U.S. patent application number 10/765754 was filed with the patent office on 2004-09-23 for systems and methods of molecular spectroscopy to provide for the diagnosis of tissue.
This patent application is currently assigned to Massachusetts Institute of Technology. Invention is credited to Baraga, Joseph J., Feld, Michael S., Rava, Richard P..
Application Number | 20040186383 10/765754 |
Document ID | / |
Family ID | 24652118 |
Filed Date | 2004-09-23 |
United States Patent
Application |
20040186383 |
Kind Code |
A1 |
Rava, Richard P. ; et
al. |
September 23, 2004 |
Systems and methods of molecular spectroscopy to provide for the
diagnosis of tissue
Abstract
Systems and methods for spectroscopic diagnosis and treatment
are employed which utilize molecular spectroscopy to accurately
diagnose the condition of tissue. Infrared Raman spectroscopy and
infrared attenuated total reflectance measurements are performed
utilizing a laser radiation source and a fourier transform
spectrometer. Information acquired and analyzed in accordance with
the invention provides accurate details of biochemical composition
and pathologic condition.
Inventors: |
Rava, Richard P.; (Palo
Alto, CA) ; Baraga, Joseph J.; (Somerville, MA)
; Feld, Michael S.; (Waban, MA) |
Correspondence
Address: |
THOMAS O. HOOVER, ESQ.
BOWDITCH & DEWEY, LLP
161 Worcester Road
P.O. Box 9320
Framingham
MA
01701-9320
US
|
Assignee: |
Massachusetts Institute of
Technology
Cambridge
MA
|
Family ID: |
24652118 |
Appl. No.: |
10/765754 |
Filed: |
January 27, 2004 |
Related U.S. Patent Documents
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Application
Number |
Filing Date |
Patent Number |
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10765754 |
Jan 27, 2004 |
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08107854 |
Oct 22, 1993 |
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6690966 |
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08107854 |
Oct 22, 1993 |
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PCT/US92/00420 |
Jan 17, 1992 |
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10765754 |
Jan 27, 2004 |
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08288990 |
Aug 11, 1994 |
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6697665 |
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08288990 |
Aug 11, 1994 |
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07661077 |
Feb 26, 1991 |
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Current U.S.
Class: |
600/473 ;
600/475; 600/477; 600/478 |
Current CPC
Class: |
A61B 2562/0242 20130101;
A61B 5/7257 20130101; A61B 5/7203 20130101; A61B 5/0071 20130101;
A61B 5/0075 20130101; A61B 5/0086 20130101; G01N 2021/656 20130101;
G01N 21/65 20130101 |
Class at
Publication: |
600/473 ;
600/475; 600/477; 600/478 |
International
Class: |
A61B 006/00 |
Goverment Interests
[0003] Funding for research conducted in connection with the
subject matter of the present application was provided under NIH
Grant No. RR 02594.
Foreign Application Data
Date |
Code |
Application Number |
Mar 8, 1993 |
WO |
PCT/US92/00420 |
Claims
What is claimed:
1. A spectroscopic diagnostic system for measuring tissue
comprising: a laser emitting radiation in an infrared spectrum; a
fiber optic cable optically coupled to the laser that delivers the
infrared radiation to a distal end of the fiber optic cable onto
tissue and collects Raman shifted radiation from the tissue for
delivery to a proximal end of the cable; a spectral analyzer that
is optically coupled to the fiber optic cable to receive the
collected Raman shifted radiation from the fiber optic cable; and a
charge coupled device detector that is optically coupled to the
spectral analyzer and that detects radiation received from the
spectral analyzer.
2. The system of claim 1 further comprising a data processor that
removes background components from the detected light to provide
corrected Raman spectral data and analyzes the corrected Raman
spectral data to diagnose a condition of the portion of the
tissue.
3. The system of claim 2 wherein the detector detectes a plurality
of Raman shifted frequencies such that the data processor analyzes
the plurality of shifted frequencies to diagnose the tissue.
4. The system of claim 1 wherein the system collects light
returning from the tissue for a period of 5 minutes or less.
5. The system of claim 1 wherein the laser emits light having an
average incident power between 2 and 20 mW.
6. A method of spectroscopic diagnosis of tissue of a patient
comprising: irradiating a portion of tissue of a patient to be
diagnosed with laser radiation directed onto the tissue through a
fiber optic cable; detecting light emitted by the portion of tissue
in response to the radiation with a charge coupled device that is
optically coupled to a proximal end of the fiber optic cable, the
device collecting the light for a period of 5 minutes or less, the
light having a Raman shifted frequency component different from the
irradiating frequency; processing the detected light to provide
corrected Raman spectral data to diagnose a condition of the
portion of tissue.
7. The method of claim 6 wherein the detecting step further
comprises detecting a plurality of Raman shifted frequency
components and background light components and the analyzing step
further comprises analyzing the plurality of Raman shifted
frequency components to diagnose the tissue.
8. The method of claim 7 further comprises removing the background
light components from the detected light to leave substantially the
Raman shifted frequency light components.
9. The method of spectroscopic diagnosis of claim 6 further
comprising coupling the laser radiation from a laser radiation
source to a fiber optic cable to transmit the laser radiation onto
the portion of tissue.
10. A method of spectroscopic of arterial tissue comprising:
positioning a probe containing a light transmitting fiber optic
cable adjacent to a portion of tissue within an artery of a patient
to be diagnosed; directing excitation light onto a portion of
tissue, the light having a frequency within an infrared spectral
range; collecting light emitted by the portion of tissue in
response to the excitation with the probe for a period of 5 minutes
or less, the light having a Raman shifted frequency different from
the frequency; transmitting the collected light to a proximal end
of the probe; detecting the collected light with a charge coupled
device that is optically coupled to the proximal end of the probe;
removing background components from the detected light to provide
corrected Raman spectral data; and analyzing the corrected Raman
spectral data received at the proximal end to diagnose a condition
of the portion of the tissue.
11. The method of spectroscopic diagnosis of claim 10 wherein the
detecting step further comprises detecting a plurality of Raman
shifted frequencies and the analyzing step further comprises
analyzing the plurality of shifted frequencies to diagnose the
tissue.
12. The method of spectroscopic diagnosis of claim 10 wherein the
fiber optic cable receives light from the tissue and delivers the
received light to a spectrometer.
13. The method of claim 10 further comprising providing a laser
light source that emits light having an average incident power in a
range from 2 to 20 mW.
14. A spectroscopic diagnostic system for analyzing tissue of a
patient comprising: a laser system emitting laser radiation at an
excitation wavelength in a range between 750 nm and 1050 nm; a
fiber optic cable coupled to the laser to deliver the laser
radiation to a distal end of the fiber optic cable and onto tissue
and to collect Raman shifted radiation having a component
wavelength different from said excitation wavelength that is
emitted by the tissue for a period of 5 minutes or less for
delivery to a proximal end of the cable; a spectral analyzer that
is optically coupled to the fiber optic cable to receive the
collected Raman shifted radiation, the spectral analyzer comprising
a spectrometer that generates a spectrum of the collected Raman
shifted radiation and a charge coupled device that detects the
generated spectrum; and a data processor that processes the
detected Raman shifted radiation to provide corrected Raman
spectral data.
15. The system of claim 14 wherein the laser emits an average
incident power in a range from 2 to 20 mW.
16. The system of claim 14 further comprising a plurality of
filters positioned to filter the collected light.
17. The system of claim 14 wherein the system collects returning
light for 8 seconds or less to generate spectral data for a region
of interest.
18. The system of claim 14 wherein the excitation wavelength is in
a range of 800 to 900 nm.
Description
CROSS REFERENCES TO RELATED APPLICATIONS
[0001] This application is a continuation application of U.S. Ser.
No. 08/107,854, filed Oct. 22, 1993, which is a U.S. National Entry
of PCT/US92/00420, filed Jan. 17, 1992 and is a
continuation-in-part of 08/288,990, filed Aug. 11, 1994, which is a
File Wrapper Continuation of U.S. Ser. No. 07/661,077, filed on
Feb. 26, 1991, now abandoned. The entire contents of the above
applications are incorporated herein by reference in entirety.
[0002] This application is also related to U.S. Ser. No.
07/661,072, filed on Feb. 26, 1991, now U.S. Pat. No. 5,280,788.
The entire contents of the above applications are incorporated
herein by reference in entirety.
BACKGROUND OF THE INVENTION
[0004] In the United States heart attacks, almost entirely
attributable to coronary atherosclerosis, account for 20-25% of all
deaths. Several medical and surgical therapies are available for
treatment of atherosclerosis; however, at present no in situ
methods exist to provide information in advance as to which lesions
will progress despite a particular medical therapy.
[0005] Objective clinical assessments of atherosclerotic vessels
are at present furnished almost exclusively by angiography, which
provides anatomical information regarding plaque size and shape as
well the degree of vessel stenosis. The decision of whether an
interventional procedure is necessary and the choice of appropriate
treatment modality is usually based on this information. However,
the histological and biochemical composition of atherosclerotic
plaques vary considerably, depending on the stage of the plaque and
perhaps also reflecting the presence of multiple etiologies. This
variation may influence both the prognosis of a given lesion as
well as the success of a given treatment. Such data, if available,
might significantly assist in the proper clinical management of
atherosclerotic plaques, as well as in the development of a basic
understanding of the pathogenesis of atherosclerosis.
[0006] At present biochemical and histological data regarding
plaque composition can only be obtained either after treatment, by
analyzing removed material or at autopsy. Plaque biopsy is
contraindicated due to the attendant risks involved in removing
sufficient arterial tissue for laboratory analysis. Recognizing
this limitation, a number of researchers have investigated optical
spectroscopic methods as a means of assessing plaque deposits. Such
"optical biopsies" are non-destructive, as they do not require
removal of tissue, and can be performed rapidly with optical fibers
and arterial catheters. With these methods, the clinician can
obtain, with little additional risk to the patient, information
that is necessary to predict which lesions may progress and to
select the best treatment for a given lesion.
[0007] Among optical methods, most attention has centered on
ultraviolet and/or visible fluorescence. Fluorescence spectroscopy
has been utilized to diagnose disease in a number of human tissues,
including arterial wall. In arterial wall, fluorescence of the
tissue has provided for the characterization of normal and
atherosclerotic artery. However the information provided is limited
by the broad line width of fluorescence emission signals.
Furthermore, for the most part, fluorescence based methods provide
information about the electronic structure of the constituent
molecules of the sample. There is a need for non-destructive real
time biopsy methods which provide more complete and accurate
biochemical and molecular diagnostic information. This is true for
atherosclerosis as well as other diseases which affect the other
organs of the body.
SUMMARY OF THE INVENTION
[0008] The present invention relates to vibrational spectroscopic
methods using Fourier transform infrared (FT-IR) attenuated total
reflectance (ATR) and near-infrared (IR) FT-Raman spectroscopy.
These methods provide extensive molecular level information about
the pathogenesis of disease. Both of these vibrational techniques
are readily carried out remotely using fiber optic probes. In
particular, a preferred embodiment utilizes FT-Raman spectra of
human artery for distinguishing normal and atherosclerotic tissue.
Near IR FT-Raman spectroscopy can provide information about the
tissue state which is unavailable from fluorescence methods. In
situ vibrational spectroscopic techniques allow probing of the
molecular level changes taking place during disease progression.
The information provided is used to guide the choice of the correct
treatment modality.
[0009] These methods include the steps of irradiating the tissue to
be diagnosed with radiation in the infrared range of the
electromagnetic spectrum, detecting light emitted by the tissue at
the same frequency, or alternatively, within a range of frequencies
on one or both sides of the irradiating light, and analyzing the
detected light to diagnose its condition. Both the Raman and ATR
methods are based on the acquisition of information about molecular
vibrations which occur in the range of wavelengths between 3 and
300 microns. Note that with respect to the use of Raman shifted
light, excitation wavelengths in the ultraviolet, visible and
infrared ranges can all produce diagnostically useful information.
Near IR FT-Raman spectroscopy is ideally suited to the study of
human tissue.
[0010] Raman spectroscopy is an important method in the study of
biological samples, in general because of the ability to this
method to obtain vibrational spectroscopic information from any
sample state (gas, liquid or solid) and the weak interference from
the water Raman signal in the "fingerprint" spectral region. The
FT-spectrometer furnishes high throughput and wavelength accuracy
which might be needed to obtain signals from tissue and measure
small frequency shifts that are taking place. Finally, standard
quartz optical fibers can be used to excite and collect signals
remotely.
[0011] Near IR FT-Raman spectroscopy provides the capability to
probe biological substituents many hundred microns below the tissue
surface. In particular, for atherosclerotic tissue, calcified
deposits below the tissue surface are easily discerned. Thus, it
becomes possible to detect pathologic conditions which would not be
apparent using angioscopic methods, as well as to study the
detailed molecular basis of the pathology.
[0012] In contrast with electronic techniques, the bands in a
vibrational spectrum are relatively narrow and easy to resolve.
Vibrational bands are readily assigned to individual molecular
groups.
[0013] The ATR technique offers several features especially suited
to sampling of human tissue in vivo. Being a surface technique, the
ATR method can non-destructively probe internal human tissue either
by direct contract in a hollow organ (e.g. artery), or by insertion
of a needle probe. In the mid-IR region, strong water absorption
dominates the spectra of highly hydrated samples such as arterial
tissue, obscuring the absorption from other tissue components (see
FIG. 8). Accurate subtraction of the strong water absorption from
FT-IR ATR spectra is relatively easy and very reliable with the
high dynamic range, linearity, stability, and wavelength precision
of available FT spectrometers. Furthermore, high quality mid-IR
spectra of aqueous protein solutions can be collected with fiber
optic ATR probes. Such probes are easily adaptable to existing
catheters for remote, non-destructive measurements in vivo. The
mid-IR ATR technique allows clinicians to gather precise
histological and biochemical data from a variety of tissues during
standard catheterization procedures with minimal additional
risk.
[0014] The present methods relate to infrared methods of
spectroscopy of various types of tissue and disease including
cancerous and pre-cancerous tissue, non-malignant tumors or lesions
and atherosclerotic human artery. Examples of measurements on human
artery generally illustrate the utility of these spectroscopic
techniques for clinical pathology. Results obtained demonstrate
that high quality, reproducible FT-IR ATR spectra of human artery
can be obtained with relative ease and speed. In addition,
molecular level details can be reliably deduced from the spectra,
and this information can be used to determine the biochemical
composition of various tissues including the concentration of
molecular constituents that have been precisely correlated with
disease states to provide accurate diagnosis.
[0015] Another preferred embodiment of the present invention uses
two or more diagnostic procedures either simultaneously or
sequentially collected to provide for a more complete diagnosis.
These methods can include the use of fluorescence of endogenous
tissue, Raman shifted measurements and/or ATR measurements.
[0016] Yet another preferred embodiment of the present invention
features a single stage spectrograph and charge-coupled device
(CCD) detector to collect NIR Raman spectra of the human artery.
One particular embodiment employs laser light in the 810 nm range
to illuminate the tissue and thereby provide Raman spectra having
frequency components in a range suitable for detection by the CCD.
Other wavelengths can be employed to optimize the diagnostic
information depending upon the particular type of tissue and the
type and stage of disease or abnormality. Raman spectra can be
collected by the CCD at two slightly different illumination
frequencies and are subtracted from one another to remove broadband
fluorescence light components and thereby produce a high quality
Raman spectrum. The high sensitivity of the CCD detector combined
with the spectra subtraction technique allow high quality Raman
spectra to be produced in less than 1 second with laser
illumination intensity similar to that for the FT-Raman system also
described herein.
BRIEF DESCRIPTION OF THE DRAWINGS
[0017] FIGS. 1A-1C are schematic illustrations of preferred systems
for providing the spectroscopic measurements of the invention.
[0018] FIG. 2 graphically illustrates FT-Raman spectra of human
aorta: a) normal artery; b) atheromatous plaque; c) FT-Raman
spectrum solid cholesterol (Sigma).
[0019] FIG. 3 graphically illustrates FT-Raman spectra of normal
human aorta: a) irradiated from intimal side (spectrum multiplied
by 3); and b) irradiated from adventitial side (primary adipose
tissue). c) NIR FT-Raman spectrum of triglyceride, triolein.
[0020] FIG. 4 graphically illustrates FT-Raman spectra from human
aorta: a) fibrous plaque; and b) atheromatous plaque. c) FT-Raman
spectrum of cholesterol monohydrate powder.
[0021] FIG. 5 graphically illustrates FT-Raman spectra of calcified
human aorta: a) calcified with fibrous cap; b) excised
calcification from a different plaque; c) spectra of the same
tissue as in a) taken from adventitial side.
[0022] FIG. 6 graphically illustrates FT-Raman spectra of calcified
human aorta: a) calcified plaque with a fibrous cap (spectrum
multiplied by 8); and b) exposed calcification.
[0023] FIG. 7 graphically illustrates the measured NIR Raman
intensity of the 960 cm.sup.-1 band (A(960 cm.sup.-1) indicates the
area of this band) in a calcified deposit as a function of depth
below the irradiated surface. The dashed curve corresponds to the
fit of an exponential function to the data with an exponent of 2.94
mm.sup.-1.
[0024] FIG. 8 graphically illustrates FT-IR ATR spectra (4000-700
cm.sup.-1) of (a) normal aorta, intimal surface; and (b) buffered
saline (0.14M NaCl, pH 7.4).
[0025] FIG. 9 graphically illustrates FT-IR ATR spectra (1800-800
cm.sup.-1) after water subtraction: 9a) Normal aorta, intimal
surface; (b) Sub-adventitial fat; (c) Saline rinsed from the
intimal surface of normal aorta.
[0026] FIG. 10 graphically illustrates FT-IR ATR spectra (1800-800
cm.sup.-1): (a) Two consecutive water-subtracted spectra of normal
aorta, intimal surface, collected immediately after placement on
ATR element (solid line) and 10 minutes later (dashed line); (b)
Same two spectra as in (a) after lipid subtraction, scaled to have
equal maxima.
[0027] FIG. 11 graphically illustrates FT-IR ATR spectra (1800-800
cm.sup.-1), water- and lipid-subtracted: (a) Normal aorta, media
layer; (b) Atherosclerotic plaque, intimal surface; (c) Athermanous
plaque with intact fibrous cap, intimal surface.
[0028] FIG. 12 graphically illustrates FT-IR ATR spectra (1800-800
cm.sup.-1): (a) Necrotic core of atheromatous plaque, water- and
lipid-subtracted; (b) Dry film of cholesterol.
[0029] FIG. 13 graphically illustrates scatter plot for all samples
of the area, A(1050), of the 1050 cm.sup.-1 cholesterol band
(integrated from 1075 to 1000 cm.sup.-1) ratioed to the area,
A(1550) of the 1548 cm.sup.-1 protein amide II band (integrated
from 1593 to 1485 cm.sup.-1). The intensities were calculated from
the water- and lipid-subtracted spectra. NORMAL denotes normal
aorta specimens, intimal side, FIBROUS includes atherosclerotic and
atheromatous plaques with intact fibrous caps, and NECROTIC
includes exposed necrotic atheroma cores and necrotic material
isolated from atheromatous plaques.
[0030] FIG. 14 graphically illustrates FT-IR ATR spectra (1800-800
cm.sup.-1): (a) Second derivative spectrum of normal aorta intima
(FIG. 8a); (b) Water-subtracted spectrum of same normal aorta
intima specimen (same as FIG. 9a).
[0031] FIG. 15 graphically illustrates a scatter diagram for all
the specimens of the area, A(1050) of the 1050 cm.sup.-1
cholesterol band plotted versus the area, A(1382), of the 1382
cm.sup.-1 cholesterol band. Both cholesterol bands have been
normalized to the area, A(1050), of the protein amide II band. All
band intensities were calculated from the water- and
lipid-subtracted spectra. Tissue categories are the same as in FIG.
13. The solid line represents a linear least squares fit to the
data.
[0032] FIGS. 16A and 16B are additional preferred embodiments of
ATR probes adapted to make the diagnostic measurements of the
present invention.
[0033] FIG. 17 is a schematic diagram of the system of FIG. 1A
modified to use the spectrograph/CCD Raman detector of the present
invention.
[0034] FIG. 18 is a schematic diagram of a preferred system for
implementing the spectrograph/CCD Raman detector of the present
invention.
[0035] FIG. 19 graphically illustrates spectrograph/CCD-Raman
spectra of normal human aorta: A) Raman plus fluorescence spectrum
produced by illuminating the tissue sample with 810 nm laser light;
B) Raman difference spectrum produced by subtracting spectra
produced by illuminating the tissue sample with 810 and 812 nm
laser light; C) resulting Raman spectrum produced by integrating
the Raman difference spectrum of B).
[0036] FIG. 20 graphically illustrates spectrograph/CCD-Raman
spectra of an atherosclerotic plaque with a calcified deposit
exposed at the surface: A) Raman plus fluorescence spectrum
produced by illuminating the tissue sample with 810 nm laser light;
B) Raman difference spectrum produced by subtracting spectra
produced by illuminating the tissue sample with 810 and 812 nm
laser light; C) resulting Raman spectrum produced by integrating
the Raman difference spectrum of B).
[0037] FIG. 21 graphically illustrates a spectrograph/CCD-Raman
spectrum of adventitial adipose tissue.
DETAILED DESCRIPTION OF THE INVENTION
[0038] The spectroscopic methods of the present invention can be
performed on a system such as that for laser treatment of
atherosclerosis which is illustrated in FIG. 1A. FIG. 1A includes
separate block diagrams for the system of the invention which
utilizes laser light for spectroscopic diagnosis as well as for
treatment and/or removal of tissue. The ablation laser 225, pulse
stretcher 229 and the pulse filler/multiplexer 231, 233 produce an
output laser ablation pulse of sufficient energy and intensity to
remove tissue and sufficient pulse duration to propagate through a
fiber optic laser catheter delivery system without damaging the
fibers. These systems and methods are more fully described in
co-pending application U.S. Ser. No. 07/644,202 filed on Jan. 22,
1991, now U.S. Pat. No. 5,312,396, which is incorporated herein by
reference.
[0039] To remove plaque, a device 219 is used to contact the tissue
such as multiple-fiber laser catheter 10 of FIG. 1B with an optical
shield. The catheter 10 is inserted into the artery and the distal
end of the catheter is brought into contact with the lesion. Next,
a determination is made as to the type of tissue at which each
optical fiber 20a-c is aimed. Only fibers aimed at diseased tissue
are activated. Thus, selective tissue removal is obtained.
Furthermore, this technique is also applicable for guiding surgical
procedures in other organs and tissue types such as the colon and
bladder.
[0040] The present invention relates to systems and methods of
performing spectral diagnostics to diagnosis the tissue in front of
each fiber. A preferred embodiment a laser light source 207 that is
coupled to the fibers. The diagnostic light is sent to the fiber of
choice by the optical fiber selector 217.
[0041] The diagnostic light exits the selected optical fiber and
falls on the tissue. The tissue absorbs the light and a fraction of
the absorbed light is re-emitted, by Rayleigh fluorescence, Raman
or other elastic or inelastic light scattering processes. This
light is returned to the optical fibers and exits from selector
217, and is detected by a photodiode, photomultiplier or other
detector 203. Returning light could use different optical fibers
than those employed for illumination. Diagnostic subsystem produces
the entire spectral signal which is coupled to computer 215.
[0042] The computer stores the information in a memory as a
spectrum, which is a graph of light intensity vs. wavelength. This
can be displayed immediately on the video display 82 or compared to
an existing spectrum stored in the computer and the difference
displayed on the spectral display 86. Temporal display 88 can
display corrections made for the wavelength dependent sensitivities
of the source. Information from either the temporal or spectral
display can be stored in the computer 80. The comparative data is
shown on numerical display 84 to provide a quantitative measure of
the health of the tissue observed.
[0043] With a multichannel detector and a computer, or with
appropriate multiple filters and detectors, it is possible to
gather this information in a fraction of a second. Thus, a spectral
or numerical display is provided which indicates the type of tissue
at which the fiber of interest is aimed. If the tissue is plaque,
and is to be removed, then fiber selector 217 will align this fiber
with the output beam of the high power laser 225. Then, the high
power laser 225 is turned on and an appropriate power level is
selected for a predetermined amount of time to remove a certain
amount of diseased tissue. The beam of laser 225 is transmitted to
pulse stretcher 229 through coupling optics 227 and pulse
filler/multiplexer 231, 233 to properly adjust the beam
fluence.
[0044] The procedure is repeated for different fibers using
switcher 235. Where diseased tissue is detected, it is quickly
removed. The laser catheter 10 nibbles away at the plaque, leaving
the healthy artery wall intact.
[0045] It the artery 30 makes a bend 31 as shown by FIG. 1B, the
laser catheter 10 will tend to make contact with artery wall 32 at
the outside wall of the bend. To prevent the catheter from
contacting the artery wall, the optical fiber 20c is not fired. The
lesion is removed asymmetrically. This allows the laser catheter 10
to follow the lumen 39, 39a around the bend. Thus, the artery wall
32 is not irradiated and is not perforated. Additional details of
this fiber optic catheter 10 are disclosed in U.S. Pat. No.
4,913,142, the contents of which are incorporated herein by
reference.
[0046] In both Raman and ATR methods, information is contained in
the spectral lines which are observed in their intensities, and
also their linewidths and center frequencies (and how they change
in different environments). Further, Raman and infrared ATR have
different "selection rules". Some vibrations seen in infrared ATR
won't show up in Raman, and vice versa. In other cases the same
vibration can be detected by both techniques, but with different
relative intensities (e.g. a strong Raman line will be weak in
ATR). So in this sense the two techniques provide complementary
information and combining the two techniques (or using either or
both with laser induced fluorescence) is valuable in diagnosing
pathology.
[0047] The methods can utilize Fourier transform detection to
observe the radiation thereby providing improved signal/noise
ratios. Other techniques (e.g. diode array detection and CCD
detection) can also be used.
[0048] As described in more detail below contributions from major
tissue constituents can be "subtracted out" to reveal information
about molecules which are present in small concentrations. For
example, in ATR water contributions are removed before the "dry"
tissue constituents could be studied. Also, derivative spectroscopy
is used to eliminate background signals and low frequency filters.
Note that these techniques deconvolute the observed spectra into
its individual constituents, an essential step for optimal
extraction of diagnostic information.
[0049] While Raman can sample deeply into tissue, ATR samples only
a very thin layer (a few microns). Thus, ATR is "naturally" suited
to probe surface disease, such as the superficial cancers of the
bladder and GI tract, whereas Raman is well suited to providing
information about conditions deep inside tissue (such as breast
cancer or stones). This is important for 3D imaging. Furthermore,
the ATR tissue sampling depth can be controlled by properly
matching the probe surface material to the tissue type.
[0050] Generally, the ATR signal is very sensitive to the surface
of the waveguide or probe. For example, if the probe surface has an
affinity for lipids in the tissue, lipids can migrate to the probe
surface, creating a thin lipid layer and producing a large signal.
This can be a problem (it can give misleading information if not
properly recognized and guarded against). Conversely, it can be
used to advantage: Probes with special surfaces can be developed to
prevent this effect or to promote it, in order to search for
particular substances in the tissue.
[0051] In a preferred method one can adjust depth probed by ATR by
varying refractive index of ATR probe. Alternatively as discussed
below one can use a "waveguide needle" to get subsurface
information.
[0052] Raman diagnostic methods permit adjustment of Raman depth by
varying the wavelength. Penetration depth is wavelength dependent,
and can be varied by choosing different excitation wavelengths
between about .lambda.=700 nm and 2 .mu.m. Another potentially
important way of adjusting Raman depth is to very the collection
angle. In the near IR, incident (exciting) light is strongly
scattered out of the forward direction into larger angles, so Raman
signals sampled at smaller angles come from tissue closer to the
surface. Therefore, the Raman sampling depth can be controlled to a
large extent by probe design.
[0053] Depth information is important if one desires to provide
imaging by creating 3D images of small tumors in the brain or
breast. Differential techniques based on the ideas of the preceding
paragraph might allow accurate localization of such tumors in three
dimensions. Near-IR Raman can be combined with a sound wave
technique (time of flight or standing waves set up in the
tissue)--the sound wave would modulate the Raman signal emanating
from a point in the tissue when it passes that point, and the
modulated signal could be used to establish the depth of the tissue
producing the signal.
[0054] A system employed for the collection of Raman spectral data
from excised tissue samples is illustrated in FIG. 1C. FT-Raman
spectra were measured from 0-4000 cm.sup.-1 below the laser
excitation frequency with a FT-IR interferometer 40 equipped with a
FT-Raman accessory. The accessory employed at 180 back scattering
geometry and a cooled (77K) InGaAs detector 42.
[0055] A 1064 nm CW Nd:YAG laser 44 can be used for irradiating a
sample 46: utilizing 500 mW to 1 W laser power in a 1.0 to 2.5 mm
spot 48 at the sample 46 to collect Raman data. Alternatively, a
pulsed laser source can also be employed. Laser 44 generated a beam
47 that is directed through plasma filter 48, mirrors 50, 52,
focusing lens 54 and mirror or prism 56 before irradiating the
sample 46. The radiation received by sample 46 undergoes various
mechanisms of absorption, reflection and scattering including Raman
scattering. Some of the light emitted by the tissue is directed
toward lens 60 and then through one or more Rayleigh filters 62.
The collecting lens 60 collects this backscattered light 64 and
collimates it. The Rayleigh filters 62 removes the elastically
scattered light and transmits the inelastically scattered,
frequency shifted (Raman) light. The Raman scattered light enters
the interferometer 40. No visible sample degradation was observed
under these conditions.
[0056] Excised human aorta was chosen of atherosclerotic artery
tissue. Samples were obtained at the time of post mortem
examination, rinsed with isotonic saline solution (buffered at PH
7.4), snap-frozen in liquid nitrogen, and stored at -85C until use.
Prior to spectroscopic study, samples were passively warmed to room
temperature and were kept moist with the isotonic saline. Normal
and atherosclerotic areas of tissue were identified by gross
inspection, separated, and sliced into roughly 8.times.8 mm
pieces.
[0057] The tissue samples were placed in a suprasil quartz cuvette
with a small amount of isotonic saline to keep the tissue moist,
with one surface in contact with the irradiated by the laser 44.
The spectra shown in FIGS. 2 through 6 were collected with 512
scans at 8 cm.sup.-1 resolution (approximately 35 minutes total
collection time).
[0058] Human aorta is composed of three distinct layers: intima,
media, and adventitia. The intima, normally less than 300 .mu.m
thick, is the innermost layer and provides a non-thrombogenic
surface for blood flow. It is mainly composed of collagen fibers
and ground substance. The medial layer, typically about 500 .mu.m
thick, is quite elastic and serves to smooth the pulsatile blood
flow from the heart. The structural protein elastin is the major
component of aortic media, with some smooth muscle cells present as
well. The outermost adventitial layer serves as a connective tissue
network which loosely anchors the vessel in place, and is mainly
made up of lipids, lipoproteins and collagen. During the
atherosclerotic process, the intima thickens due to collagen
proliferation, fatty necrotic deposits accumulate under and within
the collagenous intima, and eventually, calcium builds up, leading
to calcium hydroxyapatite deposits in the artery wall.
[0059] FIG. 2, curve a, shows the FT-Raman spectrum of a full
thickness section of aorta grossly identified as normal. Laser
irradiation was on the intimal side. The dominant bands appear at
1669 cm.sup.-1 and 1452 cm.sup.-1 and can be assigned to an amide I
backbone and C--H in-plane bending vibration from proteins,
respectively. Weaker bands at 1331 and 1225 cm.sup.-1 are assigned
to C--H wagging and amide III vibrations from proteins,
respectively. The frequencies of amide I and III are consistent
with those observed for disordered proteins.
[0060] Another example of a typical NIR FT Raman spectrum from
normal aorta is shown in FIG. 3. When irradiated from the intimal
side, FIG. 3, curve a, the major vibrational bands observed in
normal aorta are all attributable to protein vibrations: the band
at 1658 cm.sup.-1 is assigned to the amide I vibration of the
polypeptide chain, the 1453 cm.sup.-1 band to a C--H bending mode
of proteins, and the 1252 cm.sup.-1 band to the amide III
vibration. The spectrum of normal aorta is at lease 25% weaker than
any of the pathologic samples. The peak frequency of the C--H
bending band, which averaged for all the normal specimens is
1451.+-.1 cm.sup.-1, is specific to the protein C--H bending mode
(See below). The weak band near 1335 cm.sup.-1, which appears as a
shoulder in many of the normal specimens, appears to be specific to
elastin, and the weak band at 1004 cm.sup.-1 is likely due to
phenylalanine residues. In general, the relative intensities of the
bands in the region between 1250 and 1340 cm.sup.-1 appears very
much like that observed in the FT Raman spectrum of elastin. This
observation is consistent with the thin collagenous intima in
normal aorta, the elastic nature of the media of aorta, and the
deep penetration depth of 1064 nm radiation. Band assignments for
all tissue types presented here are listed in Table 2.
[0061] FIG. 3, curve b, displays the NIR FT Raman spectrum of the
adventitial side of normal aorta. In this case, the irradiated
adventitial surface consisted of several millimeters of visible
adipose tissue. In contrast with the spectrum collected from the
intimal side, the bands observed in this adipose material appear to
be mainly due to lipid, and in particular triglyceride, with almost
no contribution from protein. This is not unexpected, as the
triglyceride content of adipose tissue is on the order of 60%. The
sharp band at 1655 cm.sup.-1 is due to stretching of C.dbd.C groups
in unsaturated fatty acid chains. This band is distinguished from
amide I by its peak frequency and its width, which in this case is
17 cm.sup.-1 FWHM. Amide I, in contrast, is roughly 60 cm.sup.-1
wide. The dominant C--H bending band is shifted to 1440 cm.sup.-1,
characteristic of lipids. This band is about 3 times more intense
in adipose tissue than in normal intima, probably a result of the
greater number of C--H groups per unit volume in triglycerides. The
bands as 1301/1267 cm.sup.-1 and 1080 cm.sup.-1 are assigned to
C--H bending and C--C stretching vibrations of fatty acids,
respectively.
[0062] The 1746 cm.sup.-1 band, assigned to the C.dbd.O stretching
mode of the triglyceride ester linkages, indicates that most of the
lipid observed in the adventitial adipose tissue is of the
triglyceride form. Specifically, the integrated intensity of this
band relative to the C--H bending band at 1440 cm.sup.-1 is equal
to 0.103, while this same ratio for triolein is 0.136, which
indicates that roughly 75% of the C--H band is due to triglyceride.
The NIR FT Raman spectrum of triolein (a triglyceride containing
fatty acid chains of 18 carbons and a single double bond) is shown
for comparison in FIG. 3, curve c. Additional molecular information
regarding the state of the fatty acid chains is readily deduced
from the adventitial adipose spectrum. For example, the relative
intensity of the C.dbd.C band at 1655 cm.sup.-1 indicates that
there are on average roughly 0.7 unsaturated double bonds per fatty
acid chain, assuming 16-18 carbon fatty acids. In addition, the
frequencies and structures of the C--H bending and C--C stretching
bands suggest that most of the fatty acid chains are in the gauche
conformation. The sharp 1129 cm.sup.-1 band, characteristic of
all-trans chains, is not observed in the spectrum.
[0063] The FT-Raman spectrum obtained from a diseased artery, an
atheromatous plaque, with a thick fibrous connective tissue cap and
an underlying necrotic core is shown in FIG. 2, curve b. The
necrotic core of an atheromatous plaque contains cellular debris as
well as large accumulations of oxidized lipids and cholesterol. The
band in the amide I region, peaking at 1665 cm.sup.-1, is
distinctly narrow in this spectrum compared to normal aorta. In
addition, the in-plane C--H bend, at 1444 cm.sup.-1, is relatively
more intense and has a distinct shoulder at higher frequency. The
two weaker bands at 1307 and 1267 cm.sup.-1 are shifted in
frequency from those found in the spectrum of normal aorta. The
band frequencies and shapes in the FT-Raman spectrum of
cholesterol, shown in FIG. 2, curve c, coincide with some of those
observed in the atheromatous plaque, consistent with the expected
composition of the necrotic core.
[0064] The NIR FT Raman spectra of other fibrous plaque specimens
exhibit a range of features as shown in FIGS. 4 and 5. FIG. 4,
curve a, shows a representative spectrum from one of the types of
fibrous plaques. These fibrous plaque spectra are quite similar in
both relative and absolute band intensities to the spectra of
normal aorta. The most significant differences are that the C--H
bending band, peaking near 1448 cm.sup.-1 on average, is shifted by
2 cm.sup.-1 to a slightly lower frequency. This may be the result
of a small increase in the lipid content of these plaques relative
to normal aorta. In addition, the band near 1340 cm.sup.-1,
attributed to elastin in normal aorta, is decreased relative to
amide III at 1265 cm.sup.-1. The putative explanation is that the
collagenous intima is thickened in these specimens, so that the
spectral contribution from the elastic media is reduced relative to
that of normal aorta.
[0065] The NIR FT Raman spectra of other fibrous plaque specimens
appeared similar to atheromatous plaques' spectra (FIG. 2, curve
b). These are substantially different than either normal aorta, or
adipose tissue. In these cases, the intense C--H bending band
occurs at 1440 cm.sup.-1, characteristic of lipid material. This
band is roughly twice as intense as the C--H bending band in normal
aorta. The complete absence of a band at 1746 cm.sup.-1 indicates
that this lipid is not triglyceride. In fact, this lipid appears to
be predominantly cholesterols, as identified by the sharp,
characteristic band at 700 cm.sup.-1 and comparison to the
cholesterol spectrum shown in FIG. 4, curve c. Again, this is not
surprising, since cholesterols accumulate in high concentrations in
atherosclerotic lesions. Several of the bands between 1000 and 500
cm.sup.-1 are assignable to vibrational modes of the sterol rings.
These include the bands at 959, 882, 844, 805, 700, 605, and 546
cm.sup.-1. In addition, the 1666 cm.sup.-1 band is attributed in
part to the C.dbd.C stretching vibration of the steroid
nucleus.
[0066] The presence of fatty acid chains in the atheromatous plaque
spectra is evidence by bands at 1300/1262 cm.sup.-1 and 1130/1088
cm.sup.-1, due to C--H bending and C--C stretching vibrations,
respectively. These bands may contain contributions from
cholesterol as well. The relative intensities of the fatty acid
band at 1300 cm.sup.-1 and the sterol ring bands suggest a mixture
of free cholesterol and cholesterol-fatty acid esters. Moreover,
the relative intensities of the 1130 cm.sup.-1 C--C stretching and
the 700 cm.sup.-1 sterol bands indicate that most of the fatty acid
chains are in the gauche conformation, consistent with the
predominance of unsaturated fatty acid chains in the cholesterol
esters in these plaques. It is particularly noteworthy in the
atheromatous plaques that the cholesterol deposits are detected
from material below a thick fibrous cap indicating the ability of
the NIR FT Raman method to probe materials several hundred microns
below the tissue surface.
[0067] In addition to the cholesterol and cholesterol ester bands,
the NIR FT Raman spectra of some of the fibrous plaques contained
two unique bands, at 1519 and 1157 cm.sup.-1. The intensities of
these bands are highly correlated, which suggests that they are due
to a single component. These bands, which have been previously
observed in visibly-excited Raman spectra of atherosclerotic
plaques, are assigned to carotenoids. The amount of carotenoid in
these plaques is probably much smaller than the amounts of
cholesterols or proteins, but may be strongly pre-resonance
enhanced (14). The carotenoid bands are observed only in this
subset of fibrous plaques.
[0068] In an advanced plaque, calcium may begin to accumulate
leading to the formation of calcium hydroxyapatite crystals in the
tissue as shown by the spectra of FIGS. 5 and 6. The FT-Raman
spectrum of a calcified plaque with a thick (several hundred
microns) fibrous connective tissue cap overlying a calcified
deposit is shown in FIG. 5, curve a. The spectrum clearly indicates
bands due to the protein in the fibrous cap, amide I and III at
1665 and 1255 cm.sup.-1, respectively. However, additional bands
are observed between 1250 and 1350 cm.sup.-1 and around 1100
cm.sup.-1, as well as a strikingly sharp band at 961 cm.sup.1. The
latter is readily assigned to the symmetric phosphate stretching
vibration associated with phosphate groups in the calcium
hydroxyapatite deposits, while the band around 1100 cm.sup.-1 is an
asymmetric phosphate stretch. These assignments are confirmed by
excising the solid "rock" from a different calcified plaque, and
obtaining its spectrum as shown in FIG. 5, curve b. A strong Raman
signal from the phosphate stretching vibration in solid
calcifications in advanced atherosclerotic plaques can also be
observed utilizing standard visible Raman instrumentation. The
ability to detect the calcifications several hundred microns below
the tissue surface when using near IR FT-Raman spectroscopy,
however, is a diagnostic measurement which can be utilized to guide
treatment.
[0069] A measurement was attempted to determine if the
calcification might be detected when the tissue was irradiated from
the adventitial side. The resulting FT-Raman spectrum is shown in
FIG. 5, curve c. No evidence of the strong phosphate vibration is
apparent. In contrast, sharp vibrational bands at 1745, 1656, 1444,
1303, 1267 and 1082 cm.sup.-1 are observed which are mainly
associated with the lipid material that constitutes the majority of
the adventitia.
[0070] The NIR FT Raman spectrum of calcified plaque, containing a
subsurface calcified deposit and an overlying soft fibrous cap,
exhibits an intense, sharp, new band at 960 cm.sup.-1 (FIG. 6,
curve a). This band, specific to calcified tissue, is assigned to
the symmetric stretching vibration of phosphate groups (15), which
are present in high concentrations in the solid calcium salts. The
weaker phosphate antisymmetric stretch is also present at 1072
cm.sup.-1. A symmetric stretching vibration of carbonate groups may
also contribute to this latter band. The phosphate vibrations are
easily observed from subsurface deposits in the calcified plaques:
the 960 cm.sup.-1 band can be observed from deposits up to 1.5 mm
beneath a soft tissue cap with the current signal-to-noise level
(See below). The calcified plaque also displays protein vibrations
from the fibrous tissue cap. These include amide I at 1664
cm.sup.-1 and amide III near 1257 cm.sup.-1. The C--H bending band
at 1447 cm.sup.-1 suggests a mixture of protein and lipid, and the
weak band at 699 cm.sup.-1 is likely due to cholesterol that is
either in the fibrous cap, the calcified deposit, or both.
[0071] The NIR FT Raman spectra of exposed calcifications (FIG. 6,
curve b) display a range of features. In all cases, the major bands
are due to the calcium salts. These include the 960 cm.sup.-1
phosphate and 1072 cm.sup.-1 phosphate/carbonate bands as well as
the band at 587 cm.sup.-1, which is assigned to another phosphate
vibrational mode. On the other hand, several differences are
apparent in the weaker bands, which are presumably due to soft
tissue components which are embedded in the calcification. In some
cases (not shown), the C--H bending band occurs at 1450 cm.sup.-1,
and the band at 1663 cm.sup.-1 is similar in shape to amide I for
some of the calcifications, indicative of protein vibrational
modes. In other calcified plaques, such as that in FIG. 5, curve b,
the C--H bending band occurs at 1440 cm.sup.-1, and the 1667
cm.sup.-1 band, which is much sharper, is more like due to C.dbd.C
stretching vibrations. In this plaque, the bands are due to lipid,
in particular cholesterols, as evidenced by the 700 cm.sup.-1 and
1300 cm.sup.-1 bands.
[0072] In our histological examinations of aorta, two distinct
types of exposed calcifications have been noted. In one type, the
fibrous tissue cap is calcified. In the other, the necrotic core of
an atheromatous plaque is calcified, and the calcified deposit is
exposed by ulceration of the soft tissue fibrous cap. A positive
explanation for the two spectral types of exposed calcifications is
that the specimens which exhibit protein bands are of the former
histologic type, while the specimens which exhibit lipid bands are
of the latter type.
[0073] The present methods provide an IR FT-Raman technique for
differentiating various stages of atherosclerosis in human aorta.
They demonstrate that molecular level information is available
using these methods. This information is useful for following the
pathogenesis of the disease and in guiding the treatment of
different lesions. The near IR FT-Raman method, with its relatively
deep penetration depth, is able to obtain spectroscopic signals
from below the tissue surface, yielding details about the
atheromatous necrotic tissue and sub-surface calcifications. These
signals can be utilized with an optical fiber based imaging system
to determine the content and composition of different
atherosclerotic plaques in vivo.
[0074] With the observation that several of the biochemical species
important in the atherosclerotic process, including cholesterol and
calcium hydroxyapatite, can be easily detected below the tissue
surface, we wished to determine the depth limit of detection using
the NIR FT Raman technique. In order to address this question, ten
200 .mu.m sections of aortic media were cut and place one at a time
over a large calcified deposit (6.times.6.times.3 mm), and the FT
Raman spectra of the 960 cm.sup.-1 band monitored as a function of
depth below the surface. As indicated by the plot of FT Raman
intensity versus depth shown in FIG. 7, the signal from the
calcified deposit was detectable until the deposit was greater than
1.6 mm below the irradiated surface. Even slightly deeper depths
could be probed if the focus of the collection optics was moved
into the tissue.
[0075] The two dimensional resolution of the NIR FT Raman signal
for material below the tissue surface was then tested by placing 1
mm of aortic media above another calcified deposit, and moving the
tissue transversely in two dimensions through the laser beam and
collection lens. The FT Raman signal was observed to drop-off
rapidly as the beam and collection optics moved from the calcified
deposit. The detected FT Raman signal closely followed the geometry
of the calcified deposit below the surface, despite the significant
scattering of the overlying layer of tissue. This result suggest
that the Raman scattered light may be utilized for imaging objects
below the tissue surface with minimal image blurring due to elastic
scattering in the tissue.
[0076] A second spectroscopic method is also used to obtain
molecular vibration information, attenuated total reflective (ATR)
of infrared light.
[0077] Human aorta was chosen as an example to illustrate the
diagnosis of atherosclerotic artery tissue. As in the samples
obtained for the Raman spectral measurements human aorta samples
were obtained for ATR measurements at the time of post mortem
examination. Sample storage and preparation procedures are
identical to those set forth for the Raman spectral measurements.
These reflectance measurements can be used by themselves to provide
diagnostic data in conjunction with either the Raman spectroscopic
measurements described above or with fluorescence measurements, or
with both types of measurements to enhance diagnosis for specific
applications.
[0078] The medial layers of a normal arteries and the necrotic
cores of atheromatous plaques were exposed by blunt dissection and
spectroscopically examined. ATR spectra were also collected from
several purified tissue components including collagen, elastin, and
cholesterol to assist in analysis of the spectra.
[0079] Mid-infrared ATR spectra were measured from 4000 to 700
cm.sup.-1 with a commercially available FT-IR spectrometer and a
horizontal ATR accessory. The sampling area was purged with dry
nitrogen gas to control background absorption from atmospheric
water vapor and CO.sub.2. Spectra were collected at 4 cm.sup.-1
resolution with 50 scans. The artery specimens, kept
physiologically moist with isotonic saline (buffered at pH 7.4),
were placed in contact with the ATR element (ZnSe crystal 45 ends).
A 5 gram weight placed on the tissue sample ensured uniform sample
contact with the ATR element. An ATR spectrum of the saline
solution with absorbance similar to that of the artery samples was
also obtained and used for subtraction of spectral components due
to water. Collagen (Calbiochem: type I, bovine Achilles tendon) and
elastin (Sigma: bovine neck ligament) were prepared as saline
slurries. Cholesterol (Sigma) was prepared as a dry film on the ATR
element by evaporation of a benzene solution. These elements can be
clearly identified in the resulting spectra.
[0080] The ATR sampling crystal is a rod of high refractive index
material which acts as a waveguide for the infrared sampling beam.
This waveguide can be in the form of a needle that is adapted for
penetration into the tissue to be diagnosed. Alternatively, the
probe will have a geometry suitable for contacting the surface of
exposed tissue sites or for contacting internal locations with a
catheter.
[0081] The devices shown in FIGS. 16A and 16B illustrate preferred
embodiments of the invention adapted for ATR diagnostic
measurements within the human body. In FIG. 16A a single-ended
probe 100 is shown where one or more optical fibers 102 both the
incident light to, and the transmitted (reflected) light from, the
ATR element 104. A 100% infrared reflector 106 such as gold is
placed at the distal surface 108 of the ATR element 104 functions
to return the transmitted light back through the same fiber as well
as to provide double pass sampling. The ATR element 104 can be a
separate component optically fastened to the optical fibers 102, or
alternatively, it can be constructed from the end of the optical
fiber by removing the cladding material. Sampling is provided by
placing the ATR element in contact with the tissue 110 of interest.
Radiation is transmitted 112 and collected 114 in a radial
direction from element 104. The probe can either be inserted
through a standard endoscope or catheter to sample a hollow organ,
or, if made with sufficiently thin optical fiber, it can be
directly inserted directly into a solid organ as in the case of a
needle biopsy. In this particular embodiment the distal top 108 is
in the form of a needle. The cone or needle configuration on the
end of the catheter can be long or shallow.
[0082] A double-ended probe is illustrated in FIG. 16B. Incident IR
beam from FT-IR is transmitted through IR optical fiber 122 to ATR
element 128 positioned at the distal end of catheter body 120. The
ATR element is placed in contact with tissue 126 surface to be
sampled. Transmitted light is conducted through a second IR optical
fiber 124 back to an IR detector. The ATR element may be a separate
component optically fastened to the two optical fibers 122, 124, or
it may be simply a region of a single optical fiber in which the
fiber cladding material has been removed. The entire apparatus can
be inserted through a standard endoscope or outer catheter.
[0083] For methods of measuring excised samples, the specimen to be
sampled is placed in optical contact with the surface of the
waveguide or ATR element. The evanescent wave which extends outside
of the waveguide surface is absorbed by the sample in proportion to
its absorption coefficient. The penetration depth of the evanescent
wave into the sample depends on the wavelength of the infrared
radiation and the refractive indices of the waveguide and the
sample; for a ZnSe-water interface, this depth is roughly 1 .mu.m
from 1800 to 700 cm.sup.-1. The 1/e penetration depth of the
evanescent wave into the sample is given by
.lambda./2.pi.(n.sub.z.sup.2sin.sup.2.THETA.-n.sub.w.sup.2).sup.1/2,
where .lambda. is a wavelength, .THETA. is an angle of incidence
and n.sub.z and n.sub.w are the refractive indices of ZnSe and
water respectively. Consequently, only tissue that is in good
optical contact with the ATR element will be sampled. In addition,
individual components in the sample can exhibit different
affinities for the waveguide material (ZnSe in this case), which
can influence the relative concentrations of the components at the
waveguide surface. Despite relatively high concentrations in the
bulk tissue, components with poor optical contact can be difficult
to measure in the ATR spectrum.
[0084] FIG. 8 shows FT-IR ATR spectra of (a) normal aorta (intimal
side) and (b) buffered saline. A comparison of these spectra shows
that a majority of the IR absorption of normal intima can be
attributed to water, which comprises roughly 80% of the tissue by
weight. The large, broad bands peaking at 3300 cm.sup.-1 and 1636
cm.sup.-1 are due to the O--H stretching and H--O--H bending
vibrations, respectively, of water, and the weak band at 2120
cm.sup.-1 is due to a water combination vibration. The 3300
cm.sup.-1 and 1636 cm.sup.-1 bands also include contributions from
the N--H stretching and amide I vibrations. The relatively flat
absorption between 1500 and 900 cm.sup.-1 and the rising absorption
below 900 cm.sup.-1 is also due primarily to water; however, in the
intima, a number of very weak bands due to other tissue components
are also present in this region.
[0085] Most biomolecules give rise to IR absorption bands between
1800 and 700 cm.sup.-1, which is known as the "fingerprint region"
or primary absorption region. The dominant absorption of tissue
water in this region obscures the absorption bands from other
tissue components. To observe the IR bands from these components,
one must eliminate the water interference. With the ATR method,
spectral deconvolution or subtraction of the water component is
particularly easy. By using the 2120 cm.sup.-1 band, which is due
solely to water, as an internal intensity standard the spectrum of
buffered saline (FIG. 8, curve b) can be accurately and reliably
subtracted from the spectrum of aorta intima (FIG. 8, curve a),
yielding a water-subtracted spectrum of intima (FIG. 9, curve
a).
[0086] In the water-subtracted spectrum, the previously weak bands
are easily observed. Band assignments, based on the spectra of the
major tissue components are listed in Table I. Most of the
vibrational bands observed in the spectrum of normal intima (FIG.
9, curve a) can be divided into two broad categories:
lipid-associated bands and protein-associated bands. All of the
strong bands in normal intima are associated with one of these
moieties (see Table I). This can be seen as a simple consequence of
the large concentrations of these two materials. Aside from water,
a large fraction of tissue can be divided into one of these two
groups. Moreover, both protein and lipid components contain
repeating molecular units which are common to all members of the
group. For protein, the polypeptide backbone of repeating amide
groups is the dominant element. In lipids, the repeating
hydrocarbon chain is the defining quality. The end result is that
these molecular units are present in very large concentrations, and
their vibrational bands tend to dominate the spectrum. Note that
this does not imply that no further level of detail is derivable
from the IR spectrum of tissue. For example, the frequencies of the
amide group vibrations are sensitive to protein configuration and
conformation. Therefore, shifts in protein makeup might be expected
to produce observable changes in the amide bands.
[0087] The water-subtracted spectrum of sub-adventitial fat shown
in FIG. 9, curve b, more clearly illustrates the division of bands
into lipid and non-lipid categories. This fat can be considered as
the model of the lipid component. Protein contributions, as judged
from the intensities of the amide I and II bands, are negligible
for the purposes of this model. All of the bands observed in the
fat spectrum can be attributed to the lipid component. These
include the strong bands at 1744 cm.sup.-1 (C.dbd.O stretch), 1456
cm.sup.-1 (C--H bend), 1161 cm.sup.1 (CH.sub.2 wag, C--O--C
stretch), as well as the weaker bands at 1378 cm.sup.-1, 1239
cm.sup.-1, 1118 cm.sup.-1, 1099 cm.sup.-1, 966 cm.sup.-1, and 722
cm.sup.-1.
[0088] The bands observed in the water-subtracted spectrum of
intima constitute less than 30% of the total absorption, the rest
being due to water. Any conclusions regarding these relatively weak
bands depends critically upon the accuracy of the water
substraction. The accuracy of this subtraction can be judged from
the reproducibility of spectra obtained sequentially from the same
sample. Two consecutive water-subtracted spectra collected 10
minutes apart from a sample of normal aorta (intimal side) are
shown in FIG. 10, curve a, (solid and dashed curves). Most of the
IR bands exhibit a substantial increase in absorbance with time.
This trend continues for consecutive spectra collected more than an
hour after the placement of the sample on the ATR element. However,
not all of the bands change by the same fraction, so that the
relative intensities differ between consecutive spectra. For
instance, in FIG. 10, curve a, the C.dbd.O band at 1744 cm.sup.-1
is relatively constant, while the amide bands at 1650 cm.sup.-1 and
1547 cm.sup.-1 increase by 50% in the later spectrum. Although
these changes might seem to indicate that the water subtraction is
inaccurate, the changes with time are systematic and predictable,
which suggests that the optical contact between the sample and the
ATR element is changing regularly with time.
[0089] In fact, the constancy of the 1744 cm.sup.-1 C.dbd.O band,
which is due solely to lipid, and the increases in the amide bands,
which are due to protein, indicate that the lipid contributions to
the IR absorption remain unchanged while the non-lipid
contributions increase between consecutive scans. This is confirmed
by subtracting the spectrum of lipid (FIG. 9, curve b) from the
water-subtracted spectra of aorta intima (FIG. 10, curve a), using
the 1744 cm.sup.-1 band for normalization. The resulting
lipid-subtracted spectra of aorta intima are shown, normalized to
peak absorbance, in FIG. 10, curve b. As can be seen, the relative
peak absorbencies and spectral bandshapes in the lipid-subtracted
spectra reproduce quite well, reflecting the accuracy of both the
water and the lipid-subtraction procedures.
[0090] The constancy of the lipid bands and the variation of the
non-lipid bands between successive scans may seem somewhat
puzzling. An explanation of this apparent anomaly can be inferred
from a water-subtracted spectrum of saline solution which is rinsed
off the surface of the tissue (FIG. 9, curve c). This spectrum,
aside from the weak amide I and II bands, matches quite closely
with that of adventitial fat. The lipid component observed in the
tissue appears to be due to free lipid particles that have
equilibrated with the tissue surface water, forming a thin
water-lipid film on the tissue surface which is in full optical
contact with the ATR element immediately after the tissue specimen
is placed upon the crystal. The tissue components beneath this film
presumably achieve better optical contact with the ATR crystal as
the sample settles. As a result, the content of lipid in a spectrum
of aorta intima or media may be influenced by the presence of
sub-adventitial fat in the specimen, and the relative lipid-protein
absorbencies are accurate to 50% at best with the present
experimental design. For the reason, all of the remaining spectra
show are both water and lipid subtracted.
[0091] With the lipid bands removed, assessment of the non-lipid
bands in the spectrum of normal intima (FIG. 10, curve b) is
greatly simplified. The major bands in the spectrum may be assigned
to protein backbone vibrations. These include the bands at 1648
cm.sup.-1 (amide I), 1549 cm.sup.-1 (amide II), 1455 cm.sup.-1
(C--H bend), 1401 cm.sup.-1 (amide C--N stretch), and 1244
cm.sup.-1 (amide III). The frequency of the amide I peak (1648
cm.sup.-1), which is sensitive to protein secondary structure, may
indicate contributions from .alpha.-helix, disordered, and collagen
helix conformations. This band also exhibits a shoulder near 1634
cm.sup.-1, which may be due to the .beta.-sheet regions of proteins
or water. The protein C--H bending band at 1455 cm.sup.-1 is
distinct from the corresponding vibration in lipid, which occurs as
a double-peaked band at 1465/1457 cm.sup.-1. Note that all of these
bands may include contributions from other moieties. For instance,
the symmetric stretch of carboxylate groups and the antisymmetric
stretch of phosphate groups may also contribute, respectively, to
the 1401 cm.sup.-1 and 1244 cm.sup.-1 bands. This correlation of
components is summarized in Table I above.
[0092] A typical spectrum of the medial layer of normal aorta is
shown in FIG. 11, curve a. A comparison of this spectrum to that of
normal intima (FIG. 10, curve b) fails to reveal any significant
differences. This result is somewhat surprising, considering that
normal aorta intima and media have significantly different
compositions. Typical spectra of an atherosclerotic plaque and a
non-ulcerated atheromatous plaque are shown in FIGS. 11, curve b
and 11, curve c, respectively. For these plaques, only the intact
fibrous cap at the intimal surface is probed due to the short
penetration depth (1 .mu.m) of the beam. Any necrotic, atheromatous
material beneath this fibrous cap is not sampled. Even so, the
fibrous caps of these plaques are known to be compositionally
different than normal intima and one might expect these differences
to be reflected in the IR ATR spectrum. However, as in the case of
media, no consistent differences are observed in the spectra of
these plaques (FIGS. 11, curve b and 11, curve c) and normal intima
(FIG. 10, curve b). This issue will expanded upon in the discussion
below.
[0093] On the other hand, substantial differences are obvious in
the spectrum of the necrotic, atheromatous core of an atheromatous
plaque (FIG. 12, curve a) as compared with the corresponding
spectra of normal intima (FIG. 10, curve b) as well as those of
intact atherosclerotic (FIG. 11, curve b) and atheromatous (FIG.
11, curve c) plaques. In this case, the necrotic core was
presumably exposed in vivo as disease progressed by ulceration of
the overlying intimal fibrous tissue cap. (The spectrum of necrotic
core exposed by dissecting away the fibrous cap of a non-ulcerated
atheromatous plaque is similar.) A new band appears at 1050
cm.sup.-1, with a secondary peak at 1023 cm.sup.1. In addition, the
necrotic core spectrum exhibits an increase and frequency shift in
the 1466 cm.sup.-1 bank as compared with the 1455 cm.sup.-1 protein
band in normal intima as well as a set of unique bands near 1382
cm.sup.-1. These characteristic bands are found in the spectra of
all the exposed necrotic core samples and in none of the other
samples (see below).
[0094] The source of these unique bands in the necrotic core
spectra may be cholesterol, which is known to accumulate in large
amounts in atheromatous cores. An ATR spectrum of cholesterol (dry
film) is shown in FIG. 12, curve b. The three major bands unique to
the necrotic core, near 1463 cm.sup.-1, 1382 cm.sup.-1, and 1050
cm.sup.-1, match closely in position and relative intensities with
the three main cholesterol bands at 1466 cm.sup.-1, 1377 cm.sup.-1,
and 1056 cm.sup.-1. Each of the main cholesterol bands has a
secondary peak, which also appear to be present in the necrotic
core bands. These secondary peaks occur at 1445/1436 cm.sup.-1, and
1023 cm.sup.-1 in the cholesterol spectrum and at 1441 cm.sup.-1,
1367 cm.sup.-1, and 1023 cm.sup.-1 in the necrotic core spectrum.
In addition, several of the weak bands in the necrotic core
spectrum, including the peaks at 1334 cm.sup.-1, 1109 cm.sup.-1,
954 cm.sup.-1, and 797 cm.sup.-1, are associated with the weaker
cholesterol bands near these frequencies. Other components in the
necrotic core may also contribute to some of these distinct
bands.
[0095] The consistency of the spectral differences which are
attributed to cholesterol between the necrotic core specimens and
the normal, atherosclerotic, and non-ulcerated atheromatous
specimens are illustrated in the scatter plot in FIG. 13. This plot
depicts the integrated intensities (areas) of the 1050 cm.sup.-1
cholesterol band ratioed to the total protein content, as measured
by the area of the amide II band at 1548 cm.sup.-1. The 1050
cm.sup.-1 band was integrated from 1075 to 1000 cm.sup.-1 and
baseline subtracted using these endpoints, and the amide II band
was integrated from 1593 to 1485 cm.sup.-1 with a similar baseline
subtraction. This ratio is a measure as the relative cholesterol
contribution to the spectrum, and is proportional to the relative
cholesterol concentration of the sample with the assumption that
the area of the 1050 cm.sup.-1 band is due solely to cholesterol.
As can be seen in FIG. 13, this ratio is higher for all the exposed
necrotic core specimens than for all the other specimens. The
consistent results of this sample analysis, which is possible
because of the separation and molecular identification of the
bands, indicates the potential of IR spectroscopy for tissue
characterization.
[0096] Investigations of human arteries and atherosclerosis by
mid-IR spectroscopy have been limited to date. It has been reported
that ATR spectra have been recorded from partially dried human
artery, among other tissues. In comparing a normal aorta from an
infant to an atherosclerotic plaque in an adult, they observed
increases in several bands in the atherosclerotic aorta. Most of
these bands were associated with lipids and lipoproteins. IR
spectroscopy has been employed to determine the chemical
composition of calcified atherosclerotic deposits. A more detailed
IR study of atherosclerotic aorta involves recorded IR transmission
spectra from thin layers sectioned at different depths into the
arterial wall. Results showed increased absorption near 1739
cm.sup.-1 in the fatty (atheromatous) regions of plaque, which was
attributed to absorption by cholesterol esters in the plaque. IR
spectra from the fibrous tissue cap at the surface of the plaques
were similar to normal intima.
[0097] One of the main difficulties in measuring mid-infrared
spectra of intact human tissue is the intense water absorption,
which dominates and obscures the absorption of other tissue
components of interest. In most of the studies cited above, the
water absorption was not eliminated, limiting the quality and
amount of information available from the spectra. With the ATR
sampling method, this water interference is easily removed (see
FIG. 9). The ATR method is also naturally amenable to sampling with
fiber optic probes in vivo. Water interference in fiber optic probe
ATR spectra of aqueous protein solutions has been accurately
eliminated with a water subtraction procedure similar to the one
employed in the present study.
[0098] While the ATR method is well suited to in vivo sampling and
to accurate subtraction of the water signal, spectra collected with
the ATR method are not equivalent to IR absorption spectra, but
depend on properties of the ATR material and the sample in addition
to the sample absorption coefficient. For instance, the penetration
depth of the evanescent sampling wave depends on the refractive
indices of the ATR material and the sample. However, the refractive
indices of both ZnSe and human tissue are expected to vary slowly
with frequency between 1800 and 700 cm.sup.-1 and such variations
will at most affect the relative intensities of bands at different
frequencies. All of the structure observed in the tissue spectra is
attributed to absorption bands in the tissue.
[0099] The component absorptions observed in an ATR spectrum also
depends upon the optical contact of the sample and ATR element. The
small penetration depth of the evanescent wave into the tissue
sample implies that only a 5 .mu.m thick layer, and preferably
about 1 micron, of material at the surface is observed. This is
referred to as the near surface region of the tissue for the
purposes of this application. The tissue deeper than 5 microns from
the surface is defined as the sub-surface region. This thin,
sampled near-surface layer may differ in composition with the bulk
sample. For example, a film of free water may be present on the
surface of wet tissue, with different levels of some molecular
species of the tissue relative to their concentrations in the bulk
tissue. In addition, the varied affinities for the ATR material of
different moieties in the tissue may play an important role in the
intensities of the observed bands.
[0100] These considerations may explain the lack of substantial
differences among the ATR spectra of normal intima, plaque fibrous
cap, and media. For instance, normal aorta intima is composed of
roughly 25% collagen (dry weight) and 20% elastin, while aorta
media has 20% collagen and 50% elastin. The ATR spectra of purified
collagen and purified elastin (not shown) differ substantially. In
particular, amide I/II occur at 1657/1556 cm.sup.-1 in hydrated
collagen (type I) and 1653/1543 cm.sup.-1 in hydrated elastin
(spectra not shown).
[0101] One might expect these differences to be reflected in the
intima and media ATR spectra. A possible explanation of why this is
not the case is that the thin layer in optical contact wit the ATR
element is compositionally different from the bulk tissue, and
collagen and elastin make only a minor contribution to the IR ATR
bands of this layer. Such an effect may also explain the lack of
significant differences among the plaque fibrous cap intima and
normal intima ATR spectra. In ATR elements made of other substances
with different biochemical affinities, the spectral differences
among these tissues can be substantially enhanced depending on the
tissue type.
[0102] The results of the present investigation demonstrate that
high quality water-subtracted spectra can be readily obtained from
human tissue with ATR technique. Similar results have been obtained
in other mammalian tissues. Accurate removal of the water
interference is crucial to isolating the relatively weak tissue
absorption bands of lipid, protein, as well as other tissue
components. It is worth noting that the observation of these
relatively weak bands via spectral subtraction depends entirely
upon quality of the tissue and saline spectra. For instance, the
absorbance of the normal intima specimen (FIG. 8, curve a) between
1500 and 900 cm.sup.-1 is approximately 0.06. In the
water-subtracted spectrum (FIG. 9, curve a), the peak absorbencies
range from 0.018 (30%) for the strongest bands to 0.003 (5%) for
the weakest ones. The detection of a 0.003 absorbance peak in a
subtracted spectrum with a 0.06 absorbance background requires a
signal-to-noise ratio of 700 or better in the 100% baseline. Such a
signal-to-noise is easily achieved with an FT-spectrometer. The
high linearity, baseline stability, and wavelength precision of the
FT-spectrometer are also obviously critical for accurate background
subtraction.
[0103] While water subtraction is relatively easy and accurate with
ATR, it may be substantially more difficult with other clinically
applicable sampling techniques such as diffuse reflectance or
photoacoustic sampling. These alternative sampling techniques are
clinically useful, however, because of their longer tissue
penetration depths (approximately 10 .mu.m). As an alternative to
water subtraction, one can exploit the properties of the spectral
lineshape of water to eliminate the water signal by other
computational methods. Specifically, the spectral lineshape of
water varies rather slowly with frequency over much of the region
of interest, especially between 1500 and 700 cm.sup.-1. Therefore,
any method which filters this slower variation and spares the
sharper features of the non-water bands can separate the water and
non-water components.
[0104] One such method is second derivative spectroscopy.
Differentiation of a spectrum is typically used to narrow
absorption bands and resolve overlapping peaks. Differentiation
also tends to deemphasize broad bands relative to sharper ones. In
IR spectra of artery, the board, featureless absorption of water
can be nearly eliminated in favor of the sharper non-water bands by
computing the second derivative of the spectra. This is clearly
demonstrated in FIG. 14, which shows the second derivative of a
spectrum of normal aorta intima (FIG. 14, curve a), along with the
water-subtracted spectrum of the same specimen (FIG. 14, curve b).
Essentially only the 1633 cm.sup.-1 water band is left, partially
obscuring the amide I band. Elsewhere in this spectrum, the water
contribution is minimal. All of the bands identified in the
water-subtracted spectrum are easily observed in the second
derivative spectrum.
[0105] In addition to elimination of water interference, several of
the unresolved double peaks and shoulders in the water-subtracted
spectrum appear as distinct peaks in the second derivative
spectrum. For example, the amide II band in normal intima (FIG. 14,
curve b) has a very weak shoulder near 1518 cm.sup.-1, and the C--H
bending region near 1468 cm.sup.-1 appears to include two
overlapping peaks. In the second derivative spectrum (FIG. 14,
curve a), the 1518 cm.sup.-1 band is clearly visible, and the C--H
region exhibits two separate peaks at 1469 and 1456 cm.sup.-1.
Moreover, by sharpening the bands, the second derivative spectrum
allows a more precise determination of peak frequencies, so that
relatively small frequency shifts are observed. Such frequency
shifts can be of importance in detecting and characterizing subtle
molecular alterations involved in certain tissue conditions.
[0106] The observation of individual, resolved bands in the artery
IR ATR spectra is of considerable interest, since separation of
bands is the first step determining the composition of a sample
from its spectrum. Once a band has been isolated, its integrated
intensity is proportional to the concentration of the moiety
responsible for that band. In particular, since the amide I and II
bands are due entirely to protein, these bands can be used to
isolate the overall protein content in the spectrum. The sharp,
well resolved 1744 cm.sup.-1 C.dbd.O ester band appears to be due
to solely to lipid, and the integrated intensity of this band
should be proportional to the relative lipid content are technique
should largely eliminate the inaccuracies. Finally, it should be
remembered that the relative water content of the tissue sample is
automatically computed from the 2120 cm.sup.-1 band in the water
subtraction algorithm. However, as noted earlier, the composition
of tissue as determined from an ATR spectrum may not be precisely
identical to the composition of the bulk tissue.
[0107] The tissue composition can also be determined from
overlapping bands by first deconvolving the bands of interest into
their individual components. This is especially easy if one
component has an additional, isolated band elsewhere in the
spectrum. An example is the 1465 cm.sup.-1 C--H bending region,
which is due to different tissue components with distinct spectral
features in this region. In the normal intima spectrum (FIG. 9,
curve a), this band is attributed to a combination of lipid and
protein components. Since the lipid component also exhibits the
isolated 1744 cm.sup.-1 band, this band can be used to subtract the
lipid C--H bending component and isolate the protein C--H bending
component at 1455 cm.sup.-1 (FIG. 10, curve b), effectively
deconvolving this band. Note that this deconvolution depends on
having a reliable spectrum of one of the individual components,
which, in this example, is the lipid spectrum in FIG. 9, curve
b.
[0108] The detection of distinct bands attributed to cholesterol in
necrotic core may provide a useful means of determining cholesterol
concentrations in vivo. Both the 1050 cm.sup.-1 and 1382 cm.sup.-1
cholesterol bands appear to be fairly isolated in the necrotic core
spectrum after lipid-subtraction (FIG. 12). If these two bands are
due to a single component, namely cholesterol, the ratio of their
integrated intensities should be a constant for all the samples.
The baseline-subtracted area of the 1050 cm.sup.-1 band, A(1050),
is plotted versus that of the 1382 cm.sup.-1 band, A(1382), for all
the samples, normalized to the protein content in FIG. 15. As can
be seen in the plot, there is a roughly linear relationship between
A(1050) and A(1382). A linear least squares fit to this data yields
the line shown in the plot, with a high regression coefficient of
r=0.967. The slope of this line 2.8, while the ratio
A(1050)/A(1382) for the pure cholesterol ATR spectrum is 2.3. The
reasonable agreement between these two numbers provides additional
evidence for the assignment of these bands to cholesterol.
Moreover, it indicates that the relative spectral content of
cholesterol is reasonably approximated by the integrated
intensities of either of these bands. FIG. 15 also shows that the
ATR spectra of all the specimens other than exposed necrotic core
exhibit almost no intensity in both the 1050 and 1382 cm.sup.-1
bands, in contrast to the necrotic specimens, all of which have
significant bands at both frequencies.
[0109] The present systems and methods demonstrate that infrared
spectra of moist, bulk tissues can be reliably obtained with the
ATR technique. Although water is the dominant absorber throughout
much of the mid-infrared region, the high quality spectra acquired
with the FT-IR ATR technique allow for accurate subtraction of the
water signal. Elimination of the water interference is critical for
identifying and assigning the absorption bands of other tissue
species. The isolation and designation of these relatively sharp
bands provides a means of analyzing spectroscopically the
composition of arterial tissue non-destructively. These methods are
also applicable to the study and diagnosis of other tissues and
tissue conditions, such as neoplasia.
[0110] The observation of both lipids and cholesterol in the
spectra of necrotic atheromatous core samples is particularly
exciting, because lipids and cholesterol are thought to play major
roles in the pathogenesis of atherosclerosis. The spectral
observation of these components, cholesterol in particular,
provides a reliable means of detecting and characterizing advanced
atheromatous plaques in which ulceration of the fibrous cap has
occurred (as demonstrated in FIGS. 13 and 15). Intimal
accumulations of lipid and cholesterol occur early in the
atherogenic process. Therefore, the mid-IR ATR technique can also
be useful in detecting and studying the early fatty streak
lesion.
[0111] Spectrograph/CCD System for NIR Raman Spectra
[0112] NIR Raman spectroscopy using a single stage spectrograph and
a charge coupled device (CCD) detector offers superior sensitivity
over the Nd:YAG excited FT-Raman system of FIGS. 1A and 1C. By
shifting the wavelength of the laser excitation from 1064 nm to the
800-900 nm region, a CCD can be used to detect the Raman scattered
signals while still avoiding fluorescence excitation in most
molecules. The system can operate usefully in the range of 750 nm
to 1050 nm. Although the fluorescence emission from tissue is
significantly higher with 810 nm than with 1064 nm excitation, the
Raman signals are readily observed. This is because the dominant
noise source in the spectrograph/CCD system is shot noise
associated with the fluorescence emission, which is 2-3 orders of
magnitude smaller than the dark current noise of the InGaAs
detector, which is the dominant noise source in the FT-Raman
system.
[0113] FIG. 17 shows the laser diagnosis and treatment system of
FIG. 1A modified to use the spectrograph/CCD system of this
invention. The diagnostic subsystem 201 includes a single stage
spectrograph 310 and charge-coupled device (CCD) detector 312 for
collecting near-infrared (NIR) Raman spectra from intact human
arterial tissue. With 810 nm laser light excitation, preferably
pulsed, the fluorescence emission from human artery tissue is
sufficiently weak to observe Raman bands more rapidly with the
spectrograph/CCD system than with the 1064 nm excited FT-Raman
system of FIGS. 1A and 1C. A method for removing the broadband
emission from the spectra by computing the difference of two
emission spectra collected at slightly different excitation
frequencies was used to enhance observation of the Raman bands.
This method relies on the stability, linearity, and low noise
characteristics of the CCD detector. The results indicate that high
quality NIR Raman spectra can be collected in under 1 second with
the spectrograph/CCD system and an optical fiber probe, as compared
to 30 minutes with the FT-Raman system at similar laser power
levels, further improving the use of the technique for in vivo
clinical applications.
[0114] A preferred embodiment of a spectrograph/CCD system 300
employed for the collection of near infrared (NIR) Raman spectral
data from excised tissue samples using a spectrograph and a charge
coupled device (CCD) array is illustrated in FIG. 18. NIR Raman
spectra were measured from 100-2000 cm.sup.-1 below the laser
excitation frequency with a single stage imaging spectrograph 310
(Acton Model ARC275, 0.25 m, F/3.8) and a CCD array 312 (Princeton
Instruments EEV Model 88130).
[0115] System 300 can use a NIR 810 nm Nd:YAG pumped pulsed dye
laser 314 operating at 10 Hz for irradiating a sample 46.
Alternatively, a CW or diode laser source can also be employed.
Laser 314 generated a laser beam 316 which is directed by mirror
318 through focusing optics 320 to impinge on sample 46 mounted
behind a transparent window 321. The laser beam was focused on the
sample at a 70.degree. angle of incidence, yielding a spot size of
0.7.times.2 mm on the tissue surface. The average incident power at
the sample was maintained at 20 mW to avoid excessive peak
intensities during an individual pulse. The spectral signals were
observed to be linear over a range of average incident powers from
2 to 20 mW.
[0116] A portion of the scattered light 322 emitted by sample 46
was collected by collecting optics 324 at a 90.degree. angle
relative to the incident laser beam. Collecting optics 324
collimates and F/matches the collected light for the spectrograph
310. Prior to entering the entrance slit of the spectrograph 310,
the collected light was passed through a series of Schott glass
filters 326 which attenuated the elastically scattered component of
the collected light. The combined effect of the Schott glass
filters provided an optical density of 7 at 810 nm, a transmission
of 20% at 850 nm (580 cm.sup.-1 from 810 nm), and a transmission of
85% above 900 nm (1200 cm.sup.-1).
[0117] The spectrograph 310 utilized a 200 .mu.m slit width and a
600 groove/mm grating blazed at 1 .mu.m and could be scanned to
provide spectral coverage over different wavelength regions. The
200 .mu.m slit width provided a resolution of roughly 15
cm.sup.-1.
[0118] The CCD array 312 consisted of 298 (column) by 1152 (row)
pixel elements having a total active area of 6.7 mm.times.26 mm,
with the short axis parallel to the slit. The CCD array was cooled
to -110.degree. C. to eliminate dark current. Each row of pixels
was binned to reduce readout noise. Commercially available CCD
detectors offer extremely low detector noise and usable quantum
efficiencies out to 1050 nm and provide substantial advantages over
InGaAs and other NIR detectors. These advantages outweigh the lower
throughput of the grating spectrograph, provided that broadband
fluorescence interference is not too great with the shorter
excitation wavelengths.
[0119] Excised human aorta samples 46 obtained at the time of
post-mortem examination were rinsed with isotonic saline solution
(buffered at pH 7.4), snap-frozen in liquid nitrogen, and stored at
-85.degree. C. Prior to spectroscopic examination, samples were
passively warmed to room temperature and were kept moist with the
saline solution. Normal and atherosclerotic areas of tissue were
identified by gross inspection, separated, and sliced into roughly
8.times.8 mm pieces.
[0120] The tissue samples 46 were placed in a suprasil quartz
cuvette with a small amount of isotonic saline to keep the
specimens moist, and with one surface in contact with the
transparent window 321 and irradiated by the laser 314.
[0121] Raman spectra were typically measured between 100 cm.sup.-1
and 2000 cm.sup.-1 below the laser excitation frequency. Each
spectrum was background subtracted to remove the DC offset of the
A/D converter of the CCD controller. In addition, hot pixels due to
high energy radiation events were removed from the recorded
spectrum by applying a median filter having a 7 pixel wide window
as to each spectrum. Raman frequencies were calibrated with the
spectra of benzene and barium sulfate powder and are accurate to
.+-.5 cm.sup.-1. The spectra were not corrected for the wavelength
dependent response of the filters, spectrograph, and CCD. For each
spectrum shown in the following Figures, Raman signals were
accumulated for 5 minutes. Substantially shorter collection times
can also be used as described herein.
[0122] FIG. 19A shows the Raman spectra of a normal aorta sample
excited with 810 nm laser light and collected with the
spectrograph/CCD system 300. In this case, the broadband background
emission, which is presumably due to tissue fluorescence, is
roughly five times more intense than the strongest Raman bands at
1650, 1451, 1330, and 1253 cm.sup.-1. In contrast, the 1064 nm
FT-Raman study of normal human aorta shown in FIG. 2a exhibited
Raman signals with the peak intensities of the strongest bands,
amide I at 1650 cm.sup.-1 and C--H bend at 1451 cm.sup.-1, being
roughly three times larger than the broadband background emission.
However, this background emission in the spectrograph/CCD system is
relatively weak with respect to the Raman signals (i.e., on the
order of the Raman signals) and therefore the shot noise associated
with detecting this background emission is substantially smaller
than the Raman signals, allowing the Raman bands to be made
distinct after the background emission signals are removed through
filtering or subtraction. The shot noise is typically random noise
exhibiting a Poisson distribution and is associated with the
detector and/or the background emission itself.
[0123] In contrast, with visible excitations, the fluorescence
background emission from the arterial pathology tissue types
described is 3 to 4 orders of magnitude larger than the Raman
signals, and the shot noise associated with this stronger
background emission completely obscures the Raman bands even after
the background emissions are removed. However, certain other types
of tissue, e.g., colon and bladder, do not exhibit such high level
fluorescence reactions at visible excitation frequencies, and
therefore can probably operate with visible excitations.
[0124] The signal-to-noise ratio of the spectrum of normal aorta
collected with the spectrograph/CCD system 300 with 20 mW incident
power and 5 minutes collection time (FIG. 19A) is similar to that
obtained with the FT-Raman system of FIG. 1C with 500 mW incident
power and 35 minute collection time. Since the observed spectral
signal-to-noise ratios are similar, we estimate that the noise
level observed with the CCD detector 312 of FIG. 18 is roughly 3400
times less than that observed with the InGaAs detector 42 of FIG.
1C. For the InGaAs detector, the major noise source is the shot
noise of the dark current, while with the CCD detector the dominant
noise source is the shot noise of the broadband tissue emission, as
the dark current and readout electrons of the CCD are much smaller
than this emission.
[0125] This simple analysis has several important implications.
First, since the major noise source encountered with the
spectrograph/CCD system is shot noise from broadband emission by
the tissue sample, the spectral signal-to-noise ratio is
proportional to the square root of the product of incident
intensity and the collection time.
[0126] The FT-Raman and spectrograph/CCD systems can be compared as
follows. For the FT-Raman system, the incident intensity is 640
mW/mm.sup.2. The quantum efficiency of the InGaAs detector at 1200
nm is 0.7, and the FT-spectrometer throughput is 1.1 mm.sup.2 sr,
and the transmission efficiency of the FT-spectrometer and filters
is roughly 0.062. For the spectrograph/CCD system, the incident
intensity is 14 mW/mm.sup.2. The CCD quantum efficiency is 0.15 at
900 nm, the spectrograph throughput is 0.043 mm.sup.2 sr, and the
transmission efficiency of the spectrograph and filters is 0.24.
Combining these factors and taking into account the v.sup.4
dependence of the Raman cross-sections, the signal level measured
by the FT-Raman spectrum is estimated to be 3400 times greater than
that of the spectrograph/CCD spectrum.
[0127] Therefore, if the laser intensity is increased to the level
employed in the FT-Raman experiments, the collection time could be
reduced by a factor of 40, to 8 seconds, with no change in the
spectral signal-to-noise ratio. Second, the noise level can be
further reduced by using longer excitation wavelengths which
minimize the tissue fluorescence emission. However, such reductions
in fluorescence emission must be balanced against the decreasing
quantum efficiency of the CCD at longer wavelengths, and the
optimum excitation wavelength also depends on the fluorescence
excitation profile of the tissue. For tissue types that exhibit
little fluorescence emission at visible wavelengths, such as colon
and bladder tissue, the CCD can be operated at visible or near
visible wavelengths to take advantage of increased quantum
efficiency of the CCD at these wavelengths. Finally, the throughput
of a 500 .mu.m core, 0.2 numerical aperture fused silica optical
fiber is 0.03 mm.sup.2 sr, which is roughly the same as that of the
spectrograph/CCD system. This means that the present lens
collection system can be replaced with an optical fiber probe, as
is required for in vivo operation, with no additional loss in
signal.
[0128] FIG. 19A shows that although the shot noise due to the
broadband tissue emission is relatively small, the sloping
broadband fluorescence emission still obscures the sharper Raman
signals and complicates determination of peak frequencies and
identification of weak bands. Furthermore, given the complexity of
human tissue, it is likely that this broadband emission will be
significant throughout the useful range of the CCD. Any
quantitative analysis of the Raman bands in FIG. 19A requires that
this broadband emission be first removed from the spectrum. The
standard methods of removing fluorescence emission from Raman
spectra utilize mathematical filters, which rely upon the
fluorescence emission being relatively featureless. In an
alternative method the excitation frequency is varied over a narrow
range (10-30 cm.sup.-1). The Raman band positions vary directly
with the excitation frequency, while the fluorescence emission
remains fairly constant with such small changes in excitation
frequency, allowing it to be efficiently subtracted out. In
contrast with mathematical filters, this operation requires no
assumptions about the emission lineshape.
[0129] To implement this method, the Raman spectrum of the normal
aorta specimen is recoded with excitation wavelengths of 810 nm
(FIG. 19A) and 812 nm. The Raman bands shift with the excitation
frequency by 30 cm.sup.-1, while the fluorescence emission remains
fairly constant. By subtracting these two spectra, the broadband
emission is greatly reduced, and the Raman bands are more readily
observed (FIG. 19B). This operation is mathematically analogous to
taking the derivative of the Raman spectrum, so that the original
Raman spectrum can be recovered by integrating the difference
spectrum, as shown in FIG. 19C. The fluorescence background is
greatly reduced in FIG. 19C as compared with FIG. 19A, allowing
easier identification of the Raman bands and their peak
frequencies. The integration also smooths the Raman spectrum over a
bandwidth similar to the excitation frequency shift and causes some
linewidth broadening, as is evident from FIG. 19C. Note that the
accuracy of this method depends upon the high linearity and
stability of the CCD array.
[0130] The NIR Raman spectrum of an atherosclerotic plaque with a
calcified deposit exposed at the surface collected with the
spectrograph/CCD system is shown in FIG. 20A. In this case, the
broadband emission is nearly 10 times greater than that observed in
normal aorta (FIG. 19A), resulting in increased noise. However, the
intense phosphate stretching vibration at 960 cm.sup.-1, due to the
calcified salts, is readily identified. This band is sufficiently
intense to be observed in real time and was used in aligning the
collection optics. Some weaker bands may also be identified, such
as the phosphate/carbonate band at 1070 cm.sup.-1, although these
are obscured by the large fluorescence background. By subtracting
out this fluorescence (FIG. 20B), as above, these bands are much
more easily distinguished. The Raman spectrum obtained by
integrating the difference spectrum is shown in FIG. 20C. The
broadband emission is reduced by a factor of 50 relative to Raman
bands, and several weaker bands are readily observed. This spectrum
is remarkably similar to that of FIG. 5a which was observed with
the FT-Raman system and 1064 nm excitation.
[0131] As another example of the sensitivity of the
spectrograph/CCD system 300, the Raman spectrum of adventitial
adipose tissue is shown in FIG. 21, which can be compared to the
FT-Raman spectrum shown in FIG. 5c. The broadband emission is
similar to that of normal aorta, while the Raman bands, due mainly
to triglycerides in the tissue, are very strong, resulting in an
excellent spectral signal-to-noise ratio.
[0132] Thus, the spectrograph/CCD system with 810 nm excitation
offers a faster alternative to FT-Raman with 1064 nm excitation and
which has greater sensitivity. Even in complex mixtures such as
human tissue, the level of background emission observed with 810 nm
excitation is low enough to observe the Raman signals. This
fluorescence emission does not excessively degrade the
signal-to-noise ratio. By subtracting two spectra collected at
slightly different excitation wavelengths, and then integrating the
difference spectrum, this broadband emission is rejected, yielding
high quality Raman spectra. Deconvolution techniques can also be
used to selectively remove, or reduce, Raman, fluorescence, or
noise light components. Improvements such as using a CW laser to
increase the incident intensity and a back-thinned CCD having
better red response allows Raman spectra to be collected from
intact human tissue in under 1 second. Longer excitation
wavelengths may reduce the background emission further.
Implementation of the spectrograph/CCD system with a high power
diode laser and an optical fiber probe will provide a compact,
mobile system for rapidly acquiring NIR Raman spectra remotely from
human tissues and will provide a powerful tool for in vivo clinical
applications.
[0133] Equivalents
[0134] Those skilled in the art will recognize, or be able to
ascertain using routine experimentation, many equivalents to the
specific embodiments of the invention described herein. These and
all other equivalents are intended to be encompassed by the
following claims.
1TABLE I Preliminary assignments of IR absorption peaks in the ATR
spectra of normal aorta intima. V Preliminary Vibrational
Associated Tissue (+1 cm.sup.-1) Assignment Components 2923(s) C--H
stretch Lipid, Protein, Others 2853(s) C--H stretch Lipid, Protein,
Others 1744(s) C.dbd.O (ester) stretch Lipid 1651(s) Amide I
Protein 1635(sh) Amide I, H--O--H bend Protein, Water 1548(s) Amide
II Protein 1465(s) CH.sub.2 bend Lipid 1457(s) CH.sub.2 bend,
CH.sub.3 antisymmetric Lipid deformation 1454(m) CH bend, CH.sub.3
antisymmetric Protein, others deformation 1417(w) CH.sub.2 bend
adjacent to C.dbd.O Lipid 1401(m) COO.sup.- symmetric stretch,
Protein, others amide C--N stretch 1378(w) CH.sub.3 symmetric Lipid
deformation 1244(m) Amide III, PO.sub.2.sup.- antisymmetric
Protein, others stretch 1239(m) CH.sub.2 wag, PO.sub.2.sup.-
antisymmetric Lipid stretch 1159(s) CH.sub.2 wag, C--O--C Lipid
antisymmetric stretch 1117(w) C--C stretch, O--C--O Lipid stretch
1096(w) Lipid 1083(w) PO.sub.2.sup.- symmetric stretch Protein,
others 1030(w) Lipid 965(w) C.dbd.CH deformation (trans) Lipid
722(m) CH.sub.2 rock Lipid
[0135]
2TABLE II Peak frequencies of selected bands in normal and
atheroscleriotic aorta. Fibrous Fatty Exposed Exposed Normal
Adventitia.sup.a Plaque Plaque Calcif. I Calcif. II Assignments
1746w C.dbd.O (ester) stretch 1667m 1667m 1667m C.dbd.C Stretch
Lipid 1658s 1663m Amide I (8) 1655m C.dbd.C stretch Fatty Acids
1519w Carotenoid (12) 1451s 1440s 1440s 1440s 1450s 1440s C--H bend
(8) Protein Lipid 1301m 1301w 1301w 1300w Lipid C--H 1267w 1264w
1262w 1262w bend (CH.sub.2) Lipid C--H bend (.dbd.C--H) 1252m 1261w
Amide III (8) 1157w Carotenoid (12) 1080m 1131w 1130w 1128w C--C
stretch 1086w 1088w Lipid 1071s 1071s Phosphate antisymmetric
stretch Calcium salts (15) 1004w 1004w 1004w Phenylalanine (8)
960vs 960vs Phosphate symmetric stretch Calcium salts (15) 957w
959w 878w Cholestreols 882w 882w 850w (11) 842w 841w 804w 803w 801w
699m 700m 700m 547w 606w 606w 546w 546w 587m 587m Phosphate Calcium
salts (15) Peak frequencies of typical specimens, accurate to .+-.1
cm.sup.-1. Abbreviations: vs = very strong, s = strong, m = medium,
and w = weak relative band intensity. .sup.aAdventitia specimen is
mainly adipose tissue.
* * * * *