U.S. patent application number 11/601904 was filed with the patent office on 2007-07-12 for method and apparatus for rapid detection and diagnosis of tissue abnormalities.
This patent application is currently assigned to SpectRx, Inc.. Invention is credited to Shabbir Bakir Bambot, Mark Faupel, David Jonathan Mongin.
Application Number | 20070161876 11/601904 |
Document ID | / |
Family ID | 38233589 |
Filed Date | 2007-07-12 |
United States Patent
Application |
20070161876 |
Kind Code |
A1 |
Bambot; Shabbir Bakir ; et
al. |
July 12, 2007 |
Method and apparatus for rapid detection and diagnosis of tissue
abnormalities
Abstract
A method and apparatus are provided that interrogate, receive,
and analyze full emission spectra for at least one fluorescence
excitation wavelength and for at least one reflectance measurement
to determine tissue characteristics and correlate same to
photographic images. Further, the system and method accomplish this
measurement rapidly by increasing the light throughput by
integrating optics into a hand held unit and avoiding the need for
a coherent fiber optic bundle being used. The method includes
illuminating a first portion of a target tissue with optical
energy, forming a first image of the target tissue, illuminating a
second portion of the target tissue with optical energy, performing
spectroscopic measurements on optical energy reflected and/or
emitted by the target tissue upon illumination of the second
portion of the target tissue with optical energy, and determining
tissue characteristics of the target tissue based on the results of
the spectroscopic measurements.
Inventors: |
Bambot; Shabbir Bakir;
(Norcross, GA) ; Faupel; Mark; (Norcross, GA)
; Mongin; David Jonathan; (Norcross, GA) |
Correspondence
Address: |
ALTERA LAW GROUP, LLC
6500 CITY WEST PARKWAY
SUITE 100
MINNEAPOLIS
MN
55344-7704
US
|
Assignee: |
SpectRx, Inc.
Norcross
GA
|
Family ID: |
38233589 |
Appl. No.: |
11/601904 |
Filed: |
November 20, 2006 |
Related U.S. Patent Documents
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Application
Number |
Filing Date |
Patent Number |
|
|
60737949 |
Nov 18, 2005 |
|
|
|
Current U.S.
Class: |
600/310 ;
600/407 |
Current CPC
Class: |
G01J 3/027 20130101;
A61B 5/0068 20130101; G01J 3/4406 20130101; G01J 3/0208 20130101;
G01J 3/28 20130101; A61B 5/0059 20130101; G01J 3/0291 20130101;
G01J 2003/1213 20130101; G01J 3/0232 20130101; A61B 5/0075
20130101; A61B 5/0071 20130101; G01J 3/02 20130101; G01J 3/0229
20130101; A61B 5/0077 20130101; G01J 3/0218 20130101; G01N
2021/6484 20130101; G01J 3/0264 20130101; G01N 2201/0221 20130101;
G01J 3/0283 20130101; G01N 21/274 20130101; G01N 21/31 20130101;
G01N 2021/6417 20130101; G01J 2003/2866 20130101; G01N 21/474
20130101; G01N 2021/6423 20130101; G01J 3/0272 20130101; G01N
21/645 20130101; G01J 3/0262 20130101; A61B 5/4331 20130101; G01J
3/10 20130101 |
Class at
Publication: |
600/310 ;
600/407 |
International
Class: |
A61B 5/00 20060101
A61B005/00; A61B 5/05 20060101 A61B005/05 |
Claims
1. An apparatus for determining tissue characteristics, comprising:
a base unit comprising an illumination unit, a separate tissue
interface unit comprising and excitation unit for delivering
illumination from the illumination unit to the tissue and a
detection unit capable of detecting responses in the tissue
resulting from the illumination, said excitation and detection
units being confocal so that they converge at the tissue.
2. The apparatus according to claim 1, further including a hollow
stand off tube, configured to space the separate interface unit a
predetermined distance from the tissue.
3. The apparatus according to claim 1 further including an imaging
device capable of recording images of the tissue.
4. The apparatus of claim 3 wherein said image device is located at
least in part in said separate interface unit and is confocal with
said detection unit so that the image of said image unit can be
spacially correlated with data from said detection unit.
5. The apparatus according to claim 3 wherein said stand off tube
includes a removeable calibration target.
6. The apparatus according to claim 4 wherein said stand off tube
is sized to be of equal or larger diameter relative to the target
tissue size.
7. The apparatus according to claim 1, wherein the illumination
unit comprises an illumination source and an illumination filter
wheel.
8. The apparatus according to claim 1, wherein the illumination
unit further comprises a mask that provides for selective
illumination of the target tissue.
9. The apparatus according to claim 1, wherein the detection unit
comprises a collection filter wheel coupled to the pathway that
couples the base unit and the tissue interface unit.
10. The apparatus according to claim 1, wherein the detection unit
further comprises a spectrograph coupled to the collection filter
wheel.
11. The apparatus according to claim 1, wherein the detection unit
further comprises a reimaging device coupled to the collection
filter wheel and the spectrograph that reimages the collected
optical energy prior to the collected optical energy entering the
spectrograph.
12. A method of detecting movement during a measurement of tissue
characteristics, comprising the steps of: a) forming a first image
of the target tissue; b) illuminating a target tissue with optical
energy; c) performing spectroscopic measurements on optical energy
received from the target tissue from illumination; d) forming a
second image of the target tissue, e) comparing the first and
second images to determine the degree of movement of the tissue, if
any; f)comparing the degree of movement to a predetermined
standard; g) if the movement is less than said standard,
determining tissue characteristics of the target tissue based on
the results of the spectroscopic measurements.
13. The method of claim 12, wherein the step of performing
spectroscopic measurements includes illuminating a first portion of
a target tissue with optical energy from a first illumination
source and illuminating a second portion of the target tissue with
optical energy comprises illuminating a second portion of the
target tissue with optical energy with a second illumination
source.
14. A method of determining locating and differentiating normal and
abnormal cervical tissue by optical interrogation of target tissue
comprising the steps of: a) illuminating a portion of the target
tissue along a band of spaced apart, generally horizontal line of
points and measuring optical energy received from said points, b)
shifting said line of points generally vertically, illuminating a
new portion of the target tissue, vertically offset from the
previous illumination, and measuring optical energy received from
said points, to create a matrix of measured points c) repeating the
step of shifting generally vertically until at least a portion of
the cervix has been interrogated; d) determining the presence of
abnormal tissue by comparing adjacent measured horizontal and
vertical points based on the assumption that abnormal tissue is
more likely to spread vertically than horizontally.
15. A method of detecting tissue characteristics, comprising:
dividing an area of target tissue into a plurality of detection
points arranged in columns; illuminating the plurality of detection
points one column at a time; performing spectroscopic measurements
on optical energy received from the target tissue; and determining
tissue characteristics of the target tissue based on the results of
the spectroscopic measurements including the step detecting tissue
characteristics using the rule that the adjacent horizontal points
are more likely to be the boundary between normal and abnormal
tissue than adjacent vertical points.
16. The method of claim 15, wherein the plurality of detection
points are separated from each other by approximately 3 mm.
17. The method of claim 15, wherein the plurality of detection
points are illuminated using a probe positioned a predetermined
distance from the target tissue.
Description
CROSS REFERENCE TO RELATED APPLICATIONS
[0001] This application claims the benefit of U.S. Provisional
Application No. 60/737949 filed 18 Nov. 2005, and claims the
benefit of Ser. No. 10/647222 filed 26 Aug. 2003 which is a
continuation in part of Ser. No. 10/611,917, filed 3 Jul. 16, 2003
now U.S. Pat. No. 7,006,220, which is a continuation in part of
Ser. No. 10/603,597, filed 26 Jun. 2003 now U.S. Pat. No.
6,975,899, which is a continuation in part of Ser. No. 10/446,857,
filed 29 May 2003 now U.S. Pat. No. 6,870,620, which is a
continuation in part of Ser. No. 10/337,687, filed 8 Jan. 2003,
which is continuation in part of PCT/US02/06350, which is a
continuation of Ser. No. 09/786,781, filed 9 Mar. 2001, now
abandoned, which claims priority from provisional application Ser.
No. 60/272,458, filed 2 Mar. 2001, which is a continuation of Ser.
No. 09/700,538, filed Nov. 16, 2000 now U.S. Pat. No. 6,590,651,
which is a division of Ser. No. 09/533,817 filed 24 Mar. 2000 now
U.S. Pat. No. 6,577,391, which is a continuation of Ser. No.
09/434, 518 filed 5 Nov. 1999 now abandoned, which is the U.S.
National Stage Entry of PCT/US99/20646 which is the U.S. National
Stage Entry of PCT/US99/10947, which claims priority to provisional
application Ser. No. 60/126,056 filed 23 Mar. 1999 hereby is
incorporated herein by reference in its entirety.
FIELD OF THE INVENTION
[0002] The invention relates to medical diagnostics, more
particularly, identification of normal and diseased tissue.
BACKGROUND
[0003] Cervical Cancer is the second most common type of cancer in
women worldwide and the leading cause of cancer related mortality
in women in developing countries. Early detection and diagnosis can
saves lives and reduce the burden on the national healthcare
system. We have built a non-invasive research-prototype
point-of-care device to detect early cancerous conditions of the
uterine cervix. We have tested this prototype in a multi-center
national study accruing data from 600 women for training our
algorithm. Through conversations with the FDA we now have in place
a pivotal trial protocol for validating our algorithm. We will use
a pre-production version of our device in this pivotal study. The
pre-production device is will be a cost and size reduced, portable,
rugged and user-friendlier device and is intended to be identical
to the device intended for sale without incurring the tooling costs
necessary to enter the production phase.
[0004] Epithelial cancers collectively constitute about 90% of all
cancer occurrences. Common epithelial cancers include skin,
cervical, GI tract, colon and oral cancer. While the technology we
have developed is generic and applicable to any accessible
epithelial cancer, we have chosen cervical cancer as our first
diagnostic target. This is because the cervix is easily accessible
and the pathophysiology of cancer progression in it is well
understood. Cervical cancer is a leading cause of cancer-related
mortality in women in developing countries and the second most
common type of cancer in women worldwide. The American Cancer
Society estimates that there will be 12,200 new cases of invasive
cervical cancer diagnosed in 2003 and about 4,100 women will die
from the disease in the US this year. Worldwide, there are
approximately 500,000 cases of cervical cancer diagnosed annually
and approximately 230,000 deaths per year. Estimates show the
market potential for non-invasive cervical cancer detection to be
at $1.25 billion annually in the US and Europe.
[0005] Cervical cancer screening: The Pap test is currently the
most widely used tool to screen women for cervical cancer or
neoplasia. While its contribution to reducing patient mortality is
widely acknowledged, it is prone to errors from low screening
frequency, insufficient cell sampling, inadequate sample
preparation, lack of exfoliation of abnormal cells, and technician
reading error. The discrimination performance of this test is
therefore limited, resulting in a tradeoff between sensitivity and
specificity as illustrated in a landmark meta-analysis conducted by
Fahey. Current practice sets the sensitivity at 51% in order to
achieve a specificity of 97%. Thus, Pap tests have been used as a
means to `rule out` rather than `rule in` disease. One rationale
behind this is to limit the large number of false positives that
would inadvertently burden downstream health management systems.
While this may be true, this also results in a deference of
diagnosis at an annual cost of nearly $2 billion. Improvements to
the traditional Pap test such as ThinPrep.RTM. are becoming
increasingly popular with physicians. This test contributes to
lower intermediate classifications such as ASCUS and increases the
percentage of LSIL+ patients sent to colposcopy. Using a different
approach, the Digene HPV test used for ASCUS triage appears to be
better than a repeat Pap test in finding patients with CIN3 who are
referred to colposcopy. In a recent study of 8,170 screened women
HPV detected 93.3% of CIN3. However, the sensitivity for CIN2
disease was only 72% so that the overall sensitivity of HPV for
HSIL (CIN 2/3 and higher) was 81.8%. Moreover, a low-test
specificity results in an increase in false positives. Also, while
the FDA has approved computerized aids to Pap test screening such
as AutoPap and Papnet, the evidence regarding the impact of these
technologies on the screening process is not yet available.
[0006] Cervical cancer diagnosis: A positive first or second Pap
test is followed by a colposcopic examination. This involves
visualization of the cervix under low power magnification by a
trained clinician who looks for visual cues attributable to
neoplasia. The clinician then takes a tissue biopsy from that
location. The amount of tissue biopsied varies according to the
extent of the assessed lesion and, in some cases, the entire cervix
is removed in what is known as a Loop Electrosurgical Excision
Procedure (LEEP). A pathologist whose diagnosis is considered the
gold standard examines the biopsy specimens. Since suspect areas
are identified visually, colposcopy requires extensive training,
experience, and a significant effort toward maintenance of
skills.
[0007] A key disadvantage of the current methods is the significant
time delay in obtaining the results. A patient and care provider
must wait 14 weeks for the results of the Pap test. Quite often the
colposcopy, biopsy and histology sequence has to be repeated in
order to localize and diagnose the disease definitively. A point of
care approach in new technology will be a significant
advantage.
[0008] The problem is further compounded by the performance limits
of colposcopy. A meta-analysis of colposcopy summarizing the
results of 9 studies lists the average sensitivity and specificity
at 85% and 69% respectively for separating LSIL and lower (CIN 1
and lower) from HSIL (CIN 2/3 and higher). More recent studies show
a much lower sensitivities of 53% and 56%.
[0009] Therefore, a strong need clearly exists for better
differentiation at any point along the entire screening to
diagnosis path. The low specificity and sensitivity numbers result
in a large number of patients undergoing unnecessary biopsy and/or
a large number of patients with cancer going untreated. Moreover,
the one to four weeks required to obtain a Pap test result or a
histology evaluation results in increased patient anxiety and/or
reduced patient commitment to seeking aggressive treatment. This is
especially problematic in treating patients in developing countries
and with indigent populations in the US and other developed
countries. Given that these are the same populations with the
highest prevalence of cervical disease, a point-of-care approach
would have greater value in overall disease management. Similar
approaches have successfully emerged in other areas of diagnosis
and testing such as `stat` blood gas and blood chemistry analysis
as well as in home immunochemistry assays. In addition there is a
need for a less traumatic diagnostic method.
SUMMARY
[0010] The present invention is directed to improvements on
non-contact methods of diagnosing tissue abnormalities preferably
by optical methods. Reference should be had to the detailed
description and the claims for complete disclosures.
BRIEF DESCRIPTION OF THE FIGURES
[0011] The invention will be described in detail with reference to
the following drawings, in which like reference numerals refer to
like elements, and wherein:
[0012] FIG. 1A is a schematic side view of a tissue interface unit
of a system for determining tissue characteristics according to one
embodiment of the invention;
[0013] FIG. 1B is a schematic diagram of a base unit of a system
for determining tissue characteristics according to one embodiment
of the invention;
[0014] FIG. 1C is a schematic diagram of a tube according to one
embodiment of the invention having a clear annulus at a distal end
thereof;
[0015] FIG. 2 is a front view of an end plate of a body structure
of the tissue interface unit of FIG. 1A;
[0016] FIG. 3 shows an exemplary arrangement of illumination
optical fibers on an end plate of a body structure of a tissue
interface unit according to one embodiment of the invention;
[0017] FIG. 4 shows an exemplary arrangement of bundles of optical
fibers located at one end of an illumination pathway adjacent an
illumination unit according to one embodiment of the invention;
[0018] FIG. 5 shows an exemplary columnar arrangement of
illumination optical fibers on an end plate of a body structure of
a tissue interface unit according to one embodiment of the
invention;
[0019] FIG. 6 is a chart of exemplary spectrographic measurements
to be taken to determine tissue characteristics according to one
embodiment of the invention;
[0020] FIGS. 7A and 7B are schematic drawings of a system for
determining tissue characteristics according to another embodiment
of the invention;
[0021] FIG. 8A is a schematic drawing of a docking unit of a system
for determining tissue characteristics according to another
embodiment of the invention;
[0022] FIG. 8B is a schematic drawing of a system interface and
controller of a system for determining tissue characteristics
according to another embodiment of the invention;
[0023] FIG. 9 is a front view of a tissue interface unit of a
system for determining tissue characteristics according to another
embodiment of the invention;
[0024] FIG. 10 is a side perspective view of the tissue interface
unit of a system for determining tissue characteristics according
to another embodiment of the invention;
[0025] FIG. 11 is a side view of a tissue interface unit of a
system for determining tissue characteristics according to another
embodiment of the invention;
[0026] FIG. 12 is a front perspective view of a tissue interface
unit of a system for determining tissue characteristics according
to another embodiment of the invention;
[0027] FIG. 13 is a schematic drawing showing an exemplary
arrangement of detection points on a subject tissue;
[0028] FIG. 14 is a drawing schematically showing how measurements
of columns of detection points are sequentially taken across a
subject tissue according to one embodiment of the invention;
[0029] FIG. 15 is a drawing schematically showing an exemplary
arrangement of a column of detection points on a CCD camera
according to the invention;
[0030] FIG. 16 is a drawing schematically showing the projection of
an image across a CCD camera according to the invention;
[0031] FIG. 17 is a drawing schematically showing an image of a
left side of a cervix projected onto a CCD camera according to the
invention;
[0032] FIG. 18 is a drawing schematically showing an image of a
right side of a cervix projected onto a CCD camera according to the
invention;
[0033] FIG. 19 is an exemplary arrangement of detection points for
a cervix according to the invention;
[0034] FIG. 20A is a schematic drawing of an illumination or target
end of a fiber optic bundle;
[0035] FIG. 20B is a schematic drawing of a collection end of the
fiber optic bundle of FIG. 20A, illustrating a collection approach
according to the invention;
[0036] FIG. 21 is a schematic drawing of a setup for absolute
calibration of a system embodying the invention;
[0037] FIGS. 22-23 are tables of instrument settings for each of
eight software driven measurements that account for each of eight
column positions on a target;
[0038] FIG. 24 is a table of instrument settings for measurements
made in three sets using a different excitation and emission
wavelength for each set.
[0039] FIG. 25 is a view of an optical system of an alternative
embodiment;
[0040] FIG. 26 is a schematic illustration of on embodiment having
a coherent fiber bundle connecting the base and hand held
units;
[0041] FIG. 27 is a pair of related images, the left being a
schematic view of the points of interrogation and on the right, a
photographic image of a diseased cervix with the point of
interrogation overlayed such as with the composite presentation of
the image and measurement systems herein;
[0042] FIG. 28 shows three block components of the system. The
middle block (Hand Held Unit) may be part of either the left or
right hand block depending on the embodiment;
[0043] FIG. 29 shows a pair of schematic illustrations of the
horizontal scanning of a plurality of vertical interrogation
points;
[0044] FIG. 30 is a schematic drawing of an optical system of a
first embodiment;
[0045] FIG. 31 is a schematic drawing of an optical system of a
second embodiment;
[0046] FIG. 32 is a schematic drawing of an optical system of the
preferred embodiment with the system located in the Hand Held
Unit;
[0047] FIG. 33 is a CAD image of a Hand Held Unit taken from right
and left sides (mirror images);
[0048] FIG. 34 is a CAD representation of a commercial unit of
base, hand held unit and probe;
[0049] FIG. 35 is a schematic representation of the confocual
arrangement of the excitation and collection /camera units;
[0050] FIG. 36 is a schematic view of the system shown in FIG. 35
with the confocual arrangement of excitation and collection units
at 5 degrees from each other;
[0051] FIG. 37 is diagramatic view of an improved embodiment where
the hand held unit contains the mask, motor, illumination optics,
video system, collection optics and detection system.
[0052] Before the present systems, methods and apparatus are
disclosed and described, it is to be understood that the
terminology used herein is for the purpose of describing particular
embodiments only and is not intended to be limiting. It must be
noted that, as used in the specification and the appended claims,
the singular forms "a", "an" and "the" include plural referents
unless the context clearly dictates otherwise.
[0053] Ranges may be expressed herein as from "about" or
"approximately" one particular value and/or to "about" or
"approximately" another particular value. When such a range is
expressed, another embodiment comprises from the one particular
value and/or to the other particular value. Similarly, when values
are expressed as approximations, by use of the antecedent "about,"
it will be understood that the particular value forms another
embodiment.
[0054] Quantitative optical spectroscopy improves on prior art
technology in the four major areas listed below. If realized, these
can solve the majority of problems that exist with cervical cancer
care.
[0055] Various embodiments of the present invention include
systems, methods and apparatus that may be utilized to determine
tissue characteristics by applying and measuring optical energy,
including but not limited to visible, infrared and/or UV light. It
should be understood that the term "illumination" according to the
invention means "to give optical energy to", the term optical
energy again, including but not limited to visible, infrared and/or
UV light.
[0056] In several embodiments, the present invention comprises a
base unit, a tissue interface unit and a pathway that couples the
base unit and the tissue interface unit. In one particular
embodiment, the present invention is comprised of a tissue
interface unit that is optically and electronically coupled to a
base unit, as shown in FIGS. 1A-1B.
[0057] FIG. 1A is a schematic side view drawing of a tissue
interface unit according to an embodiment of the present invention.
The tissue interface unit 70 includes a base structure 80. The base
structure 80 may include a handle 74 attached thereto and
configured to be graspable by a user; however, other configurations
may also be appropriate.
[0058] A tube 72 may be configured to be removably attachable to
the base structure 80. The tube 72 functions as a barrier to
exclude, for example, room light. The tube 72 is not necessarily
tubular or cylindrical in shape; other configurations may also be
appropriate.
[0059] The tube 79 connects to base structure 80 via plate 80b. An
end face 80a of plate 80b is shown in FIG. 2. The end face 80a
contains at least one opening for respective pathways 73a, 73b,
73c, 73d. These pathways are connected to and selectively share the
tube 72 in such a way that no interference occurs between the
respective pathways. For example, illumination pathway 73b delivers
to a subject tissue illumination energy or light received from the
base unit 20 along illumination pathway 44. The collection pathway
73c receives energy or light reflected and/or emitted by a subject
tissue and guides it to collection pathway 60, which guides the
collected light to the base unit 20.
[0060] The tissue interface unit 70 may further include an
illumination source 76 and a second illumination pathway 73d.
Additionally, the tissue interface unit 70 may include an imaging
device 78 and an imaging pathway 73a. The imaging device could take
the form of a digital camera, or a CCD based imaging device,
although other imaging devices could also be used. The second
illumination pathway 73d delivers illumination energy or light from
the illumination source 76 to the subject tissue. This illumination
energy or light is reflected off the subject tissue as image energy
or light. The image energy or light is received into the imaging
pathway 73a where it is directed to an imaging device 78. The
imaging device is then used to provide a user with an image of all
or a portion of the subject tissue.
[0061] The second illumination pathway 73d and the imaging device
7S comprise the imaging channel (not shown). The imaging device 78
allows the user to position the distal end 72a of the tube 72 in
the proper and otherwise desired contact with the tissue and to
verify that such contact has been accomplished. Moreover, the
imaging channel allows the user to acquire a digital or other image
of the tissue with the help of the imaging device 78. This image
can serve as an additional visual tissue diagnosis tool.
[0062] The tissue interface unit 70 may also contain various lens
assemblies (not shown) that direct optical energy from the
illumination pathways 73b, 73d onto the subject tissue, and that
direct energy or light from the subject tissue into the collection
pathway 73c and the imaging pathway 73a. For example, the various
lens assemblies may comprise a set of achromatic lens doublets. The
matched set of achromatic lens doublets may be provided in each
pathway. The doublets are generally those commonly used in the art,
such as a BK.sup.7/SF.sup.2 glass biconvex/planoconcave combination
available off the shelf from Edmund Scientific, OptoSigma and
Melles Griot; although other lenses may also be appropriate. The
material of the lenses may be used to limit irradiation and
collection in the UV to a desired wavelength range, such as for
example a minimum wavelength of approximately 350 nm wavelength
range. According to embodiments of the present invention, the
lenses may provide magnification/demagnification in the
excitation/collection paths, respectively.
[0063] The tube 72 can function to fix the lens assemblies a
predetermined distance from the subject tissue. In addition, if the
subject tissue is surrounded by the end 72a of the tube 72, the
tube can function to exclude ambient optical energy from
illuminating the subject tissue. The tube makes contact with the
tissue setting the focal distance so the tissue to lens distance is
correct.
[0064] The illumination pathway 73b may include, for example, a
custom designed bundle of optical fibers. In one example, 52
optical fibers approximately 2 meters long, having a numerical
aperture (NA) of approximately 0.12, and having a core diameter of
approximately 100 .mu.m is utilized to form the illumination
pathway 73b. This fiber bundle may be only part of the illumination
pathway. The illumination pathway may also have lenses. According
to one embodiment, the tube 72 comprises a clear annulus 72b at a
distal end thereof opposite to an end plate 72c that allows the
tube 72 to be attached, removably according to certain embodiments,
to the base structure 80, as shown in FIG. 1C. Contact of the tube
72 to the surface of the tissue will be visible through the annulus
72b, which provides the user with visual confirmation that the tube
72 is properly positioned.
[0065] One exemplary arrangement of optical fibers is illustrated
in FIG. 3. The tissue end of the optical fiber bundle is held in
the tissue interface unit 70 behind a pair of achromatic lens
doublets (not shown). At the tissue end, the optical fibers 21 are
arranged as shown in FIG. 3, in 8 columns 22. At the opposite end
of the illumination pathway 44, the optical fibers 21 for each
column 22 shown in FIG. 3 may be collected into a separate bundle
23a, as shown in FIG. 4. This means that there will be eight
bundles 23 of optical fibers 21 at the opposite end of the
illumination pathway. When constructed in this manner, if optical
energy is fed into a single bundle 23a at a time, a single column
22 of optical fibers 21 will illuminate the target tissue, as
discussed below in detail.
[0066] The collection pathway 73c may be, for example, another
custom designed coherent bundle of optical fibers. In one example,
several thousand, e.g. 5000, optical fibers that were approximately
2 meters long, having a NA of approximately 0.12, and having a core
diameter of approximately 50 .mu.m were arranged in an
approximately 5-mm diameter aperture, in a coherent fashion, to
provide a one to one image transfer from the tissue interface unit
to the detection sub-unit of the base unit. The tissue end of the
bundle of optical fibers is held in the tissue interface unit
behind a pair of achromatic lens doublets. Since one column of
spots is illuminated on the tissue, for example, cervix, at a time,
as is later discussed, returned radiation from the same column is
transferred by the coherent bundle to the detection sub-unit of the
base unit. This returned radiation, which will be arranged in a
column of spots, acts as a virtual vertical slit that is then
spectrally resolved in the horizontal dimension by the detection
sub-unit of the base unit, as is later discussed.
[0067] As mentioned above, some embodiments of the device may
include an illumination device 76 and image detector 78. Together,
these items allow the device operator to obtain a real-time image
of the target tissue, which can help to properly orient the tissue
interface unit with respect to the target tissue. These items are
not required in all embodiment of the invention, and could be
completely eliminated. In other embodiments of the invention, these
items could be replaced with a sighting mechanism which simply
allows the device operator to look down the tube 72 to view the
target tissue.
[0068] In embodiments of the invention that include an illumination
device 76 and imaging device 78, the illumination source 76 may be,
for example, a 4.25V or 2 W halogen lamp manufactured by Welch
Allyn, Inc. in Skeneateles, N.Y. This exemplary lamp has an
integrated parabolic reflector that projects the optical energy
onto the tissue and provides a uniform illumination on the tissue.
The imaging device 78 may be, for example, a 1/4'' format Panasonic
color board camera with 480 horizontal TV lines. This camera has a
C mount adaptor, into which a focusing lens doublet may be mounted.
The camera may be mounted offset from the illumination and
collection pathways due to space constraints, and the image
transfer accomplished using a pair of reflectors 78a.
[0069] The tissue interface unit may be designed in conjunction
with a vaginal speculum configured for insertion into a patient's
vagina during the examination procedure. The unit is held fixed
with respect to the vaginal speculum (not shown) according to
certain embodiments. However, according to other embodiments, the
unit may be used without such a speculum.
[0070] Prior to conducting tissue measurements, some embodiments of
the instrument may be calibrated by malting one or more
measurements on a disposable calibration target 78a that mounts on
the distal tissue end of the tube 72. This disposable calibration
target could be used to take a reference or a calibration
measurement, or possibly both. Moreover, in various embodiments,
these measurements may be a reflectance and/or fluorescent
measurements.
[0071] FIG. 1B is a schematic diagram of a base unit according to
one embodiment of the invention. The base unit 20 according to the
invention is small enough to be portable or mobile. For example,
the base unit 20 could be provided on a movable cart (not
shown).
[0072] The base unit 20 comprises an illumination sub-unit 30, a
detection sub-unit 50 and a control sub-unit 45. The illumination
sub-unit 30 includes an illumination source 32. For example, the
illumination source may be a 175 W short arc Xe lamp provided with
an integrated parabolic reflector, which produces a near collimated
beam. Such a lamp is manufactured by ORC lighting products, a
division of PerkinElmer Optoelectronics (Azusa, Calif.). Other
lamps may also be appropriate. In addition, the illumination source
32 could also take the form of one or more lasers or LEDs. The
illumination source 32 may be housed inside a fan cooled heat sink
assembly (not shown) to limit dissipation of heat to the
illumination sub-unit's other components.
[0073] Optically coupled to the illumination source 32 is an
illumination filter wheel 38. The illumination filter wheel 38
provides for selective wavelength filtering and may be motorized.
For example, the illumination filter wheel may be an eight-position
filter wheel manufactured by ISI Systems (Santa Barbara, Calif.).
An example of filters that could be used in one embodiment of the
invention are listed in FIG. 6. The illumination filter wheel is
mounted within the illumination sub-unit 30, as shown in FIG. 5B,
and the control unit 45 selects the appropriate filter to be
brought into the light path.
[0074] A cold mirror 34 may be provided between the illumination
source 32 and the illumination filter wheel 38. In another
embodiment of the invention, an IR absorbing glass/filter may be
used instead of a cold mirror. A near collimated light beam from
the illumination source 32 is directed through the filter. For
example, in one embodiment of the invention, Applicants utilized a
KGI glass filter available off the shelf from Melles Griot. The
filter transmitted wavelengths in the range of approximately
340-700 nm. Because of its high absorption of IR wavelengths, the
filter helps protect downstream components from excessive heat and
also minimizes stray light in the detection sub-unit.
[0075] The illumination sub-unit 30 may also include a safety
shutter 36, in particular where a continuously operating
illumination source is utilized. In such a case, illumination would
only be allowed into the unit and through to the tissue for the
duration of the spectroscopic measurements, even though the
illumination source would be continuously operating. Software in
the control unit 45 would control actuation of the normally closed
shutter.
[0076] The illumination sub-unit 30 may also include a focusing
lens 40, for example, a single approximately 28 mm diameter,
approximately 100 mm FL, plano-convex lens. The focusing lens 40
focuses the illumination optical energy or light onto the
illumination pathway 44.
[0077] A mask 42, motorized using an encoded stepper motor (not
shown) and controlled by the control sub-unit 45, may be provided
at an entrance to the illumination pathway 44. The mask 42 is used
to control the optical energy so that the optical energy will only
pass into certain portions of the illumination pathway, for
example, into certain ones of the optical fibers, at any given
time. The mask 42 blocks the illumination optical energy from
entering the remaining portions of the illumination pathway, for
example, certain remaining optical fibers.
[0078] By way of an example, one embodiment of the illumination
sub-unit end 23 of the illumination pathway 44 is shown in FIG. 4.
As previously discussed, it has a collection of eight bundles 23a
of optical fibers, where the optical fibers in each bundle 23a
corresponding to different respective columns 22 of individual
optical fibers at the tissue end 80a of the illumination pathway
44, as shown in FIG. 5. Thus, the optical fibers in bundle number 1
at the illumination sub-unit end of the illumination pathway 44, as
shown in FIG. 4, correspond to the optical fibers 21 arranged in
column 1 of the tissue end 80a, as shown in FIG. 5.
[0079] The mask 42 has a single hole (not shown) that can be
selectively aligned with only a single bundle 23a of the optical
fibers shown in FIG. 4. The control sub-unit 45 will control
movement of the mask 42 so that each bundle 23, in turn, is
illuminated. This will cause the illumination optical energy to be
emitted from one of the columns 22 shown in FIG. 5, and as the mask
42 moves, different ones of the columns 22 of optical fibers will
illuminate the target tissue.
[0080] The detection sub-unit 50 may comprise a re-imaging device
52, a collection filter wheel 54, a spectrograph 56 and a CCD
camera 58. The collection filter wheel 54 is optically coupled to
the spectrograph and holds a plurality of filters (not shown) for
filtering the collected optical energy before it is sent into the
spectrograph 56. Exemplary filters for multiple spectral
measurements are listed in FIG. 6. Filtering can be used to reduce
artifacts due to reflected excitation from the target tissue. When
attempting to measure fluorescent emissions from the target tissue,
which have a very low amplitude, a reduction in the reflected
excitation energy or light amount is quite helpful. The insertion
of filters, however, can change the light path between the optical
fibers and the spectrograph entrance slit.
[0081] The collected optical energy, which has traveled through,
for example, optical fibers to the detection sub-unit, is re-imaged
at the entrance slit of the spectrograph 56 by a re-imaging device
52, such as, for example, an FC446-30 from Roper Scientific-Acton
Research (Acton Mass.), which does this without introducing
chromatic aberrations and astigmatism. Such a re-imaging device may
include a spacer (not shown) which allows insertion of a motorized
collection filter wheel, such as, for example, an FA-448-2 filter
wheel also from Roper Scientific-Acton Research. The re-imaging
device permits simple, straightforward insertion of the filter
wheel.
[0082] The collection pathway 60, which may be, for example, a
coherent bundle of optical fibers 60, which carries optical energy
collected from the target tissue, is placed at the entrance of the
re-imaging assembly. At any given time, the illumination pathway
44, which may be, for example, optical fibers, will only illuminate
a column of positions on the target tissue. Thus, optical energy
collected into the collection pathway 60 will only be from
approximately the same column of positions on the target tissue.
The result is that, at any given time, the optical energy entering
the spectrograph 56 from the return optical fibers 60, will be
arranged in a virtual vertical slit.
[0083] The spectrograph 56 takes the vertical slit of returned
optical energy, and resolves the optical energy into different
wavelengths by separating the energy or light in the horizontal
direction. The result is an energy or light pattern having two
dimensions, wherein the vertical dimension corresponds to different
positions on the target tissue, and wherein the horizontal
dimension corresponds to different wavelengths. The two dimensional
energy or light pattern is then recorded on a camera, for example,
a CCD camera 58.
[0084] The spectrograph may be, for example, a customized,
approximately 300 mm focal length, f#4, Czerny-Turner configuration
spectrograph, such as the SpectraPro SP-3061, manufactured by Roper
Scientific-Acton Research (Acton Mass.). According to one
embodiment of the invention, the grating of the spectrograph has
the following specifications:
Grooves/mm: 100 nm/mm
Dispersion: 32 nm/mm
Blaze angle: 1 17'
[0085] Field of view: 365 nm
[0086] The camera may be a CCD camera, for example, a
thermoelectrically cooled CCD camera, such as the NTE/CCD-512SB
manufactured by Roper Scientific-Princeton Instruments (Princeton,
N.J.) with a SITE 512.times.512, square format, approximately 24 m
pixel, back illuminated detector, along with the ST-133 high speed
DMA serial interface controller. The A/D converter in the
controller allows a 1.0 MHZ A/D scan rate. However, other types of
cameras commonly known to those skilled in the art may be used.
[0087] In one embodiment, the control unit 45 is a
software/hardware package comprised of an instrument control
section, a graphical user interface, and data storage capabilities.
For example, a compact PC with adequate ports and bays to
accommodate the requisite interfaces and PCI cards may be used for
this purpose. The control unit 45 provides control over actuation
of the illumination and collection filter wheels 38, 54, the safety
shutter 36, the camera shutter (not shown), the camera controller
(not shown), data conversion and transfer to the PC (not shown),
the spectrograph grating adjust motor (not shown), the imaging
camera 68 and corresponding illumination source 32 and the stepper
motor (not shown) for the motorized mask. Control is provided
according to a schedule template that can be modified by the
user.
[0088] In addition, the software may provide graphical feedback to
the user showing images (video and spectroscopy) that are used to
make real time determinations of measurement adequacy. The program
stores the measured data, which may include tissue particulars,
measurement particulars and/or images. The measured data for each
tissue can then be downloaded and stored in a portable recording
medium (not shown) such as a magnetic or optical disk.
[0089] The above described embodiment, which includes a
spectrograph for spectrally resolving the light returning from the
target tissue, is but one way to accomplish the spectral
resolution. In other embodiments of the invention, other devices
such as prisms or transmissive gratings, for example, could be
utilized to spectrally resolve light returning from a location on a
target tissue into different wavelengths. Yet, even further, other
devices known to those of ordinary skill in the art could be
utilized. For purposes of discussion and example only, a
spectrograph will be discussed as the spectrally resolution
device.
[0090] In addition, in some embodiments of the device, it may prove
more advantageous to take measurements at a plurality of locations
on a target tissue to measure a single narrow wavelength band of
returned light during a first measurement cycle. Another
measurement cycle could then be conducted at the same locations on
the target tissue for one or more different wavelength bands.
[0091] Furthermore, in the embodiment described above, the
illumination light was conveyed to the target tissue such that it
sequentially illuminated several different columns of positions on
the target tissue. In other embodiments of the invention, the
illuminated positions on the target tissue need not be illuminated
in a column arrangement. In fact, it some embodiments of the
invention) it may be advantageous to arrange the optical fibers
such that each sequential illumination and measurement cycle
measures the characteristics of widely separated locations across
the target tissue. Once all measurements have been taken, the
measurement results could be re-combined by the device operating
software to present an image indicative of the target tissue
characteristics. A device configured in this manner would greatly
reduce the occurrence of cross-talk between illuminated
positions.
[0092] The systems, methods and apparatus according to the present
invention use the hyperspectral imaging approach discussed by J.
Marno in "Hyperspectral imager will view many colors of earth,"
Laser Focus World, August 1996, p. 85. This involves measuring
intensities of optical energy emitted from tissue at high spectral
and spatial resolution.
[0093] Systems, methods and apparatuses embodying the present
invention should be designed to ensure that, as between measurement
speed, spectral resolution, and spatial resolution, the most
important characteristics are measured with the highest resolution
in the shortest possible time period.
[0094] In order to obtain spectra free of environmental or system
artifacts, one approach would be to calibrate the system embodied
by the present invention. The calibration procedures are as
follows: (1) provide an absolute scale to the intensity
measurements at each wavelength; (2) provide an absolute wavelength
scale; (3) correct for fluctuations in lamp intensity and spectral
shifts; (4) correct for spectrograph/grating performance
limitations due to stray light; (5) correct for background light;
(6) correct for noise; and/or (7) correct for variance and temporal
changes in optical properties, spectral transmittance, reflectance
lenses and fibers. Providing an absolute scale to the intensity
measurements at each wavelength calibrates the detection elements
of the system and provides an absolute scale to the intensity
measurements. This will also allow identification of performance
variations in the source and detection system.
[0095] With respect to noise, there exists categories of potential
noise that might typically occur with measurements comes from
several possible sources. Without limitation, they include shot
noise, instrument noise, clinical noise, and physiological
noise.
[0096] Shot noise is equal to I and refers to the inherent natural
variation of the incident photon flux. Photoelectrons collected by
a CCD exhibit a Poisson distribution which have this square root
relationship between signal and noise. Instrument noise includes
several individual noise types classified according to their
sources such as CCD noise including the read noise and dark noise
and dependent on the A/D transfer rates and the temperature of the
CCD, respectively. Additional sources of instrument noise include,
without limitation, variability in lamp intensity, variability in
the transmittance of optical components such as fibers filters and
lenses, and variability in the transmittance of fibers due to fiber
bending.
[0097] Clinical noise is the noise that arises from the clinical
measurement procedure such as the distance/angle between the target
tissue and the device, presence of blood and mucus as well as
patient/device movement.
[0098] Physiological noise is the non-diagnostic natural
variability of the biochemical and morphological properties of
tissue. The physiological noise can be one of the most challenging
to address. To alleviate this noise source is to normalize or
compare the intensities measured at any tissue site with the
intensity from a `clinically normal` site. The normal site is
identified using simple tests such as the maximum or minimum
intensity or intensity ratio.
[0099] The signal to noise ratio of a measuring device is simply
SNR = I .sigma. .function. ( I ) ##EQU1## where I is the measured
signal intensity, and .sigma.(I) is the noise or standard deviation
of the measured intensity. We have taken steps to ensure that
signal corruption in our device from the cumulative effects of
these noise sources is reduced or eliminated. The specific steps
include: [0100] A. Obtaining a high enough signal intensity such
that the noise in the measurement in dominated by the shot noise.
The shot noise is an inherent property of the CCD response and
given that it increases as I with increase in I, its proportion as
a percentage of I decreases with increase in I. At a high value of
I the contribution of shot noise is negligible. We have attempted,
as listed below, to reduce other noise sources to a value below
that of shot noise i.e. the instrument operates in the shot noise
dominated regime; [0101] B. Keeping the temperature and the A/D
transfer rate at the lowest optimum, thus minimizing read and dark
noise; [0102] C. Measuring the lamp power simultaneously with the
tissue measurement. The tissue measurement is then normalized by
this measured lamp intensity. This removes/corrects for the noise
in the measured intensity due to variability in lamp intensity and
variability in the transmittance of optical components such as
fibers, filters and lenses; [0103] D. Using ratios of intensities
at different wavelengths rather than straight intensities since
this method internally corrects for changes in transmittance and
also corrects for variations in light coupling due to changes in
the way the target tissue is oriented with respect to the light
beam. This method is limited to transmittance changes that do not
vary across the spectrum; and [0104] E. Optimizing the clinical
procedure to minimize the clinical noise. This includes an adequate
tissue cleaning procedure and keeping the device weight and shape
conducive to holding it without significant motion artifact.
[0105] Next, the horizontal dimension of the CCD, measured in pixel
number is used to mark the wavelength of the measured intensity. A
wavelength number is assigned to each pixel. Establishing these
absolute scales contribute to the calibration of the present
invention.
[0106] Calibration standards may include those commonly used by
ones skilled in the art. For example, spectral irradiance standards
may utilize a NIST traceable Quartz Tungsten halogen lamp for
wavelengths greater than approximately 400 nm. For wavelengths less
than approximately 400 nm, a NIST traceable Deuterium lamp may be
used. Wavelength calibration standards may include, without
limitation, mercury lamps and NRCC traceable Erbium Oxide lamps.
With respect to the former, these lamps have narrow, discrete
spectral lines over UV and visible wavelengths that provide a
metric for wavelength calibration. For example, for diffuse
reflectance standards, a NIST traceable Spectralon..TM.. from
LabSphere, Inc. (North Sutton, N.H.) may be utilized. The
reflectance of these standards is highly lambertian over their
spectral range. They also have a spectrally flat reflectance
profile, i.e. the percent of radiation reflected at each wavelength
(within the usable wavelength range) is constant. For diffuse
fluorescence standards, ones such as those produced by LabSphere,
Inc. may be used. These standards are also made of Spectralon.TM.
and one further embedded with inorganic fluorophores that provide a
highly stable, reproducible fluorescence.
[0107] In addition to absolute scale, calibration must correct for
variances and potential external and/or internal interferences.
Fluctuations in lamp intensity and spectral shifts may need to be
corrected for, since arc lamps such as the ones used according to
certain embodiments of the present invention are known to display
fluctuations in energy output based on lamp life, duration of use
and ambient conditions. Since the present invention determines
tissue characteristics based on intensity measurements, such
variations should be taken into consideration and accounted for by
appropriate calibration. Similarly, it is helpful to correct for
stray light that may result from the inability of a monochromator
grating to perfectly separate light of different wavelengths.
Grating efficiency, inadequate baffling and the use of short
optical path lengths needed to make a compact instrument all
contribute to stray light and therefore, should also be accounted
for by appropriate calibration. In addition to stray light,
background light may also be a factor to consider. Light leakage
into the system that results in erroneously higher intensities must
be measured and subtracted.
[0108] Finally, in addition to absolute scales and internal and/or
external light factors, calibration of the present invention may
also include accounting for dark noise and variance and temporal
changes in optical properties, spectral transmittance, reflectance
of lenses and fibers. With respect to dark noise, this issue
primarily arises as a result of thermal, non-thermal and readout
noise characteristics of the CCD detector. Although embodiments of
the invention use a PET cooled detector, the noise can be
significant and needs to be subtracted out. With respect to factors
effecting optical properties, spectral transmittance, reflectance
of lenses and fibers, each spot/location of light projected on the
tissue varies in intensity. This variance may be due to the axial
position of the spot and small differences in individual fibers and
mask apertures. The intensities of the spots/locations may change
with time due to changes in alignment and component
degradation.
[0109] The present invention utilizes at least one calibration
during its operation. One type of calibration is before the initial
operation of a device embodying the invention or when the device
needs maintenance and/or repair. This will be referred to as
pre-operative calibration. Pre-operative calibration may comprise
an absolute calibration protocol and a wavelength protocol.
[0110] Absolute calibration applies irradiance standards to
establish performance benchmarks and to provide an absolute scale
to the intensities measured. The irradiance standards allow the
coupling of known intensity levels into fibers or apertures. A
schematic diagram of a setup for performing calibration is shown in
FIG. 21, where 300 designates an aperture mask, 301 designates a
light source, and 302 designates a black absorbing material. The
aperture mask 300 shown may be replaced with an excitation fiber
bundle, where the light is coupled into fibers at one end of the
bundle. The light emerging from the other side of these apertures
or the other end of the fiber bundle can be imaged by the detection
system and the measured intensity calibrated against the known
intensity to arrive at a correction factor which will be further
taught below.
[0111] To calibrate wavelengths, wavelength calibrations are used.
The light source 301, such as a calibrated mercury arc lamp, is
positioned between a focusing lens (not shown) and the mask 300 in
FIG. 21 while the arc lamp is off or the safety shutter is closed
to ensure only illumination from the mercury lamp enters the
system. A reflectance target is held before the sight tube, taking
care to seal off and prevent room light from entering the system.
The columns of illuminated spots are spectrally resolved on the
CCD. The known natural peaks of the mercury spectrum, when the
embodiment is a mercury lamp, are captured and are used for
calculating a wavelength scale for each image. A set of eight
software driven measurements that account for each of the eight
column positions on the target, are made as show in the table of
FIG. 22 using the instrument settings indicated for each
measurement.
[0112] In addition to, or alternatively, the present invention can
be calibrated prior to each measurement. This calibration will be
referred to as "operative calibration". This calibration corrects
for both short-term system, intermediate and long-term
fluctuations, such as lamp degradation, for example. The method
that performs this calibration may be embodied in a software
program using the instrument settings listed in the tables of FIGS.
23 and 24.
[0113] The operative calibration comprises a reflectance
calibration, a fluorescence calibration, and a background and dark
noise calibration. The reflectance calibration may, according to
certain embodiments, comprise of positioning the Spectralon.TM.
diffuse reflectance target before the sight tube so as to exclude
room light from the system. A series of measurements given the
instrument settings listed in the table of FIG. 23 are made. During
a fluorescence calibration, the Spectralon.TM. (or other
comparable) fluorescence target is positioned before the sight tube
taking care to exclude room light or other superfluous light from
the system. A series of measurements given the instrument settings
listed in the table of FIG. 24 are made according to one embodiment
of the present invention. According to this embodiment, the
measurements may be made in sets of three where each set may use a
different excitation and emission wavelength selected by choosing a
different filter set.
[0114] Finally, background and dark noise calibration may be
incorporated into the fluorescence and reflectance and calibrations
above as well as into each subject target tissue/area measurement.
According to certain embodiments, the first measurement of each
sequence of 8+1 measurements in the tables of FIGS. 23 and 24 is a
background measurement where the safety shutter is held closed.
This measurement accounts for the error that may be caused due to
room light and/or other electronic noise sources that may result in
the CCD reading an intensity signal. This type of result may be
defined as background noise and is subtracted from each of the
calibration and tissue measurements.
[0115] The data collected from pre-operative and/or operative
calibrations are used to calculate a set of correction factors for
absolute calibration as follows: C(f, .lamda.)=T(f, .lamda.)/M(f,
.lamda.) [0116] where f is the position/spot number or aperture
location in the target area and .lamda. is the wavelength
(.about.400-700 nm). T(f, .lamda.) is the true intensity from the
standard coupled into the aperture at least one wavelength, and M
(f,.lamda.) is the intensity measured by the system from that
aperture at that at least one wavelength. All spectra acquired with
the same detection system can then be multiplied point-for-point by
these correction factors in order to eliminate effects of the
non-uniform response (spectral and spatial) of the detection
system.
[0117] In calibrating wavelengths, the measured spectrum of a
mercury light source contains sharp peaks which correspond to the
spectral lines of the source. The wavelength of each corresponding
spectral line can be assigned to the pixel number along the
horizontal axis of the CCD for each position of the peak. With two
or more peaks present in the spectrum, a linear interpolation is
then used to determine the wavelength values for all the
pixels.
[0118] For operative calibration, the protocol comprises a
reflectance intensity calibration, a fluorescence intensity
calibration, and a stray light or other superfluous light
calibration. Intensity calibration measurements for reflectance
spectra are performed by normalizing the spectrum measured from
each spot on a tissue with the spectrum measured from the same spot
on the reflectance calibration target. This is done after
subtracting the background light from each measurement. This
procedure eliminates any error from spot-to-spot variations in
excitation intensity and can be expressed as follows: R(f,
.lamda.)={[RS(f, .lamda.)-BS(f, .lamda.)]/[-RR(f,.lamda.)-BR(f,
.lamda.)]}=TR(f, .lamda.), [0119] where RS(f,.lamda.) is the
reflected intensity spectrum measured from the subject target area
and RR(f, .lamda.) is the reflected intensity spectrum measured
from a reference whose true reflectance TR(f, .lamda.) is known.
This true reflectance is provided by a diffuse reflectance standard
whose reflectance is substantially constant for all wavelengths
used in the system taught by the present invention.
[0120] BS(f, .lamda.) is the background measurement corresponding
to the tissue reflectance measurement, e.g. tissue background
measurement taken using the same instrument settings as the tissue
measurement, but with the safety shutter closed. BR(f,.lamda.) is
the background measurement corresponding to the reference
measurement. With these measurements, a meaningful estimate of
tissue reflectance R(f, .lamda.) may be obtained.
[0121] For fluorescence spectra, intensity calibration involves
normalizing the fluorescence spectrum from each location on the
target area by the fluorescence intensity from the same location
when measuring on the fluorescence calibration target. Then, either
the integral or the peak of each position's intensity spectrum may
be used to normalize spectrum using the following formula: F(f,
.lamda.)=[FS(f, .lamda.)-BS(f, .lamda.)]/[FR(f,v.)-BR(f, .lamda.)]
[0122] where FS(f, .lamda.) is the fluorescence spectrum measured
on subject target area and BS(f, .lamda.) is the corresponding
background measurement taken using the same instrument settings as
the subject target area measurement but with the safety shutter
closed, FR(f, .lamda.) is the measurement on the fluorescence
reflectance standard, BR(f, .lamda.) is the corresponding
background measurement, and F(f, .lamda.) is the corrected
fluorescence spectrum.
[0123] With respect to stray light or superfluous light
calibration, correcting each fluorescence spectrum for the stray
light output of the excitation monochromator involves subtracting
the stray light spectrum reflected from the tissue from the
measured fluorescence spectrum of the tissue. This correction
employs the principle that the absolute reflectance (as a function
of wavelength) is independent of the spectrum used for
illumination. This principle can be expressed as an extension of
the immediately preceding equation as follows: {RS[I.sub.1(f,
.lamda.)]-BS[I.sub.1(f,
.lamda.)]}/{RR[I.sub.1(f,.lamda.)]-BR[I.sub.1(f,
.lamda.)]}={R.sub.S[I.sub.2(f,
.lamda.)]}/{RR[I.sub.2(f,.lamda.)]-B.sub.R[I.sub.2(f,
.lamda.)]}.
[0124] Here, I.sub.1 is the standard, broadband output of the
illumination system used to measure reflectance of tissue, for
example, and I.sub.2 is the stray/superfluous light of the
illumination system that accompanies the monochromatic excitation
used for tissue fluorescence measurements. Thus, tissue calibration
may be achieved by normalizing this procedure with the standard
reflectance. The result is a calibration factor, as follows:
{RS[I.sub.1(f,
.lamda.)-BS[I.sub.1(f,.lamda.)]]}/{RR[I.sub.1(f,.lamda.)-BR[I.sub.1(f,.la-
mda.)]]} which when multiplied by the stray/superfluous light
spectrum measured on the standard from supposedly monochromatic
excitation gives the stray/superfluous light inadvertently measured
along with tissue fluorescence. This is illustrated by rearranging
the equation such that: RS[I.sub.1(f, .lamda.)]=({RS[I.sub.1(f,
.lamda.)-BS[I.sub.1(f, .lamda.)]]}/{RR[I.sub.1(f,
.lamda..)-BR[I.sub.2(f, .lamda.)]]})={RR[I.sub.2(f,
.lamda.)]-BR[I.sub.2(f, .lamda.)]]}.
[0125] RR[I.sub.2(f,.lamda.)] and the corresponding
BR[I.sub.2(f,.lamda.)] are measured in a similar way as discussed
in the previous section for intensity calibration of fluorescence
spectra. The reflectance standard is illuminated with monochromatic
light (and associated stray light), and the measurement focuses on
wavelengths at which stray light is present (i.e. longer than the
excitation wavelength) rather than the excitation bandwidth.
RR[I.sub.2(f,.lamda.)] is then subtracted from the measured
fluorescence spectrum.
[0126] After the present invention has been calibrated, the tube 72
of the tissue interface unit 70 may be first inserted into the
patient's vagina so that the end of the tube is immediately
adjacent, or covering the patient's cervix. The cervix is then
illuminated by the illumination source 76. Collected optical energy
transmitted and/or reflected from the tissue is directed to the
imaging device 78, which, in this embodiment, is located in the
tissue interface unit 70. The imaging device 78 sends a video
signal that is viewed with a computer or video monitor (not shown).
Thus, the imaging device 78 provides the user with a view of the
patient's cervix, which assists the physician in properly aligning
and situating the tube 72 with respect to the patient's cervix. The
imaging device 78 may also be used to capture still images of the
cervix, which may be digitally stored and used for later data
analysis.
[0127] The tube 72 is appropriately placed such that a good view of
the subject target area can be seen through the imaging device 78,
the tissue interface unit is fixed in place relative to the subject
target area. At this point, a still picture of the subject target
area may be taken with the imaging device. The illumination device
76 is then turned off, and the spectroscopic measurements are
started. As described above, a series of measurement cycles would
be conducted. During each measurement cycle, a column of positions
on the subject target area would be illuminated, and the light
returning from the subject target area would be detected by the
detection sub-unit. During each measurement cycle, the spectrograph
would spectrally resolve the column of positions into a
two-dimensional image that is captured by the camera 58. Each two
dimensional image would be arranged such that one axis is
indicative tissue position, and the other perpendicular axis would
be indicative of wavelength. The two dimensional images recorded
during the measurement cycles would then be recorded and analyzed
by the device operating software in the control sub-unit 45.
[0128] FIGS. 7A and 7B are schematic drawings of a system for
determining tissue characteristics according to another embodiment
of the invention. The system 110 includes a tissue interface unit
170, which may be configured as a handheld probe-type unit, and a
docking unit 120. The tissue interface unit 170 and the docking
unit. 120 communicate with each other via communication pathway
177, which may comprise one or more optical fibers or other type of
signal cable.
[0129] The docking unit 120 may include a stand or cradle 119 for
docking or holding the tissue interface unit 170 when not in use.
The docking unit 120 may also include one or more pathways 182 for
outputting or receiving signals to or from additional system
components, such as a image recording device, such as a VCR or
other type image recording device 183 or monitor 184, such as a
color TV monitor (shown in FIG. 7B).
[0130] As shown in FIG. 5A, the docking unit 120 may further
include a processor 190, a power supply 191, an illumination source
132 and an illumination source controller 132a. The docking unit
120 may also include a light guide (not shown), such as a liquid
light guide that guides optical energy from the illumination
source, for example, into an optical fiber or other type cable to
be delivered to the tissue interface unit 170.
[0131] As shown in FIG. 7B, the tissue interface unit 170 includes
illumination pathways 173b, 173d, which may comprise a single or
multiple pathways. These pathways 173b, 173d may include one or
more light guides 130 that receive optical energy from the
illumination source 132 disposed in the docking unit via
communication pathway 177b. The tissue interface unit may also
include an illumination lens assembly 131 and an illumination
aperture/filter 131a that provides for selective wavelength
filtering and a shutter function, also shown in FIG. 7B.
[0132] The tissue interface unit 170 further includes a collection
pathway 173c, which guides optical energy reflected and/or emitted
by a subject tissue to a device for making spectroscopic
measurements 175. The device for making spectroscopic measurements
175 may include a diffraction grating 157, a camera 158, and camera
controller 158a. The camera and camera controller may be, for
example, a CCD camera controlled by a CCD camera circuit card
assembly. The spectroscopic measurements may be sent to the
processor 190 disposed within the docking unit 120 via
communication pathway 177a for processing. According to one
embodiment of the invention, the system is capable of detecting
reflectance information between approximately 360 nm and 660 nm at
a resolution and fluorescence information at 2 or 3 wavelength
bands. According to various embodiments, the resolution and
wavelength bands can range from 2 nm to 30 nm. According to one
embodiment, the resolution is at 20 nm as are the wavelength bands.
Each frame of data is transferred from the tissue interface unit to
the docking unit for processing.
[0133] The collection pathway 173c may include a shutter 156 that
blocks out illumination optical energy when spectroscopic
measurements are not being made, a filter 159 that provides for
selective filtering of wavelengths not of interest, and a
collection lens assembly 155.
[0134] The tissue interface unit 170 may further include an imaging
pathway 173a that guides reflected optical energy to an imaging
device 187. The image pathway 173a may also include a lens assembly
178b. The imaging device 187 comprises, for example, camera 178 and
camera controller 178a. The camera and camera controller may be,
for example, a video camera and video camera controller, or any
similar type image recording device. The video imaging channel
according to one embodiment may have a resolution of 300 TV lines
(NTSC analog output for video recording and display) with fixed
magnification and focus, a field of view of approximately 25 mm,
and a depth of field of approximately +/-5 mm. The imaging device
187 allows a user to view the subject tissue in order to position
the tissue interface unit 170 with respect to the subject tissue.
The tissue interface unit 170 may include a monitor, or may
communicate with a separate monitoring device to permit viewing of
the tissue by a user. Additionally, the tissue interface unit 170
may include a user interface (not shown) that provides for entry of
patient information, for example.
[0135] The tissue interface unit further includes a power monitor
199 and a system interface and controller 195, as shown in FIG. 7B.
As shown in FIG. 8B, according to one embodiment of the invention,
the system interface and controller 195 includes a data interface
unit 503 that controls the exchange of data signals between the
tissue interface unit 170 and the docking unit 120. The system
interface and controller 195 may further include a discrete
interface unit 504 that controls the system's respective power and
switches, and an analog interface unit 505 that controls the
systems interface with an external image recording device 183. The
system interface and controller 195 may also include a shutter
controller 502 that controls operation of shutter 159 and an
illumination aperture/filter controller 501 that controls operation
of the motor of the illumination filter.
[0136] An example of one embodiment of a hand-held tissue interface
unit according to the invention is shown in side view in FIG. 11.
The tissue interface unit 170 includes housing 186, a handle 174
configured to be graspable by a user, a tube 172 that delivers
illumination optical energy to a subject tissue, and optical energy
reflected and/or emitted by the subject tissue to the viewing
device and/or the spectroscopic measurement device, and a liquid
light guide 130 that guides optical energy received from a docking
unit 120 into the tube 172. The tube 172 may be removable, as
discussed below, and may be disposable. As shown in FIG. 12, the
tissue interface unit 170 may also include a heat sink 199 that
maintains the tissue interface unit within an acceptable
temperature range.
[0137] FIG. 9 is a front view of a tissue interface unit according
to the invention without outer casing 186, and handle 174. FIG. 10
is a side perspective view of the tissue interface unit of FIG. 9
without outer casing 186 and tube 172. As shown in FIGS. 9 and 10,
the tube 172 connects to the base structure 180 via a plate 180b.
The plate 180b has an endface 180a. The endface 180a includes
openings for illumination pathways 173b, 173d, collection pathway
173c and imaging pathway 173a. These pathways share tube 172 in
such a way that no interference occurs between pathways. Tube 172
may be attached to endface 180a by some type of attachment means
180c, as shown in FIG. 9.
[0138] In an embodiment of the invention configured to detect
tissue characteristics of a patient's cervix, the tube 172 of the
tissue interface unit 170 may be first inserted into the patient's
vagina so that the end of the tube is immediately adjacent,
circumscribing or covering the patient's cervix. The cervix is then
illuminated by the illumination source 132 via illumination pathway
173d. Collected optical energy transmitted and/or reflected from
the tissue is directed to the imaging device 187, which is located
in the tissue interface unit 170. The imaging device 187 sends a
video signal that is viewed with a computer or video monitor. Thus,
the imaging device 187 provides the user with a view of the
patient's cervix, which assists the physician in properly aligning
and situating the tube 172 with respect to the patient's cervix.
The imaging device 187 may also be used to capture still images of
the cervix, which may be digitally stored and used for later data
analysis.
[0139] Once the tube 172 is appropriately placed such that a good
view of the cervix can be seen through the imaging device 187, the
tissue interface unit would be fixed with respect to the patient's
cervix. At this point, a still picture of the cervix may be taken
with the imaging device. The image signal is output to the docking
unit or directly to a monitor provided within the tissue interface
unit, or as a separate component. For example, the image, along
with relevant text, could be displayed on a hand-held LCD unit or a
LCD unit attached to the tissue interface unit. The spectroscopic
measurements are then started. The spectroscopic measurement
results are sent to the processor 190 in the docking unit 120 for
processing. For example, the results can be utilized to categorize
the spectroscopic measurement data, and thus the subject tissue, as
"Normal", "Non-Dysplastic", "Low Grade SIL", and "High Grade
SIL."
[0140] The systems, methods and apparatus of the present invention,
may conduct both fluorescence and reflectance spectroscopy using
both visible and UV light or any combination thereof. This is
generally referred to as multimodal spectroscopy. Cervical cancer,
being a form of epithelial dysplasia, provides an ideal target for
diagnosis using the epithelium down to the germinative layer, since
it undergoes minimum absorption and scattering from non-specific
interactions and obtains the largest possible diagnostic
information on its biochemical and morphological state. Other areas
with similar qualities that may serve as comparable targets for
diagnosis include, without limitation, oral cancer and colon
cancer.
[0141] Fluorescence and reflectance spectra may be made at several
locations on the target area by the present invention. Such
locations may be equispaced. Obtaining measurements across the
entire target area, for example, may allow for differential
diagnosis between dysplasia and surrounding tissue depending on the
embodiment.
[0142] Many investigators have pointed to the large biological
variation in the spectroscopic signature of normal tissue. This
natural variation is often higher than the variation seen in the
spectroscopic signatures going from normal to dysplasia tissue in
the same patient, for example. One cannot, therefore, assign an
absolute spectral intensity or signature to disease state. Rather,
all measurements must be normalized or baselined to "normal" tissue
in the same patient, and it is this relative measure or change that
has diagnostic relevance. Given our inability to determine "a
priori" the location of abnormal and normal tissue with certainty,
the logical alternative is to measure substantially the entire
target area.
[0143] A reflectance measurement is made by measuring the intensity
of light returned from the tissue at the same wavelength as that
used to irradiate the tissue. Reflectance measures the
morphological changes associated with dysplasia progression.
Although biochemical changes precede the morphological changes that
occur as a result of the former, in reality, varying degrees of
morphological change accompany the biochemical changes.
Morphological changes appear later in the course of dysplasia
progression and are defined as any change in average cell nuclei,
cell size, cell appearance, cell arrangement, and the presence of
non-native cells. In addition, effects of the host response such as
increased perfusion from angiogenesis result in an overall
difference in tissue appearance.
[0144] The morphological changes add more complexity to the
fluorescence measurement by absorbing and scattering both the
excitation and fluorescent light, thereby altering the true
fluorescence signal. Thus, it is difficult to make a fluorescence
measurement that is truly independent of the effects of scattering
and absorption. At the same time, both measurements provide
information that is partially independent of one another.
[0145] In reflectance spectroscopy, the tissue properties of
absorption and scattering dictate the amount of radiation measured
at the detector. For example, the increased vascularization due to
angiogenesis causes increased blood absorption of visible light.
Light propagating through and re-emitted from tissue is also
strongly affected by light scattering interactions. For example,
dysplasia cells have enlarged nuclei and since nuclei have a
different refractive index from that of the cell cytoplasm, they
serve as efficient light scatters. Thus, dysplasia tissue can
display increased light scattering.
[0146] While the absorption and scattering properties of tissue
correlate quantitatively with disease, by knowing the absorption
and scattering at each site on the tissue the corresponding error
that these effects produce in the fluorescence yields can also be
corrected for. This is the crux of the multimodal spectroscopy
approach. In order to reap this advantage, both measurements must
be made on the same site at the same time so as to ensure nearly
identical conditions.
[0147] The use of near LTV and UV wavelengths elicits the
fluorescence and reflectance response of intrinsic markers shown to
be highly indicative of biological and morphological changes caused
by pre-dysplastic conditions in tissue. Accordingly, the systems,
methods and apparatus according to the invention may be configured
to acquire broad absorption and fluorescent spectra (approximately
340 nm to 700 nm). Particular examples of illumination and
collection wavelengths are shown in FIG. 6. Although these
wavelengths have shown promise, the invention is in no way limited
to the use of these wavelengths.
[0148] The measurements are made from a predetermined standoff
distance from the tissue. In one embodiment constructed by the
inventors to detect abnormalities on cervical tissue, the standoff
distance was set to approximately 175 mm (17.5 cm) to the first
optical surface of the tissue. This standoff distance can be
defined by and maintained by the length of the tube 72, 172 on the
tissue interface unit 70, 170.
[0149] In order to capture high-resolution spectral data from
several locations in a short time (hyperspectral imaging) design
compromises are required. By compromising on the spatial resolution
and measurement time, fluorescence and reflectance spectra can be
captured at approximately 10 nm spectral resolution according to
certain embodiments.
[0150] In one embodiment of the invention used to take measurements
on a subject tissue, for example, a cervix, the system uses a
line-scan approach to collect data from a plurality of detection
points. After positioning, measurements are made at, for example,
52, approximately 0.5-mm circular spots nominally separated from
each other by approximately 3.0 mm, as shown in FIG. 5. The subject
tissue is first flooded with illumination optical energy. Optical
energy returned by the subject tissue is fed to a viewing device,
which provides a user with an image of the tissue so that the user
can appropriately position the system with respect to the subject
tissue. Next, a single line or column of points on the tissue is
illuminated with optical energy. According to one embodiment, the
optical energy is illuminated in a range of approximately 340-700
nm. The radiation/light returned from the target tissue is
collected using a coherent fiber bundle. The result is that the
collected optical energy is formed into a virtual slit at the
entrance of the spectrograph. The spectrograph is then used to
spectrally resolve the optical energy. Given the spectral
resolution required, and the dispersion by the spectrograph, in
this embodiment, a single column is measured at any given time. The
system sequentially scans through all eight columns shown in FIG.
5, acquiring both fluorescence and reflectance spectra in a total
time duration of approximately 2 minutes.
[0151] According to another embodiment of the invention, the system
uses a flood illumination approach. The subject tissue is first
flooded with illumination optical energy. Optical energy returned
by the subject tissue is fed to a viewing device, which provides a
user with an image of the tissue so that the user can appropriately
position the system with respect to the subject tissue. After
positioning, the subject tissue is again flooded with illumination
optical energy, for example, in a range of approximately 340-700
nm.
[0152] The optical energy reflected and/or transmitted with respect
to the subject target area is imaged with the help of a set of
optics onto the face of a fiber bundle (target end) as shown in
FIG. 20A. This end of the fiber bundle has fibers arranged at
discrete points, as shown in FIG. 20A, and the light imaged onto
the bundle at these points is transferred via the fibers to the
other end of the bundle, as shown in FIG. 20B. The other end of the
bundle has all of the fibers arranged in a single column. This
column serves as the entrance slit of the spectrograph, which is
then able to spectrally resolve, in the horizontal direction, the
light in this column.
[0153] In another embodiment, the optical energy is directed to the
subject target area with the help of a set of optics that images a
mask of apertures onto the tissue. This is an alternative
embodiment to those embodiments taught and described in FIGS. 4 and
5. The apertures are arranged in a column on the mask. The mask can
be horizontally moved to scan the entire subject target area while
presenting at least one single column of light at the entrance of
the spectrograph at a given instant. The spectrograph is then able
to spectrally resolve, in the horizontal direction according to an
embodiment, the light collected by this column of apertures.
[0154] The optical energy reflected and/or transmitted by the
subject tissue is then collected and directed to a diffraction
grating, which separates the light spectrally. Wavelengths not of
interest may be filtered out. For example, the illumination
wavelength may be filtered out. The collected light is then
reflected onto a device for making spectrographic measurements,
such as a CCD camera and controller.
[0155] As in the previous embodiments, the spectrograph only makes
measurements at a single column 200 of detection points 210 at a
time on a subject tissue 205, as shown in FIG. 13. According to an
embodiment, reflectance measurements and fluorescence measurements
are made at fifty-six points on the cervix with a separation of
approximately 3 mm. However, depending on the embodiment, the
number of points can vary to any number of possible points at a
separation sufficient to avoid optical cross-talk/interference
among the points. Reference numeral 215 represents a center of the
subject tissue, in the case of a cervix this would be the Os.
Measurements for various columns are then sequentially made, as
shown in FIG. 14.
[0156] FIG. 15 schematically shows what would be recorded by a CCD
camera coupled to the output of a spectrograph. The light returning
from a column of locations on the cervix would be spectrally
resolved into different wavelengths that extend away from the
column in a perpendicular direction. In other words, the pixels of
the CCD camera extending to the left and the right of a single
measurement position would received light of different wavelengths
returned from the measurement position. The intensity of the light
received at each pixel is indicative of the intensity at a
particular range of wavelengths. Thus, examining the values
registered at each pixel on the CCD array allows the device to
determine the intensity of the light returned from each position on
the illuminated column of positions at a plurality of different
wavelengths.
[0157] FIG. 16 schematically shows how a series of measurements
would be taken during different measurement cycles. Each
measurement cycle would provide information about the light
returned from a different column of illuminated positions on the
target tissue.
[0158] Note, the spectrograph would separate the light from each
illuminated measurement position 210 into a +1 Order Spectra and a
-1 Order Spectra. Each Spectra would contain essentially the same
spectral information. Thus, when interrogating a column of
positions 210 on the left side of the cervix, as shown in FIG. 17,
the device could utilize the +1 Order Spectra, which illuminates
pixels within the CCD array. When interrogating a column of
positions 210 on the right side of the cervix, as shown in FIG. 18,
the device could utilize the -1 Order Spectra.
[0159] In cases where the entire spectral bandwidth is not
available in either the +1 or the -1 order spectra, appropriate
wavebands from both orders will be combined to form a complete
spectral set.
[0160] FIG. 19 schematically shows the ultimate arrangement of
detection points 210 collected for an entire cervix using this
system.
[0161] In a further embodiment, an improved non-invasive device and
method are disclosed. Some of the objectives may be:
1. The potential for a truly non-invasive test. Replacing physical
biopsy and histology of tissue with "optical biopsy",
[0162] 2. The potential for providing results at the point of care.
The present invention does not require reading cytology or
histology slides following sample collection from the patient.
Since the tissue in the patient is interpreted using the present
invention algorithm at the point of care, follow-up consultation
with the test results in hand is made possible.
3. Improved detection and diagnosis. There is improved
discrimination using fluorescence and reflectance spectroscopy.
[0163] 4. The test can be performed by a `non-specialist`. The
performance of the present invention will be compared to colposcopy
and biopsy/pathology. The present invention can remove or alleviate
this `subjectivity` if used in an adjunctive or triage mode.
5. Cost Effective Approach. In order to make an impact on cervical
or other disease management, this new technology is economically
viable in order to be accepted by potential users.
[0164] In the preferred embodiment we use both fluorescence and
reflectance (multimodal) spectroscopy with visible and UV light
though they can be used independently. Cervical cancer being an
epithelial cancer provides an ideal target for diagnosis using both
spectroscopic methods. This is because of the short path that light
energy must travel (100 .mu.m-1 mm) to fully penetrate the
epithelium down to the germinative layer. Consequently, light
undergoes minimum absorption and scattering from non-specific
interactions while obtaining information on the biochemical and
morphological state of tissue.
[0165] Both spectroscopic techniques may be used simultaneously in
an imaging mode allowing the entire cervix to be interrogated if
desired. Many investigators have pointed to the large
patient-to-patient variation in the spectroscopic signature of
normal tissue. This patient-to-patient variation is often higher
than the variation between the spectroscopic signatures of normal
and diseased tissue in the same patient. As a result, absolute
intensities are of little value and it is necessary to baseline or
normalize all measurements on a subject to those made on normal
tissue in the same subject. It is this relative measure or change
that has diagnostic relevance. Given the inherent inability to
determine `a priori` the location of abnormal and normal tissue
with certainty, the logical alternative is to measure the entire
cervix. So in short, it is preferably to measure a large portion of
the cervix in order that a base line of normal tissue and be
observed, interrogated and compared to other, potentially diseased
tissue. Identifying healthy baseline tissue is achieved by 1)
recognizing that it is likely that most of the cervix is healthy
and thus by measuring all or large part of the cervix, the majority
of tissue can be assumed to be healthy and 2) we have found that
most abnormalities tend to spread vertically therefore, by scanning
substantially horizontally, on most scans, we are likely to detect
both healthy an diseased tissue in a most passes. This provides for
easier differentiation and eliminates the problems with attempting
to calibrate the system across patients, which is prone to
significant error.
[0166] Possible measurements include: 1) blood profusion
(angiogenesis), 2)epithelial thickening, 3) nuclear site and
content, 4) cell orientation. In the preferred embodiment, the
system (sensor) is non-contact. This has the significant advantage
that the tissue is not disturbed in any way by contact with the
system. Contact can dramatically skew the test results. Further,
the preferred embodiment does not require the use of tissue
preparation such as with acetic acid. Acetic acid pre-treatment of
tissue will enhance detection of nuclear size and content, by
increasing reflectance but tend to suppress all other test
measurement mentioned above.
[0167] Fluorescence measurement. A fluorescence measurement is made
by measuring the intensity of light emitted from the tissue at a
wavelength red-shifted (longer) from that of light used to
irradiate the tissue, and preferably filtering (blocking) the
irradiation light frequencies). Fluorescence measures biochemical
changes, i.e., the earliest changes that occur in the course of
normal cells becoming malignant. The natural fluorophores present
in tissue are the aromatic amino acids tyrosine, phenylalanine and
tryptophan, the metabolites NAD(H), FAD and FAD(H) and structural
proteins collagen and elastin. The fluorescence from these
molecules depends upon their physiochemical environment including
pH, salvation and oxidation state. For example, the reduced form
NAD(H) fluoresces while the oxidized form does not. The reverse is
true for FAD(H). The action of various proteases secreted by tumor
cells on structural proteins renders the fluorophores (tryptophan,
phenylalanine etc.) exposed to a different local environment
(different solvation, viscosity and hydrophobicity), thus altering
their fluorescence.
[0168] Reflectance measurement. This measurement is made by
measuring the intensity of light returned from the tissue at the
same wavelength as that used to irradiate the tissue. Reflectance
measures the morphological changes associated with cancer
progression. Although biochemical changes precede the morphological
changes that occur as a result of the former, varying degrees of
morphological change, in reality, accompany the biological changes.
Morphological changes appear later in the course of tumor
progression and are defined as any change in cell nuclei, cell
size, cell appearance, cell arrangement and the presence of non
native cells. In addition, effects of the host response such as
increased perfusion from angiogenesis result in an overall
difference in tissue appearance. The morphological changes add more
complexity to the fluorescence measurement by scattering and
absorbing both the excitation and fluorescent light thereby
altering the true fluorescence signal. Thus, it is difficult to
make a fluorescence measurement that is truly independent of the
effects of scattering and absorbance.
[0169] Multimodal Spectroscopy. The interactive nature of the
information gathered from fluorescence and reflectance modes makes
it preferable to use both modes to correct for interferences from
one mode to the other. For example, by knowing the absorption and
scattering at each site on the tissue, the corresponding error that
these effects produce in the fluorescence yield can be corrected
for. In addition, as explained earlier, the information content of
each mode is partly exclusive with fluorescence being sensitive to
earlier biochemical changes and reflectance being sensitive to
later morphological changes. Thus by combining the two modes a
better measurement is made. This is the crux of the multimode
spectroscopy advantage. In order to gain this advantage, however,
both measurements must be made on the same site at preferably the
same or nearly same time so as to ensure identical conditions.
[0170] In addition to detection, the present invention may include
a camera (still and or video) and spectrograph together comprise
the detection system.
[0171] This integrated camera-spectrograph is shown in FIG. 25.
Layout of a preferred spectrographic system includes: convex
aberration corrected
[0172] grating and concave mirror. The entrance slit and the
position of the CCD are as shown. (right panel) A CCD camera where
the sensor is placed at the focus of the spectrograph as shown and
analog data is carried to the A/D converter via a cable. The A/D
converter is placed with the CCD preamplifier and clock driver on
one of 4 boards modularized in order to provide flexibility of
placement FIG. 25.
[0173] FIG. 26, this device includes two parts: (a) A hand held
patient interface that is electrically and optically connected to
(b) the electro-optic instrumentation located on a movable cart.
The hand held unit (HHU) looks like a hair dryer and has a
removable snout called the `contact tube`. The contact tube is
designed, in conjunction with a vaginal speculum, for placement in
the patient's vagina during the examination procedure. Prior to
subject measurement the instrument is calibrated by making
measurements on fluorescent and reflective calibration targets.
[0174] FIG. 26 Simplified schematic of the research prototype
showing key component details of a system with a coherent fiber
optic bundle.
[0175] During probe insertion, the subject's cervix is illuminated
by a small lamp and viewed through a video-imaging camera, both of
which are located in the HHU. This camera provides a `video view`
of the subject's cervix on a monitor and assists the physician in
properly aligning and positioning the contact tube and helps
determine if there has been any movement during the test. The
contact tube makes circumferential contact with the periphery of
the cervix. The spectroscopic interrogation, however, is done in a
stand-off manner on the area enclosed by the contact tube. After
the contact tube is satisfactorily positioned, the video camera is
used to capture a still image of the cervix, which is digitally
stored and used for later data analysis. The video lamp is then
automatically turned off and the spectroscopic measurement
started.
[0176] FIG. 27 In one embodiment, illumination fiber bundle design
showing how spots, representing multiple fibers each, at the lamp
end map to corresponding columns at the patient end. Also shown is
a view channel image of the cervix showing overlaid spots where the
cervix is spectroscopically interrogated and the dimensions of the
spots and the pattern. The cervix shown therein is a plastic
medical replica of an average adult female.
[0177] The research prototype spectroscopically interrogates the
cervix in a structured manner from a standoff distance of approx.
176 mm that is maintained by the length of the contact tube.
Measurements are made of 56, 0.5-mm circular spots nominally
separated from each other by 2.75 mm as shown in FIG. 4 and FIG. 5.
We use the line scan approach to gather data from all 56 points. In
this method a line or column of points is illuminated at any given
time and the returned radiation from the tissue is collected using
a coherent fiber bundle. In another embodiment, the coherent light
bundle is dispensed with by moving the sensor to the HHU. This
produces many advantages as will be explained below. The column of
light is transferred through the coherent fiber bundle and acts as
a virtual slit at the entrance of the spectrograph used to
spectrally resolve the light. Given the spectral resolution
required and the light dispersion by the spectrograph, only one
column can be measured at a given time. The system sequentially
scans through all eight rows shown in FIG. 27 and FIG. 28,
acquiring both fluorescence and reflectance spectra in a total time
duration of 4.5 minutes.
[0178] FIG. 28 shows a scanning method by which spots on tissue are
illuminated and imaged (after spectral resolution) onto the CCD.
The first row shows how a total of 56 spots are illuminated in a
sequence of 8 shots. The actual sequence is different from that
shown and is as follows: 4,5,3,6,2,7,1,8 or an inside to outside
horizontal scan pattern. Given that disease occurs primarily in the
center of the field near the os and the squamo-columnar junction
and extends when it does in a top to bottom direction, this scan
pattern ensures that the center of the field is imaged first the
likelihood of simultaneous sampling of both normal and diseased
tissue is maximized. The bottom row shows how light from each row
of spots after spectral decomposition is imaged on a rectangular
portion of the CCD.
[0179] The components of one embodiment of the HHU are described
below. The Contract tube snaps onto the HHU and serves as a light
barrier to exclude room light, a channel for providing an
unobstructed view of the cervix and fixes the object (cervix)
distance from the lens assembly in the HHU to 176 mm, the focal
length of the optics. [0180] a) Video imaging camera. A 1/4''
format color CCD board camera is placed behind a dedicated lens
set. [0181] b) b) Lamp light source for video imaging: A 4.25 W
halogen lamp with an integrated elliptical reflector is used with a
GG295 filter (suppresses any energy less the 295 nm) to provide
uniform illumination on the cervix for video viewing [0182] c)
Illumination fiber bundle. Illustrated in FIG. 27, this custom
bundle accomplishes the line scan approach. It uses 56, 2 meter
long, 0.12 NA, 100 .mu.m core diameter fibers. Fibers from each one
of eight rows at the patient end, maps to one of eight spots at the
lamp end as shown. Lamplight (from the 300 W arc lamp described
below) is sequentially coupled into one of the eight spots of the
lamp end ferrule causing one column at the patient end to light up
at a time. [0183] d) Collection fiber bundle. This is another
custom bundle that uses 0.43 NA, 2 meter long, 10 .mu.m element
fibers arranged in a 6.times.6-mm square aperture in a coherent
fashion to provide a one to one image transfer from the HHU to the
spectrograph. The spectrograph end is rotated 90.degree. so that
each row imaged at the patient end serves as a virtual column or
slit at the entrance to the spectrograph. [0184] e) Lens sets for
excitation and collection. In order to focus the illumination for
spectroscopy on the cervix and for collecting tissue emission a
matched set of achromatic lens doublets is placed in front of the
patient ends of the excitation and collection bundles respectively.
The doublets are BK7/SF2 glass biconvex/planoconcave combinations.
The material choice limits irradiation to greater than 300 nm and
collection to greater than 400 nm.
[0185] Electro-Optic Instrumentation. The electro-optic
instrumentation is located on a movable cart and consists of the
illumination, detection, control instrumentation, user interface
and data storage. The electro-optic instrumentation is further
divided into illumination, detection and control subsystems.
[0186] The illumination subsystem: the following component are
listed in their preferred order of appearance in the light
path.
1. Lamp assembly. This is a 300 W short-arc Xe lamp with an
integrated parabolic reflector, which produces a near collimated
beam.
2. Hot mirror. The near collimated lamp light beam is directed at a
"hot mirror" placed in the beam path. The mirror transmits
wavelengths in the range of 250-700 nm and absorbs/reflects the IR
wavelengths.
[0187] 3. Motorized excitation filter wheel. This eight-position
filter wheel is mounted inside the lamp enclosure as shown in FIG.
26. The filters used in each specific measurement are listed in
Table 2. TABLE-US-00001 TABLE 2 Spectral measurement parameters.
Spectral Measure- # Measurement Excitation Collection Range ment
time R Reflectance OD filter OD filter as 400-700 nm 0.5 secs as
needed needed to F Fluorescence 340 nm 385 nm 450-700 nm 5 secs (40
nm Long-pass F Fluorescence 400 nm 435 nm 500-700 nm 5 secs (30 nm
Long-pass F Fluorescence 460 nm 495 nm 500-700 nm 5 secs (20 nm
Long-pass
4. Motorized safety shutter. Although the lamp operates
continuously, this shutter allows Illumination into the system and
through to the patient only for the duration of the spectroscopic
measurements. 5. Focusing lens. A custom built aspheric lens is
used to focus light into the excitation fiber bundle. 6. Motorized
mask. A custom designed mask that is actuated using an encoded
stepper motor and controller translates the `lamp end ferrule of
the excitation bundle` to position each spot, as illustrated in
FIG. 27, coaxially with respect to the lamp/lens illumination
output.
[0188] The collection subsystem: The components are listed in the
preferred order of appearance in the light path. [0189] 1.
Collection filter wheel and re-imaging assembly. The collection
wheel is populated with filters as listed in Table 2. The
re-imaging assembly re-images the fiber column at the spectrograph
entrance and corrects chromatic aberrations and astigmatism caused
by the filters being present between the spectrograph and the
detection end of the coherent bundle. [0190] 2. Imaging
Spectrograph. The imaging spectrograph has a 300-mm focal length
with a 40-lines/mm plane grating in a CzernyTurner arrangement. The
spectrograph with this grating allows us to capture a spectral
range of 885 nm. [0191] 3. CCD Camera. A thermoelectrically cooled
CCD camera with a SITE 512.times.512, square format, 24 .mu.m
pixel, back illuminated detector along with the ST-133 high speed
DMA serial interface controller. The A/D converter in the
controller allows a 1.0 MHz A/D scan rate.
[0192] Key device features that determine data quality and device
discrimination performance.
[0193] a. Spectroscopic measurements of the cervix while at a
standoff from the cervix. Since the optical properties of tissue as
well as the efficiency of light coupling into tissue are altered,
standoff and contact measurements are qualitatively different.
[0194] b. Spectroscopic interrogation of 56 points on the cervix of
spot size and spacing as shown in FIG. 27. Cross talk between spots
illuminated simultaneously as well as the spatial resolution of the
measurement is determined by this factor, which must be
preserved.
[0195] c. Multimodal spectral measurement (1 reflectance and 3
fluorescence) at 10 nm spectral resolution as listed in Table 2. A
key finding was that 10 nm was the optimal spectral resolution for
best device performance. In addition spectral measurement
parameters are as shown in Table 2 with the exception of
measurement times as explained in item `g` below.
d. A spectroscopy light source with an excitation spectrum of a
typical Xenon arc lamp. This is necessary in order to preserve the
relative intensities at each wavelength as well as use certain lamp
spectral peaks for calibration purposes.
[0196] e. Include a video imaging channel. This is to facilitate
proper positioning of the cervix in preparation for spectroscopy as
well as to provide a color picture of the portion of the cervix
that is the same as the portion measured spectroscopically. In
addition the video imaging channel to be co-aligned with
spectroscopy imaging channel (common image plane). The video/still
image can be aligned and overlayed on the spectrographic analysis
so that the user can identify visually where the system believes
the abnormal tissue is located without further intervention.
Furthermore, the system can take "before" and "after" still images
and compare to insure that movement during the test was not so
great as to compromise the test results. This could occur, for
example if the contact tube is found to have moved from its initial
position to its final position in the before and after images. This
would suggest that during the tissue examination by the system, the
tissue or system had moved.
[0197] f. Nominal and maximum power per spot. The nominal power per
spot on tissue is listed in Table 3. The maximum exposure time can
be increased 10.times., across the board, without exceeding the
safety thresholds specified by the American Conference of
Governmental Industrial Hygienists (ACGIH). This provided the same
exposure times listed are used. A shorter exposure time may permit
higher power levels and will be prorated accordingly.
[0198] Table 3. Nominal power per spot measured on tissue for each
excitation mode. The power should not drop to below 50% of the
stated nominal power over the life of the device. The power levels
indicated are irradiated over the integration times shown.
TABLE-US-00002 TABLE 3 e Nominal Power on Cervix 7.3 .mu.W 24.6
.mu.W 28.1 .mu.W 98 .mu.W Exposure time 5 secs 5 secs 5 secs 0.5
secs
[0199] Measurement time. The measurement time for the research
prototype is 4.5 minutes. This time is about equally distributed
between actual CCD image acquisition
[0200] g. time and time spent in moving stepper motors in
preparation for the next measurement. While some reduction in the
latter can be achieved by more efficient motion and control, the
CCD integration time (measurement time in Table 3) as well as the
digitization/data transfer time must also be reduced. We are
capturing at least 36 images (8 mask positions.times.4 modes+4 dark
images) and any reduction in image integration and transfer time
would therefore be significant. We will have reduced image transfer
time from 500 ms currently to 200 ms by using a faster A/D and
camera interface in the JY camera. Integration time can be reduced
by improving system throughput, a goal that is addressed for all
components affecting throughput in this grant application.
[0201] h. Instrument Signal to Noise Ratio (SNR): The SNR is a
performance metric that is partially affected by system throughput.
An increase in throughput increases SNR. Given that our device is a
multichannel instrument (measuring multiple spatial points
simultaneously) the SNR is divided into two components as shown.
The requirements shown are based on average results from multiple
copies of the research prototype. We require that the performance
of the pre-production device be equal to or higher than the numbers
shown. Note that SNR is an issue for fluorescence measurements
only. Reflectance measurements typically have an orders of
magnitude higher SNR and therefore do not have an SNR problem. Any
device changes that result in higher fluorescence SNR however, will
also further increase reflectance SNR. [0202] Single Channel SNR.
The SNR over multiple (static) measurements of the same point at
intensity levels measured on a temporally stable calibration target
for 340, 400 and 460 nm fluorescence emissions. SNR = Mean .times.
.times. of .times. .times. multiple .times. .times. intensities
.times. .times. measured .times. at .times. .times. a .times.
.times. single .times. .times. point .times. .times. on .times.
.times. a .times. .times. Calibration .times. .times. target
Standard .times. .times. deviation .times. .times. of .times.
.times. the .times. .times. same = 80 .+-. 10 ##EQU2## [0203]
Multi-channel SNR. The SNR over all 56 points in a single
measurement of a flat (spatially uniform) calibration target. This
measurement is also made at intensity levels measured on tissue for
340, 400 and 460 nm fluorescence emissions. SNR = Mean .times.
.times. of .times. .times. calibrated .times. .times. intensities
.times. .times. from 56 .times. .times. points .times. .times.
measured .times. .times. on .times. .times. a .times. .times.
Calibration .times. .times. target .times. Standard .times. .times.
deviation .times. .times. of .times. .times. the .times. .times.
same = 13 .+-. 2 ##EQU3##
[0204] i. Human factors. Certain human factors related parameters
based on marketing and human factors studies described in Sections
4.4.7 and 4.4.8, must be preserved as listed in Table 4.
TABLE-US-00003 TABLE 4 Key Human factor parameters 1. Weight of
Hand Held Unit 5 lbs 2. Dimensions of HHU Industrial design model
is available and optimized for usable 3. Length of Contact tube 176
mm. 4. Max diameter of contact tube 1.14 inch. at distal end 5.
Length of Distal end. 4 inch. 6. Max diameter of contact tube 1.26
inch. at proximal end. 7. Length of proximal end 2.8 inch
(including step down to 1.26 inch at the distal end)
[0205] FIG. 28 shows a functional block diagram of our research
prototype. The base unit of this device contains the excitation and
detection subsystems. The detection subsystem contains the
spectrometer and camera, the cost and performance drivers of this
device. They account for 65% of the overall device cost of the
research prototype. Without a priori knowledge of the functional
requirements of these two components, we chose a scientific grade
state-of-the-art camera and spectrograph for the research prototype
placing cost and size at a lower priority and deferring any effort
to reduce cost and size to when functional requirements were known.
In Phase I we will have designed, built and tested a size and cost
reduced integrated camera-spectrograph allowing us to move it into
the HHU (FIG.28 right panel). As a result of this, we will be able
to dispense with the coherent imaging bundle and the associated
cost since we no longer need an image transfer mechanism. An
immediate advantage of doing this is that our system throughput
will be increased two-fold since the coherent bundle has a
transmittance of 50%. Also as shown in Table 8, the device cost
share of the camera and spectrograph is now reduced to 40%.
[0206] FIG. 28 and FIG.35 illustrate an alternative embodiments
which may be preferred. The illustration on the left is a
functional block diagram of the research prototype showing how the
coherent imaging bundle connects the HHU to the detection subsystem
in the base unit. In an alternative embodiment the detection
subsystem is now placed in the HHU and we have dispensed with the
coherent fiber bundle. The excitation filter wheel and mask
assembly have also been transferred into the HHU. The items being
relocated in the new design are shown in grey.
[0207] FIG. 36 is a schematic view of the system shown in FIG. 35
with the confocual arrangement of excitation and collection units
at 5 degrees from each other;
[0208] FIG. 37 is diagramatic view of an improved embodiment where
the hand held unit contains the mask, motor, illumination optics,
video system, collection optics and detection system.
[0209] Excitation lamp and lamp housing. The excitation lamp used
in the research prototype is a 300 W short-arc Xe lamp. This lamp
uses bulb with an integrated parabolic reflector. A reduction in
lamp power is desirable in order to have an inexpensive, compact,
durable and rugged design with lower cooling demand. We, in our
device, have an apparent trade off between lamp power and
measurement SNR. An exception to this is to use a lamp with a
geometry that allows for more efficient focusing of light into a
given spot at a given Numerical Aperture (NA). This increases the
energy coupled into the fibers at the lamp end of the excitation
fiber bundle as illustrated in FIG. 29. The excitation fibers are
of a low Numerical Aperture or angle of light cone (NA) of
0.12.
[0210] Maximizing low NA light coupling is important to match the
low NA of the excitation optics in the HHU. The optics in turn use
a low NA since a high NA is prone to stray light generation. Also
our size constraints in the HHU require the use of small clear
apertures (CA). All high NA light will be rejected by these optics
and if present it will add to the stray light. We have used low NA
fibers to reject high NA light at fiber entrance thus minimizing
the generation of stray light in the optics. The disadvantage of a
low NA is that the light throughput is lowered. This require a
longer tissue examination, but elimination of the coherent fiber
bundle provides such a dramatic improvement in the throughput, that
the test time is actually shortened as seen in Table 6 below
[0211] Excitation fiber bundle and mask assembly. There is the only
fiber bundle that remains in the pre-production device design.
Minimizing the use of fiber optics in a product is necessary for
device ruggedness and for reducing the possibility of device damage
from fiber breakage. For patient safety reasons the lamp, the only
remaining key component, has been retained in the base unit and
away from the patient. A fiber bundle is necessary to transfer
light from the lamp in the base unit to the optics in the HHU. A
six-foot length is required on this fiber bundle for HHU
maneuverability and ease of use. We have to embodiments for this
fiber bundle as illustrated in FIG. 29. Option 1 appears to be the
preferred choice where both ends of the fiber bundle are fixed and
a moving mechanical mask selects the rows of spots that are
illuminated on tissue. The tradeoff is a lower power coupling per
fiber as shown in Table 5. A stepper motor under software control
will be used to move the mask. The structured end of the fiber
bundle will be fixed in the HHU using appropriate strain relief.
The mask and the mask stepper motor are both located in the
HHU.
[0212] FIG. 29. Excitation fiber bundle geometry options. In option
1 all 56, 100 .mu.m diameter fibers are illuminated. In option 2
only 9 fibers are illuminated providing for a smaller spot diameter
at the lamp end. However this option requires moving the fiber
bundle ferrule in the HHU to scan the tissue.
[0213] Excitation Optics. The excitation optics is located in the
HHU. These optics magnify and focus the spots of light produced by
the excitation fibers onto cervical tissue. A preliminary design is
illustrated in FIG. 30. The magnification factor chosen is 4.75
yielding nominally 500 .mu.m spots on the tissue. A similar design
is used in the existing research prototype. However, being limited
to available off-the-shelf lenses, we suffered significant
vignetting and transmittance loss which resulted in the excitation
optics throughput being <60%. Vignetting is a particularly
challenging issue that arises from our use of small lens clear
apertures (CA) in order to meet the size constrains of the HHU. The
design shown in FIG. 30 will use custom lenses and UV transmissive
glass (Schoft UBK7 glass). Custom lens prescriptions will also be
optimized to reduce vignetting and thus increase throughput. The
three lenses will be held in a 12 mm ID tube that is interrupted by
a filter wheel (excitation FW) as shown. Zemax analyses have shown
that it is possible to obtain a >90% throughput. In addition a
low divergence beam will be maintained through the excitation band
pass filters to eliminate wavelength shifts from non-normal light
incidence. We will optimize the design to maintain a high spot size
and intensity uniformity over the 25 mm diameter tissue as well as
a depth of focus of +/-5 mm.
[0214] Collection Optics. The collection optics is arranged
alongside the excitation optics in the HHU in a confocal
arrangement. The confocal arrangement makes it possible to locate
the optics and sensors outside the patient yet in the hand held
unit (HHU) and avoid the need for a coherent light fiber optic. The
focus of both is at the tissue located 176 mm from the HHU and
their primary axes form a 5 degree angle with respect to each
other. A preliminary design is illustrated in FIG. 31. This is a
design know as the Cooke triplet. The lenses are interspersed with
a collection filter (collection filter wheel) and a grating for
spectrally resolving the light. Early in our design efforts, we
chose a transmission grating for spectral resolution as illustrated
in FIG. 31. In the preferred embodiment we use a reflective grating
for superior stray light performance. We show the transmission
grating only to illustrate the design concept. Lens 1 demagnifies
and projects an intermediate real image at a field stop. Lens 2 and
3 respectively serve to collimate (necessary prior to spectral
splitting by the grating) and focus the spectrally resolved light
on the CCD.
[0215] FIG. 31 shows a transmission grating, an option we
previously considered. Although our current plan to use a
reflective grating the figure serves to illustrate the Cooke
Triplet design we have chosen.
[0216] In the reflective grating version, Lenses 2 and 3 are
replaced by a single concave mirror that performs the same
functions and images the field stop onto the CCD. The same design
using a reflective grating is illustrated in FIG. 32. A triplet is
chosen for lens 1 to correct for chromatic and geometric
aberrations. The design shown is preliminary and has superior
performance to that shown in FIG. 31 in terms of field distortion,
stray light and cost. FIG. 32 illustrate a Cooke triplet collection
optics design using a reflective convex
[0217] aberration corrected grating. Dimensions shown are in mm.
The location of the field stop is as shown 40 mm to the right of
the triplet. Clear Aperture of the triplet is to be determined.
[0218] Magnification. The image on tissue is demagnified
(magnification=0.25) by lens 1 (triplet in FIG. 32). This
demagnification is required and must be sufficient for spectra from
all spots to be imaged within the dimensions of our chosen CCD. The
CCD is chosen in the Phase I effort and subsequent design in Phase
II is therefore dependent on this choice. [0219] Intermediate real
image at a field stop. This intermediate imaging is important since
it allows the use of a spatial aperture that can be adjusted at the
factory to eliminate any mirroring artifacts from the inside of the
contact tubes. Although we will use a `flat black` material for the
contact tubes, some mirroring is expected and must be removed.
[0220] NA of light incident upon the CCD. The sensor chosen is an
interline CCD with microlens integrated onto the silicon substrate.
This limits the NA acceptance of the CCD with responsivity
declining upon increasing NA. We chose an NA <0.2 where the
responsivity is >80% of the maximum. This in turn limits the NA
of light accepted onto the concave mirror at the field stop to
<0.2. NA matching at the field stop is therefore important for
both, maximizing throughput and minimizing stray light.
[0221] Video imaging Channel. The video imaging channel provides
the user with feedback on proper device positioning on cervical
tissue. It also allows the user to capture still images of tissue
similar to those obtained by available video imaging colposcopes
currently available. This feature will permit the procedure
conducted with our device to be reimbursed using the colposcopy CPT
code (please see market analysis in Table 7).
[0222] We have included this video imaging channel in the
opto-mechanical layout of the pre-production HHU concept as
illustrated in FIG. 33.
[0223] Optical throughput comparison. Among the key device
requirements that determine data quality and discrimination
performance are the following: [0224] i. Maintain the nominal power
per spot at or higher than those listed in Table 3. [0225] ii.
Maintain an instrument SNR (single and multi channel) [0226] iii.
Reduce measurement time to less than 3 minutes.
[0227] To illustrate this we conducted an optical throughput
analysis using measured power and throughput efficiency for key
optical components in the research prototype and compared these
with those measured from the proposed components for the
pre-production device as shown in Table 6.
[0228] The power in each fiber is taken from Table 5 (option 1 for
one embodiment). The power at tissue is a measured value. This
energy was reflected off a 10% reflective Spectrolon calibration
target (Labsphere Inc. North Sutton, N.H.). Due to the lambertian
reflectance profile from Spectrolon, which is similar to that off
tissue, as well as the small aperture of the collection optics only
0.15% of the energy in each case is collected. This is a common
problem in any tissue spectroscopy device and only a small fraction
of energy emitted all around (in 4.about. steradians) is measured.
We have assumed the same research prototype collection optics
transmittance for both devices since the collection optics for the
proposed device does not exist although we believe that
transmittance in the design detailed in Section 4.4.4 will be
higher and can be improved by using custom coatings and UV
transmissive glasses. Our largest gains come from the elimination
of the coherent bundle and the use of the integrated spectrograph
design illustrated in FIG. 32. This results in a power at the CCD
that is 10 fold higher than that currently seen in the research
prototype. We can thus afford to lose some of this increase in
power by reducing the CCD integration time, thus reducing the
overall measurement time.
[0229] Table 6. Optical throughput comparison between the research
prototype and the pre-production device. Starting with 460 nm (20
nm band pass) light coupled into a 100 .mu.m excitation fiber and
measuring the power exiting each component along the optical path
we arrived at the power incident upon the CCD. Embodiment 2 shows a
10-fold increase in power arriving at the CCD compared to
embodiment 1. TABLE-US-00004 TABLE 6 OPTICAL THROUGHPUT COMPARISON
Embodiment 1 (using coherent Power in each excitation FIBER (.mu.W)
28 Excitation optics thruput (transmittance, 57% vignetting) Power
at tissue (.mu.W) 16 Target Reflectance (Spectralon) 10% Light
gathering of collection optics 0.15% Collection optics throughput
67% Collection filter transmittance 92% Coherent bundle
transmittance 46% NA missmatch at Spectrograph 54% Acton
spectrograph transmittance 32% Power at CCD (nano W) 0.12 460 nm
excitation (20 nm band FWHM) 460 nm collection (20 nm band FWHM)
Embodiment 2 (no coherent bundle) Power in each excitation FIBER
(.mu.W) 25 Excitation optics thruput (transmittance, 90%
vignetting) Power at tissue (.mu.W) 23 Target Reflectance
(Spectralon) 10% Light gathering of collection optics 0.15%
Collection optics throughput 67% Collection filter transmittance
92% Reflective spectrograph transmittance 60% Power at CCD (nano W)
1.25
[0230] The ectocervical probe (component 1 in FIG. 34) contains a
single-use contact tube (the black cylinder) that is positioned to
contact the patient's cervix. It is shown as conical but could be
tapered so long as it is not less that the system NA so that it
does not impinge on the cone of light. The contact tube ensures
proper positioning of the cervix relative to the optics in the
probe body, attenuates ambient light sources, contains an integral
calibration standard used to calibrate the instrument prior to
patient measurements, and prevents cross-contamination between
patients. The back of the ectocervical probe contains the controls
and display used to operate the instrument. The endocervical probe
(component 2 in FIG. 34) is a separate, slender probe that it used
to measure the endocervical tissue. A single-use protective sheath
is placed over the probe prior to measurement to prevent
cross-contamination. The base unit (component 3 in the FIG. 34)
contains common electronics, power supplies, light sources,
printer, etc.
[0231] The following method may be used to perform a measurement.
[0232] 1) Attach a new disposable contact tube to the ecto-cervical
hand held unit (HHU) and enter patient data via the controls on the
back of the ecto-cervical HHU. At this point, an automatic
calibration is performed using a disposable calibration standard
attached to the contact tube. [0233] 2) Remove and discard the
calibration material from the contact tube. [0234] 3) Gently press
the contact tube against the cervix by placing it inside a standard
vaginal speculum that is already positioned inside the patient. The
operator is aided in this positioned step by viewing a live video
image of the cervix on the display on the back of the ecto-cervical
HHU. Once the probe is properly positioned, the instrument
automatically scans the cervix and collects the spectroscopic data
in approximately three to four minutes. [0235] 4) Remove the
ecto-cervical HHU from patient, remove and discard the contact
tube, and place the HHU back in the base unit. [0236] 5) If an
endo-cervical canal measurement is desired, place a new protective
sheath on the endo-cervical probe and gently insert the probe into
the patient's endo-cervical canal. Hold the probe in place for
approximately one minute while the instrument collects
spectroscopic data. Still and video images may be taken and the
test will be run with a final still image preferably taken at the
end of the test. When the data collection is complete, remove the
probe, discard the sheath, and place the probe in the base unit.
[0237] 6) The instrument automatically displays the test results on
the ecto-cervical probe display and optionally prints a hardcopy on
the printer. Results are shown using a numerical scale that ranges
from 0 to 100, with higher values indicating greater probability of
CIN2+ cervical disease. The device will also display a disease
localization map, which will used a color coding to indicate areas
on the cervix with of highest likelihood of having CIN2+ cervical
disease. The map may be overlayed on to the still or video images
which have been taken.
[0238] The continuous output scale displayed to the physician will
be evaluated at two thresholds in order to determine sensitivity
for detecting CIN2+ cervical disease and specificity for ruling out
benign lesions on the cervix.
[0239] The foregoing embodiments and advantages are merely
exemplary and are not to be construed as limiting the present
invention. The present teaching can be readily applied to other
types of apparatuses and applications that may be common to those
of ordinary skill in the art. The description of the present
invention is intended to be illustrative, and not to limit the
scope of the claims. Many alternatives, modifications, and
variations will be apparent to those skilled in the art. In the
claims, means-Plus-function clauses are intended to cover the
structures described herein as performing the recited function and
not only structural equivalents but also equivalent structures.
* * * * *