U.S. patent application number 09/922659 was filed with the patent office on 2002-10-03 for acousto-optic monitoring and imaging in a depth sensitive manner.
Invention is credited to Fee, Michale Sean, Schnitzer, Mark J..
Application Number | 20020140942 09/922659 |
Document ID | / |
Family ID | 26953782 |
Filed Date | 2002-10-03 |
United States Patent
Application |
20020140942 |
Kind Code |
A1 |
Fee, Michale Sean ; et
al. |
October 3, 2002 |
Acousto-optic monitoring and imaging in a depth sensitive
manner
Abstract
A system monitors or images a portion of a sample. The system
includes an optical interferometer with a measurement arm, a
reference arm, and an optical splitter. The arms are coupled to
receive light from the optical splitter. One of the arms includes
an acousto-optical modulator. The interferometer is configured to
interfere light output from the two arms. The system also includes
a detector that receives the interfered light and uses the received
light to determine a depth-dependent quantity characterizing a
portion of the interior of the sample.
Inventors: |
Fee, Michale Sean; (New
Vernon, NJ) ; Schnitzer, Mark J.; (Summit,
NJ) |
Correspondence
Address: |
Docket Administrator
Lucent Technologies Inc.
600 Mountain Avenue
P.O. Box 636
Murray Hill
NJ
07974-0636
US
|
Family ID: |
26953782 |
Appl. No.: |
09/922659 |
Filed: |
August 6, 2001 |
Related U.S. Patent Documents
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Application
Number |
Filing Date |
Patent Number |
|
|
60269586 |
Feb 17, 2001 |
|
|
|
Current U.S.
Class: |
356/477 |
Current CPC
Class: |
G02B 6/262 20130101;
A61B 5/0084 20130101; A61B 5/0086 20130101; G02B 6/3512 20130101;
G02B 6/32 20130101; A61B 5/0066 20130101; G02B 6/264 20130101; G02B
6/14 20130101 |
Class at
Publication: |
356/477 |
International
Class: |
G01B 009/02 |
Claims
What is claimed is:
1. A system for monitoring or imaging a sample, comprising: an
optical interferometer comprising a measurement arm, a reference
arm, and an optical splitter, the arms being coupled to receive
light from the optical splitter, the interferometer being
configured to interfere light outputted by the arms, one of the
arms having an acousto-optical modulator; and a detector configured
to receive the interfered light and to use the received light to
determine a depth-dependent quantity characterizing a portion of
the interior of the sample.
2. The system of claim 1, wherein the quantity is representative of
one of a signed displacement and a velocity of the portion of the
interior of the sample
3. The system of claim 1, wherein the quantity is representative of
a signed displacement of the portion of the interior of the
sample.
4. The system of claim 2, further comprising: an optical source
coupled to transmit light to the measurement and reference arms and
capable of producing light with a coherence length of less than 1
centimeter.
5. The system of claim 2, further comprising: an optical source
coupled to transmit light to the measurement and reference arms and
capable of producing light with a coherence length of less than 1
millimeter.
6. The system of claim 2, where in one of the reference arm and the
measurement arm has a variable optical path length.
7. A system for medical monitoring or imaging of a patient or
animal, comprising: an optical interferometer having a measurement
arm, a reference arm, and an optical splitter, the arms being
coupled to receive light from the optical splitter and configured
to cause light outputted by the arms to interfere; and an
interference detector coupled to receive a portion of the
interfering light and configured to determine information
representative of a location, an orientation, or a velocity of a
portion of the patient or animal from the received light; and a
controller coupled to receive the information and to adjust
collected data on the animal or patient in a manner responsive to a
change in a relative location, orientation, or velocity between a
probe and a portion of the interior of a tissue in the animal or
patient.
8. The system of claim 7, wherein the controller is configured to
adjust collected image data to correct the image data for motion of
the interior of the tissue in the animal or patient.
9. The system of claim 7, wherein the controller is configured to
control the position of the probe.
10. The system of claim 7, wherein one of the reference arm and the
measurement arm has a variable optical path length.
11. The system of claim 7, wherein one of the reference arm and the
measurement arm includes an acousto-optical modulator.
12. The system of claim 7, wherein the detector is configured to
determine one of a velocity and a signed displacement of the
portion of the interior of the tissue based on the received
interfering light.
13. The system of claim 7, wherein the measurement arm includes an
optical endoscope for sending light to and receiving light from the
portion of the interior of the tissue of the patient or animal.
14. The system of claim 13, wherein the endoscope includes an
optical fiber configured to perform the sending and receiving.
15. A process for monitoring a sample, comprising: transmitting
light to measurement and reference arms of an interferometer;
acousto-optically frequency shifting light in one of the reference
arm and the measurement arm; collecting light from the measurement
arm in response to the light scattering off a portion of the
interior of the sample; and interfering light from the reference
arm with the collected light.
16. The process of claim 15, further comprising: determining one of
a velocity and a signed displacement of the portion of the interior
of the sample based on a measurement of the interfering light.
17. The process of claim 15, wherein the transmitted light has a
coherence length of less than 1 centimeter.
Description
[0001] This application claims the benefit of U.S. Provisional
Application No. 60/269,586, filed Feb. 17, 2001.
BACKGROUND OF THE INVENTION
[0002] 1. Field of the Invention
[0003] This invention relates to optical monitoring and
imaging.
[0004] 2. Discussion of the Related Art
[0005] Contemporary medical technology uses x-rays, sound waves,
and visible light to produce in vivo images of biological tissues.
Visible light and infrared (IR) light imaging has a better
potential resolution than imaging produced by sound waves, because
visible and IR light has a shorter wavelength than sound waves. In
spite of this advantage of visible and IR light, in vivo imaging
systems often use sound waves, because visible and IR light does
not penetrate thick tissues. Consequently, many in vivo imaging
systems do not have the image resolutions obtainable in imaging
systems based on visible or IR light.
[0006] The resolution of in vivo examination systems is also
limited by tissue motion. For many organisms internal tissue
motions are always present, and these tissue motions interfere with
tissue examinations that require more observation time than the
time scale associated with the internal tissue motions. The
internal tissue motions can cause image scans to produce multiple
or smeared images. The internal tissue motions also cause the
displacement of probes placed in the tissues. These displacements
cause the probes to measure smeared out tissue properties.
BRIEF SUMMARY OF THE INVENTION
[0007] Various embodiments include systems that process optical
imagining data to obtain information on the position and/or
velocity of a region inside a sample being imaged. In some
embodiments, the position and/or velocity data is used to
compensate for tissue motions during in vivo examinations of living
organisms. The compensation corrects for smearing of scan images or
measured properties that would otherwise result due to relative
motions between the tissue and a monitoring probe.
[0008] One system according to principles of the invention images a
portion of a sample. The system includes an optical interferometer
with a measurement arm, a reference arm, and an optical splitter.
The arms are coupled to receive light from the optical splitter.
One of the arms includes an acousto-optical modulator. The
interferometer is configured to interfere light output from the two
arms. The system also includes a detector that receives the
interfered light and uses the received light to determine a
depth-dependent quantity characterizing a portion of the interior
of the sample.
[0009] In some such systems, the detector uses the interfered light
to determine a signed displacement or velocity of a portion of the
sample. Such a system monitors body tissue movements and may
further include a controller that receives displacement or velocity
information. The controller adjusts data collection from the tissue
in a manner that is responsive to changes in the tissue's location,
orientation, or velocity relative to a monitoring probe.
BRIEF DESCRIPTION OF THE FIGURES
[0010] FIG. 1A shows a system that optically monitors or images a
sample;
[0011] FIG. 1B is a flow chart for a process that uses the system
of FIG. 1A;
[0012] FIG. 2 shows a graded index (GRIN) fiber-size lens used in
some embodiments of the probe of FIG. 1A;
[0013] FIG. 3A shows a conventional GRIN fiber lens;
[0014] FIG. 3B shows a refractive index profile for one embodiment
of the GRIN fiber-size lens of FIG. 2;
[0015] FIGS. 4A and 4B show embodiments of the system of FIG. 1A
that incorporate optical interferometers;
[0016] FIG. 4C illustrates an acousto-optical modulator used in the
interferometers of FIGS. 4A and 4B; and
[0017] FIG. 5 shows a medical diagnostic system based on the system
of FIGS. 4A and 4C.
DETAILED DESCRIPTION OF THE EMBODIMENTS
[0018] 1. Optical Micro-Probe and Imagining System
[0019] A co-pending patent application describes optical
micro-probes and systems used in some embodiments of the invention
of the present application.
[0020] FIG. 1A shows a system 10 for optically monitoring or
imaging a region of a sample 12, e.g., for endoscopic viewing of a
biological tissue. Various embodiments of the system 10 determine
the velocity and/or three-dimensional position of the region being
monitored or imaged, e.g., via tomography. Such monitoring or
imaging functions are useful for medical diagnostics and treatment,
e.g., invasive imaging of anomalous tissue structures in vivo and
monitoring of tissue motion during other medical procedures.
[0021] The system 10 includes a source 14 of IR, visible, or
ultraviolet light, an optical splitter or circulator 16, an optical
micro-probe 18, and a light detector 20. Exemplary sources 14
include monochromatic sources or multi-chromatic sources, e.g., a
pulsed Ti-sapphire laser with a low coherence time of about
10.sup.-15-10.sup.-13 seconds. The optical splitter or circulator
16 directs a portion of the light from the source 14 to the optical
micro-probe 18. The optical micro-probe 18 has a distal end 22
located either above or below the surface 23 of the remote sample
12. The optical micro-probe 18 delivers the source light to a
region of the sample 12. The optical micro-probe 18 also returns to
the splitter or circulator 16 a portion of the light scattered or
emitted by the region of the sample 12 illuminated by the optical
micro-probe 18. The optical splitter or circulator 16 redirects the
returned light to detector 20. The detector 20 uses the returned
light to determine a scattering or emission characteristic of the
region of the sample 12 that produced the light. Some detectors 20
are configured to determine the distance of the region from the
optical micro-probe 18 and/or the velocity of the region.
[0022] FIG. 2 shows one embodiment 18' of optical micro-probe 18
shown in FIG. 1A. The optical micro-probe 18' includes a
single-mode optical fiber 24 that transports light to and from the
sample 12. The distal end 22 of the fiber 24 is fused to a GRIN
fiber-size lens 26, which has the same outer diameter as the
optical fiber 24. In some embodiments, the GRIN fiber-size lens 26
also has a rounded end face 28 that facilitates insertion of the
end 22 of the optical micro-probe 18' into samples such as
biological tissues. In some embodiments, a portion of the GRIN
fiber-size lens 26 adjacent the end face 28 has a conical taper
(not shown). The taper also facilitates insertion of the optical
micro-probe 18 into sample 12, i.e., the taper functions like a
needle point.
[0023] The GRIN fiber-size lens 26 collimates light from fiber 24
into a focused beam 30. The collimated beam 30 illuminates a region
of the sample 12 located forward of the lens 26. Points 32 in the
illuminated region scatter or emit light in response to being
illuminated. The backscattered or emitted light is useable for
imaging or monitoring.
[0024] The beam collimation enables resolving transverse locations
of the points 32 with respect to the axis of the GRIN fiber-sized
lens 26, because points 32 producing backscattered or emitted light
are located within the region illuminated by the beam 30.
[0025] In some embodiments, a mechanical driver (not shown) drives
the distal end 22 of optical micro-probe 18 to execute scanning
motions parallel and/or transverse to the axis of the GRIN
fiber-size lens 26. These scanning motions enable system 10 to
collect optical data for two-dimensional or three-dimensional
images of the sample 12, i.e., a planar or full 3D image.
[0026] Illumination beam 30 has a width that varies with distance
from the end surface 28 of the GRIN fiber-size lens 26. The beam
width has a minimum value at an approximate focal point 34 of the
GRIN fiber-size lens 26, i.e., at a distance "f" from end face 28.
Typically, the distance "f" has a value from about 0.2 millimeters
(mm) to about 1.5 mm, and exemplary values of "f" are greater than
about 0.8 mm. The beam 30 has a divergence that is characterized by
a rayleigh range "z". Herein, the rayleigh range is half the length
of the portion of the beam 30 that has a width less than about
{square root}2 times the minimum width at the approximate focal
point 34. An exemplary GRIN fiber-size lens 26 has a rayleigh range
greater than about 200 microns (.mu.), e.g., z.gtoreq.300.mu. or 8
mm.gtoreq.z.gtoreq.300.mu..
[0027] The focal distance and rayleigh range of GRIN fiber-size
lens 26 depend on the radial profile of the refractive index in the
GRIN lens and on the length of the GRIN lens. GRIN fiber-size lens
26 is either a conventional GRIN fiber-size lens or a new GRIN
fiber-size lens with a gentler refractive index profile.
[0028] Conventional GRIN fiber lenses are described in U.S. Pat.
No. 4,701,011, which is incorporated herein by reference in its
entirety. FIG. 3A shows the radial refractive index profile of one
such GRIN fiber lens. The refractive index is constant over a range
of values of the radius that correspond to the fiber's outer
cladding and varies over values of the radius that correspond to
the fiber's core. Restricting the refractive index variations to
the core typically produces a GRIN fiber lens with a short focal
length, less than about 0.7 mm, and a short rayleigh range, e.g.,
less than 200.mu..
[0029] FIG. 3B shows a radial refractive index profile of a new
GRIN fiber-size lens 26 for which the profile's radial curvature is
smaller in magnitude than in conventional GRIN fiber-size lenses.
The smaller magnitude curvature causes the new GRIN fiber-size lens
to have a longer focal length than the conventional GRIN fiber lens
associated with the profile of FIG. 3A. The new GRIN fiber-size
lenses are described in co-pending U.S. patent application Ser. No.
09/896,789, filed Jun. 29, 2001, which is incorporated herein by
reference in its entirety.
[0030] In the profile of FIG. 3B, the refractive index varies over
the whole diameter of the lens. Thus, the new GRIN fiber-size lens
has no outer cladding. The absence of cladding increases the radial
range over which the refractive index varies permitting a larger
optical mode, which results in the associated GRIN fiber-size lens
having a longer rayleigh range than the GRIN fiber-size lens
associated with the profile FIG. 3A.
[0031] Refractive index profiles are characterized by a parameter
"g" that measures the radial curvature of the profile in the core
of a GRIN fiber lens. In particular, the parameter g is defined as:
1 g = - 1 n 0 2 P ( r ) r 2 r = 0
[0032] Here, "r" is radial distance for the axis of the GRIN fiber
lens, no is the value of the refractive index on the axis of the
GRIN fiber lens, and P(r) is the value of the refractive index at
the distance "r" from the axis of the fiber lens.
[0033] Exemplary new GRIN fiber-size lenses have refractive index
profiles whose radial curvatures are smaller in magnitude than
those disclosed in Table 1 of "Analysis and Evaluation of
Graded-Index Fiber-Lenses", Journal of Lightwave Technology, Vol.
LT-5, No. 9 (September 1987), pages 1156-1164, by W. L. Emkey et
al, which is incorporated by reference herein in its entirety. The
new GRIN fiber-size lenses 26 have a "g" that is less than
1.7.times.10.sup.-6 .mu.m.sup.-2, preferable less than about
0.9.times.10.sup.-6 .mu.m.sup.2 and more preferably less than about
5.0.times.10.sup.-7 .mu.m.sup.-2. For 125 .mu.m--diameter GRIN
fiber lenses 18, values of "g" are selected from the range
1.7.times.10.sup.-6 .mu.m.sup.-2 to 5.0.times.10.sup.-7
.mu.m.sup.-2 and preferably in the range 0.9.times.10.sup.-6
.mu.m.sup.-2 to 5.0.times.10.sup.-7 .mu.m.sup.-2 to provide good
beam collimation.
[0034] Referring again to FIG. 2, the above-disclosed refractive
index profiles produce focal lengths and rayleigh ranges for GRIN
fiber-size lens 26 that are consistent with the above-recited
values. Some embodiments of optical micro-probe 18' use a GRIN
fiber-size lens 26 with a profile similar to that of FIG. 3B,
because such a profile provides a longer rayleigh range. The longer
rayleigh range provides a larger usable depth range for sample
probing. Typically, the usable depth of the optical micro-probe 18'
is about 1 to 8 rayleigh ranges from the focal point 34.
[0035] FIG. 1B is a flow chart for a process 40 that uses system 10
of FIGS. 1A and 2. The process 40 includes positioning distal end
22 of the optical micro-probe to monitor a selected portion of
sample 12 (step 42). The positioning includes selecting an
orientation of the optical micro-probe 18 with respect to the
sample surface 23 and selecting a lateral position and depth for
the distal end 22 with respect to the sample surface 23. After
positioning the optical micro-probe 18, source 14 transmits source
light to the optical micro-probe 18 via splitter or circulator 16
(step 44). The transmitted source light passes through GRIN
fiber-size lens 26, which focuses the light into beam 30 (step 46).
The region illuminated by the beam 30 produces the scattered or
emitted light. The GRIN fiber-size lens 26 collects a portion of
the light that is scattered or emitted by the illuminated region of
the sample (step 48). The optical micro-probe 18 returns the light
collected by the GRIN fiber-size lens 26 to the optical splitter or
circulator 16, which redirects a portion of the returned light to
optical detector 20 (step 50). The detector 20 determines the
scattering or emission characteristics of the region of the sample
12 from the light redirected thereto (step 52). Since the beam 30
has an intensity that varies with the beam width, the detector 20
primarily receives light from a region of the sample 12 that has a
volume limited by the boundary of the beam 30. The volume includes
sample points within about 1 to 8 rayleigh ranges of focal point
34. The light from the sample points and known position and
orientation of optical micro-probe 18 enable using data from
detector 20 to determine lateral positions and depths of the sample
points backscattering or emitting light in some embodiments of
system 10.
[0036] 2. Interferometric Optical Monitoring and Imaging
[0037] Various embodiments according to principles of the invention
of this application include monitoring and imagining systems that
determine depth and/or velocity data for a region of a sample that
backscatters or emits light delivered by an optical micro-probe.
Exemplary micro-probes include both probe 18' with attached GRIN
fiber-sized lens 26, as shown in FIG. 2, and a single mode optical
fiber without attached terminal GRIN fiber-sized lenses. FIGS. 4A
and 4B show two such embodiments 60, 60'. To determine depths of
sample regions, the systems 60, 60' use "low-coherence
interferometry" a method known to those of skill in the art.
[0038] Each system 60, 60' includes an interferometer with a
measurement arm 62 and a reference arm 64. The two arms 62, 64
receive light from a multi-chromatic source 66, i.e., a
low-temporal coherence source. Typically, source 66 is spatially
coherent. The measurement arm 62 outputs light scattered by sample
points in response to being illuminated by source light.
[0039] Each system 60, 60' interferometrically combines the light
output by measurement arm 62 and reference arm 64. The combined
light provides an output signal sensitive to optical path
differences between the two arms 62, 64. Because of the
low-coherence nature of source 66, depth resolution is provided by
the rapid fall off of the amplitude of the interference signal with
increasing path length difference. Interferometric combining of
light requires that the difference in the optical path lengths
traversed by the light being combined, e.g., the path difference
between the two arms 62, 64, be less than the light's coherence
length, e.g., the coherence length of the source 66. Interference
detector 74 uses the interferometrically combined light to
determine one or more characteristics of the region of the sample
12 that produced scattered light, e.g., the intensity of the light
scattered back into the optical micro-probe 18. Thus, the
sensitivity to optical path differences makes detector 74 sensitive
to the depth of sample points 32 producing scattered light. The
detector 74 is only sensitive to light produced by sample points 32
that are located within the sample depth range for which the
optical path length difference between the measurement and
reference arms 62, 64 is less than the coherence length of the
source 66.
[0040] To increase depth resolution, a less coherent source 66,
e.g., a pulsed Ti-sapphire laser, is used in systems 60, 60'. The
source 66 has a coherence length that is at least less than one
centimeter and typically is less than one millimeter. In some
embodiments, the source 66 has a coherence length that is as small
as 100 microns or even 1 micron. Since interferometric combination
only occurs if some optical path length differences between the
measurement and reference arms 62, 64 are less than about one
coherence length, this condition defines the depth resolution of
the system 10. For a sample depth resolution of 10 microns, the
source 14 should produce an output beam that is only coherent for a
time equal to about 10.sup.-5 meters/{3.times.10.sup.8
meters/second} =3.times.10.sup.-14 seconds.
[0041] The systems 60 and 60' of FIGS. 4A and 4B include a
Michelson interferometer and a Mach-Zehnder interferometer,
respectively. Each system 60, 60' has an optical splitter/combiner
68 that couples to one end of the measurement and reference arms
62, 64. The optical splitter/combiner 68 transmits mutually
coherent light from low-coherence source 66 to the measurement and
reference arms 62, 64. The measurement arm 62 includes optical
micro-probe 18. In the system 60' of FIG. 4B, the probe 18 connects
to the measurement arm 62 through an optical circulator 65. The
probe 18 illuminates a sample region with source light and also
collects light scattered produced by the illuminated sample region.
In some embodiments, the optical micro-probe is a single-mode fiber
24 having a GRIN fiber-size lens 26 fused to its distal end 22. The
reference arm 64 includes a moveable reflector 76, e.g. a moving
mirror, and an acousto-optical modulator (AOM) 70. The moveable
reflector allows an operator to change the optical path length of
the reference arm 64, i.e., to scan different sample depths by
moving the reflector 76. The AOM 70 acoustically frequency shifts
the source light received from the splitter/combiner 68 and enables
velocities of sample points 34 to be measured (see below). Some
embodiments include dispersion compensator 72 that corrects
differences in chromatic dispersion or pulse broadening between
light propagating in the measurement and reference arms 62, 64. The
construction of dispersion compensators is known to those of skill
in the art.
[0042] The interference detector 74 receives frequency-shifted
light from the reference arm 64 and light scattered by the sample
from the measurement arm 62. The arms 62, 64 have optical path
lengths that are equal to within about one coherence length of
source 66 so that some light from the two arms 62, 64
interferometrically combines in the detector 74, i.e., light
produced by scattering at some sample depth. The detector 74
determines characteristics of regions of the sample producing light
that interferometrically combines with light from the reference arm
64. The moving reflector 76 enables an operator to adjust the
optical path length difference between the reference and
measurement arms 64, 62 so that sample depths can be scanned by the
interference detector 74. Through such scans, the systems 60, 60'
are able to generate images of the sample 12 as a function of
sample depth.
[0043] FIG. 4B also shows an exemplary interference detector 74.
The exemplary interference detector 74 includes a 50/50 optical
splitter/combiner 73 that produces signals with a 180.degree. phase
difference on its two output terminals. From the 50/50 optical
splitter/combiner 73, the 180.degree. out of phase optical signals
go to separate intensity detectors 75. Outputs of the intensity
detectors 73 couple to the inputs of a differential amplifier 77
whose output signal is representative of optical interference
between signals from the reference and measurement arms 64, 62.
[0044] Referring to FIG. 4C, AOM 70 includes a radio frequency (RF)
source 78 and an optical medium 80. The RF source 78 excites sound
waves, i.e., phonons, in the optical medium 80. The sound waves are
directed along direction "Y" and have the source's RF frequency. A
voltage oscillator 81 drives the RF source 78. In some embodiments,
the oscillator 81 is variable so that the phonon frequency is
variable.
[0045] Referring to FIGS. 4A-4C, a photon in reference arm 64 may
absorb or emit a phonon while propagating through the optical
medium 80. Absorption or emission of a phonon produces both a
frequency-shift, i.e., .+-..DELTA./h, and a direction-change for
the photon. Thus, the acoustically-driven medium 80 produces both
directionally undeviated output light, i.e., photons that have
neither absorbed nor emitted a phonon, and directionally deviated
output light, i.e., photons that have absorbed or emitted a phonon.
Momentum conservation fixes the directions of the deviated output
light to be different from the direction of the undeviated output
light. The AOM 70 is configured to deliver deviated output light of
one frequency to the detector 74 and to not deliver the undeviated
output light to the detector 74. One embodiment screens out the
undeviated light by imaging only deviated light, which has a new
propagation direction, on an optical fiber that delivers light to
moving reflector 72. Thus, the interference detector 74 receives
light whose frequency has been shifted by absorption or emission of
a phonon in the AOM 70.
[0046] In some embodiments, the light makes two passes through the
AOM 70, and the AOM 70 screens out light whose frequency has not
been shifted by the absorption or emission of two phonons. Then,
this AOM 70 produces light whose frequency is shifted with respect
to the optical source 66 by twice the frequency of the RF source
78.
[0047] Referring again to FIGS. 4A-4B, the detector 74 obtains
information on the displacement or velocity of the region of the
sample 12 that backscatters source. The displacement or velocity
information is encoded in the size of the Doppler shift caused by
the velocity of the scattering region of the sample 12. The AOM 70
enables detection of such Doppler shifts through phase-sensitive
detection, which are known to those of skill in the art. In some
embodiments, this detection technique enables a determination of
both the sign and the magnitude of the velocities of scattering
sample particles along the axis of optical micro-probe 18. In other
embodiments, this detection technique enables a determination of
both the sign and the magnitude of displacements of scattering
sample particles along the axis of optical micro-probe 18.
[0048] The AOM 70 provides light outputted by the reference arm 64
with a different frequency from the frequency of light outputted by
the measurement arm 62. In the absence of sample motion, this
frequency difference is equal to the frequency of the RF energy
driving the AOM 70, i.e., if the reference arm 64 produces photons
that absorb or emit one phonon. Sample motion at the depth for
which the path difference between the measurement and reference
arms 62, 64 vanishes changes the frequency difference between the
light from the two arms 62, 64, i.e., due to Doppler shifting. The
detector 74 uses the magnitude of the change in the frequency
difference between the light from the two arms 62, 64 to determine
the speed of a sample particle producing scattering. The detector
74 uses the phase of the change in frequency difference, i.e.,
positive or negative, to determine the sign of the sample motion,
i.e., towards or away from the optical micro-probe 18. Standard
electronic or optical techniques are known for determining both the
magnitude and sign of the frequency difference between the light
from the two arms 62, 64.
[0049] The systems 60, 60' use the AOM 70 to determine information
representative of velocities of sample points at a selected sample
depth. Information representative of velocities of sample points
includes signed displacements and velocities of the sample points
along the axis of probe 18. The systems 60, 60' are also able to
select different optical path lengths for the reference arm 64,
i.e., by moving reflector 76. By scanning such optical path
lengths, systems 60, 60' are able to select different sample depths
for which interferometric combination of scattered light from the
measurement arm 62 and light from the reference arm 64 occurs.
During such a scan, detector 74 determines sample region velocities
as a function of distance from end 22 of optical micro-probe 18,
i.e., as a function of sample depth. This type of scan of sample
velocities as a function of depth enables, e.g., for mapping blood
flow rates in an artery of an animal or patient.
[0050] The AOM 70 shifts light in the reference arm 64 by a single
frequency. This simple form of the frequency shift enables the
detector 74 to determine velocities in the sample 12.
[0051] FIG. 5 shows a medical diagnostic system 90 based on system
60 of FIGS. 4A and 4C. The system 90 includes a differential
amplifier 91 and an electronic filtering chain 92 that amplify and
remove input noise, respectively. The system 90 also includes a
multiplier 93 that combines a signal representative of the
interferometrically combined optical signals from the measurement
and reference arms 62, 64 with a signal representative of the RF
signal driving RF source 78. The output of the multiplier 93 goes
to a fringe counter 94 that determines both the magnitude and sign
of the velocity of a monitored portion of the sample. To determine
the sign of the velocity, i.e., towards or away from optical
micro-probe 18, the counter 94 compares the signal from the
multiplier 93 when the multiplier receives, i.e., via line 95,
different signals representative of the RF signal driving source
78. The different signals are out of phase by 90.degree..
[0052] The fringe counter 94 couples to a feed forward circuit 96
that in turn transmits information on the velocity and/or position
of sample 12 to a controller 97. The controller 97 is connected to
a second diagnostic probe 98, e.g., a monitoring electrode or a
scanner for the same sample 12. The controller 97 uses the
information fed forward by circuit 96 to correct data that is
output by the probe 98 for the effects of sample motion. In some
embodiments, the controller 97 mechanically adjusts the position of
the second diagnostic probe 98 to eliminate relative motion between
the sample and probe 98. In other embodiments, the controller 97
corrects the data collected by the second diagnostic probe 98 to
compensate for the motion of the sample 12, e.g., by displacing
image scan data to eliminate motion induced smearing.
[0053] Other embodiments of the invention will be apparent to those
skilled in the art in light of the specification, drawings, and
claims of this application.
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