U.S. patent application number 13/667808 was filed with the patent office on 2013-10-31 for dual modality imaging system for coregistered functional and anatomical mapping.
The applicant listed for this patent is Seno Medical Instruments, Inc.. Invention is credited to Peter Brecht, Bryan Clingman, Andre Conjusteau, Sergey A. Ermilov, Donald G. Herzog, Vyacheslav Nadvoretskiy, Alexander A. Oraevsky, Richard Su, Jason Zalev.
Application Number | 20130289381 13/667808 |
Document ID | / |
Family ID | 49477868 |
Filed Date | 2013-10-31 |
United States Patent
Application |
20130289381 |
Kind Code |
A1 |
Oraevsky; Alexander A. ; et
al. |
October 31, 2013 |
DUAL MODALITY IMAGING SYSTEM FOR COREGISTERED FUNCTIONAL AND
ANATOMICAL MAPPING
Abstract
A real-time imaging system that provides ultrasonic imaging and
optoacoustic imaging coregistered through application of the same
hand-held probe to generate and detect ultrasonic and optoacoustic
signals. These signals are digitized, processed and used to
reconstruct anatomical maps superimposed with maps of two
functional parameters of blood hemoglobin index and blood
oxygenation index. The blood hemoglobin index represents blood
hemoglobin concentration changes in the areas of diagnostic
interest relative to the background blood concentration. The blood
oxygenation index represents blood oxygenation changes in the areas
of diagnostic interest relative to the background level of blood
oxygenation. These coregistered maps can be used to noninvasively
differentiate malignant tumors from benign lumps and cysts.
Inventors: |
Oraevsky; Alexander A.;
(Houston, TX) ; Ermilov; Sergey A.; (Houston,
TX) ; Conjusteau; Andre; (Houston, TX) ;
Brecht; Peter; (Santa Monica, CA) ; Nadvoretskiy;
Vyacheslav; (Houston, TX) ; Su; Richard;
(Sugar Land, TX) ; Herzog; Donald G.;
(Collingswood, NJ) ; Clingman; Bryan; (Chandler,
AZ) ; Zalev; Jason; (Thornhill, CA) |
|
Applicant: |
Name |
City |
State |
Country |
Type |
Seno Medical Instruments, Inc. |
San Antonio |
TX |
US |
|
|
Family ID: |
49477868 |
Appl. No.: |
13/667808 |
Filed: |
November 2, 2012 |
Related U.S. Patent Documents
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Application
Number |
Filing Date |
Patent Number |
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13507217 |
Jun 13, 2012 |
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13667808 |
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13341950 |
Dec 31, 2011 |
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13507217 |
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13287759 |
Nov 2, 2011 |
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13341950 |
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Current U.S.
Class: |
600/407 |
Current CPC
Class: |
A61B 5/14552 20130101;
A61B 2090/306 20160201; A61B 8/14 20130101; A61B 5/726 20130101;
A61B 5/14542 20130101; A61B 8/0825 20130101; A61B 5/0095 20130101;
A61B 5/14546 20130101; A61B 8/5261 20130101; A61B 8/4416 20130101;
A61B 8/0891 20130101; A61B 8/085 20130101; A61B 8/4444 20130101;
A61B 5/0035 20130101; A61B 8/463 20130101; A61B 2562/0238 20130101;
A61B 5/7425 20130101 |
Class at
Publication: |
600/407 |
International
Class: |
A61B 5/00 20060101
A61B005/00 |
Claims
1. An imaging system for visualization of slices into the depth of
tissue of at least a portion of a body, comprising: a hand-held
imaging probe comprising a light emitting portion and an array of
ultrasonic transducers; a processing system configured to receive
data originating from the hand-held imaging probe and to process at
least three independent images based at least in part upon said
data, the three independent images together comprising: a first
functional image reflecting distribution of total hemoglobin
concentration; a second functional image reflecting distribution of
blood oxygen saturation; and a morphological image of tissue
structures; the processing system being further configured to
substantially co-register the first functional image, the second
functional image, and the morphological image in time and space,
and to output a substantially co-registered image.
2. The system of claim 1, in which the light emitting portion and
the array of ultrasonic transducers in the hand-held imaging probe
are arranged in a generally flat linear shape.
3. The system of claim 1, in which the light emitting portion and
the array of ultrasonic transducers in the hand-held imaging probe
are arranged in a curved concave arc shape.
4. The system of claim 1, in which the hand-held imaging probe is
configured to produce at least two optical beams, one on each side
of the array of ultrasonic transducers, so as to deliver optical
energy to a skin surface at such angle and such distance between
them that the optical beams merge into one beam within a distance
of skin thickness under the array of ultrasonic transducers.
5. The system of claim 1, further comprising one or more
dual-wavelength short-pulse lasers.
6. The system of claim 1, further comprising a plurality of
single-wavelength short pulse lasers.
7. The system of claim 1, further comprising a fiberoptic light
delivery system.
8. The system of claim 1, in which the system is configured to
present images substantially in real time by operating at a video
frame rate.
9. The system of claim 1, in which the hand-held optoacoustic probe
is configured to deliver optical energy from either under a face of
said array of transducers or its side.
10. The system of claim 1, in which the hand-held imaging probe
comprises an acoustic lens.
11. The system of claim 10, in which the acoustic lens comprises
optically reflective materials.
12. The system of claim 11, in which the optically reflective
materials comprise a thin, highly optically reflective metallic
layer that removes image artifacts associated with light
interactions with the acoustic lens.
13. The system of claim 11, in which the acoustic lens is formed
from a white opaque material.
14. The system of claim 11, in which the thin, highly optically
reflective metallic layer comprises aluminum, gold, or silver.
15. The system of claim 1, in which the hand-held imaging probe
comprises an output fiber bundle with multiple sub-bundles.
16. The system of claim 15, in which said multiple sub-bundles are
shaped to provide even illumination of the image plane and smooth
illumination edges so as to reduce edge-related optoacoustic
artifacts.
17. The system of claim 1, in which the array of ultrasonic
transducers comprises ultrasonic transducers having an ultrawide
ultrasonic frequency band of sensitivity, with bandwidth of up to
200% from the central frequency.
18. The system of claim 1, in which the hand-held imaging probe
comprises an input fiber bundle that is circular in shape to match
an incident laser beam.
19. The system of claim 1, in which the hand-held imaging probe
comprises an input fiber bundle having a thermally fused fiber
bundle tip such that substantially all fibers in the bundle are
reshaped to avoid loss of light through spaces between fibers.
20. The system of claim 1, in which the hand-held imaging probe
comprises a fiber bundle that is divided into at least two
sub-bundles, with fibers in each sub-bundle being randomized such
that two neighboring fibers at an input appear in different
sub-bundles of the output fiber bundle.
21. The system of claim 1, in which the hand-held imaging probe
comprises a fiber bundle that is divided into at least two
sub-bundles to form fiber bundle paddles, with at least one paddle
placed on each side of the ultrasonic transducer array, each
paddle, in turn, being divided into smaller sub-bundles, each
smaller sub-bundle being in a slot in said paddle so as to provide
controlled profile of an optical beam.
22. The system of claim 1, in which the hand-held imaging probe
comprises fiber bundle, which produces an optical beam that is
shaped to complement a size and shape of the ultrasonic transducer
array.
23. The system of claim 1, in which the hand-held imaging probe
comprises an output fiber bundle having triangular shaped ends so
as to allow an output beam to have smooth edges of optical fluence
after passing through a light diffuser.
24. The system of claim 1, in which the hand-held imaging probe
comprises a plurality of optical windows, each comprising one or
more anti-reflection-coated plates with acoustic impedance matching
that of tissues to be imaged.
25. The system of claim 24, in which the anti-reflection-coated
windows comprise glass, polymer or other solid optically
transparent material.
26. The system of claim 1, in which the hand-held imaging probe
comprises: first and second light diffusers; first and second
optical windows; at least two output fiber bundles arranged such
that optical beams respectively emerging therefrom pass through the
respective light diffusers, then pass through the respective
optical windows, then merge at least partially.
27. The system of claim 1, further comprising a three-dimensional
positioning system configured to control position of the hand-held
probe so as to allow assembly of three-dimensional volumetric
images of the body from two-dimensional slices made though a depth
of tissue obtained by scanning the hand-held probe along the
surface of at least a portion of the body.
28. The system of claim 1, in which the handheld imaging probe
further comprises an acoustic lens formed from a material that
allows it to reflect and scatter light from illumination components
with substantially no absorption of such light, and yet be
optically opaque.
29. The system of claim 28, in which the acoustic lens is formed
from silicon rubber.
30. The system of claim 29, in which the silicon rubber is filled
with titanium dioxide.
31. The system of claim 29, in which the silicon rubber is filled
with barium sulfate powder.
32. The system of claim 1, in which the handheld imaging probe
further comprises a housing that provides hypo-echoic encapsulation
of the probe.
33. The system of claim 32, in which internal or external parts of
the housing comprise materials that do not absorb near-infrared
laser light.
34. The system of claim 33, in which internal or external parts of
the housing comprise materials having low thermal expansion
properties such that the housing does not emit ultrasound after
absorption of laser light.
35. The system of claim 1, in which an assembly of the array of
ultrasonic transducers is made of hypo-echoic material.
36. The system of claim 1, further comprising a layer of
hypo-echoic material between an assembly of the array of ultrasonic
transducers and a fiberoptic assembly to avoid generation of
ultrasound upon interaction of light with assembly of the array of
ultrasonic transducers.
37. The system of claim 1, in which the hand-held imaging probe
comprises a fiber bundle that is divided into at least two
sub-bundles to form at least one fiber bundle paddle, said at least
one fiber bundle paddle being located on one side of the ultrasonic
transducer array.
Description
[0001] This application is a continuation-in-part of U.S. patent
application Ser. No. 13/507,217, filed Jun. 13, 2012, entitled
"System and Method for Acquiring Optoacoustic Data and Producing
Parametric Maps Thereof" and is a continuation-in-part of U.S.
patent application Ser. No. 13/341,950, filed Dec. 31, 2011,
entitled "System and Method for Adjusting the Light Output of an
Optoacoustic Imaging System" and is a continuation-in-part of U.S.
patent application Ser. No. 13/287,759, filed Nov. 2, 2011,
entitled "Handheld Optoacoustic Probe." The entire disclosures of
those applications, including the appendices thereof, are
incorporated herein by reference.
FIELD OF THE TECHNOLOGY
[0002] At least some embodiments disclosed herein relate, in
general, to systems for biomedical imaging, and more particularly,
to real-time imaging systems that visualize thin tissue slices
noninvasively through skin.
BACKGROUND
[0003] Medical ultrasound imaging is a well-established imaging
technology for visualization of tissue morphology in various organs
that provides diagnostic information based on analysis of anatomy.
Optoacoustic imaging is used in medical applications for in vivo
and in vitro mapping of animal and human tissues and organs based
on variation in tissue optical properties. Optoacoustic tomography
can provide anatomical, functional and molecular imaging, but the
most significant value of optoacoustic imaging is in its capability
to provide quantitative functional information based on endogenous
contrast of molecular constituents of red blood cells. The essence
of functional imaging is to provide the physician with a map of
blood distribution and its level of oxygenation, so that the
physician can determine whether particular tissue functions
normally or not. For example, a map of total hemoglobin
distribution simultaneously showing an area with increased
concentration and decreased oxygen saturation indicates potential
malignancy. The essence of molecular imaging is to provide maps of
distributions and concentrations of various molecules of interest
for a specific health condition. For example, distribution of
specific protein receptors in cell membranes gives insight into
molecular biology or cells that aids in designing drugs and
therapeutic methods to treating human diseases.
SUMMARY
[0004] In an embodiment, the invention provides a real-time imaging
system that visualizes thin tissue slices noninvasively through
skin and provides three independent and co-registered images with
biomedically significant information. Specifically, images of deep
biological tissue structures are precisely superimposed with images
of tissue functional state, such as the total hemoglobin
concentration and blood level of oxygen saturation. The invention
in this embodiment thus combines ultrasonic imaging and
optoacoustic imaging technologies in a novel manner. These
technologies can advantageously be combined given the complementary
nature of information provided by them, and the fact that the same
set of ultrasonic/pressure detectors and the same set of analog and
digital electronics can be used to acquire both types of signals
from tissues. In order to achieve a high level of accuracy of
quantitative information and present it substantially in real time
(i.e. substantially as it occurs), a design is disclosed that
utilizes one or more dual-wavelength short-pulse lasers or a
plurality of single-wavelength short pulse lasers, a fiberoptic
light delivery system, a hand-held imaging probe, other electronic
hardware and processing software.
[0005] In an embodiment, an imaging system is disclosed for
visualization of slices into the depth of tissue of at least a
portion of a body. The system includes a processing subsystem that
produces three independent images, including two functional images
showing distribution of the total hemoglobin concentration and
distribution of the blood oxygen saturation and one morphological
image of tissue structures, the images being co-registered in time
and space by utilizing one and the same hand-held imaging probe.
The system may include a three-dimensional positioning system which
provides the capability of assembling three dimensional volumetric
images of said body from two-dimensional slices made though the
depth of tissue obtained by scanning a hand-held probe along the
surface of at least portion of the body.
[0006] In an embodiment, an imaging method provides coregistered
functional and anatomical mapping of tissue of at least a portion
of a body. Ultrasonic pulses are delivered into the tissue and
backscattered ultrasonic signals reflected from various structural
tissue boundaries associated with body morphology are detected. Two
optical pulses having different spectral bands of electromagnetic
energy are delivered, and transient ultrasonic signals resulting
from selective absorption of different energy fractions from each
of the two optical pulses by hemoglobin and oxyhemoglobin of blood
containing tissues are detected. Detected ultrasonic signals are
processed to remove noise, to revert signal alterations in the
course of signal propagation through tissue and through the
detection system components, and to restore the temporal shape and
ultrasonic spectrum of the original signals. Image reconstruction
and processing is performed to generate morphological images of
tissue structures coregistered and superimposed with partially
transparent functional images of the total hemoglobin concentration
and blood oxygen saturation. The above steps of the process are
repeated with a video frame rate so that real-time images can
display tissue functional and morphological changes substantially
as they occur.
BRIEF DESCRIPTION OF THE DRAWINGS
[0007] The disclosed embodiments are illustrated by way of example
and not limitation in the figures of the accompanying drawings in
which like references indicate similar elements.
[0008] FIG. 1A illustrates an embodiment of an optoacoustic probe
with illumination of tissue through skin by a scattered light beam
formed in tissue by merging two optical beams.
[0009] FIG. 1B illustrates how laser illumination light and an
acoustic signal from an optoacoustic probe can be scattered from
the skin towards an acoustic lens of a probe.
[0010] FIGS. 2A and 2B illustrate optoacoustic signals showing the
impact of lateral ultrasonic waves induced by laser pulses in skin
using optical beams on each side of an ultrasound transducer array,
and using detection by transducers tilted at large angle relative
to the plane of images generated therefrom.
[0011] FIG. 3 illustrates an embodiment of manifestation of image
artifacts associated with the edge effect of optical illumination
beam having abrupt changes of the optical fluence.
[0012] FIGS. 4A-4C illustrate embodiments wherein optical
illumination of tissue is accomplished using a hand-held
optoacoustic probe that delivers optical energy from either under
the optoacoustic probe or on the side of the probe at different
distances.
[0013] FIGS. 5A and 5B illustrate two embodiments of a hand-held
optoacoustic ultrasonic probe protected from optical illumination
of the acoustic lens.
[0014] FIG. 6 illustrates optoacoustic images using a probe with an
acoustic lens that is not totally optically reflective and with a
probe having an optically reflective layer of gold, which removes
lens related image artifacts.
[0015] FIG. 7A illustrates an embodiment of an optical beam with
sharp edges that may produce edge effects of acoustic waves and
related artifacts and an optical beam with smooth edges producing
reduced edge-related artifacts.
[0016] FIGS. 7B and 7C illustrate designs of an output fiber bundle
with multiple sub-bundles shaped to provide even illumination of
the image plane and reduce edge related optoacoustic artifacts.
[0017] FIG. 8 illustrates the effect of optical illumination for
two probes where two fiber bundles on each side of the respective
probes are oriented to illuminate skin directly under the
probe.
[0018] FIG. 9A illustrates embodiments of ultrasonic probes having
flat, concave or convex arc shapes.
[0019] FIG. 9B shows a hand-held optoacoustic probe having a
concave arc shape.
[0020] FIG. 9C illustrates details of a hand-held optoacoustic
probe having a concave arc shape.
[0021] FIG. 9D shows an optoacoustic image of three spherical
objects and demonstrates that within the field of view of the arc
spatial (and especially lateral) resolution is excellent even for a
large object.
[0022] FIG. 9E illustrates an alternate embodiment of an
optoacoustic/ultrasonic hand-held probe design.
[0023] FIGS. 10A-10C show examples of the impulse response of an
ultrasonic transducer with a relatively narrow ultrasonic frequency
band of sensitivity, the impulse response of an ultrawide-band
ultrasonic transducer, and the ultrasonic spectra of transducer
sensitivity as a function of frequency for ultrawide-band and
narrow band resonant transducers.
[0024] FIGS. 11A-11B provides an illustrative example of the
deconvolution of impulse response of transducers from the detected
optoacoustic signals where deconvolution restores the original,
unaltered, N-shaped pressure signals.
[0025] FIGS. 12A-12C provide an illustrative example of wavelet
filtered N-shaped optoacoustic signals restored to their original
rectangular pressure profile by summation of all scales
corresponding to frequency ranges from low to high for five scales,
seven scales and nine scales.
[0026] FIG. 13 provides an illustrative diagram of radial
backprojection where each transducer element aperture is weighted
and normalized for the total aperture of the transducer array.
[0027] FIGS. 14A and 14B provide an illustrative example of
optoacoustic tomographic images of an imaging slice through tissue
with a small artery, larger vein and a rectangular grid allowing
estimation of system performance in visualization of
microvessels.
[0028] FIGS. 15A and 15B provide an illustrative example of
optoacoustic tomographic images of a point spread function as
visualized with a flat linear probe using a backpropagation
algorithm and an aperture normalized backprojection algorithm.
[0029] FIGS. 16A and 16B provide an illustrative example of
optoacoustic images of a phantom with hairs embedded at different
depths where the first image was created using an embodiment of a
standard palette and the second image was created using an
embodiment of a depth-normalized palette.
[0030] FIGS. 17A and 17B provide an illustrative example of
optoacoustic images of a phantom of a spherical simulated tumor
obtained with a flat linear probe.
[0031] FIG. 18 shows a diagram illustrating tumor differentiation
based on absorption coefficients at two wavelengths, 757 nm and
1064 nm, which match the local maximum of hemoglobin absorption
like in totally in hypoxic blood (757 nm) and minimum of the ratio
of absorption by hypoxic hemoglobin to absorption by oxyhemoglobin
like in normally oxygenated blood (1064 nm).
[0032] FIG. 19 illustrates tumor differentiation based on
absorption coefficients at two wavelengths in a phantom simulating
benign (box) and malignant (sphere) tumors.
[0033] FIG. 20A shows an optoacoustic image of two intersecting
tubes filled with blood having different levels of blood [SO2].
[0034] FIG. 20B shows a photograph of an experimental setup that
includes artificial blood vessels placed in milk solution and
imaged using arc-shaped optoacoustic probe.
[0035] FIG. 20C shows coregistered 2D cross-sectional anatomical
and functional images of blood vessel tubes showing six image
panels with different anatomical and functional images.
[0036] FIGS. 21A and 21B show optoacoustic signal amplitude as a
function of blood oxygen saturation (with constant hematocrit)
under laser illumination at a wavelength of 1064 nm in FIG. 21A and
at 757 nm in FIG. 21B. These plots illustrate that blood oxygen
saturation can be monitored with optoacoustic imaging.
[0037] FIG. 22 illustrates optical absorption spectra of the main
tissue chromophores absorbing optical energy in the near-infrared
range: hemoglobin, oxyhemoglobin and water.
[0038] FIGS. 23A and 23B illustrate coregistered functional and
anatomical imaging of breast tumors in phantoms accurately
replicating optical and acoustic properties of an average breast
with tumors.
[0039] FIGS. 24A and 24B illustrate coregistered functional and
anatomical imaging of breast tumors.
DETAILED DESCRIPTION
[0040] The following description and drawings are illustrative and
are not to be construed as limiting. Numerous specific details are
described to provide a thorough understanding. However, in certain
instances, well-known or conventional details are not described in
order to avoid obscuring the description. References to one or an
embodiment in the present disclosure are not necessarily references
to the same embodiment; and, such references mean at least one.
[0041] Reference in this specification to "an embodiment" or "the
embodiment" means that a particular feature, structure, or
characteristic described in connection with the embodiment is
included in at least an embodiment of the disclosure. The
appearances of the phrase "in an embodiment" in various places in
the specification are not necessarily all referring to the same
embodiment, nor are separate or alternative embodiments mutually
exclusive of other embodiments. Moreover, various features are
described which may be exhibited by some embodiments and not by
others. Similarly, various requirements are described which may be
requirements for some embodiments but not other embodiments.
System Overview
[0042] In at least some embodiments, the present disclosure is
directed to a dual-modality ultrasonic/optoacoustic system for
medical diagnostics that uses a hand-held probe for scanning along
the skin surface of an organ and provides two types of
two-dimensional maps into the depth of tissue, anatomical
(morphological) and functional (blood hemoglobin index and blood
oxygenation index). In an embodiment, these two maps are spatially
coregistered by using the same array of ultrasonic transducers and
temporally coregistered by acquiring the two types of images in
real time, faster than any physiological changes can occur in the
tissue of diagnostic interest. The blood hemoglobin index
represents blood hemoglobin concentration changes in the areas of
diagnostic interest relative to the background blood concentration.
The blood oxygenation index represents blood oxygenation changes in
the areas of diagnostic interest relative to the background level
of blood oxygenation. These coregistered maps can be used to
noninvasively differentiate malignant tumors from benign lumps and
cysts.
[0043] In an embodiment, the dual-modality ultrasonic/optoacoustic
system of the present disclosure provides two-dimensional imaging
of a body utilizing delivery of optical energy and acoustic
detection of resulting transient pressure waves using
interchangeable hand-held probes, one of which is flat and used to
perform a translational scan through at least a flat portion of the
body under examination, and the second of which is curved being
shaped as a concave arc to perform a translational scan through at
least a cylindrical or curved portion of the body under
examination, both scans contribute to a more complete understanding
normal or pathological functions in the body.
[0044] In an embodiment, at least a portion of the body under
examination contain molecules of blood constituents such as
hemoglobin and oxyhemoglobin responsible for body functions or
receptors in cells responsible for cell functioning, water, lipids
or other constituents.
[0045] In an embodiment, optical energy produced using at least one
laser beam is used for body illumination with at least one
wavelength of light. In an embodiment, the optical energy is
pulsed, with the pulse duration shorter than the time of ultrasound
propagation through the distance in the body equal to the desirable
spatial resolution. In an embodiment, the optical energy is within
the spectral range from 532 nm to 1064 nm. In an embodiment, the
optical energy is replaced with other electromagnetic energy with
the wavelength from 1 nm to lm.
[0046] In an embodiment, electronic signals produced by ultrasonic
transducers are amplified using low noise wide band electronic
amplifiers with high input impedance. In an embodiment, analog
electronic signals are digitized by a multi-channel
analog-to-digital converter and further processed utilizing a field
programmable gate array. In an embodiment, the ultrasonic
transducers are ultrawide-band transducers that detect ultrasonic
signals with no or minimal reverberations. In an embodiment, the
system is integrated with an ultrasound imaging system used to
enhance visualization of acoustic boundaries in the body and parts
of the body with different density and/or speed of sound.
[0047] In an embodiment, quantitative measurements of
concentrations of target molecules, cells or tissues is made
through characterization of optical energy propagation and
absorption combined with processing of digital electronic signals
by deconvolution of the hardware transfer function in order to
obtain intrinsic optoacoustic amplitude and profile of such signals
and the distribution of the optical absorption coefficient in the
body.
[0048] In an embodiment, an optoacoustic contrast agent is used to
visualize a portion of the body of interest or characterize
distribution of certain molecules, cells or tissues in the
body.
[0049] In an embodiment, the system comprises at least a laser, a
light delivery system, an optoacoustic probe, an electronic system,
a computer and an image display.
Laser
[0050] In an embodiment, the laser is capable of emitting short,
nanosecond pulses of near infrared light at two (or more) different
toggling wavelengths, i.e. two different spectral bands. In an
embodiment, one of the wavelengths is preferentially absorbed by
hemoglobin of blood and the other is preferentially absorbed by
oxyhemoglobin of blood. In an embodiment, illumination of the organ
under examination with the first laser pulse at one wavelength
(spectral band) and detection of the first optoacoustic signal
profile resulting from the first illumination, followed by the
illumination with the second laser pulse at the second wavelength
band and detecting of the second optoacoustic signal profile, can
provide data that can be used for reconstruction of two
coregistered tomographic images that can be used for generation of
functional maps of the areas of diagnostic interest based on (i)
blood hemoglobin index and (ii) blood oxygenation index.
Light Delivery System
[0051] In an embodiment, the light delivery system comprises
bundles of optical fibers. In an embodiment, the input of the
optical fiber bundle is circular to match the incident laser beam,
while the output of the fiber bundle is rectangular to match the
size and shape of the ultrasonic transducer array. In an
embodiment, each fiber has a small diameter (e.g., down to 50
micron) to provide excellent flexibility of the bundles. In an
embodiment, the input tip of the fiber bundle is fused to shape the
bundle into a hexagon and to eliminate spaces between the fibers in
the bundle, thereby providing up to 20% better transmission of the
laser energy. In an embodiment, the output tip of the fiber bundle
is fully randomized such that fibers that appear close to each
other at the input will appear far from each other at the output or
even in different branches of the bifurcated fiber bundle.
Optoacoustic Probe
[0052] The probe is designed to provide high contrast and
resolution of both optoacoustic and ultrasonic images. In an
embodiment, the probe is a hand-held probe with an array of
ultrasonic/optoacoustic transducers, which can be designed to be
single dimensional, 1.5 dimensional or two-dimensional. In an
embodiment, the transducers detect acoustic waves within an
ultra-wide band of ultrasonic frequencies and the ultra-wide band
is shaped to match the spectrum of optoacoustic signals emitted by
tissue of diagnostic interest. In an embodiment, the transducers
are also designed to emit acoustic waves as short pulses of
ultrasound with short ring-down time and minimal reverberations of
gradually decreasing magnitude.
[0053] To achieve such a design, transducer material can be chosen
from, for example, piezoelectric ceramics (such as PZT, PMN-PT, and
PZNT), piezoelectric single crystals (such as PZT, PMN-PT, and
PZNT), piezoelectric polymers (such as PVDF and copolymer PVDF
copolymer), composite polymer-ceramic and polymer-crystal
piezoelectric materials and capacitive micromachined ultrasonic
transducers (CMUT). In an embodiment, the thickness of the
transducer elements that provide the central frequency and
materials of the backing layer and the front surface matching layer
of the transducers are optimized.
[0054] In various embodiments, the shape of the ultrasound
transducer array may be either flat or concave arc. A flat design
is suited to scanning of the surface of an organ under examination
that has radius of curvature much larger than the size of the
probe, such as a human body. A concave arc-shaped design provides
the largest aperture for the optoacoustic signal detection with
minimal physical dimensions. The large aperture, in turn, provides
for improved lateral resolution within the angle of the field of
view formed by lines connecting the arc's focal point with each
edge transducer of the array. The arc-shaped probe is often the
most effective for scanning body surfaces that are curved with a
radius approximately matching that of the probe (such as the
average sized breast, neck, arms and legs).
[0055] FIG. 1A illustrates an embodiment of an optoacoustic probe
that provides illumination of tissue (TS) through skin (SK) by the
scattered light (SL) beam formed in tissue by merging two optical
beams (OB) emerging from fiber bundles (FB) expanding and passing
through light diffusers (LD) then passing through optical windows
(OW). Acoustic waves (AW) generated in blood vessels or tumors (BV
or TM) by the scattered light (SL) in tissue propagate through
acoustic lens (AL) to transducers (TR) and converted into
electrical signals being transmitted by electrical cables (EC)
through the backing material (BM) to the electronic amplifiers.
[0056] In an embodiment, the design of the optical fiber bundle is
as follows. The input of the fiber bundle is circular with fused
fiber tips to avoid loss of light through spaces between the
fibers. The fiber diameter may be approximately 200 microns for
good flexibility, and a fiber diameter of 100 microns or even 50
microns may be desirable in a particular application. This fiber
bundle is Y-split into two half-bundles and fully randomized, so
that substantially any two neighboring fibers from the input appear
in different half-bundles. At least a majority of the neighboring
fibers should be randomized in this regard. Each half-bundle is
preferably split into multiple sub-bundles, and each sub-bundle is
placed in its slot/niche to form fiber bundle "paddles". The two
paddles are placed on each side of the ultrasonic transducer (TR)
array assembly. As is discussed below with reference to FIGS. 7B
and 7C, the output shape of each fiber bundle paddle may be
rectangular for the width of the field of view, typically 40 mm,
and have triangular shaped ends. Such triangular shape allows the
output beam to have smooth edges after passing through light
diffuser (LD), FIG. 1A. Finally, the optical beam from fiber bundle
paddles exit from the probe into the skin (SK) through optical
windows (OW) that comprise thin anti-reflection-coated glass plates
or anti-reflection-coated polymer or plastic plates with acoustic
impedance matching that of tissues to be imaged.
[0057] There are a number of objectives for the present
optoacoustic probe design: (i) substantially no light should
propagate either through the acoustic lens (AL) or through the
optical block acoustic damper (OBAD) on the sides of the probe,
(ii) substantially no acoustic waves should be generated in the
acoustic lens or the optical block acoustic damper materials
through absorption of light; acoustic waves in a wide range of
ultrasonic frequencies from 0.1 MHz to 15 MHz should be able to
pass through (AL) with no attenuation, and no acoustic waves should
be able to pass through OBAD; (iii) the optical beams (OB) exiting
through optical windows (OW) should have smooth edges of the
optical fluence, and these optical beams should enter the skin as
close to each other as necessary to merge due to optical scattering
within the skin and enter underlying tissue under the array or
transducers providing maximum fluence in the image plane.
[0058] In an embodiment, the light delivery system directs light
underneath the transducer elements, not through the array of
transducer elements. In an embodiment, the design of the
optoacoustic probe is based on an array of ultrasonic transducers
with fiber optic delivery systems on each side of the ultrasonic
array, positioned as close to the transducers as possible and with
dimensions in the elevation axis of the transducer as small a
possible, considering the need to focus ultrasonic beams to the
depth of the most probable target. In an embodiment, the fiber
optic delivery system is designed to allow penetration of the
optical energy of the near infrared light into the organ being
imaged, such as a breast, and minimum opto-thermo-mechanical
interaction of the light beam with skin.
[0059] Another alternative design of the light delivery system
delivers light to a mirror or prism(s) placed underneath of the
ultrasonic transducers in order to reflect the light orthogonally
to the skin surface of an organ being imaged. In such embodiments,
a standoff can be placed between the transducer elements and the
skin/tissue. These alternative embodiments may be combined within
the scope of the invention.
Detailed Description of Aspects of System Components
Optical Illumination and Probe Design
[0060] An acoustic lens is typically placed on transducers within
an optoacoustic probe for purposes of focusing ultrasonic beams.
While a probe could be provided without an acoustic lens, if there
is no lens then ultrasonic transducers may be directly exposed to
light and absorb such light, which can result in very large
artifact ultrasonic signals, especially where such light is pulsed.
Optical illumination of the lens on an ultrasonic probe causes very
strong transient acoustic waves that result in image artifacts. Up
to 50% of near infrared light can be diffusely scattered by skin,
depending on skin color. Mismatch of acoustic impedance between the
lens of the transducer elements can cause reverberations with long
ring down time. Therefore, an embodiment of a probe design includes
a white strongly scattering opaque lens. If such lens is not needed
due to curved shape of each transducer element, then a white
strongly scattering front matching layer should be employed to
protect transducer elements from near-infrared light.
[0061] FIG. 1B illustrates how laser illumination light 110 and 120
from an optoacoustic probe can be scattered 130 from the skin 140
towards an acoustic lens 150 of a probe.
[0062] Furthermore, (laser) optical pulses can have a direct impact
on ultrasonic transducers of the acoustic waves induced by strong
interaction of the optical pulses with skin of the organ being
imaged that laterally traverse along the skin surface in a
direction orthogonal to the image plane. When detected by the array
of transducers, spatial distributions of these acoustic waves are
projected onto the optoacoustic image at a depth equal to the
lateral distance between the array of transducers and the optical
beams on the skin surface, creating artifacts. Furthermore,
acoustic waves generated in skin through reverberation of the
acoustic lens and the housing of the probe can further affect the
quality of imaging.
[0063] FIGS. 2A and 2B illustrate exemplary optoacoustic signals
showing impact of lateral ultrasonic waves induced by laser pulses
in skin using optical beams on each side of an ultrasound
transducer array. The signals shown are generated by transducers in
the direction almost orthogonal to the plane of images generated
therefrom. Such transducers may receive signals at large oblique
angle (up to 90 deg) relative to the plane of images generated
therefrom, which is undesirable. Therefore, the design of the
transducer array includes means to reject signals coming out of the
image plane. Such means include but not limited to concave arc
shape of the transducer elements and acoustic lens and delivery of
the optical beam underneath the transducers. The detected
optoacoustic signals 210 in FIG. 2A were generated using an
effective acoustic coupling agent, in this case water. The signals
220 in FIG. 2B were generated in the absences of such acoustic
coupling agent, i.e., using only air space to couple the acoustic
signals to the transducer array.
[0064] Furthermore, the finite dimensions of the optical beam can
affect the acoustic waves generated in response to impingement of
the optical beam on tissue. Such acoustic waves can be generated at
the sharp edges of the optical beam, propagate towards the array of
transducers and result in artifacts. Since a system that utilizes
flat linear array of ultrasonic transducers is configured such that
the first and the last transducer in the array detect these waves
first and the central transducers detect these waves the last, this
"edge effect" results in v-shaped artifacts on a sinogram of
optoacoustic signals and low frequency bulk artifacts on
optoacoustic images.
[0065] FIG. 3 illustrates an example of manifestation of v-shaped
artifacts 310 on a sinogram 300 of optoacoustic signals and
associated artifacts 320 on optoacoustic images. Since these
acoustic waves are associated with the edge effect of the optical
illumination beams having abrupt changes of the optical fluence, in
an embodiment, one can see a V-shaped bright signals on the
sinogram and associated series of artifact waves on the
opto-acoustic image.
[0066] Furthermore, the illumination geometry of optical beams
projected by an optoacoustic probe can affect image quality. Where
the optical beams of an optoacoustic probe are placed too far
apart, this can result in a gradual transition from the dark field
illumination (two separate optical beams on each side of the probe
resulting in the absence of direct light under the probe in the
image plane) into the bright field of illumination (one beam under
the probe going into the depth of tissue along the image plane).
This transition creates a problem in the image brightness map
making the map not quantitatively accurate and causes artifacts at
the depth equal to the initial width between separate optical
illumination beams on each side of the probe.
[0067] FIGS. 4A-4C illustrate an embodiment wherein optical
illumination of tissue is accomplished using a hand-held
optoacoustic probe 410, 420, 430 that delivers optical energy from
either under the optoacoustic probe or on the side of the probe at
different distances. In the embodiment of FIG. 4A, when the optical
beams are delivered under the ultrasonic probe, the distribution of
the optical energy in the image plane has a smooth gradient with a
maximum at the skin surface. This optical distribution is
beneficial for high contrast of optoacoustic images. In the
embodiment of FIG. 4B, when the optical beams are delivered close
to a thin optoacoustic probe, the two beams can merge due to
optical scattering within the skin, so that the distribution of the
optical energy in tissue under the skin can be made similar to the
embodiment of FIG. 4A. In the embodiment of FIG. 4C, when the
optical beams are separated by a large distance, they merge only at
significant depth within tissue, creating the optical distribution
in the image plane with a dark zone (no light) in a subsurface
layer of the tissue and a bright zone in the depth of the tissue,
which is detrimental to the contrast of optoacoustic images,
especially considering projection of brightly illuminated areas of
skin onto the optoacoustic image plane at a depth equal to the
separation distance of the two beams.
[0068] Thus, in the embodiments illustrated in FIGS. 4A-4C, the
image brightness map 412, 422 and 432 of the tissue being scanned
is optimized where the illumination of the skin is directly under
the probe 410. As the distance between the center of the
transducers and the center of the optical beams increases, as shown
at 420 and 430, the image brightness map 422 and 432 of the tissue
being scanned becomes progressively more uneven.
[0069] Lastly, the reflection from boundaries of tissue structures
(such as tumor, vessels or tissue layers) of laser-induced
ultrasound waves launched into the tissue after being generated in
skin, can also lead to image artifacts represented by lines, curves
and overall noise.
[0070] In an embodiment, the acoustic lens of the probe is designed
such that the lens reflects and scatters, but does not absorb,
light from the illumination components, yet it is optically opaque.
In various embodiments, such lens can be made either using strongly
optically scattering material such as silicon rubber filled with
titanium dioxide or barium sulfate powder, or using a thin metallic
highly reflective layer such as aluminum or gold or a combination a
white opaque lens material and a metal layer. In an embodiment, to
avoid peel-off of the thin metallic layer from the front surface of
the acoustic lens, in case of a combination of diffusely scattering
material of the lens and a thin reflective layer (foil), the
metallic reflective layer can be placed between the two layers of
diffusely scattering material. As it is difficult to make a
material with absolutely zero optical absorption, and such
absorption may generate ultrasound in thermoelastic materials, the
lens material can be made from thermoplastic materials having
minimal thermal expansion, which produces minimal or no ultrasound
in response to the absorbed optical energy.
[0071] FIGS. 5A and 5B illustrate two embodiments, respectively, of
hand-held optoacoustic ultrasonic probes 510 and 520 that are
protected from optical illumination of the acoustic lens. In FIG.
5A, a totally reflective opaque white lens is utilized, and in FIG.
5B a partially reflective white lens is utilized, with light
reflection capability of the lens enhanced by a gold layer or
coating.
[0072] FIG. 6 illustrates optoacoustic images using a probe with a
non-reflective acoustic lens 610 and a probe with reflective layer
of gold 620. The probe utilizing a reflective layer of gold 620
produces an image with reduced artifacts 612 and 614.
[0073] In an embodiment, the probe housing serves as hypo-echoic
encapsulation of the probe, which means that the probe housing is
made from materials that (i) do not absorb laser light (more
specifically near-infrared light), but if a small absorption is
unavoidable, the materials having low thermal expansion do not emit
ultrasound after absorption of the laser light, (ii) strongly
attenuate and dampen ultrasonic waves and do not reverberate. The
transducer assembly inside the probe housing is also made of the
hypo-echoic material. Alternatively, a layer of said hypo-echoic
material is placed between the transducer assembly and the
fiberoptic assembly to avoid generation of any ultrasound upon
interaction of light with transducer assembly. In various
embodiments, such materials can be chosen, for example, from white
color porous and anechoic heterogeneous composites for baffles,
foams, polymers, rubbers and plastics (such as CR-600 casting resin
available from Micro-Mark of Berkeley Heights, N.J., or AM-37
available from Syntech Materials of Springfield, Va.), and others.
In an embodiment, any such materials are electrically
non-conducting insulators to, inter alia, protect the probe from
external electromagnetic emissions.
[0074] In an embodiment, the optical illumination subsystem is
configured to deliver optical beams with smooth intensity edges. In
an embodiment, the width of the optical beams is equal to that of
the array of ultrasonic transducers within the optoacoustic probe
(for example, about 5 mm). This is achieved by designing the bundle
of optical fibers to have a gradually decreasing density of fibers
at the edges. This design enables one to deliver laser illumination
to the skin of the organ under examination so that the beam does
not generate sharp edge-related acoustic waves, and such
laser-induced acoustic waves do not produce V-shaped artifacts in
the sinogram of optoacoustic images.
[0075] FIG. 7A illustrates an embodiment of an optical beam with
sharp edges 710 that may produce edge effects of acoustic waves and
related artifacts and an optical beam 720 with smooth edges
producing reduced edge-related artifacts. FIGS. 7B and 7C
illustrate that an optical beam with smooth edges of fluence can be
produced using a fiberoptic bundle design having multiple
sub-bundles and a triangular shape in each end of the fiber bundle
assembly.
[0076] In an embodiment, the fiber bundle is positioned at a
distance from skin that is sufficient for the optical beam to
expand to a desirable width. Where the dimensions of the probe are
compact, the fibers used in the fiber bundle can be selected to
have a higher numerical aperture (e.g., >0.22). In an
embodiment, in order to achieve better coupling of the optical beam
into the skin, the beam is delivered through an optical window. In
such embodiment, the optical window touches the skin, making its
surface flat for better light penetration, simultaneously removing
any excess of coupling gel from the skin surface being imaged. In
an embodiment, the fiber bundle and the optical window are
incorporated into the probe housing, so that the air gap between
the fiber bundle and the window is protected.
[0077] In an embodiment, the optical window is designed is to allow
minimal interactions of both the optical beam and the laser-induced
acoustic waves with such window. In an embodiment, the window is
very thin and made of optically transparent material with
antireflection (AR) optical coating. In an embodiment, such
material has anechoic acoustic properties. These anechoic acoustic
properties and the fact that the illuminated skin is being
depressed upon optoacoustic scanning results in dampening of
ultrasonic waves laterally propagating from the laser-illuminated
skin surface to the transducer array, thereby reducing associated
artifacts.
[0078] In an embodiment, the probe is designed such that the
optical beams are very close to the transducer elements on each
side of the ultrasonic probe, which is made as thin as
technologically possible. In an embodiment, the thickness of the
probe is so small (e.g., 5 mm) that the light beams delivered to
skin at this distance, d, from the probe center will merge into one
beam within the thickness of the skin (about z=5 mm), and the
tissue of the organ under examination receives one beam underneath
the transducer elements. This is called bright field illumination.
In an embodiment, the optoacoustic probe is designed such that the
optical light beam is delivered to the skin directly underneath the
transducer elements.
[0079] FIG. 8 illustrates the effect of optical illumination for
two probes 810 and 820 where two fiber bundles on each side of the
respective probes are oriented to illuminate skin directly under
the probe 812 and on either side of the probe 822. Where the skin
is illuminated directly under the probe 812, a tumor 814 is clearly
discernable and there is no clutter on the image background 816.
Where the skin is illuminated on either side of the probe 822, the
tumor is not discernable 824 and there are numerous artifacts on
the image background 826.
[0080] In an embodiment, the optical beam width is designed to
deliver increased light into the slice of tissue being imaged. In
an embodiment, the beam is homogeneous, such that it has a constant
fluence through the beam, as a heterogeneous beam generates
acoustic sources of heterogeneities, which in turn produce
artifacts in optoacoustic images. The fluence level is defined by
the ANSI laser safety standards for the laser illumination of skin.
The beam width is limited by the capability of the optical
scattering in tissue to deliver photons of light into the central
slice located underneath the transducer elements (the slice being
imaged). In an embodiment, the length of the optical beam is equal
to the length of the transducer array. In an embodiment, the
optical beam also has smooth edges, that is to say, gradually
reduced fluence at the edges, since sharp edges produce strong edge
artifacts on optoacoustic images.
[0081] In an embodiment, design features of the optical
illumination system and optoacoustic probe of the present
disclosure can be summarized in the following Table:
TABLE-US-00001 TABLE 1 Summary of optical illumination and probe
design. System Feature Advantages Arc hand-held probe Higher
aperture - lower distortions Light delivery into Improves
optoacoustic image contrast and the imaging plane decreases
artifacts by increasing the ratio of useful information (from the
imaging plane) to noise (outside of the imaging plane) Optical
shielding of Reduces artifacts from direct and scattered light the
probe striking the acoustic lens, probe housing, etc. Acoustic
shielding of Acoustic shielding of the probe's housing reduces the
probe artifacts (clutter) from acoustic waves propagating through
the probe's housing Using ultrawide band Allows to have the same
array working in transducers for both ultrasonic and optoacoustic
imaging ultrasound and optoacoustic imaging
[0082] In various embodiments, the shape of the ultrasonic
transducer array for the combined optoacoustic/ultrasonic imaging
can be either flat or convex arc-shaped. In an embodiment, the
probe shape for optoacoustic imaging is concave arc-shaped. Such a
concave shape provides a large aperture with minimal physical
dimensions, wider field of view of an object being imaged, which in
turn provides for improved lateral resolution and better
reconstruction of the shape of the object being imaged.
[0083] FIGS. 9A-9C illustrate embodiments of
optoacoustic/ultrasonic hand-held probes having flat or concave arc
shapes 910 (FIG. 9A) and a hand-held transrectal probe with a
linear shape 920 (FIG. 9B). FIG. 9C illustrates details of the
optoacoustic/ultrasonic hand-held probe design with its face
showing ultrasonic transducers assembly, two layers of hypo-echoic
light reflecting and ultrasound damping material on each side, and
two optical windows for delivery of the optical beam.
[0084] FIG. 9C illustrates details of a hand-held optoacoustic
probe having a concave arc shape. Electrical cables 930 are
provided for bi-directional communication to and from the probe,
and fiberoptic bundles 940 are provided for delivering light to the
probe. An array of wide-band ultrasonic transducers 950 send and
receive acoustic energy. The transducer array 950 is covered by an
opaque white cylindrical lens (not shown for clarity purposes) that
extensively scatters and reflects near-infrared light. Optical
windows 960 provide optical beam outputs. In the embodiment FIG.
9C, the ultrasonic transducers within the probe may be designed so
as not to be sensitive to lateral acoustic (ultrasonic) waves and
to be free of reverberations, especially in the lateral direction.
This can be achieved by the selection of the piezoelectric
composite material, the shape of piezoceramic elements in the
matrix and anechoic properties of the matrix. In an embodiment, the
ultrasonic transducers are also designed to possess high
sensitivity within an ultrawide band of ultrasonic frequencies.
This in turn results in minimal reverberations that cause artifacts
on optoacoustic/ultrasonic images.
[0085] FIG. 9D shows an optoacoustic image that illustrates
advantages of the concave-arc shaped hand-held probe in terms of
resolution in optoacoustic images. As presented in this embodiment,
the shape and sharp edges of a large sphere are well depicted in
cases where the object is within the field of view of the probe
aperture. Outside the probe aperture resolution and accuracy of
shape reproduction decreases, however remains better than those for
flat linear probes of similar width.
[0086] FIG. 9E illustrates an alternate embodiment of an
optoacoustic/ultrasonic hand-held probe design that is capable of
two-dimensional imaging within the plane going parallel to the skin
surface at various selected depths, and three-dimensional images as
well.
[0087] In an embodiment, a hand-held probe that is scanned along
the skin surface producing real-time two-dimensional images of
tissues in the body under examination also has a component serving
for accurate global 3D positioning of the probe. This design allows
the imaging system to remember positions of all tissue slices and
to reconstruct three-dimensional images at the end of the scanning
procedure.
Electronic Data Acquisition System
[0088] In an embodiment, the present disclosure is directed to an
optoacoustic imaging system having an electronic data acquisition
system that operates in both optoacoustic and ultrasonic modes and
can rapidly switch between such modes. In an embodiment, this is
achieved with firmware that controls functions of a Field
Programmable Gate Array (FPGA), the main microprocessor on the
electronic data acquisition system. In an embodiment, a
reprogrammable FPGA can toggle between optoacoustic and ultrasound
operation modes in real-time, thus enabling co-registration of
ultrasound and optoacoustic images, which can be used for
diagnostic imaging based on functional and anatomical maps. In an
embodiment, FPGA functions include controlling, acquiring, and
storing optoacoustic and/or ultrasound data, signal processing and
transferring data for real-time image reconstruction and
processing. In an embodiment, the FPGA may also be employed in
ultrasound beam forming and image reconstruction.
[0089] In an embodiment, the electronic data acquisition system
design utilizes one or more multi-core Graphical Processor Units
(GPU) for image reconstruction and processing. In the ultrasound
mode, in an embodiment, the FPGA controls the ultrasound
transmission and it performs both ultrasound and optoacoustic data
acquisitions on a multi-channel board. In order to enhance
operation of the memory of the FPGA, an external memory buffer can
be used. In an embodiment, the FPGA allows rapid reprogramming of
ultrasonic data acquisition with signal/frame repetition rate of
about 2 to 20 kHz to optoacoustic data acquisition with
signal/frame repetition rate of about 10-20 Hz, and also configures
the structure of gates and internal memory structure and size to
allow real-time switching between ultrasound emission and
detection, laser synchronization, and system controls. In an
embodiment, multiple FPGAs can be used to enhance the system
performance. In an embodiment, in the ultrasound and optoacoustic
modes, the FPGA clock can be changed by the appropriate
time-division multiplexing (TDM). In an embodiment, the design of a
multi-channel electronic data acquisition system can be based on
modules, with a module typically being from 16 to 128 channels,
although 256 channels or more may be appropriate in some
applications. In an embodiment, the design of a multi-channel
electronic data acquisition system has 64 channels.
[0090] In order to achieve dual modality operation of the
optoacoustic/ultrasonic system, a separate optoacoustic electronic
system could also be combined with a separate ultrasonic electronic
system through a single probe. In an embodiment, the probe has a
cable that has a Y-split to connect the probe to optoacoustic and
ultrasonic electronic systems. In an embodiment, a programmable
electronic switch allows one to send the detected signal from the
probe (transducer array) either to the optoacoustic electronics (to
operate in optoacoustic mode) or to the ultrasonic electronics and
from the ultrasonic electronics to the probe (to operate in
ultrasound mode). In an embodiment, a synchronization trigger
signal is sent to the optoacoustic and ultrasonic systems
sequentially, so that the optoacoustic and ultrasonic images are
acquired one after the other.
Processing, Reconstruction and Display of Images
Signal Processing
[0091] In various embodiments, a goal of the diagnostic imaging
procedure is to display each pixel with a brightness that correctly
replicates the originally generated signals in each voxel of tissue
displayed on the image. On the other hand, intrinsic pressure
signals generated by optical pulses within tissues may be
significantly altered in the course of propagation through tissue
and especially in the course of detection and recording by the
ultrasonic transducers and the electronics subsystem.
[0092] In an embodiment, detected signals are processed to reverse
alterations and restore original signals. In an embodiment, such
reversal can be achieved through deconvolution of the impulse
response (IR) of the system. In an embodiment, the impulse response
can be measured by recording and digitizing a delta-function
ultrasonic signal generated by short (nanosecond) laser pulses in a
strongly absorbing optical medium with high thermoelastic expansion
coefficient.
[0093] One component of the impulse response is the
acousto-electrical impulse response, which provides for the
optoacoustic or ultrasonic signal distortions due to the properties
of the ultrasonic transducers, cables and analog electronics. A
second part of the impulse response is the spatial impulse response
that provides for the signal distortions associated with finite
dimensions of the ultrasonic transducers. In various embodiments,
large transducers can integrate ultrasonic waves incident at an
angle, whereas point-source-like transducers can provide perfect or
near perfect delta-function spatial impulse response.
[0094] In an embodiment, any distortions in acousto-electrical
impulse response can be reversed by the impulse response
deconvolution from the detected signals. However, possible
distortions in the spatial impulse response can be avoided by
designing transducers with small dimensions within the image plane.
In an embodiment, the dimensions of the transducers are much
smaller than the shortest wavelength of the ultrasound that may be
detected or emitted by the transducers.
[0095] FIGS. 10A-10C show examples of the impulse response of an
ultrasonic transducer with a relatively narrow band of sensitivity
1010, the impulse response of an ultrawide-band ultrasonic
transducer 1020 and ultrasonic spectra of the transducer
sensitivity as a function of frequency for ultrawide-band and
narrow band resonant transducers 1030.
[0096] In an embodiment, the first step in processing an
optoacoustic signal in an imaging system that produces
two-dimensional optoacoustic images is deconvolution of the
acousto-electrical impulse response.
[0097] FIGS. 11A and 11B provide an illustrative example of the
deconvolution of impulse response of transducers from the detected
optoacoustic signals 1110, where deconvolution restores the
original, unaltered, N-shaped pressure signals 1120.
[0098] In an embodiment, the second step in processing an
optoacoustic signal is signal filtering to remove noise using a
signal filter. In an embodiment, the signal filter is based on a
wavelet transform that operates simultaneously in the frequency and
time domain. In an embodiment, such a wavelet filter is capable of
filtering certain frequency components of the signal that belong to
noise and appear at a given time, while preserving similar
frequency components of the useful signal that appear at a
different time. In an embodiment, the frequency spectrum of a
wavelet filter replicates the frequency band of a typical N-shaped
optoacoustic signal while simultaneously providing smooth window
edges which do not cause signal distortions upon convolution.
[0099] In an embodiment, such a wavelet filter is useful in
optoacoustic imaging in its capability to restore the original
pressure profile generated in tissue prior to pressure propagation.
In the course of propagation through tissue, the originally
positive pressure signal converts into a bipolar
(compression/tension) profile. Therefore, reconstruction of an
image of absorbed optical energy (optoacoustic image) requires a
transformation that starts with bipolar signals and provides for
all-positive values of the optoacoustic image intensities. In an
embodiment, a multi-scale wavelet filter, for example, a filter
that simultaneously integrates the signal over time and provides
summation of a number of frequency bands present in the signal, can
convert bipolar pressure signals into monopolar signal representing
thermal energy or originally generated positive pressure.
[0100] FIGS. 12A-12C provide an illustrative example of wavelet
filtered N-shaped optoacoustic signals restored to their original
rectangular pressure profile by summation of all scales
corresponding to frequency ranges from low to high for five scales
1210, seven scales 1220 and nine scales 1230.
[0101] In various embodiments, wavelet filtering permits
enhancements of objects on the image within certain range of
dimensions. An imaging operator (ultrasonic technician or
diagnostic radiologist) typically desires to better visualize a
tumor with specific dimensions and other objects, such as blood
vessels, with their specific dimensions. In an embodiment, the
wavelet filter allows the operator to apply specific selection of
scales of the wavelet filter than would enhance only objects of
certain sizes and suppress object of other unimportant sizes. In an
embodiment, boundaries can be well visualized for objects of any
size, so the high-frequency wavelet scales are beneficial for the
image quality and are included in the selection of scales. In an
embodiment, for a mathematically correct tomographic
reconstruction, a ramp filter can be applied to the signal, which
can linearly enhance contribution of higher frequencies.
Image Reconstruction
[0102] In various embodiments, image reconstruction typically uses
radial back-projection of processed and filtered signals onto the
image plane. However, due to the limited field of view available
from small hand-held probes, only an incomplete data set can be
obtained. As a result, the 2D optoacoustic images may include
artifacts distorting the shape and brightness of the objects
displayed on the images. In an embodiment, aperture integrated
normalized radial back projection is used to correct some of the
reconstruction artifacts that are observed in limited aperture
optoacoustic tomography.
[0103] FIG. 13 provides an illustrative diagram of radial
backprojection where each transducer element aperture is weighted
and normalized for the total aperture of the transducer array.
[0104] In an embodiment, T.sub.k,-T.sub.k+4, 1311-1315, are the
transducers 1310 in the array, B.sub.i,j is the brightness
(intensity) of a pixel with coordinates (i,j), .omega..sub.i,j,k
1320, 1330 is the angular portion of optoacoustic wave front
emitted by the pixel (i,j) as it is visualized by the transducer
#k, .OMEGA..sub.i,j.SIGMA..omega..sub.i,j,k (sum of all
.omega..sub.i,j,k) is the portion of the optoacoustic wave front
emitted by pixel (i,j) as it is visualized by the entire transducer
array, and S.sub.i,j,k is the sample of the optoacoustic signal
measured by k.sup.th transducer and used in reconstruction of the
brightness in the pixel (i,j). Various backpropagation algorithms
can be used to normalize an optoacoustic image.
[0105] In an embodiment, a backpropagation algorithm can be
expressed as:
B.sub.i,j=.SIGMA..sub.kS.sub.i,j,k (1)
[0106] However, in at least some embodiments, aperture normalized
backprojection produces superior image results. In an embodiment,
the aperture normalized backprojection can be expressed as:
B i , j = 1 .OMEGA. i , j k S i , j , k .OMEGA. i , j , k ( 2 )
##EQU00001##
[0107] FIGS. 14A and 14B provide an illustrative example of
optoacoustic tomographraphic images 1410 and 1420 of an imaging
slice through a tumor angiogenesis model. In the first image 1410,
a backpropagation algorithm, such as the first algorithm
immediately above, is used to normalize the image. The resulting
image has strong, bright arc-shaped artifacts 1412 around the blood
vessels 1414 that are close to array surface. In the second image
1420, aperture normalized backprojection algorithm, such as the
second algorithm immediately above, is used to normalize the image.
As can be seen, the aperture normalized backprojection algorithm
corrects image brightness and reduces the arc-shaped artifacts.
[0108] FIGS. 15A and 15B provide an illustrative example of
optoacoustic tomographic images 1510 and 1520 of a point spread
function as visualized with flat linear probe using 1510 a
backpropagation algorithm, such as the first algorithm immediately
above, and 1520 an aperture normalized backprojection algorithm,
such as the second algorithm immediately above. As can be seen, the
aperture normalized backprojection algorithm corrects image
brightness and reduces artifacts.
Image Processing and Display
[0109] In an embodiment, the optoacoustic image palette is
equalized to diminish effects of light distribution within tissue.
Such equalization transforms the dynamic range of optoacoustic
images for better visualization of both shallow and deep
objects.
[0110] FIGS. 16A and 16B provide an illustrative example of
optoacoustic images 1610 and 1620 of a phantom with hairs embedded
at different depths where the first image 1610 was created using a
an embodiment of a standard palette and the second image 1620 was
created using an embodiment of a depth-normalized palette. As can
be seen, utilizing the depth-normalized palette enhances visibility
of deep objects in the illustrated embodiment.
[0111] In an embodiment, principal component analysis (PCA) on a
single optoacoustic image acquisition (different channels) is used
to remove cross-correlated signal noise. Principal component
analysis on a dataset of optoacoustic signals can remove correlated
image clutter. Principal component analysis on optoacoustic frames
can also remove correlated image clutter.
[0112] FIGS. 17A and 17B provide an illustrative example of
optoacoustic images 1710 and 1720 of a phantom of a spherical
simulated tumor obtained with flat linear probe. The first image
1710 is a raw image that was not subjected to principal component
analysis processing. The second image 1720 has been subjected to
principal component analysis processing with first principal
component deconvolution. As can be seen, utilizing principal
component analysis processing enhances image quality by, inter
alia, reducing artifacts.
[0113] In an embodiment, design features of signal and image
processing of the present disclosure can be summarized in the Table
2 as follows:
TABLE-US-00002 TABLE 2 Summary of signal and image processing.
System Feature Advantages Operator-assisted Can improve
quantitative optoacoustic diagnostics by evaluating the boundary
tracking on diagnostic parameters within the tumor boundary defined
on US images ultrasonic and Diagnostics can be enhanced by
morphological analysis of the tumor optoacoustic images boundary
Aperture integrated Corrects some of the reconstruction artifacts
that are observed in a normalized radial limited aperture
optoacoustic tomography back projection Equalization of the
Transforms dynamic range of optoacoustic images for better
optoacoustic image visualization of both shallow and deep objects
palette to diminish effects of light distribution within the tissue
Principal component PCA on a single optoacoustic acquisition
(different channels) is a fast analysis (PCA) of the and efficient
way to remove cross-correlated signal noise optoacoustic signal PCA
on a dataset of optoacoustic signals removes correlated image data
clutter PCA on optoacoustic frames removes correlated image clutter
Optoacoustic imaging Cancer diagnostics based on those parameters
or a single malignancy system with index (tHb * water/oxygenation)
with respect to average background quantitative assessment of total
hemoglobin, blood oxygenation, and water Wavelet transform Operator
can easily select the maximum size of the objects to be that
enhances images enhanced on the image. Everything larger will be
filtered out of objects within certain dimension range Adaptive
beam- Allows individual reconstruction on a family of radial
wavelet sub- forming for bands optoacoustic imaging
Diagnostic Image Reprocessing
[0114] The principles of functional diagnostic imaging can be based
on the tumor pathophysiology. For example, malignant tumors have
enhanced concentration of the total hemoglobin and reduced level of
oxygen saturation in the hemoglobin of blood. In an embodiment,
optoacoustic images can be reprocessed and converted into, inter
alia, images of (i) the total hemoglobin [tHb] and (ii) the oxygen
saturation of hemoglobin [SO2]. FIG. 18 demonstrates an example of
two breast tumors
[0115] FIG. 18 shows a diagram illustrating tumor differentiation
based on absorption coefficients at two wavelengths, 755 nm, 1810,
and 1064 nm, 1820, which match the local maximum (757 nm) and
minimum (1064 nm) of the ratio of absorption by hemoglobin (hypoxic
blood) to absorption by oxyhemoglobin. As can be seen, a malignant
tumor, 1830, has a higher absorption coefficient at 757 nm than a
benign tumor, 1840, whereas the benign tumor, 1840, has a higher
absorption coefficient at 1064 nm than a malignant tumor, 1830.
[0116] FIG. 19 illustrates tumor differentiation by optoacoustic
imaging based on absorption coefficients at two wavelengths 1910
and 1920 in a phantom. At 757 nm, 1920, a model of a malignant
tumor is clearly visible, 1922, whereas the model of the malignant
tumor, 1922, is not visible at 1064 nm, 1910.
[0117] FIG. 20A shows an optoacoustic image of two intersecting
tubes filled with blood having different levels of blood [SO2] (98%
in the left tube, and 31% in the right tube). The tubes were placed
in 1% fat milk with optical properties similar to those found in
the human breast. The wavelength of laser illumination used for
this image is 1064 nm. FIG. 20B shows a photograph of an
experimental setup that includes artificial blood vessels placed in
milk solution and imaged using arc-shaped optoacoustic probe. FIG.
20C shows coregistered 2D cross-sectional anatomical and functional
images of blood vessel tubes showing six image panels: (1--upper
left) ultrasound image depicting anatomy of the body with vessels;
(2--upper right) optoacoustic image obtained at the wavelength of
757 nm; (3--lower right) optoacoustic image obtained at the
wavelength of 1064 nm; (4--lower left) functional image of the
total hemoglobin [tHb]; (5--lower center) functional image of the
blood oxygen saturation [SO2]; (6--upper center) functional image
of the blood oxygen saturation presented only in the area of
maximum concentration of the total hemoglobin. Raw optoacoustic
images depicted in FIG. 20C in the upper right and lower right
panels demonstrate different brightness of blood vessels having
blood with different level of the total hemoglobin concentration
[tHb] and blood oxygen saturation [SO2], accurate quantitative
measurements could be performed under conditions of normalized
fluence of the optical illumination of tissue in the body as a
function of depth. These optoacoustic images were used to
reconstruct functional images of the total hemoglobin [tHb] and the
blood oxygenation [SO2]. All functional images displayed in FIG.
20C are coregistered and superimposed with the anatomical image of
tissue structure for better correlation of features.
[0118] FIGS. 21A and 21B show optoacoustic signal amplitude as a
function of blood oxygen saturation (with constant hematocrit)
under laser illumination at the wavelength of 1064 nm in FIG. 21A
and at 757 nm in FIG. 21B. These plots illustrate that blood oxygen
saturation can be monitored with optoacoustic imaging.
Specifically, this embodiment illustrates quantitative data based
on measurements of the optoacoustic signal amplitude in blood
having various levels of oxygen saturation (from 30% to 98%) and
hematocrit of 38 g/dL of hemoglobin [tHb] in erythrocytes. As
predicted by the published absorption spectra of blood, the
optoacoustic signal amplitude at 1064 nm illumination increases
with increased level of oxygen saturation, while the optoacoustic
signal amplitude decreases with increased blood oxygenation at 757
nm illumination wavelength.
[0119] FIG. 22 illustrates optical absorption spectra of the main
tissue chromophores absorbing optical energy in the near-infrared
range: hemoglobin, oxyhemoglobin and water. Preferred laser
wavelengths for functional imaging are 757 nm and 1064 nm matching
max and min ratio of [HHb]/[O2Hb], while the wavelength of 800 nm
is the best for calibration purposes through measurements of the
total hemoglobin [tHb].
[0120] FIGS. 23A and 23B illustrate coregistered functional and
anatomical imaging of breast tumors in phantoms accurately
replicating optical and acoustic properties of an average breast
with tumors. FIG. 23A shows 2D images of: model of malignant tumor
morphology based on ultrasound (left), the same anatomical image
coregistered with functional image of the total hemoglobin
concentration (center) and with functional image of the blood
oxygenation (right). FIG. 23B shows 2D images of a model benign
tumor: morphology image based on ultrasound (left), the same
anatomical image coregistered with functional image of the total
hemoglobin concentration (center) and with functional image of the
blood oxygenation (right).
[0121] FIGS. 24A and 24B illustrate coregistered functional and
anatomical imaging of breast tumors. FIG. 24A shows 2D images of
invasive ductal carcinoma, a malignant tumor with rough boundaries,
heterogeneous morphology, high concentration of total hemoglobin
and low oxygen saturation (hypoxia). The malignant tumor morphology
is based on ultrasound in the left image, and the same anatomical
image coregistered with functional image of the blood oxygenation
in the center image and with functional image of the total
hemoglobin concentration in the right image. FIG. 24B shows 2D
images of a breast with Fibroadenoma, a benign tumor with
relatively round boundaries, normal concentration of oxyhemoglobin
and relatively low total hemoglobin. Breast morphology is based on
ultrasound in the left image, and the same anatomical image is
coregistered with a functional image of the blood oxygenation in
the center image and with a functional image of the total
hemoglobin concentration in the right image.
CONCLUSION
[0122] While some embodiments can be implemented in fully
functioning computers and computer systems, various embodiments are
capable of being distributed as a computing product in a variety of
forms and are capable of being applied regardless of the particular
type of machine or computer-readable media used to actually effect
the distribution.
[0123] At least some aspects disclosed can be embodied, at least in
part, in software. That is, the techniques described herein may be
carried out in a special purpose or general purpose computer system
or other data processing system in response to its processor, such
as a microprocessor, executing sequences of instructions contained
in a memory, such as ROM, volatile RAM, non-volatile memory, cache
or a remote storage device.
[0124] Routines executed to implement the embodiments may be
implemented as part of an operating system, firmware, ROM,
middleware, service delivery platform, SDK (Software Development
Kit) component, web services, or other specific application,
component, program, object, module or sequence of instructions
referred to as "computer programs." Invocation interfaces to these
routines can be exposed to a software development community as an
API (Application Programming Interface). The computer programs
typically comprise one or more instructions set at various times in
various memory and storage devices in a computer, and that, when
read and executed by one or more processors in a computer, cause
the computer to perform operations necessary to execute elements
involving the various aspects.
[0125] A machine-readable medium can be used to store software and
data which when executed by a data processing system causes the
system to perform various methods. The executable software and data
may be stored in various places including for example ROM, volatile
RAM, non-volatile memory and/or cache. Portions of this software
and/or data may be stored in any one of these storage devices.
Further, the data and instructions can be obtained from centralized
servers or peer-to-peer networks. Different portions of the data
and instructions can be obtained from different centralized servers
and/or peer-to-peer networks at different times and in different
communication sessions or in a same communication session. The data
and instructions can be obtained in entirety prior to the execution
of the applications. Alternatively, portions of the data and
instructions can be obtained dynamically, just in time, when needed
for execution. Thus, it is not required that the data and
instructions be on a machine-readable medium in entirety at a
particular instance of time.
[0126] Examples of computer-readable media include but are not
limited to recordable and non-recordable type media such as
volatile and non-volatile memory devices, read only memory (ROM),
random access memory (RAM), flash memory devices, floppy and other
removable disks, magnetic disk storage media, optical storage media
(e.g., Compact Disk Read-Only Memory (CD ROMS), Digital Versatile
Disks (DVDs), etc.), among others.
[0127] In general, a machine readable medium includes any mechanism
that provides (e.g., stores) information in a form accessible by a
machine (e.g., a computer, network device, personal digital
assistant, manufacturing tool, any device with a set of one or more
processors, etc.).
[0128] In various embodiments, hardwired circuitry may be used in
combination with software instructions to implement the techniques.
Thus, the techniques are neither limited to any specific
combination of hardware circuitry and software nor to any
particular source for the instructions executed by the data
processing system.
[0129] Although some of the drawings illustrate a number of
operations in a particular order, operations that are not order
dependent may be reordered and other operations may be combined or
broken out. While some reordering or other groupings are
specifically mentioned, others will be apparent to those of
ordinary skill in the art and so do not present an exhaustive list
of alternatives. Moreover, it should be recognized that the stages
could be implemented in hardware, firmware, software or any
combination thereof.
[0130] In the foregoing specification, the disclosure has been
described with reference to specific exemplary embodiments thereof.
It will be evident that various modifications may be made thereto
without departing from the broader spirit and scope as set forth in
the following claims. The specification and drawings are,
accordingly, to be regarded in an illustrative sense rather than a
restrictive sense.
* * * * *