U.S. patent application number 15/102900 was filed with the patent office on 2016-10-20 for devices and methods for parameter measurement.
This patent application is currently assigned to THE BOARD OF REGENTS OF THE UNIVERSITY OF TEXAS SYSTEM. The applicant listed for this patent is The Board of Regents of the University of Texas System. Invention is credited to Mauli AGRAWAL, Steven BAILEY, Peter STARR.
Application Number | 20160302729 15/102900 |
Document ID | / |
Family ID | 53371800 |
Filed Date | 2016-10-20 |
United States Patent
Application |
20160302729 |
Kind Code |
A1 |
STARR; Peter ; et
al. |
October 20, 2016 |
DEVICES AND METHODS FOR PARAMETER MEASUREMENT
Abstract
A thin-film, diaphragm based device is disclosed which can be
used to perform an array of sensing and actuating operations where
a very thin profile is desired, such as in millimeter, micrometer,
or nanometer tight spaces.
Inventors: |
STARR; Peter; (San Antonio,
TX) ; BAILEY; Steven; (San Antonio, TX) ;
AGRAWAL; Mauli; (San Antonio, TX) |
|
Applicant: |
Name |
City |
State |
Country |
Type |
The Board of Regents of the University of Texas System |
Austin |
TX |
US |
|
|
Assignee: |
THE BOARD OF REGENTS OF THE
UNIVERSITY OF TEXAS SYSTEM
Austin
TX
|
Family ID: |
53371800 |
Appl. No.: |
15/102900 |
Filed: |
December 10, 2014 |
PCT Filed: |
December 10, 2014 |
PCT NO: |
PCT/US2014/069525 |
371 Date: |
June 9, 2016 |
Related U.S. Patent Documents
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Application
Number |
Filing Date |
Patent Number |
|
|
61914473 |
Dec 11, 2013 |
|
|
|
Current U.S.
Class: |
1/1 |
Current CPC
Class: |
A61B 5/0031 20130101;
A61B 5/686 20130101; A61B 5/0084 20130101; A61B 2562/16 20130101;
A61B 5/02158 20130101; A61B 2017/00345 20130101; A61B 2562/0204
20130101; A61B 2562/0247 20130101; G01L 9/0072 20130101; A61M 1/122
20140204; A61B 5/026 20130101; A61F 2/24 20130101; B29L 2031/753
20130101; A61B 2562/12 20130101; A61M 1/101 20130101; A61B 5/145
20130101; A61M 1/1037 20130101; A61B 2562/0233 20130101; A61B
2562/04 20130101; A61B 8/0891 20130101; A61B 5/0004 20130101; A61B
5/6862 20130101; A61M 1/125 20140204; A61B 5/032 20130101; A61B
5/6817 20130101; A61B 2562/028 20130101; A61M 1/1053 20130101; A61B
5/1038 20130101; A61M 1/106 20130101; A61B 5/1473 20130101; A61B
2562/0285 20130101; A61M 1/1086 20130101; B29C 65/48 20130101; A61B
5/0215 20130101; A61B 5/1036 20130101; A61B 5/6852 20130101; A61B
5/03 20130101; A61B 2562/164 20130101; A61M 1/12 20130101; B81B
2201/0264 20130101; A61B 7/023 20130101 |
International
Class: |
A61B 5/00 20060101
A61B005/00; A61B 5/103 20060101 A61B005/103; B29C 65/48 20060101
B29C065/48; A61B 5/1473 20060101 A61B005/1473; A61M 1/12 20060101
A61M001/12; A61F 2/24 20060101 A61F002/24; A61B 5/0215 20060101
A61B005/0215; A61B 7/02 20060101 A61B007/02 |
Claims
1. A device comprising: a substrate; and a diaphragm coupled to the
substrate, wherein the diaphragm is a thin film capacitive
transducer less than 1 mm thick.
2. The device of claim 1 wherein the thin film capacitive
transducer is between 10 .mu.m and 20 .mu.m thick.
3. The device of claim 1 wherein the diaphragm is coupled to the
substrate via an adhesive or other bonding method.
4. The device of claim 1 further comprising a chamber structure
between the diaphragm and the substrate.
5. The device of claim 4 wherein: the diaphragm is coupled to the
substrate via an adhesive; the chamber structure comprises a
bonding pad around the perimeter of the chamber structure; and the
chamber structure is positioned between the diaphragm and the
adhesive layer.
6. The device of claim 1 wherein the substrate is electrically
conductive.
7. The device of claim 1 wherein the substrate and diaphragm are
configured as a wireless resonant pressure sensor sized for
implantation in a human artery.
8. The device of claim 1 wherein the diaphragm is approximately 15
.mu.m thick.
9. The device of claim 1 wherein the substrate is approximately 50
.mu.m thick.
10. The device of claim 1 wherein the substrate is configured as an
antenna.
11. The device of claim 1 wherein the device is configured to
measure pressure with a linear sensitivity of approximately four
percent between 0 and 400 mm Hg.
12. The device of claim 1 wherein the substrate and the diaphragm
are biocompatible.
13. The device of claim 1 wherein the device is configured as a
pressure sensor.
14. The device of claim 1 wherein the device is configured as an
audio wave sensor.
15. The device of claim 1 wherein the device is configured as a
chemical sensor.
16. The device of claim 1 wherein the device is configured as a
biological sensor.
17. The device of claim 1 wherein the device is configured as an
optical sensor.
18. The device of claim 1 wherein the device is configured as a
pump.
19. The device of claim 1 wherein the device is configured as a
valve.
20. The device of claim 1 further comprising a first electrode
coupled to the diaphragm and a second electrode coupled to the
substrate.
21. A method of fabricating a thin film capacitive transducer, the
method comprising; providing a substrate; providing a diaphragm,
wherein the diaphragm is between 10 .mu.m and 20 .mu.m thick; and
coupling the diaphragm to the substrate.
22. The method of claim 21 wherein coupling the diaphragm to the
substrate comprises using adhesive to couple the diaphragm to the
substrate.
23. The method of claim 21 further comprising inserting a chamber
structure between the diaphragm and the substrate before coupling
the diaphragm to the substrate.
24. The method of claim 23 wherein the diaphragm and chamber
structure are constructed using photolithography.
Description
[0001] The application claims priority to U.S. Provisional Patent
Application No. 61/914,473 filed Dec. 11, 2013, which is
incorporated herein by reference in its entirety.
BACKGROUND INFORMATION
[0002] Most MEMS sensors are built onto silicon based wafers of
approximately 500 .mu.m thickness. While these thick substrates
confer stability during fabrication and over the long term, the
thickness limits applications in tight spaces, which includes many
biomedical and industrial conditions. The rigidity and
biocompatibility of silicon based sensors are additional limiting
factors. Overcoming these issues is particularly challenging for
diaphragm based sensors, due to the tight control required to build
three-dimensional cavities and diaphragms at such a small
scale.
[0003] The active region of many silicon based sensors is the
deflecting diaphragm near the surface of the sensor. Typically, the
active region ranges from the low-micron to sub-micron scale, which
is a small fraction of the overall sensor thickness. The sensor
profile can be significantly reduced if the inactive substrate is
replaced with a thinner substrate or if the active region is
integrated into the device package.
[0004] Passive ultrasonic sensors, methods and systems for their
use are described in U.S. Pat. No. 6,770,032. Specifically, passive
acoustic sensors having at least two flat parallel acoustically
reflecting surfaces. At least one reflecting surface is on a member
which is movable such that the distance between the reflecting
surfaces varies as a function of a physical variable to be
determined Preferably, the sensor is made such that the intensity
of a first portion of incident acoustic waves which is reflected
from one reflecting surface is equal or substantially similar to
the intensity of a second portion of the incident acoustic waves
which is reflected from the other reflecting surface. The first
portion and the second portion interfere to form a returning
acoustic signal having one or more maximally attenuated frequencies
which is correlated with the value of the physical variable. The
internal acoustic signal is received and processed to determine the
value of the physical variable from one or more of the maximal
attenuation frequencies. Methods and systems for using the passive
sensors are disclosed.
[0005] Existing systems such as those described above use an
ultrasound probe, with a limited transmitting/receiving bandwidth,
which permitted limited sensing of resonators, because most
feasible mechanical resonators have natural frequencies in the
audible or just above audible range under physiologically relevant
pressure.
[0006] A passive sensor system using ultrasonic energy is also
described in PCT Patent Publication WO 1995020769. In particular, a
passive sensor system (14) utilizing ultrasonic energy is
disclosed. The passive sensor system includes at least one
ultrasonically vibratable sensor (10) and an ultrasonic activation
and detection system (20, 22, 24, 25). The sensor (10) has at least
one vibration frequency which is a function of a physical variable
to be sensed. The ultrasonic activation and detection system (20,
22, 24, 25) excites the sensor and detects the vibration frequency
from which it determines a value of the physical variable. The
sensor includes (see FIG. 2-4) a housing, a membrane which is
attached to the housing and which is responsive to the physical
variable, a vibratable beam attached to the housing at one end and
a coupler, attached to the membrane and to a small portion of the
vibratable beam, which bends the vibratable beam in response to
movement of the membrane.
[0007] The ability to measure pressure locally can be used in the
analysis of certain conditions. Diabetics are prone to foot
ulceration, with a population prevalence of approximately 8% and a
lifetime risk of up to 25% (Margolis, Boulton). Loss of innervation
due to diabetic peripheral neuropathy induces muscle laxity and
associated skeleton deformities, as well as loss of sensation. This
increased risk of focal stress points and reduced ability to
accommodate to the initiating trauma greatly contribute to the
formation of ulcers, which can progress in severity to the point
where amputation is necessary. Critically, prevention and
management by proper monitoring of foot conditions could reduce
amputations by 50% (Driver).
[0008] Treatment of an ulcer is difficult after formation due to
repetitive damage and compromised healing in diabetics. Over a
third of the direct expenditure on diabetes in the US ($116
billion) is on ulcer treatment, with each treatment costing on
average $28,000. Prevention by careful monitoring of the condition
of the feet is considered to be the best approach and is thought to
potentially avert half of the amputations due to ulceration
(Driver).
[0009] Space constrictions limit conventional sensing devices in
many environments, such as in shoes or insoles. Wires, power
supplies, circuitry, and antennas in conventional approaches are
all too large and cumbersome to fit without disruption.
Electromagnetic resonance sensing offers a solution because these
simple wireless systems only require a coil and a capacitor to
operate. As such, they can be made small enough to wirelessly sense
otherwise inaccessible environments. They are interrogated
wirelessly by magnetic coupling. In some cases, the resonant system
can entirely replace the conventional radio link system; in other
cases, it can be used together with a radio link to extend the
sensing range. The mechanism of resonance sensing is not widely
used or known, probably because most sensing environments are
accessible via wired sensors. Presently, to synthesize this
mechanism of sensing with an application in foot pressure sensing
requires a breadth of knowledge in a numerous disparate fields,
including physics, mechanics, electrical engineering, and clinical
medicine.
[0010] Peripheral neuropathy contributes to the high prevalence of
foot ulceration in diabetics. Several systems, integrated into shoe
insoles and socks, are currently available for monitoring foot
pressures to prevent ulceration. However, these systems have
practical limitations and inconveniences for end user, such as
dangling wires or tenuous electronics.
[0011] Embodiments of the present disclosure offer a clean solution
through a resonant wireless system in a shoe insole. The sensing
insole is physically simple and durable, requires no on-site power
supply or circuitry, and can wirelessly transmit pressure signals
to a nearby device with radio link capability, such as a clip on
the outer shoe, an anklet, or a waist belt. To our knowledge, no
resonant wireless sensing system has been applied to measuring foot
pressures in the patent or scientific literatures. An embodiment
has been enabled with a thin film capacitive pressure transducer
which demonstrates functionality and excellent pressure
sensitivity.
SUMMARY
[0012] Described herein is a thin-film, diaphragm based device
which can be used to perform an array of sensing and actuating
operations anywhere where a very thin profile is desired, such as
in millimeter, micrometer, or nanometer tight spaces. The device
has a diaphragm and can operate by capacitive, resistive, and
resonant mechanisms. Due to its general structure, applications
include: mechanical sensing and actuation, chemical-biological
sensing, and optical sensing.
[0013] The device can be bonded to any substrate, allowing for
device integration. Additionally, the device can be constructed
from flexible materials, which allows for applications which
require flexibility, conformation to a nonflat or mobile surface,
or in three dimensional configurations. The device can also be
fabricated as an array of diaphragms to measure single factors, to
measure multiple factors simultaneously, or to measure surface maps
of factors. The fields of application are wide ranging, from
biomedical to industrial. The thin film sensor can be considered a
platform technology for low profile MEMS sensing due to its general
structure and utility.
[0014] The dimensions and materials of commercially available
pressure transducers limit their applications for intravascular and
implantable blood pressure sensing. Here, a high fidelity pressure
transducer is presented which is .about.10 um thick and can be
embedded into any surface, including cardiovascular catheters,
guide-wires, and stents. The transducer is micro-fabricated from
various polyimides, and is bonded onto 50 um thick 316 L stainless
steel foil for prototyping.
[0015] The static and dynamic characteristics of the transducer are
excellent. The transducer signal has high linearity (R2>0.99),
and resolution <<1 mmHg which is limited only by the system
noise. The operating frequency range is from 0 to >1 kHz, which
is well over the necessary limit for dynamic cardiovascular
applications, even in small animals with rapid heart rates.
Additionally, theoretical analysis indicates that both static and
dynamic performance of the transducer can be further improved with
optimization. Stability studies of the transducer in a pulsatile
flow environment with saline and serum show little drift in
transducer characteristics over a four week period.
[0016] Exemplary embodiments of the present disclosure relate to a
thin-film sensing or actuating device. In certain embodiments, the
device can be configured as a general sensor with broad ranging
applications, as described in more detail below.
[0017] Exemplary embodiments include a thin film sensor which can
be integrated onto any substrate, using methods that are compatible
with a range of materials and sensing mechanisms. One embodiment is
about 15 um, which has been bonded to a 50 um thick stainless steel
substrate.
[0018] Exemplary embodiments include a device comprising: a
substrate; and a diaphragm coupled to the substrate, wherein the
diaphragm is a thin film capacitive transducer less than 1 mm
thick. In particular embodiments, the thin film capacitive
transducer is between 10 .mu.m and 20 .mu.m thick. In certain
embodiments, the diaphragm is coupled to the substrate via an
adhesive or other bonding method. Particular embodiments further
comprise a chamber structure between the diaphragm and the
substrate.
[0019] In specific embodiments, the diaphragm is coupled to the
substrate via an adhesive; the chamber structure comprises a
bonding pad around the perimeter of the chamber structure; and the
chamber structure is positioned between the diaphragm and the
adhesive layer. In certain embodiments, the substrate is
electrically conductive. In particular embodiments, the substrate
and diaphragm are configured as a wireless resonant pressure sensor
sized for implantation in a human artery. In some embodiments, the
diaphragm is approximately 15 .mu.m thick, and in particular
embodiments, the substrate is approximately 50 .mu.m thick. In
particular embodiments, the substrate is configured as an antenna.
In specific embodiments, the device is configured to measure
pressure with a linear sensitivity of approximately four percent
between 0 and 400 mm Hg.
[0020] In certain embodiments, the substrate and the diaphragm are
biocompatible. In particular embodiments, the device is configured
as a pressure sensor. In some embodiments, the device is configured
as an audio wave sensor. In specific embodiments, the device is
configured as a chemical sensor. In some embodiments, the device is
configured as a biological sensor. In certain embodiments, the
device is configured as an optical sensor. In particular
embodiments, the device is configured as a pump. In some
embodiments, the device is configured as a valve. Particular
embodiments further comprise a first electrode coupled to the
diaphragm and a second electrode coupled to the substrate.
[0021] Exemplary embodiments also include a method of fabricating a
thin film capacitive transducer, the method comprising; providing a
substrate; providing a diaphragm, wherein the diaphragm is between
10 .mu.m and 20 .mu.m thick; and coupling the diaphragm to the
substrate. In certain embodiments, coupling the diaphragm to the
substrate comprises using adhesive to couple the diaphragm to the
substrate. In particular embodiments, the method further comprises
inserting a chamber structure between the diaphragm and the
substrate before coupling the diaphragm to the substrate. In
specific embodiments, the diaphragm and chamber structure are
constructed using photolithography.
[0022] Described is a thin-film, diaphragm based device which can
be used to perform an array of sensing and actuating operations
anywhere where a very thin profile is desired, such as in
millimeter, micrometer, or nanometer tight spaces. The device has a
diaphragm and can operate by capacitive, resistive, and resonant
mechanisms. Due to its general structure, applications include:
mechanical sensing and actuation, chemical-biological sensing, and
optical sensing.
[0023] The device can be bonded to any substrate, allowing for
device integration. Additionally, the device can be constructed
from flexible materials, which allows for applications which
require flexibility, conformation to a nonflat or mobile surface,
or in three dimensional configurations. The device can also be
fabricated as an array of diaphragms to measure single factors, to
measure multiple factors simultaneously, or to measure surface maps
of factors. The fields of application are wide ranging, from
biomedical to industrial. The thin film sensor can be considered a
platform technology for low profile MEMS sensing due to its general
structure and utility.
[0024] A mechanical resonator and system for acoustic wireless
interrogation of the resonator are also disclosed. In certain
embodiments, the resonator is micron-scale, with a resonance
frequency that is strongly dependent on external pressure. Methods
for interrogation of an implanted resonator include a skin piezo
device which sends an impulse to the resonator. Induced resonance
returns to the piezo a pressure wave at the pressure dependent
frequency of the resonator. High resonance frequencies, >1 kHz,
permit hundreds of pressure samples per second, which enables a
dense recreation of the blood pressure waveform. Additional factors
can be measured by the sensor, including temperature, local
gas--fluid environment, and local viscosity.
[0025] In one embodiment, for example, it can be used in an
implantable blood pressure sensing device where an ultra thin
profile is important to successful implementation. Cardiovascular
problems can be addressed by one embodiment of the sensor, wherein
a wireless implantable pressure sensor addresses the ubiquitous
need for blood pressure monitoring and control, given the many
conditions which hypertension negatively affects. Rates of heart
attack, stroke, heart failure, and cardiac arrhythmias are all
significantly increased at higher blood pressure levels. Exemplary
embodiments of the device could serve as a monitoring of blood
pressure for patients with a chronic cardiovascular condition to
ensure compliance with treatment and as a warning system for an
acute event. The potential demand is large as, according to the
AHA, cardiovascular disease accounts for nearly $500 billion in
cost, $75 billion of which is exclusive to hypertension and
sequelae.
[0026] In certain embodiments, the device may comprise a sensing
diaphragm that is approximately 15 .mu.m thick, which is bonded to
a stent material (e.g., 50 .mu.m stainless steel). Exemplary
embodiments can provide a linear sensitivity of about 4% over 400
mmHg Exemplary embodiments provide good dynamic fidelity, and have
been shown to accurately measure frequencies up to 10 kHz (and
possibly higher, as higher frequencies have not been tested). In
vitro studies are currently underway to characterize the robustness
of the sensor over time. In vitro studies with the sensor and
antenna in wireless mode are also planned in the future.
[0027] The device may be used in many applications, including for
example: arrays of force/pressure sensors could be used as a
tactile sensor, for diabetic patients with nerve damage and
foot/skin ulcers, or for robotics applications. In addition, two
pressure sensors spaced in a tube/artery can measure fluid flow
rates by the pressure drop. In certain embodiments, the device can
be used for mechanical sensing and actuation; bio-chemical sensing;
optical sensing; implantable intravascular pressure monitoring; and
cardio-vascular implants; and applications in implants for hearing
loss.
[0028] In a specific embodiment, the device may be configured as a
pressure sensor in an inductor-capacitor (LC) resonator for a
wireless implantable blood pressure sensor. Such a device relates
to a wireless implantable blood pressure sensor that reduces the
thickness of the transducing element for its implementation in
medium to small arteries, including the peripheral arteries. One
aspect of the device replaces the thick silicon wafer onto which
most pressure sensors are built with a very thin substrate or the
surface of an existing device or implant. This substitution of
platforms can save hundreds of micrometers of thickness. In
addition, using the shape-memory NiTi as an antenna allows for an
antenna that can be radially compressed and self-expand during a
percutaneous catheter delivery of the device.
[0029] A wireless implantable pressure sensor that addresses the
ubiquitous need for blood pressure monitoring and control and could
serve as a monitoring of blood pressure (BP) for patients with a
chronic cardiovascular condition to ensure compliance with
treatment and as a warning system for an acute event. However,
exemplary embodiments of the pressure transducer have applications
beyond an implantable sensor.
[0030] With a diaphragm thickness of approximately 15 .mu.m, the
device could be bonded to the tip of a catheter for intravascular
pressure sensing during operations. Biomedical applications beyond
cardiovascular include ocular pressure sensing, compartment
(syndrome) sensing, and integration into Lab-on-Chip (LOC) systems.
Industrial applications include locations with heavy space
constraints and/or need for physical flexibility, including
robotics and tire pressure systems.
[0031] A thin film diaphragm sensor is described herein with
multiple applications, including: mechanical sensing and actuation,
chemical-biological sensing, and optical sensing. Exemplary
embodiments of the device are approximately 10-20 .mu.m thick and
can be bonded to virtually any substrate. Exemplary embodiments may
comprise a deflecting diaphragm mechanism which can be used under a
variety of sensing and actuating mechanisms.
[0032] Certain exemplary embodiments of the device may be
configured as a pressure sensor or an acoustic sensor. In the
former, its thin profile can allow for implantable endovascular
blood pressure monitoring. When coupled with a self-expanding coil
composed of shape memory metal, it can be deployed conveniently
through percutaneous catheterization and interrogated with a small
coil near the skin surface. In an acoustic application, the device
can provide for high transduction fidelity through the audible
range.
[0033] Exemplary embodiments of the diaphragm device can be
configured as closed cells or channels or as open cells or
channels. The former configuration primarily serves in physical,
mechanical, and resonance sensing and some forms of actuation. The
latter configuration primarily serves in permittivity based sensing
for biological and chemical factors, and some forms of
actuation.
[0034] Closed cells are critical for establishing a pressure
difference between the device chambers and the outside, which then
allows for diaphragm deflection. Open cells are critical for
allowing biological or chemical factors for permeating the
inter-electrode space during permittivity based sensing.
Additionally, access to the inter-electrode space is necessary in
some forms of actuation, such as in pneumatic actuation of the
diaphragm. Modes of Operation
[0035] Exemplary embodiments of the disclosed diaphragm based
device can be used in capacitive mode (two overlapping electrodes),
in resistive mode (resistors on or within the diaphragm), in
resonance mode (diaphragm is driven into mechanical resonance), or
as a mechanical actuator. As described more fully below:
[0036] (1) Capacitive Mode
Capacitance=.epsilon. A/z [0037] .epsilon.--electrical permittivity
of material/space between the electrodes [0038] A--Area of
overlapping electrodes [0039] z--gap between electrodes
[0040] Mechanisms of Capacitive Sensing
[0041] A factor that modifies any of these three properties can be
sensed by a capacitive sensor.
[0042] (i) The most common sensing mechanism is by shifting the
electrode gap (.DELTA.z) by diaphragm deflection. In this mode,
force, pressure, and acoustic signals are typical measured which
are directly sensed. Numerous other factors can be indirectly
sensed by a deflecting diaphragm. For instance, flow can also be
measured with two pressure sensors in series. Additionally,
biochemical factors and analytes can be sensed if a swelling smart
material, for instance a receptor conjugated hydrogel, fills the
electrode gap. Also, optical sensing can be achieved in a Golay
cell configuration, described later.
[0043] (ii) Changing the permittivity (.DELTA..epsilon.) is an
additional sensing mechanism. Typically, a permeable material fills
the space between electrodes and absorbs the factor or analyte.
Absorption alters the permittivity and changes capacitance.
Humidity and pH are commonly sensed by this mechanism, but an
analyte specific material such as a receptor conjugated polymer
(e.g., hydrogel) can allow for specific biochemical analytes to be
sensed by this method. Additionally, a swelling hydrogel may
combine the effects of permittivity shifts and diaphragm
deflection.
[0044] (iii) Changing the area of overlapping electrodes (.DELTA.A)
is another sensing mechanism of capacitors with certain moving
parts. Shear forces and acceleration by comb-drives or other
arrangements are measured. However, since our capacitive sensor has
fixed borders around the diaphragm, this mechanism does not
apply.
[0045] (2) Resistive Mode
[0046] The most common pressure sensor is a deflecting diaphragm
with a bridge of resistive sensors, either thin metal films or
semiconductors. Deflection strains the diaphragm and its associated
resistors, which then modifies their resistance.
[0047] In this mode, applications would likely be limited to force
or pressure sensing, and their derivatives such as flow or acoustic
sensing.
[0048] (3) Resonance Mode
[0049] The basic mechanism is that the diaphragm is driven into
mechanical resonance and this resonance frequency is monitored. A
factor which modifies this resonance frequency can then be
detected.
[0050] A primary application of this sensing mechanism is for
biological or chemical sensing. In this mode, the exposed surface
of the diaphragm is conjugated with a receptor for the measured
factor. Depending on its configuration, the diaphragm can be driven
into resonance by various means, including electrostatically (if it
contains parallel electrodes), thermally, or an applied pressure
via an acoustic signal or a pop-test (a step drop in pressure,
which induces resonance in the diaphragm). The resonance frequency
can be monitored electrically by various methods which depend on
whether the diaphragm device is acting as a variable capacitor or a
variable resistor. When the biological or chemical factor binds the
receptor, it mass loads the diaphragm and thereby shifts its
resonance frequency.
[0051] (4) Mechanical Actuator
[0052] Whereas the sensing diaphragm moves in response to a signal,
the diaphragm can alternatively be driven into movement to achieve
a mechanical goal. Methods to induce mechanical actuation include
pneumatic, electrostatic, thermal, among others. Most applications
of a thin film mechanical actuator will likely lie in microfluidics
devices, where the actuator can serve as a valve, a pump or other
pressurizing device.
[0053] Sensing Types and Possible End Applications
[0054] Exemplary embodiments of the disclosed diaphragm based
device can be used to achieve mechanical and physical sensing,
mechanical actuation, biological and chemical sensing, and optical
sensing, among others.
[0055] (1) Mechanical and Physical Sensing
[0056] The diaphragm device can operate as a force sensor under
numerous configurations and conditions where a thin profile or
flexibility is desired.
[0057] (i) Biomedical applications of force or pressure sensors
include cardiovascular blood pressure or flow sensing (e.g.,
hypertension, heart failure), ocular pressure sensing (e.g.,
glaucoma), pulmonary pressure sensing (e.g., chronic obstructive
pulmonary disease), pleural cavity pressure sensing (e.g.,
pneumothorax), urinary pressure sensing (e.g., incontinence),
gastrointestinal pressure sensing (e.g., incontinence), peritoneal
cavity pressure sensing (e.g., ascites), cerebro-spinal fluid
pressure sensing (e.g., hydrocephalus), muscular pressure sensing
(e.g., compartment syndrome), orthopedic pressure sensing (e.g.,
joint, disc, and/or implant pressures), podiatric pressure sensing
(e.g., for diabetic ulcers) among others.
[0058] One particular use of high value would be on a catheter tip
for intravascular blood pressure sensing or for urological sensing.
Currently, in the hospital wards, fluid filled catheters transmit
pressures from inside the body to an external pressure sensor. This
arrangement has significant sensitivity and drift errors, in
addition to artifacts such as the catheter whip effect, which
reports artificially high spikes in pressure when the fluid filled
catheter moves. Silicon microsensors do exist, but are very
expensive (>$1 k) and sterilization for re-use between patients
is not common for safety reasons. An inexpensive, thin, flexible
sensor which could be positioned at the tip of a disposable
catheter could considerably improve pressure sensing accuracy.
Additionally, a very thin profile would allow for measuring
pressures in tighter spaces than are currently possible.
[0059] An additional application of high value for a thin pressure
sensor is in implantable blood pressure sensing devices. Currently,
silicon based microsensors are built onto silicon or silica chips
which are at least several hundred microns in thickness. This
thickness precludes applications in all but the largest arteries,
since most medium and small arteries, such as the coronaries and
peripheral arteries, are <4 mm in internal diameter. An
implantable blood pressure sensing device has value in direct,
continuous, and chronic monitoring of hypertension and heart
failure, and can additionally serve as warning system for acute
cardiovascular events. Such a device could be constructed as an
inductor-capacitor (LC) system, with the thin pressure transducer
in capacitive mode. It could also be constructed alternatively,
where the transducer operates in either resistive or capacitive
mode.
[0060] (ii) Capacitive microphones are very common for transducing
audio signals. One biomedical application of a very thin audio
transducer is in an unobtrustive hearing device, such as an inner
ear implant, a cochlear implant, or hearing aid. The thinness is of
particular relevance, as the tympanic membrane is .about.50 um
thick. Additionally, the sensing range is not necessarily limited
to the audible range, however, and applications may include the
sub-audible and ultrasound ranges.
[0061] (iii) Industrial applications for force and pressure sensing
include automotive (e.g., tire pressure sensing, force sensors for
monitoring shock, misalignment), machines and robotics (e.g.,
monitoring shock, misalignment), among others.
[0062] (iv) Robotics applications include artificial skin for
tactile sensing. An array of diaphragms would allow for sensing a
two dimensional surface map. Such an artificial skin could be used
in a sensing skin for artificial intelligence robotics or in a
prosthetic for sensory loss in humans.
[0063] (2) Mechanical Actuation
[0064] Most applications of a thin film mechanical actuator will
likely lie in microfluidics devices, where the actuator can serve
as a valve, a pump or other pressurizing device. As a valve, one or
multiple diaphragms can situated as walls of a micro-channel. The
diaphragm can be driven outward or inward for either by an
electrostatic signal across the two electrodes or by a pneumatic
signal from within the inter-electrode space. The valve state will
be closed when the diaphragm is driven out and contacts the
opposing wall, thereby occluding the channel. The valve state will
be open when the diaphragm is driven in.
[0065] As a pump, the diaphragm can be driven to induce pressure to
drive flow in an adjacent chamber or channel. Upstream the pump can
be a one-way valve which blocks backwards flow, such that the pump
only drives forward flow.
[0066] The diaphragm device could also operate as a miniature
capacitive speaker, either in the sub-audible, audible, or
ultrasound range. This mode of operation, the diaphragm would
likely be driven electrostatically. As with the audio transducer
operating in an unobtrusive hearing device, a miniature speaker
could also be used in such a device for amplification of the audio
signal.
[0067] (3) Biological and Chemical Sensing
[0068] Biological and chemical sensing in tight spaces has
biomedical applications, among others in high technology. Possible
transduction mechanisms include resonance or capacitive modes. An
array of diaphragms could allow for monitoring of multiple
markers.
[0069] This transducer could be used in an implantable device for
monitoring biomarkers, for monitoring the status of either chronic
disease or cancer. If configured to give surface map data, the
device could be used as an artificial tasting or smelling device
(smart tongue or smart nose).
[0070] (4) Optical Sensing
[0071] Indirect optical sensing can be achieved in a Golay cell
configuration, whereby an air chamber with an optical filter sits
atop the deflecting diaphragm. The optical signal enters the top
chamber, changes its temperature, which induces expansion or
contraction of the chamber volume, and thereby changes the applied
pressure to the diaphragm of the capacitive sensor below.
[0072] If the sensor is configured as an array of frequency
specific optical transducers, an optical camera can be achieved for
imaging applications. A specific biomedical application of such an
optical camera includes a retinal implant for restoring vision. The
thin, flexible nature of the sensor confers a particular advantage
for conforming to the curved topography of the eye.
[0073] In the following, the term "coupled" is defined as
connected, although not necessarily directly, and not necessarily
mechanically.
[0074] The use of the word "a" or "an" when used in conjunction
with the term "comprising" in the claims and/or the specification
may mean "one," but it is also consistent with the meaning of "one
or more" or "at least one." The terms "about", "approximately" or
"substantially" means, in general, the stated value plus or minus
5%. The use of the term "or" in the claims is used to mean "and/or"
unless explicitly indicated to refer to alternatives only or the
alternative are mutually exclusive, although the disclosure
supports a definition that refers to only alternatives and
"and/or."
[0075] The terms "comprise" (and any form of comprise, such as
"comprises" and "comprising"), "have" (and any form of have, such
as "has" and "having"), "include" (and any form of include, such as
"includes" and "including") and "contain" (and any form of contain,
such as "contains" and "containing") are open-ended linking verbs.
As a result, a method or device that "comprises," "has," "includes"
or "contains" one or more steps or elements, possesses those one or
more steps or elements, but is not limited to possessing only those
one or more elements. Likewise, a step of a method or an element of
a device that "comprises," "has," "includes" or "contains" one or
more features, possesses those one or more features, but is not
limited to possessing only those one or more features. Furthermore,
a device or structure that is configured in a certain way is
configured in at least that way, but may also be configured in ways
that are not listed.
[0076] Other objects, features and advantages of the present
invention will become apparent from the following detailed
description. It should be understood, however, that the detailed
description and the specific examples, while indicating specific
embodiments of the invention, are given by way of illustration
only, since various changes and modifications within the spirit and
scope of the invention will be apparent to those skilled in the art
from this detailed description.
BRIEF DESCRIPTION OF THE DRAWINGS
[0077] The following drawings form part of the present
specification and are included to further demonstrate certain
aspects of the present disclosure. The invention may be better
understood by reference to one of these drawings in combination
with the detailed description of specific embodiments presented
herein.
[0078] The patent or application file contains at least one drawing
executed in color. Copies of this patent or patent application
publication with color drawing(s) will be provided by the Office
upon request and payment of the necessary fee.
[0079] FIG. 1 shows an exploded view of one embodiment of a device
according to the present disclosure.
[0080] FIG. 2 shows a section view of the embodiment of FIG. 1.
[0081] FIG. 3 shows a graph of capacitance versus pressure for one
embodiment of a device according to the present disclosure.
[0082] FIGS. 4-9 illustrate properties of the embodiment of FIG. 3
as measured over period of several days.
[0083] FIGS. 10-11 illustrate measurements of the embodiment of
FIG. 3 of dynamic signals from inside a flow loop with pulsatile
pressure.
[0084] FIG. 12 shows a schematic of one embodiment of a device
configured as an audio sensor.
[0085] FIG. 13 shows data recorded with the embodiment of FIG.
16.
[0086] FIGS. 14-25 show circuits and data for a specific embodiment
for insole pressure measurement.
[0087] FIGS. 26-28 illustrate data for exemplary embodiments of
four sensors according to the present disclosure over one month in
saline under pulsatile pressure.
[0088] FIGS. 29-32 contain regressions and drift of parameters over
the one month period for the data illustrated in FIGS. 26-28.
[0089] FIGS. 33-39 contain data from one sensor which addresses the
source of drift in the parameters for the data illustrated in FIGS.
29-32.
[0090] FIG. 40 illustrates a schematic of an exemplary embodiment
of a resonator according to the present disclosure.
[0091] FIG. 41 illustrates resonance frequency signals at different
pressures for exemplary embodiments of devices according to the
present disclosure.
[0092] FIGS. 42-43 illustrate data showing the pressure dependence
of diaphragm resonance frequency.
[0093] FIGS. 44-45 illustrate schematics of exemplary embodiments
of resonator devices anchored to a structure according to the
present disclosure.
[0094] FIG. 46 illustrates a schematic of acoustic interrogation of
an exemplary embodiment of a mechanical resonator according to the
present disclosure.
[0095] FIG. 47 illustrates schematics for wireless sensing
modalities for exemplary embodiments of implantable sensors
according to the present disclosure.
[0096] FIG. 48 illustrates a coordinate system and a schematic of
an exemplary embodiment of an analytical model according to the
present disclosure.
[0097] FIGS. 49-51 illustrate an experimental setup used to obtain
results previously shown in FIGS. 41-43.
[0098] FIGS. 52-53 illustrate frequency versus pressure data in the
audible range obtained from exemplary embodiments of resonators
according to the present disclosure.
[0099] FIGS. 54-55 illustrate data showing the penetration of
audible acoustic waves in soft tissue.
[0100] FIG. 56 illustrates the level of acoustic energy that can be
delivered to a resonator for different materials according to the
present disclosure.
[0101] FIG. 57 illustrates the reflected pressure ratio and
reflected power ratio for soft tissue in combination for different
materials.
[0102] FIGS. 58-60 illustrate a schematic of an experimental set up
for ultrasonic measurements and data obtained from the
experiment.
DETAILED DESCRIPTION OF ILLUSTRATIVE EMBODIMENTS
[0103] Referring initially to FIGS. 1-2, an exemplary embodiment of
a device 100 configured as a thin film sensor comprises a diaphragm
110, a chamber structure 120, an adhesive 130 and a substrate 140.
In the embodiment shown, diaphragm 110 is configured as a thin film
diaphragm transducer between 10 .mu.m and 20 .mu.m thick and is
bonded to substrate 140 via adhesive 130. In particular
embodiments, diaphragm 110 is approximately 15 .mu.m thick and
substrate 140 is approximately 50 .mu.m thick. In the present
disclosure, the thickness of a material is measured across the
primary plane of the material (i.e. the minimum dimension for a
given layer of material, as would be measured in a vertical
direction in the configuration shown in FIG. 2).
[0104] In the illustrated embodiment, chamber structure 120
comprises a bonding pad 125 around its perimeter and chamber
structure 120 is positioned between diaphragm 110 and substrate
140. In exemplary embodiments of device 100, substrate 140 can be
electrically conductive, and in certain embodiments can be
configured as an antenna.
[0105] Exemplary embodiments of device 100 may be fabricated by
constructing a thin sensing film, which comprises of an array of
diaphragms 110 enclosed by bonding pads 125. In certain
embodiments, to construct the sensing film, multiple layers of
photolithography with various polyimides can be performed on a
carrier substrate. The diaphragm can be defined in one step, the
chamber walls can be defined in a second step, and a thin adhesive
film applied in a third step. The sensing film can then be released
from the carrier.
[0106] In exemplary fabrication techniques, the sensing film can
then be bonded to the substrate of choice. In certain embodiments,
the thin adhesive can be deposited onto a conductive substrate. If
the substrate is not inherently conductive, a thin conductive film
may be deposited to provide a bottom electrode of the diaphragm
sensor. The sensing film can then be bonded to the substrate under
pressure and temperature.
[0107] In certain embodiments, the final fabrication step is to
sputter an electrode and bond lead wires. For example, a thin
conductive film can be deposited on top of the sensing film to
define the top electrode of the diaphragm sensor. Lead wires can
then be bonded onto the top and bottom electrodes.
[0108] In certain embodiments of the sensor, the sensing film (e.g.
diaphragm 110) is 10-15 .mu.m thick and substrate 140 is 50 .mu.m
thick stainless steel. In certain embodiments, diaphragms 110 form
a sensing film that is 3 mm.times.10 mm, but it can be of arbitrary
size to suit the application.
[0109] In certain embodiments, substrate 140 may be formed by
polymers processing techniques. Other microfabrication techniques
could produce a similarly-structured device composed of other
materials, including traditional microfabrication ceramics such as
silicon, silica, quartz, silicon nitrides, other nitrides, other
oxides, and other insulating or semiconducting materials.
[0110] During operation of device 100, deflection of diaphragm 110
toward and away from substrate 140 can be measured by changes in
electrical properties and correlated to environmental conditions or
parameters affecting device 100. For example, in certain
embodiments, the capacitance of device 100 (measured between
diaphragm 110 and substrate 140) can be correlated to pressure.
Referring now to FIG. 3, one example illustrates a substantially
linear relationship between the measured capacitance (in pF) versus
the pressure on diaphragm 110 (measured in mmHg). FIGS. 4-9
illustrate other properties of the embodiment of FIG. 3 as measured
over period of several days. FIGS. 10 and 11 illustrate
measurements of the embodiment of FIG. 3 of dynamic signals from
inside a flow loop with pulsatile pressure. FIG. 10 illustrates
waveforms from device 100 and a reference sensor. As illustrated,
the average difference is approximately 1 mm Hg.
[0111] Referring now to FIGS. 12 and 13, a schematic of device 100
(and resulting data) are shown for an embodiment configured as an
audio sensor. In this embodiment, device 100 senses an audio wave
150, which causes deflection of diaphragm 110 (not labeled in FIG.
6 for purposes of clarity; see FIGS. 1 and 2 for view depicting
diaphragm 110). Diaphragm 110 deflections cause a change in the
measured capacitance/voltage across device 100, which can be viewed
as a waveform on display 160.
[0112] In the graph shown in FIG. 13, audio frequencies were
recorded with high fidelity up to 10 kHz, indicating certain
embodiments of device 100 may be suitable for use for hearing aid
implants. In addition, the ability to record frequencies up to 10
kHz also indicate the potential utility of device 100 in
cardiovascular applications due to the ability to faithfully record
high frequency information in the pressure waveform.
[0113] Device 100 can be used in many different applications. For
example, device 100 can be configured for use as a sensor,
including a pressure, acoustic, force or flow sensor. Device 100
may also be configured as a mechanical actuating device, including
for example an electrostatically (or pneumatically)-driven membrane
that can be used as a pump or valve in microfluidics applications.
For example, in a valve configuration, diaphragm 110 can be
deflected outward (e.g. away from substrate 140) to occlude flow
and toward substrate 140 to allow flow to pass over diaphragm
110.
[0114] In still other embodiments, device 100 can be configured a
capacitive microphone, including for example configuration a
hearing aid.
[0115] In certain embodiments, device 100 can be configured as a
chemical or biological sensor. For example, chamber structure 120
can be configured as a polymer or hydrogel with selective
absorption that can swell and deflect diaphragm 110 in the presence
of certain analytes.
[0116] In particular embodiments, device 100 may also be used for
detecting chemical or biological analytes by mass loading of the
sensing diaphragm, which changes its resonance frequency. The
sensing diaphragm can have analyte receptors bound to its surface
and the resonance frequency of the sensing diaphragm can be
monitored by actuating device 100 electrostatically or thermally.
Detection of the analyte occurs by recording the shift in resonant
frequency of the diaphragm.
[0117] In specific embodiments, device 100 may be configured for
indirect sensing by principles similar to those used in a Golay
cell. For example, chamber structure 120 may be filled with a gas
that expands with increased temperature and causes deflection of
diaphragm 110. In particular embodiments, diaphragm 110 may be
coated with a bandpass filter to provide for specific detection of
light wavelengths or color. Such configurations could be used in
imaging or retinal implant applications.
[0118] In certain embodiments, device 100 can be configured as a
thin-film pressure sensor in an inductor-capacitor (LC) resonator
for a wireless implantable blood pressure sensor. In particular
embodiments, device 100 can operate by capacitive, resistive, and
resonant mechanisms. In exemplary embodiments, device 100 can sense
a broad range of factors, individually and multiple simultaneously.
Device 100 can be configured as an electrical inductor-capacitor
(LC) resonator that measures pressure by a thin film capacitive
transducer that resonates with a stent-like antenna.
[0119] In exemplary embodiments, the thin active region of the
sensor is decoupled from a thick inactive substrate. Certain
embodiments can incorporate the use of a shape-memory NiTi as an
antenna for percutaneous catheter delivery of the device. In
certain embodiments, movements in local pressure change the
transducer capacitance and thus shift the resonance frequency. In
particular embodiments, the resonance frequency can be monitored
externally by magnetic coupling to determine intravascular
pressure.
[0120] In specific embodiments, the sensor can be bonded to a thin
metallic substrate and coupled to a flexible NiTi stent-antenna
(inductor), and the diaphragm sensor and inductive antenna form an
electrical inductor-capacitor (LC) resonator.
[0121] In certain exemplary embodiments, device 100 has a thin
profile, is wireless, biocompatible, implantable, and allows for
intravascular implantation for blood pressure sensing. In
particular embodiments, device 100 can be fabricated with
biocompatible materials, is flexible and due to thin profile allows
for 3-D conformations of sensor in vivo, allows for implementation
in medium to small arteries, including the peripheral arteries.
[0122] In particular embodiments, device 100 can be bonded to
virtually any substrate, and be integrated or embedded into various
devices. The thin and flexible profile of device 100 is suitable
for implantation into constrained spaces which were previously
inaccessible for sensors.
[0123] The replacement of a thick silicon wafer onto which most
pressure sensors are built with a very thin substrate (or the
surface of an existing device or implant, including e.g. a stent)
can save hundreds of micrometers of thickness which can be critical
in particular applications. For example, one embodiment enables the
development of a wireless resonant pressure sensor which is
suitable for implantation in a large, medium, or small sized
artery. As described in the literature reviews on endovascular
blood pressure sensing devices and on pressure transducers,
transducer size has been a limiting factor in the development of
small implantable devices.
[0124] As described previously, exemplary embodiments of the
present disclosure substitute the platform for the sensing
diaphragm to reduce sensor thickness. Commercially available
pressure sensors use silicon wafer as substrates with a thickness
of about 500 .mu.m, most of which can be eliminated by integrating
the sensing element onto a robust surface of the device.
Test Data
[0125] FIGS. 26-28 provide raw data on four sensors over one month
in saline under pulsatile pressure. The data includes all tracked
parameters, and the sensors had an initial two week immersion
period in saline to allow parameter values to settle. Those values
were then measured two times per week.
[0126] FIGS. 29-32 contain regressions and drift of parameters over
the one month period. A graph at the end shows average drift in
each parameter.
[0127] FIGS. 33-39 contain data from one sensor which addresses the
source of drift in the parameters. Pressure was increased to 400
mmHg and pressure sensitivity curves were recorded; this was
repeated for ten consecutive cycles. Some drift in sensor
parameters are noted (for instance, 0.15% increase in baseline
capacitance). The sensor was left alone for a twelve hour break,
and then ten more cycles were performed. For almost all of the
parameters, after the twelve hour break, the parameter value
returned to the original value from day one, indicating that the
drift in parameters was not permanent (e.g., a hysteresis effect
which can be addressed during development and commercial
design).
[0128] Design and fabrication of exemplary embodiments requires
detailed knowledge and synthesis of multiple fields including
microelectronics, microfabrication, cardiovascular medicine, and
biomaterials. Additionally, silicon wafers are the epicenter of the
microelectronics and microfabrication fields; departing from this
fabrication orthodoxy is difficult.
[0129] Embodiments of the current invention include a class of
resonant sensors which can be used in a shoe insole for monitoring
foot pressures. The general sensor is a resistor-inductor-capacitor
(RLC) resonant circuit, which allows for either capacitive sensing
or resistive sensing. FIG. 14 shows circuit schematics of these two
possible configurations. In both cases, an external device with a
small coil and a radio link (e.g., Bluetooth), such as a clip on
the outer shoe, an anklet, or a waist belt, can interrogate the
sensor and transmit the pressure signals to a smart device,
computer, or wireless network.
[0130] In the capacitive design, a planar conductive coil is
electrically connected to a capacitive pressure transducer to form
an RLC tank, which is then embedded into an insole. The resonance
frequency of the tank depends on the applied pressure. The sensor
can be interrogated by an external coil which sweeps across a
specified frequency range to monitor shifts in the resonance
frequency.
[0131] In the resistive design, a planar conductive coil is
electrically connected to a capacitor and a resistive transducer to
form an RLC tank, which is then embedded into an insole. In this
case, the resonance frequency of the tank is fixed, but the quality
of resonance (quality factor Q) depends on the applied pressure.
The sensor can be interrogated by an external coil at a fixed
frequency by monitoring the strength of the magnetically coupled
signal.
[0132] The capacitive design of an RLC sensor has been enabled.
Referring back now to FIG. 15, a prototype insole is shown with a
thin film capacitive transducer and an embedded 2-turn coil. The
prototype has a strong and linear pressure sensitivity (40 kHz/PSI;
R2=0.993). Further experimental data are attached in an appendix. A
video of the sensor in operation will be sent electronically. The
particular type of capacitive transducer is non-essential to the
invention. A variety of thin transducers could easily be used, from
capacitive microsensors to custom capacitive sensors made from a
sandwich of thin metal foil with a compressible dielectric in the
middle.
[0133] FIGS. 17-19 provide pressure data for various embodiments,
while FIG. 20 provides a schematic showing different features of
existing systems and an embodiment of the present disclosure using
a magnetic coupling and a radio link.
[0134] FIGS. 21-25 provide data on a wireless insole reading range
according to exemplary embodiments of the present disclosure. For
all sensors, data was acquired with a non-optimized sensor and a
non-optimized interrogation system. The impedance analyzer operated
at 0.5V across the interrogating coil (instrument limit). Industry
RFID interrogators frequently use >10V to increase sensitivity,
and frequently have interrogation distances of >1 m for resonant
tags approximately 1 cm.
[0135] It is understood that the above-described devices and
methods are merely non-limiting examples of embodiments of the
devices and methods disclosed herein.
[0136] Exemplary embodiments of the present disclosure include
resonators that operate in the audible acoustic range. Existing
systems typically stipulate stimulation in the ultrasound
range.
[0137] Bandwidth of the acoustic transmitter and/or receiver in
exemplary embodiments of the present disclosure is much lower than
standard ultrasound crystals. In certain embodiments, a unique
probe may be developed for this application in the 1-20 kHz
range.
[0138] Mechanical resonators are most sensitive to gauge pressure,
and only to the first several hundred mmHg, after which sensitivity
drops considerably. Therefore, it is not possible simply to use or
test any commercially available pressure sensor with a
micromachined diaphragm, which have chambers underneath which are
frequently hermetically sealed under vacuum. In the case of vacuum
sealed, commercially available pressure sensors, the gauge pressure
across the diaphragm at the physiological range is >800 mmHg,
which offers negligible pressure sensitivity if used as a
mechanical resonator.
[0139] An exemplary embodiment of a prototype resonator is square
polyimide diaphragm (500 um long, 5 um thick) over a closed air
chamber, as shown in FIG. 40. In this embodiment, a square
diaphragm made of polyimide over an air chamber is bonded to
stainless steel substrate.
[0140] Modification of the standard equation for determining the
resonance frequency of such a diaphragm (Roark) yields the
following expression for resonance frequency with a strong pressure
dependence
F ( p ) = 36 2 .pi. Dg a 4 ( w + p ) ##EQU00001##
where
D = Et 3 12 ( 1 - v 2 ) ##EQU00002##
is the flexural rigidity of the diaphragm, v is the poisson ratio
of the diaphragm material, E is the elastic modulus of the
diaphragm material, t is the diaphragm thickness, a is the square
diaphragm length, g is the gravitational constant, w is the weight
of the diaphragm per unit area, and p is externally applied
pressure (gauge pressure across the diaphragm).
[0141] Experimental testing shows good agreement with the theory.
An impulse test was applied to the prototype diaphragms at various
pressures to induce resonance. FIG. 41 clearly shows the strong
pressure dependence of the resonance on local pressure. FIGS. 42
and 43 show that the experimental data matches the theory well.
Response of the mechanical resonator to an impulse response is
shown at different pressures. Increasing pressure reduces the
resonance frequency. There is a good match between experimental vs
predicted resonance frequency at various pressures, for a square
polyimide diaphragm 500 um long and 5 um thick.
[0142] Further theoretical analysis shows that ceramic resonators
should given even better pressure responses, due to their rigidity.
Additionally, the outstanding mechanical stability of ceramics,
particularly monocrystalline ceramics of silicon and SiO.sub.2
(quartz), should lend excellent robustness and long term sensing
stability.
[0143] FIGS. 44 and 45 illustrate variations of a conceived ceramic
resonator, anchored to a stent or stent-like structure. In FIG. 44,
the resonator is bonded to the stent surface, while in FIG. 45 the
resonator is embedded into the stent.
[0144] Once percutaneously implanted, the resonator can be
interrogated wirelessly by an acoustic impulse test. FIG. 46
illustrates how a piezo device at the skin surface can send a pulse
to the resonator, induce vibration, and read the frequency of the
vibration. In section 1 of FIG. 46, the piezo sends an impulse,
either a square wave or a sine wave near the resonance frequency of
the resonator. In section 2, the impulse stimulates vibration of
the resonator, which produces a pressure wave with an oscillating
decay at its resonance frequency. The piezo switches to listen
mode, or a second receiving piezo is used, to record the resonator
pressure wave.
[0145] If the resonance frequency of the resonator is sufficiently
high (>1 kHz), >>100 samples of blood pressure samples can
be taken during the pressure wave cycle. This should allow for a
dense recreation of the blood pressure waveform.
[0146] In the past several decades, many wireless sensing platforms
have been developed which utilize a radio link to transmit the
sensed data. Currently, with small Bluetooth-like radio links and
smart devices, these platforms are still in full force. There is a
miniaturization limit, however, due the numerous components such as
power sources, circuitry, and antennas. Resonance based systems
offer an alternative for wireless sensing, because resonators are
typically very simple structures, can be made small, and
efficiently receive and transmit energy within a certain frequency
range. FIG. 47 compares these wireless sensing modalities for
implantable sensors, and optical methods could also be included in
this comparison. In wireless modalities for implantable sensors,
conventional wireless systems are bulky, with many components.
Resonant systems have fewer components, no power requirements, and
can be interrogated magnetically or acoustically
[0147] While well known in the physics, mechanics, and electrical
literatures for over a century, resonance based sensing systems
have become more intensively investigated since the 1990s, with a
particular focus on electrical resonators. These electrical
resonators require only a capacitive sensor and a coil to operate,
and can be interrogated magnetically. Mechanical resonators,
however, have not been intensively investigated for stand-alone
sensing purposes.
[0148] The scientific literature is full of discussion of
electrical and mechanical resonance and resonators. The engineering
literature has several well recognized instances, the most
prominent one being the class of resonant pressure sensors in
silicon microsensors. Incidentally, these resonant pressure sensors
are known to have sensitivity and stability at least an order of
magnitude great than piezoresistive and capacitive sensors. In this
class, a micro-beam lies on a deflecting diaphragm and is induced
into resonance. Pressure deflects the diaphragm and changes the
strain on the beam, whose resonance frequency then shifts. This
shift is monitored by piezoresistors on the beam, which are then
processed by circuitry on or near the transducer chip. An important
aspect is that most declared "resonant sensors" operate similarly
to this class of sensors and are not stand-alone, passive resonant
sensors which can be wirelessly interrogated.
[0149] Significant intellectual property exists on the class of
stand-alone, passive mechanical resonators. Included is an appendix
table with examples of differences between the disclosed invention
and the relevant patents. The significant point is that the
inventions are largely undeveloped and, without a known exception,
utilize a very different acoustic frequency range. The work in
patents is done predominantly in the medical ultrasound range
(MHz), whereas the disclosed invention here operates in the audible
range (<20 kHz). Additionally, and related to this distinction,
the method of interrogation of the patented inventions is
frequently different from that of this disclosed invention.
[0150] Exemplary embodiments of the present disclosure provide
numerous non-obvious advantages over existing systems. For example,
the analytic solutions for resonance frequency of diaphragms and
beams do not contain explicit pressure terms, and thus the pressure
dependence is not obvious. Minor modifications of the formulas
readily yield pressure dependence, but the insight to make them
must first be had. In addition, the mechanism of sensing is
fundamentally different from that of most silicon-based resonant
pressure sensors. Most silicon-based resonant pressure sensors
focus on inducing a pressure dependent strain on a resonating beam.
This is typically done by deflecting the mechanical base on which
the beam lies, or by deflecting another mechanical member onto the
beam. IE, the resonance frequency of the sensing element is not
directly shifted by local pressure. In our case, resonance
frequency of our disclosed invention is directly shifted by local
pressure.
[0151] Furthermore, the acoustic frequency range of the disclosed
invention is fully audible (<20 kHz) rather than very high
ultrasound (MHz). The largely undeveloped inventions covered in the
scientific and patent literature typically operate in the medical
ultrasound frequency range, which is 2 to 4 orders of magnitude
higher than that of the disclosed invention here. The interrogation
systems for this prior art are typically standard medical
ultrasound probes, which limits the frequency range of the
implantable sensors. Additionally, embodiments of the disclosed
invention are only sensitive to low levels (several 100 mmHg) of
gauge pressure across the diaphragm. That is, silicon transducers
with diaphragms over vacuum sealed chambers (most of them) will not
exhibit significant pressure dependence of their resonance
frequency; at sensing levels, gauge pressure across the diaphragm
is >800 mmHg For example, the theory of the disclosed invention
must be understood, and additionally, an off the shelf transducer
cannot be used to empirically validate that theory.
[0152] To date, numerous systems exist for attempting to measure
intravascular blood pressure, but all have significant limitations.
For implantable devices, miniaturization and powering are the key
limitations. For noninvasive devices (optical, tonometry), blood
pressure waveforms can easily be generated, but scaling them with
accurate systolic and diastolic values has been a persistent
challenge.
[0153] Embodiments of the disclosed invention offer a solution, by
providing simple passive sensor which can be anchored onto
stent-like structure and be acoustically interrogated. The sensor
can be made extremely small (low micron), and can be made of
extremely stable ceramics (SiO2) to confer long term sensing
stability. Additionally, the device has strong pressure
sensitivity, enabling tenths of mmHg to be accurately measured
[0154] In one example, a mechanical resonator can be configured as
an implantable blood pressure sensor capable of measuring varying
low, medium, and high pressure ranges and operating in one of the
wireless modalities shown in FIG. 47.
[0155] FIG. 48(a) illustrates a coordinate system, while FIG. 48(b)
provides an illustration for an analytical model. In this example,
2a=length of square diaphragm; t=thickness; d=deflection.
Analytic Expression for Resonance Frequency
[0156] In an air environment, the analytical solution for the
natural resonance frequency of a square plate with damped edges
f 0 = 36 2 .pi. ( D a 4 ) ( 1 q ) ( 11 a ) ##EQU00003##
where
D = Et 3 12 ( 1 - v 2 ) ##EQU00004##
Is the flexural rigidity of the diaphragm, a is the square
diaphragm length, and q is the load on the diaphragm including its
weight per unit area and applied pressure. Here, the spring
constant of the diaphragm is
K = ( 324 D .pi. 2 a 4 ) . ##EQU00005##
[0157] FIGS. 49-51 show an experimental setup used to obtain
results previously shown in FIGS. 41-43, and FIGS. 52-53 illustrate
frequency versus pressure data obtained prototyping and development
in the audible range.
[0158] FIGS. 54-55 illustrate the high penetration of audible
acoustic waves in soft tissue. FIG. 54 demonstrates ultrasound
attenuation occurs exponentially with penetration depth, and
increases with increased frequency. The curves show the relative
intensity of ultrasound at a particular frequency as a function of
penetration depth in a medium with an attenuation coefficient of
(0.5 db/cm)/MHz. The total distance traveled by the ultrasound
pulse and echo is twice the penetration depth.
[0159] FIG. 55 demonstrates an attenuation function of a phantom
measured using the pulse-echo substitution method. As shown, in the
frequency range of 1.2 to 4.4 MHz, the least squares line is
y=0.0767+0.692 x. The linear correlation coefficient is 0.9996.
FIG. 56 is a chart showing the high level of acoustic energy that
can be delivered to the resonator for different materials. FIG. 57
shows the reflected pressure ratio and reflected power ratio for
soft tissue in combination with glass, stainless steel, and air,
where:
Reflected Pressure = Z 2 - Z 1 Z 2 + Z 1 ##EQU00006## Reflected
Intensity ( Power Area ) = ( Reflected Pressure ) 2
##EQU00006.2##
[0160] FIGS. 58-60 illustrate a schematic of an experimental set up
for ultrasonic measurements and data obtained from the experiment,
as disclosed in M. W. Borner et. al., Sensors and Actuators A 46-47
(1995) 62-65. FIG. 58 illustrates the schematic for the
measurements, while FIG. 59 illustrates amplitude versus time data
for (a) a micromembranes supported by a nickel honeycomb structure;
(b) a membrane without the microstructure; and (c) the honeycomb
structure alone. As shown in the figure, the echo from the
micromembranes consists of two parts, with the first representing
the initial signal and the second part attributed to vibrations of
the membranes. As shown, echoes from the membrane or microstructure
alone do not show the second part of the signal. FIG. 60 shows a
Fourier transform of the signal, where the resonance frequency of
the micromembranes can be seen.
[0161] As demonstrated herein, resonators used as implantable
sensors provide numerous advantages, including no on-site power
source or circuitry requirements, very small, and a robust design.
Mechanical resonators provide numerous advantages (e.g. over
electrical resonators), including the fact that non-electrical,
extremely small mechanical resonator sensors can be implanted. In
addition, mechanical resonators provide incredible sensitivity,
given how sensitive mechanical resonance is to external pressure,
and can be tailored to specific pressure ranges. Mechanical
resonators theoretically excellent readout range given how well
acoustic signals travel through the body. In addition, mechanical
resonators have much more sensing stability over time, again
because electronics are not necessary, and an elastic ceramic
(quartz, glass, silicon, whatever) will not plastically deform over
time. Furthermore, mechanical resonators provide for pulsewave
recreation because the resonance frequency is high enough to permit
dozens of samples per second in an unoptimized sensor, and possibly
hundreds per second in an optimized sensor.
[0162] In addition, mechanical resonators provide audible acoustic
(<10 kHz) interrogation rather than ultrasound and inexpensive
piezoelectrics can be used instead of expensive ultrasound crystals
and devices. Mechanical resonators provide much simpler readout
electronics with inexpensive piezoelctrics and without frequency
sweeps utilizing a simple, one-time acoustic pulse and then listen
for the resonance echo. Furthermore, mechanical resonators can be
configured with a very small size (.mu.m range in any dimension).
In certain embodiments, mechanical resonators can be sized small
enough to be coupled to a stent and/or for percutaneous delivery to
implantation size.
[0163] All of the apparatus, devices, systems and/or methods
disclosed and claimed herein can be made and executed without undue
experimentation in light of the present disclosure. While the
devices, systems and methods of this invention have been described
in terms of particular embodiments, it will be apparent to those of
skill in the art that variations may be applied to the devices,
systems and/or methods in the steps or in the sequence of steps of
the method described herein without departing from the concept,
spirit and scope of the invention. All such similar substitutes and
modifications apparent to those skilled in the art are deemed to be
within the spirit, scope and concept of the invention as defined by
the appended claims.
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