U.S. patent application number 12/589636 was filed with the patent office on 2010-07-08 for detection, imaging and characterization of brain tissue.
Invention is credited to Britton Chance.
Application Number | 20100174160 12/589636 |
Document ID | / |
Family ID | 26755485 |
Filed Date | 2010-07-08 |
United States Patent
Application |
20100174160 |
Kind Code |
A1 |
Chance; Britton |
July 8, 2010 |
Detection, imaging and characterization of brain tissue
Abstract
An optical examination technique employs an optical system for
in vivo non-invasive transcranial examination of brain tissue of a
subject. The optical system includes an optical module arranged for
placement on the exterior of the head, a controller and a
processor. The optical module includes an array of optical input
ports and optical detection ports located in a selected geometrical
pattern to provide a multiplicity of photon migration paths inside
the biological tissue. Each optical input port is constructed to
introduce into the examined tissue visible or infrared light
emitted from a light source. Each optical detection port is
constructed to provide light from the tissue to a light detector.
The controller is constructed and arranged to activate one or
several light sources and light detectors so that the light
detector detects light that has migrated over at least one of the
photon migration paths. The processor receives signals
corresponding to the detected light and forms at least two data
sets, a first of said data sets representing blood volume in the
examined tissue region and a second of said data sets representing
blood oxygenation of the examined tissue. The processor is arranged
to correlate the first and second data sets to detect abnormal
tissue in the examined tissue.
Inventors: |
Chance; Britton; (Marathon,
FL) |
Correspondence
Address: |
IVAN DAVID ZITKOVSKY PH.D PC
5 MILITIA DRIVE
LEXINGTON
MA
02421
US
|
Family ID: |
26755485 |
Appl. No.: |
12/589636 |
Filed: |
October 26, 2009 |
Related U.S. Patent Documents
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Application
Number |
Filing Date |
Patent Number |
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10982542 |
Nov 5, 2004 |
7610082 |
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12589636 |
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09622112 |
Nov 20, 2000 |
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PCT/US99/03030 |
Feb 11, 1999 |
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10982542 |
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60074296 |
Feb 11, 1998 |
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60098018 |
Aug 26, 1998 |
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Current U.S.
Class: |
600/323 ;
600/476 |
Current CPC
Class: |
A61B 5/14553 20130101;
A61B 5/0042 20130101; A61B 2562/0233 20130101; A61B 2562/046
20130101; A61B 5/0073 20130101; A61B 2562/0242 20130101; A61B
5/4064 20130101 |
Class at
Publication: |
600/323 ;
600/476 |
International
Class: |
A61B 5/1455 20060101
A61B005/1455; A61B 6/00 20060101 A61B006/00 |
Claims
1. An optical system for in vivo, non-invasive transcranial
examination of brain tissue of a subject comprising: an optical
module including an array of optical input ports and detection
ports located in a selected geometrical pattern to provide a
multiplicity of photon migration paths inside an examined region of
the biological tissue, each said optical input port being
constructed to introduce visible or infrared light emitted from a
light source, each said optical detection port being constructed to
receive photons of light that have migrated in the examined tissue
region from at least one of said input ports and provide said
received light to a light detector; a controller constructed and
arranged to control operation of said light source and said light
detector to detect light that has migrated over at least one of
said photon migration paths; and a processor connected to receive
signals from said detector and arranged to form at least two data
sets, a first of said data sets representing blood volume in the
examined tissue region and a second of said data sets representing
blood oxygenation in the examined tissue region; said processor
being arranged to correlate said first and second data sets to
detect abnormal tissue in the examined tissue region.
2. The optical system of claim 1 wherein said second data set
includes hemoglobin deoxygenation values.
3. The optical system of claim 1 wherein said processor is arranged
to form a third data set being collected by irradiating a reference
tissue region.
4-8. (canceled)
9. The optical system of claim 1 wherein said processor is
programmed to order said first and second data sets as
two-dimensional images and to determine said congruence using the
following formula: 1 ( maximum overlap residual maximum selected
tissue signal ) .times. 100 ##EQU00003##
10-35. (canceled)
35. An optical method for in vivo, non-invasive transcranial
examination of brain tissue of a subject comprising: providing an
optical module including an array of optical input ports and
detection ports located in a selected geometrical pattern to
provide a multiplicity of photon migration paths inside an examined
region of the tissue; placing said optical module on the exterior
of the head of the subject; introducing visible or infrared light
from at least one said optical input port into an examined tissue
region and receiving photons of light that have migrated in the
examined tissue region to at least one of said detection ports;
detecting said received photons by at least one optical detector
optically coupled to said least one said detection port;
controlling said introducing and detecting steps to collect optical
data corresponding to photons of light that have migrated between
selected input and detection ports; processing said optical data to
form at least two data sets, a first of said data sets representing
blood volume in the examined tissue region and a second of said
data sets representing blood oxygenation in the examined tissue
region; and correlating said first and second data sets to detect
abnormal tissue in the examined tissue region.
36. The optical method of claim 35 including ordering said first
and second data sets as two-dimensional images and determining said
congruence using said two-dimensional images.
37. The optical method of claim 35 including ordering said first
and second data sets as two-dimensional images and determining said
congruence using a formula: 1 - ( maximum overlap residual maximum
selected tissue signal ) .times. 100 ##EQU00004##
38. The optical method of claim 35 including determining a location
of said abnormal tissue within the examined tissue region.
Description
[0001] This application is a continuation of U.S. application Ser.
No. 10/982,542, filed on Nov. 5, 2004, now U.S. Pat. No. 7,610,082,
which is a continuation of U.S. application Ser. No. 09/622,112,
filed on Nov. 20, 2000, which is a 371 of PCT/US99/03030, filed on
Feb. 11, 1999, which claims priority from U.S. Provisional
Application 60/074,294 filed on Feb. 11, 1998 and from U.S.
Provisional Application 60/098,018 filed on Aug. 26, 1998, all of
which are incorporated by reference as if fully set forth
herein.
THE FIELD OF THE INVENTION
[0002] The present invention relates to imaging and qualitative or
quantitative characterization of biological tissue using visible or
infra-red radiation, and more particularly to imaging and
characterization of brain tissue.
BACKGROUND
[0003] Traditionally, X-rays or .gamma.-rays has been used to
examine and image biological tissue. This radiation propagates in
the tissue on straight, ballistic tracks, i.e., scattering of the
radiation is negligible. Thus, imaging is based on evaluation of
the absorption levels of different tissue types. For example, in
roentgenography the X-ray film contains darker and lighter spots.
In more complicated systems, such as computerized tomography (CT),
a cross-sectional picture of human organs is created by
transmitting X-ray radiation through a section of the human body at
different angles and by electronically detecting the variation in
X-ray transmission. The detected intensity information is digitally
stored in a computer which reconstructs the X-ray absorption of the
tissue at a multiplicity of points located in one cross-sectional
plane.
[0004] Near infra-red radiation (NIR) has been used to study
non-invasively the oxygen metabolism in tissue (for example, the
brain, finger, or ear lobe). Using visible, NIR and infra-red (IR)
radiation for medical imaging could bring several advantages. In
the NIR or IR range the contrast factor between a tumor and a
tissue is much larger than in the X-ray range. In addition, the
visible to IR radiation is preferred over the X-ray radiation since
it is non-ionizing and thus, potentially causes fewer side effects.
However, the visible or IR radiation is strongly scattered and
absorbed in biological tissue, and the migration path cannot be
approximated by a straight line, making inapplicable certain
aspects of cross-sectional imaging techniques.
[0005] Computerized Tomography using MR spectrometry has been used
for in vivo imaging. This technique utilizes NIR radiation in an
analogous way to the use of X-ray radiation in an X-ray CT. The
X-ray source is replaced by several laser diodes emitting light in
the NIR range. The NIR-CT uses a set of photodetectors that detect
the light of the laser diodes transmitted through the imaged
tissue. The detected data are manipulated by a computer similarly
as the detected X-ray data would be in an X-ray CT. Different
NIR-CT systems have recognized the scattering aspect of the
non-ionizing radiation and have modified the X-ray CT algorithms
accordingly.
[0006] The above-mentioned X-ray or .gamma.-ray techniques have
been used to detect a tissue tumor. Under the term "angiogenesis" I
mean the generation of new blood vessels into a tissue or organ.
Under normal physiological conditions humans or animals undergo
angiogenesis only in very specific restricted situations. For
example, angiogenesis is normally observed in wound healing, fetal
and embryonal development and formation of the corpus luteum,
endometrium and placenta.
[0007] Both controlled and uncontrolled angiogenesis are thought to
proceed in a similar manner. Persistent, unregulated angiogenesis
occurs in a multiplicity of disease states, tumor metastasis and
abnormal growth by endothelial cells and supports the pathological
damage seen in these conditions. The diverse pathological disease
states in which unregulated angiogenesis is present have been
grouped together as angiogenic dependent or angiogenic associated
diseases. The hypothesis that tumor growth is angiogenesis
dependent was first proposed in 1971. (Folkman J., Tumor
Angiogenesis: Therapeutic Implications., N. Engl. Jour. Med. 285:
1182-1186, 1971) In its simplest terms it states: "Once tumor
`take` has occurred, every increase in tumor cell population must
be preceded by an increase in new capillaries converging on the
tumor." Tumor `take` is understood to indicate a prevascular phase
of tumor growth in which a population of tumor cells occupying a
few cubic millimeters volume and not exceeding a few million cells,
can survive on existing host microvessels. Expansion of tumor
volume beyond this phase requires the induction of new capillary
blood vessels. This explanation was directly or indirectly observed
and documented in numerous publications.
[0008] There is still a need for a non-invasive, relatively
inexpensive technique that can detect, image and characterize a
tumor alone or in conjunction with the above-mentioned techniques.
Furthermore, there is still a need for a non-invasive, relatively
inexpensive technique that can characterize brain tissue to detect
a disease or functional abnormality.
SUMMARY
[0009] The present invention includes different novel apparatuses
and methods for examination of biological tissue and specifically
for transcranial optical examination or monitoring of the brain
using visible or infra-red light. The optical examination technique
can be used alone to detect and characterize a brain tissue anomaly
or can be used in combination with X-ray techniques (including CT),
magnetic resonance imaging (MRI or fMRI), or PET.
[0010] The novel techniques can employ different a single optical
module placed on the head, or several optical modules placed on the
right or left brain hemisphere of a patient that may be alert or
even unconscious. If a suspicious structure in the head is
detected, the technique can non-invasively characterize the
structure (e.g., tissue mass, fluid volume) by taking optical data
at different wavelengths and by generating one or several tissue
specific characteristics related to the tissue metabolism (or
hypermetabolism), biochemistry, pathophysiology (including
angiogenesis) or another characteristic of a pathological tissue
condition.
[0011] In one aspect, the optical examination technique employs an
optical system for in vivo non-invasive examination of a volume of
biological tissue of a subject. The optical examination system
includes an optical module, a controller and a processor. The
optical module includes an array of optical input ports and optical
detection ports located in a selected geometrical pattern to
provide a multiplicity of source-detector paths of photon migration
inside the biological tissue. Each optical input port is
constructed to introduce into the tissue volume visible or infrared
light emitted from a light source. Each optical detection port is
constructed to provide light from the tissue to a light detector.
The controller is constructed and arranged to activate one or
several light sources and light detectors so that the light
detector detects light that has migrated over at least one of the
source-detector migration paths. The processor receives signals
corresponding to the detected light and creates a defined spatial
image of the examined tissue.
[0012] The optical examination system may generate single
wavelength or multiple wavelength images of the examined brain
tissue, wherein the used wavelength is sensitive to absorption or
scattering by a tissue constituent (e.g., an endogenous or
exogenous pigment, tissue cells) or is sensitive to structural
changes in the tissue. The optical images may display tissue
absorption, tissue scattering or both. The optical imaging system
may also generate blood volume, hemoglobin oxygenation images, and
hemoglobin deoxygenation images (or images of any other tissue
constituent) based on a single wavelength optical data or a
multiple wavelength optical data. A processor may use different
image processing and enhancing algorithms known in the art. The
processor may correlate several images to detect a suspicious
structure and then characterize the detected structure. The
correlation includes determining congruency of the detected
structures. The processor may employ different types of combined
scoring, based on several optical images alone, or in combination
with X-ray techniques, MRI or PET, to characterize a suspicious
tissue mass.
[0013] The optical examination system may generate the
above-described images of symmetrical tissue regions of the right
brain and the left brain, symmetrical tissue regions of the brain
lobes of the right brain and the left brain, or may generate images
of both the entire right brain and the entire left brain. The
optical examination system may also generate "model images" of by
irradiating a model constructed to have scattering and absorptive
propertied of a selected tissue region. The optical examination
system may also separately calibrate its sources and detectors on a
model. To identify and characterize a suspicious tissue mass, the
processor may employ the different types of combined scoring by
correlating the different images mentioned above.
[0014] The optical imaging system may collect single wavelength or
multiple wavelength data of a brain tissue model for calibration,
or for detection of background data. In the calibration procedure,
the optical module is placed on the model and the imaging system
can collect a limited number of optical data or can collect optical
data using the same sequences used during the tissue examination.
The system may either collect and store the model data for a
subsequent digital processing, or may adjust the source or detector
gains to detect optical data according to a selected pattern. The
imaging system may use different head models having the same
scattering coefficient and the same absorption coefficient as the
normal brain tissue and the same scattering coefficient and the
same absorption coefficient as the normal skull. The model tissue
may have the scattering and absorption coefficient of infected
cerebral tissue, tissue with cerebral vasculitis, Parkinson's
disease, Alzheimer's disease, or multiple sclerosis. Furthermore,
the models may have different sizes and shapes.
[0015] To characterize the examined tissue, the imaging system can
correlate several images of blood volume, hemoglobin oxygenation,
hemoglobin deoxygenation, or images sensitive to an optical
contrast agent. The imaging system can correlate images of the same
tissue region taken at different times. The correlation of the
images identifies pathological tissue regions, such as tumors
undergoing angiogenetic growth wherein the tumor area exhibits an
increased blood volume and decreased hemoglobin oxygenation.
Furthermore, the correlation of the images can be used to monitor
inhibition of angiogenesis during or after drug treatment.
[0016] The described optical systems can also provide amplitude or
phase cancellation patterns that demonstrated for single or
multiple source-detector pairs remarkable sensitivity and were used
to detect small objects. Using back-projection algorithms or other
known imaging algorithms, the described optical systems can image
sensorimotor activation of adult and pre- and full-term neonate
human brain function and achieve two dimensional resolutions of
less than 1 cm. In addition, the optical system records rapidly and
accurately sensorimotor responses in pre- and full-term infants.
The present systems and methods can be used in evaluation of
cerebral dysfunctions or pathologies of adults, children, infants
or neonates.
[0017] According to another aspect, the optical examination
technique employs an optical system for in vivo, non-invasive
examination of biological tissue of a subject. The optical system
includes an optical module, a controller, and a processor. The
optical module includes an array of optical input ports and
detection ports located in a selected geometrical pattern to
provide a multiplicity of photon migration paths inside an examined
region of the biological tissue. Each optical input port is
constructed to introduce visible or infrared light emitted from a
light source. Each optical detection port is constructed to receive
photons of light that have migrated in the examined tissue region
from at least one of the input ports and provide the received light
to a light detector. The controller is constructed and arranged to
control operation of the light source and the light detector to
detect light that has migrated over at least one of the photon
migration paths. The processor is connected to receive signals from
the detector and arranged to form at least two data sets, a first
of the data sets representing blood volume in the examined tissue
region and a second of the data sets representing blood oxygenation
in the examined tissue region. The processor is arranged to
correlate the first and second data sets to detect abnormal tissue
in the examined tissue region.
[0018] Preferably, the second data set includes hemoglobin
deoxygenation values. The processor may be arranged to form a third
data set being collected by irradiating a reference tissue
region.
[0019] According to another aspect, the optical examination
technique employs an optical system for in vivo, non-invasive
examination of biological tissue of a subject. The optical system
includes an optical module, a controller, and a processor. The
optical module includes an array of optical input ports and
detection ports located in a selected geometrical pattern to
provide a multiplicity of photon migration paths inside an examined
region of the biological tissue. Each optical input port is
constructed to introduce visible or infrared light emitted from a
light source. Each optical detection port is constructed to receive
photons of light that have migrated in the tissue from at least one
of the input ports and provide the received light to a light
detector. The controller is constructed and arranged to control
operation of the light source and the light detector to detect
light that has migrated over at least one of the photon migration
paths. The processor is connected to receive signals from the
detector and arranged to form at least two data sets, a first of
the data sets being collected by irradiating an examined tissue
region of interest and a second of the data sets being collected by
irradiating a reference tissue region having similar light
scattering and absorptive properties as the examined tissue region.
The processor is arranged to correlate the first and second data
sets to detect abnormal tissue in the examined tissue region.
[0020] According to another aspect, the optical examination
technique employs an optical system for in vivo, non-invasive
examination of biological tissue of a subject. The optical system
includes an optical module, a controller, and a processor. The
optical module includes an array of optical input ports and
detection ports located in a selected geometrical pattern to
provide a multiplicity of photon migration paths inside an examined
region of the biological tissue or a model representing biological
tissue. Each optical input port is constructed to introduce visible
or infrared light emitted from a light source. Each the optical
detection port is constructed to receive photons of light that have
migrated in the tissue or the model from at least one of the input
ports and provide the received light to a light detector. The
controller is constructed and arranged to control operation of the
light source and the light detector to detect light that has
migrated over at least one of the photon migration paths. The
processor is connected to receive signals from the detector and
arranged to form at least two data sets of two tissue regions, a
first of the data sets being collected by irradiating an examined
tissue region and a second of the data sets being collected by
irradiating a region of a tissue model having selected light
scattering and absorptive properties. The processor is arranged to
correlate the first and second data sets to detect abnormal tissue
in the examined tissue region.
[0021] Preferred embodiments of these aspects of the inventions
have one or more of the following features.
[0022] The processor may be arranged to correlate the first and
second data sets by determining congruence between data of the two
data sets.
[0023] The processor may be programmed to order the first and
second data sets as two-dimensional images and to determine the
congruence using the two-dimensional images. The processor may be
programmed to order the first and second data sets as
two-dimensional images and to determine the congruence using the
following formula:
1 - ( maximum overlap residual maximum selected tissue signal )
.times. 100 ##EQU00001##
[0024] The processor may be further arranged to determine a
location of the abnormal tissue within the examined tissue
region.
[0025] The processor may be adapted to produce from the data set an
image data set by implementing an optical tomography algorithm. The
optical tomography algorithm may use factors related to determined
probability distribution of photons attributable to the scattering
character of the tissue being imaged.
[0026] The controller may be arranged to activate the source and
the detector to obtain a first selected distance between the input
and detection ports, and the processor may be arranged to form the
data set for the first distance. The processor may produce an image
data set from the data set formed for the first distance. The
controller may further be arranged to activate the source and the
detector to obtain a second selected distance between the input and
detection ports and is arranged to form another data set for the
second distance.
[0027] The optical system may further include a display device
constructed to receive the image data set from the processor and to
display an image.
[0028] The optical system may further include a first oscillator
and a phase detector. The first oscillator is constructed to
generate a first carrier waveform at a first frequency on the order
of 10.sup.8 Hz, the first frequency having a time characteristic
compatible with the time delay of photon migration from the input
port to the detection port. The light source is coupled to the
first oscillator and constructed to generate the light modulated by
the first carrier waveform. The phase detector is constructed to
determine change in waveform of the detected light relative to the
waveform of the introduced light and measure therefrom the phase
shift of the detected light at the wavelength, wherein the
phase-shifted light is indicative of scattering or absorptive
properties of the examined tissue region. The processor is arranged
to form the data set based on the measured phase shift. This
optical system may further include a second oscillator constructed
to generate a second waveform at a second frequency. The detector
is then arranged to receive a reference waveform at a reference
frequency offset by a frequency on the order of 10.sup.3 Hz from
the first frequency and to produce a signal, at the offset
frequency, corresponding to the detected radiation. The phase
detector is adapted to compare, at the offset frequency, the
detected radiation with the introduced radiation and to determine
therefrom the phase shift.
[0029] The optical system may further include an oscillator, a
phase splitter, and first and second double balanced mixers. The
oscillator is constructed to generate a first carrier waveform of a
selected frequency compatible with time delay of photon migration
from the input port to the detection port. The light source is
connected to receive from the oscillator the carrier waveform and
is constructed to generate optical radiation modulated at the
frequency. The phase splitter is connected to receive the carrier
waveform from the oscillator and produce first and second reference
phase signals of predefined substantially different phases. The
first and second double balanced mixers are connected to receive
from the phase splitter the first and second reference phase
signals, respectively, and are connected to receive from the
detector the detector signal and to produce therefrom a in-phase
output signal and a quadrature output signal, respectively. The
processor being connected to the double balanced mixers and
arranged to receive the in-phase output signal and the quadrature
output signal and form therefrom the data set.
[0030] The processor may be arranged to calculate a phase shift
(.THETA..sub..lamda.) between the light introduced at the input
port and the light detected at the detection port prior to forming
the data set.
[0031] The processor may arranged to calculate an average migration
pathlength of photons scattered in the examined tissue between the
optical input port and the optical detection port prior to forming
the data set.
[0032] The processor may further employ the pathlength in
quantifying hemoglobin saturation (Y) of the examined tissue.
[0033] The processor may be arranged to calculate a signal
amplitude (A.sub..lamda.) determined as a square root of a sum of
squares of the in-phase output signal and the quadrature output
signal prior to forming the data set.
[0034] The optical system may further include a narrow band
detector connected to receive from the optical detector the
detector signal and to produce a DC output signal therefrom. The
processor then further determines a modulation index
(M.sub..lamda.) as a ratio of values of the signal amplitude and
the signal amplitude plus the DC output signal.
[0035] The optical system may further include at least one
oscillator connected to at least one light source. The oscillator
is constructed to generate a carrier waveform of a selected
frequency. The light source generate slight of a visible or
infrared wavelength being intensity modulated at the frequency to
achieve a known light pattern. The controller is constructed to
control the emitted light intensity or phase relationship of
patterns simultaneously introduced from multiple input ports,
wherein the introduced patterns form resulting radiation that
possesses a substantial gradient of photon density in at least one
direction. This resulting radiation is scattered and absorbed over
the migration paths. The detector is constructed and arranged to
detect over time the resulting radiation that has migrated in the
tissue to the detection port. The processor is further arranged to
process signals of the detected resulting radiation in relation to
the introduced radiation to create the data sets indicative of
influence of the examined tissue upon the substantial gradient of
photon density of the resulting radiation.
[0036] The optical system may further include a phase detector
constructed to detect the phase of the detected radiation and
provide the phase to the processor.
[0037] The optical system may further include an amplitude detector
constructed to detect the amplitude of the detected radiation and
provide the amplitude to the processor.
[0038] The phase relationship of light patterns introduced from two
input ports may be 180 degrees.
[0039] The optical system may be constructed as described in U.S.
Pat. No. 5,119,815 or 5,386,827. This system includes a light
source constructed to generate pulses of radiation of the
wavelength, the pulses having a known pulse wave form of a duration
on the order of a nanosecond or less. An optical detector is
constructed to detect over time photons of modified pulses that
have migrated in the tissue from the input ports. This system also
includes an analyzer connected to the detector and adapted to
determine a change in the pulse waveform shape of the detected
pulses relative to the introduced pulses, at the employed
wavelength. The processor then creates the data set based on the
determined pulse waveform change. The processor may also be
constructed and arranged to calculate the effective pathlength of
photons of the wavelength migrating between the input and detection
ports in conjunction with creating the data set. The processor may
also be constructed and arranged to calculate the scattering
coefficient at the wavelength in conjunction with creating the
image data set The processor may also be constructed and arranged
to calculate the absorption coefficient at the wavelength in
conjunction with creating the data set.
[0040] The optical system may use the light source that produces
relatively long light pulses and the processor that forms the data
set by subtracting amplitude of two the pulses emitted from two
input ports located symmetrically relative to one detection
port.
[0041] The optical system may be constructed to introduce and
detect photons at two wavelengths selected to be sensitive to a
tissue constituent. The tissue constituent may be an endogenous
pigment or an exogenous pigment. The endogenous pigment may be
hemoglobin. The exogenous pigment may be a selected contrast
agent.
[0042] According to another aspect, an optical system for in vivo,
non-invasive imaging of tissue change includes an optical module
including an array of input ports and detection ports located in a
selected geometrical pattern to provide a multiplicity of arrayed
single source, single detector pairs engaged directly with the
subject. The optical system also includes a spectrophotometer with
a light source means constructed to introduce electromagnetic
radiation of visible or infra-red wavelength into the examined
tissue successively at the input ports, the wavelength being
sensitive to a constituent of the imaged tissue, and detector means
constructed to detect, at the detection ports, radiation of the
selected wavelength that has migrated in the tissue from respective
input ports. The spectrophotometer also includes a processor
connected to receive signals of the detected radiation from the
detector means and constructed to create a defined spatial image of
the tissue by effectively producing from signals from the
multiplicity of arrayed single source, single detector pairs, a
succession of data sets representing, from a selected view, a
succession of spatial images of the tissue, and an image data set
related to differences between data of the successive data
sets.
[0043] According to another aspect, an optical system is provided
for in vivo, non-invasive functional neuroimaging of brain tissue.
The optical system include a stimulator constructed to stimulate a
selected functional activity of neural tissue of interest, an
optical module including an array of input ports and detection
ports located in a selected geometrical pattern to provide a
multiplicity of arrayed single source, single detector pairs
engaged directly with the subject, a spectrophotometer including
light source means constructed to introduce electromagnetic
radiation of visible or infra-red wavelength into the examined
neural tissue successively at the input ports, the wavelength being
sensitive to a tissue constituent associated with a physiological
response of the imaged functional activity, detector means
constructed to detect, at the detection ports, radiation of the
selected wavelength that has migrated in the stimulated neural
tissue from respective input ports, and a processor receiving
signals of the detected radiation from the detector means, and
constructed and arranged to create a defined spatial image of the
functional activity of neural tissue by effectively producing from
the signals from the multiplicity of arrayed single source, single
detector pairs, a first data set representing, from a selected
view, a spatial image of the neural tissue at rest, a second data
set representing, from the same selected view, a spatial image of
the neural tissue during stimulation, and a functional image data
set that is related to the differences between the first and second
data sets, over the sets.
[0044] According to another important aspect, an instrument is
provided for functional imaging of brain activity of a subject
comprising an imager constructed and arranged to image
oxyhemoglobin, deoxyhemoglobin or blood volume. The imager includes
an array of sources of near infrared or visible photons, and array
of detectors positioned to receive photons from the sources
following migration of photons from the sources through the tissue.
The imager enables numerous readings of migrated photons to be
taken systematically for different source-detector positions
relative to the tissue, and a processor employing data sets taken
during rest and during stimulation, with an imaging algorithm that
is based on respectively different probabilities for a given
source-detector position, for photons from the source passing
through different regions of the volume of the scattering tissue
that are located at different positions distributed laterally from
a straight reference line between source and detector.
[0045] Preferred embodiments of these aspects of the inventions
have one or more of the following features.
[0046] The optical module is constructed to maintain a selected
distance between the input and detection ports for the respective
source-detector pairs during the production of the first and second
data sets, the distance being selected according to the tissue
depth desired to be imaged.
[0047] To characterize the examined tissue, the imaging system can
correlate several images of blood volume, hemoglobin oxygenation,
hemoglobin deoxygenation, or images sensitive to an optical
contrast agent, prior and after stimulation. The imaging system can
also correlate the images taken over time without stimulation. The
correlation of the images identifies pathological tissue regions or
dysfunctional tissue regions of the brain.
[0048] The optical module or an associated set of the modules is
constructed to detect light that has migrated in the tissue at
different depths to produce 3D data sets from which an image to
data set may be produced.
[0049] The processor is adapted to produce the image data set by
implementing an optical tomography algorithm.
[0050] The optical tomography algorithm preferably employs factors
related to determined probability distribution of photons
attributable to the scattering character of the tissue being
imaged.
[0051] The optical system is constructed to form the image data set
from a part of the head. In particular embodiments the optical
system is constructed to form the functional image data set from
below the surface region of the cortex.
[0052] The stimulator is constructed to stimulate the visual
cortex, the cognitive cortex, the sensory motor cortex, or spinal
tissue. In various embodiments, the stimulator is constructed to
deliver electrical signals to selected tissue, apply an electrical
field to selected tissue, or deliver magnetic signals to selected
tissue.
[0053] In various embodiments the image set is related to at least
one of the group consisting of blood volume, hemoglobin oxygenation
or deoxygenation, photon absorption coefficient, photon scattering
coefficient, refractive index, change in magnetic field, change in
electric field, production of or change of a specific tissue
constituent, and production of or change in the concentration of a
tissue constituent. The tissue constituent may be an endogenous
pigment, for example hemoglobin, or an exogenous pigment, for
example a selected optical contrast agent.
[0054] The source means, the detector means, the source to detector
distance, and the rate of excitation and detection are selected to
enable an image data set to be obtained within a short time, i.e.,
within minutes, preferably within a minute or less.
[0055] Each source is laterally displaced from its detector or
detectors (or each detector is laterally displaced from its source
or sources) on the surface of a subject at a side by side spacing
between about 1 cm and 10 cm (preferably 1.5 cm and 7 cm) to
establish a banana-shaped probability gradient of migrating photons
in the tissue that extends from source to detector.
[0056] The invention also features methods of producing an image
from a volume of light-scattering tissue of a living subject
comprising, providing and employing on the subject an imaging
instrument according to any of the foregoing aspects. In certain
preferred embodiments of the methods an optical contrast agent or a
drug is introduced to the blood stream of the subject, and the
instrument is employed to produce an image data set for the tissue
while the contrast agent or drug is present in blood circulating in
the tissue of the subject or is present in localized tissue.
[0057] Other advantages and features of the invention will be
apparent from the following description of the preferred embodiment
and from the claims.
BRIEF DESCRIPTION OF THE DRAWINGS
[0058] FIGS. 1 and 1A show an optical module located on the
forehead of a subject.
[0059] FIGS. 2 and 2A show another embodiment of the optical module
located on the forehead of the subject.
[0060] FIGS. 3 and 3A show diagrammatically respective single
wavelength and dual wavelength phase cancellation imaging systems
that employ the optical module of FIG. 1A or FIG. 2A.
[0061] FIG. 3B is a timing diagram used by the imaging system of
FIGS. 3 and 3A.
[0062] FIGS. 4 and 4A show diagrammatically another embodiment of
the phase cancellation imaging system employing the optical module
of FIG. 1A or FIG. 2A.
[0063] FIG. 5 shows diagrammatically another embodiment of the
phase cancellation imaging system employing the optical module of
FIG. 1A or FIG. 2A.
[0064] FIG. 6 shows schematically an amplitude cancellation imaging
system using another embodiment of the optical module shown in FIG.
6A.
[0065] FIGS. 7, 7A and 7B show different embodiments of a cooling
module used with a broad band light source such as a tungsten light
bulb.
[0066] FIG. 8 shows diagrammatically another embodiment of the
amplitude cancellation imaging system employing the optical module
of FIG. 2A.
[0067] FIG. 8A shows a circuit configuration for one element of the
amplitude cancellation imaging system of FIG. 8.
[0068] FIG. 8B is a timing diagram used by the imaging system of
FIG. 8.
[0069] FIG. 8C shows diagrammatically one channel of the amplitude
cancellation imaging system of FIG. 8.
[0070] FIG. 8D shows diagrammatically another embodiment of the
amplitude cancellation imaging system of FIG. 8.
[0071] FIG. 9 is an example of a "four" dimensional graph that
could be used to summarize optical examination of suspicious
masses.
[0072] FIGS. 10 and 10A show an experimental optical image obtained
by the imaging system of FIG. 3 with contralateral, parietal finger
touching as a stimulation.
[0073] FIG. 11 shows co-registration of optical and NMR signals in
sensory motor simulation.
[0074] FIGS. 12A through 13D show prefrontal cortex optical images
detected by the optical imaging system of FIG. 3 during a cognitive
activity of subjects.
[0075] FIGS. 14A and 14B are histograms of the positions on the
forehead for two subjects.
[0076] FIGS. 15A and 15B show optical images during the functional
activation of pre- and full-term neonates.
DESCRIPTION OF THE PREFERRED EMBODIMENTS
[0077] Referring to FIGS. 1, 1A, 2 and 2A the brain tissue of a
subject 8 is examined using an imaging system connected to an
optical module 12 or 14. Optical modules 12 and 14 include a
multiplicity of light sources (e.g., laser diodes, LEDs, flashlight
bulbs) providing light in the visible to infrared range and light
detectors (e.g., photo multiplier tubes, Si diode detector, PIN,
avalanche or other diode detectors), which may also include
interference filters. The light sources and the light detectors are
arranged to form selected geometrical patterns that provide a
multiplicity of source-detector paths of photon migration inside
the brain tissue. An optical examination system provides an in vivo
optical data of the examined tissue, and the data may be processed
to create an image. The image can show a location and size of an
abnormal structure in the tissue, such as a tumor or bleeding.
Furthermore, the optical data can provide a qualitative and
quantitative measure (e.g., metabolism, metabolic biochemistry,
pathophysiology) of an abnormal tissue structure. (Alternatively,
an optical module includes a multiplicity of optical fibers
connected to one or several light sources, and a multiplicity of
optical detection fibers connected to one or several light
detectors as described in the PCT applications PCT/US96/00235 and
PCT/US96/11630 (filed Jan. 2, 1996 and Jul. 12, 1996).)
[0078] In one embodiment, optical module 12 includes nine laser
diodes S.sub.1, S.sub.2, . . . , S.sub.9 and four photo multiplier
tubes (PMTs) D.sub.1, D.sub.2, D.sub.3, D.sub.4. The laser diodes
and PMTs are embedded in a pliable rubber-like material positioned
in contact with the scalp. There may be a Saran.RTM. wrap or
similar material located between the laser diodes and the skin, and
between the PMTs and the skin. Similarly, optical module 14
includes four laser diodes S.sub.1, S.sub.2, S.sub.3, S.sub.4 and
27 silicon diode detectors D.sub.1, D.sub.2, D.sub.27 embedded in a
pliable rubber-like material. The optical systems shown in FIGS. 3
through 7 may be interfaced with optical module 12 or 14 for
imaging of the brain tissue. Optical modules 12 and 14 have pairs
of optical input ports symmetrically located (or equidistantly
located) relative to an optical detection port, or have pairs of
optical detection ports symmetrically located relative to an
optical input port. In general, however, the ports do not have to
be positioned symmetrically. The optical systems can vary the
source or detector gain to account for any positional asymmetry or
can introduce a selected asymmetry by adjusting the source or
detector gain.
[0079] Furthermore, the systems shown in FIGS. 3 through 7, may be
interfaced with two identical optical modules (12 or 14) located on
symmetrical brain tissue, such as the right brain hemisphere and
the left brain hemisphere for lateralization, that is, comparative
tissue examination the right brain hemisphere and the left brain
hemisphere. The comparative examination may be performed on the
individual brain lobes, such as the right temporal lobe and the
left temporal lobe, the right occipital lobe and the left occipital
lobe, or the right parietal lobe and the left parietal lobe of the
brain. Alternatively, the comparative examination may be performed
on symmetric tissue of the same lobe, such as the frontal lobe. For
calibration, the optical module may also be placed on one or
several models of the head having the same scattering coefficient
and the same absorption coefficient as the normal brain tissue
including the skull.
[0080] Referring to FIGS. 1A and 3, a phased array imaging system
15 is connected to optical module 12 with nine laser diodes
S.sub.1, S.sub.2, . . . , S.sub.9 and four PMTs D.sub.1, D.sub.2,
D.sub.3, D.sub.4 (e.g., Hamamatsu R928, Hamamatsu R1645u, TO8)
powered by a high voltage supply (not shown). Four laser diodes
surround each PMT forming an equidistant arrangement (for example,
different optical modules may use distances of 3.5, 7 and 10.5 cm).
A switch 18 connects laser diodes S.sub.1, S.sub.2, . . . , S.sub.9
to a phase splitter 20, which provides to the diodes an RF
modulation signal having both a 0 degree phase and a 180 degree
phase. Imaging system 15 also includes a 50 MHZ single side band
transmitter 22 connected by a phase lock loop 24 to a 50 MHZ single
side band receiver 26. Single side band (SSB) transmitter 22 is
connected to a 1 kHz oscillator 28, which provides a reference
signal 30 to a phase detector 32. SSB receiver 26 is connected to a
switch 27, which connects one of the four PMTs (0.5 .mu.V
sensitivity) depending on control signals from a controller 19. The
SSB transmitter-receiver pair can operate in the frequency region
of 10-1000 MHZ (preferably 50-450 MHZ). The SSB receiver detects
signal levels on the order of microvolts in a 2 KHz bandwidth. The
phase noise of this apparatus is less than about 0.1.degree..
However, this narrow bandwidth limits the spread of switching of
various light sources to approximately 1.0 msec, and thus the
sequencing time for an entire image of 16 source detector
combinations can be .about.1 sec. The system uses a 1 sec averaging
time.
[0081] Controller 19, connected to a personal computer (not shown),
sequences laser diodes S.sub.1, S.sub.2, . . . , S.sub.9 so that
two diodes receive 0.degree. phase and 180.degree. phase signals
from splitter 20, every 0.1 sec. At the same time, controller 19
connects a symmetrically located PMT to SSB receiver 26. As shown
in a timing diagram 40 (FIG. 3B), phased array imaging system 15
triggers two sources so that they emit modulated light of a
0.degree. phase and a 180.degree. phase for about 100 msec, and at
the same time triggers a symmetrically located PMT. For example,
when laser diodes 1 (S.sub.1) and 2 (S.sub.2) emit light of a
0.degree. and 180.degree. phase, respectively, and detector 1
(D.sub.1) detects light that has migrated in the examined tissue.
SSB receiver 26, which is phase locked with SSB transmitter 22,
receives signal from detector 1 and provides output signal 34 to
phase detector 32. Phase detector 32 measures the phase (36) of the
detected light, and SSB receiver 26 provides the amplitude (38) of
the detected light. This phase detection circuit was described in
U.S. Pat. No. 4,972,331, which is incorporated by reference.
[0082] In the next cycle, controller 19 directs switch 18 to
connect laser diodes 2 (S.sub.2) and 3 (S.sub.3), which emit
modulated light of a 0.degree. phase and a 180.degree. phase,
respectively, and detector 2 (D.sub.2) detects light that has
migrated in the examined tissue. Controller 19 also directs switch
27 to connect detector 2 to SSB receiver 26, which receives
detection signal corresponding to the photons that have migrated
from laser diodes 2 and 3 to detector 2. Again, phase detector 32
measures the phase (36) of the detected light, and SSB receiver 26
provides the amplitude (38) of the detected light. The duration of
each pair of light flashes is 100 msec. The complete set of data
for all source detector combinations is collected every 30 sec. A
computer (not shown) stores the phase values and the amplitude
values measured for the different combinations shown in timing
diagram 40 and employs these values to create images of the
examined tissue, as is described below. The computer uses the
ADA2210 board for data acquisition.
[0083] Before or after the above-described measurement, phased
array imaging system 15 may be calibrated on a model of the skull
and brain tissue. In the calibration procedure, the optical module
is placed on the model and the imaging system collects the phase
data and the amplitude data using the sequences shown in the timing
diagram 40. The imaging system may use different models having the
same scattering coefficient and the same absorption coefficient as
the normal brain tissue, a brain that suffered trauma manifested as
cerebral edema, cerebral contusion, intracranial hemorrhage. The
model tissue may have scattering and absorption coefficient of
infected cerebral tissue, tissue with cerebral vasculitis,
Parkinson's disease, Alzheimer's disease or multiple sclerosis.
Furthermore, the models may have different sizes and shapes.
[0084] Phased array imaging system 15 generates a "model" image for
each wavelength employed. The model image may later be subtracted
from the brain images to calibrate the system and also account for
the boundary conditions of the light migrating in the tissue.
Alternatively, phased array imaging system 15 is calibrated prior
to taking measurement data and the gain on the light sources or the
detectors is adjusted to obtain selected values.
[0085] Referring to FIGS. 1A and 3A, a dual wavelength phased array
imaging system 45 is connected to optical module 12 with nine 780
nm laser diodes S.sub.1, S.sub.2, . . . , S.sub.9, nine 830 nm
laser diodes S.sub.1a, S.sub.2a, . . . , S.sub.9a, and the four
PMTs D.sub.1, D.sub.2, D.sub.3, and D.sub.4 powered by a high
voltage supply (not shown). Pairs of laser diodes S.sub.1 and
S.sub.1a, S.sub.2 and S.sub.2a, . . . , S.sub.9 and S.sub.9a are
located next to each other and arranged to introduce modulated
light at almost the same tissue locations. A switch 48 connects
laser diodes S.sub.1, S.sub.2, . . . , S.sub.9 to a phase splitter
50, which provides to the laser diodes an RF modulation signal
having both a 0 degree phase and a 180 degree phase. Similarly, a
switch 48a connects laser diodes S.sub.1a, S.sub.2a, . . . ,
S.sub.9a to a phase splitter 50a, which provides to the laser
diodes an RF modulation signal having both a 0 degree phase and a
180 degree phase. A 52 MHZ SSB transmitter 52 is connected by a
phase lock loop 54 to a 52 MHZ SSB receiver 56, and a 50 MHZ SSB
transmitter 52a is connected by a phase lock loop 54a to a 50 MHZ
SSB receiver 56a. Both SSB transmitters 52 and 52a are connected to
a 1 kHz oscillator 58, which provides a reference signal 60 to
phase detectors 62 and 62a. SSB receivers 56 and 56a are connected
one of the four PMTs by a switch 57 depending on control signals
from controller 49. Controller 49, connected to a personal
computer, sequences the laser diodes so that two pairs of the laser
diodes receive 0.degree. phase and 180.degree. phase signals from
splitters 50 and 50a, and at the same time controller 49 connects a
symmetrically located detector to SSB receivers 56 and 56a.
[0086] As shown in timing diagram 40 (FIG. 3B), phased array
imaging system 45 triggers for each wavelength two sources that
emit simultaneously modulated light of a 0.degree. phase and a
180.degree. phase for about 100 msec and, at the same time,
controller 49 connects the symmetrically located PMT. For example,
switch 48 connects SSB transmitter 52 to 780 nm laser diode 4
(S.sub.4) to emit 52 MHZ modulated light of a 180.degree. phase and
connects 780 nm laser diode 5 (S.sub.5) to emit 52 MHZ modulated
light of a 0.degree. phase. At the same time, switch 48a connects
SSB transmitter 52a to 830 nm laser diode 4a (S.sub.4a) to emit 50
MHZ modulated light of a 180.degree. phase and connects 830 nm
laser diode 5a (S.sub.5a) to emit 52 MHZ modulated light of a
0.degree. phase. Simultaneously, switch 57 connects detector 1
(D.sub.1) to SSB receivers 56 and 56a to receive the detection
signal corresponding to photons of both wavelengths that have
migrated in the examined tissue.
[0087] Phase detector 62 provides the phase (66) of the detected
780 nm light, and phase detector 62a provides the phase (66a) of
the detected 830 nm light for the selected geometry. Similarly, SSB
receiver 56 measures the amplitude (68) of the detected 780 nm
light and SSB receiver 56a measures the amplitude (68a) of the
detected 830 nm light. This operation is repeated for all
combinations of sources and detectors shown in timing diagram 40. A
computer (not shown) stores the phase values and the amplitude
values (at each wavelength) measured for the different combinations
shown in timing diagram 40. The computer then uses the measured
values to create images using algorithms included the enclosed
source code. Initially, the system takes quick pictures to find the
area of interest so that the optical module can be moved around to
find an optimal geometry. Once found, the 780 nm and 830 nm data
(i.e., both the phase and amplitude data) is acquired and saved on
a disk.
[0088] Several phased array systems were described in the PCT
application PCT/US 93/05868 (published as WO 93/2514 on Dec. 23,
1993), which is incorporated by reference. This PCT publication
also describes the basic principles of phase and amplitude
cancellation. The phased array imaging system uses a detector for
detecting light emitted from equidistant sources located
symmetrically with respect to the detector (or one source and
several equidistant detectors located symmetrically). If two
sources S.sub.1 and S.sub.2 emit modulated light having equal
amplitude and a 0.degree. phase and a 180.degree. phase, detector
D.sub.1 located in the middle detects a null in the amplitude
signal and detects a crossover between the 0.degree. and
180.degree. phase, i.e., a 90.degree. phase, for substantially
homogeneous tissue. That is, the detector is located on the null
plane. In heterogeneous tissue, the null plane is displaced from
the geometric midline. Nevertheless, the null establishes an
extremely sensitive measure to perturbation by an absorber or
scatterer. Furthermore, at the null condition, the system is
relatively insensitive to amplitude fluctuations common to both
light sources, and insensitive to inhomogeneities that affect a
large tissue. The system has a high sensitivity to scattering
provided that the scattering contrast is the same as the absorbing
contrast. The system can readily observe shifts of 50 to 60.degree.
of phase under altered blood volume or blood oxygenation
conditions, where the phase noise is less than a 0.1.degree.
(s/n>400) for a 1 Hz bandwidth. The amplitude signal is little
less useful in imaging since the position indication is somewhat
ambiguous, i.e., an increase of signal is observed regardless of
the displacement of the absorbing object with respect to the null
plane, although this is remedied by further encoding of the
sources.
[0089] As described in the PCT application PCT/US 93/05868, the
light sources excite a photon diffusion wave, due to cancellation
effects, which has a relatively long wavelength (--10 cm),
determined by the scattering (.mu..sub.s'=10 cm.sup.-1) and
absorption (.mu..sub.a=0.04 cm.sup.-1) properties of the tissue.
The photon diffusion wavelength of about 10 cm provides imaging in
the "near field." The imaging system may use light sources of one
or several optical wavelengths in the visible to infrared range,
depending on the characteristic to be imaged (i.e., blood volume,
blood oxygenation, a distribution of a contrast agent in the
tissue, an absorbing constituent of the tissue, a fluorescing
constituent of the tissue, or other). The phase signal at zero
crossing detection is essentially a square wave "overloaded"
signal. It is moderately insensitive to the changes of signal
amplitude that may occur in imaging from proximal to distal
source-detector pairs and is also moderately insensitive to ambient
light.
[0090] Referring to FIG. 4, in another embodiment, a phased array
imaging system 100 is used instead of imaging systems 15 or 45.
Imaging system 100, connected to optical module 12 (shown in FIG.
1A) having nine laser diodes S.sub.1, S.sub.2, . . . , S.sub.9 and
four PMTs D.sub.1, D.sub.2, D.sub.3, and D.sub.4, employs homodyne
phase detection. A switch 102 connects laser diodes S.sub.1,
S.sub.2, . . . , S.sub.9 to a phase splitter 104, which provides to
the diodes an RF modulation signal having both a 0 degree phase and
a 180 degree phase. Imaging system 100 also includes a 200 MHZ
oscillator 106 providing RF signal to a driver 108, which is
connected to phase splitter 104. (Alternatively, an oscillator in
the range of 10-1000 MHZ, preferably 50-500 MHZ, may be used.) A
phase shifter 114 receives the drive signal (112) from driver 108
and provides the signal of a selected phase (e.g., a 0.degree.
phase change) to a 90.degree. phase splitter 116. Phase splitter
116 provides a 0.degree. phase signal (118) and a 90.degree. phase
signal (120) to double balance mixers (DBM) 122 and 124,
respectively.
[0091] A controller 140, connected to a personal computer,
sequences laser diodes S.sub.1, S.sub.2, . . . , S.sub.9 using
switch 102 so that two diodes receive modulate signal at a
0.degree. phase and a 180.degree. phase from splitter 104. At the
same time, a controller 140 connects a symmetrically located PMT
using a switch 130 to an amplifier 134. Amplifier 134 provides a
detection signal (136) to double balance mixers 122 and 124, and to
a DC detector 138. Double balance mixer 122 receives the detection
signal (136) and the 0.degree. phase reference signal (118) and
provides an in-phase signal 1 (144). Double balance mixer 124
receives the detection signal (136) and the 90.degree. phase
reference signal (120) and provides a quadrature signal R (142). DC
detector 138 provides DC signal (146). The in-phase signal I and
quadrature signal R specify the phase (.theta.=tan.sup.-1I/R) of
the detected optical radiation and the amplitude
(A=(R.sup.2+I.sup.2).sup.-1/2) of the detected optical radiation.
This phase detection circuit was described in U.S. Pat. No.
5,553,614, which is incorporated by reference.
[0092] Similarly as for imaging systems 15 and 45, imaging system
100 directs controller 140 to sequence the laser diodes and the PMT
detectors using timing diagram 40. The computer stores the phase
value and the amplitude value measured for each of the combinations
and generates images described below.
[0093] FIG. 4A shows diagrammatically one portion of phase
cancellation, phased array imaging system 100. The depicted portion
of imaging system 100 includes two laser diodes LD.sub.1, and
LD.sub.2 and a light detector D.sub.1, which are included in
optical module 12 or 14. Oscillator 106 provides carrier waveform
having a frequency in range of 30 to 140 MHz. The carrier waveform
frequency is selected depending on the operation of the system.
When time multiplexing the light sources using switch 102, then the
carrier waveform is modulated at a lower frequency, e.g., 30 MHz to
afford switching time.
[0094] When no time multiplexing is performed, oscillator 106
operates in the 100 MHz region. Splitter 104 splits the oscillator
waveform into 0.degree. and 180.degree. signals that are then
attenuated by digitally controlled attenuators 107A and 107B by 0%
to 10% in amplitude. The phase of the attenuated signals is
appropriately shifted by digitally controlled phase shifters 109A
and 109B in the range of 10.degree.-30.degree. and preferably
20.degree. in phase. Laser drivers 108A and 108B drive LD.sub.1 and
LD.sub.2, respectively, which emit light of the same wavelength,
for example, 780 or 800 nm. After the introduced light migrates in
the examined tissued, a PMT detector D.sub.1 amplifies the detected
signals having initially the 0 and 180.degree. phases. As described
above, for homogeneous tissue and symmetric locations of LD.sub.1,
LD.sub.2 and D.sub.1, the output of the PMT is 90.degree., i.e.,
halfway between 0.degree. and 180.degree. and the amplitude is
close to zero. The personal computer (PC) adjusts the attenuation
provided by attenuator 107B and the phase shift provided by phase
shifter 109B so that detector D.sub.1 detects phase nominally
around 25.degree. and amplitude nominally around .ltoreq.10
millivolts for homogeneous tissue. This signal is connected to
amplifier 134 and to the IQ circuit 139. The cosine and sine
signals are fed into the personal computer, which takes the
amplitude (the square root of the sum of the squares of I and Q)
and the phase angle (the angle whose tangent is I/Q) to give
outputs of phase around 25.degree. and amplitude signals around 10
millivolts. The personal computer also adjusts the reference signal
to the IQ to have the phase .PHI..sub.3 between 10.degree. to
30.degree. and preferably around 25.degree., i.e., phase shifter
114 provides to the IQ circuit 139 the reference phase having a
value selected by the combination of phase shifters 109A and
109B.
[0095] In a currently preferred embodiment, splitter 104 is a two
way 180.degree. power splitter model number ZSCJ-2 1, available
from Mini-Circuits (P.O. Box 350186, Brooklyn, N.Y. 11235-0003).
The phase shifters 109A, 109B and 114 and attenuators 107A, and
107B are also available from Mini-Circuits, wherein the attenuators
can be high isolation amplifier MAN-IAD. IQ demodulator 139 is a
demodulator MIQY-140D also available from Mini-Circuits.
[0096] The system obtains the initial values of attenuator 107B
(A.sub.2) and phase shifter 109B (.PHI..sub.2) on a model or a
symmetric tissue region (e.g., the contralateral brain lobe that is
tumor free). The entire optical probe is calibrated on a tissue
model by storing the calibration values of A.sub.2 and .PHI..sub.2
for the various source-detector combinations (i.e., the baseline
image). The probe is then moved to the exterior of the head, for
example, and the phases and amplitudes are detected for the various
source and detector combinations. When the contralateral tumor free
brain lobe is used as a model, the probe is transferred to the
contralateral lobe (taking note to locate the probe on the
symmetrical tissue considering the brain physiology) and then the
images are read out from all the source-detector combinations to
acquire the tissue image. There is no limitation on multiplexing as
long as the bandwidth of F.sub.1 and F.sub.2 is recognized as being
the limiting condition in the system normalization. It should be
noted that normalization must be accurate and without "dither" and
therefore, a significant amount of filtering in F.sub.1 and
F.sub.2, i.e., less than 10 Hz bandwidth. If .PHI..sub.2 is
adjusted over a large range, there will be an amplitude-phase
crosstalk. Thus, the system may adjust phase and then amplitude and
repeat these adjustments iteratively because of the amplitude phase
crosstalk. The control of A.sub.1 and .PHI..sub.1 provides even a
greater range of control, where obviously inverse signals would be
applied to them, i.e., as the A.sub.1.PHI..sub.1 signals are
increased, the A.sub.2, .PHI..sub.2 signals would be decreased.
Both A.sub.2 and .PHI..sub.2 can be controlled by PIN diodes, to
achieve an extremely wideband frequency range. However, since
signal processing controls the bandwidth of the feedback system,
that either PIN diode or relay control of the phase and amplitude
is feasible for automatic compensation. If, in addition, dual
wavelength or triple wavelength sources are used, each one of them
must be separately calibrated because no two light sources can be
in the same position relative to the imaged tissue (unless, of
course, they are combined with optical fibers).
[0097] Referring to FIG. 5, in another embodiment, a dual
wavelength phased array imaging system 150 is used instead of
imaging systems 15, 45 or 100. Imaging system 150, connected to
optical module 12 (shown in FIG. 1A) having nine 760 nm laser
diodes S.sub.1, S.sub.2, . . . , S.sub.9, nine 840 nm laser diodes
S.sub.1a, S.sub.2a, . . . , S.sub.9a and four PMTs D.sub.1,
D.sub.2, D.sub.3, and D.sub.4 is based on heterodyne phase
detection. A switch 152 connects the laser diodes to a phase
splitter 154, which provides to the diodes an RF modulation signal
having both a 0 degree phase and a 180 degree phase. Imaging system
150 employs a mixer 165 connected to a 200 MHZ oscillator 160 and
200.025 MHZ oscillator 162 (Alternatively, oscillators operating in
the range of 10-1000 MHZ, preferably 50-500 MHZ, may be used.)
Mixer 165 provides a 25 kHz reference signal (168) to an adjustable
gain controller 177. Oscillator 162 connected to power amplifier
163 provides a 200.025 MHZ reference signal (170) to the second
dynode of each PMT detector for heterodyne detection. Each PMT
detector provides a 25 kHz detection signal (172) to a switch 178,
which in turn provides the signal to a 25 kHz filter 180. A phase
detector 184 is connected to an adjustable gain controller 182,
which provides a filtered and amplified detection signal (186) and
to adjustable gain controller 177, which provides the reference
signal (188). Phase detector 184, connected to a switch 190,
provides the detected phase value for each wavelength. This phase
detection circuit was described in U.S. Pat. No. 5,187,672, which
is incorporated by reference. Another type of phase detection
circuit was described in U.S. Pat. No. 5,564,417, which is
incorporated by reference.
[0098] Similarly as described above, controller 175, connected to a
personal computer, sequences laser diodes S.sub.1, S.sub.2, . . . ,
S.sub.9 or laser diodes S.sub.1a, S.sub.2a, . . . , S.sub.9a using
switch 152 so that two diodes emitting the same wavelength receive
0.degree. phase and 180.degree. phase signals from splitter 154. At
the same time, controller 175 connects a symmetrically located PMT
using a switch 178 to filter 180 and adjustable gain controller
182. Phase detector 184 provides the measured phase. Imaging system
employs timing diagram 40 (FIG. 3B); however, since the two
wavelength light is not frequency encoded, laser diodes S.sub.1,
S.sub.2, . . . , S.sub.9 or laser diodes S.sub.1a, S.sub.2a, . . .
, S.sub.9a are triggered in each sequence. The computer stores the
phase values measured for the different combinations and generates
images described below.
[0099] Referring to FIG. 6, in another embodiment, an amplitude
cancellation imaging system 200 uses an optical module 212 shown in
FIG. 6B. Optical module 212 includes twelve light sources S1, S2, .
. . , S12 and four light detectors D1, D2, D3, and D4 mounted on a
plastic or rubber foam material. The light sources and the light
detectors are located on a geometrical pattern that provides
sixteen source-detector combinations (C1, C2, . . . , C16) having a
selected source-detector separation. The separation may be 2.5 cm
to produce about 1.25 cm average light penetration. (Several
modules with different source-detector separations may be used to
obtain several two dimensional images of different tissue depths.
Alternatively, a single module may include source detector
combinations providing different separations.) The light sources
are 1 W tungsten light bulbs, which emit broad band non-modulated
light. The light detectors are silicon diodes, each equipped with
an interference filter transmitting a 10 nm wide band centered at
760 nm and 850 nm. The 760 nm and 850 nm wavelengths are selected
to detect oxyhemoglobin and deoxyhemoglobin in the examined
tissue.
[0100] Optical module 212 is connected to an analog circuit 202,
which includes a source circuit 204 for controlling sources S1, S2,
. . . S12. Optical module 212 is connected to a detector circuit
206, which controls diode detectors D1, D2, D3 and D4. In general,
imaging system 200 can turn ON each source for a selected period in
the range of 10.sup.-6 sec. to 0.1 sec., and one or several
symmetrically located detectors are turned on simultaneously or
sequentially to collect optical data. Specifically, as provided in
Appendix B, one of sources S1, S2, . . . S12 is turned ON for 500
msec and the emitted light is introduced into the tissue from the
corresponding input port. The introduced photons migrate over
banana shaped paths in the examined tissue to a detection port. The
corresponding detector is triggered 200 msec. after the source and
collects light for 200 msec. Detector circuit 206 receives a
detector signal from the diode detector. Detection circuit 206
enables correction for the dark current/noise that comprises
background light, DC offset of the operational amplifiers,
photodiode dark current, temperature effects on the outputs of
individual components and variations due to changing
environment.
[0101] Imaging system 200 performs data acquisition in four steps
synchronized by its internal oscillator. The first step is
performed by having the light sources OFF. The detector output is
directed to an integrator 216 and integration capacitor 218 is
charged to the dark level voltage. In the second step, the light
source is turned ON and after 200 msec the preamplifier output that
corresponds to the intensity of the detected light is directed to
integrator 216 in a way to charge capacitor 218 with current of
polarity opposite to the polarity of the charging current in the
first step. This is achieved using an appropriate ON/OFF
combination of switches A and B. The voltage of capacitor 218 is
charging to a value that, after 200 msec., represents the total
detected intensity minus the dark level noise signal. In the third
step, both switches A and B are turned OFF to disconnect both the
positive unity gain and the negative unity gain operational
amplifiers (220 and 222). Then, the output of integrator 218 is
moved via switch C to an analog-to-digital converter and the
digital signal is stored in the memory of a computer. In the fourth
step, the switches A, B and C are open and switch D is closed in
order to discharge capacitor 218 through a 47K resistor. At this
point, the circuit of integrator 216 is reset to zero and ready for
the first step of the detection cycle.
[0102] Alternatively, analog circuit 202 may be replaced by a
computer with an analog-to-digital converter and appropriate
software that controls the entire operation of optical module 212.
An algorithm controls the sources and the detectors of optical
module 212 in a similar way as described above. The detected dark
level noise signal is digitally subtracted from the detected
intensity of the introduced light.
[0103] The collected data sets are processed using an imaging
algorithm. The imaging algorithm calculates the blood volume of the
examined tissue for each source-detector combination for each data
set. The imaging algorithm can also calculate the oxygenation of
the examined tissue for each source-detector combination.
[0104] The blood volume or oxygenation images can be subtracted
from "model" images. The blood volume image can be subtracted from
the oxygenation image to create congruence data to localize and
characterize a tissue anomaly. That is, the imaging algorithm
creates an image using the differential image data sets. Prior to
creating the image, an interpolation algorithm is employed to
expand the differential image data set, containing 16 (4.times.4)
data points, to an imaging data set containing 32.times.32 image
points.
[0105] Alternatively, the computer uses a back-projection algorithm
known in computed tomography (CT) modified for light diffusion and
refraction and the banana like geometry employed by the optical
imaging system. In the optical back-projection algorithm, the
probabilistic concept of the "photon migration density" replaces
the linear relationship of ballistically transmitted X-rays, for
the beam representing pixels. The photon migration density denotes
a probability that a photon introduced at the input port will
occupy a specific pixel and reach the detection port. For different
types of tissue, the phase modulation spectrophotometer provides
the values of the scattering and absorption coefficients employed
in the probability calculations. In the image reconstruction
program, the probability is translated into a weight factor, when
it is used to process back-projection. A back-projection algorithm
known in X-ray CT may be used. The back-projection averages out the
values of information that each beam carries with the weighting in
each pixel. A weighting algorithm for creating a photon density
image may be used in the back-projection reconstruction algorithm
mentioned above.
[0106] A method for correcting blurring and refraction used in the
back-projection algorithm was described by S. B. Colak, H.
Schomberg, G. W.'t Hooft, M. B. van der Mark on Mar. 12, 1996, in
"Optical Back-projection Tomography in Heterogeneous Diffusive
Media" which is incorporated by reference as if fully set forth
herein. The references cited in this publication provide further
information about the optical back-projection tomography and are
incorporated by reference as if fully set forth herein.
[0107] Another embodiment of the amplitude cancellation imaging
system 200 uses optical module 14 shown in FIG. 2A. In this
arrangement, four centrally located light sources S1, S2, S3, and
S4 and 21 detectors D1, D2, . . . , D21 provide a multiplicity of
symmetric photon migration paths for each source. For example,
source S1 is turned ON for a period in the range of 10.sup.-6 sec.
to 0.1 sec. The source emits non-modulated light into the examined
tissue. Symmetrically located detectors D1 and D11 are ON
simultaneously to collect introduced photons migrating over
substantially symmetric paths. For normal brain tissue, detectors
D1 and D11 detect light of the same intensity, and thus the
differential signal is zero, i.e., the detected amplitude are
canceled. Imaging system 200 collects the differential data for a
multiplicity of symmetric photon migration paths and generates an
image of the examined tissue. Imaging system 200 may collect
optical data for several wavelengths and generate blood volume
images and blood oxygenation images for the examined tissue.
Amplitude cancellation imaging system 200 may also use a second
identical optical module 14 placed to examine a symmetrical brain
region, for example, the opposite lobe of the brain. The blood
volume or oxygenation images collected for the two symmetric brain
regions may be subtracted to provide a differential image, which
will further emphasize a tissue abnormality located in one brain
region.
[0108] Alternatively, the amplitude cancellation imaging system
uses light modulated at frequencies in the range of 0.1 to 100 kHz.
The system employs the above-described algorithm, but the light
sources emit frequency modulated light and the detectors, each
connected to a lock-in amplifier, detect light modulated at the
same frequency. This lock-in detection may further increase the
signal to noise ratio by eliminating external noise. The detected
light intensities are processed the same way as described above to
image the examined tissue.
[0109] FIGS. 7, 7A and 7B show different embodiments of a cooling
module used with a broad band light source or light guides, where
they are positioned close to the skin. In this arrangement, there
is trapped heat that frequently causes an uncomfortable
temperature. FIG. 7 depicts a cooling module 230, which surrounds
light sources 232A and 232B. Cooling module 230 includes a fan 234
and a set of air passages 236. In a similar design, two fans are
juxtaposed on each side of one or more light bulbs to form an "open
frame" so that the fans blow not only upon the light sources, but
upon the skin itself. The cooling module enables a power increase
on the light sources, but no increase of heat upon the skin itself,
which remains under comfortable conditions.
[0110] FIG. 7A depicts a cooling module 240 for cooling light
guides. Light guides 242 deliver light and heat to the skin. A
cooling ring 244 includes an air inlet 246 and a set of air
passages 248 (or jets) for providing air flow to the irradiation
location. FIG. 7B depicts a cooling module 250 constructed to air
cool a light barrier 252. Light barrier 252 has similar optical
properties as the light bather described in the PCT application
PCT/US92/04153 (published on Nov. 26, 1992 as WO 92/20273), which
is incorporated by reference. This embodiment utilizes the
advantages of the light barrier and enables the use of higher light
intensities. Cooling module 250 includes air inlets 252A and 252B,
which provide air to a set of conduits and openings that deliver
air to the skin near light source 254. Compressed air may also be
used.
[0111] The intensity regulations for delivering continuous
otherwise noncoherent light to the skin often depend on the
temperature rise of the skin itself. For examination of large
tissue volumes or deep tissues (i.e., where there is a large
separation between the optical input and optical detection ports)
relatively large light intensities are needed. Under conditions of
prolonged even low level illumination, the skin may become
uncomfortably warm and may blister. However, the erythemic effects
are much smaller in the NIR, where the delivered heat is a factor,
than they are in UVA and UVB, where cancer-producing damage may
occur (but is not known for the NIR). The effect of the cooling air
is not just convection of warm air away from the skin, but it
enhances the evaporation of perspiration from the skin. Thus, as
soon as the skin temperature rises and perspiration is initiated,
greatly enhanced cooling is obtained with the forced air increasing
the evaporation
[0112] Referring to FIG. 8, an amplitude cancellation imaging
system 260 is used instead of imaging systems 15, 45, 100, 150, or
202. Dual wavelength amplitude cancellation imaging system 260 is
connected to optical module 14 shown in FIG. 2A and includes four
750 nm laser diodes S.sub.1, S.sub.2, S.sub.3, and S.sub.4, four
830 nm laser diodes S.sub.1a, S.sub.2a, S.sub.3a, and S.sub.4a, and
twenty-one silicon diode detectors D.sub.1, D.sub.2, . . . ,
D.sub.21. Each detector is connected to a preamplifier and an
adjustable gain controller that may be used initially for
calibration. The detector outputs are switched by a switch 262 by a
controller 264 so that analog-to-digital converters 266 and 266a
receive 750 nm and 830 nm data, respectively, from two
symmetrically located detectors. A computer 270 stores the detected
values measured for the different combinations using algorithms
employed by the enclosed source code. The computer also generates
images described below. Another type of amplitude detection circuit
was described in FIGS. 11 through 13 and the corresponding
specification of U.S. Pat. No. 5,673,701, which is incorporated by
reference as if fully set forth herein.
[0113] Also referring to FIGS. 8A and 8B, the controller sequences
an oscillator 261 so that each source emits a 50 .mu.sec light
pulse as shown in timing diagram 272. The system sequences through
the various source/detector combinations in approximately one msec,
and averages the imaged data over 8 sec to get a very high signal
to noise ratio. The circuit configuration for one element of
imaging system 260, i.e., 754 nm sources S.sub.1, S.sub.2 and 830
nm sources S.sub.1a, S.sub.2a, and two symmetrically positioned
detectors D.sub.3 and D.sub.11, is shown in FIG. 8A. The light
intensities detected for the symmetrical locations are subtracted
in a digital or analog way. The computer stores all data detected
for the two wavelengths for generating tissue images.
[0114] FIG. 8C shows diagrammatically a single channel 260A of the
time multiplex imaging system 260. Detector D.sub.1 detects light
emitted from light source S.sub.1 emitting light pulses of the
duration of about 50 .mu.sec. The detector signal is amplified and
provided to a sample-and-hold circuit and filter. Detector D.sub.1
is a silicon diode detector that has the detection area of about
4.times.4 mm and includes a pre-amplifier. The filtered signal 272
is provided to an AGC 274, which adjusts the amplitude of the
signal based on a control signal from a personal computer. The
personal computer has normalization amplitudes for the individual
source-detector combinations.
[0115] Amplitude cancellation imaging system 260 is normalized on a
tissue model by detecting signals for the individual
source-detector combinations and appropriately normalizing the
detected signal using the AGC control. The individual
normalization/calibration amplitudes form a baseline image that is
stored in the computer. As described above, the baseline image may
also be acquired on a symmetric tissue region, such as the
contralateral brain tissue for brain tissue examination, or the
contralateral tissue in general for any tissue examination. The
normalization process can be repeated several times to account for
drifts in the individual elements. During the measurement process,
the personal computer can adjust the gain of each AGC 314 based on
the calibration values that account only for the electronic drift.
Then, the defected image is subtracted from the baseline image of
the examined tissue. Alternatively, while collecting the
measurement data on the examined tissue, the measurement image is
subtracted from the baseline image to create the tissue image that
includes any tissue in homogeneities such as a tumor or bleeding.
The sample-and-hold circuit maybe an analog circuit or the
sample-and-hold function, including the filtering, may be performed
digitally.
[0116] FIG. 8D shows diagrammatically an amplitude cancellation
imaging system employing a frequency multiplex method. Amplitude
cancellation system 300 includes 21 oscillators 302 operating a
frequencies in the range of 1 kHz to 100 kHz. Each oscillator 302
drives a light source 304 (for example, a laser diode or LED),
which emits an intensity modulated light into the examined tissue.
Each light detector 306 (for example, a photomultiplier, an
avalanche photodiode PIN detector or a silicon detector) detects
the intensity modulated light and provides a detector signal to an
amplifier 308. The amplified detector signal is provided to a
processing channel 310, which includes a band pass filter 312, an
AGC 314, a lock-in amplifier 316, and a filter 318. Filter 312
filters the detector signal, and AGC 314 adjusts the amplitude
according to the input signal from a personal computer. Lock-in
amplifier 316 receives the amplified signal 315 and a reference
signal 320 from oscillator 302. Lock-in amplifier 312 provides
amplitude signal 317 to filter 318. Processing channel 310 may be
an analog channel or a digital channel.
[0117] In the amplitude cancellation system 310, all light sources
emit light at the same time into a selected tissue region. Each
light source is modulated at a distinct frequency in the range of 1
kHz to 100 kHz. In order to resolve the modulated light signals and
attribute them to the individual light sources, the oscillators
operate at frequencies 1 kHz, 2 kHz, 4 kHz, 8 kHz, 16 kHz, . . . .
Filters 312 and 318 are designed to provide only the detection
signal from a selected light source, and lock-in amplifier 312
provides the amplitude of the signal at the selected frequency.
Frequency multiplex system 300 is calibrated the same way as the
time multiplex system 260, and the normalization/calibration
amplitude values are also stored in the personal computer. The
images are processed as described above.
[0118] All above-described imagers will achieve a higher spacial
resolution of the imaged tissue by increasing the number of sources
and detectors. Furthermore, the sources and detectors may form
various 1 dimensional, 1.5 dimensional, or 2 dimensional arrays as
described in the above-referenced documents.
[0119] Before examination of a selected brain region, the imager is
first calibrated on a brain model. During the examination, the
patient or the attendant holds optical probe 12 over a designated
portion of the head. The mirror image region on the contralateral
brain region may also be recorded. The images can be acquired by
taking advantage of a priori information obtained by X-ray
tomography, an MRI or PET scan. The optical images were created
using a back-projection algorithm with or without correction for
non-ballistic photon propagation (i.e., tissue absorption or
scattering) as provided in Appendix A-5. The images may be
displayed in the format of the left brain hemisphere data minus the
model data, the right brain hemisphere data minus the model data,
for each wavelength (e.g., 750 and 830 nm). Alternatively, the
model calibration may be performed by adjusting the detector gains
prior to the brain tissue measurements. Furthermore, the images may
be the differential between the right brain region and the left
brain region, for each wavelength, to emphasize any tissue
difference, such as a suspicious structure, which is unlikely
located symmetrically in both brain regions.
[0120] The optical images may also be processed to image blood
volume and blood oxygenation of the examined tissue of each brain
region. The blood volume image is the sum of 0.3 times the 750 nm
data and 1.0 times the 830 nm data. The blood deoxygenation image
is the difference of the 750 nm and the 830 nm data. The above
coefficients were derived from blood tests in model systems. The
images have the highest specificity and sensitivity when the
contralateral brain region data is used as a baseline and both the
blood volume data and the hemoglobin deoxygenation data is imaged
and positionally compared.
[0121] The blood volume and hemoglobin deoxygenation images provide
an important tool in characterizing a suspicious anomaly in the
examined brain. While the blood volume and hemoglobin deoxygenation
images, as well as the single wavelength images, are useful in
locating an abnormal tissue region (i.e., detecting the abnormal
structure), these images are also used to characterize the
metabolism or pathology of the suspicious tissue anomaly.
Specifically, an increased blood volume signal is observed with
respect to the adipose tissue background due to the increased
vascularity of a tumor as a consequence of angiogenetic factors.
These factors include actively metabolizing regions and
necrotic/apoptotic regions of the tumor. On the other hand, the
hemoglobin deoxygenation signal is related to metabolic intensity.
That is, the balance between oxygen delivery and oxygen uptake,
which in tumors is usually balanced in favor of oxygen uptake
exceeding oxygen delivery. The increased oxygen uptake occurs
particularly for those tumors that are aggressively growing, and
may as well be metastatic.
[0122] By selecting an appropriate wavelength, or several
wavelengths, sensitive to an optically active tissue property, the
imaging system can non-invasively characterize a tissue anomaly.
The above-mentioned wavelengths are sensitive to hemoglobin and
hemoglobin oxygenation, but other wavelengths sensitive to
absorption by any tissue constituent may be used. Furthermore, an
optical contrast agent (e.g., cardiogreen, indocianine green) may
be injected intravenously. The imaging system will then use a
wavelength sensitive to the administered contrast agent. The
regions of increased blood volume will also have a higher content
of the contrast agent.
[0123] Alternatively, differences in tissue scattering may be
imaged. Due to differences in the optical refractive index,
different types of tissue and different tissue solutes scatter
light differently. The above-described imaging systems are also
sensitive to scattering changes. The imaging system may use a
wavelength that does not exhibit absorption changes for different
types of tissue and different tissue solutes, but exhibits
differences in scattering.
[0124] The non-invasive characterization of the brain tissue may be
performed by combining the data from the above described images.
For example, a two dimensional data chart may display blood volume
(i.e., vasculogenesis) vs. blood deoxygenation (i.e.
hypermetabolism) for a "suspicious structure" using the
contralateral brain region data as a reference, or using the model
data as a reference.
Quantitation of Co-registration of Several Images
[0125] In principle, vasculogenesis (blood volume) and
hypermetabolism (tissue hypoxia) occur in similar and often
identical tissue volumes. The vascular volume signal can be
reinforced by the blood volume signal. I can evaluate the
congruence of the two images in order to further reinforce the
identity of a suspicious region, for example, by quantitation of
the congruence evaluated pixel by pixel. The first step is the
normalization of the two images to equalize the maximum signals.
Appropriate computer programs exist for selecting the area and
obtaining the integrated value for the spatial congruence residual
and for the blood volume signal. Then, subtraction pixel-by-pixel
gives an image that provides a residual on which to base an
estimate of the congruence of the two shapes, blood volume and
deoxygenation. This has been carried out for those shapes which
appear by inspection to be congruent and the integral of the
residual non-zero pixels is compared to the total signal. A simpler
procedure is to take the maximum value of the difference and divide
it by the maximum value of the normalized value for the two
images.
[0126] Referring to FIG. 9, a "four" dimensional graph may be used
to summarize images of suspicious regions (FIG. 9 is only a
hypothetical summary and not actual brain tissue data). The blood
volume (Volts) is plotted on the abscissa and deoxygenation (Volts)
on the ordinate. The measured size of image is depicted as the
circle diameter and the percentage congruence between the blood
volume image and the deoxygenation image is shown by a color scale.
Color coding the percentage of congruence signals may be given a
color scale based on the following formula:
1 - ( maximum overlap residual maximum blood volume signal )
.times. 100 ##EQU00002##
The "four" dimensional diagram is summarized as follows: 1. The
size of the image of suspicious mass (plotted as one half its
longest dimension). 2. The congruence of blood volumes and blood
deoxygenation plotted in a color. 3. The blood volume in the
congruent region measured in volts (scale of the abscissa). 4.
Blood deoxygenation in the congruent region (scale of the
ordinate).
[0127] A brain model was constructed to test the above-described
imaging techniques and calibrates the imaging systems. The model
included a 4.times.8.times.8 millimeter cellophane chamber
connected to a source of oxygenated or deoxygenated blood. The
chamber was placed 2.5 cm deep within the solid brain model with
the absorption coefficient .mu..sub.a=0.04 cm.sup.-1, and the
scattering coefficient .mu..sub.a'=10 cm.sup.-1. The chamber was
filled with blood of appropriate concentrations and could be moved
to various positions within the model. An accurate determination of
relative changes of blood concentration was obtained in error of 2
.mu.M from 50 to 160 .mu.M (covering the physiological range). The
position errors of .+-.2 mm was determined by comparing the image
obtained from the back-projection algorithm with the real position.
The phased array system has shown a very high positional accuracy
and object detection at a depth of 3 cm.
Functional Imaging
[0128] In another important embodiment of the invention, the
above-described imaging systems are used to image the functional
activity of a selected brain region. The functional imaging alone,
or in combination with the above described structural imaging or
tissue characterization imaging, detects a brain anomaly. A
functional imaging system includes one of the above described
optical imaging systems and a stimulation unit that is constructed
to stimulate a specific neural function of the examined subject.
The optical module is placed to examine the stimulated tissue
region (for example, on the parietal bone of the skull to observe
the surface of the parietal cortex). The stimulator, operating in
unity with the imaging system, emits mechanical, electrical,
thermal, sound or light signals designed to stimulate selected
neural activity in the tissue region probed by visible or infrared
light. The neural activity is induced by sensory stimuli, such as
visual, auditory, or olfactory stimuli, taste, tactile
discrimination, pain and temperature stimuli, or proprioceptive
stimuli. The functional imaging is also described in U.S. Pat. No.
5,853,370 issued Dec. 29, 1998, which is incorporated by reference
as if fully set forth herein.
[0129] The functional imaging can examine and image numerous
centers of the neural activity. For example, the optical module may
be attached to the temporal bone of the skull to examine the
surface of the temporal lobe. Then, the stimulator stimulates the
auditory function while the optical tomography system images
neurofunctional activity of the auditory area of the temporal lobe.
The optical system may also image the auditory association cortex
of Wernicke in the temporal lobe before and after stimulation by
the stimulator.
[0130] Another neurofunctional examination includes placing the
optical module to the frontal bone of the skull to examine the
frontal lobe. Then, the stimulator stimulates the motor speech
function while the optical tomography system images neurofunctional
activity of the motor speech area of Broca before and during
stimulation. Additionally, the optical module may be attached to
the right parietal bone to examine the neurofunctional activity of
the general sensory area before and during stimulation of pain, hot
or cold sensation, or vibrational sensation on the left
extremities, and vice versa.
[0131] Alternatively, the stimulation unit is constructed to induce
physiologic and pathologic reflexes in the cerebral or spinal
tissue. The stimulation unit stimulates pupillary reflexes, corneal
reflexes, oculocephalic reflexes, oculovestibular reflexes, deep
tendon reflexes, abdominal reflex, cremasteric reflexes, postural
reflexes, gag reflex, infantile reflexes (such as blinking reflex,
cochleopalpebral reflex, palmar grasp reflex, digital response
reflex, rooting reflex, Galant's reflex, tonic neck reflex, Perez
reflex, startle reflex).
[0132] The stimulator stimulates a selected region of the nervous
system. The corresponding neurologic impulses, transmitted by the
neurons, are detected and imaged at different points of their
paths, for example, in the nerves, in the spinal cord, in the
thalamus, or in the cerebral cortex. For example, when the
stimulator causes a cold or hot stimulation on the little finger of
the left hand, this thermal stimulation produces impulses that
travel in the right lateral spinothalamic tract of the cervical
spinal cord, to the thalamic sensory nuclei and end in the right
postcentral gyrus of the parietal lobe.
[0133] In a clinical study, provided here only for illustration
purposes, the optical tomography system was used to image the
cognitive activity in the prefrontal cortex of a subject. High
school students together with teachers and three University of
Pennsylvania undergraduate mentors underwent a study using 50 MHZ
phased array imaging system 15, shown in FIG. 3, to explore the
repeatability of their cognitive responses and their geometric
distribution on the forehead. A large number of cognitive tests
were studied by selected groups of four. The simplicity and
versatility of the backwards spelling was selected by the student
group. Thus, each member of the team was tested by the other
members for three to four episodes of rest (30 sec). The students
spelled five letter words backwards (usually 5) for 30 sec, then
rested for 30 sec, etc. Each time a new word was used from a word
list unknown to the subject. The subject was not scored on the
correctness of their response, and as soon as one spelling had been
achieved, another word was given, no prompting was involved in the
protocol.
[0134] The total population studied exceeded 18 participants, but
as shown here the prefrontal data are highly individualized and not
suitable for global averaging. Instead, for individual subjects,
extensive longitudinal studies made in 20 days with 125 tests of 5
subjects were completed, and the results displayed here are based
on .about.25 studies of each of the five and .about.50 more tests
were conducted on the remaining 14 students. There was no selection
of subjects in this study
[0135] The back-projection images were processed using Matlab
software to produce the phase and amplitude images. The phase image
was robust and unambiguous. The data presentations are in the form
of histogram displays of dated data accumulated over six weeks. The
blood volume responses were scored by their position on the
forehead, being divided up into nine areas of a 4 cm.sup.2 area.
Responses>20.degree. in the particular areas were used to create
histograms shown below, indicating the frequency of responses in
particular areas for particular individuals. These voxels could
contain at least one and possibly two responses in view of our
.about.1 cm.sup.2 resolution. However, choice of nine areas seems
adequate at present.
Brain Studies: Parietal Region
[0136] FIGS. 10 and 10A show an experimental optical image obtained
by the imaging system of FIG. 3 with contralateral, parietal finger
touching as a stimulation. These figures illustrate the resolution
obtainable with contralateral, parietal finger touching as a
stimulation. The more intense part of the image is 1.5 by 0.7 cm.
The intensity is profiled on the right hand side of the figure, and
the peak is approximately 4 mm in diameter. Importantly, the phase
scale indicates over 40.degree. phase shift for the peak of the
parietal stimulation with a noise background of less than a few
degrees of phase, confirming the very high signal to noise ratio of
the phase cancellation system shown in FIGS. 3 through 5.
Independent recordings of the amplitude changes measure the
absorbance increase in the focal region, which is due to increased
blood concentration. This wavelength will also register changes of
hemoglobin oxygenation, which may accompany the blood concentration
increase. Thus, the phase shift signal is a composite of increased
absorbance due to blood concentration increase and a smaller
decrease of absorbance due to replacement of deoxygenated blood by
more oxygenated blood, the net change being increased absorbance
and a shortening of the optical pathlength or phase delay.
[0137] FIG. 11. shows co-registration of optical and NMR signals in
sensory motor simulation. The ability to co-register the optical
images (PAI) and the MRI images is within the accuracy of the
optical method for the blood volume or oxygenation changes. Thus,
the maximal blood concentration increase as measured by the phased
array images is congruent with maximal decrease of deoxyhemoglobin
as measured by the fMRI (FIG. 11). However, the shape is elliptical
rather than rectangular, 2 cm.times.1 cm. Such differences may be
verified in future studies in which the fractional deoxygenation of
hemoglobin and the blood concentration are observed rather than the
incremental change of deoxyhemoglobin.
[0138] FIGS. 12A through 13D show images detected by the above
described cognitive study performed by high school students,
wherein the optical tomography system was used to image the
cognitive activity in the prefrontal cortex of a subject. The nine
source, 4 detector system operating at 780 nm on a 9 cm.times.4 cm
optical pad was located between the eyebrow and the hair line. The
optical data was detected while the subject was performing backward
spelling and at rest. Referring to FIGS. 13A through 13D, a second
subject (KW) showed image on the other side of midline, 1.5.times.1
cm of varying intensity, and in the fourth repetition, a facing out
of the pattern and an emergence of a pattern similar to that of the
first subject. Referring to FIGS. 12A, 12B and 12C, the responses
of one subject (DIPTI) to repeated tests are almost identical in
position and in intensity, roughly a 1.5.times.3 cm area along the
forehead.
[0139] FIGS. 14A and 14B are histograms of the positions on the
forehead for the two subjects DIPTI and KW, respectively.
Variability in response was observed, particularly in the younger
members of the group who showed changes of position of the response
maximum. For this reason it was considered that histograms of the
position of the response on the forehead would be a better
representation of the individual responses. Signals above
20.degree. were selected and their positions were scored in nine
spaces (FIG. 5) which are abbreviated, bottom right (BR), lower
left (LL), center right (CR), upper middle (UM), etc. Referring to
FIGS. 14A and 14B, the two individuals exhibit different portions
of the prefrontal cortex in responding to the task of spelling
backwards. FIG. 14A employs mainly the center left region, and the
upper middle region to a small extent. FIG. 14B exhibits the upper
middle region to a much greater extent than the center left region.
These two cases are exemplary of the many subjects studied, and
define what may be a novel and important element of pre-frontal
cognitive response. Note that both subjects nearly always responded
in the dominant position over the four test intervals.
Neonates Imaging
[0140] In another clinical study, provided here for illustration
only, this technique has been applied to the functional activation
of pre- and full-term, non-white neonates as shown in FIGS. 15A and
15B. In this case, prefrontal activity could not yet be tested,
instead, the sensorimotor region was tested. The probe was held on
the head of the infant for 30 sec to acquire a rest image. Touching
the right finger evokes a response imaged over 30 sec. The
magnitude is large; a 100.degree. phase shift and an image
approximately 2 cm in size in the contralateral hemisphere is
indicated. When, however, the right finger is touched, a similar
area of response is obtained in the contralateral hemisphere
displaced laterally. In addition, the right leg kicked
spontaneously.
[0141] The examined infant was a 26 week gestation, 1 kilogram,
premature infant, which was studied at age four weeks. The
stimulation, in this case, was touching the baby's right finger
(FIG. 15B). A distinct image was obtained somewhat on the
contralateral side (approximately 1.times.1.5 cm in size), and of a
large magnitude (over 100.degree. phase shift). Stimulation of the
left finger (FIG. 15B A) gave a distinctive image of the same size
but laterally displaced in the right hemisphere. At the same time,
the spontaneous kicking was clearly resolved and the image was
displaced laterally. Thus, images of the voluntary and involuntary
responses of the pre-term neonate brain were obtained.
[0142] While these data are preliminary, they are remarkable for
the large amplitude of the responses, in fact as large as observed
in the high school student population. This large response of the
neonates as compared with the high school population is due to in
part to their thinner skull and smaller CSF space. This works well
for the detection of diminished response of dysfunctional infants
who may have had hypoxia/ischemia or other traumatic events either
pre-partum or intra-partum.
[0143] The formation of a well resolved image of brain function
using multiple light sources and detectors in the NIR region, with
either continuous or modulated light, opens up a fertile field of
study of visual sensorimotor and prefrontal functions in adults,
full- and pre-term neonates. The methods are intentionally
over-simplified to afford a fast, simple, straight-forward, safe
and affordable method for studying brain function. Sufficient
studies of the high school students were made to ensure validity of
the stability of the signal and of the position of the maximal
response for a given individual and of the variability of the
response among individuals. The striking results are the plasticity
of the response of the pre-frontal area observed in the series of
the high school student studies on the one hand, and ease with
which evoked signals in the parietal region are obtained with the
pre- and full-term infant on the other. Thus, we present a
preliminary report on these studies in order to stimulate further
research, here and elsewhere.
[0144] These results make practical and affordable for large
populations the complex technology of measuring brain function.
While lacking the resolution of MRI or the chemical specificity of
PET, it has the capability of multi-wavelength operation to give
enhanced sensitivity for oxy-, deoxy-hemoglobin and for light
scattering changes. More importantly, the method opens up new
fields of study of the human population, in adults under conditions
of simulated or real stress that may have important effects upon
functional performance, or in other cases, where the subject cannot
be well controlled as in the full- and pre-term neonate and those
not fully responsive due to accidents or due to disease.
Higher Resolution Images
[0145] In previous studies, optical tomography has attempted to
mimic the X-ray image by a 2D projection of absorbance usually in 2
planes. The success of this techniques is based upon the ability of
the radiologist to identify the structures of either scattering or
absorbing material that differ from the normal tissue. However, a
high resolution is required to delineate such structural features
on which identification of malignant tissue is usually based. High
resolution is time intensive as well as apparatus intensive, i.e.,
numerous source detector combinations are required to achieve
resolution comparable to PET/MRI. In the above systems, imaging
resolution is employed mainly to increase the signal to noise ratio
in quantifying optical properties of the tumor with respect to
normal tissue or a model of a normal tissue. However, the blood
volume, oxygenation and deoxygenation data collected by the optical
systems do not depend critically upon high resolution imaging.
[0146] An optical system with an increased number of sources and
detectors will render higher spatial resolution. Furthermore, a
larger source-detector separation (i.e., the input port to
detection port separation) achieves deeper penetration of the
introduced optical radiation. By using selected separation values,
the above-described imaging systems can collect three-dimensional
optical data that are used for three dimensional
reconstruction.
[0147] Additional embodiments are within the following claims:
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