U.S. patent application number 09/954332 was filed with the patent office on 2002-03-21 for photoacoustic breast scanner.
This patent application is currently assigned to Optosonics, Inc.. Invention is credited to Kruger, Robert A..
Application Number | 20020035327 09/954332 |
Document ID | / |
Family ID | 24891163 |
Filed Date | 2002-03-21 |
United States Patent
Application |
20020035327 |
Kind Code |
A1 |
Kruger, Robert A. |
March 21, 2002 |
Photoacoustic breast scanner
Abstract
Methods and apparatus for measuring and characterizing the
localized electromagnetic wave absorption properties of biologic
tissues in vivo, using incident electromagnetic waves to produce
resultant acoustic waves. Multiple acoustic transducers are
acoustically coupled to the surface of the tissue for measuring
acoustic waves produced in the tissue when the tissue is exposed to
a pulse of electromagnetic radiation. The multiple transducer
signals are then combined to produce an image of the absorptivity
of the tissue, which image may be used for medical diagnostic
purposes. In specific embodiments, the transducers are moved to
collect data from multiple locations, to facilitate imaging.
Specific arrangements of transducers are illustrated. Also,
specific mathematical reconstruction procedures are described for
producing images from transducer signals.
Inventors: |
Kruger, Robert A.;
(Indianapolis, IN) |
Correspondence
Address: |
WOOD, HERRON & EVANS, L.L.P.
2700 Carew Tower
Cincinnati
OH
45202
US
|
Assignee: |
Optosonics, Inc.
|
Family ID: |
24891163 |
Appl. No.: |
09/954332 |
Filed: |
September 17, 2001 |
Related U.S. Patent Documents
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Application
Number |
Filing Date |
Patent Number |
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09954332 |
Sep 17, 2001 |
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09604752 |
Jun 27, 2000 |
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6292682 |
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09604752 |
Jun 27, 2000 |
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09076968 |
May 13, 1998 |
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6102857 |
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09604752 |
Jun 27, 2000 |
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PCT/US97/17832 |
Oct 1, 1997 |
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PCT/US97/17832 |
Oct 1, 1997 |
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08719736 |
Oct 4, 1996 |
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5713356 |
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Current U.S.
Class: |
600/437 |
Current CPC
Class: |
A61B 5/0091 20130101;
A61B 5/4312 20130101; A61B 8/4281 20130101; A61B 5/7239 20130101;
A61B 5/0507 20130101; A61B 5/0095 20130101 |
Class at
Publication: |
600/437 |
International
Class: |
A61B 008/00 |
Claims
What is claimed is:
1. A method of imaging tissue structures in a three-dimensional
volume of tissue by detecting localized absorption of
electromagnetic waves in said tissue, comprising providing a source
of electromagnetic radiation in proximity to said tissue; providing
a plurality of acoustic sensors; acoustically coupling said
plurality of acoustic sensors to said tissue via a coupling media
chosen from one or more of: water, salinated water, alcohol, oil
and mineral oil; irradiating said three-dimensional volume of
tissue with a pulse of diffuse electromagnetic radiation from said
source to generate resultant pressure waveforms within said
three-dimensional volume of tissue; detecting said resultant
pressure waveforms arriving at said acoustic sensors and storing
data representative of said waveforms; combining a plurality of
said detected pressure waveforms to derive a measure of pressure
waveforms originating at a point distant from said acoustic
sensors; and repeating said combining step for a plurality of
points to produce an image of structures in said tissue.
2. The method of claim 1 wherein said step of combining a plurality
of detected pressure waveforms to derive a measure of pressure
waveforms originating at a point comprises determining a distance
between said point and a pressure sensor, computing a value related
to the time rate of change in a pressure waveform detected by said
pressure sensor at a time which is a time delay after said pulse of
electromagnetic radiation, said time delay being equal to the time
needed for sound to travel said distance through said tissue;
repeating said determining and computing for additional pressure
sensors and pressure sensor waveforms; and accumulating said
computed values to form said measure of pressure waveforms
originating at said point.
3. The method of claim 2 wherein said step of providing a plurality
of acoustic sensors comprises providing a differentiating acoustic
sensor responsive to a pressure waveform by producing an electrical
output representative of a time rate of change of said pressure
waveform, and said step of computing a value of the time rate of
change in a pressure waveform, comprises computing a value of said
electrical output of said differentiating pressure sensor.
4. The method of claim 3 wherein said differentiating acoustic
sensor includes a piezoelectric crystal which produces an analog
signal, and producing an electrical output representative of a time
rate of change comprises combining a delayed version of said analog
signal with said analog signal to produce said electrical
output.
5. The method of claim 2 wherein computing a value related to the
time rate of change in a pressure waveform at a time delay, further
comprises multiplying said time rate of change by a factor
proportional to said time delay to produce said value, whereby to
compensate for diffusion of acoustic energy radiated from said
point.
6. The method of claim 1 wherein said step of combining a plurality
of detected pressure waveforms to derive a measure of pressure
waveforms originating at a point, comprises determining a distance
between said point and a pressure sensor; computing a value related
to a sum of the pressure waveform detected by said pressure sensor
over a time period, said time period beginning substantially
contemporaneous with said pulse of electromagnetic radiation and
said time period having a duration equal to the time needed for
sound to travel said distance through said tissue; repeating said
determining and computing for additional pressure sensors and
pressure sensor waveforms; and accumulating said sums to form said
measure of pressure waveforms originating at said point.
7. The method of claim 6 wherein computing a value related to a sum
of the pressure waveform detected by a pressure sensor over a time
period, further comprises multiplying said sum by a factor
proportional to the duration of said time period to produce said
value, whereby to compensate for diffusion of acoustic energy
radiated from said point.
8. The method of claim 1 wherein providing said plurality of
sensors comprises providing a surface and positioning said sensors
evenly spaced across said surface.
9. The method of claim 8 wherein said steps of irradiating said
tissue and detecting said pressure waveforms are performed while
said surface and said sensors are at a first position, and further
comprising the steps of moving said surface and said sensors to a
second position, repeating said irradiating step, repeating said
detecting step, and combining waveforms collected by said sensors
in said first and said second positions to generate said image of
said tissue.
10. The method of claim 9 wherein moving said surface comprises
moving said surface in a rectilinear fashion.
11. The method of claim 9 further comprising moving said
electromagnetic radiation source in synchrony with said surface and
said sensors.
12. The method of claim 9 wherein moving said surface comprises
rotating said surface.
13. The method of claim 12 wherein said sensors are positioned on
said surface along a spiral path.
14. The method of claim 1 wherein said acoustic coupling media has
an acoustic characteristic impedance which is substantially similar
to that of said tissue to reduce reflections of acoustic waves
impinging into said media from said tissue.
15. The method of claim 14 further comprising providing a flexible
film containing said acoustic coupling media, and pressing said
tissue upon said flexible film to couple acoustic waves from said
tissue into said acoustic coupling media.
16. The method of claim 14 wherein said electromagnetic coupling
media has an electromagnetic characteristic impedance which is
substantially similar to that of said tissue to reduce reflections
of electromagnetic waves impinging into said tissue from said
electromagnetic coupling media.
17. The method of claim 16 further comprising providing a flexible
film enclosing electromagnetic coupling media, and pressing said
tissue upon said flexible film to couple electromagnetic waves from
said electromagnetic coupling media into said tissue.
18. The method of claim 14 further comprising immersing said
electromagnetic radiation source in said acoustic coupling media,
wherein said acoustic coupling media has a characteristic
electromagnetic impedance which is substantially similar to that of
said tissue, to reduce reflections of electromagnetic waves
impinging into said tissue from said media.
19. The method of claim 1 wherein irradiating said tissue comprises
irradiating said tissue with a laser generating electromagnetic
radiation in the near-infrared band.
20. The method of claim 1 wherein irradiating said tissue comprises
irradiating said tissue with a Xenon flash lamp.
21. The method of claim 1 wherein irradiating said tissue comprises
irradiating said tissue with an electrically conductive coil
generating microwave frequency radiation.
22. The method of claim 21 wherein said microwave frequency is
substantially four hundred and thirty-three MHZ.
23. The method of claim 21 wherein said microwave frequency is
substantially nine hundred and fifteen MHZ.
24. Apparatus for imaging tissue structures in a three-dimensional
volume of tissue by detecting localized absorption of
electromagnetic waves in said tissue, comprising an electromagnetic
radiation source; a plurality of acoustic sensors arrayed across a
surface, said surface being acoustically coupled to said tissue via
a coupling media chosen from one or more of: water, salinated
water, alcohol, oil and mineral oil; power circuitry pulsing said
electromagnetic radiation source to produce a pulse of diffuse
electromagnetic radiation from said source irradiating said
three-dimensional volume of tissue to generate resultant pressure
waveforms within said three-dimensional volume of tissue; and
computing circuitry detecting resultant pressure waveforms arriving
at said acoustic sensors, storing data representative of said
waveforms, and combining a plurality of said detected pressure
waveforms to derive an image, points in said image being derived by
combining measures of pressure waveforms originating at points
within said tissue.
25. The apparatus of claim 24 wherein said sensors are
piezoelectric transducers having a largest dimension smaller than
four times the distance traveled by sound in tissue over the time
duration of said pulse of electromagnetic radiation.
26. The apparatus of claim 24 wherein said sensors are evenly
spaced across said surface.
27. The apparatus of claim 24 further comprising a motor coupled to
said surface for moving said surface and said sensors to generate
said image of said tissue.
28. The apparatus of claim 27 wherein said motor moves said surface
in a rectilinear fashion.
29. The apparatus of claim 28 further comprising a second motor
coupled to said electromagnetic radiation source for moving said
source in synchrony with said surface and said sensors.
30. The apparatus of claim 27 wherein said motor rotates said
surface.
31. The apparatus of claim 30 wherein said sensors are positioned
on said surface along a spiral path.
32. The apparatus of claim 24 further comprising a tank containing
said acoustic coupling media, said surface being positioned inside
of said tank and immersed in said acoustic coupling media, whereby
said acoustic coupling media has an acoustic characteristic
impedance which is substantially similar to that of said tissue to
reduce reflections of acoustic waves impinging into said media from
said tissue.
33. The apparatus of claim 32 wherein said tank includes an open
top surface whereby said tissue may be received into said acoustic
coupling media.
34. The apparatus of claim 32 wherein said tank further comprises a
flexible film cover enclosing said tank to contain said acoustic
coupling media, whereby said tissue may be pressed upon said
flexible film to couple acoustic waves into said acoustic coupling
media.
35. The apparatus of claim 32 further comprising a second tank
containing an electromagnetic coupling media, said electromagnetic
radiation source being positioned inside of second tank and
immersed in said electromagnetic coupling media, whereby said
second tank may be filled with an electromagnetic coupling media
having an electromagnetic characteristic impedance which is
substantially similar to that of said tissue to reduce reflections
of electromagnetic waves impinging into said tissue from said
electromagnetic coupling media.
36. The apparatus of claim 35 wherein said second tank further
comprises a flexible film cover enclosing said tank to contain said
electromagnetic coupling media, whereby said tissue may be pressed
upon said flexible film to couple electromagnetic waves from said
electromagnetic coupling media into said tissue.
37. The apparatus of claim 32 wherein said electromagnetic
radiation source is positioned inside of said tank and immersed in
said acoustic coupling media, whereby said acoustic coupling media
in said tank may be selected to have a characteristic
electromagnetic impedance which is substantially similar to that of
said tissue, to reduce reflections of electromagnetic waves
impinging into said tissue from said media.
38. The apparatus of claim 24 wherein said electromagnetic
radiation source is a laser.
39. The apparatus of claim 38 wherein said laser emits
electromagnetic radiation in the near-infrared band.
40. The apparatus of claim 38 wherein said laser is a Nd:YAG
laser.
41. The apparatus of claim 24 wherein said electromagnetic
radiation source is a flash lamp.
42. The apparatus of claim 41 wherein said flash lamp is a Xenon
flash lamp.
43. The apparatus of claim 24 wherein said electromagnetic
radiation source is an electrically conductive coil.
44. The apparatus of claim 43 wherein said power circuitry pulses
said coil at a microwave frequency.
45. The apparatus of claim 44 wherein said microwave frequency is
substantially four hundred and thirty-three MHZ.
46. The apparatus of claim 44 wherein said microwave frequency is
substantially nine hundred and fifteen MHZ.
47. (New) A method of imaging tissue structures by detecting
localized absorption of electromagnetic waves in said tissue,
comprising providing a source of electromagnetic radiation in
proximity to said tissue; providing a plurality of acoustic
sensors; acoustically coupling said plurality of acoustic sensors
to said tissue via a coupling media chosen from one or more of:
water, salinated water, alcohol, oil and mineral oil; irradiating
said tissue with a pulse of electromagnetic radiation from said
source to generate resultant pressure waveforms within said tissue;
detecting said resultant pressure waveforms arriving at said
acoustic sensors and storing data representative of said waveforms;
combining a plurality of said detected pressure waveforms to derive
a measure of pressure waveforms originating at a point distant from
said acoustic sensors, by determining a distance between said point
and a pressure sensor, computing a value related to a sum of the
pressure waveform detected by said pressure sensor over a time
period, said time period beginning substantially contemporaneous
with said pulse of electromagnetic radiation and said time period
having a duration equal to the time needed for sound to travel said
distance through said tissue, repeating said determining and
computing for additional pressure sensors and pressure sensor
waveforms, and accumulating said sums to form said measure of
pressure waveforms originating at said point; and repeating said
combining step for a plurality of points to produce an image of
structures in said tissue.
48. (New) The method of claim 47 wherein computing a value related
to a sum of the pressure waveform detected by a pressure sensor
over a time period, further comprises multiplying said sum by a
factor proportional to the duration of said time period used to
produce said value, whereby to compensate for diffusion of acoustic
energy radiated from said point.
49. (New) A method of imaging tissue structures by detecting
localized absorption of electromagnetic waves in said tissue,
comprising providing a source of electromagnetic radiation in
proximity to said tissue; providing a surface and positioning a
plurality of acoustic sensors spaced across said surface in a
spiral path; acoustically coupling said plurality of acoustic
sensors to said tissue via a coupling media chosen from one or more
of: water, salinated water, alcohol, oil and mineral oil;
positioning said surface and said sensors in a first position;
irradiating said tissue with a pulse of electromagnetic radiation
from said source to generate resultant pressure waveforms within
said tissue; detecting said resultant pressure waveforms arriving
at said acoustic sensors and storing data representative of said
waveforms; rotating said surface and said sensors to a second
position; repeating said irradiating step; repeating said detecting
step; combining a plurality of said detected pressure waveforms
collected by said sensors in said first and said second positions
to derive a measure of pressure waveforms originating at a point
distant from said acoustic sensors; and repeating said combining
step for a plurality of points to produce an image of structures in
said tissue.
50. (New) A method of imaging tissue structures by detecting
localized absorption of electromagnetic waves in said tissue,
comprising providing an acoustic coupling media adjacent said
tissue, having an acoustic characteristic impedance which is
substantially similar to that of said tissue to reduce reflections
of acoustic waves impinging into said media from said tissue, said
coupling media chosen from one or more of: water, salinated water,
alcohol, oil and mineral oil; providing an electromagnetic coupling
media adjacent said tissue, said electromagnetic coupling media
having an electromagnetic characteristic impedance which is
substantially similar to that of said tissue to reduce reflections
of electromagnetic waves impinging into said tissue from said
electromagnetic coupling media, said coupling media chosen from one
or more of: water, salinated water, alcohol, oil and mineral oil;
providing a source of electromagnetic radiation in proximity to
said tissue and immersing said source in said electromagnetic
coupling media to electromagnetically couple said source to said
tissue; providing a plurality of acoustic sensors and immersing
said sensors in said acoustic coupling media to acoustically couple
said plurality of acoustic sensors to said tissue; irradiating said
tissue with a pulse of electromagnetic radiation from said source
to generate resultant pressure waveforms within said tissue;
detecting said resultant pressure waveforms arriving at said
acoustic sensors and storing data representative of said waveforms;
combining a plurality of said detected pressure waveforms to derive
a measure of pressure waveforms originating at a point distant from
said acoustic sensors; and repeating said combining step for a
plurality of points to produce an image of structures in said
tissue.
51. (New) The method of claim 50 further comprising providing a
flexible film enclosing electromagnetic coupling media, and
pressing said tissue upon said flexible film to couple
electromagnetic waves from said electromagnetic coupling media into
said tissue.
52. (New) A method of imaging tissue structures by detecting
localized absorption of electromagnetic waves in said tissue,
comprising providing an coupling media adjacent said tissue, said
coupling media having an acoustic characteristic impedance which is
substantially similar to that of said tissue to reduce reflections
of acoustic waves impinging into said media from said tissue, and
having a characteristic electromagnetic impedance which is
substantially similar to that of said tissue, to reduce reflections
of electromagnetic waves impinging into said tissue from said
media, and chosen from one or more of: water, salinated water,
alcohol, oil and mineral oil; providing a source of electromagnetic
radiation in proximity to said tissue and immersing said source in
said coupling media to electromagnetically couple said source to
said tissue; providing a plurality of acoustic sensors and
immersing said sensors in said coupling media to acoustically
couple said plurality of acoustic sensors to said tissue;
irradiating said tissue with a pulse of electromagnetic radiation
from said source to generate resultant pressure waveforms within
said tissue; detecting said resultant pressure waveforms arriving
at said acoustic sensors and storing data representative of said
waveforms; combining a plurality of said detected pressure
waveforms to derive a measure of pressure waveforms originating at
a point distant from said acoustic sensors; and repeating said
combining step for a plurality of points to produce an image of
structures in said tissue.
53. (New) Apparatus for imaging tissue structures by detecting
localized absorption of electromagnetic waves in said tissue.
comprising an electromagnetic radiation source; a plurality of
acoustic sensors arrayed across a surface along a spiral path, said
surface being acoustically coupled to said tissue via a coupling
media chosen from one or more of: water, salinated water, alcohol,
oil and mineral oil; a motor coupled to said surface for rotating
said surface and said sensors; power circuitry pulsing said
electromagnetic radiation source to produce a pulse of
electromagnetic radiation from said source within said tissue; and
computing circuitry detecting resultant pressure waveforms arriving
at said acoustic sensors, storing data representative of said
waveforms, and combining a plurality of said detected pressure
waveforms to derive an image, points in said image being derived by
combining measures of pressure waveforms originating at points
within said tissue.
54. (New) Apparatus for imaging tissue structures by detecting
localized absorption of electromagnetic waves in said tissue,
comprising a first tank containing an acoustic coupling media
having an acoustic characteristic impedance which is substantially
similar to that of said tissue to reduce reflections of acoustic
waves impinging into said media from said tissue, and chosen from
one or more of: water, salinated water, alcohol, oil and mineral
oil; a second tank containing an electromagnetic coupling media
having an electromagnetic characteristic impedance which is
substantially similar to that of said tissue to reduce reflections
of electromagnetic waves impinging into said tissue from said
electromagnetic coupling media, and chosen from one or more of:
water, salinated water, alcohol, oil and mineral oil; a plurality
of acoustic sensors positioned within said first tank and immersed
in said acoustic coupling media; an electromagnetic radiation
source positioned inside of said second tank and immersed in said
electromagnetic coupling media; power circuitry pulsing said
electromagnetic radiation source to produce a pulse of
electromagnetic radiation from said source within said tissue; and
computing circuitry detecting resultant pressure waveforms arriving
at said acoustic sensors, storing data representative of said
waveforms, and combining a plurality of said detected pressure
waveforms to derive an image, points in said image being derived by
combining measures of pressure waveforms originating at points
within said tissue.
55. (New) The apparatus of claim 54 wherein said second tank
further comprises a flexible film cover enclosing said tank to
contain said electromagnetic coupling media, whereby said tissue
may be pressed upon said flexible film to couple electromagnetic
waves from said electromagnetic coupling media into said
tissue.
56. (New) Apparatus for imaging tissue structures by detecting
localized absorption of electromagnetic waves in said tissue,
comprising a tank containing a coupling media having an acoustic
characteristic impedance which is substantially similar to that of
said tissue to reduce reflections of acoustic waves impinging into
said media from said tissue, and having an electromagnetic
characteristic impedance which is substantially similar to that of
said tissue to reduce reflections of electromagnetic waves
impinging into said tissue from said coupling media, and chosen
from one or more of: water, salinated water, alcohol, oil and
mineral oil; an electromagnetic radiation source positioned inside
of said tank and immersed in said coupling media; a plurality of
acoustic sensors positioned within said tank and immersed in said
coupling media; power circuitry pulsing said electromagnetic
radiation source to produce a pulse of electromagnetic radiation
from said source within said tissue; and computing circuitry
detecting resultant pressure waveforms arriving at said acoustic
sensors, storing data representative of said waveforms, and
combining a plurality of said detected pressure waveforms to derive
an image, points in said image being derived by combining measures
of pressure waveforms originating at points within said tissue.
Description
[0001] This application is a divisional of U.S. Ser. No.
09/076,968, filed May 13, 1998, which is a divisional of P.C.T.
Application Ser. No. US97/17832, filed Oct. 1, 1997, which is a
continuation of U.S. Ser. No. 08/719,736, filed Oct. 4, 1996.
BACKGROUND OF THE INVENTION
[0002] The present invention relates to imaging properties of
tissue based upon differential absorption of electromagnetic waves
in differing tissue types by photo-acoustic techniques.
[0003] It is well established that different biologic tissues
display significantly different interactions with electromagnetic
radiation from the visible and infrared into the microwave region
of the electromagnetic spectrum. While researchers have
successfully quantified these interactions in vitro, they have met
with only limited success when attempting to localize sites of
optical interactions in vivo. Consequently, in vivo imaging of
disease at these energies has not developed into a clinically
significant diagnostic tool.
[0004] In the visible and near-infrared regions of the
electromagnetic spectrum, ubiquitous scattering of light presents
the greatest obstacle to imaging. In these regions, scattering
coefficients of 10-100 mm.sup.-1 are encountered. Consequently,
useful numbers of unscattered photons do not pass through more than
a few millimeters of tissue, and image reconstruction must rely on
multiply-scattered photons. While efforts persist to use visible
and infrared radiation for imaging through thick tissue (thicker
than a few centimeters), clinically viable imaging instrumentation
has not been forthcoming.
[0005] In the microwave region (100-3000 MHZ), the situation is
different. Scattering is not as important, since the wavelength (in
biologic tissue) at these frequencies is much greater than the
"typical" dimension of tissue inhomogeneities (.apprxeq.1 .mu.m).
However, the offsetting effects of diffraction and absorption have
forced the use of long wavelengths, limiting the spatial resolution
that can be achieved in biologic systems. At the low end of the
microwave frequency range, tissue penetration is good, but the
wavelengths are large. At the high end of this range, where
wavelengths are shorter, tissue penetration is poor. To achieve
sufficient energy transmission, microwave wavelengths of roughly
2-12 cm (in tissue) have been used. However, at such a long
wavelength, the spatial resolution that can be achieved is no
better than roughly 1/2 the microwave length, or about 1-6 cm.
[0006] In vivo imaging has also been performed using ultrasound
techniques. In this technique, an acoustic rather than
electromagnetic wave propagates through the tissue, reflecting from
tissue boundary regions where there are changes in acoustic
impedance. Typically, a piezoelectric ceramic chip is electrically
pulsed, causing the chip to mechanically oscillate at a frequency
of a few megahertz. The vibrating chip is placed in contact with
tissue, generating a narrow beam of acoustic waves in the tissue.
Reflections of this wave cause the chip to vibrate, which
vibrations are converted to detectable electrical energy, which is
recorded.
[0007] The duration in time between the original pulse and its
reflection is roughly proportional to the distance from the
piezoelectric chip to the tissue discontinuity. Furthermore, since
the ultrasonic energy is emitted in a narrow beam, the recorded
echoes identify features only along a narrow strip in the tissue.
Thus, by varying the direction of the ultrasonic pulse propagation,
multi-dimensional images can be assembled a line at a time, each
line representing the variation of acoustic properties of tissue
along the direction of propagation of one ultrasonic pulse.
[0008] For most diagnostic applications, ultrasonic techniques can
localize tissue discontinuities to within about a millimeter. Thus,
ultrasound techniques are capable of higher spatial resolution than
microwave imaging.
[0009] The photoacoustic effect was first described in 1881 by
Alexander Graham Bell and others, who studied the acoustic signals
that were produced whenever a gas in an enclosed cell is
illuminated with a periodically modulated light source. When the
light source is modulated at an audio frequency, the periodic
heating and cooling of the gas sample produced an acoustic signal
in the audible range that could be detected with a microphone.
Since that time, the photoacoustic effect has been studied
extensively and used mainly for spectroscopic analysis of gases,
liquid and solid samples.
[0010] It was first suggested that photoacoustics, also known as
thermoacoustics, could be used to interrogate living tissue in
1981, but no subsequent imaging techniques were developed. The
state of prior art of imaging of soft tissues using photoacoustic,
or thermoacoustic, interactions is best summarized in Bowen U.S.
Pat. No. 4,385,634. In this document, Bowen teaches that ultrasonic
signals can be induced in soft tissue whenever pulsed radiation is
absorbed within the tissue, and that these ultrasonic signals can
be detected by a transducer placed outside the body. Bowen derives
a relationship (Bowen's equation 21) between the pressure signals
p(z,t) induced by the photoacoustic interaction and the first time
derivative of a heating functions, S(z,t), that represents the
local heating produced by radiation absorption. Bowen teaches that
the distance between a site of radiation absorption within soft
tissue is related to the time delay between the time when the
radiation was absorbed and when the acoustic wave was detected.
[0011] Bowen discusses producing "images" indicating the
composition of a structure, and detecting pressure signals at
multiple locations, but the geometry and distribution of multiple
transducers, the means for coupling these transducers to the soft
tissue, and their geometrical relationship to the source of
radiation, are not described. Additionally, nowhere does Bowen
teach how the measured pressure signals from these multiple
locations are to be processed in order to form a 2- or
3-dimensional image of the internal structures of the soft tissue.
The only examples presented are 1-dimensional in nature, and merely
illustrate the simple relationship between delay time and distance
from transducer to absorption site.
SUMMARY OF THE INVENTION
[0012] The present invention improves upon what is disclosed by
Bowen in two ways. First, the present invention uses multiple
transducers to collect photoacoustic signals in parallel, and then
combines these signals to form an image. This approach represents a
significant advance over Bowen in that the use of multiple,
parallel transducers, substantially reduces the time needed to
collect sufficient information for imaging. Furthermore, while
Bowen fails to suggest methodologies for creating multidimensional
images, the present invention provides specific methodologies for
reconstructing multidimensional images of internal tissues through
the combination of multiple pressure recordings. As part of
achieving these advances over Bowen, the present invention details
the frequencies that might be used, the size of the multiple
transducers, their geometrical relationship to one another and to
the tissue, and structures for coupling sensors to the tissue.
[0013] Specifically, in one aspect, the invention provides a method
of imaging tissue structures by detecting localized absorption of
electromagnetic waves in the tissue. An image is formed by
irradiating the tissue with a pulse of electromagnetic radiation,
and detecting and storing resultant pressure waveforms arriving at
the acoustic sensors. Multiple detected pressure waveforms are then
combined to derive a measure of the extent to which pressure
waveforms are originating at a point distant from the acoustic
sensors. This step can then be repeated for multiple points to
produce an image of structures in the tissue.
[0014] In a disclosed particular embodiment, the multiple pressure
waveforms are combined to form an image at a point by determining a
distance between the point and a pressure sensor, and then
computing a value related to the time rate of change in the
pressure waveform, at a time which is a time delay after the pulse
of electromagnetic radiation--this time delay being equal to the
time needed for sound to travel through the tissue from the point
to the pressure sensor. This process, determining a distance and
time delay, and then computing a value for time rate of change, is
repeated for each additional pressure sensor and its pressure
waveform, and the computed values are accumulated to form the
measure of the pressure waveforms originating at the point. These
point measurements may then be collected into a multi-dimensional
image.
[0015] In one specific embodiment, the pressure sensor signal is
processed by appropriate electrical circuitry so that the
electrical output of the sensor is representative of the time rate
of change of the pressure waveform. As a result, the value
representing the time rate of change of pressure is directly
available from the sensor output. To create an appropriate output,
delayed versions of the output of the sensor are combined with the
output of the sensor, which produces an electrical output
representative of the time rate of pressure change.
[0016] In an alternative embodiment discussed below, a measure of
pressure waveforms originating at a point, is generated by
computing a value related to a sum of the pressure waveform
detected by the acoustic transducer over the time period--where
again the time period begins simultaneous with the electromagnetic
irradiating pulse, and has a duration equal to the time needed for
sound to travel through the tissue from the point to the pressure
sensor. These steps can then be repeated for additional pressure
sensors and their waveforms, and the results accumulated as
discussed above to form the measure of pressure waveforms
originating at the point.
[0017] In either approach, it is useful to multiply the computed
time rate of change, or computed time period sum, of an acoustic
transducer signal, by a factor proportional to the time delay used
to produce the value. Doing so compensates for the diffusion of
acoustic energy radiated from the point as it travels through the
tissue to the transducer.
[0018] In apparatus for carrying out these imaging methods, the
sensors are positioned on a surface and relatively evenly spaced
across the surface so as to, in combination, produce sharp
multi-dimensional images through the tissue. To reduce the number
of sensors required, the sensors may be moved to multiple positions
while producing an image. Specifically, while the sensors are in a
first position, the tissue is irradiated and the pressure waveforms
from the sensors are recorded. Then the sensors are moved to a
second position and the irradiation and waveform storage are
repeated. In this way, each sensor can be moved to a number of
positions to generate multiple waveforms. All of the stored
waveforms can then be combined to generate an image of the
tissue.
[0019] The sensors may be positioned on a plane and moved in a
rectilinear fashion, in which case the electromagnetic irradiation
source may be moved in synchrony with the sensors. Alternatively,
the sensors may be positioned on a spherical surface (having a
center of curvature approximately in the center of the tissue
region to be imaged) which is rotated to multiple positions. In
this latter case, the sensors can be advantageously positioned on
the spherical surface along a spiral path, so that rotation of the
sensors produces a relatively even distribution of sensor locations
across the spherical surface.
[0020] To enhance acoustic coupling to the tissue, the sensors may
be immersed in an acoustic coupling media, having an acoustic
characteristic impedance which is substantially similar to that of
the tissue to reduce reflections of acoustic waves impinging into
the media from the tissue. A flexible film may be used to contain
the acoustic coupling media so that the tissue can be pressed upon
the flexible film to couple acoustic waves from the tissue into the
acoustic coupling media.
[0021] Similarly, the electromagnetic radiation source may be
immersed in an electromagnetic coupling media having an
electromagnetic characteristic impedance which is substantially
similar to that of the tissue to reduce reflections of
electromagnetic waves impinging into the tissue from the
electromagnetic coupling media. Here again, a flexible film is used
to couple electromagnetic waves from the electromagnetic coupling
media into the tissue.
[0022] In one particular embodiment, both the electromagnetic
radiation source and the acoustic transducers are immersed in the
same coupling media, and the coupling media has a characteristic
acoustic and electromagnetic impedance which is substantially
similar to that of the tissue.
[0023] The electromagnetic radiation may be laser-generated
radiation in the ultraviolet, visible or near-infrared band, light
generated by a Xenon flash lamp, or microwave frequency radiation
from a microwave antenna such as a coil. In the latter case, a
microwave frequency of four hundred and thirty-three or nine
hundred and fifteen MHZ may be advantageous since these frequencies
are FCC approved and fall within a frequency band in which
malignant and normal tissue exhibit substantially different
absorptivities.
[0024] The above and other objects and advantages of the present
invention shall be made apparent from the accompanying drawings and
the description thereof.
BRIEF DESCRIPTION OF THE DRAWINGS
[0025] The accompanying drawings, which are incorporated in and
constitute a part of this specification, illustrate embodiments of
the invention and, together with a general description of the
invention given above, and the detailed description of the
embodiments given below, serve to explain the principles of the
invention.
[0026] FIG. 1 is a functional block diagram of a photoacoustic
scanner for scanning breast tissue in accordance with a first
embodiment of the present invention;
[0027] FIG. 2 is a top view of one embodiment of a transducer array
for the scanner of FIG. 1;
[0028] FIG. 3 illustrates the waveforms produced in the scanner of
FIG. 1;
[0029] FIG. 4 illustrates the spatial response of pressure
transducers used in a photoacoustic scanner such as that of FIG.
1;
[0030] FIG. 5 is a second embodiment of a photoacoustic breast
scanner in accordance with the present invention, using a laser or
flash tube source of electromagnetic energy;
[0031] FIG. 6 is an embodiment of a transducer array and
electromagnetic source for a scanner such as that of FIG. 1,
configured for rectilinear scanning motion;
[0032] FIG. 7 is an embodiment of a transducer array and
electromagnetic source for a scanner such as that of FIG. 1,
configured for rotational scanning motion;
[0033] FIG. 8 is a particular embodiment of a rotationally scanning
transducer array, formed on a spherical surface, illustrating the
positioning of the transducers on the spherical surface of the
array;
[0034] FIG. 9 illustrates the axial alignment of the transducers on
the spherical surface of the array of FIG. 8;
[0035] FIG. 10 illustrates the locus of transducer positions
brought about through rotational scanning of the array of FIG.
8;
[0036] FIGS. 11A and 11B are a third embodiment of a photoacoustic
breast scanner in accordance with the present invention, using an
acoustic coupling tank configured to permit placement of a
rotationally scanning acoustic transducer array in close proximity
to a human breast;
[0037] FIG. 12 is a circuit diagram of an integral transducer
signal amplifier for a photoacoustic breast scanner;
[0038] FIG. 13 is a fourth embodiment of a photoacoustic breast
scanner in accordance with the present invention, using an acoustic
coupling tank configured to permit a rotationally scanning acoustic
transducer array to surround a human breast;
[0039] FIG. 14 illustrates the geometric relationships involved in
the reconstruction methodologies used to generate a tissue
image;
[0040] FIG. 15 illustrates a reconstruction methodology for forming
a tissue image from acoustic transducer signals;
[0041] FIG. 16 is an experimental apparatus used to generate an
image of an absorption phantom generally in accordance with the
methodology of FIG. 15, and FIG. 17 is the image created
therefrom;
[0042] FIG. 18 illustrates a second reconstruction methodology for
forming a tissue image from acoustic transducer signals;
[0043] FIG. 19 illustrates the ideal impulse response of a
transducer which produces an electrical output signal indicative of
the first temporal derivative of an incident pressure signal;
[0044] FIG. 20 illustrates a simulated actual impulse response and
a methodology for converting this impulse response to an
approximation of the ideal response illustrated in FIG. 19; and
[0045] FIG. 21 is a circuit diagram of a circuit for performing the
conversion methodology of FIG. 20.
DETAILED DESCRIPTION OF SPECIFIC EMBODIMENTS
[0046] FIG. 1 illustrates a photoacoustic breast scanner 10 in
accordance with one embodiment of the present invention, which
displays several key elements for successful photoacoustic scanning
of the female human breast.
[0047] A human breast 12 is compressed between two coupling tanks
14, 16. Coupling tank 14 contains fluid or semi-solid media 18
having dielectric properties which are close to that of "average"
breast tissue at the microwave (or radio wave) frequencies used to
stimulate photoacoustic emission within the breast 12. Examples
would be salinated water, alcohol or mineral oil. The media 18 is
contained within tank 14 by a flexible sheet 19, for example of
polyethylene, on the surface of the tank coupled to the breast 12.
Sheet 19 ensures good mechanical contact between the tissue of
breast 12 and the media 18 in tank 14.
[0048] Within the top coupling tank is a microwave antenna 20. A
microwave generator 22, i.e., a source of pulse microwave or radio
wave energy, is coupled to antenna 20 through a transmission line
24. (One suitable microwave generator is a Hewlett-Packard model
8657B tunable generator, coupled to a 200 Watt RF amplifier
available from AMP Research.) Antenna 20 is large enough to
irradiate all or a large fraction of the breast volume to be
imaged. A cylindrically-shaped coil antenna, three to nine inches
in diameter would be suitable. Further details on waveguides which
can be used as microwave radiators can be found in Fang et al.,
"Microwave Applicators for Photoacoustic Ultrasonography", Proc.
SPIE 2708: 645-654, 1996, which is incorporated by reference herein
in its entirety.
[0049] The purpose of dielectric coupling media 18 and sheet 19 is
to improve the penetration of the microwave energy into the breast
tissue. Because breast 12 is compressed against the surface of tank
14, there is a continuous interface between coupling media 18 and
the tissue of breast 12, uninterrupted by air gaps. An air gap, or
any other physical discontinuity having a corresponding
discontinuity in dielectric properties, will cause a large fraction
of the microwave energy to reflect away from the interface (and
thus away from the surface of the breast), rather than penetrate
into the breast. By matching the dielectric properties of the
breast and media 18, and eliminating air gaps, such discontinuities
are reduced, improving microwave penetration into breast 12.
[0050] As noted above, microwave generator 22 delivers
short-duration pulses of radiation to breast 12. These bursts
should last anywhere from 10 nanoseconds to one microsecond, e.g.,
0.5 microseconds. Each radiation burst causes localized heating and
expansion of the breast tissue exposed to the microwave energy.
Tissue heating and expansion will be greatest in those regions of
the breast tissue which are most absorptive of the microwave
energy. If a region of tissue within breast 12 (e.g., a tumor) is
particularly more absorptive than the surrounding tissue, the
region will expand relatively more rapidly and extensively than the
surrounding tissue, creating an acoustic wave which will propagate
through the tissue. These acoustic waves are manifested as
longitudinal pressure waves, containing acoustic frequencies
ranging from very low frequencies to approximately the reciprocal
of the electromagnetic pulse length. For a one-half microsecond
irradiation pulse, this maximum acoustic frequency would be 2
million cycles per second, or two megaHertz (MHZ).
[0051] Any of several different microwave frequencies may be used,
but frequencies in the range of 100-1000 MHZ are likely to be
particularly effective. At these frequencies, energy penetration is
good, absorption is adequate, and differential absorption between
different types of tissue, e.g. fat and muscle, is high. It has
also been reported that the ratio of absorbed energy in cancerous
relative to normal breast tissue is enhanced in this frequency
range, peaking at 2-3 between about 300-500 MHZ. (See, e.g.,
Joines, W. T. et al, "The measured electrical properties of normal
and malignant human tissues from 50-900 MHZ", Medical Physics,
21(4):547-550, 1994.) The frequency of 433 MHZ, specifically, has
been approved by the FCC for use in hyperthermia treatments, and
accordingly is available and may be used in photoacoustic imaging
in accordance with the present invention. Imaging might also be
performed at the FCC approved frequency of 915 MHZ. Furthermore, it
has been reported that the electrical conductivity of malignant
tissue and normal tissue may vary by a factor of fifty.
Accordingly, low frequency electromagnetic radiation could also be
used to stimulate varied energy absorption and acoustic responses
in tissue.
[0052] FIG. 1 illustrates the acoustic wavefronts 26 produced by
electromagnetic irradiation of three absorptive regions 28 within
the breast 12. It will be understood that the acoustic waves
produced by regions 28 are omnidirectional; however, for clarity
only those wavefronts directed toward coupling tank 16 have been
illustrated. These acoustic waves travel through the tissue at a
velocity of sound propagation .nu..sub.s which is approximately 1.5
mm/.mu.s.
[0053] Coupling tank 16 is filled with media 29 having an acoustic
impedance and velocity of sound propagation which are close to that
of a "typical" human breast. Distilled and deionized water is an
effective media for this purpose. Media 29 is retained within tank
16 by a thin sheet 30, such as polyethylene. Breast 12 is
compressed against sheet 30, thus ensuring good mechanical coupling
from breast 12 to media 29 within tank 16, and allowing acoustic
energy to freely pass from breast 12 into tank 16. As with sheet 19
for tank 14, good mechanical coupling through sheet 30 and the
similar acoustic characteristics of breast 12 and media 29 enhances
transmission of acoustic signals out of breast 12 and into media 29
and reduces acoustic wave reflections at the surface of breast
12.
[0054] An array 32 of N acoustic transducers is located in tank 16.
Several useful array geometries are discussed herein and can be
used successfully in the embodiment of d be at least about two
inches across, and might for some applications be as large as
twelve inches across. The transducers should be evenly spaced
across the array. FIG. 2, for example, is a view illustrating an
essentially planar array 32, approximately three inches square,
bearing forty-one individual transducers 33 which can be used as
the transducer array 32 in tank 16 of FIG. 1. Other arrangements of
transducers will be discussed below.
[0055] Transducers in array 32 detect acoustic pressure waves that
are induced within the breast by the short irradiation pulse, and
travel from emission sites (e.g., regions 28) at the velocity of
sound in tissue. The transducers are fabricated so as to be most
sensitive to sonic frequencies just below the maximum frequency
stimulated by the irradiation pulse noted above.
[0056] The N transducers in array 32 are coupled through N
electronic signal lines 34 to a computer circuit 36. Computer 36 is
further connected through a control line 38 to activate microwave
generator 22 to produce a pulse of microwave energy. Following each
pulse of radiation, the time-dependent, acoustic pressure signals
recorded by each of the N transducer elements are electronically
amplified, digitized and stored within computer 36. The recorded
pressure signal from transducer i will be referenced hereafter as
p.sub.i(t).
[0057] For sufficient resolution, the pressure signals should be
digitized to a resolution of 8-12 bits at a sampling rate of at
least 5-20 MHZ, but higher resolutions and sampling rates could be
used. The amplifier should have sufficient gain so that the analog
thermal noise from the transducer is greater than 1/2 LSB of the
span of the analog-to-digital converter, or greater. Assuming the
amplifier/transducer circuit has an equivalent resistance of 50
Ohms, and the amplifier has a bandwidth of approximately 4 MHZ,
thermal noise will produce a signal magnitude of approximately 2
.mu.volts. Suitable resolution can be achieved by amplifying
transducer signals with a 5 MHZ, 54 dB preamplifier available from
Panametrics, and digitizing the amplified signals with an 8-bit, 20
MHZ sampling rate analog-to-digital converter with a .+-.0.2 volt
input span, manufactured by Gage Electronics. Additionally,
adjustable high pass filtering at 0.03, 0.1 and 0.3 can be added as
needed to achieve desired signal to noise performance.
[0058] As an example, FIG. 3 illustrates the pressure signals
p.sub.i(t) that might be produced by four hypothetical transducers
in response to pressure waves produced by a short duration of
electromagnetic irradiation of tissue. FIG. 3 shows the signal E(t)
produced by computer circuit 36 (FIG. 1) on control line 38, which
has a brief pulse, which causes microwave generator 22 to produce a
corresponding pulse of microwave energy. The resulting acoustic
signals produced within breast 12 are subsequently received by each
of the transducers, producing signals p.sub.i(t) having differing
relative magnitudes and timing, as illustrated.
[0059] It is important that the transducers be small enough so that
they are sensitive to sonic waves that impinge upon the transducers
from a wide angle. Referring to FIG. 4, three hypothetical
absorbing regions 28a, 28b and 28c are shown in greater detail,
along with the respectively corresponding wavefronts 26a, 26b and
26c emitted by these regions, toward a transducer 33. Upon
irradiation, each region 28 is the origin of an acoustic pressure
wave that travels in all directions. Part of each wave reaches
transducer 33 after a delay time.
[0060] Transducer 33 is a piezoelectric ceramic chip (or a suitable
alternative) having a cross-sectional diameter d exposed to regions
28a, 28b and 28c. Electrical contacts (not shown) attached to the
exterior of transducer 33 detect an electrical waveform produced by
the chip in response to mechanical vibration, as a result of the
piezoelectric property of the ceramic chip.
[0061] Because the acoustic energy is transmitted in a wave,
transducer 33 is not equally sensitive to the pressure waves from
the three absorptive regions. The transducer is most sensitive to
acoustic waves from region 28c, which lies on axis 40 of transducer
33 (axis 40 being defined by the direction that lies at a
90.degree. angle to the front surface of transducer 33). Transducer
33 is less sensitive to acoustic waves from region 28b because this
region is off of axis 40. Past a certain maximum angle, .theta.,
away from axis 40, transducer 33 is substantially insensitive to
pressure waves such as those from region 28a.
[0062] Maximum angle .theta. is given approximately by the
relationship sin(.theta.).apprxeq.v.sub.s.tau./d, where v.sub.s is
the velocity of sound in the relevant medium (here, tissue), .tau.
is the irradiation pulse length and d is the diameter of the
transducer. If a relatively large volume is to be imaged, then
.theta. should be as large as possible (small d), but if d is too
small, the transducer will produce a signal too weak to be
electrically detectable without excessive noise. In general, the
transducer diameter should be in the range of
.nu..sub.s.tau.<d<4.nu..sub.s.tau.. The velocity of sound in
tissue is approximately 1.5 mm/.mu.s. Thus, for a nominal pulse
width, .tau., of 1 .mu.s, d should be in the range of approximately
1.5 to 6.0 millimeters.
[0063] FIG. 5 illustrates a second embodiment of the invention
identical in structure to FIG. 1 with the exception that a pulsed
source 44 of visible or infrared radiation 46 is used to irradiate
the breast 12 instead of a microwave antenna. Also, a coupling
media may not be needed due to the close m1.064 .mu.m, pulse
width<10 nsec, 250 mJ/pulse), positioned approximately 50 mm
from the regions in the tissue to be imaged and collimated to a
25-100 mm diameter beam. Alternatively, radiation source 44 may be
a flashtube energized by a pulsing power supply, such as a xenon
flashtube and power supply from Xenon Corp., Woburn, Mass., which
can produce a radiation pulse with a 1 .mu.sec rise time, followed
by a decaying tail with a 4 .mu.sec time constant. A cylindrically
curved, reflective surface (e.g., from Aluminum foil) may be used
with the flash tube to direct radiation from the flash tube into
the breast 12.
[0064] As noted above, array 32 is preferably of a sufficient size
to image a substantial area of tissue. In some applications,
however, the tissue to be imaged may be larger than array 32.
Referring to FIG. 6, in such situations, array 32 and the radiation
source (antenna 22 or laser or flashlamp 44) may be synchronously
scanned in a rectilinear fashion as indicated by arrows 46 and 48.
At each respective position of the radiation source and array 32,
photoacoustic data is collected and used to develop a corresponding
image. The images may then be combined or superimposed to produce a
complete image of the breast 12. In this embodiment, scanning the
transducer array produces the effect of increasing the transducer
array size, and increases the angular sampling of the breast by the
transducer array.
[0065] Referring to FIG. 7, in another alternative embodiment of
the present invention, the transducer array 32 is rotated during
the data acquisition, as indicated by arrow 50. Here again, the
breast 12 is irradiated by microwave, visible or infrared radiation
from an antenna 22, or laser or flash tube 44. At each angular
position of the transducer array, photoacoustic data is collected
by the transducers and used to develop a corresponding image. The
images may then be combined or superimposed to produce a complete
image of the breast 12. In this embodiment, rotating the array 32
has the effect of increasing the effective number of transducer
elements.
[0066] FIG. 8 illustrates a specific embodiment of a rotating
spherically curved surface 52. The radius of curvature of the
surface 52 is R and the diameter of the array is D.
[0067] The position of each of the transducers in the spiral array,
relative to the center C of curvature of surface 52, can be
detailed with reference to FIG. 8. The position of each transducer
33 is given by three spherical coordinates (r,.theta.,.phi.) as is
illustrated in FIG. 8. Each of the N transducers 33 is on the
spherical surface (at a radius R), located at a unique
(.theta.,.phi.), and is oriented on the surface with its axis 40
(see FIG. 4) passing through the center C of the radius of
curvature of the spherically curved surface 52. The .phi. positions
of the transducers 33 range from a minimum angle of .phi..sub.min
to a maximum angle of .phi..sub.max. It is desirable to maximize
this range of angles, i.e., so that .phi..sub.max-.phi..sub.min is
as large as possible, since doing so will enhance the extent to
which features in the imaged tissue can be reconstructed in
multiple dimensions. (In some embodiments,
(.phi..sub.max-.phi..sub.min typically must be less than
45.degree.; however, in the embodiment of FIG. 13,
.phi..sub.max-.phi..sub.min approaches 90.degree..)
[0068] The spiral array will be rotationally stepped to each of M
positions during data acquisition, uniformly spanning
0<.theta.<360.degree.. The (.theta.,.phi.) positions of each
of the N transducers are chosen so that after scanning, the locus
of N.times.M transducer locations produced by the M rotational
steps are distributed approximately uniformly over the spherical
surface.
[0069] To accomplish uniform distribution of transducer locations
over the spherical surface of the array, the .theta.-positions of
the transducers are given as
.theta..sub.i=i.multidot.(360/N).multidot.(k+(sin
.theta..sub.min/sin .theta..sub.max)), where .theta..sub.i is the
.theta.-position of the i-th transducer (1.ltoreq.i.ltoreq.N), and
k is an arbitrary integer. The .phi.-positions of the transducers
are given recursively as
.phi..sub.i+1=.phi..sub.i+(.alpha./sin(.phi..sub.i), where .alpha.
is a constant that depends on the radius of curvature of the
spherical array and the diameter of the transducer, and
.phi..sub.i=.phi..sub.min.
[0070] Two features of the rotationally scanned, spherical-spiral
array are illustrated in FIGS. 9 and 10. FIG. 9 illustrates the
convergence of the axes 40 of the N transducers 33 to a single
point within the breast. The convergence insures that the regions
to which each of the N transducers is most sensitive (see FIG. 4)
will have a high degree of overlap, in an area 54 centered within
the tissue under study. Also evident is the wide range of angles
.phi. spanned by the transducer array.
[0071] FIG. 10 illustrates the nearly uniform distribution of the
locus of transducer locations produced by rotation of a spherically
curved surface 52 containing an array of N=32 transducers arranged
in a spiral, when stepped to 32 evenly spaced angles of rotation
.theta. in accordance with the foregoing. Referring to FIG. 10, one
position of the 32 transducer elements is shown in cross-hatching.
The remaining 31 positions of the transducers arrived at by .theta.
rotation of surface 52, are illustrated in outline. As is apparent
from FIG. 10, a nearly uniform distribution of the transducer
locations across the spherical surface is achieved.
[0072] FIGS. 11A and 11B illustrate a more specific embodiment of
the invention, incorporating a spherically curved spiral transducer
array. Tank 16 containing acoustic media is shaped to allow the
tank to be brought alongside the body 56 of a patient to be
examined. The breast 12 of the patient is compressed against the
flexible sheet 30 to facilitate acoustic imaging. A source of
radiation, either microwave, visible or infrared, is placed in
contact with the opposite side of the breast 12 to stimulate
photoacoustic waves from the breast tissue. Transducers 33 are
mounted on a spherically curved surface 52 such that their axes are
directed toward the center of the radius of curvature of the
surface 52, resulting in a large region of sensitivity overlap as
previously illustrated in FIG. 9.
[0073] The spherical array 52 is rotated by a stepper motor on a
support shaft 50 which is journalled within tank 16. A suitable
stepper motor controller (PC board) can be obtained from New
England Affiliated Technologies. The transducer array may be
formulated from a monolithic, annular array of five mm diameter
elements, arranged in a spiral pattern as discussed above.
Satisfactory results have been achieved using low-Q ceramic
transducers having a wide band frequency response from 200 kHz to 2
MHZ, falling to zero near 4 MHZ.
[0074] The annular array is encased in an aluminum-shielded housing
in which preamplifiers and line drivers are incorporated. Referring
to FIG. 12, a suitable amplifier circuit can be constructed from a
JFET 57 and bipolar transistor 59 arranged in a dual-stage
amplifier. Signals output from the integral amplifier/line drivers
are led outside of tank 16 using ultra-thin coaxial cable cables,
to an external amplifier and analog-to-digital converter.
[0075] FIG. 13 illustrates another embodiment of the invention,
specifically adapted for human breast imaging, in which the angle
.phi..sub.max-.phi..sub.min of spherically curved surface 52 is
substantially larger than in the preceding embodiment. In this
embodiment, the microwave source is a helical, "end-launch" antenna
20, for which the spherically curved, conductive surface of the
spherical transducer array 52 serves as a ground plane. Surface 52
also serves as a tank for containing an acoustic and
electromagnetic coupling media 18/29. (Distilled and deionized
water serves as a suitable acoustic/electromagnetic coupling
media.) The breast is suspended vertically into the coupling media
18/29 as illustrated, to permit coupling of both microwave energy
into the breast and acoustic energy out of the breast. The
individual transducers 33 are arranged as a spherical, spiral array
as previously described, and the surface 52 is rotated on shaft 50
to collect an even distribution of samples from the
transducers.
[0076] After sonic pressure waves are recorded using one of the
embodiments of the invention described above, photoacoustic images
must be "reconstructed" from multiple-pressure signals. The aim is
to reconstruct some property of the breast from an ensemble of
pressure measurements made externally to the breast. In this case,
these measurements are time-dependent pressure signals recorded
subsequent to object-irradiation by a short pulse of radiation.
[0077] The generalized reconstruction geometry is illustrated in
FIG. 14. The excess pressure p(r,t) that arrives at position r,
where transducer 33 is located, at time t, is the sum of the
pressure waves produced at all positions within the tissue. This
sum can be expressed as a volume integral: 1 p ( r , t ) = 4 dr ' r
- r ' 2 T ( r ' , t ' ) t 2 ( 1 )
[0078] where .rho. is the mass density and .beta. is the
coefficient of thermal expansion of the tissue, the volume integral
is carried out over the entire r'-space where the temperature
acceleration .differential..sup.2T(r',t')/.differential.t'.sup.2 is
non-zero, and where
t'=t-.vertline.r-r'.vertline./.nu..sub.s(.vertline.r-r'.vertline./.-
nu..sub.s being the time delay for an acoustic pressure wave to
propagate from position r' to position r at the speed of sound in
tissue .nu..sub.s).
[0079] Under the assumption that the radiation pulse which causes
the temperature acceleration is of a duration .tau. which is short
enough (.tau.<1 .mu.s) to generate an adiabatic expansion of
absorptive tissue, the preceding equation can be rewritten in terms
of a regional heat absorption function S(r',t): 2 p ( r , t ) = 4 C
S ( r ' , t ' ) t ' dr ' r - r ' ( 2 )
[0080] where C is the specific heat of tissue. We can further write
the heating-function as the product of a purely spatial and a
purely temporal component, i.e.,
[0081] where I.sub.0 is a scaling factor proportional to the
incident radiation intensity and R(r') represents the fractional
energy absorption of r'. Defined in this way I.sub.0T(t') describes
the irradiating field and R(r') describes the absorption properties
of the medium (breast). The excess pressure can then be written as:
3 p ( r , t ) = I 0 4 C R ( r ' ) T ( t ' ) t ' dr ' r - r ' ( 4
)
[0082] Equation 4 expresses how the time-sequential information
conveyed by the pressure signal delivers spatial information about
the absorption properties of the medium.
[0083] To further simplify, both sides of equation (4) are
integrated in time and multiplying factors are moved to the left,
to obtain: 4 4 C I 0 0 t p ( r , t " ) t " = R ( r ' ) dr ' r - r '
T ( t ' ) ( 5 )
[0084] Now, assuming that the temporal distribution of the
irradiating field is of unit height and duration .tau. (see the
function E(t) illustrated in FIG. 3), T(t') has a value of 1 only
from t'=0 to t'=.tau.. As a result, the integrand on the right side
of equation (5) will have a value of zero everywhere except along a
thin, spherical "shell" of inner radius v.sub.st surrounding point
r, where 0<t'<.tau., i.e., where
.vertline.r-r'.vertline./v.sub.s<t<.t-
au.+.vertline.r-r'.vertline./v.sub.s. This thin "shell" has a
thickness of v.sub.s.tau.; accordingly, the volume integral for
this thin "shell" can be approximated by v.sub.s.tau. multiplied by
the surface integral, over the inner surface of the "shell", i.e.,
where .vertline.r-r'.vertline./v.- sub.s=t, i.e.: 5 4 C I 0 0 t p (
r , t " ) t " v s t = r - r ' . v s R ( r ' ) dr ' r - r ' ( 6
)
[0085] Finally, noting that .vertline.r-r'.vertline.=v.sub.st, and
rearranging terms, we can define the "projection" at the position
r, S.sub.r(t), as 6 S r ( t ) = 4 Ct I 0 0 t p ( r , t " ) t " t =
r - r ' . v s R ( r ' ) dr ' ( 7 )
[0086] Equation (7) shows that the integral of all pressure waves
received at a transducer at position r and up to time t, is
proportional to the sum of the absorption function over a spherical
surface a distance v.sub.st from the transducer. Accordingly, an
image of R(r') can be reconstructed by mapping integrated pressure
data acquired at multiple transducers, over spherical surfaces (to
create three-dimensional image) or co-planar arcs (to create a
two-dimensional image).
[0087] Specifically, referring to FIG. 15, this method of image
reconstruction comprises:
[0088] 1. Positioning transducers acoustically coupled to the
tissue under study (step 60).
[0089] 2. Positioning an electromagnetic source electromagnetically
coupled to the tissue under study (step 62).
[0090] 3. Irradiating the tissue with a brief pulse of
electromagnetic energy E(t) at time t=0 to induce acoustic signals
in the tissue (step 64).
[0091] 4. Sampling and storing pressure measurements P.sub.i(t) at
each transducer i beginning at time t=0 (step 66).
[0092] 5. Computing the sums 7 S i ( t ) = t t ' = 0 t p i ( t '
)
[0093] of pressure signals (step 68).
[0094] 6. For a point r' in the tissue to be imaged, determining
the time delay t.sub.1 needed for sound to travel from point r' to
the position r.sub.1 of transducer i (step 70), selecting the value
of the sum S.sub.i(t.sub.i) (generated from transducer i) which
occurs at time t.sub.i (step 72), repeating these steps for each
transducer i (step 74), and then accumulating the selected values
S.sub.i(t.sub.i) to generate a value K(r') at position r' according
to 8 K ( r ' ) = A i = 1 N S 1 ( r i - r ' v s )
[0095] (8) (step 76).
[0096] 8. Repeating step 7 for each point r' to be imaged (step
78).
[0097] 9. Spatially filtering the resulting values of K(r') to
obtain values for R(r'). This filtering can be performed in the
frequency domain using a function having a response proportional to
the square of frequency. Alternatively, filtering may be performed
by computing the Laplacian of the three-dimensional spatial
function K(r'), i.e., R(r')=A..gradient..sup.2K(r') (9) (Step
70).
[0098] 9. Plotting the values of R(r') as an image of the tissue
(step 82).
[0099] This reconstruction methodology was generally tested for a
two-dimensional image, by constructing the simplified experimental
test bed illustrated in FIG. 16. The test bed included a wideband
transducer 82 with a center frequency of 2 MHZ, mounted on a 150 mm
arm that was rotated along a circular path 84 under stepper-motor
control. The transducer was 50 mm (height).times.6 mm (width) and
had a radius of curvature of 150 mm along the long dimension. The
transducer was asymmetrical and focused in one dimension radially
inwardly with respect to path 84; accordingly, the transducer was
most responsive to acoustic signals received over a wide angle
within the horizontal plane of circular path 84.
[0100] The scanning mechanism was immersed in a 50 ml/l
concentration Intralipid-10%, a fatty emulsion frequently used as a
tissue-mimicking scattering medium. The scattering coefficient
(.mu..sub.s) for Intralipid-10% @ 1.064 .mu.m was measured as 0.015
mm.sup.-1/ml/l. This is close to the 0.013 mm.sup.-1/ml/l reported
by van Staveren. (See van Staveren, H. J., et al., "Light
scattering in Intralipid-10% in the wavelength range of 400-1100
nm", Applied Optics, 31(1):4507-4514 (1991).) Using a value of 0.48
for the mean cosine of scatter (g), as reported by van Staveren,
and the scattering coefficient measured in our laboratory, the 50
ml/l concentration of Intralipid-10% produced a reduced scattered
coefficient .mu..sub.s'=0.39 mm.sup.-1
[.mu..sub.s'.ident.(1-g).mu..sub.s]. At this wavelength, the
absorption of Intralipid-10% is due almost entirely to the
absorption of water, .mu..sub.a.ident.0.0164 mm.sup.-19. These
values are a factor of 2-3 less than those measured in vitro for
different types of breast tissue at 900 nm.
[0101] A 50 mm diam laser beam from a pulsed Nd:YAG laser
(.lambda.=1.064 .mu.m, pulse width<10 ns, 20 Hz repetition rate,
250 mJ/pulse) illuminated the scattering medium from below. The
imaging plane of path 84 was normal to the laser beam and was
located 47.5 mm above the bottom surface of the scattering medium.
The laser beam axis and rotational axis of the transducer scanning
arm were coincident.
[0102] Data acquisition proceeded as follows: The transducer was
stepped through 360.degree. at 2.degree. increments along path 84.
At each angle, the temporal acoustic signal recorded by the
transducer was digitized to 12 bits at a rate of 10 MHZ for a total
of 1024 samples. The sampling interval was synchronized to the
pulsing of the laser. At each angle, the temporal acoustic signal
for 16 consecutive pulses were averaged. This procedure was
repeated for 180 angles.
[0103] The absorption phantom illustrated in FIG. 16 was used in
imaging. It consisted of a 4 mm diam, black, latex ball 86 and a
black, rubber cylinder 88 suspended on two, 0.35 mm diam, clear,
polyethylene threads. The dimensions for the cylinder were 8.5 mm
outside diameter 5.0 mm inside diameter and 4 mm length.
[0104] Image reconstruction proceeded using an adaptation of the
integrated, filtered-back projection algorithm described above,
applicable to a two-dimensional image. The S.sub.r(t) were computed
for each of the 180 transducer angles, backprojected over
appropriate arcs and summed. A value of .nu..sub.s=1.5 mm/.mu.s was
assumed. The next step was to apply a 2-D filter. Filtering was
performed in the frequency domain using a linear ramp function a
cosine-weighted apodizing window, i.e.,
F(.function.)=.vertline..function./.function..sub.n.vertline.*(1-co-
s((.pi..function./.function..sub.n))/2, where .function. is the
spatial frequency and .function..sub.n is the Nyquist frequency
associated with the reconstruction matrix. In this instance,
.function..sub.n=3 cycles/mm. The center 30 mm region of the
reconstruction is displayed in FIG. 17.
[0105] The basic relationship between an acoustic signal and a
heterogeneous distribution of absorbed energy is given by Equation
7. At any moment in time following an irradiating optical pulse,
the temporally weighted and temporally integrated acoustic pressure
up to that time is proportional to a surface integral of the
absorbed heat distribution R(r) within the object being imaged.
This relationship is true, provided the irradiating optical pulse
is short enough and sharp enough. This condition is met for optical
pulses less than 1 .mu.s duration.
[0106] In order to "reconstruct" R(r') from a set of acoustic
measurements, data must be sampled over at least 2.pi. steradians.
In the restricted case, where significant optical absorption takes
place within a narrow plane, R(r') can be reconstructed using a set
of co-planar acoustic data acquired over 360.degree.. The image
displayed in FIG. 17 was reconstructed under these conditions. This
image reflects what one would expect: a "cut" through the center of
a spherical and cylindrical absorber. It is of note that a "halo"
artifact surrounds the image of the latex ball 86. This originates
from the decreased velocity of sound within the latex ball (1.0
mm/.mu.s) compared to the Liposyn-10% solution (1.5 mm/.mu.s).
[0107] Were R(r') distributed throughout a larger volume, it would
have been necessary to obtain acoustic data over the surface of a
hemisphere in order to adequately reconstruct R(r'). Such an
operation can be performed by the transducer geometries described
above.
[0108] Further details on the above experimental arrangement can be
found in Kruger et al., "Photoacoustic ultrasound
(PAUS)--Reconstruction tomography", Medical Physics
22(10):1605-1609 (October 1995), incorporated by reference herein
in its entirety.
[0109] A second methodology for image generation can also be
derived from Equations (8) and (9). Specifically, it can be shown
that the Laplacian of the back-projection of the time-weighted,
integrated pressure signals is approximately equal to the
back-projection of the first time derivative of the pressure
signal, if the radius R of any imaged object is small, i.e., where
.vertline.r-r'.vertline.>>R, as follows: 9 R ( r ' ) = A i =
1 NxM t i p i ( t i ) t ( 10 )
[0110] where t.sub.i=.vertline.r.sub.i-r'.vertline./.nu..sub.s, r'
is a vector that denotes the location within the tissue, r.sub.1 is
a vector that denotes the location of transducer i, .nu..sub.s is
the velocity of sound, A is a constant, and p.sub.i(t.sub.i) is the
samples of the pressure signal that reaches the i-th
transducer.
[0111] Referring to FIG. 18, using this approximation, the steps in
the reconstruction process are as follows:
[0112] 1. Positioning transducers acoustically coupled to the
tissue under study (step 114).
[0113] 2. Positioning an electromagnetic source electromagnetically
coupled to the tissue under study (step 116).
[0114] 3. Irradiating the tissue with brief pulse of
electromagnetic energy E(t) at time t=0 to induce acoustic signals
in tissue (step 118).
[0115] 4. Sampling and storing pressure measurements p.sub.i(t) at
each transducer beginning at time t=0 (step 120).
[0116] 5. Calculating the time-weighted, first temporal derivative
of p.sub.i(t.sub.i), i.e., t.sub.i(dp.sub.i(t.sub.i)/dt), for each
of the i transducers (step 122).
[0117] 6. For each position, r', in the tissue, summing the
selected values of the time-weighted first temporal derivatives of
the pressure signals from each transducer as indicated in Equation
9 (steps 124-132).
[0118] 7. Generating an image of the tissue from computed values of
R(r') (step 134).
[0119] This reconstruction procedure produces three-dimensional
images of the energy deposition within the interior of the tissue,
which is representative of the differential absorption of the
irradiating energy by the different types of tissues within the
tissue.
[0120] To perform the above calculation, it is necessary to obtain
the first time-derivative of the pressure signal that reaches each
transducer. It should be noted, however, that a transducer produces
a characteristic "ringing" in its electrical response to an
externally-applied pulse of pressure, which distorts the shape of
the electrical output of the transducer away from that of the
pressure waveform. Referring to FIG. 19, this ringing response 136
approximates the impulse response of the transducer 33, i.e., the
electrical signal as a function of time that is produced when a
very abrupt pressure impulse 138 strikes the transducer.
[0121] If a transducer were fabricated to produce a simple biphasic
(or "doublet") response to an impulse of pressure, that is one
positive lobe, followed a short time later by one negative lobe (an
ideal response 136 is illustrated in FIG. 19), then the electrical
output of the transducer would be approximately proportional to the
first time-derivative of the input pressure signal. This would be
desirable, because it would eliminate the necessity of computing
the first time-derivative of the input pressure signal; rather, the
time derivative would be produced by the transducer in the first
instance.
[0122] For any real transducer, however, such a response would be
difficult to achieve. Rather, the impulse response of a transducer
is closer to a damped sinusoid, as is illustrated in waveform 140
(p(t)) in FIG. 20. In this example, the impulse response of the
transducer is assumed to be of the form
p(t)=sin(2.pi..function.t)e.sup.aft. Such a response displays a
periodic component of a characteristic temporal frequency
.function., that decays exponentially with time.
[0123] In this case, an approximate "differential" transducer
response can by synthesized by delaying the originally recorded
pressure waveform, p(t), by varying amounts, weighing the delayed
pressure signals, and summing the delayed pressure signals together
with the original waveform. An example is illustrated in FIG. 20,
which shows two weighted, time-delayed waveforms (Ap(t-.DELTA.t)
142 and Bp(t-2.DELTA.t) 144 (where .DELTA.t is 1/2f) generated from
the assumed impulse response 140 of the transducer. When the
time-delayed waveforms 142 and 144 are added to the response 140 of
the transducer, the resulting waveform 146 synthesizes a biphasic
impulse response S(t).
[0124] Thus, to implement the reconstruction algorithm described
above, the transducer responses can be synthesized to be
differential in nature using the methodology illustrated in FIG.
20, after which the output of each transducer will be proportional
to dp(t)/dt.
[0125] Referring to FIG. 21, a circuit for performing such a
reconstruction includes an analog-to-digital converter 148 for
converting the analog signal from the transducer to an equivalent
digital signal, an amplifier 149 and cache 150 for receiving and
temporarily storing samples from A/D converter 148 and outputting
the sample which was stored .DELTA.t earlier multiplied by a gain
factor A, a second amplifier 151 and cache 152 for storing samples
and outputting the sample which was stored 2.DELTA.t earlier
multiplied by a gain factor B, and a digital accumulator 154 for
summing the outputs of caches 148 and 150 with the current sample
from the A/D converter to produce an output digital signal S which
is representative of dp(t)/dt.
[0126] Using a circuit such as that shown in FIG. 21, steps 120 and
122 of the reconstruction process described by FIG. 18 can be
accomplished in a single operation by hardware rather than in
software computations, increasing the scanning and imaging rate of
the apparatus.
[0127] While the present invention has been illustrated by a
description of various embodiments and while these embodiments have
been described in considerable detail, it is not the intention of
the applicants to restrict or in any way limit the scope of the
appended claims to such detail. Additional advantages and
modifications will readily appear to those skilled in the art. The
invention in its broader aspects is therefore not limited to the
specific details, representative apparatus and method, and
illustrative example shown and described. Accordingly, departures
may be made from such details without departing from the spirit or
scope of applicant's general inventive concept.
* * * * *