U.S. patent application number 13/919809 was filed with the patent office on 2013-10-24 for image-guided thermotherapy based on selective tissue thermal treatment.
The applicant listed for this patent is Tomophase Corporation. Invention is credited to Peter E. Norris, Andrei Vertikov.
Application Number | 20130282083 13/919809 |
Document ID | / |
Family ID | 43527741 |
Filed Date | 2013-10-24 |
United States Patent
Application |
20130282083 |
Kind Code |
A1 |
Vertikov; Andrei ; et
al. |
October 24, 2013 |
IMAGE-GUIDED THERMOTHERAPY BASED ON SELECTIVE TISSUE THERMAL
TREATMENT
Abstract
Devices and techniques for thermotherapy based on optical
imaging.
Inventors: |
Vertikov; Andrei; (Westwood,
MA) ; Norris; Peter E.; (Cambridge, MA) |
|
Applicant: |
Name |
City |
State |
Country |
Type |
Tomophase Corporation |
Burlington |
MA |
US |
|
|
Family ID: |
43527741 |
Appl. No.: |
13/919809 |
Filed: |
June 17, 2013 |
Related U.S. Patent Documents
|
|
|
|
|
|
Application
Number |
Filing Date |
Patent Number |
|
|
12770713 |
Apr 29, 2010 |
8467858 |
|
|
13919809 |
|
|
|
|
61173921 |
Apr 29, 2009 |
|
|
|
Current U.S.
Class: |
607/100 ;
607/96 |
Current CPC
Class: |
A61B 2018/00023
20130101; A61B 18/12 20130101; A61B 2017/00084 20130101; A61B 90/30
20160201; A61B 5/14532 20130101; A61B 18/1815 20130101; A61N 5/00
20130101; A61B 18/24 20130101; A61B 2018/1861 20130101; A61B 90/36
20160201; A61B 5/14558 20130101; A61B 90/361 20160201 |
Class at
Publication: |
607/100 ;
607/96 |
International
Class: |
A61N 5/00 20060101
A61N005/00 |
Claims
1. A device for thermotherapy, comprising: a catheter including a
working channel configured for insertion into a passage of a body
to reach a target tissue inside the body; an optical imaging module
comprising (1) an imaging optic fiber having a portion inserted
into the working channel and (2) an optical probe head coupled to
an end of the imaging optic fiber and located inside the working
channel, the optical imaging module operable to direct probe light
to and collect reflected light from the target tissue in the body
through the imaging optic fiber and the optical probe head and to
obtain imaging information of the target tissue from the collected
reflected light; a thermotherapy module having a power delivery
waveguide having a portion inserted into the working channel to
deliver thermal energy to the target tissue; a control unit that
controls the optical imaging module to extract the imaging
information from the collected reflected light, to obtain a spatial
distribution of diseased locations of the target tissue, and to
obtain a temperature map of the target tissue for thermotherapy
based on the spatial distribution of the diseased locations of the
target tissue, the control unit controlling the thermotherapy
module to control a location and an amount of thermal energy
delivery to each of the diseased locations based on the temperature
map to perform thermotherapy.
2. The device as in claim 1, comprising: a liquid cooling unit
coupled to the catheter to direct a cooling liquid that cools a
surface of the target tissue.
3. The device as in claim 1, wherein the imaging optic fiber is
structured to support light in a first propagation mode and a
second, different propagation mode, and the optical imaging module
includes: a light source to produce the probe light, wherein the
imaging optic fiber receives and guides the probe light in the
first propagation mode, wherein the optical probe head is coupled
to the imaging optic fiber to receive the light from the imaging
optic fiber and to reflect a first portion of the light back to the
imaging optic fiber in the first propagation mode and direct a
second portion of the light to a target location of the target
tissue, the probe head collecting reflection of the second portion
from the target location and exporting to the imaging optic fiber
the reflection as a reflected second portion in the second
propagation mode; an optical differential delay unit to produce and
control a relative delay between the reflected first portion and
the reflected second portion received from the imaging optic fiber
in response to a control signal; a detection module to receive the
reflected first portion and the reflected second portion from the
imaging optic fiber and to extract information of the target area
carried by the reflected second portion; and an imaging control
unit, which produces the control signal to the optical differential
delay unit, to set the relative delay at two different bias values
to select a layer of material inside the target area to measure an
optical absorption of the selected layer.
4. The device as in claim 3, wherein the detection module
comprises: an optical device to direct light in the first
propagation mode along a first optical path and light in the second
propagation mode along a second, different optical path; a first
optical element in the first optical path to separate light into a
first set of different beams at different wavelengths; a plurality
of first light detectors to respectively receive and detect the
first set of different beams from the first optical element; a
second optical element in the second optical path to separate light
into a second set of different beams at the different wavelengths;
and a plurality of second light detectors to respectively receive
and detect the second set of different beams from the second
optical element.
5. The device as in claim 4, wherein the first and second optical
elements are optical gratings.
6. The device as in claim 3, wherein the detection module comprises
a digital signal processor to process information of the target
area in the reflected second portion and to generate spectral
absorbance data of the target area.
7. The device as in claim 3, wherein the optical differential delay
unit comprises: a mode splitting unit to separate received light
into a first beam in the first propagation mode and a second beam
in the second propagation mode; and a variable optical delay
element in one of the first and the second beams to adjust an
optical delay between the first and the second beams in response to
the control signal.
8. The device as in claim 3, wherein the first and second
propagation modes are two orthogonal polarization modes supported
by the probe optic fiber, and wherein the detection module
comprises: an optical detector; and an optical polarizer to receive
and mix the reflected first and second portions to produce an
optical output to the optical detector.
9. The device as in claim 3, wherein the optical imaging module
further comprises: a plurality of light sources emitting light at
different wavelength bands centered at different wavelengths as the
probe light into the probe optic fiber, wherein the optical probe
head reflects a first portion of the probe light back to the
imaging optic fiber in the first propagation mode and directs a
second portion of the probe light to the target area, and wherein
the probe head collects reflection of the second portion from the
target area and exports to the imaging optic fiber the reflection
as a reflected second portion in a second propagation mode
different from the first propagation mode; an optical differential
delay unit to produce and control a relative delay between the
reflected first portion and the reflected second portion received
from the single waveguide in response to a control signal; a
detection module to receive the reflected first portion and the
reflected second portion and to extract information of the target
area carried by the reflected second portion; and a probe control
unit, which produces the control signal to the optical differential
delay unit, to set the relative delay at two different bias values
to select a layer of material inside the target area to measure an
optical absorption of the selected layer at each and every
wavelength from the different light sources.
10. The device as in claim 3, wherein the optical imaging module
further comprises: a plurality of tunable laser sources emitting
light within different wavelength bands centered at different
wavelengths as the probe light to the probe optic fiber, wherein
the imaging optic fiber is configured to receive and guide the
probe light at the different wavelength bands in a first
propagation mode, wherein the probe head is to reflect a first
portion of the probe light back to the imaging optic fiber in the
first propagation mode and direct a second portion of the light to
the target area, and wherein the probe head collects reflection of
the second portion from the target area and exports to the imaging
optic fiber the reflection as a reflected second portion in a
second propagation mode different from the first propagation mode;
a detection module to receive the reflected first portion and the
reflected second portion in the waveguide and to extract
information of the target area carried by the reflected second
portion; and a probe control unit to tune each tunable laser in a
corresponding laser emitting wavelength band to obtain absorption
measurements of the target area at different wavelengths within
each corresponding wavelength band.
11. The device as in claim 1, comprising: a liquid cooling unit
coupled to the catheter to direct a cooling liquid that cools a
surface of the target tissue, the liquid cooling unit including a
balloon that receives the cooling liquid and is located in contact
with the surface of the target tissue.
12. A method for thermotherapy, comprising: directing an imaging
optical beam to a target tissue to obtain image information;
processing the obtained image information of the target tissue to
obtain a spatial distribution of the diseased locations of the
target tissue; generating a temperature map of the target tissue
for thermotherapy based on the spatial distribution of the diseased
locations of the target tissue; and controlling the thermal energy
delivery to each of the diseased locations and cooling at the
surface of each diseased locations based on the temperature map to
perform the thermotherapy.
Description
CROSS-REFERENCE TO RELATED APPLICATIONS
[0001] This patent document is a continuation application of U.S.
patent application Ser. No. 12/770,713, filed on Apr. 29, 2010,
which claims the benefit of U.S. Provisional Patent Application No.
61/173,921, entitled "Image-Guided Optical Thermotherapy [IGOTT]
based on Selective Tissue Thermal Treatment Using a Combination of
Optical Tissue Imaging, Targeted Energy Deposition and Thermal
Mapping," and filed on Apr. 29, 2009. The entire contents of the
before-mentioned patent applications are incorporated by references
as part of the disclosure of this application.
BACKGROUND
[0002] This document relates to devices and techniques for treating
tissues by thermotherapy.
[0003] Thermotherapy is treatment of a diseased tissue by heat.
When the amount of thermal energy absorbed by a diseased tissue
exceeds a certain threshold, a desired therapeutic effect in the
diseased tissue can be achieved to lessen or mitigated the disease
condition of the tissue. Thermotherapy for treatment of tissue can
use radio frequency (RF) energy, microwave energy, laser radiation
and ultrasound as heating energy sources.
SUMMARY
[0004] This document describes devices and techniques for
thermotherapy based on optical imaging. In one aspect, a method for
thermotherapy includes directing an imaging optical beam to a
target tissue to obtain image information; processing the obtained
image information of the target tissue to obtain a spatial
distribution of the diseased locations of the target tissue;
generating a temperature map of the target tissue for thermotherapy
based on the spatial distribution of the diseased locations of the
target tissue; and controlling the thermal energy delivery to each
of the diseased locations and cooling at the surface of each
diseased locations based on the temperature map to perform the
thermotherapy.
[0005] This and other aspects, features, associated advantages, and
implementation variations are described in detail in the attached
drawings, the description, and the claims.
BRIEF DESCRIPTION OF THE DRAWINGS
[0006] FIG. 1 shows an example of a conventional optical sensing
device based on the well-known Michelson interferometer with
reference and sample beams in two separate optical paths.
[0007] FIG. 2 shows one example of a sensing device according to
one implementation.
[0008] FIG. 3 shows an exemplary implementation of the system
depicted in FIG. 2.
[0009] FIG. 4 shows one exemplary implementation of the probe head
and one exemplary implementation of the polarization-selective
reflector (PSR) used in FIG. 3.
[0010] FIGS. 5A and 5B illustrate another exemplary optical sensing
system that use three waveguides and a light director to direct
light in two modes to and from the probe head in measuring a
sample.
[0011] FIG. 6 illustrates the waveform of the intensity received at
the detector in the system in FIGS. 5A and 5B as a function of the
phase where the detected light intensity exhibits an oscillating
waveform that possesses a base frequency and its harmonics.
[0012] FIG. 7 shows one exemplary operation of the described system
in FIG. 5B or the system in FIG. 3 for acquiring images of optical
inhomogeneity.
[0013] FIGS. 8A and 8B illustrate one exemplary design of the
optical layout of the optical sensing system and its system
implementation with an electronic controller where light in a
single mode is used as the input light.
[0014] FIG. 9 shows another example of a system implementation
where the optical probe head receives light in a single input mode
and converts part of light into a different mode.
[0015] FIGS. 10A and 10B show two examples of the possible designs
for the probe head used in sensing systems where the input light is
in a single mode.
[0016] FIG. 11 shows one implementation of a light director that
includes a polarization-maintaining optical circulator and two
polarization beam splitters.
[0017] FIG. 12 illustrates an example of the optical differential
delay modulator used in present optical sensing systems where an
external control signal is applied to control a differential delay
element to change and modulate the relative delay in the
output.
[0018] FIGS. 12, 12A and 12B illustrate two exemplary devices for
implementing the optical differential delay modulator in FIG.
12.
[0019] FIGS. 13A and 13B illustrate two examples of a mechanical
variable delay element suitable for implementing the optical
differential delay modulator shown in FIG. 12B.
[0020] FIG. 14A shows an exemplary implementation of the delay
device in FIG. 12B as part of or the entire differential delay
modulator.
[0021] FIG. 14B shows a delay device based on the design in FIG.
14A where the mirror and the variable optical delay line are
implemented by the mechanical delay device in FIG. 13A.
[0022] FIG. 15 illustrates an optical sensing system as an
alternative to the device shown in FIG. 5B.
[0023] FIG. 16 shows a system based on the design in FIG. 2 where a
tunable filter is inserted in the input waveguide to filter the
input light in two different modes.
[0024] FIG. 17 shows another exemplary system based on the design
in FIG. 8A where a tunable filter is inserted in the input
waveguide to filter the input light in a single mode.
[0025] FIG. 18 illustrates the operation of the tunable bandpass
filter in the devices in FIGS. 16 and 17.
[0026] FIG. 19A illustrates an example of a human skin tissue where
the optical sensing technique described here can be used to measure
the glucose concentration in the dermis layer between the epidermis
and the subcutaneous layers.
[0027] FIG. 19B shows some predominant glucose absorption peaks in
blood in a wavelength range between 1 and 2.5 microns.
[0028] FIG. 20 illustrates one exemplary implementation of the
detection subsystem in FIG. 3 where two diffraction gratings are
used to separate different spectral components in the output light
beams from the polarizing beam splitter.
[0029] FIGS. 21 and 22 shows examples of optical sensing devices
that direct light in a single mode to the optical probe head and
direct output light from the probe head in the same single
mode.
[0030] FIG. 23 shows an example of a design for the optical probe
head for the devices in FIGS. 21 and 22 where the optical probe
head does not change the mode of light.
[0031] FIG. 24 illustrates two selected surfaces underneath a
surface of a body part in optical spectral absorbance mapping
measurements.
[0032] FIGS. 25, 26 and 27 show examples of devices that use
multiple light sources at fixed center emitting wavelengths for
spectral absorbance mapping measurements.
[0033] FIG. 28 shows one example of an optical multiplexer to
combine beams from different light sources into a common waveguide
or optical path.
[0034] FIGS. 29A and 29B show another example of an optical
multiplexer with dichroic filters to combine beams from different
light sources into a common waveguide or optical path and the
spectral properties of the dichroic filters.
[0035] FIG. 30 shows an exemplary device that uses multiple light
sources at fixed center emitting wavelengths for spectral
absorbance mapping measurements, where an optical switch is used to
sequentially direct different beams from different light sources
into a common waveguide or optical path.
[0036] FIG. 31 illustrates an example for using different beams at
different wavelengths to detect an absorption feature in a sample
in spectral absorbance mapping measurements.
[0037] FIG. 32 shows an exemplary device that uses multiple tunable
light sources for spectral absorbance mapping measurements.
[0038] FIG. 33 shows an example of an integrated system that
combines an X-ray CT scan module, a reference cross-sectional
tissue imaging module, and a laser treatment module to provide a
complete diagnostic and treatment platform for treating lung
cancer.
[0039] FIG. 34 shows one exemplary use of the system in FIG. 33 in
detecting and treating lung cancer.
[0040] FIG. 35 shows a tubular unit or sheath for holding the probe
fiber and the waveguide together as a single unit inserted inside
the working channel shown in FIG. 33.
[0041] FIG. 36 shows an example of a thermotherapy device based on
optical imaging and controlled delivery of thermal energy guided by
the optical imaging.
[0042] FIG. 37 shows an example of a catheter design with a cooling
mechanism for thermotherapy.
[0043] FIGS. 38A and 38B show another example of a catheter design
for thermotherapy.
[0044] FIG. 39 shows an example method of thermotherapy based on
optical imaging.
DETAILED DESCRIPTION
[0045] The thermotherapy techniques and devices described in this
document use an optical imaging mechanism to obtain images of a
target tissue and to obtain spatial distribution of diseased
locations in the target tissue by processing the images of the
target tissue. Based on the spatial distribution of diseased
locations in the target tissue, a temperature map of the target
tissue for thermotherapy can be generated. This temperature map is
then used as a guide to control the delivery of the therapeutic
thermal energy to each of the diseased locations in the target
tissue to perform the thermotherapy. In addition, the surface at
each diseased location is cooled in a controlled manner to minimize
undesired damage to the surface tissue and the surrounding
tissues.
[0046] The imaging of the target issue is performed in real time
during the thermotherapy process and the imaging information is
used as a guide to precisely deliver the therapeutic thermal energy
to each diseased location, making sure that adjacent tissue is not
damaged while targeted tissue is treated. The real time imaging
guidance provides accurate mapping of targeted tissue and meets the
need for selective thermal treatment of specific tissue and for
achieving high efficacy of subsequent heat treatment.
[0047] The optical imaging techniques described in this document
can be implemented in ways to map the targeted tissue for the case
of internal organs when tissue to be treated is located below the
surface. For example, in the case of bronchial thermoplasty [BT],
the targeted tissue is airway smooth muscle (ASM) located inside
the bronchial lumen (e.g., about 3-10 mm in diameter) and separated
from interior of the lumen by other tissue layers, e.g. epithelium
or submucosa. The present optical imaging techniques provide needed
sub-surface imaging capability and can be used to obtain imaging
data in small caliber airways. The present optical imaging
techniques can also be implemented to provide the spatial
resolution necessary to visualize airway microstructures, e.g.,
approximately 10 microns in some cases. The optical imaging guides
the application of thermal energy and thus can greatly reduce
complication risks in comparison to unguided delivery of the
thermal energy in other thermotherapy devices and procedures.
[0048] Notably, there is often large variability in the diseased
tissue location from patient to patient for a given procedure and
large variability in the diseased tissue location between different
procedures. For example, in contrast with tumor tissues typically
located in epithelium layer, i.e. on the (internal surface) of a
lumen, ASM is located below the surface and is not uniformly
distributed within the airway. Application of thermal energy
without precise guidance based on the actual image information of
the target tissue may produce poor overlap of resultant temperature
distribution with the targeted tissue, causing overheating of
adjacent tissue and/or increasing procedural duration. The optical
imaging techniques described here can be used to control the
spatial distribution of heating energy to generate pre-determined
temperature profile that closely matches the targeted tissue for
effective thermotherapy.
[0049] In addition, tissues tend to have significant variability of
thermal, optical and electrical properties and this can cause
significant variability in the temperature profiles generated by
the thermotherapy even for reproducible spatial distribution of
heating energy. The optical imaging mechanism in the present
thermotherapy devices is used to measure, in-vivo, actual
temperature profiles to provide procedural feedback for real-time
focused heat deposition.
[0050] The disclosed image-guided thermotherapy techniques can be
used to provide selective thermal treatment of airway tissue for
downstaging the severity of persistent asthma. Such image-guided
thermotherapy techniques can be implemented to reduce the number of
physician and ER visits, as well as hospitalizations of severely
asthmatic patients. As a result, the clinical outcome and quality
of life of the severely asthmatic patient can be improved and to
reduce healthcare cost. A severe asthmatic is the asthma that is
poorly controlled by inhaled anti-inflammatory or cortico-steroid
drugs. Severe asthmatics are responsible for approximately 1.9
million ER visits and hospitalizations annually. The cost
differential between severe asthmatics and mild chronic asthmatics
is significant, e.g., in excess of $8,000 per patient per year by
some estimate. The present image-guided thermotherapy techniques
can be sued for downstaging severe asthma to mild persistent asthma
and significantly reduce the costs of healthcare related to asthma
patients.
[0051] For example, a device for thermotherapy based on the present
optical imaging technique can include a catheter that includes a
working channel configured for insertion into a passage of a body
to reach a target tissue inside the body; and an optical imaging
module that includes (1) an imaging optic fiber having a portion
inserted into the working channel and (2) an optical probe head
coupled to an end of the imaging optic fiber and located inside the
working channel. The optical imaging module is operable to direct
probe light to and collect reflected light from the target tissue
in the body through the imaging optic fiber and the optical probe
head and to obtain imaging information of the target tissue from
the collected reflected light. This device includes a thermotherapy
module having a power delivery waveguide having a portion inserted
into the working channel to deliver thermal energy to the target
tissue. A control unit is provided to control the optical imaging
module to extract the imaging information from the collected
reflected light, to obtain a spatial distribution of diseased
locations of the target tissue, and to obtain a temperature map of
the target tissue for thermotherapy based on the spatial
distribution of the diseased locations of the target tissue. The
control unit controls the thermotherapy module to control a
location and an amount of thermal energy delivery to each of the
diseased locations based on the temperature map to perform
thermotherapy.
[0052] Examples of various implementations of various components
and features are described in U.S. Publication No.
US-2007-0103683-A1 entitled "OPTICALLY MEASURING SUBSTANCES USING
PROPAGATION MODES OF LIGHT" for U.S. application Ser. No.
10/567,185, PCT Publication No. WO 2009/108950 entitled
"TEMPERATURE PROFILE MAPPING AND GUIDED THERMOTHERAPY" for PCT
Application No. PCT/US2009/035773, and U.S. Publication No. US
2006-0079762 A1 entitled "Integrated Disease Diagnosis and
Treatment System" for U.S. application Ser. No. 11/253,242. All of
these patent documents are incorporated by reference as part of the
disclosure of this document.
[0053] The optical imaging module can be implemented in various
configurations. Specific examples are provided below for
non-invasive optical imaging.
[0054] Investigation of substances by non-invasive and optical
means has been the object of many studies as inhomogeneity of
light-matter interactions in substances can reveal their
structural, compositional, physiological and biological
information. Various devices and techniques based on optical
coherence domain reflectometry (OCDR) may be used for non-invasive
optical probing of various substances, including but not limited to
skins, body tissues and organs of humans and animals, to provide
tomographic measurements of these substances.
[0055] In many OCDR systems, the light from a light source is split
into a sampling beam and a reference beam which propagate in two
separate optical paths, respectively. The light source may be
partially coherent source. The sampling beam is directed along its
own optical path to impinge on the substances under study, or
sample, while the reference beam is directed in a separate path
towards a reference surface. The beams reflected from the sample
and from the reference surface are then brought to overlap with
each other to optically interfere. Because of the
wavelength-dependent phase delay the interference results in no
observable interference fringes unless the two optical path lengths
of the sampling and reference beams are very similar. This provides
a physical mechanism for ranging. A beam splitter may be used to
split the light from the light source and to combine the reflected
sampling beam and the reflected reference beam for detection at an
optical detector. This use of the same device for both splitting
and recombining the radiation is essentially based on the
well-known Michelson interferometer. The discoveries and the
theories of the interference of partially coherent light are
summarized by Born and Wolf in "Principles of Optics", Pergamon
Press (1980).
[0056] FIG. 1 illustrates a typical optical layout used in many
fiber-optic OCDR systems described in U.S. Pat. No. 6,421,164 and
other publications. A fiber splitter is attached to two optical
fibers that respectively guide the sampling and reference beams in
a Michelson configuration. Common to many of these and other
implementations, the optical radiation from the low-coherence
source is first physically separated into two separate beams where
the sampling beam travels in a sample waveguide to interact with
the sample while the reference beam travels in a reference
waveguide. The fiber splitter than combines the reflected radiation
from the sample and the reference light from the reference
waveguide to cause interference.
[0057] Lung cancer is one of the most deadly cancers in the United
States. Patients with lung cancer have a relatively low 5-year
survival rate of only 10-15% after diagnosis. The lung cancer in
many patients is already in the second or third stage and has
metastasized to other sites or organs by the time they begin to
exhibit symptoms and seek medical treatment. Few are diagnosed in
early stages where the survival rate can be much higher,
approaching 85% for the stage 1 lung cancer. The conventional
annual chest X-ray examination has not shown sufficient sensitivity
to reveal the isolated, small (e.g., less than 1 centimeter in
diameter) tumors typically found in the stage 1 lung cancer.
[0058] Recently, emphasis has shifted to early stage detection in
major European and Japanese studies. In the US, a major new trial,
the National Lung Screening Trial (NLST), has begun and is aimed at
evaluating the efficacy of thoracic Computed Tomography (CT) scans
in detecting early stage lung cancer. The NLST will compare a
randomly selected group of high risk subjects (ex-smokers) who
receive annual CT scans to a control group of subjects receiving
chest x-rays.
[0059] The results of early studies have shown that thoracic CT
scans often revealed a substantial number of solitary pulmonary
nodules (SPNs). Biopsies have shown that approximately 80% or
greater (e.g., 98%) of these SPNs were calcified and benign.
However, the CT scan could not distinguish between calcified SPNs
and active SPNs. The inability of the CT scans to distinguish
malignancies from benign SPNs has led to a vigorous debate as to
the efficacy of the CT scans in early screening for lung
cancer.
[0060] A remedy to this defect of CT scans is to perform one or
more pulmonary biopsies in order to further examine the nature of
the SPNs identified by the CT scans. Pulmonary biopsies, however,
can be risky. Statistics show that one in four pulmonary biopsies
results in pneumothorax, a punctured lung. Also, the elderly and
patients on blood thinners are at substantial risk of bleeding
during pulmonary biopsies. In addition, pulmonary biopsies are
relatively expensive. These and other factors have lead to search
for alternative diagnostic methods to replace pulmonary
biopsies.
[0061] The non-invasive optical probing techniques and devices
described in this application may be used to detect and diagnose
lung diseases in humans and animals including lung cancer. The
optical probe head described in various implementations may be
inserted into the lung to optically measure various parts of the
lung without taking physical samples from the lung. The following
sections first describe the specific implementations of
non-invasive optical probing based on spectral responses of tissues
or parts and interactions of different optical modes in the probe
light. Next, examples of integrated lung disease diagnosis and
treatment systems that combine CT scan with optical probing and
laser treatment are described.
[0062] Spectral responses of materials and substances are important
in many applications. For example, some distinct material
properties are reflected in their spectral responses and can be
detected or measured via the spectral responses. A detected or
measured distinct property may be used for, e.g., identifying and
locating a region or area such as a body part of a person or
animal. Next, the identified body part may be further analyzed. As
a more specific example, cancer tumors or other conditions can be
detected and located using the measured spectral responses. Various
non-invasive optical techniques described in this application may
be used to measure spectral responses of a targeted body part of a
person or animal. An optical probe head is used to scan a probing
beam through the body part to optically measure the optical
responses of the targeted body part to obtain a map. At each
location within the targeted body part, light at different optical
wavelengths is used to obtain optical absorption responses at these
different wavelengths. Notably, the spectral absorption features of
a target layer underneath the surface may be optically selected and
measured by rejecting contributions to the reflected probe light
made by the tissues outside the boundaries of the target layer.
[0063] In some implementations, a single broadband light source may
be used for the acquisition of the spectral information within the
emission spectral range of the light source. A tunable optical
filter may be used to single out the spectral response of a narrow
wavelength band within the emitted spectrum of the light source.
When an absorbance feature to be measured or various targeted
absorbance features in a body part under measured occupy a broad
spectral range beyond the emission spectral bandwidth of a single
light source, the light source may be implemented by combining two
or more light sources for the acquisition of spectral absorbance
mapping (SAM) in tissues and other samples.
[0064] The following sections first describe various techniques and
devices for non-invasive optical probing using a single light
source and then describe devices and techniques that combine two or
more different light sources at different spectral ranges for the
SAM measurements.
[0065] Energy in light traveling in an optical path such as an
optical waveguide may be in different propagation modes. Different
propagation modes may be in various forms. States of optical
polarization of light are examples of such propagation modes. Two
independent propagation modes do not mix with one another in the
absence of a coupling mechanism. As an example, two orthogonally
polarization modes do not interact with each other even though the
two modes propagate along the same optical path or waveguide and
are spatially overlap with each other. The exemplary techniques and
devices described in this application use two independent
propagation modes in light in the same optical path or waveguide to
measure optical properties of a sample. A probe head may be used to
direct the light to the sample, either in two propagation modes or
in a single propagation modes, and receive the reflected or
back-scattered light from the sample.
[0066] For example, one beam of guided light in a first propagation
mode may be directed to a sample. A first portion of the first
propagation mode may be arranged to be reflected before reaching
the sample while a second portion in the first propagation mode is
allowed to reach the sample. The reflection of the second portion
from the sample is controlled in a second propagation mode
different from the first propagation mode to produce a reflected
second portion. Both the reflected first portion in the first
propagation mode and the reflected second portion in the second
propagation mode are directed through a common waveguide into a
detection module to extract information from the reflected second
portion on the sample.
[0067] In another example, optical radiation in both a first
propagation mode and a second, different propagation mode may be
guided through an optical waveguide towards a sample. The radiation
in the first propagation mode is directed away from the sample
without reaching the sample. The radiation in the second
propagation mode is directed to interact with the sample to produce
returned radiation from the interaction. Both the returned
radiation in the second propagation mode and the radiation in the
first propagation mode are coupled into the optical waveguide away
from the sample. The returned radiation in the second propagation
mode and the radiation in the first propagation mode from the
optical waveguide are then used to extract information of the
sample.
[0068] In these and other implementations based on the disclosure
of this application, two independent modes are confined to travel
in the same waveguide or the same optical path in free space except
for the extra distance traveled by the probing light between the
probe head and the sample. This feature stabilizes the relative
phase, or differential optical path, between the two modes of
light, even in the presence of mechanical movement of the
waveguides. This is in contrast to interferometer sensing devices
in which sample light and reference light travel in different
optical paths. These interferometer sensing devices with separate
optical paths are prone to noise caused by the variation in the
differential optical path, generally complex in optical
configurations, and difficult to operate and implement. The
examples described below based on waveguides are in part designed
to overcome these and other limitations.
[0069] FIG. 2 shows one example of a sensing device according to
one implementation. This device directs light in two propagation
modes along the same waveguide to an optical probe head near a
sample 205 for acquiring information of optical inhomogeneity in
the sample. A sample holder may be used to support the sample 205
in some applications. Light radiation from a broadband light source
201 is coupled into the first dual-mode waveguide 271 to excite two
orthogonal propagation modes, 001 and 002. A light director 210 is
used to direct the two modes to the second dual-mode waveguide 272
that is terminated by a probe head 220. The probe head 220 may be
configured to perform at least the following functions. The first
function of the probe head 220 is to reverse the propagation
direction of a portion of light in the waveguide 272 in the mode
001; the second function of the probe head 220 is to reshape and
deliver the remaining portion of the light in mode 002 to the
sample 205; and the third function of the probe head 220 is to
collect the light reflected from the sample 205 back to the second
dual-mode waveguide 272. The back traveling light in both modes 001
and 002 is then directed by light director 210 to the third
waveguide 273 and further propagates towards a differential delay
modulator 250. The differential delay modulator 250 is capable of
varying the relative optical path length and optical phase between
the two modes 001 and 002. A detection subsystem 260 is used to
superpose the two propagation modes 001 and 002 to form two new
modes, mutually orthogonal, to be received by photo-detectors. Each
new mode is a mixture of the modes 001 and 002.
[0070] The superposition of the two modes 001 and 002 in the
detection subsystem 260 allows for a range detection. The light
entering the detection subsystem 260 in the mode 002 is reflected
by the sample, bearing information about the optical inhomogeneity
of the sample 205, while the other mode, 001, bypassing the sample
205 inside probe head 220. So long as these two modes 001 and 002
remain independent through the waveguides their superposition in
the detection subsystem 260 may be used to obtain information about
the sample 205 without the separate optical paths used in some
conventional Michelson interferometer systems.
[0071] For the simplicity of the analysis, consider a thin slice of
the source spectrum by assuming that the amplitude of the mode 001
is E.sub.001 in a first linear polarization and that of the mode
002 is E.sub.002 in a second, orthogonal linear polarization in the
first waveguide 271. The sample 205 can be characterized by an
effective reflection coefficient r that is complex in nature; the
differential delay modulator 250 can be characterized by a pure
phase shift .GAMMA. exerted on the mode 001. Let us now superpose
the two modes 001 and 002 by projecting them onto a pair of new
modes, E.sub.A and E.sub.B, by a relative 45-degree rotation in the
vector space. The new modes, E.sub.A and E.sub.B, may be expressed
as following:
{ E A = 1 2 ( j .GAMMA. E 001 + rE 002 ) ; E B = 1 2 ( j .GAMMA. E
001 - rE 002 ) . ( 1 ) ##EQU00001##
It is assumed that all components in the system, except for the
sample 205, are lossless. The resultant intensities of the two
superposed modes are
{ I A = 1 2 [ E 001 2 + E 002 2 + r E 001 E 002 cos ( .GAMMA. -
.PHI. ) ] ; I B = 1 2 [ E 001 2 + E 002 2 - r E 001 E 002 cos (
.GAMMA. - .PHI. ) ] , ( 2 ) ##EQU00002##
where .phi. is the phase delay associated with the reflection from
the sample. A convenient way to characterize the reflection
coefficient r is to measure the difference of the above two
intensities, i.e.
I.sub.A-I.sub.B=|r|E.sub.001E.sub.002 cos(.GAMMA.-.phi.). (3)
[0072] If .GAMMA. is modulated by the differential delay modulator
250, the measured signal, Eq. (3), is modulated accordingly. For
either a periodic or a time-linear variation of .GAMMA., the
measured signal responds with a periodic oscillation and its
peak-to-peak value is proportional to the absolute value of r.
[0073] For a broadband light source 201 in FIG. 2, consider the two
phases, .GAMMA. and .phi. to be dependent on wavelength. If the two
modes 001 and 002 experience significantly different path lengths
when they reach the detection system 260, the overall phase angle,
.GAMMA.-.phi., should be significantly wavelength dependant as
well. Consequently the measured signal, being an integration of Eq.
(3) over the source spectrum, yields a smooth function even though
.GAMMA. is being varied. The condition for a significant
oscillation to occur in the measured signal is when the two modes
001 and 002 experience similar path lengths at the location of
their superposition. In this case the overall phase angle,
.GAMMA.-.phi., becomes wavelength independent or nearly wavelength
independent. In other words, for a given relative path length set
by the modulator 250, an oscillation in the measured signal
indicates a reflection, in the other mode, from a distance that
equalizes the optical path lengths traveled by the two modes 001
and 002. Therefore the system depicted in FIG. 2 can be utilized
for ranging reflection sources.
[0074] Due to the stability of the relative phase between the two
modes, 001 and 002, phase-sensitive measurements can be performed
with the system in FIG. 2 with relative ease. The following
describes an exemplary method based on the system in FIG. 2 for the
determination of the absolute phase associated with the radiation
reflected from the sample 205.
[0075] In this method, a sinusoidal modulation is applied to the
differential phase by the differential delay modulator 250, with a
modulation magnitude of M and a modulation frequency of .OMEGA..
The difference in intensity of the two new modes is the measured
and can be expressed as follows:
I.sub.A-I.sub.B=|r|E.sub.001E.sub.002 cos [M sin(.OMEGA.t)-.phi.].
(4)
It is clear from Eq. (4) that the measured exhibits an oscillation
at a base frequency of .OMEGA. and oscillations at harmonic
frequencies of the base frequency .OMEGA.. The amplitudes of the
base frequency and each of the harmonics are related to .phi. and
|r|. The relationships between r and the harmonics can be derived.
For instance, the amplitude of the base-frequency oscillation and
the second harmonic can be found from Eq. (4) to be:
A.sub..OMEGA.=E.sub.001E.sub.002J.sub.1(M)|r| sin .phi.; (5a)
A.sub.2.OMEGA.=E.sub.001E.sub.002J.sub.2(M)|r| cos .phi., (5b)
where J.sub.1 and J.sub.2 are Bessel functions of the first and
second order, respectively. Eq. (5a) and (5b) can be used to solve
for |r| and .phi., i.e. the complete characterization of r. We can
therefore completely characterize the complex reflection
coefficient r by analyzing the harmonic content of various orders
in the measured signal. In particular, the presence of the
base-frequency component in the measured is due to the presence of
.phi..
[0076] FIG. 3 shows an exemplary implementation of the system
depicted in FIG. 2. The spectrum of source 201 may be chosen to
satisfy the desired ranging resolution. The broader the spectrum is
the better the ranging resolution. Various light sources may be
used as the source 201. For example, some semiconductor
superluminescent light emitting diodes (SLED) and amplified
spontaneous emission (ASE) sources may possess the appropriate
spectral properties for the purpose. In this particular example, a
polarization controller 302 may be used to control the state of
polarization in order to proportion the magnitudes of the two
modes, 001 and 002, in the input waveguide 371. The waveguide 371
and other waveguides 372 and 373 may be dual-mode waveguides and
are capable of supporting two independent polarization modes which
are mutually orthogonal. One kind of practical and commercially
available waveguide is the polarization maintaining (PM) optical
fiber. A polarization maintaining fiber can carry two independent
polarization modes, namely, the s-wave polarized along its slow
axis and the p-wave polarized along its fast axis. In good quality
polarization maintaining fibers these two modes can have virtually
no energy exchange, or coupling, for substantial distances.
Polarization preserving circulator 310 directs the flow of optical
waves according to the following scheme: the two incoming
polarization modes from fiber 371 are directed into the fiber 372;
the two incoming polarization modes from fiber 372 are directed to
the fiber 373. A polarization-preserving circulator 310 may be used
to maintain the separation of the two independent polarization
modes. For instance, the s-wave in the fiber 371 should be directed
to the fiber 372 as s-wave or p-wave only. Certain commercially
available polarization-preserving circulators are adequate for the
purpose.
[0077] The system in FIG. 3 implements an optical probe head 320
coupled to the waveguide 372 for optically probing the sample 205.
The probe head 320 delivers a portion of light received from the
waveguide 372, the light in one mode (e.g., 002) of the two modes
001 and 002, to the sample 205 and collects reflected and
back-scattered light in the same mode 002 from the sample 205. The
returned light in the mode 002 collected from the sample 205
carries information of the sample 205 and is processed to extract
the information of the sample 205. The light in the other mode 001
in the waveguide 372 propagating towards the probe head 320 is
reflected back by the probe head 320. Both the returned light in
the mode 002 and the reflected light in the mode 001 are directed
back by the probe head 320 into the waveguide 372 and to the
differential delay modulator 250 and the detection system 260
through the circulator 310 and the waveguide 373.
[0078] In the illustrated implementation, the probe head 320
includes a lens system 321 and a polarization-selective reflector
(PSR) 322. The lens system 321 is to concentrate the light energy
into a small area, facilitating spatially resolved studies of the
sample in a lateral direction. The polarization-selective reflector
322 reflects the mode 001 back and transmits the mode 002. Hence,
the light in the mode 002 transmits through the probe head 320 to
impinge on the sample 205. Back reflected or scattered the light
from the sample 205 is collected by the lens system 321 to
propagate towards the circulator 310 along with the light in the
mode 001 reflected by PSR 322 in the waveguide 372.
[0079] FIG. 4 shows details of the probe head 320 and an example of
the polarization-selective reflector (PSR) 322 according to one
implementation. The PSR 322 includes a polarizing beam splitter
(PBS) 423 and a reflector or mirror 424 in a configuration as
illustrated where the PBS 423 transmits the selected mode (e.g.,
mode 002) to the sample 205 and reflects and diverts the other mode
(e.g., mode 001) away from the sample 205 and to the reflector 424.
By retro reflection of the reflector 424, the reflected mode 001 is
directed back to the PBS 423 and the lens system 321. The reflector
424 may be a reflective coating on one side of beam splitter 423.
The reflector 424 should be aligned to allow the reflected
radiation to re-enter the polarization-maintaining fiber 372. The
transmitted light in the mode 002 impinges the sample 205 and the
light reflected and back scattered by the sample 205 in the mode
002 transmits through the PBS 423 to the lens system 321. The lens
system 321 couples the light in both the modes 001 and 002 into the
fiber 372.
[0080] In the implementation illustrated in FIG. 3, the detection
system 260 includes a polarizing beam splitter 361, and two
photodetectors 362 and 363. The polarizing beam splitter 361 is
used to receive the two independent polarization modes 001 and 002
from the modulator 250 and superposes the two independent
polarization modes 001 and 002. The beam splitter 361 may be
oriented in such a way that, each independent polarization is split
into two parts and, for each independent polarization mode, the two
split portions possess the same amplitude. This way, a portion of
the mode 001 and a portion of the mode 002 are combined and mixed
in each of the two output ports of the beam splitter 361 to form a
superposed new mode and each photodetector receives a superposed
mode characterized by Eq. (1). The polarizing beam splitter 361 may
be oriented so that the incident plane of its reflection surface
makes a 45-degree angle with one of the two independent
polarization mode, 001 or 002.
[0081] The system in FIG. 3 further implements an electronic
controller or control electronics 370 to receive and process the
detector outputs from the photodetectors 362 and 363 and to control
operations of the systems. The electronic controller 370, for
example, may be used to control the probe head 320 and the
differential delay modulator 250. Differential delay modulator 250,
under the control of the electronics and programs, generates a form
of differential phase modulation as the differential path length
scans through a range that matches a range of depth inside the
sample 205. The electronic controller 370 may also be programmed to
record and extract the amplitude of the oscillation in the measured
signal characterized by Eq. (3) at various differential path
lengths generated by the modulator 250. Accordingly, a profile of
reflection as a function of the depth can be obtained as a
one-dimensional representation of the sample inhomogeneity at a
selected location on the sample 205.
[0082] For acquiring two-dimensional images of optical
inhomogeneity in the sample 205, the probe head 320 may be
controlled via a position scanner such as a translation stage or a
piezo-electric positioner so that the probing light scans in a
lateral direction, perpendicular to the light propagation
direction. For every increment of the lateral scan a profile of
reflection as a function of depth can be recorded with the method
described above. The collected information can then be displayed on
a display and interface module 372 to form a cross-sectional image
that reveals the inhomogeneity of the sample 205.
[0083] In general, a lateral scanning mechanism may be implemented
in each device described in this application to change the relative
lateral position of the optical probe head and the sample to obtain
a 2-dimensional map of the sample. A xy-scanner, for example, may
be engaged either to the optical head or to a sample holder that
holds the sample to effectuate this scanning in response to a
position control signal generated from the electronic controller
370.
[0084] FIGS. 5A and 5B illustrate another exemplary system that use
waveguides 271, 272, and 273 and a light director 210 to direct
light in two modes to and from the probe head 320 in measuring the
sample 205. A first optical polarizer 510 is oriented with respect
to the polarization axes of the PM waveguide 271 to couple
radiation from the broadband light source 201 into the waveguide
271 in two orthogonal linear polarization modes as the independent
propagation modes. An optical phase modulator 520 is coupled in the
waveguide 271 to modulate the optical phase of light in one guided
mode relative to the other. A variable differential group delay
(VDGD) device 530 is inserted in or connected to the waveguide 273
to introduce a controllable amount of optical path difference
between the two waves. A second optical polarizer 540 and an
optical detector 550 are used here to form a detection system. The
second polarizer 540 is oriented to project both of the guided
waves onto the same polarization direction so that the changes in
optical path difference and the optical phase difference between
the two propagation modes cause intensity variations, detectable by
the detector 550.
[0085] The light from the source 201 is typically partially
polarized. The polarizer 510 may be aligned so that maximum amount
of light from the source 201 is transmitted and that the
transmitted light is coupled to both of the guided modes in the
waveguide 271 with the substantially equal amplitudes. The electric
fields for the two orthogonal polarization modes S and P in the
waveguide 271 can be expressed as:
{ E s = 1 2 E , E p = 1 2 E . ( 6 ) ##EQU00003##
where the electric field transmitting the polarizer is denoted as
E. It should be appreciated that the light has a finite spectral
width (broadband or partially coherent). The fields can be
described by the following Fourier integral:
E=.intg.E.sub..omega.e.sup.j.omega.td.omega.. (7)
For the simplicity of the analysis, a thin slice of the spectrum,
i.e. a lightwave of a specific wavelength, is considered below.
Without loosing generality, it is assumed that all the components,
including polarizers, waveguides, Router, PSR and VDGD, are
lossless. Let us designate the reflection coefficient of the sample
r, that is complex in nature. The p-wave picks up an optical phase,
.GAMMA., relative to the s-wave as they reach the second polarizer
540:
{ E s = 1 2 E , E p = 1 2 rE j .GAMMA. . ( 8 ) ##EQU00004##
The light that passes through Polarizer 540 can be expressed by
E a = 1 2 ( E s + E p ) = 1 2 E ( 1 + r j .GAMMA. ) . ( 9 )
##EQU00005##
The intensity of the light that impinges on the photodetector 550
is given by:
I = E a E a * = 1 4 E 2 [ 1 + r 2 + 2 r cos ( .GAMMA. = .delta. ) ]
. ( 10 ) ##EQU00006##
where phase angle .delta. reflects the complex nature of the
reflection coefficient of the sample 205 and is defined by
r=|r|e.sup.j.delta.. (11)
Assuming the modulator 520 exerts a sinusoidal phase modulation,
with magnitude M and frequency .OMEGA., in the p-wave with respect
to the s-wave, the light intensity received by the detector 550 can
be expressed as follows:
I = 1 + r 2 4 E 2 + r 2 E 2 cos [ M sin ( .OMEGA. t ) + .PHI. +
.delta. ] . ( 12 ) ##EQU00007##
where phase angle .phi. is the accumulated phase slip between the
two modes, not including the periodic modulation due to the
modulator 520. The VDGD 530 or a static phase shift in the
modulator 520, may be used to adjust the phase difference between
the two modes to eliminate .phi..
[0086] FIG. 6 illustrates the waveform of the intensity I received
at the detector 550 as a function of the phase. The detected light
intensity exhibits an oscillating waveform that possesses a base
frequency of .OMEGA. and its harmonics. The amplitudes of the base
frequency and each of the harmonics are related to .delta. and |r|.
The mathematical expressions for the relationships between r and
the harmonics can be derived. For instance, the amplitude of the
base-frequency oscillation and the second harmonic are found to
be:
A.sub..OMEGA.=0.5|E|.sup.2J.sub.1(M)|r| sin .delta.; (13a)
A.sub.2.OMEGA.=0.5|E|.sup.2J.sub.2(M)|r| cos .delta., (13b)
where J.sub.1 and J.sub.2 are Bessel functions of the first and
second order, respectively. Eq. (13a) and (13b) can be used to
solve for |r| and .delta., i.e. the complete characterization of
r.
[0087] The effect of having a broadband light source 201 in the
system in FIGS. 5A and 5B is analyzed below. When there is a
significant differential group delay between the two propagation
modes there must be an associated large phase slippage .phi. that
is wavelength dependent. A substantial wavelength spread in the
light source means that the phase slippage also possesses a
substantial spread. Such a phase spread cannot be eliminated by a
phase control device that does not also eliminate the differential
group delay. In this case the detected light intensity is given by
the following integral:
I = .intg. { 1 + r 2 4 E ( .lamda. ) 2 + r 2 E ( .lamda. ) 2 cos [
M sin ( .OMEGA. t ) + .PHI. ( .lamda. ) = .delta. ] } .lamda. . (
14 ) ##EQU00008##
It is easy to see that if the range of .phi.(.lamda.) is comparable
to .pi. for the bandwidth of the light source no oscillation in I
can be observed as oscillations for different wavelengths cancel
out because of their phase difference. This phenomenon is in close
analogy to the interference of white light wherein color fringes
are visible only when the path difference is small (the film is
thin). The above analysis demonstrates that the use of a broadband
light source enables range detection using the proposed apparatus.
In order to do so, let the s-wave to have a longer optical path in
the system compared to the p-wave (not including its round-trip
between Probing Head and Sample). For any given path length
difference in the system there is a matching distance between
Probing Head and Sample, z, that cancels out the path length
difference. If an oscillation in I is observed the p-wave must be
reflected from this specific distance z. By varying the path length
difference in the system and record the oscillation waveforms we
can therefore acquire the reflection coefficient r as a function of
the longitudinal distance z, or depth. By moving Probing Head
laterally, we can also record the variation of r in the lateral
directions.
[0088] FIG. 7 further shows one exemplary operation of the
described system in FIG. 5B or the system in FIG. 3 for acquiring
images of optical inhomogeneity. At step 710, the relative phase
delay between the two modes is changed, e.g., increased by an
increment, to a fixed value for measuring the sample 205 at a
corresponding depth. This may be accomplished in FIG. 5B by using
the differential delay device 530 or the bias in the differential
delay modulator 250 in FIG. 3. At step 720, a modulation driving
signal is sent to the modulator 520 in FIG. 5B or the modulator 250
in FIG. 3 to modulate the relative phase delay between the two
modes around the fixed value. At step 730, the intensity waveform
received in the detector 550 in FIG. 5B or the intensity waveforms
received in the detectors 362,363 in FIG. 3 are measured and stored
in the electronic controller 370. Upon completion of the step 730,
the electronic controller 370 controls the differential delay
device 530 in FIG. 5B or the bias in the differential delay
modulator 250 in FIG. 3 to change the relative phase delay between
the two modes to a different fixed value for measuring the sample
205 at a different depth. This process iterates as indicated by the
processing loop 740 until desired measurements of the sample at
different depths at the same location are completed. At this point,
electronic controller 370 controls the probe head 320 to laterally
move to a new location on the sample 205 and repeat the above
measurements again until all desired locations on the sample 205
are completed. This operation is represented by the processing loop
750. The electronic controller 370 processes each measurement to
compute the values of .delta. and |r| from the base oscillation and
the harmonics at step 760. Such data processing may be performed
after each measurement or after all measurements are completed. At
step 770, the computed data is sent to the display module 372.
[0089] In the above implementations, light for sensing the sample
205 is not separated into two parts that travel along two different
optical paths. Two independent propagation modes of the light are
guided essentially in the same waveguide at every location along
the optical path except for the extra distance traveled by one mode
between the probe head 320 and the sample 205. After redirected by
the probe head 320, the two modes are continuously guided in the
same waveguide at every location along the optical path to the
detection module.
[0090] Alternatively, the light from the light source to the probe
head may be controlled in a single propagation mode (e.g., a first
propagation mode) rather than two different modes. The probe head
may be designed to cause a first portion of the first mode to
reverse its propagation direction while directing the remaining
portion, or a second portion, to reach the sample. The reflection
or back scattered light of the second portion from the sample is
collected by the probe head and is controlled in the second
propagation mode different from the first mode to produce a
reflected second portion. Both the reflected first portion in the
first propagation mode and the reflected second portion in the
second propagation mode are directed by the probe head through a
common waveguide into the detection module for processing. In
comparison with the implementations that use light in two modes
throughout the system, this alternative design further improves the
stability of the relative phase delay between the two modes at the
detection module and provides additional implementation
benefits.
[0091] FIGS. 8A and 8B illustrate one exemplary design of the
optical layout of the optical sensing system and its system
implementation with an electronic controller. An input waveguide
871 is provided to direct light in a first propagation mode, e.g.,
the mode 001, from the broadband light source 201 to a light
director 810. The waveguide 871 may be a mode maintaining waveguide
designed to support at least one propagation mode such as the mode
001 or 002. When light is coupled into the waveguide 871 in a
particular mode such as the mode 001, the waveguide 871 essentially
maintains the light in the mode 001. A polarization maintaining
fiber supporting two orthogonal linear polarization modes, for
example, may be used as the waveguide 871. Similar to systems shown
in FIGS. 2, 3, 5A and 5B, dual-mode waveguides 272 and 273 are used
to direct the light. A light director 510 is used to couple the
waveguides 871, 272, and 273, to convey the mode 001 from the input
waveguide 871 to one of the two modes (e.g., modes 001 and 002)
supported by the dual-mode waveguide 272, and to direct light in
two modes from the waveguide 272 to the dual-mode waveguide 273. In
the example illustrated in FIG. 8A, the light director 810 couples
the light in the mode 001 from the waveguide 871 into the same mode
001 in the waveguide 272. Alternatively, the light director 810 may
couple the light in the mode 001 from the waveguide 871 into the
different mode 002 in the waveguide 272. The dual-mode waveguide
271 is terminated at the other end by a probe head 820 which
couples a portion of light to the sample 205 for sensing.
[0092] The probe head 820 is designed differently from the prove
head 320 in that the probe head 830 converts part of light in the
mode 001 into the other different mode 002 when the light is
reflected or scattered back from the sample 205. Alternatively, if
the light in the waveguide 272 that is coupled from the waveguide
871 is in the mode 002, the probe head 820 converts that part of
light in the mode 002 into the other different mode 001 when the
light is reflected or scattered back from the sample 205. In the
illustrated example, the probe head 820 performs these functions:
a) to reverse the propagation direction of a small portion of the
incoming radiation in mode 001; b) to reshape the remaining
radiation and transmit it to the sample 205; and c) to convert the
radiation reflected from the sample 205 to an independent mode 002
supported by the dual-mode waveguide 272. Since the probe head 820
only converts part of the light into the other mode supported by
the waveguide 272, the probe head 820 is a partial mode converter
in this regard. Due to the operations of the probe head 820, there
are two modes propagating away from the probe head 820, the mode
001 that bypasses the sample 205 and the mode 002 for light that
originates from sample reflection or back scattering. From this
point on, the structure and operations of the rest of the system
shown in FIG. 8A may be similar to the systems in FIGS. 2, 3, 5A,
and 5B.
[0093] FIG. 8B shows an exemplary implementation of the design in
FIG. 8A where an electronic controller 2970 is used to control the
differential delay modulator 250 and the probe head 820 and a
display and interface module 372 is provided. Radiation from
broadband light source 201, which may be partially polarized, is
further polarized and controlled by an input polarization
controller 802 so that only a single polarization mode is excited
in polarization-maintaining fiber 371 as the waveguide 871 in FIG.
8A. a polarization preserving circulator may be used to implement
the light director 810 for routing light from the waveguide 371 to
the waveguide 372 and from the waveguide 372 to the waveguide
373.
[0094] The probe head 820 in FIG. 8B may be designed to include a
lens system 821 similar to the lens system 321, a partial reflector
822, and a polarization rotator 823. The partial reflector 822 is
used to reflect the first portion of light received from the
waveguide 372 back to the waveguide 372 without changing its
propagation mode and transmits light to and from the sample 205.
The polarization rotator 823 is used to control the light from the
sample 205 to be in the mode 002 upon entry of the waveguide
372.
[0095] FIG. 9 shows another example of a system implementation
where the optical probe head 820 receives light in a single input
mode and converts part of light into a different mode. An input
polarizer 510 is used in the input PM fiber 272 to control the
input light in the single polarization mode. A phase modulator 520
and a variable differential group delay device 530 are coupled to
the output PM fiver 273 to control and modulate the relative phase
delay of the two modes before optical detection. An output
polarizer 540 is provided to mix the two modes and the detector 550
is used to detect the output from the output polarizer 540.
[0096] FIGS. 10A and 10B show two examples of the possible designs
for the probe head 820 including a partially reflective surface
1010, a lens system 1020, and a quarter-wave plate 1030 for
rotating the polarization and to convert the mode. In FIG. 10A, the
termination or end facet of polarization-maintaining fiber 372 is
used as the partial reflector 1010. An uncoated termination of an
optical fiber reflects approximately 4% of the light energy.
Coatings can be used to alter the reflectivity of the termination
to a desirable value. The lens system 1020 reshapes and delivers
the remaining radiation to sample 205. The other role played by the
lens system 1020 is to collect the radiation reflected from the
sample 205 back into the polarization-maintaining fiber 372. The
quarter wave plate 1030 is oriented so that its optical axis make a
45-degree angle with the polarization direction of the transmitted
light. Reflected light from the sample 205 propagates through the
quarter wave plate 1030 once again to become polarized in a
direction perpendicular to mode 001, i.e. mode 002. Alternatively,
the quarter wave plate 1030 may be replaced by a Faraday rotator.
The head design in FIG. 10B changes the positions of the lens
system 1020 and the quarter wave plate or Faraday rotator 1030.
[0097] In the examples in FIGS. 8A, 8B, and 9, there is only one
polarization mode entering the light director 810 or the
polarization-preserving circulator from waveguide 871 or 371.
Therefore, the light director 810 or the polarization preserving
circulator may be constructed with a polarization-maintaining
optical circulator 1110 and two polarization beam splitters 1120
and 1130 as shown in FIG. 11. The polarization-maintaining
circulator 1110 is used to convey only one polarization mode among
its three ports, rather than both modes as in the case shown in
FIGS. 3, 5A and 5B. The polarizing beam splitter 1120 and 1130 are
coupled to polarization-maintaining circulator 1110 so that both
polarization modes entering Port 2 are conveyed to Port 3 and
remain independent.
[0098] A number of hardware choices are available for differential
delay modulator 250. FIG. 12 illustrates the general design of the
modulator 250 where an external control signal is applied to
control a differential delay element to change and modulate the
relative delay in the output. Either mechanical or non-mechanical
elements may be used to produce the desired relative delay between
the two modes and the modulation on the delay.
[0099] In one implementation, a non-mechanical design may include
one or more segments of tunable birefringent materials such as
liquid crystal materials or electro-optic birefringent materials
such as lithium niobate crystals in conjunction with one or more
fixed birefringent materials such as quartz and rutile. The fixed
birefringent material provides a fixed delay between two modes and
the tunable birefringent material provides the tuning and
modulation functions in the relative delay between the two modes.
FIG. 12A illustrates an example of this non-mechanical design where
the two modes are not physically separated and are directed through
the same optical path with birefringent segments which alter the
relative delay between two polarization modes.
[0100] FIG. 12B shows a different design where the two modes in the
received light are separated by a mode splitter into two different
optical paths. A variable delay element is inserted in one optical
path to adjust and modulate the relative delay in response to an
external control signal. A mode combiner is then used to combine
the two modes together in the output. The mode splitter and the
mode combiner may be polarization beams splitters when two
orthogonal linear polarizations are used as the two modes.
[0101] The variable delay element in one of the two optical paths
may be implemented in various configurations. For example, the
variable delay element may be a mechanical element. A mechanical
implementation of the device in FIG. 12B may be constructed by
first separating the radiation by polarization modes with a
polarizing beam splitter, one polarization mode propagating through
a fixed optical path while the other propagating through a variable
optical path having a piezoelectric stretcher of polarization
maintaining fibers, or a pair of collimators both facing a
mechanically movable retroreflector in such a way that the light
from one collimator is collected by the other through a trip to and
from the retroreflector, or a pair collimators optically linked
through double passing a rotatable optical plate and bouncing off a
reflector.
[0102] FIGS. 13A and 13B illustrate two examples of a mechanical
variable delay element suitable for FIG. 12B. Such a mechanical
variable delay device may be used to change the optical path length
of a light beam at high speeds and may have various applications
other than what is illustrated in FIG. 12B. In addition, the
optical systems in this application may use such a delay
device.
[0103] The mechanical delay device shown in FIG. 13A includes an
optical beam splitter 1310, a rotating optical plate 1320 which may
be a transparent plate, and a mirror or reflector 1330. The beam
splitter 1310 is used as the input port and the output port for the
device. The rotating optical plate 1320 is placed between the
mirror 1330 and the beam splitter 1310. The input light beam 1300
is received by the beam splitter 1310 along the optical path
directing from the beam splitter 1310 to the mirror 1330 through
the rotating optical plate 1320. A portion of the light 1300
transmitting through the beam splitter 1310 is the beam 1301 which
impinges on and transmits through the rotating optical plate 1320.
The mirror or other optical reflector 1330 is oriented to be
perpendicular to the light beam incident to the optical plate 1310
from the opposite side. The reflected light beam 1302 from the
mirror 1320 traces the same optical path back traveling until it
encounters the Beam Splitter 1310. The Beam Splitter 1310 deflects
part of the back traveling light 1302 to a different direction as
the output beam 1303.
[0104] In this device, the variation of the optical path length is
caused by the rotation of the Optical Plate 1320. The Optical Plate
1320 may be made of a good quality optical material. The two
optical surfaces may be flat and well polished to minimize
distortion to the light beam. In addition, the two surfaces should
be parallel to each other so that the light propagation directions
on both sides of the Optical Plate 1320 are parallel. The thickness
of the Optical Plate 1320 may be chosen according to the desirable
delay variation and the range of the rotation angle. The optical
path length experienced by the light beam is determined by the
rotation angle of the Optical Plate 1320. When the surfaces of the
Optical Plate 1320 is perpendicular to the light beam (incident
angle is zero), the path length is at its minimum. The path length
increases as the incident angle increases.
[0105] In FIG. 13A, it may be beneficial to collimate the input
light beam so that it can travel the entire optical path without
significant divergence. The Optical Plate 1320 may be mounted on a
motor for periodic variation of the optical delay. A good quality
mirror with a flat reflecting surface should be used to implement
the mirror 1330. The reflecting surface of the mirror 1330 may be
maintained to be perpendicular to the light beam.
[0106] If a linearly polarized light is used as the input beam 1300
in FIG. 13A, it is beneficial to have the polarization direction of
the light parallel to the incident plane (in the plane of the
paper) as less reflection occurs at the surfaces of Optical Plate
1320 for this polarization compared to other polarization
directions. Antireflection coatings can be used to further reduce
the light reflection on the surfaces of the Optical Plate 1320.
[0107] The beam splitter 1310 used in FIG. 13A uses both its
optical transmission and optical reflection to direct light. This
aspect of the beam splitter 1310 causes reflection loss in the
output of the device due to the reflection loss when the input
light 1300 first enters the device through transmission of the beam
splitter 1310 and the transmission loss when the light exits the
device through reflection of the beam splitter 1310. For example, a
maximum of 25% of the total input light may be left in the output
light if the beam splitter is a 50/50 beam splitter. To avoid such
optical loss, an optical circulator may be used in place of the
beam splitter 1320. FIG. 13B illustrates an example where the
optical circulator 1340 with 3 ports is used to direct input light
to the optical plate 1320 and the mirror 1330 and directs returned
light to the output port. The optical circulator 1340 may be
designed to direct nearly all light entering its port 1 to port 2
and nearly all light entering its port 2 to the port 3 with nominal
optical loss and hence significantly reduces the optical loss in
the device. Commercially available optical circulators, either
free-space or fiber-based, may be used to implement the circulator
1340.
[0108] FIG. 14A shows an exemplary implementation of the delay
device in FIG. 12B as part of or the entire differential delay
modulator 250. A first optical mode splitter 1410 is used to
separate two modes in the waveguide 373 into two paths having two
mirrors 1431 and 1432, respectively. A second optical mode splitter
1440, which is operated as a mode combiner, is used to combine the
two modes into an output. If the two modes are two orthogonal
linear polarizations, for example, polarization beam splitters may
be used to implement the 1410 and 1440. A variable optical delay
line or device 1420 is placed in the upper path to control the
differential delay between the two paths. The output may be coupled
into another dual-mode waveguide 1450 leading to the detection
module or directly sent into the detection module. FIG. 14B shows a
delay device based on the design in FIG. 14A where the mirror 1432
and the variable optical delay line 1420 are implemented by the
mechanical delay device in FIG. 13A. The mechanical delay device in
FIG. 13B may also be used to implement the device in FIG. 14A.
[0109] In the above examples, a single dual-mode waveguide 272 or
372 is used as an input and output waveguide for the probe head
220, 320, or 820. Hence, the input light, either in a single mode
or two independent modes, is directed into the probe head through
that dual-mode waveguide 272 or 372, and the output light in the
two independent modes is also directed from the probe head to the
detection subsystem or detector.
[0110] Alternatively, the single dual-mode waveguide 272 or 372 may
be replaced by two separate waveguides, one to direct input light
from the light source to the probe head and another to direct light
from the probe head to the detection subsystem or detector. As an
example, the device in FIG. 2 may have a second waveguide different
from the waveguide 272 to direct reflected light in two different
modes from the optical probe head 220 to the modulator 250 and the
detection subsystem 260. In this design, the light director 210 may
be eliminated. This may be an advantage. In implementation, the
optics within the probe head may be designed to direct the
reflected light in two modes to the second waveguide.
[0111] FIG. 15 illustrates an example for this design as an
alternative to the device shown in FIG. 5B. In this design, the
probing light is delivered to the sample 205 through one dual-mode
waveguide 1510 and the reflected/scattered light is collected by
the probe head 320 and is directed through another dual-mode
waveguide 1520. With the probe head shown in FIG. 4, the mirror 424
may be oriented and aligned so that the light is reflected into the
waveguide 1520 instead of the waveguide 1510. This design may be
applied to other devices based on the disclosure of this
application, including the exemplary devices in FIGS. 2, 3, 8A, 8B
and 9.
[0112] The above-described devices and techniques may be used to
obtain optical measurements of a given location of the sample at
different depths by controlling the relative phase delay between
two modes at different values and optical measurements of different
locations of the sample to get a tomographic map of the sample at a
given depth or various depths by laterally changing the relative
position of the probe head over the sample. Such devices and
techniques may be further used to perform other measurements on a
sample, including spectral selective measurements on a layer of a
sample.
[0113] In various applications, it may be beneficial to obtain
information about certain substances, identifiable through their
spectral absorbance, dispersed in the samples. For this purpose, a
tunable bandpass filter may be used to either filter the light
incident to the probe head to select a desired spectral window
within the broadband spectrum of the incident light to measure the
response of the sample and to vary the center wavelength of the
spectral window to measure a spectral distribution of the responses
of the sample. This tuning of the bandpass filter allows a variable
portion of the source spectrum to pass while measuring the
distribution of the complex reflection coefficient of the
sample.
[0114] Alternatively, the broadband light may be sent to the
optical probe head without optical filtering and the spectral
components at different wavelengths in the output light from the
probe head may be selected and measured to measure the response of
the sample around a selected wavelength or the spectral
distribution of the responses of the sample. In one implementation,
a tunable optical bandpass filter may be inserted in the optical
path of the output light from the probe head to filter the light.
In another implementation, a grating or other diffractive optical
element may be used to optically separate different spectral
components in the output light to be measured by the detection
subsystem or the detector.
[0115] As an example, FIG. 16 shows a system based on the design in
FIG. 2 where a tunable filter 1610 is inserted in the input
waveguide 271 to filter the input light in two different modes.
FIG. 17 shows another exemplary system based on the design in FIG.
8A where a tunable filter 1710 is inserted in the input waveguide
871 to filter the input light in a single mode. Such a tunable
filter may be placed in other locations.
[0116] FIG. 18 illustrates the operation of the tunable bandpass
filter in the devices in FIGS. 16 and 17. The filter selects a
narrow spectral band within the spectrum of the light source to
measure the spectral feature of the sample.
[0117] Notably, the devices and techniques of this application may
be used to select a layer within a sample to measure by properly
processing the measured data. Referring back to the devices in
FIGS. 16 and 17, let us assume that the absorption characteristics
of a layer bounded by interfaces I and II is to be measured. For
the simplicity of description, it is assumed that the spectral
absorption of the substance in the layer is characterized by a
wavelength-dependent attenuation coefficient .mu..sub.h(.lamda.)
and that of other volume is characterized by .mu..sub.g(.lamda.).
It is further assumed that the substance in the vicinity of
interface I (II) possesses an effective and wavelength independent
reflection coefficient r.sub.I (r.sub.II). If the characteristic
absorption of interest is covered by the spectrum of the light
source, an optical filter 1610 or 1710 with a bass band tunable
across the characteristic absorption of the sample 205 may be used
to measure the spectral responses of the sample 205 centered at
different wavelengths.
[0118] In operation, the following steps may be performed. First,
the differential delay modulator 250 is adjusted so that the path
length traveled by one mode (e.g., the mode 001) matches that of
radiation reflected from interface I in the other mode (e.g., the
mode 002). At this point, the pass band of filter 1610 or 1710 may
be scanned while recording the oscillation of the measured signal
due to a periodic differential phase generated by the modulator
250. The oscillation amplitude as a function of wavelength is given
by
A.sub.I(.lamda.)=r.sub.Ie.sup.-2.mu..sup.g.sup.(.lamda.)z.sup.I
(15)
where z.sub.I is the distance of interface I measured from the top
surface of the sample 205. Next, the differential delay modulator
250 is adjusted again to change the differential delay so that the
path length traveled by the mode 001 matches that of radiation
reflected from interface II in the mode 002. The measurement for
the interface II is obtained as follows:
A.sub.II(.lamda.)=r.sub.IIe.sup.-2.mu..sup.g.sup.(.lamda.)z.sup.I.sup.-2-
.mu..sup.h.sup.(.lamda.)z.sup.II (16)
where z.sub.II is the distance of interface II measured from
interface I. To acquire the absorption characteristics of the layer
bounded by the interfaces I and II, Eq. (7) and Eq. (6) can be used
to obtain the following ratio:
A II ( .lamda. ) A I ( .lamda. ) = r II r I - 2 .mu. h ( .lamda. )
z II . ( 17 ) ##EQU00009##
Notably, this equation provides the information on the absorption
characteristics of the layer of interest only and this allows
measurement on the layer. This method thus provides a "coherence
gating" mechanism to optically acquire the absorbance spectrum of a
particular and designated layer beneath a sample surface.
[0119] It should be noted that the pass band of the optical filter
1610 or 1710 may be designed to be sufficiently narrow to resolve
the absorption characteristics of interest and at the meantime
broad enough to differentiate the layer of interest. The following
example for monitoring the glucose level by optically probing a
patient's skin shows that this arrangement is reasonable and
practical.
[0120] Various dependable glucose monitors rely on taking blood
samples from diabetes patients. Repeated pricking of skin can cause
considerable discomfort to patients. It is therefore desirable to
monitor the glucose level in a noninvasive manner. It is well known
that glucose in blood possesses "signature" optical absorption
peaks in a near-infrared (NIR) wavelength range. It is also
appreciated the main obstacle in noninvasive monitoring of glucose
is due to the fact that a probing light beam interacts, in its
path, with various types of tissues and substances which possess
overlapping absorption bands. Extracting the signature glucose
peaks amongst all other peaks has proven difficult.
[0121] The above "coherence gating" may be used to overcome the
difficulty in other methods for monitoring glucose. For glucose
monitoring, the designated layer may be the dermis layer where
glucose is concentrated in a network of blood vessels and
interstitial fluid.
[0122] FIG. 19A illustrates an example of a human skin tissue where
the coherence gating technique described here can be used to
measure the glucose concentration in the dermis layer between the
epidermis and the subcutaneous layers. The dermis layer may be
optically selected and measured with the coherence gating
technique. It is known that the superficial epidermis layer, owing
to its pigment content, is the dominant source of NIR absorption.
Because of the absence of blood, however, the epidermis yields no
useful information for glucose monitoring. The coherent gating
technique can be applied to acquire solely the absorbance spectrum
of the dermis layer by rejecting the absorptions of the epidermis
and the subcutaneous tissues. An additional advantage of this
technique is from the fact that dermis exhibits less temperature
variation compared to the epidermis. It is known that surface
temperature variation causes shifts of water absorption, hampering
glucose monitoring.
[0123] FIG. 19B shows some predominant glucose absorption peaks in
blood in a wavelength range between 1 and 2.5 microns. The width of
these peaks are approximately 150 nm. To resolve the peaks, the
bandwidth of the tunable bandpass filter may be chosen to be around
30 nm. The depth resolution is determined by the following
equation:
2 ln ( 2 ) .pi. .lamda. o 2 .DELTA. .lamda. = 60 m ( 18 )
##EQU00010##
Therefore, the coherence gating implemented with the devices in
FIGS. 16 and 17 or other optical sensing devices may be used to
determine the absorption characteristics of the glucose in tissue
layers no less than 60 .mu.m thick. As illustrated in FIG. 19A,
human skin consists of a superficial epidermis layer that is
typically 0.1 mm thick. Underneath epidermis is the dermis,
approximately 1 mm thick, where glucose concentrates in blood and
interstitial fluids. The above analysis indicates that it is
possible to use the apparatus shown in FIGS. 16 and 17 to isolate
the absorption characteristics of the dermis from that of the
epidermis and other layers.
[0124] It is clear from Eq. (18) that the product of spectral
resolution and layer resolution is a constant for a given center
wavelength .lamda..sub.0. The choice of the filter bandwidth should
be made based on the tradeoff between these two resolutions against
the specific requirements of the measurement.
[0125] The tunable bandpass filter 1610 or 1710 may be operated to
acquire the absorption characteristics of an isolated volume inside
a sample.
[0126] FIG. 20 illustrates one exemplary implementation of the
detection subsystem 260 in FIG. 3 where two diffraction gratings
2010 and 2020 are used to separate different spectral components in
the output light beams from the polarizing beam splitter 361. A
lens 2012 is positioned to collect the diffracted components from
the grating 2010 and focus different spectral components to
different locations on its focal plane. A detector array 2014 with
multiple photodetector elements is placed at the focal plane of the
lens 2012 so that different spectral components are received by
different photodetector elements. A second lens 2022 and a detector
array 2024 are used in the optical path of the diffracted
components in a similar way. In devices shown in FIGS. 5A, 5B, 8A,
and 8B where a single optical detector is used for measurements, a
single grating, a lens, and a detector array may be used.
[0127] In operation, each detector element receives light in a
small wavelength interval. The photocurrents from all elements in
an array can be summed to form a signal which is equivalent to the
signal received in each single detector without the grating shown
in FIG. 3. By selectively measuring the photocurrent from an
individual element or a group of elements in an array, the spectral
information of the sample can be obtained.
[0128] In the above described examples, the optical probe head
sends out light in two different propagation modes where light in
one of the two modes carries the information from the sample.
Alternatively, light in a single propagation mode may be used as
the input light to the optical probe head and as output light from
the optical probe head. Hence, devices based on this design not
only use a common optical path to direct light to and from the
probe head and sample but also control the light in a single mode.
In comparison with above examples where two different modes are
used for light coming out of the probe heads, this single-mode
design further eliminates or reduces any differences between
different modes that propagate in the same optical path.
[0129] FIG. 21 shows one exemplary system for acquiring information
of optical inhomogeneity and other properties in substances with
only one propagation mode inside waveguides. A broadband or
low-coherence light from Broadband Light Source 201 is directed to
a probe head 2110 by means of polarization-maintaining waveguides
271 and 272. A partial reflector inside the probe head 2110
reverses the direction of a small portion of the input light to
create a radiation wave 1 while transmitting the remainder of the
input light to the sample 205. Backscattered or reflected light
from the sample 205 becomes a second radiation wave 2 and is
collected by the probe head 2110. The probe head 2110 combines and
couples both the radiation waves 1 and 2 back into the waveguide
272. The radiation waves 1 and 2 travel in the waveguide 272
towards Light the light director 210 which directs radiation waves
1 and 2 through the waveguide 273 towards the detection module
2101. Notably, the radiation waves 1 and 2 output from the probe
head 2110 are in the same mode as the input light to the probe head
2110. the probe head 2110 does not change the mode of light when
directing the radiation waves 1 and 2 to the waveguide 272.
[0130] The detection module 2101 includes a beam Splitter 2120, two
optical paths 2121 and 2122, an optical variable delay element 2123
in the path 2122, a beam combiner 2130, and two optical detectors
2141 and 2142. The beam splitter 2120 splits the light in the
waveguide 273, which includes the radiation waves 1 and 2 in the
same mode, into two parts that respectively propagate in the two
optical paths 2121 and 2122. Notably, each of the two parts
includes light from both the radiation waves 1 and 2. The variable
delay element or delay line 2123 in the optical path 2122 is
controlled by a control signal to adjust the relative optical delay
between the two optical paths 2121 and 2122 and may be implemented
by, e.g., the exemplary delay elements described in this
application and other delay designs. The beam combiner 2130
combines the signals of the two optical paths to overlap with each
other and to output two optical signals for optical detectors 2141
and 2142, respectively. The beam combiner may be a polarization
beam splitter which splits the combined light into two parts,
orthogonal in polarization to one another.
[0131] The probe head 2110 may include a partial reflector to
produce the radiation wave 1 which does not reach the sample 205.
Assuming the single propagation mode for the light to the probe
head 2110 and the light out of the probe head 21110 is a
polarization mode, the light reflected from the partial reflector
in the probe head 2110, i.e., the radiation wave 1, has the same
polarization as the light collected from the sample, the radiation
wave 2. Therefore, both Radiation 1 and 2 travel in the same
propagation mode in the waveguides, 272 and 273. Because the
radiation waves 1 and 2 are reflected from different locations,
they experience different optical path lengths when reaching the
beam splitter 2120. The effect of variable delay element 2123 is to
add an adjustable amount of the delay in the light in the path 2122
relative to the light in the path 2121.
[0132] In operation, the variable delay element 2123 can be
adjusted so that the partial radiation 1 reaching the polarization
beam splitter 2130 through the path 2122 can be made to experience
a similar optical path length as the partial radiation 2 reaching
the beam splitter 2130 via the other path 2121. The superposition
of the two beams at the photo detectors 2141 and 2142 causes a
measurable intensity variation as their relative path length is
being varied by the variable delay element 2123. This variation can
be utilized to retrieve information on the inhomogeneity and other
properties of the sample 205.
[0133] FIG. 22 shows an exemplary implementation of the system in
FIG. 21 using polarization maintaining optical fibers. A
polarization controller 202 may be placed at the output of the
light source 201 to control the polarization of the input light in
one polarization mode. The optical head 2110 is shown to include a
lens system 2111 and a partial reflector 2112. Two mirrors 1 and 2
are used to construct the two optical paths between the beam
splitters 2120 and 2130. The optical radiation reflected from the
partial reflector 2122 and from the sample 205 travel in the
polarization-maintaining (PM) fiber 272 in the same mode. The main
portions of the radiation waves 1 and 2 are deflected to the mirror
1 while the remaining portions are directed to the mirror 2 by the
beam splitter 2120.
[0134] The incident plane of the polarizing beam splitter 2130 can
be made to have a finite angle with respect to the polarization
directions of light from both the Mirror 2 in one optical path and
the variable delay element 2123 from the other optical path. In
this configuration, light energies received by both detectors 2141
and 2142 are the superposition of the two radiations, i.e.,
Radiation 1 and Radiation 2. It should be appreciated that the
linkage between the beam splitters 2120 and 2130 can be made by
means of optical fibers or other optical waveguides to eliminate
the free space paths and the two mirrors 1 and 2.
[0135] In the examples shown in FIGS. 21 and 22, the spacing
between the optical head 2110 and the sample 205 may be greater
than the sample depth of interest so that, upon reaching the beam
splitter 2130, the partial radiation 1 experiences optical path
length similar only to that of partial radiation 2. In other words,
split parts of the same radiation do not experience similar optical
path length during the operation of the systems in FIGS. 21 and
22.
[0136] FIG. 23 shows one exemplary optical arrangement for the
probe head 2110. The partial reflector 2310 can be realized with a
partially reflective fiber termination, i.e., the end facet of the
fiber 272. An uncoated fiber tip has a reflectivity of
approximately 4% and thus may be used as this partial reflector.
Optical coating on the end facet may be used to change the
reflectivity to a desirable value.
[0137] The reflectance of the fiber termination 2310 may be chosen
based on several factors. In one respect, the radiation wave 1
should be strong enough so that its superposition with the
radiation wave 2 creates an adequate intensity variation at the two
detectors 2141 and 2142. On the other hand, the radiation wave 1
may not be too strong as it may overwhelm the photodetectors 2141
and 2142, prohibiting the use of high gain in the detection
systems. For optimized operation of the system, one may want to
choose the reflectance of the fiber termination to be comparable to
the total light collected by the fiber from the sample.
[0138] In FIGS. 21 and 22, a common waveguide 272 is used for both
sending input light into the probe head 2110 and directing output
light output the probe head 2110. Alternatively, similar to the
design in FIG. 15, the waveguide 272 may be replaced by an input
waveguide for sending input light into the probe head 2110 and an
output waveguide directing output light output the probe head 2110
to the beam splitter 2120 of the detection module 2101. In this
design, the light director 210 can be eliminated and the optical
probe head 2110 may be designed to direct output light with both
the radiation waves 1 and 2 into the output waveguide.
[0139] Similar to tuning the frequency of light in other examples
as described, in implementing the devices in FIGS. 21 and 22, a
tunable optical bandpass filter may be used to tune the frequency
band of the light to selectively measure the property of the sample
205 at the frequency band of the filter. In addition, the use of
gratings in the detection module to measure different spectral
components of the sample as shown in FIG. 20 may be used in the
module 2101 as well.
[0140] FIG. 24 further illustrates the measuring technique for
optically targeting a layer underneath the surface of a body for
its spectral absorbance. Referring to Equations (15)-(17), the
optical differential delay can be adjusted to obtain the
measurements A.sub.I and A.sub.II from the two depths I and II in
order to obtain measurement for the layer between the depths I and
II. If the center wavelength of the light source .lamda. is scanned
to obtain measurements at different wavelengths, the measured ratio
in Equation (17) can be used to obtain spectral absorption
characteristics of the substance bounded by interfaces I and II
only, i.e. .mu..sub.h(.lamda.). Therefore, this techniques
effectively isolates the substance between I and II in terms of its
spectral absorbance for the measurement. This procedure can be
carried out for all layers, by varying the depths of the interfaces
I and II, to obtain a cross-sectional spectral absorbance mapping
(SAM).
[0141] One way to obtain SAM measurements is to first obtain the
cross-sectional maps of the reflectance, A(.lamda.), at two or more
different wavelengths using light radiations centered at these
wavelengths. When a single light source is used as described above,
a tunable optical filter is used to select the different
wavelengths at each spatial location of the probe head over the
target area to obtain measurements. Upon completing measurements at
different wavelengths at one location, the probe head is moved to
the next location and the measurements repeat. This process
continues until all locations within the target area are measured.
This use of the optical differential delay at variable delay values
and the scan along the target surface in combination effectuates a
3-dimensional mapping of the spectral absorbance of the target
area.
[0142] In some applications where the sample has absorption
features in a broad spectral range, a single light source may not
be able to provide a sufficiently broad spectral coverage over
these absorption features. The following sections describe
techniques that use two or more light sources with radiations
centered at different wavelengths to provide a broad spectral
coverage in SAM measurements.
[0143] Various optical arrangements described here can be adopted
for performing SAM measurements. Several examples are described
below for using multiple light sources at different
wavelengths.
[0144] FIG. 25 shows an optical device 2500 that uses two or more
different light sources 2510 at different optical wavelengths to
obtain reflectance maps from a sample. Each light source emits
within a bandwidth .DELTA..lamda. centered at a different
wavelength from other light sources. The wavelengths of the light
sources 2510 can be selected to cover the spectral range of the
absorption features in samples to be measured. In some
applications, the wavelengths of 2510 may be selected to
effectively sample a specific absorbance feature of interest, as
shown in FIG. 31. and thus may not cover other absorption features
in the sample. In a specific implementation, the bandwidth
.DELTA..lamda. of each light source should be selected with
consideration of the depth spatial resolution desired for the
measurements. An optical multiplexer 2520 is used to receive the
optical radiations from different light sources 2510, to combine
these optical waves into a common optical path, i.e., the common
optical waveguide 271. The light director 210 directs the combined
optical radiation to the probe head 220 via a common waveguide 272.
The probe head 220, which is positioned above the sample 205, split
a portion of light from the multiplexed or combined optical
radiation as the probe light and direct this probe light to the
sample 205. The reflected light from the sample 205 is collected by
the probe head 220 and is directed to the differential delay
modulator 250 via the waveguide 272, the light director 210 and
another waveguide 273. Details, various implementations and
operations of the device 2500 are described in previous sections.
An optical demultiplexer 2530 is further used to separate the light
output from the differential delay modulator 250 spatially based on
different wavelength bands centered at the different wavelengths of
the light sources 2510. Accordingly, an array of different optical
detector modules 2540 are used to respectively receive and detect
the separated beams of different wavelength bands. As an example,
light radiation centered at the wavelength .lamda..sub.1 and within
the bandwidth of .DELTA..lamda..sub.1 from one light source is
separated from the rest and sent to the detector module 1 (D1).
Each detector module may include one or multiple optical detectors.
The differential delay modulator 250, the demultiplexer 2530 and
the detector modules 2540 form at least part of an optical
correlator which performs the optical detection of the device 2500.
The multiplexed light radiations are delivered to the tissue
through the optical waveguide 272 or fiber and the probe head 220.
Backscattered and reflected light from the tissue is collected in
part by the probe head 220 and redirected to the optical
correlator.
[0145] In practice, the probe head 220 is operated to scan the
multiplexed light radiation over the sample 205 to obtain
measurements at different wavelengths. For every designated spatial
interval the differential delay modulator 250 scans over a range to
correspond to a range of depth inside the sample. This process
repeats until all sampling locations of an area of the sample are
measured. In this implementation, cross-sectional maps for light
radiations at two or more wavelengths can be simultaneously
obtained. While the differential delay modulator 250 and the
probing light radiation are being scanned, the photocurrents from
the detector modules 2540, each receiving light radiation within a
different wavelength band associated with one of the light sources
2510, can be simultaneously recorded as the data from which the
multiple reflectance maps, A(.lamda..sub.1), A(.lamda..sub.2) and
so on, can be extracted. Each reflectance map is formed by
radiation within the band of one light source. These reflectance
maps can then be used to derive SAM using an algorithm based on the
principles outlined by Equations (15) through (17).
[0146] FIG. 26 shows one implementation 2600 of the device 2500 in
FIG. 25 where a digital signal processor (DSP) 2610 is used to
process the detector outputs from the detector module 2540 and to
produce the spectral absorption map. The DSP 2610 may be part of
the device controller 370 shown in FIG. 3. A display and user
interface module 372 is used to allow an operator to view the SAM
result and to control the device.
[0147] FIG. 27 shows an example of the device 2600 in FIG. 26 where
the demultiplexer 2530 is implemented with two gratings 2010 and
2020 and two lenses 2012 and 2014. The detector modules 2540 are
implemented with two detector arrays 2014 and 2024, i.e., each of
the detector modules 2540 includes one detector in the array 2014
and another detector in the array 2024 for detecting light at the
same wavelength and in different polarization states. The
polarization beam splitter 361 split the wavelength multiplexed
light from the differential delay modulator 250 into two beams with
mutually orthogonal polarization states where each split beam is a
mixture of light in two different modes from the probe head 220.
The polarization beam splitter 361 converts a part of received
light in the first propagation mode and a part of received light in
the second propagation mode into light in a third propagation mode
that propagates along a first optical path and to convert remaining
portions of the received light in the first and the second
propagation modes into light in a fourth propagation mode that
propagates along a second, different optical path. The third and
fourth modes are two orthogonal polarization modes of the
polarization beam splitter 361.
[0148] The gratings 2010 and 2020 separate the wavelength
multiplexed light radiation into angle intervals, each
corresponding to the light from one of the light sources 2510. The
number of photosensitive elements in one detector array can be
equal to the number of light sources used. The sensing area of each
of the photosensitive elements may be designed to be sufficiently
large so that all the light radiation within the band of one light
source can be received by one element in the array. For instance,
if three light sources are used in the system, two arrays each with
three photosensitive elements may be used.
[0149] The optical multiplexer 2520 may be implemented in various
configurations. FIG. 28 shows one example of the multiplexer 2520
where partially reflective mirrors 2810 and 2802 are used to
multiplex radiation beams from three different light sources 2510A,
2510B and 2510C. The design can be used with N partially reflective
mirrors to multiplex beams from (N+1) light sources. The partially
reflective mirrors can be manufactured by coating one side of a
glass with a thin metal layer. With this arrangement not all the
light power will be multiplexed into the optical fiber, as loss of
optical power occurs at each reflector. An optical collimator 2810
is used to couple the multiplexed light into the optical waveguide
or fiber 271.
[0150] FIG. 29A shows another example of the multiplexer 2520 which
reduces the optical loss in the design in FIG. 28 and provides an
efficient use of the available optical power. In this example,
optical dichroic filters 2901 and 2902 are used to replace the
partially reflective mirrors 2801 and 2802, respectively. A
Dichroic filter may be implemented in various forms. One
implementation is to use two short-pass interference filters as the
dichroic filters 2901 and 2902.
[0151] FIG. 29B shows the optical designs of the dichroic filters
2901 and 2902. The cut-off wavelength of the filter 2901 is set
between the radiation bands of the first and second light sources
centered at .lamda.1 and .lamda.2, respectively; that of the filter
2802 set between the radiation bands of the second and the third
light sources centered at .lamda.2 and .lamda.3, respectively. With
this arrangement, except for the imperfection of the filters, all
radiation from the three light sources are coupled into the optical
fiber 271 without significant optical loss. Interference optical
filters of this kind can be fabricated using multilayer dielectric
thin films. Other possible multiplexers include arrayed waveguide
type and grating type.
[0152] In the above devices for SAM measurements, light beams at
different wavelength bands are simultaneously directed by the probe
head 220 to the sample 205. Hence, the optical measurements at
different wavelengths are performed simultaneously. Alternatively,
the optical multiplexer 2520 may be replaced by an optical switch
3010 as shown in FIG. 30 to direct a probe beam within one of the
wavelength bands at a time so that probe light beams at different
wavelength bands are directed to the sample 205 sequentially at
different times to obtain the reflectance maps. In one
implementation, the N broad band light sources 2510 can be
sequentially linked to the optical device in FIG. 30 through a
1.times.N optical switch as the switch 3010. The reflective maps,
A(.lamda..sub.1), A(.lamda..sub.2) and so on at different
wavelength bands are obtained sequentially and are then used in the
calculation of SAM.
[0153] The choice of the broadband light sources in any of the
above device designs can be made according to the specific
absorption features to be measured. As an example, FIG. 31 shows an
absorbance amplitude spectrum 3100 of a sample where an absorption
peak 3110 is present. Three or more different broadband light
sources with the center wavelengths shown may be used to map the
peak 3110. To achieve a better spectral resolution, the number of
the light sources may be increased. In the example of three light
sources, the reflectance maps at the three different wavelengths,
i.e., A(.lamda..sub.1), A (.lamda..sub.2) and A(.lamda..sub.3), can
be used to calculate the strength of the feature for SAM.
[0154] The axial resolution (i.e., the depth resolution) of SAM is
related to the bandwidth (spectral width), .DELTA..lamda., of the
light source at a center wavelength .lamda..sub.0 is given by the
following:
.DELTA. z = 2 ln ( 2 ) .pi. .lamda. o 2 .DELTA. .lamda. ( 23 )
##EQU00011##
For a given bandwidth, .DELTA..lamda., the depth resolution of the
corresponding reflectance map is determined by the above equation.
Hence, a broad bandwidth is desirable for resolving a small spatial
feature along the direction of the probe beam, which limits the
spectral resolution as a tradeoff. For example, if one wants to map
an spectral absorbance feature that occupies a 20 nm range near an
optical wavelength of 1 .mu.m, light sources of bandwidth around 5
nm can be chosen. Under these conditions, the spatial resolution
for SAM is roughly 90
[0155] In the above multi-source SAM measurements, each light
source has a fixed emission center wavelength and a bandwidth. In
other implementations based on the above-described designs,
multiple tunable laser sources may be used to replace the fixed
light sources. Each tunable laser source may be configured to
provide highly coherent radiation over a wavelength range of
.DELTA..lamda. centered at .lamda.. Due to the same consideration
that a spectral absorbance feature of interest may be too broad for
a single tunable laser source to cover, two or more tunable laser
sources, each tunable over a wavelength range centered at a
different wavelength, can be implemented in various designs for SAM
measurements.
[0156] FIG. 32 illustrates one example of a device 3200 for SAM
measurements where two or more tunable lasers 3210 are used as the
light sources. The optical radiations from the tunable laser
sources 3210 are combined through the multiplexer 2520 before being
guided to the probe head 220. The light waves from the probe head
220, including what is collected from the tissue under examination,
are redirected towards the demultiplexer 2530 where they are
separated into separated beams in different optical paths according
to the wavelength bands, each of which is received by an optical
detector.
[0157] This arrangement may be configured to allow for the
simultaneous tuning of the wavelengths of the tunable laser sources
3210, which in turn allows for the simultaneous recording of the
light waves from the probe head 220 in the different wavelength
bands. One feature in the design in FIG. 32 is the lack of the
optical differential delay modulator 250 used in other designs
where one or more fixed light sources are used to produce the probe
light. Each tunable laser is tuned through its tuning spectral
range during the measurement and the recorded light intensity as a
function of the laser wavelength in each of the bands can be
computed to obtain a reflectance map for that band. The reflectance
maps, for the various center wavelengths, can be computed by
analyzing the photocurrents of the photodetectors as functions of
the wavelength. A variation of the photocurrent with a certain
wavelength periodicity indicates a reflection originated from a
certain distance, or depth, in the sample 205. Such a computation
is, in essence, a decomposition of the photocurrent according to
its frequency, or commonly known as a Fourier transformation. In
order for the reflectance maps to cover a range of the depth the
tunable lasers should have an adequate coherent length which is
comparable or longer than the range of the depth. Two or more
reflectance maps for two or more wavelength bands can be obtained
for the sample 205 under examination and can be used to derive the
SAM of the sample 205 based on Equations (15)-(17).
[0158] This use of the tunable lasers may be implemented in the
various device designs for SAM measurements by removing the optical
differential delay modulator 250. For example, the design in FIG.
30 may be used, without the differential delay modulator 250, to
sequentially direct light radiations from different tunable laser
sources to the sample 205. When the radiation from a particular
tunable laser is directed to the sample 205, the laser is tuned in
its laser frequency through its tuning range to obtain measurements
of the optical absorption at different wavelengths within the
tuning range.
[0159] In some precise optical phase measurements using the above
described techniques with tunable laser sources, a differential
phase modulator 250 may be inserted in the common waveguide 273 to
receive the light from the probe head in the first and second
propagation modes and to produce and modulate the relative optical
phase between the first and second propagation modes. The
modulation of the relative optical phase between the first and
second propagation modes causes the photocurrents out of the
photodetectors 2540 (or detectors in the detector arrays 2014 and
2024) to shift their peak positions and valley positions with
respect to the wavelength. This allows for accurate calculations of
the reflected optical phase of the light reflected from the sample
using mathematical analysis similar to the analysis represented by
Equations 12 and 13.
[0160] As an application of the above non-invasive optical probing
techniques and devices, FIG. 33 shows an example of an integrated
system 3300 that combines an X-ray CT scan module 3310 for locating
pulmonary nodules, a minimally invasive optical probing module
3320, and a treatment module 3330 to provide a complete diagnostic
and treatment platform for treating lung cancer. The treatment
module 3330 may be designed to use electromagnetic radiation, such
as laser radiation, RF or microwave radiation energy, to treat a
malignant condition at a selected target area. A bronchoscope 3340
is used to provide a means for inserting the optical probe for the
optical probing module 3320 into the lung to optically measure a
target area in the lung. In addition, the bronchoscope 3340 is also
used to guide the laser beam from the laser treatment module 3330
to the lung for laser treatment. As illustrated, the bronchoscope
3340 includes a working channel 3342 that is hollow and receives
the optical probe head and optical fiber 3322 for the optical
probing module 3320 and an optical power delivery waveguide 3322
for the laser treatment module 3330. The working channel 3342 is
inserted inside the lung to probe different targeted areas in the
lung. The distal end of the working channel 3342 includes an end
facet or window that transmits both the optical probe light and the
laser beam from the laser treatment module 3330. A computing and
control module 3350 is provided to control the three different
modules 3310, 3320 and 3330 and to perform analysis on the
measurements. A display and user interface module 3360, which may
include a user input interface and a display monitor, is used to
allow an operator to operate the system 3300.
[0161] The CT scan module 3310 is used to scan the lung of a
patient to detect and locate all solitary pulmonary nodules (SPNs).
Each SPN is visually located via the CT scan imaging. Next, the
optical probing module 3320 is used to measure each SPN identified
by the CT scan. This is a differential diagnosis and the optical
measurement is analyzed to determine whether each SPN is benign or
malignant. The laser treatment module 3330 is then used to treat
each malignant SPN. All three procedures can be performed in one
integrated system.
[0162] The minimally invasive optical probing module 3320 may be
implemented in various embodiments as described in this
application. As a specific example, the optical probing module 3320
may be implemented as a cross-sectional imaging module. The optical
module 3320 can be used to allow the anticipated use of CT scans in
early stage lung cancer diagnosis and, in addition, can facilitate
cancer therapy using optical methods such as Laser Hyperthermia.
The module 3320 utilizes optical correlation techniques to obtain
optical tomographs to non-destructively reveal the tissue structure
and other physiological information. The probe head of the imaging
module 3320 is fiber optic-based and is inserted into the working
channel 3342 of the bronchoscope 3340. The bronchoscope 3340 has
been previously used to visually locate the tumor inside the lung.
A sequential, in-vivo examination of the suspect tissue or SPN with
the optical probing module 3320 can distinguish a calcified, benign
SPN from a malignant one by virtue of their different structure and
optical properties. This use of the optical probing module 3320
resolves the CT scan diagnostic dilemma, enabling an minimally
invasive procedure to locate SPNs and then identify which nodules
are malignant. Notably, the use of this diagnostic sequence based
on the optical probing allows the physician to avoid most, if not
all, pulmonary biopsies, thereby significantly reducing the risks
discussed above and greatly improving chances for a successful
diagnosis without side effects.
[0163] FIG. 34 illustrates one exemplary use of the system 3300 in
FIG. 33. The CT scan is used to perform the initial examination of
the lung to scan for all SPNs, benign and malignant. After the CT
scan, the optical probing is performed at each detected SPN to
determine whether the SPN is benign or malignant. If a SPN is
determined to be malignant, the laser treatment can be performed to
treat the malignant condition of the SPN by using the laser
treatment module 3330. If no malignant SPN is found, the patient
may be scheduled for periodic CT scans to monitor the condition of
the lung.
TABLE-US-00001 TABLE 1 Assumed Tumor diameter 1 cm. Approx. Tumor
Volume 0.5 cc Optical Power delivered 0.5 watt 1.0 watt 2.0 watts
Estimated Laser Power 1.5 watt 3.0 watt 6.0 watt Temperature Rise
for: 5 sec. Exposure +5 C. +10 C. +20 C. 10 sec. Exposure +10 C.
+20 C. +40 C.
[0164] The laser treatment may be implemented in various
configurations such as laser hyperthermia treatment and laser
ablation treatment. For example, a pulmonologist may use a high
power laser in the laser treatment module 3330 and an optical
fiber-based therapeutic probe inserted into the working channel
3342 of the bronchoscope 3340 to deliver optical power to the
tumor. This procedure, called Laser Hyperthermia, has been shown to
necrotize cancerous tissue. The laser emission wavelength is chosen
so that essentially all of the light is absorbed by the tissue,
e.g., within first centimeter of tissue. Several types of high
power laser sources may be used. For example, compact, powerful
diode-pumped solid state lasers are readily available. Optical
fibers capable of transmitting substantial power levels (e.g., on
the order of watts) are also available. We estimate that coupling
of the laser optical power to the fiber can be accomplished with
approximately 33% efficiency using normal methods known to
practitioners in this field.
[0165] As an example, Table 1 lists calculated exposure times
needed to elevate the temperature of the suspect tissue for
different optical power levels delivered to the tissue. The laser
power input to the optical delivery waveguide 3332 (e.g., optic
fiber) would need to be three times higher assuming 33% coupling
efficiency. In the above estimates, it is assume that the malignant
tissue behaves thermally as if it were water (about 70% accurate)
and that the nodule is essentially in poor thermal contact with the
surrounding tissue. Researchers have found that a 10.degree. C.
rise in temperature is sufficient to kill cancer cells and that
higher temperature rises kill malignant cells more quickly. Base
upon the results of Table 1, a 3-6 watt laser should suffice to
perform Laser Hyperthermia in-vivo with a 5-10 sec. exposure.
[0166] The integrated diagnostic and therapeutic system in FIG. 33
and the technique in FIG. 34 may be implemented to allow both SPN
location/detection and bronchoscopic examination to be performed in
a single session or visit. In addition, both differential diagnosis
and laser therapy can therefore be performed during a single
bronchoscopic procedure. Therefore, The three different procedures,
SPN detection/location, malignant-benign differential diagnosis,
and remedial therapy, can be performed in a single office visit. A
CT Scan system may be modified to incorporate the much smaller
optical devices for differential diagnosis and laser therapy so
that the complete process may be performed on a single piece of
equipment. This results in very efficient use of the physician's
time and convenience for the patient. Laser therapy methods, such
as laser hyperthermia and laser ablation, do not have significant
adverse side effects on the patient under treatment and thus are
advantageous in this regard in comparison with other therapeutic
regimens such as chemotherapy and radiation. In addition, the
integrated system in FIG. 34 may be implemented to reduce any delay
in the differential diagnosis and therapy and thus such
implementation can be advantageous over other methods that use the
`wait-and-see` observation of tumor size growth protocol which is
often employed to distinguish between malignant and benign
SPNs.
[0167] The integrated system in FIG. 33 may also be implemented by
using treatment modules other than laser therapy modules. Various
electromagnetic radiation therapies using the radiofrequency (RF)
energy and microwave energy for ablation may be used. An RF or
microwave waveguide probe may be inserted into the working channel
3342 to deliver the RF or microwave energy to a targeted SPN for
treatment. For example, the laser treatment module 3330 may be
replaced by a microwave ablation therapy unit. The distal end facet
of the working channel can be made to transmit both the probe light
and the RF/microwave radiation.
[0168] In addition, the integrated design shown in FIG. 33 may also
be implemented for diagnosing and treating other illness. In one
implementation, for example, an integrated diagnostic and treatment
system may include a CT scan unit to locate ailing areas in a body
part, a referenced cross-sectional imaging unit to analyze each
ailing area, and a laser, RF or microwave irradiation therapy unit
to treat a selected area. This system may be used to diagnose and
treat lung cancer, prostate cancer and other tumors. One specific
implementation of this system is the example in FIG. 33 for
diagnosing and treating lung cancer where a bronchoscope is
inserted into the lung for deliver the probe light and the
treatment laser beam.
[0169] In implementing the system in FIG. 33, the optical probe
head 220 of the optical imaging module 3320 and the therapy
delivery waveguide 3332 may be unified as a single assembly when
inserted inside the working channel 3342 so that the treatment
radiation, which may be laser radiation, RF or microwave radiation,
can be directed to approximately the same location where the
optical probe head 220 is located. In this design, when a SPN is
identified as malignant, the treatment radiation can be delivered
to the same location where the malignant SPN is without changing
the location of the distal tip of the working channel 3342.
Therefore, this unified assembly may be used to simply the
alignment of the treatment radiation with respect to a malignant
SPN identified by the optical imaging module 3320.
[0170] FIG. 35 shows one example of a unified assembly 3500. A
tubular unit or sheath 3510 is used to hold the probe fiber 3322
and the waveguide 3332 together as a single unit. The probe head
220 at or near the end of the fiber 3322 and the distal end of the
waveguide 3332 are placed next to each other at the distal end of
the tubular unit 3510 within an end facet window 3520. As such, the
probe head 220 and the distal end of the waveguide 3332 are aimed
at the essentially the same location. The assembly 3500 is then
inserted inside the working channel 3342 to place the end facet
window 3520 at the end of the working channel 3342.
[0171] FIG. 36 shows an example of a thermotherapy device 3600
based on optical imaging and controlled delivery of thermal energy
guided by the optical imaging. A catheter includes a working
channel 3342 that is configured for insertion into a passage of a
body to reach a target tissue inside the body. An optical imaging
module 3630 includes (1) an imaging optic fiber 3322 having a
portion inserted into the working channel 3342 and an optical probe
head coupled to an end of the imaging optic fiber 3332 and located
inside the working channel 3342. The optical imaging module 3630 is
operable to direct probe light to and collect reflected light from
the target tissue in the body through the imaging optic fiber and
the optical probe head and to obtain imaging information of the
target tissue from the collected reflected light. This device 3600
includes a thermotherapy module 3640, e.g., a laser thermotherapy
module, an RF or microwave thermotherapy module, having a power
delivery waveguide having a portion inserted into the working
channel 3342 to deliver thermal energy to the target tissue. A
control unit is provided to include a computing and control module
3610 and a cooling control module 3620. The control module 3610
includes a processor that processes the image data and local memory
to store data and processing software for performing various
processing and control tasks. The control unit controls the optical
imaging module 3630 to extract the imaging information from the
collected reflected light, to obtain a spatial distribution of
diseased locations of the target tissue, and to obtain a
temperature map of the target tissue for thermotherapy based on the
spatial distribution of the diseased locations of the target
tissue. The control unit controls the thermotherapy module 3640 to
control a location and an amount of thermal energy delivery to each
of the diseased locations based on the temperature map to perform
thermotherapy. A cooling liquid circulation module 3622 is included
as part of the cooling for the device that cools the surface of the
target tissue.
[0172] FIG. 37 shows an example of a catheter design with a cooling
mechanism for thermotherapy based on laser thermotherapy. Features
in this example can be used for other thermotherapy such as RF or
microwave thermotherapy. A liquid balloon or reservoir 3710 is
provided at the tip of the catheter to receive a cooling liquid
3720 from the cooling liquid circulation module 3622. In this
example, the liquid balloon 3710 encloses the end facet of the
catheter tip or window 3520 through which the imaging light 3701
and the thermal treatment radiation such as the treatment laser
3702 are directed to the target tissue. The material for the liquid
balloon is transparent to the imaging light and the treatment
light. Various balloon designs can be used in this design,
including designs that use two or more balloons that contain the
cooling liquid. In operation, the liquid balloon is placed in
contact with the surface of the target tissue under the
thermotherapy treatment to cool the tissue surface and to prevent
undesired thermal damage to the tissue surface.
[0173] FIG. 38 shows another example of a catheter design for
thermotherapy. An external tubing is provided to house a
thermotherapy catheter that delivers the therapeutic thermal energy
to the target tissue and an optical imaging catheter for performing
the optical imaging of the target tissue. These two catheters can
be retracted in their positions within the external tubing. The
external tubing includes a narrow tip which allows only one
catheter to be present at a time. The two catheters are controlled
so that one catheter can be inserted into the narrow tip for
interacting with the target tissue while the other catheter is
retracted back in the external tubing. FIG. 38A shows the imaging
catheter is extended into the narrow tip while the thermotherapy
catheter is at its retracted position. FIG. 38B shows the imaging
catheter is retracted from the narrow tip while the thermotherapy
catheter is extended into the narrow tip.
[0174] FIG. 39 shows an example method of thermotherapy based on
optical imaging. In this example, the thermotherapy is performed by
directing an imaging optical beam to a target tissue to obtain
image information; processing the obtained image information of the
target tissue to obtain a spatial distribution of the diseased
locations of the target tissue; generating a temperature map of the
target tissue for thermotherapy based on the spatial distribution
of the diseased locations of the target tissue; and controlling the
thermal energy delivery to each of the diseased locations and
cooling at the surface of each diseased locations based on the
temperature map to perform the thermotherapy.
[0175] As a specific example, the above described devices and
techniques for thermotherapy can be used to effectively downstage
asthma. An asthmatic attack occurs when a stimulus [smoke, pollen
etc.] evokes an exacerbation in the airway smooth muscle causing it
to contract and engorge with histamines and mast cells. The
resulted decrease in airway lumen diameter interferes with normal
breathing and can be so severe as to completely block the airway,
or even cause death in extreme cases.
[0176] Referring to FIG. 36, the laser thermotherapy module 3640
can be used to heat the airway smooth muscle [ASM], imbedded in the
airway wall, to the point where it no longer responds strongly to
an asthmatic stimulus. The result is a reduction in severity of
asthmatic exacerbations, which can then be managed by inhaled drug
therapies. ASM is not uniformly distributed within the airway wall
and in larger airways may only occupy 10-15% of the wall area.
Without guidance, a `blind` thermotherapeutic procedure may miss
the ASM over 80% of the time. In addition, failure to position the
therapeutic catheter in closest proximity to the ASM may result in
insufficient treatment temperatures. The present image-guidance
provided by the optical imaging module 3630 allows optimal
positioning of the thermotherapy catheter with respect to the ASM.
Furthermore, real-time temperature profile mapping permits
controlling the temperature to avoid over treatment or
undertreatment.
[0177] In some implementations, the image-guided thermotherapy can
include microstructural imaging of airway wall cross-sections to
locate ASM, laser thermotherapy using image-guidance to treat ASM
locations which can use spiral pullback of the thermotherapy
catheter through airway lumen for a quasi-continuous therapeutic
process, and real-time thermal profile mapping [imaging] for
monitoring and or control of the laser thermal therapeutic process
temperature thereby reducing adverse events. Referring to FIG. 39,
the obtained 3D image information of the target tissue is processed
to obtain a spatial distribution of the diseased locations of the
target tissue. A modeling process can be used to construct a
temperature map of the target tissue for thermotherapy based on the
spatial distribution of the diseased locations of the target
tissue. This modeling process can be implemented in software in the
control processor of the device. As the images and the
corresponding spatial distribution of the diseased locations of the
target tissue are updated in real time during the process, the
modeling performed by the control processor generates real-time
thermal profile mapping and this information can be used for
monitoring and or control of the laser thermal therapeutic process
temperature. The thermal energy delivery to each of the diseased
locations and cooling at the surface of each diseased locations can
be controlled based on the temperature map to perform the
thermotherapy. Also, real time pre-, peri- and post-operative
imaging of treatment site can be provided. In addition, real-time
follow-up imaging at predetermined intervals can be used to monitor
therapeutic efficacy.
[0178] Real time optical imaging is provided by the optical imaging
module to obtain high resolution, real-time cross-sectional images
of the human bronchial wall. The imaging catheter illuminates the
tissue being scanned with a narrow beam of low power, near infrared
light and collects the light reflected as the beam penetrates the
tissue. The image is obtained by mechanically scanning the imaging
catheter across the region of interest while acquiring data on
reflected light. If this is done while slowly continuously
withdrawing the catheter, an helical [scan] micro image of the
airway wall can be obtained in much the same way as helical scan CT
operates on a macroscopic basis. In this manner the ASM may be
"visualized" for more efficacious laser thermotherapy.
[0179] The laser thermotherapy is integrated or compatible with the
optical imaging so that the imaging information can be used to
provide guidance in positioning the thermotherapy catheter and to
provide guidance on the temperature profile control. For example,
the laser thermotherapy module can use light at 1.3 micron by using
the radiation from a Nd:YAG laser to provide effective laser
thermotherapy. The heat depositing system can be implemented to
include a laser source producing laser light which is coupled into
the tissue imaging system at the proximal end with controllable
switch. The laser source may contain more than one wavelength. At
the distal end, the energy is deposited and image is obtained
through the same integrated catheter.
[0180] One implementation of the catheter includes a common fiber
that directs imaging radiation and laser treatment radiation via
common fiber core or collinear fiber cores. The catheter includes
re-imaging and redirecting optics common for treatment laser energy
and imaging radiation energy. The fiber can be configured to
support different modes so that treatment laser energy profile on
the fiber can be enlarged to prevent overheating and to control
energy deposition area. The catheter can also include substantially
transparent sheath or window, that might be filled with cooling
liquid. For larger airways, expandable balloons may be used for
better contact with internal surface of airway. Also direct
irrigation scheme can be used for better cooling of airway surface,
preferably with two balloons.
[0181] The above technique can be used to obtain 3D image of a
tissue of interest such as smooth muscle via rotating catheter and
pulling it back while keeping image registration to the
bronchoscope and/or external reference frame. Then pull back is
repeated with energy deposited upon the tissue of interest. Some
examples of such optical imaging devices and techniques are
described in U.S. Pat. No. 7,706,646 entitled "Delivering Light via
Optical Waveguide and Multi-View Optical Probe Head" which is
incorporated by reference as part of this document. The lateral
energy deposition pattern is controlled by synchronizing energy
level with catheter position. The axial deposition of heat is
controlled by applying cooling liquid in the lumen and selecting
spectral bandwidth of laser source and pulsing pattern as a
function of catheter position. Additionally, temperature maps can
be obtained and energy deposition pattern is fine-tuned for better
overall of temperature map with targeted tissue.
[0182] In another implementation, the catheter can include an
imaging fiber with re-imaging and directing optics and a second,
separate fiber with pattern forming optics for treatment laser.
Pattern forming optics is understood to include radially
symmetrical patterns as well as any arbitrary pattern. The imaging
optics and pattern forming optics are pre-aligned to have a fixed
spatial relationship.
[0183] Referring to FIG. 38, a catheter design can include two
separate catheters--an imaging catheter and a laser thermotherapy
catheter that move relative to each other.
[0184] In the case of laser thermotherapy for lung cancer, a
Contrast Agent can be used where IndoCyanine Green is known to be
an effective Contrast Agent in LTT [or Laser Hyperthermia]
treatment of tumors in dogs.
[0185] Various implementations can be made based on the above
described features and techniques, including an optical tissue
imaging system having an imaging catheter and an optical engine; an
energy deposition system integrated with a catheter as described
above; an energy deposition system having a laser source that emits
light of at least two wavelengths; an integrated catheter having a
liquid filled balloon in contact with the airway tissue; and an
optical tissue imaging system having means of measuring temperature
induced optical properties change for the purpose of temperature
mapping.
[0186] The above systems can be used to perform following
operations: imaging airway to locate 3D map of tissue of interest
and depositing energy selectively to the tissue of interest by
controlling energy level, balloon liquid temperature, and laser
source wavelength based on the catheter position in a
pre-calculated pattern to match energy deposition with tissue
location. This method can also include calculating energy level,
balloon temperature and source spectrum, and applying small
fraction of the source energy at a level of below the threshold for
thermotherapy to obtain a pre-treatment temperature map. Next, the
source settings can be controlled to deliver a high dosage of the
laser light for the thermotherapy at a level above the
thermotherapy threshold until the actual measured temperature map
coincides with the pre-treatment temperature map. In some
implementations, a contrast agent, such as Indocyanine Green (ICG)
or other chemical or biological contrast agents, can be applied to
the region of therapeutic interest such as a malignant Solitary
Pulmonary Nodule, so as to preferentially absorb optical energy in
the region of therapeutic interest.
[0187] While this specification contains many specific
implementation details, these should not be construed as
limitations on the scope of the invention or of what may be
claimed, but rather as descriptions of features specific to
particular embodiments of the invention. Certain features that are
described in this specification in the context of separate
embodiments can also be implemented in combination in a single
embodiment. Conversely, various features that are described in the
context of a single embodiment can also be implemented in multiple
embodiments separately or in any suitable subcombination. Moreover,
although features may be described above as acting in certain
combinations and even initially claimed as such, one or more
features from a claimed combination can in some cases be excised
from the combination, and the claimed combination may be directed
to a subcombination or variation of a subcombination.
[0188] Only a few implementations are disclosed. Variations,
modifications and enhancements of the disclosed implementations and
other implementations can be made based on what is described and
illustrated in this document.
* * * * *