16] were simply mechanical devices, it was recognized that
inductive coils could also be fabricated in this way.
This paper explores the possible use of stent-like structures
micromachined from planar foils for chronic monitoring of
intraluminal conditions within an artery, duct, fistula, or other
flow path.1 Two complementary approaches are investigated.
The first one, described in Section II, explores a structure that
integrates silicon-micromachined capacitive pressure sensors
with a dual-inductor stentenna to form a device for wireless flow
measurement. The second approach, described in Section III,
explores structures for the electromagnetic (EM) detection of
flow [22], [23] in a polarizable liquid. Section IV addresses
some of the ways that performance may be improved in particular applications, whereas Section V draws conclusions from
the overall effort.
When a liquid flows through a channel, there is a pressure
drop between two separate locations that depends on the flow
rate. A general expression for this drop in steady-state flow
is [24]
are pressures at downstream and upstream lowhere
cations, respectively, is area-averaged flow velocity in an unobstructed vessel, and
are coefficients that depend on
obstacle geometry and fluid properties. The first term is associated with a loss due to viscous shearing stress, and the second
is due to geometry variation inside a channel, such as by redeposited plaque or excess tissue growth over a stent. As the
obstruction grows, the nonlinear term tends to dominate. With
a correction for average pressure, changes in the local pressure at a site (that has been strategically selected for chronic
monitoring) can provide indication of an impending medical
problem. Further, if there are two pressure sensors that are located upstream and downstream of a blockage, they can be used
1Portions of this manuscript have been presented in conference abstract form
in [18]–[21].
differentially to determine the absolute flow, and a separate measurement of average pressure is not needed. However, if the two
pressure sensors are not used differentially but are instead used
to provide a single averaged reading, their output can be indicative of changes in flow due to blockages that are upstream or
downstream of both of them.
A. Stentenna Design
For an effective wireless link, minimal damping is desired in
the LC tank (Fig. 1). The quality factor, , is expressed as
is inductance of the stentenna,
is capacitance
is parasitic capacitance, and
is paraof the sensor,
is greater than that of
sitic resistance. The impact of the
. For a given material, and at frequencies low enough that
the skin effect is not a factor, the parasitic resistance of the stentenna is inversely related to the cross sectional area of its beams.
However, the parasitic capacitance that it contributes is propordepends on the
tional to the beam surface area. Therefore,
square of the beam diameter whereas
is simply proportional to it. Thus, it can be electrically favorable to increase the
thickness of the beams. In fact, this is favorable mechanically as
well, because it would increase the radial stiffness of the stent.
However, from the biological viewpoint, increasing the volume
of the structural elements can be undesirable, and may warrant
application-specific designs and structural optimization. In this
cross section.
effort the stentennas had beams of 50 50
This is similar to the purely mechanical stents reported in [16],
and at the lower end of the range offered in existing commercial
The planar microstructure that was used in this effort is illustrated in Fig. 2(a). It has a series of cross bands that have involute contours with a bridge to a longitudinal beam at the center
of the device. The involute bands form dual inductors, whereas
the beam is a common electrical node. At each end of the device
the bands terminate in a section that forms a ring. This provides
enhanced mechanical rigidity in the pre-expansion state as the
angioplasty balloon is inserted into it. It also provides improved
radial stiffness after expansion. Capacitive pressure sensors are
connected across the common line and a lead that is connected
to the ring, thereby implementing two LC tanks when complete.
During the deployment of a stent, the structure is pushed
against the walls of the lumen as the angioplasty balloon is
inflated, so it is necessary to consider the immunity of the
micromachined pressure sensors from overpressure in this step.
While capacitive pressure sensors have natural overpressure
protection because the diaphragm deflection is constrained by
the substrate, the range of pressures used in balloon angioplasty
(4–12 atm.) can still be of some concern. Having platforms
for the sensors in the stentenna structure not only permits
rigid bonding of the pressure sensors, but also helps to protect
them during balloon inflation. Several preliminary experiments
were conducted with dummy samples that were pressurized
by balloons inside 3-mm i.d. silicone mock arteries with
0.25-mm-thick walls, which are manufactured for the purpose
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EDM patterning. (b) A stentenna integrated with two sensor chips is mounted on a deflated angioplasty balloon.
Fig. 2. (a) Stentenna design for of testing vascular implants and have compliance similar to
human arteries (Dynatek Dalta Scientific Instruments, MO)
[25]. They showed, in fact, that the sensors had no damage and
were still functioning even after full expansion of the stent to
3.5-mm diameter. However, to provide an extra measure of
security, as shown in Fig. 2(b), the stent was designed to locate
the sensors outside the longitudinal span of the balloon, which
had a 16-mm length in this case.
B. Device Fabrication
Although stents that serve purely mechanical purposes do not
require any insulation, the stentenna plays an electrical role and
does require insulation from the surrounding fluid. Electrical
isolation is particularly needed between the cross bands of the
stentenna that might otherwise contact each other and shunt out
the inductance if the angioplasty balloon expands nonuniformly.
Further, this insulation should cover the micromachined pressure sensors and lead transfers to them, and be biocompatible
as well. Parylene-C was chosen as the coating material because
it provides a thin, uniform and conformal coating that is nonconductive, chemically inert, and biocompatible. It also has a
history of applications to biomedical devices, including cardiac
stents [26].
The majority of commercial stainless-steel stents use type
316L steel. This differs from type 304, which is one of the
most common and versatile grades of steel. Type 316 steel has
molybdenum (2%–3%), which gives it better overall corrosion
resistance than 304; type 316L is a low carbon version of 316,
which is immune to carbide precipitation at grain boundaries.
However, a comparison of the important material properties of
these three kinds of steel (Table I) shows that for in vitro tests,
they have very similar characteristics. Consequently, in this effort 304 steel was used because it was easily available in 50
thick foils.
Fig. 3 shows the process flow using a cross-sectional view at
- in Fig. 2(a). First, the steel foil is patterned by
Fig. 3. Fabrication flow. (a) stainless steel foil. (b) Glue down sensor
chips. (c) Electrically connect to terminal pads of the sensors. (d) Deposit passivation layer. (e) Release device and apply epoxy to the perimeter of the sensors.
“a”). The machined structure remains connected to the original
sheet at this point for ease of handling during subsequent steps.
Pressure sensors fabricated on a 500- -thick glass substrate
chips in advance. Two sensors are
are diced into 1.4 1.8
attached to platforms which are linked at the ends of the longitudinal beams extended from the rings (step “b”). The bonding
is performed with enamel, which offers good adhesion and mechanical strength, and it is also easily applied and cured rapidly.
Perforations in the platforms serve as escape paths for excess
enamel. The sensors are then electrically connected to the leads
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Fig. 5. A setup for wireless pressure/flow sensing.
Fig. 4. (a) Deployed device with balloon removed. (b) Equivalent circuit after
expansion completed.
with a conductive adhesive (step “c”). The device is next coated
with 0.5- -thick parylene everywhere for electrical insulation (step “d”). Since this is a vapor-phase deposition step, the
coating is conformal. The device is then mechanically released
from the sheet (step “e”). Finally, additional epoxy is applied to
the perimeter of the sensors to enhance the bonding strength.
The micromachined pressure sensor consists of a
silvacuum-sealed cavity capped by a 3.7- -thick
icon circular diaphragm with a 1-mm diameter and 5[29]. The diaphragm has a 10- -thick boss for better linearity
and an oxide layer on the backside for electrical protection in
case of contact between the diaphragm and a bottom electrode,
which is a thin film of Ti/Pt/Au patterned on a 500- -thick
Pyrex glass substrate. In the experiments described here, the
glass substrate of a pressure sensor that is upstream of the
stentenna was thinned (by wet etching) down to 100
order to create a more streamlined profile. Of course, this can
be done to the other pressure sensor as well.
The fabricated devices are then deployed inside a mock artery
using a standard angioplasty balloon with a 3.5-mm inflated
diameter. In this step, the stentenna becomes permanently deformed from a planar into a helical shape, which consists of two
separate coils with 3 and 3.5 turns in this case (Fig. 4). As noted
previously, the two sensors are located just off the region occupied by the balloon, minimizing the physical impact to both the
sensors and the tube wall.
C. Experimental Setup and Results
Fig. 5 shows the fluidic test setup used to evaluate the device in Fig. 4. A pump/flow-controller regulates the flow (of
water), and a separate meter (PS309, Validyne Engineering Co.,
CA) measures the differential pressure drop along the 8-cm-long
tube. To simulate a blockage, a dielectric rod with 1.5-mm diameter is positioned inside the stentenna. The complex input
impedance of an external coil is monitored with an HP4195
spectrum analyzer. The output power fed from the analyzer to
, which makes the power denthe external coil is
recomsity well below a biological threshold of 10
mended by ANSI for the purpose of protecting human health
[30]. The stentenna inductance is approximately 110 nH in total.
The pressure sensors have a measured response of
[see Fig. 6(a)], which reduces to
with a 1.3-thick parylene coating. Sensor 1, which has lower capacitance,
Fig. 6. (a) Measured response of uncoated capacitive pressure sensors.
(b) Measured gauge pressure versus flow rate with and without the regulation
rod that serves as a partial obstruction to the flow.
is coupled to the 3-turn inductor, and sensor 2 is paired with the
3.5-turn inductor, so that these LC tanks have different resonant
frequencies. If the quality factor of these resonant tanks was sufficiently high, they would produce discernible peaks yielding
the upstream and downstream pressures separately. However,
the present implementation offers relatively low inductance due
to limitations in machining resolution, which means that the
peaks will be merged together.
Fig. 6(b) shows gauge pressure at upstream and downstream
locations as a function of flow rate, which is measured with the
setup of Fig. 5. The plot also shows its dependence on the obstruction, which causes a quadratic dependence on flow rate that
is consistent with (1). The nonlinearity in the curve for unobstructed flow (i.e., no rod) comes from the flow resistance of the
device itself.
An impedance peak that is nominally at 239.1 MHz in a 4-mm
diameter external coil with inductance of 610 nH is shifted down
by increasing flow rate as shown in Fig. 7(a). Fig. 7(b) plots
this frequency shift and the corresponding differential pressure
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quantitatively very sensitive to the values of the parasitic elements as well as the strength of the wireless coupling between
the three inductors, and should be used with caution for predictive modeling.
The principle of the electromagnetic flow sensing is based on
Faraday’s law of magnetic induction. When polarizable liquid
passes through a magnetic field at an angle, a voltage is generated across the flow channel, in a manner analogous to the Hall
effect. This voltage is sensed by two electrodes that are placed
on the channel walls, and is given by [22]
Fig. 7. (a) Measured amplitude of input impedance Z in an external coil
shifted near 239 MHz due to flow change. (b) Measured resonant frequency and
differential pressure versus flow rate.
Fig. 8. A possible model for the microsensor system.
drop: there is a reduction of 9–31 KHz per mL/min. increase
in the flow range over 370 mL/min. (Typical coronary artery
flow is 100–200 mL/min.) The pressure response observed is
57.4 KHz/torr (at gauge pressure of 113 torr), while the corresponding sensitivity is 273 ppm/torr. These preliminary measurements were taken in close proximity, with the receiving coil
away from the stentenna.
Fig. 8 shows a qualitative model for the wireless setup in
, and
, respecFig. 5. In this model,
tively, denote inductance of an external coil, inductance for the
downstream inductor, that for the upstream inductor, capacitance of sensor 1, and that of sensor-2. The two inductors of the
device are mutually coupled to the external coil. It also shows
, and
, which are
parasitic elements
capacitance for the external coil, resistance for the coil, lumped
resistance for the downstream LC tank, and that for the upstream
tank, respectively. Other parasitics (not shown) include capacitances that are associated with the stentenna, capacitive pressure sensors, and their packaging. While this model is helpful
for the qualitative understanding of the device operation, it is
where is the diameter of the flow channel, is the magnetic
flux density and is the cross-sectional average velocity of the
flow. In a Cartesian coordinate system, if the flow is parallel to
the axis and the voltage is sensed along the axis, then the angles and are, respectively, the angles made by the magnetic
field to the – and – planes. This equation assumes that (a)
the magnetic field is spatially uniform, and (b) the flow velocity
profile is axially symmetric. The condition (b) is valid downstream of a narrowed site if the location is reasonably distanced
from the blockage. The output voltage is independent of conduc),
tivity over a wide range [31] (which is preferably
and thus EM sensing is, in principle, widely applicable. It has
been used in diverse fields including food, chemicals, paper and
pulp, water supply, and energy supply industries [32].
Blood flow measurement has been a major application of EM
flow sensors [33]–[40]. A number of devices for blood flow
sensing that were based on this detection scheme were developed in late 1960’s and early 1970’s. Some of these included
electromagnetic or permanent magnet elements for the local
application of the magnetic field, whereas others relied upon
an externally applied one. Additionally, the devices could be
perivascular, cannular, or intravascular. The earliest perivascular
devices used sense electrodes that were placed external to the
lumen. As researchers developed ways to miniaturize the lateral dimensions of devices (in part by relying on an external
magnetic field), intravascular designs were developed. Not surprisingly, the devices used three–dimensional (3-D) construction methods, and required the assembly of multiple components made of different materials. In general, they were also intended for acute usage, either because of their relatively large
size, or because they were intended to remain at the tip of a
catheter that partially occluded the flow path. For chronic implants, of course, it is necessary to insure that fouling of the
sense electrodes does not interfere with the measurements.
A drawback of this method is its dependence upon orientation
of magnetic field. For example, when the magnetic field is tilted
by 15 for both and , the signal loss is estimated to be 6.7%.
This is not necessarily a major problem for every potential application. However, to the extent that correction is needed, one
past approach has been to include a calibration sensor and its
circuitry within the catheter-based structure [35]. Another used
a built-in electromagnet so that a magnetic flux with a fixed orientation is provided to the sensing site [36].
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Fig. 10. A setup for the flow measurement with the cuff.
Fig. 9. (a) A layout of the planar cuff structure. (b) Expanded cuff remains
attached to the inner walls of the tube by pressure after the balloon is deflated
and removed. (The lead wires have been removed.)
In this part of the effort, flow sensing is facilitated primarily
by an intraluminal cuff that senses the EM voltage. Fabricated
by a lithography-compatible approach similar to the one described for the stentenna, this structure is conformal to the inner
wall of the lumen, and has a low profile. It has two electrodes
located diametrically across from each other. These electrodes
are electrically insulated from each other by dielectric links
formed in the cuff. The voltage built up between the electrodes
is sensed by a separate component. While this device is not intended specifically for vascular applications, it is not incompatible with such use.
A. Fabrication and Testing
The planar microstructure for the cuff has a pair of meander bands comprised of 50- -wide beams, electrode plates
), and two dielectric links which mechanically
(400 750
tie the bands but electrically insulate them from each other
[Fig. 9(a)]. This pattern is cut by
into 50- -thick #304
stainless steel foil so that the two bands are connected to the
original foil at tethers shown in Fig. 9(a), maintaining 100-
gaps at the links. Insulating cement is used to bridge the gaps,
and then the device is released from the foil at the tethers. This
is similar to the process flow for a Kelvin probe described in
[41]. Lead wires are bonded to the electrodes with conductive
adhesive. All surfaces of the device except the front surfaces
of the electrodes are coated with an insulating layer. (Without
this, spatial averaging could reduce the signal voltage.) The
electrode may be optionally coated with an antifouling layer.
This feasibility experiment uses two-part epoxy and enamel for
the cement and the insulation layer, respectively. The fabricated
planar structure is mounted on a deflated angioplasty balloon
of a standard angioplasty catheter so that one of the bands is
located above the balloon whereas the other band is below it.
The cuffs were tested in silicone tubes with 3-mm i.d. To implant a cuff, the angioplasty balloon was inflated up to 3.5 mm in
diameter and then deflated and removed from the tube, leaving
the expanded cuff within the tube [see Fig. 9(b)]. Tests with
flow velocities up to 2 m/s show that both the structure and its
placement are robust and immovable. (As a point of compar). The
ison, the maximum arterial flow is typically
fluidic test setup is shown in Fig. 10. A pump/flow-controller
regulated the flow of 2% wt. saline and a voltmeter measured
voltage between the electrode leads. A permanent magnet with
was used to provide magnetic
dimensions of 25 25 9
field. The field orientation was perpendicular to both flow direction and the voltage sense axis defined by the locations of
the two electrodes. The magnetic field was characterized by an
InAs Hall sensor (BH-205, F. W. Bell, FL) and measured as
at the location of the cuff. The presence of the cuff
had no detectable impact on the externally measured magnetic
field. A baseline voltage can be associated with polarization and
electrochemical effects, so the flow-dependent voltage was measured with opposing orientations of the magnetic field (Fig. 11).
The transduction was linear and symmetric, and the voltage reper cm/s and 50–70 ppm
sponse and sensitivity were 3.1–4.3
per cm/s, respectively.
Using (3) with
, the calculated output voltages are plotted with a dotted line (Cal-1)
in Fig. 12(b). Although at low flow velocity the prediction
matches well with the experimental result, it deviates as the
flow velocity is increased. Fig. 12(a) shows the actual electrode positions observed in the expanded cuff, which are not
actually on a diametrical line but are shifted by approximately
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Fig. 11. Measurement result showing linear dependence on flow velocity and
symmetric responses with opposing magnetic fields.
Fig. 13. (a) Parasitic capacitances in the actual placement of the wireless EM
sensor. (b) Measured frequency shift with the emulated V .
B. The Wireless Option
Fig. 12. (a) Electrode positions in the cuff structure. (b) Measured and projected voltages as a function of flow velocity.
50 in total during the balloon inflation. This nonuniform expansion can lower the output voltage. Two hypothetical cases
are evaluated and plotted in Fig. 12(b): in one, each of the electrodes is equally offset by 25 as shown in Fig. 12(a) (Cal-2),
while in the other one of the electrodes is offset by the entire
50 (Cal-3). The latter appears to be a better match. A further
source of error is the loss associated with the nonideal profile
of flow velocity. The presence of electric lead wires, which are
connected to the electrodes and/or the cuff itself, can disturb
the flow. The sensitivity to local flow velocity varies across
the channel and rapidly increases close to the electrodes [42].
Thus, the disturbance near an electrode due to the boundary
irregularities can potentially require a correction factor to be
introduced into the theoretical estimate, which is otherwise idealized. The use of relatively large electrodes in the setup may
have also contributed to the reduced response.
For chronic intraluminal monitoring, the EM cuff may be
combined with a microelectronic interface chip that is telemetrically powered and controlled. A stentenna that is co-located
with it can potentially serve as its wireless link. However, there
could be cases in which the size of the chip prevents its use,
and a smaller device with reduced functionality would be preferred. One possible way to implement a passive readout of a
fully intraluminal EM flow sensor is by combining the inductance of the stentenna and the capacitance of a varactor to form
an LC tank. In this scheme, the stentenna and the varactor are
connected in series, and the ends of the series pair are terminated by the two electrodes of the EM cuff [see Fig. 13(a)]. The
generated between the electrodes
varactor is biased by the
to modulate the resonant frequency of the tank. This circuit is
interrogated wirelessly by an inductive coupling between the
tank and an external coil, as used for the pressure-based device
in Section II.
As a preliminary test, the junction capacitance of a diode
(1N3595, Fairchild Semiconductor Co., ME) was used as the
varactor. Since the threshold voltage of the diode is much higher
, the series-connected tank achieves
than a typical range of
high input impedance. The stentenna structure was fabricated
by patterning a 50- -thick stainless steel sheet by
and then electroplating it with copper to reduce the parasitic
resistance of the structure and, therefore, increase the quality
factor of the tank. The resistance of the stentenna, which was
originally 14 , was reduced to about 1/10 of the value with
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a 3- -thick copper coating. The structure was then coated
thick parylene-C, leaving the electrodes exposed.
with 1The diode was then mounted on the structure and sealed with
insulating epoxy for both electrical and mechanical protection.
The assembly was then mounted on a deflated angioplasty bal(proloon and deployed inside a test lumen. An emulated
vided by a power supply) was used to modulate the varactor,
while the frequency at which the resonant phase dip occurred
was monitored in an external coil. Fig. 13(b) shows the measured response of 2.0–3.3 KHz/mV that saturates as it reaches a
turn-on voltage of the diode, which is 680 mV in this case.
The challenge in this architecture lies in finding a varactor
with sensitivity that is high enough and parasitics that are low
enough that it can be substantially modulated by the voltages
built up on the cuff. In particular, it is important for any parasitic shunt capacitance and conductance in the varactor to be as
low as possible so as not to load the voltage source. The parasitic capacitances that are particularly detrimental include
(which is a lumped equivalent of the parasitic capacitance in the
(which is the parasitic capacitance in the
stentenna), and
discrete diode) as illustrated in Fig. 13(a).
In order to improve performance of the stentenna, one of the
priorities is to increase the inductance per unit length. With respect to the pressure-based measurement that was described in
Section II, it is noted that a single resonance peak was observed
even though two pressure sensors were available. To obtain a
differential measurement from these devices, it will be necessary to increase of the quality factor of the LC tanks. Increasing
the inductance and quality factor will also increase the distance
over which the signal can be measured, which is essential for
the practical use of this technology. The wireless EM device reported in Section III can also benefit from these changes. As
indicated in (2), to increase inductance, the number of turns in a
stentenna is one of the most important factors. In the test cases
described here, this is limited because of the spatial resolution of
the machining, which was performed by
using a wire-tip
[17] can achieve
as the cutting electrode. Batch mode
finer geometries because the cutting electrodes are lithographically fabricated on a microchip. This leads to more turns per unit
length of the stentenna, which leads to higher inductance. (Batch
also offers higher precision and about 100X immode
provement in throughput under the proper circumstances.) Another variable that affects the quality factor is parasitic resistance. The parasitic resistance of the stentennas can be reduced
by coating them with a low-resistivity metal such as gold or
copper as demonstrated in the wireless EM device. Stentennas
that offer higher electrical performance will also contribute to
achieve better coupling with external circuitry at the other end
of the wireless link.
Another task that lies ahead is to evaluate the biological impact of the proposed structures in the context of various applications. Vascular applications, for example, may face challenges
such as weakened arterial walls, clot formation, migration of
the structure under certain physical conditions, etc. Further, if
the structures must be coated with drug-eluting materials, the
consequent changes in mechanical and electrical properties also
must be investigated.
This effort has demonstrated, in a preliminary manner, the
possibility of using steel structures that are micromachined
from planar foil to monitor intraluminal pressure and flow. Of
the approaches investigated, the first integrates micromachined
capacitive pressure sensors with an inductive antenna stent.
This structure retains compatibility with standard stenting tools
and procedures, as demonstrated by its the deployment inside
a mock artery using an angioplasty balloon. It can be used to
monitor both pressure and flow of both insulating and conductive liquids, and is also suitable for air passages. However, the
flow velocity is not measured directly, and must be deduced
from the pressure. The second approach exploits an EM flow
detection mechanism that can be applied to only polarizable
liquids, including blood, and provides a more direct measurement of flow velocity that is less affected by the properties of
the liquid. However, it does not provide pressure information.
Its demonstration used a micromachined intraluminal cuff that
included a pair of electrodes that were mechanically coupled
but electrically isolated by dielectric links. It is anticipated that
(with the help
in the future, the use of high-resolution
of lithographically patterned cutting tools) will allow more
intricate patterns to be developed in such structures, allowing,
for example, stentennas with higher inductance. In addition,
the integration of this fabrication process with other planar
technologies will facilitate the integration of elements such as
pressure sensors and varactors.
The authors thank Dr. A. DeHennis for fabricating the capacitive pressure sensors and providing helpful advise regarding
testing and calibration procedures. They also thank Dr. S. Mutlu
for the parylene coating of samples.
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Kenichi Takahata (M’04) received the B.S. degree
in physics from Sophia University, Tokyo, Japan, in
1990 and the M.S. and Ph.D. degrees in electrical
engineering from the University of Michigan, Ann
Arbor, in 2004 and 2005, respectively.
In 1990, he joined Matsushita Research Institute
Tokyo, Inc. (Panasonic) and was with Matsushita
Electric Industrial Co., Japan, until 2001. At Matsushita, he was engaged in research and development
of micromechanics and microfabrication technologies including microelectrodischarge machining
(EDM) partially for the Japanese National Project “Micromachine Technology.” In 1997, he was appointed Researcher in the International Joint
Research Program supported by New Energy and Industrial Technology
Development Organization (NEDO) of Japan, which explored the compatibility between EDM and deep X-ray lithography (LIGA) processes at
the University of Wisconsin, Madison. From 1999 through 2001, he held a
Visiting Scientist position at the University of Wisconsin, Madison, where
he investigated EDM techniques that utilized lithographically fabricated
electrodes. His doctoral research at the University of Michigan involved batch
manufacturing technologies based on EDM and application to MEMS with a
focus on implantable devices and sensors. From 2005 to 2006, he was a Senior
Research Engineer at 3M Company, St. Paul, MN. Presently, he is an Assistant
Professor with the Department of Electrical and Computer Engineering,
University of British Columbia, Vancouver, Canada. He currently has 25
publications, six issued patents, and 20 pending patents in the United States
and Japan. His research interests are in MEMS realized by a combination of
silicon and nonsilicon-based manufacturing technologies.
Yogesh B. Gianchandani (S’83–M’85–SM’04) received the B.S., M.S, and Ph.D. degrees in electrical
engineering, with a focus on microelectronics and
He is presently a Professor with the Department
of Electrical Engineering and Computer Science and
holds a joint appointment with the Department of Mechanical Engineering, University of Michigan, Ann
Arbor. At the University of Michigan, he serves as
the Director of the College of Engineering Interdisciplinary Professional Degree Program in Integrated
Microsystems. Prior to this, he was with the Engineering and Computer Science
Department, University of Wisconsin, Madison. He has also held industry positions with Xerox Corporation, Microchip Technology, and other companies,
working in the area of integrated circuit design. His research interests include
all aspects of design, fabrication, and packaging of micromachined sensors and
actuators and their interface circuits. He has published about 150 papers in the
field of MEMS, and has about 25 patents issued or pending.
Prof. Gianchandani is the recipient of a National Science Foundation Career
Award. He serves on the editorial boards of IOP Journal of Micromechanics
and Microengineering and Journal of Semiconductor Technology and Science,
and served as a section editor for Sensors and Actuators for five years. He also
served on the steering and technical program committees for the IEEE/ASME
International Conference on Micro Electro Mechanical Systems (MEMS), and
as a General Co-Chair for this meeting in 2002.
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Kensall D. Wise (S’61–M’69–SM’83–F’86) received the B.S.E.E. degree with highest distinction
from Purdue University, West Lafayette, IN, in
1963 and the M.S. and Ph.D. degrees in electrical
engineering from Stanford University, Stanford, CA,
in 1964 and 1969, respectively.
From 1963 to 1965 (on leave 1965–1969) and from
1972 to 1974, he was a Member of Technical Staff
at Bell Telephone Laboratories, where his work was
concerned with the exploratory development of integrated electronics for use in telephone communications. From 1965 to 1972, he was a Research Assistant and then a Research
Associate and Lecturer with the Department of Electrical Engineering at Stanford University, working on the development of integrated circuit technology
and its application to solid-state sensors. In 1974, he joined the Department of
Electrical Engineering and Computer Science at the University of Michigan,
Ann Arbor, where he is now the J. Reid and Polly Anderson Professor of Manufacturing Technology and Director of the NSF Engineering Research Center for
Wireless Integrated MicroSystems. His present research interests focus on the
development of integrated microsystems for health care, process control, and
environmental monitoring.
Dr. Wise organized and served as the first chairman of the Technical Subcommittee on Solid-State Sensors of the IEEE Electron Devices Society (EDS). He
was General Chairman of the 1984 IEEE Solid-State Sensor Conference, Technical Program Chairman of the IEEE International Conference on Solid-State
Sensors and Actuators (1985), and IEEE-EDS National Lecturer (1986). He
served as General Chairman of the 1997 IEEE International Conference on
Solid-State Sensors and Actuators. He received the Paul Rappaport Award from
the EDS (1990), a Distinguished Faculty Achievement Award from the University of Michigan (1995), the Columbus Prize from the Christopher Columbus
Fellowship Foundation (1996), the SRC Aristotle Award (1997), and the 1999
IEEE Solid-State Circuits Field Award. In 2002, he was named the William
Gould Dow Distinguished University Professor at the University of Michigan.
He is a Fellow of the AIMBE and a member of the United States National
Academy of Engineering.
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