U.S. patent application number 14/058363 was filed with the patent office on 2014-09-04 for radioactive emission detector equipped with a position tracking system.
This patent application is currently assigned to Blosensor International Group, Ltd.. The applicant listed for this patent is Blosensor International Group, Ltd.. Invention is credited to Roni Amarmi, Udi Antebi, Gal Ben-David, Yona Bouskila, Yoav Kimchy, Nick Sidorenko, Yoel Ziberstein.
Application Number | 20140249402 14/058363 |
Document ID | / |
Family ID | 31999761 |
Filed Date | 2014-09-04 |
United States Patent
Application |
20140249402 |
Kind Code |
A1 |
Kimchy; Yoav ; et
al. |
September 4, 2014 |
RADIOACTIVE EMISSION DETECTOR EQUIPPED WITH A POSITION TRACKING
SYSTEM
Abstract
A radioactive emission probe in communication with a position
tracking system and the use thereof in a variety of systems and
methods of medical imaging and procedures, are provided.
Specifically, wide-aperture collimation-deconvolution algorithms
are provided, for obtaining a high-efficiency, high resolution
image of a radioactivity emitting source, by scanning the
radioactivity emitting source with a probe of a wide-aperture
collimator, and at the same time, monitoring the position of the
radioactive emission probe, at very fine time intervals, to obtain
the equivalence of fine-aperture collimation. The blurring effect
of the wide aperture is then corrected mathematically. Furthermore,
an imaging method by depth calculations is provided, based on the
attenuation of photons of different energies, which are emitted
from the same source, coupled with position monitoring.
Inventors: |
Kimchy; Yoav; (Haifa,
IL) ; Amarmi; Roni; (Yokneam, IL) ; Bouskila;
Yona; (Maidenhead, GB) ; Antebi; Udi; (Kiryat
Bialik, IL) ; Sidorenko; Nick; (Acre, IL) ;
Ben-David; Gal; (Mitzpe Adi, IL) ; Ziberstein;
Yoel; (Herzilla, IL) |
|
Applicant: |
Name |
City |
State |
Country |
Type |
Blosensor International Group, Ltd. |
HaMilton |
|
BM |
|
|
Assignee: |
Blosensor International Group,
Ltd.
Hamilton
BM
|
Family ID: |
31999761 |
Appl. No.: |
14/058363 |
Filed: |
October 21, 2013 |
Related U.S. Patent Documents
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Application
Number |
Filing Date |
Patent Number |
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10616307 |
Jul 10, 2003 |
8565860 |
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14058363 |
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10343792 |
Feb 4, 2003 |
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PCT/IL01/00638 |
Jul 11, 2001 |
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10616307 |
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09727464 |
Dec 4, 2000 |
7826889 |
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10343792 |
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09641973 |
Aug 21, 2000 |
8489176 |
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09727464 |
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60394936 |
Jul 11, 2002 |
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60286044 |
Apr 25, 2001 |
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Current U.S.
Class: |
600/411 ;
600/424; 600/427; 600/436 |
Current CPC
Class: |
A61B 8/4245 20130101;
A61B 6/482 20130101; A61B 5/415 20130101; A61B 5/6835 20130101;
A61B 2090/392 20160201; A61B 6/4258 20130101; A61B 6/12 20130101;
A61B 8/4254 20130101; A61B 8/0833 20130101; A61B 5/064 20130101;
A61B 5/06 20130101; G01T 1/161 20130101; A61B 5/07 20130101; A61B
6/4417 20130101; A61B 5/055 20130101; A61B 6/4057 20130101; A61B
6/037 20130101; A61B 6/583 20130101; A61B 5/418 20130101; A61B
6/4423 20130101 |
Class at
Publication: |
600/411 ;
600/436; 600/427; 600/424 |
International
Class: |
A61B 6/03 20060101
A61B006/03; A61B 5/055 20060101 A61B005/055; A61B 5/06 20060101
A61B005/06; A61B 6/00 20060101 A61B006/00 |
Claims
1. An intracorporeal-imaging head, comprising: a housing, which
comprises: at least one radioactive-emission probe, mounted on said
housing, adapted to image radioactive-emission from at least two
different viewing angles of a same portion of a tissue without
movement of the housing.
2. The intracorporeal imaging head of claim 1, further comprising
an imaging system adapted to image said portion.
3. The intracorporeal imaging head of claim 2, wherein the imaging
system is one of a fluoroscope, a computed tomographer, a magnetic
resonance imager, an ultrasound imager, an impedance imager, and an
optical camera.
4. The intracorporeal imaging head of claim 1 wherein the at least
one radioactive-emission probe is at least one wide-angle
collimator probe.
5. The intracorporeal imaging head of claim 4 wherein the
wide-angle collimator probe has a viewing angle of between
81.degree. and 280.degree..
6. The intracorporeal imaging head of claim 4, further comprising:
a data processor for correcting blurring effects from said
wide-angle collimator probe.
7. The intracorporeal imaging head of claim 6, wherein said
wide-angle collimator probe is adapted to obtain image data at a
first resolution or less, and wherein said data processor is
further adapted for forming a three-dimensional model of said
portion of said tissue, wherein said three-dimensional model is in
a second resolution, higher than said first resolution.
8. The intracorporeal imaging head of claim 6, wherein said data
processor corrects the blurring effects using a
collimation-deconvolution algorithm based on account estimation of
a transfer function.
9. The intracorporeal imaging head of claim 6, wherein said data
processor corrects the blurring effects by calculating a value of
readings for each voxel based on multiple readings from view points
of said probe; and using a collimation-deconvolution algorithm to
reconstruct voxels of the radiation map with diminished or no
blurriness.
10. The intracorporeal-imaging head of claim 1, further comprising:
a position tracking system in a fixed positional relation with said
at least one radioactive-emission probe, for providing positional
information for said at least one radioactive-emission probe.
11. The intracorporeal imaging head of claim 10 wherein the
position tracking system is one of a fluoroscope, a computed
tomographer, a magnetic resonance imager, an ultrasound imager, an
impedance imager, and an optical camera.
12. The intracorporeal imaging head of claim 10, wherein the
position tracking system comprises an imaging system.
13. The intracorporeal imaging head of claim 10 wherein the
position tracking system is adapted to track acoustic
electromagnetic radiation or magnetic fields.
14. The intracorporeal imaging head of claim 10 wherein the at
least one radioactive-emission probe is at least one wide-angle
collimator probe.
15. The intracorporeal imaging head of claim 10 wherein the
wide-angle collimator probe has a viewing angle of between
81.degree. and 280.degree..
16. The intracorporeal imaging head of claim 10, wherein said
wide-angle collimator probe is adapted to obtain image data at a
first resolution or less, and wherein the intracorporeal imaging
head further comprises a data processor for forming a
three-dimensional model of said portion of said tissue by data
processing said image data at said first resolution and positional
information received from said position tracking system, wherein
said three-dimensional model is in a second resolution, higher than
said first resolution.
17. The intracorporeal imaging head of claim 1, wherein said
radioactive probe comprises a plurality of radiation detectors.
18. The intracorporeal imaging head of claim 1, wherein said
radioactive probe comprises a plurality of single pixel detector
units configured to generate image data from each of said
pixels.
19. A flexible probe, comprising: an intracorporeal imaging head
according to claim 1; and an ultrasonic imager sized for rectal
insertion for imaging of prostate.
20. A flexible probe, comprising: an intracorporeal imaging head
according to claim 1; and an optical imager sized for rectal
insertion for imaging of colon.
21. A device for obtaining an image of a radioactivity emitting
source in a system-of-coordinates, the device comprising: a
moveable radioactive emission detector; said radioactive emission
detector being moveable to perform a plurality of radioactivity
measurements of a radioactivity emitting source from a plurality of
locations and directions; a position tracking device which monitors
a position of said radioactive emission detector in relation to
said system of coordinates; and a data processor which reconstructs
an image of the radioactivity emitting source from a plurality of
said radioactivity measurements with a varying spatial resolution
and obtains a positional distribution of said radioactive emitting
source in relation to said system of coordinates using measurements
of a same area from multiple locations having different resolution
according to said position.
22. The device of claim 21, wherein said position tracking device
is comprised of a plurality of accelerometers, or a plurality of
potentiometers, or is sound wave based or radio frequency based, or
electromagnetic field based or optically based.
23. The device of claim 21, wherein said position tracking device
is responsive to a change in the position of said detector.
24. The device of claim 21, wherein said detector is carried by a
moveable arm, and said position tracking device is responsive to
movement of said arm.
25. The device of claim 24, wherein said moveable arm is
articulated.
26. The device of claim 21, wherein said detector performs said
plurality of radioactivity measurements at different
resolutions.
27. The device of claim 21, wherein said data processor
reconstructs said image according to radioactivity measurements
having different resolutions.
28. The device of claim 27, wherein said resolutions varies
according to the distance of said detector from said radiation
source.
29. The device of claim 28, wherein said spatial resolution is
higher for radioactivity measurements as the detector approaches
the radiation source.
30. The device of claim 21, wherein said radioactive emission
detector is configured for free-hand scanning, and the position of
the detector is monitored by said position tracking device in a
first system-of-coordinates, and further comprising: a surgical
instrument whose position is monitored by an additional position
tracking device in a second system-of-coordinates; and at least one
data processor designed and configured to receive data inputs from
said position tracking device, from said radioactive emission
detector and from said additional position tracking device and to
calculate the position of the surgical instrument and the
radiopharmaceutical up-taking portion of the body component in a
common system-of-coordinates.
31. The device of claim 30, wherein said first position tracking
device and said additional position tracking device are comprised
in a single unit.
32. The device of claim 30, further comprising a display device
which serves for visual co-presentation of the position of said
surgical instrument and the radiopharmaceutical up-taking portion
of the body component.
33. The device of claim 30, wherein said surgical instrument is
selected from the group consisting of laser probe, a cardiac
catheter, an angioplastic catheter, an endoscopic probe, a biopsy
needle, an ultrasonic probe, fiber optic scopes, aspiration tubes,
a laparoscopy probe, a thermal probe and a suction/irrigation
probe.
34. The device of claim 30, wherein said common system of
coordinates is a three dimensional space.
35. The device of claim 1, wherein said system of coordinates is a
three dimensional space.
36. The device of claim 1, wherein said data processor is
responsive to said plurality of said radioactivity measurements to
reconstruct a three dimensional image of said radioactivity
emitting source.
Description
RELATED APPLICATIONS
[0001] This application is a continuation of U.S. patent
application Ser. No. 10/616,307 filed on Jul. 10, 2003, which
claims the benefit of priority under 35 USC .sctn.119(e) of U.S.
Provisional Patent Application No. 60/394,936 filed on Jul. 11,
2002.
[0002] U.S. patent application Ser. No. 10/616,307 is also a
continuation-in-part (CIP) of U.S. patent application Ser. No.
10/343,792 filed on Feb. 4, 2003, which is a National Phase of PCT
Patent Application No. PCT/IL01/00638 filed on Jul. 11, 2001, which
is a continuation-in-part (CIP) of U.S. patent application Ser. No.
09/727,464 filed on Dec. 4, 2000, now U.S. Pat. No. 7,826,889,
which is a continuation-in-part (CIP) of U.S. patent application
Ser. No. 09/714,164 filed Nov. 17, 2000, now abandoned, which is a
continuation-in-part (CIP) of U.S. patent application Ser. No.
09/641,973 filed Aug. 21, 2000, now U.S. Pat. No. 8,489,176.
[0003] PCT Patent Application No. PCT/IL01/00638 also claims the
benefit of priority under 35 USC .sctn.119(e) of U.S. Provisional
Patent Application No. 60/286,044 filed on Apr. 25, 2001.
[0004] The contents of the above applications are all incorporated
by reference as if fully set forth herein in their entirety.
FIELD AND BACKGROUND OF THE INVENTION
[0005] The present invention relates to a radioactive emission
probe equipped with a position tracking system. More particularly,
the present invention relates to the functional integration of a
radioactive emission probe equipped with a position tracking system
as above with medical imaging modalities and (or) with guided
minimally-invasive surgical instruments. The present invention is
therefore useful for calculating the position of a concentrated
radiopharmaceutical in the body in positional context of imaged
portions of the body, which information can be used, for example,
for performing an efficient minimally invasive surgical procedure.
The present invention further relates to a surgical instrument
equipped with a position tracking system and a radioactive emission
probe for fine in situ localization during resection and (or)
biopsy procedures, which surgical instrument is operated in concert
with other aspects of the invention.
[0006] The use of minimally invasive surgical techniques has
dramatically affected the methods and outcomes of surgical
procedures. Physically cutting through tissue and organs to
visually expose surgical sites in conventional "open surgical"
procedures causes tremendous blunt trauma and blood loss. Exposure
of internal tissues and organs in this manner also dramatically
increases the risk of infection. Trauma, blood loss, and infection
all combine to extend recovery times, increase the extent of
complications, and require a more intensive care and monitoring
regimen. The result of such open surgical procedures is more pain
and suffering, higher procedural costs, and greater risk of adverse
outcomes.
[0007] In sharp contrast, minimally invasive surgical procedures
cause little blunt trauma or blood loss and minimize the risk of
infection by maintaining the body's natural barriers to infection
substantially intact. Minimally invasive surgical procedures result
in faster recovery and cause fewer complications than conventional,
open, surgical procedures. Minimally invasive surgical procedures,
such as laparoscopic, endoscopic, or cystoscopic surgeries, have
replaced more invasive surgical procedures in all areas of surgical
medicine. Due to technological advancements in areas such as fiber
optics, micro-tool fabrication, imaging and material science, the
physician performing the operation has easier-to-operate and more
cost-effective tools for use in minimally invasive procedures.
However, there still exist a host of technical hurdles that limit
the efficacy and increase the difficulty of minimally invasive
procedures, some of which were overcome by the development of
sophisticated imaging techniques. As is further detailed below, the
present invention offers further advantages in this respect.
[0008] Radionuclide imaging is one of the most important
applications of radioactivity in medicine. The purpose of
radionuclide imaging is to obtain a distribution image of a
radioactively labeled substance, e.g., a radiopharmaceutical,
within the body following administration thereof to a patient.
Examples of radiopharmaceuticals include monoclonal antibodies or
other agents, e.g., fibrinogen or fluorodeoxyglucose, tagged with a
radioactive isotope, e.g., .sup.99Mtechnetium, .sup.67gallium,
.sup.201thallium, .sup.111indium, .sup.123iodine, .sup.125iodine
and .sup.18fluorine, which may be administered orally or
intravenously. The radiopharmaceuticals are designed to concentrate
in the area of a tumor, and the uptake of such radiopharmaceuticals
in the active part of a tumor, or other pathologies such as an
inflammation, is higher and more rapid than in the tissue that
neighbors the tumor. Thereafter, a radiation emission detector,
typically an invasive detector or a gamma camera (see below), is
employed for locating the position of the active area. Another
application is the detection of blood clots with
radiopharmaceuticals such as ACUTECT from Nycomed Amersham for the
detection of newly formed thrombosis in veins, or clots in arteries
of the heart or brain, in an emergency or operating room. Yet other
applications include radioimaging of myocardial infarct using
agents such as radioactive anti-myosin antibodies, radioimaging
specific cell types using radioactively tagged molecules (also
known as molecular imaging), etc.
[0009] The distribution image of the radiopharmaceutical in and
around a tumor, or another body structure, is obtained by recording
the radioactive emission of the radiopharmaceutical with an
external radiation detector placed at different locations outside
the patient. The usual preferred emission for such applications is
that of gamma rays, which emission is in the energy range of
approximately 20-511 KeV. When the probe is placed in contact with
the tissue, beta radiation and positrons may also be detected.
[0010] The first attempts at radionuclide "imaging" were in the
late 1940's. An array of radiation detectors was positioned
mechanically on a matrix of measuring points around the head of a
patient. Alternatively, a single detector was positioned
mechanically for separate measurements at each point on the
matrix.
[0011] A significant advance occurred in the early 1950's with the
introduction of the rectilinear scanner by Ben Cassen. With this
instrument, the detector was scanned mechanically in a
predetermined pattern over the area of interest.
[0012] The first gamma camera capable of recording all points of
the image at one time was described by Hal Anger in 1953. Anger
used a detector comprising a NaI(Tl) screen and a sheet of X-ray
film. In the late 1950's, Anger replaced the film screen with a
photomultiplier tube assembly. The Anger camera is described in Hal
O. Anger, "Radioisotope camera in Hine GJ", Instrumentation in
Nuclear Medicine, New York, Academic Press 1967, chapter 19. U.S.
Pat. No. 2,776,377 to Anger, issued in 1957, also describes such a
radiation detector assembly.
[0013] U.S. Pat. No. 4,959,547 to Carroll et al. describes a probe
used to map or provide imaging of radiation within a patient. The
probe comprises a radiation detector and an adjustment mechanism
for adjusting the solid angle through which radiation may pass to
the detector, the solid angle being continuously variable. The
probe is constructed so that the only radiation reaching the
detector is that which is within the solid angle. By adjusting the
solid angle from a maximum to a minimum while moving the probe
adjacent the source of radiation and sensing the detected
radiation, one is able to locate the probe at the source of
radiation. The probe can be used to determine the location of the
radioactivity and to provide a point-by-point image of the
radiation source or data for mapping the same.
[0014] U.S. Pat. No. 5,246,005 to Carroll et al. describes a
radiation detector or probe, which uses statistically valid signals
to detect radiation signals from tissue. The output of a radiation
detector is a series of pulses, which are counted for a
predetermined amount of time. At least two count ranges are defined
by circuitry in the apparatus and the count range which includes
the input count is determined. For each count range, an audible
signal is produced which is audibly discriminable from the audible
signal produced for every other count range. The mean values of
each count range are chosen to be statistically different, e.g., 1,
2, or 3 standard deviations, from the mean of adjacent lower or
higher count ranges. The parameters of the audible signal, such as
frequency, voice, repetition rate, and (or) intensity are changed
for each count range to provide a signal which is discriminable
from the signals of any other count range.
[0015] U.S. Pat. No. 5,475,219 to Olson describes a system for
detecting photon emissions wherein a detector serves to derive
electrical parameter signals having amplitudes corresponding with
the detected energy of the photon emissions and other signal
generating events. Two comparator networks employed within an
energy window, which define a function to develop an output, L,
when an event-based signal amplitude is equal to or above a
threshold value, and to develop an output, H, when such signal
amplitude additionally extends above an upper limit. Improved
reliability and accuracy is achieved with a discriminator circuit
which, in response to these outputs L and H, derives an event
output upon the occurrence of an output L in the absence of an
output H. This discriminator circuit is an asynchronous,
sequential, fundamental mode discriminator circuit with three
stable states.
[0016] U.S. Pat. Nos. 5,694,933 and 6,135,955 to Madden et al.
describe a system and method for diagnostic testing of a structure
within a patient's body that has been provided with a radioactive
imaging agent, e.g., a radiotracer, to cause the structure to
produce gamma rays, associated characteristic x rays, and a
continuum of Compton-scattered photons. The system includes a
radiation receiving device, e.g., a hand-held probe or camera, an
associated signal processor, and an analyzer. The radiation
receiving device is arranged to be located adjacent the body and
the structure for receiving gamma rays and characteristic X-rays
emitted from the structure and for providing a processed electrical
signal representative thereof. The processed electrical signal
includes a first portion representing the characteristic X-rays
received and a second portion representing the gamma rays received.
The signal processor removes the signal corresponding to the
Compton-scattered photons from the electrical signal in the region
of the full-energy gamma ray and the characteristic X-ray. The
analyzer is arranged to selectively use the X-ray portion of the
processed signal to provide near-field information about the
structure, to selectively use both the X-ray and the gamma-ray
portions of the processed signal to provide near-field and
far-field information about the structure, and to selectively use
the gamma-ray portion of the processed signal to provide extended
field information about the structure.
[0017] U.S. Pat. No. 5,732,704 to Thurston et al. describes a
method for identifying a sentinel lymph node located within a
grouping of regional nodes at a lymph drainage basin associated
with neoplastic tissue wherein a radiopharmaceutical is injected at
the situs of the neoplastic tissue. This radiopharmaceutical
migrates along a lymph duct towards the drainage basin containing
the sentinel node. A hand-held probe with a forwardly disposed
radiation detector crystal is maneuvered along the duct while the
clinician observes a graphical readout of count rate amplitudes to
determine when the probe is aligned with the duct. The region
containing the sentinel node is identified when the count rate at
the probe substantially increases. Following surgical incision, the
probe is maneuvered utilizing a sound output in connection with
actuation of the probe to establish increasing count rate
thresholds followed by incremental movements until the threshold is
not reached and no sound cue is given to the surgeon. At this point
of the maneuvering of the probe, the probe detector will be in
adjacency with the sentinel node, which then may be removed.
[0018] U.S. Pat. No. 5,857,463 to Thurston et al. describes further
apparatus for tracking a radiopharmaceutical present within the
lymph duct and for locating the sentinel node within which the
radiopharmaceutical has concentrated. A smaller, straight,
hand-held probe is employed carrying two hand actuable switches.
For tracking procedures, the probe is moved in an undulatory
manner, wherein the location of the radiopharmaceutical-containing
duct is determined by observing a graphic readout. When the region
of the sentinel node is approached, a switch on the probe device is
actuated by the surgeon to carry out a sequence of squelching
operations until a small node locating region is defined.
[0019] U.S. Pat. No. 5,916,167 to Kramer et al. and U.S. Pat. No.
5,987,350 to Thurston describe surgical probes wherein a
heat-sterilizable and reusable detector component is combined with
a disposable handle and cable assembly. The reusable detector
component incorporates a detector crystal and associated mountings
along with preamplifier components.
[0020] U.S. Pat. No. 5,928,150 to Call describes a system for
detecting emissions from a radiopharmaceutical injected within a
lymph duct wherein a hand-held probe is utilized. When employed to
locate sentinel lymph nodes, supplementary features are provided
including a function for treating validated photon event pulses to
determine count rate level signals. The system includes a function
for count-rate based ranging as well as an adjustable threshold
feature. A post-threshold amplification circuit develops full-scale
aural and visual outputs.
[0021] U.S. Pat. Nos. 5,932,879 and 6,076,009 to Raylman et al.
describe an intraoperative system for preferentially detecting beta
radiation over gamma radiation emitted from a radiopharmaceutical.
The system has ion-implanted silicon charged-particle detectors for
generating signals in response to received beta particles. A
preamplifier is located in proximity to the detector filters and
amplifies the signal. The probe is coupled to a processing unit for
amplifying and filtering the signal.
[0022] U.S. Pat. No. 6,144,876 to Bouton describes a system for
detecting and locating sources of radiation, with particular
applicability to interoperative lymphatic mapping (ILM) procedures.
The scanning probe employed with the system performs with both an
audible as well as a visual perceptive output. A desirable
stability is achieved in the readouts from the system through a
signal processing approach which establishes a floating or dynamic
window analysis of validated photon event counts. This floating
window is defined between an upper edge and a lower edge. The
values of these window edges vary during the analysis in response
to compiled count sum values. In general, the upper and lower edges
are spaced apart a value corresponding with about four standard
deviations.
[0023] To compute these count sums, counts are collected over
successive short scan intervals of 50 milliseconds and the count
segments resulting therefrom are located in a succession of bins
within a circular buffer memory. The count sum is generated as the
sum of the memory segment count values of a certain number of the
bins or segments of memory. Alteration of the floating window
occurs when the count sum either exceeds its upper edge or falls
below its lower edge. A reported mean, computed with respect to the
window edge that is crossed, is developed for each scan interval
which, in turn, is utilized to derive a mean count rate signal. The
resulting perceptive output exhibits a desirable stability,
particularly under conditions wherein the probe detector is in a
direct confrontational geometry with a radiation source.
[0024] U.S. Pat. No. 5,846,513 teaches a system for detecting and
destroying living tumor tissue within the body of a living being.
The system is arranged to be used with a tumor localizing
radiopharmaceutical. The system includes a percutaneously
insertable radiation detecting probe, an associated analyzer, and a
percutaneously insertable tumor removing instrument, e.g., a
resectoscope. The radiation detecting probe includes a needle unit
having a radiation sensor component therein and a handle to which
the needle unit is releasably mounted. The needle is arranged to be
inserted through a small percutaneous portal into the patient's
body and is movable to various positions within the suspected tumor
to detect the presence of radiation indicative of cancerous tissue.
The probe can then be removed and the tumor removing instrument
inserted through the portal to destroy and (or) remove the
cancerous tissue. The instrument not only destroys the tagged
tissue, but also removes it from the body of the being so that it
can be assayed for radiation to confirm that the removed tissue is
cancerous and not healthy tissue. A collimator may be used with the
probe to establish the probe's field of view.
[0025] The main limitation of the system is that once the body is
penetrated, scanning capabilities are limited to a translational
movement along the line of penetration.
[0026] An effective collimator for gamma radiation must be several
mm in thickness and therefore an effective collimator for high
energy gamma radiation cannot be engaged with a fine surgical
instrument such as a surgical needle. On the other hand, beta
radiation is absorbed mainly due to its chemical reactivity after
passage of about 0.2-3 mm through biological tissue. Thus, the
system described in U.S. Pat. No. 5,846,513 cannot efficiently
employ high energy gamma detection because directionality will to a
great extent be lost and it also cannot efficiently employ beta
radiation because too high proximity to the radioactive source is
required, whereas body tissue limits the degree of maneuvering the
instrument.
[0027] The manipulation of soft tissue organs requires
visualization (imaging) techniques such as computerized tomography
(CT), fluoroscopy (X-ray fluoroscopy), magnetic resonance imaging
(MRI), optical endoscopy, mammography or ultrasound which
distinguish the borders and shapes of soft tissue organs or masses.
Over the years, medical imaging has become a vital part in the
early detection, diagnosis and treatment of cancer and other
diseases. In some cases medical imaging is the first step in
preventing the spread of cancer through early detection and in many
cases medical imaging makes it possible to cure or eliminate the
cancer altogether via subsequent treatment.
[0028] An evaluation of the presence or absence of tumor metastasis
or invasion has been a major determinant for the achievement of an
effective treatment for cancer patients. Studies have determined
that about 30% of patients with essentially newly diagnosed tumor
will exhibit clinically detectable metastasis. Of the remaining 70%
of such patients who are deemed "clinically free" of metastasis,
about one-half are curable by local tumor therapy alone. However,
some of these metastasis or even early stage primary tumors do not
show with the imaging tools described above. Moreover often enough
the most important part of a tumor to be removed for biopsy or
surgically removed is the active, i.e., growing part, whereas using
only conventional imaging cannot distinguish this specific part of
a tumor from other parts thereof and (or) adjacent non affected
tissue.
[0029] A common practice in order to locate this active part is to
mark it with radioactivity tagged materials generally known as
radiopharmaceuticals, which are administered orally or
intravenously and which tend to concentrate in such areas, as the
uptake of such radiopharmaceuticals in the active part of a tumor
is higher and more rapid than in the neighboring tumor tissue.
Thereafter, a radiation emission detector, typically an invasive
detector, is employed for locating the position of the active
area.
[0030] Medical imaging is often used to build computer models which
allow doctors to, for example, guide exact radiation in the
treatment of cancer, and to design minimally-invasive or open
surgical procedures. Moreover, imaging modalities are also used to
guide surgeons to the target area inside the patient's body, in the
operation room during the surgical procedure. Such procedures may
include, for example, biopsies, inserting a localized radiation
source for direct treatment of a cancerous lesion, known as
brachytherapy (so as to prevent radiation damage to tissues near
the lesion), injecting a chemotherapy agent into the cancerous site
or removing a cancerous or other lesions.
[0031] The aim of all such procedures is to pin-point the target
area as precisely as possible in order to get the most precise
biopsy results, preferably from the most active part of a tumor, or
to remove such a tumor in its entirety, with minimal damage to the
surrounding, non affected tissues.
[0032] This goal is yet to be achieved, as most of the common
imaging modalities such as fluoroscopy, CT, MRI, mammography or
ultrasound demonstrate the position and appearance of the entire
lesion with anatomical modifications that the lesion causes to its
surrounding tissue, without differentiating between the non-active
mass from the physiologically active part thereof.
[0033] Furthermore, prior art radiation emission detectors and (or)
biopsy probes, while being suitable for identifying the location of
the radiation site, leave something to be desired from the
standpoint of facilitating the removal or other destruction of the
detected cancerous tissue, with minimal trauma.
[0034] The combination of modalities, as is offered by the present
invention, can reduce the margin of error in locating such tumors.
In addition, the possibility of demonstrating the position of the
active part of a tumor superimposed on a scan from an imaging
modality that shows the organ or tumor, coupled with the
possibility to follow a surgical tool in reference to the afflicted
area during a surgical procedure will allow for a more precise and
controlled surgical procedures to take place, minimizing the
aforementioned problems.
[0035] The present invention addresses these and other issues which
are further elaborated hereinbelow, and offers the physicians and
patients more reliable targeting, which in turn will result in less
invasive and less destructive surgical procedures and fewer cases
of mistaken diagnoses.
SUMMARY OF THE INVENTION
[0036] The present invention successfully addresses the
shortcomings of the presently known configurations by providing a
radioactive emission probe in communication with a position
tracking system and the use thereof in a variety of systems and
methods of medical imaging and procedures. Specifically,
wide-aperture collimation-deconvolution algorithms are provided,
for obtaining a high-efficiency, high resolution image of a
radioactivity emitting source, by scanning the radioactivity
emitting source with a probe of a wide-aperture collimator, and at
the same time, monitoring the position of the radioactive emission
probe, at very fine time intervals, to obtain the equivalence of
fine-aperture collimation. The blurring effect of the wide aperture
is then corrected mathematically. Furthermore, an imaging method by
depth calculations is provided, based on the attenuation of photons
of different energies, which are emitted from the same source,
coupled with position monitoring.
[0037] The present invention has many other applications in the
direction of therapeutics, such as, but not limited to, implanting
brachytherapy seeds, ultrasound microwave radio-frequency
cryotherapy and localized radiation ablations.
[0038] Implementation of the methods and systems of the present
invention involves performing or completing selected tasks or steps
manually, automatically, or a combination thereof. Moreover,
according to actual instrumentation and equipment of preferred
embodiments of the methods and systems of the present invention,
several selected steps could be implemented by hardware or by
software on any operating system of any firmware or a combination
thereof. For example, as hardware, selected steps of the invention
could be implemented as a chip a circuit. As software, selected
steps of the invention could be implemented as a plurality of
software instructions being executed by a computer using any
suitable algorithms. In any case, selected steps of the method and
system of the invention could be described as being performed by a
data processor, such as a computing platform for executing a
plurality of instructions.
BRIEF DESCRIPTION OF THE DRAWINGS
[0039] The invention is herein described, by way of example only,
with reference to the accompanying drawings. With specific
reference now to the drawings in detail, it is stressed that the
particulars shown are by way of example and purposes of
illustrative discussion of the preferred embodiments of the present
invention only, and are presented in the cause of providing what is
believed to be the most useful and readily understood description
of the principles and conceptual aspects of the invention. In this
regard, no attempt is made to show structural details of the
invention in more detail than is necessary for a fundamental
understanding of the invention, the description taken with the
drawings making apparent to those skilled in the art how the
several forms of the invention may be embodied in practice.
[0040] In the drawings:
[0041] FIG. 1 is a black box diagram of a system according to the
teachings of the present invention;
[0042] FIG. 2 is a perspective view of an articulated arm which
serves as a position tracking system shown carrying a radioactive
emission probe in accordance with the teachings of the present
invention;
[0043] FIG. 3 is a schematic depiction of a radioactive emission
probe carrying a pair of three coaxially aligned accelerometers
which serve as a position tracking system in accordance with the
teachings of the present invention;
[0044] FIG. 4 is a schematic presentation of a radioactive emission
probe communicating with yet another type of a position tracking
system in accordance with the teachings of the present
invention;
[0045] FIG. 5 is a simplified cross-sectional view of a narrow or
wide angle radioactive emission probe used to implement an
embodiment of the present invention;
[0046] FIG. 6 is a presentation of a scanning protocol which can be
effected with the detector of FIG. 5;
[0047] FIG. 7 is a simplified cross-sectional view of a spatially
sensitive radioactive emission probe, e.g., a gamma camera, used to
implement another embodiment of the present invention;
[0048] FIG. 8 is a presentation of a scanning protocol which can be
effected with the detector of FIG. 7;
[0049] FIG. 9 demonstrates a system in accordance with the
teachings of the present invention which employs four position
tracking systems for co-tracking the positions of a patient, a
radioactive emission probe, an imaging modality and a surgical
instrument;
[0050] FIG. 10 demonstrates the use of a pair of radiation emission
detectors connected therebetween via a connector, preferably a
flexible connector or a flexible connection to the connector
according to the present invention;
[0051] FIG. 11 is a schematic diagram of a surgical instrument and
accompanying system elements according to the teachings of the
present invention;
[0052] FIG. 12 is a simplified pictorial illustration of an imaging
system constructed and operative in accordance with a preferred
embodiment of the present invention, including a radiation probe
and position sensor, position tracking system, medical imaging
system and coordinate registration system;
[0053] FIG. 13 is a simplified pictorial illustration of a single
dimension image formation with a nuclear radiation probe attached
to a position tracking system of the system of FIG. 12, in
accordance with a preferred embodiment of the present
invention;
[0054] FIG. 14 is a simplified pictorial plot of detecting a
radiation point source with the nuclear radiation probe of the
system of FIG. 12, without further processing, in accordance with a
preferred embodiment of the present invention;
[0055] FIG. 15 is a simplified flow chart of an averaging algorithm
used in the imaging system of FIG. 12, in accordance with a
preferred embodiment of the present invention;
[0056] FIG. 16 is a simplified pictorial plot of detecting a
radiation point source with the nuclear radiation probe of the
system of FIG. 12, with averaging processing, in accordance with a
preferred embodiment of the present invention;
[0057] FIGS. 17 and 18 are simplified pictorial illustrations of
hot cross and hot bar phantom images, respectively, of images
produced by a gamma radiation probe of the system of FIG. 12;
[0058] FIG. 19 is a simplified flow chart of a minimizing algorithm
used in the imaging system of FIG. 12, in accordance with a
preferred embodiment of the present invention;
[0059] FIG. 20 is a simplified pictorial plot of detecting a
radiation point source with the nuclear radiation probe of the
system of FIG. 12, with minimizing processing, in accordance with a
preferred embodiment of the present invention;
[0060] FIG. 21 is a simplified pictorial illustration of an image
reconstruction system constructed and operative in accordance with
a preferred embodiment of the present invention, which produces a
combined image made up of medical images, the position of the peak
radiation location and the location of a therapeutic
instrument;
[0061] FIG. 22 is a simplified flow chart of a radiation map
reconstruction algorithm, in accordance with a preferred embodiment
of the present invention;
[0062] FIGS. 23A and 23B are illustrations of radiolabeled patterns
observed in images produced by the system of the invention and by a
conventional gamma camera, respectively, of an autonomous adenoma
of a thyroid;
[0063] FIGS. 24A and 24B are illustrations of radiolabeled patterns
observed in images produced by the system of the invention and by a
conventional gamma camera, respectively, of suspected Paget's
disease of a humerus;
[0064] FIGS. 25A and 25B are illustrations of radiolabeled patterns
observed in images produced by the system of the invention and by a
conventional gamma camera, respectively, of chronic
osteomyelitis;
[0065] FIGS. 26A and 26B are illustrations of radiolabeled patterns
observed in images produced by the system of the invention and by a
conventional gamma camera, respectively, of skeletal metastasis
from medulloblastoma;
[0066] FIGS. 27A-27I demonstrate the operation of an algorithm
provided by the present invention for estimating the distribution
of radiation sources in a control volume;
[0067] FIGS. 28A-28F schematically illustrate a handheld probe, in
accordance with preferred embodiments of the present invention;
[0068] FIGS. 29A-29B schematically illustrate the manner of
calibrating the handheld probe of FIGS. 28A-28F, in accordance with
a preferred embodiment of the present invention;
[0069] FIG. 30. schematically illustrates the manner of
synchronizing event and position readings of the handheld probe of
FIGS. 28A-28F, in accordance with a preferred embodiment of the
present invention;
[0070] FIGS. 31A-31C describe a spatial resolution bar-phantom test
of a radioactive emission probe, in accordance with a preferred
embodiment of the present invention;
[0071] FIGS. 32A-32D describe a spatial resolution bar-phantom test
of a prior-art probe;
[0072] FIGS. 33A-33B illustrate the energy resolution of a single
pixel of a radioactive emission probe, in accordance with the
present invention;
[0073] FIGS. 34A-34C illustrate endoscopic radioactive emission
probes, in accordance with the present invention;
[0074] FIG. 35 illustrate a method of calculating the depth of a
radiation source, in accordance with the present invention; and
[0075] FIG. 36 illustrate a two-dimensional image of the
radioactivity emitting source, produced by a free-hand scanning of
a cancerous prostate gland, ex vivo, in accordance with the present
invention.
DESCRIPTION OF THE PREFERRED EMBODIMENTS
[0076] The present invention relates to a radioactive emission
probe in communication with a position tracking system and the use
thereof in a variety of systems and methods of medical imaging and
procedures. Specifically, wide-aperture collimation-deconvolution
algorithms are provided, for obtaining a high-efficiency, high
resolution image of a radioactivity emitting source, by scanning
the radioactivity emitting source with a probe of a wide-aperture
collimator, and at the same time, monitoring the position of the
radioactive emission probe, at very fine time intervals, to obtain
the equivalence of fine-aperture collimation. The blurring effect
of the wide aperture is then corrected mathematically. Furthermore,
an imaging method by depth calculations is provided, based on the
attenuation of photons of different energies, which are emitted
from the same source, coupled with position monitoring.
[0077] The principles and operation of the present invention may be
better understood with reference to the drawings and accompanying
descriptions.
[0078] Before explaining at least one embodiment of the invention
in detail, it is to be understood that the invention is not limited
in its application to the details of construction and the
arrangement of the components set forth in the following
description or illustrated in the drawings. The invention is
capable of other embodiments or of being practiced or carried out
in various ways. Also, it is to be understood that the phraseology
and terminology employed herein is for the purpose of description
and should not be regarded as limiting.
[0079] Functional imaging, or the use of radioactive materials to
tag physiologically active tissue within the body of a patient, for
determining the tissue's localization and demarcation by
radioactive emission probes has been disclosed in the medical
literature for at least forty years. Functional imaging shows the
metabolic activity of body tissue, since dying or damaged body
tissue absorbs radiopharmaceuticals at a different rate from a
healthy tissue. The functional image may be used for example, to
study cardiac rhythm, or respiratory rhythm. However, a functional
image may not show structural, or anatomic details.
[0080] Significant developments in the localization and demarcation
of tissue bearing radioactive isotope tags for diagnostic and (or)
therapeutic purposes have occurred since that time. In fact, it is
now becoming an established practice in the diagnosis and (or)
treatment of certain diseases, e.g., cancer, blood clots,
myocardial infarct and abscesses, to introduce monoclonal
antibodies or other agents, e.g., fibrinogen, fluorodeoxyglucose
labeled with a radioactive isotope (e.g., .sup.99.sup.MTechnetium,
.sup.67Gallium, .sup.201Thallium, .sup.111Indium, .sup.123Iodine,
.sup.18Fluorine and .sup.125Iodine) into the body of the patient.
Such radiopharmaceuticals tend to localize in particular tissue or
cell type, whereas uptake or binding of the specific
radiopharmaceutical is increased in more "physiologically active"
tissue such as the active core of a cancerous tissue, so that the
radiation emitted following nuclear disintegrations of the isotope
can be detected by a radiation detector to better allocate the
active portion of a tumor. Such radiation may be, for example,
.alpha., .beta..sup.-, .beta..sup.+ and (or) .gamma. radiation.
[0081] In another type of applications radioactive substances are
used to determine the level of flow of blood in blood vessels and
the level of perfusion thereof into a tissue, e.g., coronary flow
and myocardial perfusion.
[0082] Referring now to the drawings, FIG. 1 illustrates a system
for calculating a position of a radioactivity emitting source in a
system-of-coordinates, in accordance with the teachings of the
present invention, which system is referred to hereinbelow as
system 20.
[0083] System 20 includes a radioactivity emission detector 22.
System 20 according to the present invention further includes a
position tracking system 24. System 24 is connected to and (or)
communicating with radioactive emission probe 22 so as to monitor
the position of detector 22 in a two- or three-dimensional space
defined by a system-of-coordinates 28 in two, three or more, say
four, five or preferably six degrees-of-freedom (X, Y, Z, .rho.,
.theta. and .phi.). System 20 further includes a data processor 26.
Data processor 26 is designed and configured for receiving data
inputs from position tracking system 24 and from radioactive
emission probe 22 and, as is further detailed below, for
calculating the image of the radioactivity emitting source in
system-of-coordinates 28. As shown in FIG. 10, a pair (or more) of
detectors 22 connected therebetween via a physical connector, each
of detectors 22 is position tracked, can be used for calculating
the image of the radioactivity emitting source in
system-of-coordinates 28. If more than a single detector 22 is
used, detectors 22 are preferably connected there between via a
connector 29. Connector 29 is preferably flexible. In the
alternative, the connections of detectors 22 to connector 29
provide the required flexibility.
[0084] It will be appreciated that system 20 of radioactivity
emission detector 22 and position tracking system 24 is inherently
different from known SPECT and PET imaging systems, as well as from
other imaging systems such as X-ray, Mammography, CT, and MRI,
since the motion of detector 22 of the present invention is not
limited to a predetermined track or tracks, with a respect to an
immovable gantry. Rather, detector 22 of the present invention is
adapted for a variable-course motion, which may be for example,
free-hand scanning, variable-course motion on a linkage system,
motion within a body lumen, endoscopic motion through a trocar
valve, or another form of variable-course motion.
[0085] Position tracking systems per se are well known in the art
and may use any one of a plurality of approaches for the
determination of position in a two- or three-dimensional space as
is defined by a system-of-coordinates in two, three and up to six
degrees-of-freedom. Some position tracking systems employ movable
physical connections and appropriate movement monitoring devices
(e.g., potentiometers) to keep track of positional changes. Thus,
such systems, once zeroed, keep track of position changes to
thereby determine actual positions at all times. One example for
such a position tracking system is an articulated arm.
[0086] FIG. 2 shows an articulated arm 30 which includes six arm
members 32 and a base 34, which can therefore provide positional
data in six degrees-of-freedom. Monitoring positional changes may
be effected in any one of several different ways. For example,
providing each arm member 32 with, e.g., potentiometers or optical
encoders 38 used to monitor the angle between adjacent arm members
32, to thereby monitor the angular change of each such arm member
with respect to adjacent arm members, so as to determine the
position in space of radioactive emission probe 22, which is
physically connected to articulated arm 30.
[0087] As is shown in FIG. 3 other position tracking systems can be
attached directly to radioactive emission probe 22 in order to
monitor its position in space. An example of such a position
tracking system is an assortment of three triaxially (e.g.,
co-orthogonally) oriented accelerometers 36 which may be used to
monitor the positional changes of radioactive emission probe 22
with respect to a space. A pair of such assortments, as is
specifically shown in FIG. 3, can be used to determine the position
of detector 22 in six-degrees of freedom.
[0088] As is shown in FIGS. 4 and 10, other position tracking
systems re-determine a position irrespective of previous positions,
to keep track of positional changes. Such systems typically employ
an array of receivers/transmitters 40 which are spread in known
positions in a system-of-coordinates and transmitter(s)/receiver(s)
42, respectively, which are in physical connection with the object
whose position being monitored. Time based triangulation and (or)
phase shift triangulation are used in such cases to periodically
determine the position of the monitored object, radioactive
emission probe 22 in this case. Examples of such a position
tracking systems employed in a variety of contexts using acoustic
(e.g., ultrasound) electromagnetic radiation (e.g., infrared, radio
frequency) or magnetic field and optical decoding are disclosed in,
for example, U.S. Pat. Nos. 5,412,619; 6,083,170; 6,063,022;
5,954,665; 5,840,025; 5,718,241; 5,713,946; 5,694,945; 5,568,809;
5,546,951; 5,480,422 and 5,391,199, which are incorporated by
reference as if fully set forth herein.
[0089] Radioactive emission probes are well known in the art and
may use any one of a number of approaches for the determination of
the amount of radioactive emission emanating from an object or
portion thereof. Depending on the type of radiation, such detectors
typically include substances which when interacting with
radioactive decay emitted particles emit either electrons or
photons in a level which is proportional over a wide linear range
of operation to the level of radiation impinging thereon. The
emission of electrons or photons is measurable and therefore serves
to quantitatively determine radiation levels. Solid-state detectors
in the form of N-type, P-type, PIN-type pixellated or unpixellated
include, for example, Ge, Si, CdTe, CdZnTe, CdSe, CdZnSe,
HgI.sub.2, TlBrI, GaAs, InI, GaSe, Diamond, TlBr, Pb.sub.2, InP,
ZnTe, HgBrI, a-Si, a-Se, BP, GaP, CdS, SiC, AlSb, PbO, BiI.sub.3
and ZnSe detectors. Gas (e.g., CO.sub.2 CH.sub.4) filled detectors
include ionization chamber detectors, proportional chamber
detectors and Geiger chamber detectors. Scintillation detectors
include organic scintillator crystals and liquids, such as
C.sub.14H.sub.10, C.sub.14H.sub.12, C.sub.10H.sub.8, etc.,
Plastics, NE102A, NE104, NE110, Pilot U and inorganic scintillator
crystals, such as NaI, CsI, BGO, LSO, YSO, BaF, ZnS, ZnO,
CaWO.sub.4 and CdWO.sub.4. Also known are scintillation fiber
detectors. Scintillator coupling include photomultiplier tube (PMT)
of the following types: side-on type, head-on type, hemispherical
type, position sensitive type, icrochannel plate-photomultiplier
(MCP-PMTs) and electron multipliers, or photodiodes (and
photodiodes arrays), such as Si photodiodes, Si PIN photodiodes, Si
APD, GaAs(P) photodiodes, GaP and CCD.
[0090] FIG. 5 shows a narrow angle or wide angle radioactive
emission probe 22'. Narrow or wide angle radioactive emission probe
22' includes a narrow slit (collimator) so as to allow only
radiation arriving from a predetermined angular direction (e.g.,
1.degree.-280.degree. wide angle, preferably
1.degree.-80.degree.--narrow angle) to enter the detector. Narrow
or wide angle radioactive emission probes especially suitable for
the configuration shown in FIG. 10 are manufactured, for example,
by Neoprobe, Dublin, Ohio (www.neoprobe.com), USA, Nuclear Fields,
USA (www.nufi.com) IntraMedical Imaging, Los Angeles, Calif., USA
(www.gammaprobe.com).
[0091] As is shown in FIG. 6, such a detector is typically used to
measure radioactivity, point by point, by scanning over the surface
of a radioactive object from a plurality of directions and
distances. In the example shown, scans from four different
directions are employed. It will be appreciated that if sufficient
radioactivity records are collected from different angles and
distances, and the orientation and position in space of detector
22' is simultaneously monitored and recorded during such scans, a
three-dimensional model of a radioactive region can be
reconstituted and its position in space determined. If two or more
detectors are co-employed, as shown in the configuration of FIG.
10, the results may be collected faster.
[0092] FIG. 7 shows another example of a radioactive emission
probe, a spatially sensitive (pixellated) radioactive emission
probe 22'' (such as a gamma camera). Detector 22'', in effect,
includes an array of multitude narrow angle detector units 23. Such
an arrangement is used in accordance with the teachings of the
present invention to reduce the amount of measurements and angles
necessary to acquire sufficient data so as to reconstitute a
three-dimensional model of the radioactive object. Examples of
spatially sensitive radioactive emission probes employed in a
variety of contexts are disclosed in, for example, U.S. Pat. Nos.
4,019,057; 4,550,250; 4,831,262; and 5,521,373; which are
incorporated by reference as if set forth herein. An additional
example is the COMPTON detector
(http://www.ucl.ac.uk/MedPhvs/posters/giulia/giulia.htm). FIG. 8
shows a scan optionally made by spatially sensitive radioactive
emission probe 22'' (such as a gamma camera).
[0093] A radioactive emission detector of particular advantages for
use in context of the present invention is the Compton gamma probe,
since, in the Compton gamma probe, spatial resolution is
independent of sensitivity and it appears possible to exceed the
noise equivalent sensitivity of collimated imaging systems
especially for systems with high spatial resolution. The Compton
probe is a novel type of gamma-probe that makes use of the
kinematics of Compton scattering to construct a source image
without the aid of mechanical collimators. Compton imaging
telescopes were first built in the 1970s for astronomical
observations [V. Schoenfelder et al., Astrophysical Journal 217
(1977) 306]. The first medical imaging laboratory instrument was
proposed in the early 1980s [M. Singh, Med. Phys. 10 (1983) 421].
The potential advantages of the Compton gamma probe include higher
efficiency, 3-D imaging without detector motion, and more compact
and lightweight system. In the Compton gamma probe, high-energy
gamma rays are scattered from a first detector layer (or detectors
array) into a second detector layer array. For each gamma, the
deposited energy is measured in both detectors. Using a line drawn
between these two detectors, the Compton scattering equation can be
solved to determine the cone of possible direction about this axis
on which the gamma ray must have entered the first detector. The
intersection of cones from many events is then developed to locate
gamma ray sources in the probe's field-of-view. Obviously only
coincident events are considered, and the more accurately their
energy can be determined, the less uncertainty there is in the
spatial angle of the arrival cone. The probe's electronic system is
combining coincidence measurements across many detectors and
detectors layers with a very good energy resolution. The choice of
the geometry and the material of the first layer detector plays a
major role in the system imaging capability and depends on (i)
material efficiency of single Compton events, in relation to other
interactions; (ii) detector energy resolution; and (iii) detector
position resolution. In particular, the overall angular resolution
results from the combination of two components, related to the
energy resolution and to the pixel volume of the detector.
[0094] Thus, as now afforded by the present invention, connecting a
radioactive emission probe to a position tracking system, permits
simultaneous radioactivity detecting and position tracking at the
same time. This enables the accurate calculation of the shape, size
and contour of the radiating object and its precise position in a
system-of-coordinates.
[0095] The present invention thus provides a method for defining a
position of a radioactivity emitting source in a
system-of-coordinates. The method is effected by (a) providing a
radioactive emission probe which is in communication with a
position tracking system; and (b) monitoring radioactivity emitted
from the radioactivity emitting source, while at the same time,
monitoring the position of radioactive emission probe in the
system-of-coordinates, thereby defining the image of the
radioactivity emitting source in the system-of-coordinates.
[0096] It will be appreciated by one of skills in the art that the
model produced by system 20 is projectable onto any of the other
systems-of-coordinates, or alternatively, the system-of-coordinates
defined by position tracking system 24 may be shared by other
position tracking systems, as is further detailed hereinbelow, such
that no such projection is required.
[0097] Thus, as is further shown in FIG. 1, system 20 of the
present invention can be used for calculating a position of a
radioactivity emitting source in a first system-of-coordinates 28
and further for projecting the image of the radioactivity emitting
source onto a second system-of-coordinates 28'. The system includes
radioactive emission probe 22, position tracking system 24 which is
connected to and (or) communicating with radioactive emission probe
22, and data processor 26 which is designed and configured for (i)
receiving data inputs from position tracking system 24 and from
radioactive emission probe 22; (ii) calculating the image of the
radioactivity emitting source in the first system-of-coordinates;
and (iii) projecting the image of the radioactivity emitting source
onto the second system-of-coordinates.
[0098] A method for calculating a position of a radioactivity
emitting source in a first system-of-coordinates and for projecting
the image of the radioactivity emitting source onto a second
system-of-coordinates is also offered by the present invention.
This method is effected by (a) providing a radioactive emission
probe being in communication with a position tracking system; and
(b) monitoring radioactivity being emitted from the radioactivity
emitting source, while at the same time, monitoring the position of
the radioactive emission probe in the first system-of-coordinates,
thereby defining the image of the radioactivity emitting source in
the first system-of-coordinates and projecting the image of the
radioactivity emitting source onto the second
system-of-coordinates.
[0099] It will be appreciated that the combination of a radioactive
emission probe and a position tracking system connected thereto and
(or) communicating therewith allows a suitable data processor to
generate a two- or three-dimensional image of the radioactivity
emitting source. An algorithm can be used to calculate image
intensity based on, for example, a probability function which
averages radiation counts and generates an image in which the
shorter the time interval between radioactive counts, the brighter
the image and vise versa, while down-compensating when a location
is re-scanned. A free-hand scanning with a directional detector can
be employed for this purpose.
[0100] In one embodiment, when scanning a body area with the
detector, the detector is made to follow a three-dimensional
surface, which defines the body curvature and in effect is used
also as a position tracking pointer. This information can be used
to define the position of the radioactive source with respect to
the outer surface of the body, so as to create a three-dimensional
map of both the radioactive source and of the body curvature. This
approach can also be undertaken in open surgeries, such as open
chest surgeries so as to provide the surgeon in real time with
information concerning the functionality of a tissue.
[0101] The radioactive emission probe, which can be used in context
of the present invention can be a beta emission detector, a gamma
emission detector, a positron emission detector or any combination
thereof. A detector that is sensitive to both beta and (or)
positron and gamma emission can be used to improve localization by
sensing for example gamma emission distant from the source and
sensing beta or positrons emission closer to the source. A beta
detector is dedicated for the detection of either electrons from
sources such as .sup.131Iodine, or positrons from sources such as
.sup.18Fluorine. A gamma detector can be designed as a single
energy detector or as a detector that can distinguish between
different types of energies, using the light intensity in the
scintillator as a relative measure of the gamma energy. Also, the
detector can be designed to utilize coincidence detection by using
detectors facing one another (180 degrees) with the examined organ
or tissue in-between. The radiation detector can have different
collimators with different diameters. A large bore will be used for
high sensitivity with lower resolution while a small bore
collimator will have higher resolution at the expense of lower
sensitivity.
[0102] Another possibility is to have a the collimator moving or
rotating with the opening eccentric so that a different solid angle
is exposed to the incoming photons at any one time, thus gathering
the photons from overlapping volumes at different time intervals.
The rest of the image processing is similar if the probe moves or
if the collimator eccentric opening moves.
[0103] System 20 of the present invention can be used in concert
with other medical devices, such as, but not limited to, any one of
a variety of imaging modalities and (or) surgical instruments.
[0104] Structural imaging modalities, which provide anatomic, or
structural maps of the body, are well known in the art. The main
modalities that serve for two-(projectional or cross sectional) or
three- (cosequtive cross sectional) dimensional imaging are a
planer X-ray imager, a fluoroscope, a computerized tomography
scanner, a magnetic resonance imager an ultrasound imager, an
impedance imager, and an optical camera.
[0105] Medical images taken of the human body are typically
acquired or displayed in three main orientations (i) coronal
orientation: in a cross section (plane), for example, across the
shoulders, dividing the body into front and back halves; (ii)
sagittal orientation: in a cross section (plane), for example, down
the middle, dividing the body into left and right halves; and (iii)
axial orientation: in a cross section (plane), perpendicular to the
long axis of the body, dividing the body into upper and lower
halves. Oblique views can also be acquired and displayed.
[0106] Various types of X-ray imaging are central to diagnosis of
many types of cancer. Conventional X-ray imaging has evolved over
the past 100 years, but the basic principal is still the same as in
1895, when first introduced. An X-ray source is turned on and
X-rays are radiated through the body part of interest and onto a
film cassette positioned under or behind the body part. The energy
and wavelength of the X-rays allows them to pass through the body
part and create the image of the internal structures like bones. As
the X-rays pass through the hand, for instance, they are attenuated
by the different density tissues they encounter. Bone attenuates a
great deal more of the X-rays than the soft tissue surrounding it
because of its grater density. It is these differences in
absorption and the corresponding varying exposure level of the film
that creates the images. In fact, X-ray imaging results in a
projection of the integrated density of column-voxels defined by
the X-rays as they pass through the body.
[0107] Fluoroscopy is a method based on the principals of film
X-ray that is useful for detecting disorders and tumors in the
upper gastro-intestinal (GI) system (for example, the stomach and
intestines). Fluoroscopic imaging yields a moving X-ray picture.
The physician can watch the screen and see an image of the
patient's body (for example the beating heart). Fluoroscopic
technology improved greatly with the addition of television cameras
and fluoroscopic "image intensifiers". Today, many conventional
X-ray systems have the ability to switch back and forth between the
radiographic and fluoroscopic modes. The latest X-ray systems have
the ability to acquire the radiograph or fluoroscopic movie using
digital acquisition.
[0108] Computed Tomography (CT) is based on the X-ray principal,
where the film is replaced by a detector that measures the X-ray
profile. Inside the covers of the CT scanner is a rotating frame
which has an X-ray tube mounted on one side and the detector
mounted on the opposite side. A fan beam of X-ray is created as the
rotating frame spins the X-ray tube and detector around the
patient. Each time the X-ray tube and detector make a 360.degree.
rotation, an image or "slice" has been acquired. This "slice" is
collimated to a thickness between 1 mm and 10 mm using lead
shutters in front of the X-ray tube and X-ray detector.
[0109] As the X-ray tube and detector make this 360.degree.
rotation, the detector takes numerous profiles of the attenuated
X-ray beam. Typically, in one 360.degree. lap, about 1,000 profiles
are sampled. Each profile is subdivided spatially by the detectors
and fed into about 700 individual channels. Each profile is then
backwards reconstructed (or "back projected") by a dedicated
computer into a two-dimensional image of the "slice" that was
scanned.
[0110] The CT gantry and table have multiple microprocessors that
control the rotation of the gantry, movement of the table (up/down
and in/out), tilting of the gantry for angled images, and other
functions such as turning the X-ray beam on an off. The CT contains
a slip ring that allows electric power to be transferred from a
stationary power source onto the continuously rotating gantry. The
innovation of the power slip ring has created a renaissance in CT
called spiral or helical scanning. These spiral CT scanners can now
image entire anatomic regions like the lungs in a quick 20 to 30
second breath hold. Instead of acquiring a stack of individual
slices which may be misaligned due to slight patient motion or
breathing (and lung/abdomen motion) in between each slice
acquisition, spiral CT acquires a volume of data with the patient
anatomy all in one position. This volume data set can then be
computer-reconstructed to provide three-dimensional models such as
of complex blood vessels like the renal arteries or aorta. Spiral
CT allows the acquisition of CT data that is perfectly suited for
three-dimensional reconstruction.
[0111] MR Imaging is superior to CT in detecting soft tissue
lesions such as tumors as it has excellent contrast resolution,
meaning it can show subtle soft-tissue changes with exceptional
clarity. Thus, MR is often the method of choice for diagnosing
tumors and for searching for metastases. MR uses magnetic energy
and radio waves to create single or consequtive cross-sectional
images or "slices" of the human body. The main component of most MR
systems is a large tube shaped or cylindrical magnet. Also, there
are MR systems with a C-shaped magnet or other type of open
designs. The strength of the MR systems magnetic field is measured
in metric units called "Tesla". Most of the cylindrical magnets
have a strength between 0.5 and 1.5 Tesla and most of the open or
C-shaped magnets have a magnetic strength between 0.01 and 0.35
Tesla.
[0112] Inside the MR system a magnetic field is created. Each total
MR examination typically is comprised of a series of 2 to 6
sequences. An "MR sequence" is an acquisition of data that yields a
specific image orientation and a specific type of image appearance
or "contrast". During the examination, a radio signal is turned on
and off, and subsequently the energy which is absorbed by different
atoms in the body is echoed or reflected back out of the body.
These echoes are continuously measured by "gradient coils" that are
switched on and off to measure the MR signal reflecting back. In
the rotating frame of reference, the net magnetization vector
rotate from a longitudinal position a distance proportional to the
time length of the radio frequency pulse. After a certain length of
time, the net magnetization vector rotates 90 degrees and lies in
the transverse or x-y plane. It is in this position that the net
magnetization can be detected on MRI. The angle that the net
magnetization vector rotates is commonly called the `flip` or `tip`
angle. At angles greater than or less than 90 degrees there will
still be a small component of the magnetization that will be in the
x-y plane, and therefore be detected. Radio frequency coils are the
"antenna" of the MRI system that broadcasts the RF signal to the
patient and (or) receives the return signal. RF coils can be
receive-only, in which case the body coil is used as a transmitter;
or transmit and receive (transceiver). Surface coils are the
simplest design of coil. They are simply a loop of wire, either
circular or rectangular, that is placed over the region of
interest.
[0113] A digital computer reconstructs these echoes into images of
the body. A benefit of MRI is that it can easily acquire direct
views of the body in almost any orientation, while CT scanners
typically acquire cross-sectional images perpendicular or nearly
perpendicular to the long body axis.
[0114] Ultrasound imaging is a versatile scanning technique that
uses sound waves to create images of organs or anatomical
structures in order to make a diagnosis. The ultrasound process
involves placing a small device called a transducer, against the
skin of the patient near the region of interest, for example,
against the back to image the kidneys. The ultrasound transducer
combines functions of emitting and receiving sound. This transducer
produces a stream of inaudible, high frequency sound waves which
penetrate into the body and echo off the organs inside. The
transducer detects sound waves as they echo back from the internal
structures and contours of the organs. Different tissues reflect
these sound waves differently, causing a signature which can be
measured and transformed into an image. These waves are received by
the ultrasound machine and turned into live pictures with the use
of computers and reconstruction software.
[0115] Ultrasound scanning has many uses, including: diagnosis of
disease and structural abnormalities, helping to conduct other
diagnostic procedures, such as needle biopsies etc.
[0116] There are limitations to some ultrasound techniques: good
images may not be obtained in every case, and the scan may not
produce as precise results as some other diagnostic imaging
procedures. In addition, scan results may be affected by physical
abnormalities, chronic disease, excessive movement, or incorrect
transducer placement.
[0117] Both two- (cross sectional) and three- (consequtive
cross-sectional) ultrasound imaging techniques are available
nowadays. Worth mentioning is the Dopler three-dimensional
ultrasound imaging.
[0118] In many cases imaging modalities either inherently include
(e.g., fluoroscope, CT, MRI) and (or) are integrated with
position-tracking-systems, which enable the use of such systems to
reconstruct three-dimensional image models and provide their
position in a system-of-coordinates.
[0119] It will be appreciated that, similar to the vision system,
also an optical camera can be used to generate three-dimensional
imagery date according to the present invention by imaging a body
from a plurality (at least two) directions. This type of imaging is
especially applicable in open chest surgeries or other open
surgeries. Software for calculating a three-dimensional image from
a pair of stereoscopic images is well known in the art.
[0120] Thus, as used herein and in the claims section that follows,
the phrase "three-dimensional imaging modality" refers to any type
of imaging equipment which includes software and hardware for
generating a three-dimensional image. Such an equipment can
generate a three-dimensional image by imaging successive
cross-sections of a body, e.g., as if viewed from a single
direction. Alternatively, such an equipment can generate a
three-dimensional image by imaging a body from different angles or
directions (typically two angles) and thereafter combining the data
into a three-dimensional image.
[0121] In accordance with the present invention, a structural
imaging probe, for example, an ultrasound probe, may be combined
with a position tracking system, in a manner analogous to System 20
of radioactivity emission detector 22 and position tracking system
24 (FIG. 1). Similarly, a miniature MRI probe, for example, as
taught by U.S. Pat. No. 5,572,132, to Pulyer, et al., entitled,
"MRI probe for external imaging," whose disclosure is incorporated
herein by reference, which teaches an MRI catheter for endoscopical
imaging of tissue of the artery wall, rectum, urinal tract,
intestine, esophagus, nasal passages, vagina and other biomedical
applications, may be used.
[0122] Additionally, a structural imaging probe, for example, an
ultrasound probe, may be combined with system 20 of radioactivity
emission detector 22 and position tracking system 24 (FIG. 1), so
as to have:
[0123] i. structural imaging;
[0124] ii. functional imaging; and
[0125] iii. position tracking,
simultaneously. Similarly, other structural imaging modalities, as
known, may be combined with system 20 of FIG. 1.
[0126] Furthermore, the structural imaging modality may have an
additional, dedicated position tracking system, or share position
tracking system 24 of radioactivity emission detector 22.
[0127] Position tracking may also be accomplished by using imaging
information from the structural imaging system. This may be
accomplished by tracking relative changes from one image to another
to determine relative motion.
[0128] Surgical instruments are also well known in the art and may
use any one of a plurality of configurations in order to perform
minimally-invasive surgical procedures. Examples include laser
probes, cardiac and angioplastic catheters, endoscopic probes,
biopsy needles, aspiration tubes or needles, resecting devices,
ultrasonic probes, fiber optic scopes, laparoscopy probes, thermal
probes and suction/irrigation probes. Examples of such surgical
instruments employed in a variety of medical contexts are disclosed
in, for example, U.S. Pat. Nos. 6,083,170; 6,063,022; 5,954,665;
5,840,025; 5,718,241; 5,713,946; 5,694,945; 5,568,809; 5,546,951;
5,480,422 5,391,199, 5,800,414; 5,843,017; 6,086,554; 5,766,234;
5,868,739; 5,911,719; 5,993,408; 6,007,497; 6,021,341; 6,066,151;
6,071,281; 6,083,166 and 5,746,738, which are incorporated by
reference as if fully set forth herein.
[0129] For some applications, examples of which are provided in the
list of patents above, surgical instruments are integrated with
position-tracking-systems, which enable to monitor the position of
such instruments while placed in and guided through the body of a
treated patient.
[0130] According to a preferred embodiment of the present
invention, the surgical instrument is equipped with an additional
radioactive emission probe attached thereto or placed therein. This
additional detector is used, according to preferred embodiments of
the invention, to fine tune the location of radioactive emission
from within the body, and in closer proximity to the radioactive
source. Since the surgical tool is preferably in communication with
a position-tracking system, the position of the additional detector
can be monitored and its readouts used to fine tune the position of
the radioactive source within the body. Thus, according to this
aspect of the present invention, at least one extracorporeal
detector and an intracorporeal detector are used in concert to
determine the position of a radioactive source in the body in
highest precision. The extracorporeal detector provides the general
position of the source and is used for directing the surgical
instrument thereto, whereas the intracorporeal detector is used for
reassuring prior to application of treatment or retrieval of biopsy
that indeed the source was correctly targeted at the highest
precision.
[0131] While according to a presently preferred embodiment of the
invention two detectors, one extracorporeal and one intracorporeal,
are employed as described above, for some applications a single
intracorporeal detector may be employed, which detector is attached
to or integrated with a surgical instrument whose position is
tracked.
[0132] The use of intracorporeal and extracorporeal detectors calls
for careful choice of the radioactive isotope employed with the
radiopharmaceutical. While the extracorporeal detector can be
constructed with a suitable collimator for handling strong
radiation, such as gamma radiation, the intracorporeal detector is
miniature by nature and is limited in design and construction by
the construction of the surgical instrument with which it is
employed. Since collimators for high energy (80-511 KeV) gamma
radiation are robust in nature, they are not readily engageable
with miniature detectors. Electron (beta) and positron radiation
are characterized by: (i) they highly absorbed by biological tissue
as they are of lower energy and higher chemical reactivity; and
(ii) they are readily collimated and focused by thin metal
collimators. It is also possible to use low energy gamma radiation
(10-30 KeV) for intracorporeal applications since the collimation
of these gamma photons can be achieved with thin layers of Tantalum
or Tungsten. As such, the radio pharmaceutical of choice is
selected to emit both gamma and beta and (or) positron radiation,
whereas the extracorporeal detector is set to detect the high
energy gamma radiation, whereas the intracorporeal detector is set
to detect the low energy gamma, beta and (or) positron radiation.
Isotopes that emit both high energy gamma and (or) low energy
gamma, beta and (or) positron radiation and which can be used per
se or as a part of a compound as radiopharmaceuticals include,
without limitation, .sup.18F, .sup.11 In and .sup.123I in
radiopharmaceuticals, such as, but not limited to,
2-[.sup.18F]fluoro-2-deoxy-D-glucose (.sup.18FDG),
.sup.111In-Pentetreotide
([.sup.111In-DTPA-D-Phe.sup.1]-octreotide),
L-3-[.sup.123I]-Iodo-alpha-methyl-tyrosine (IMT),
O-(2-[.sup.18F]fluoroethyl)-L-tyrosine (L-[.sup.18F]FET),
.sup.111In-Capromab Pendetide (CYT-356, Prostascint) and
.sup.111In-Satumomab Pendetide (Oncoscint).
[0133] FIG. 11 illustrates a system in accordance with this aspect
of the present invention. A surgical instrument 100 is shown
connected to a resection/aspiration control element 102 as well
known in the art. Surgical instrument 100 includes a radioactive
emission probe 104, which has a collimator 106 for collimating low
energy gamma, beta and (or) positron radiation. In some
embodiments, as indicated by arrow 108, detector 104 may be
translated within instrument 100. A position tracking system having
one element thereof 110 attached to instrument 100 and another
element thereof 112 at a fixed location serves to monitor the
position of instrument 100 at all times in two, three and up to six
degrees of freedom. Radioactive emission probe 104 communicates
with a counter 114 for counting low energy gamma, beta and (or)
positron radiation. All the data is communicated to, and processed
by, a processor 116. The 2D or 3D data may be projected and
displayed along with 2D or 3D imaging data derived from an imaging
modality using a shared presentation device as described elsewhere
herein. A real or virtual image of the surgical instrument itself
may also be co-displayed. Examples of commercially available
radiation emission detectors that can fit inside, for example, a
biopsy needle include scintillating plastic optical fibers like
S101 and S104, manufactured by PPLASTIFO or an optical fiber
communicating with a scintillator (either detector paint or
scintillation crystal) at the fiber edge. The level of detected
radiation can be reported visually or by an audio signal, as is
well known in the art.
[0134] Thus, a surgical instrument equipped with a radiation
emission detector and which is connected to and (or) communicating
with a position tracking system forms one embodiment of this aspect
of the present invention. Such a design acting in concert with
either conventional imaging modalities and (or) extracorporeal
radiation emission detectors form other embodiments of this aspect
of the invention. In all cases, a surgical instrument equipped with
a radiation emission detector and which is connected to and (or)
communicating with a position tracking system serves for in situ
fine tuning of a radioactive source in the body.
[0135] It will be appreciated that in some minimally-invasive
procedures even the position of the patient him or herself is
monitored via a position tracking system, using, for example,
electronic or physical fiducial markers attached at certain
locations to the patient's body.
[0136] Thus, as is further detailed hereinbelow, by projecting the
three-dimensional data and positions received from any of the above
mentioned devices into a common system of coordinates, or
alternatively, employing a common position tracking system for all
of these devices, one can integrate the data into a far superior
and comprehensive presentation.
[0137] An example to this desired outcome is shown in FIG. 9. In
the embodiment shown, four independent position tracking systems
50, 52, 54 and 56 are used to track the positions of a patient 58,
an imaging modality 60, a radioactive emission probe 62 and a
surgical instrument 64 in four independent systems-of-coordinates
66, 68, 70 and 72, respectively. If the patient is still, no
tracking of the patient's position is required.
[0138] It will be appreciated that any subset or all of the
position tracking systems employed may be integrated into one or
more common position tracking systems, and (or) that any subset or
all of the position tracking systems employed may share one or more
systems-of-coordinates, and further that any positional data
obtained by any of the position tracking systems described in any
of the systems-of coordinates may be projected to any other system
of coordinates or to an independent (fifth) system of coordinates
74. In one preferred embodiment, applicable for applications at the
torso of the patient, the system of coordinates is a dynamic system
of coordinates which takes into account the chest breathing
movements of the patient during the procedure.
[0139] As indicated at 76, the raw data collected by detector 62 is
recorded and, as indicated at 78, the position and the radioactive
data records are used to generate a three-dimensional model of a
radiopharmaceutical uptaking portion of a body component of the
patient.
[0140] Similarly, as indicated at 80, the imagery data collected by
imaging modality 60 is recorded and the position and the imagery
data records are used to generate a three-dimensional model of the
imaged body component of the patient.
[0141] All the data collected is then fed into a data processor 82
which processes the data and, as indicated at 84, generates a
combined or superimposed presentation of the radioactive data and
the imagery data, which is in positional context with patient 58
and surgical instrument 64.
[0142] Instrument 64, which by itself can be presented in context
of the combined presentation, may then be used to perform the
procedure most accurately. Processor 82 may be a single entity or
may include a plurality of data processing stations which directly
communicate with, or even integral to, any one or more of the
devices described.
[0143] Additionally or alternatively, a structural imaging probe,
for example, an ultrasound probe, or another structural probe, as
known, may be incorporated with the surgical instrument.
[0144] The present invention provides a major advantage over prior
art designs because it positionally integrates data pertaining to a
body portion as retrieved by two independent imaging techniques,
conventional imaging and radioactive imaging, to thereby provide a
surgeon with the ability the fine point the portion of the body to
be sampled or treated.
[0145] It will be appreciated that subsets of the devices described
in FIG. 9 may be used as stand-alone systems. For example, a
combination of detector 62 with its position-tracking system and
instrument 64 with its position-tracking-system may in some
instances be sufficient to perform intrabody procedures. For mere
diagnostic purposes, without biopsy, a combination of detector 62
position-tracking-system and modality 60 position-tracking-system
are sufficient.
[0146] Reference is now made to FIG. 12, which illustrates an
imaging system 200 constructed and operative in accordance with a
preferred embodiment of the present invention. Imaging system 200
preferably includes a radiation probe 202, such as the narrow angle
radioactive emission probe 22' described hereinabove with reference
to FIGS. 5 and 10.
[0147] A position sensor 204 is provided for sensing the position
of radiation probe 202. Position sensor 204 may be physically
attached to radiation probe 202, or may be distanced therefrom.
Position sensor 204 transmits the sensed position data to a
position tracking system 206. Position tracking system 206 may be a
system like position tracking system 24, described hereinabove with
reference to FIG. 1, and position sensor 204 may be any kind of
sensor applicable for such position tracking systems.
[0148] Another method which can be used to locate the source of
radiation emission is by using a small hand held gamma camera 205
(such as the DigiRad 2020tc Imager TM, 9350 Trade Place, San Diego,
Calif. 92126-6334, USA), attached to position sensor 204.
[0149] Position tracking system 206 enables radiation probe 202 to
freely scan back and forth in two- or three-dimensions over the
area of interest of the patient, preferably incrementing a short
distance between each scan pass. Position tracking system 206
tracks the position of radiation probe 202 with respect to a
position tracking coordinate system, such as X.sub.p, Y.sub.p and
Z.sub.p, with an origin O.sub.p.
[0150] Imaging system 200 also includes a medical imaging system
208, such as, but not limited to, computed or computerized
tomography (CT), magnetic resonance imaging (MRI), ultrasound
imaging, positron emission tomography (PET) and single photon
emission computed tomography (SPECT), for example. Medical imaging
system 208 provides images of a patient 209 with respect to a
medical imaging coordinate system, such as X.sub.m, Y.sub.m and
Z.sub.m, with an origin O.sub.m.
[0151] Imaging system 200 also includes a coordinate registration
system 210, such as that described in commonly owned, now
abandoned, U.S. patent application Ser. No. 09/610,490, entitled,
"method for registering coordinate systems," the disclosure of
which is incorporated herein by reference. Coordinate registration
system 210 is adapted to register the coordinates of the position
tracking coordinate system with those of the medical imaging
coordinate system.
[0152] Position tracking system 206, medical imaging system 208 and
coordinate registration system 210 are preferably in wired or
wireless communication with a processing unit 212 (also referred to
as a data processor 212).
[0153] In operation of imaging system 200, after administration of
a radiopharmaceutical to patient 209, a clinician/physician/surgeon
(not shown) may move or scan radiation probe 202 about a target
area under examination. A physiological activity map of the target
area is obtained by measuring the radiation count rate with
radiation probe 202, and by correlating the count rate with the
count rate direction with position tracking system 206, which
follows the motion of the moving or scanning radiation probe
202.
[0154] Reference is now made to FIG. 13, which illustrates image
formation with radiation probe 202, in accordance with a preferred
embodiment of the present invention. For the purposes of
simplicity, the example shown in FIG. 13 is for a single dimension
image formation, but it is readily understood that the same
principles hold true for any other dimensional image formation.
[0155] In one example of carrying out the invention, radiation
probe 202 may be a gamma ray detector probe that comprises a
collimator 211 and radiation detector 213. Gamma rays are projected
through the probe collimator 211 onto radiation detector 213, which
produces electronic signals in accordance with the radiation
detected. Radiation probe 202 sends pulses to a probe counter 215
which may include a pulse height analyzer circuit (not shown). The
pulse height analyzer circuit analyzes the electronic signals
produced by radiation detector 213. If the electronic signals are
within a selected energy window, the level of radiation, i.e.,
number of radiation counts, is counted by probe counter 215.
[0156] Examples of suitable radiation detectors include a solid
state detector (SSD) (CdZnTe, CdTe, HgI, Si, Ge, and the like), a
scintillation detector (NaI(Tl), LSO, GSO, CsI, CaF, and the like),
a gas detector, or a scintillating fiber detector (S101, S104, and
the like).
[0157] The position sensor 204 associated with the radiation probe
202 senses the position of radiation probe 202, and position
tracking system 206 calculates and monitors the motion of radiation
probe 202 with respect to the position tracking coordinate system.
The motion is calculated and monitored in two, three and up to six
dimensions--the linear directions of the X, Y and Z axes as well as
rotations about the X, Y and Z axes, i.e., rotational angles .rho.,
.theta. and .phi., respectively.
[0158] Examples of suitable position tracking systems include a
measurement mechanical arm (FaroArm,
http://www.faro.com/products/faroarm.asp), optical tracking systems
(Northern Digital Inc., Ontario, Canada NDI-POLARIS passive or
active systems), magnetic tracking systems (NDI-AURORA), infrared
tracking systems (E-PEN system, http://www.e-pen.com), and
ultrasonic tracking systems (E-PEN system), for example.
[0159] Processing unit 212 combines the radiation probe count rate
from probe counter 215 together with the positional information
from position tracking system 206, and uses an imaging software
algorithm 217 to form a two-dimensional or three-dimensional
radiotracer-spread image of the target area inside the patient's
body. The spatial probe positions together with the spatial count
rates may be stored in memory or displayed on a computer monitor
214 as a pattern of marks corresponding to the spatial and count
rate position.
[0160] An example of such a pattern is shown in FIG. 14, which
illustrates a single-dimensional, unprocessed simulation of a
radiation point source 218 (FIG. 13), 30 mm deep inside the human
body, detected by using a 10 mm nuclear radiation probe 202 coupled
to position tracking system 206. The graph of FIG. 14 indicates to
a physician that there is a peak count rate of about 500 in the
probe position of about 50 mm.
[0161] In one embodiment of the invention, the imaging software
algorithm 217 employs an averaging process to refine the curve of
FIG. 14. This averaging process will now be described with
reference to FIG. 15.
[0162] Probe counter 215 feeds probe count rate information N(Xc,
Yc, Zc, .rho., .theta., .phi. to processing unit 212 (step 301).
Position sensor 204 feeds probe position information (Xc, Yc, Zc,
.rho., .theta., .phi.) to processing unit 212 (step 302). Probe
parameters (such as its physical size, dx, dy, dz) are also input
into processing unit 212 (step 303).
[0163] Processing unit 212 then finds all the voxels (i.e., volume
pixels) that represent the probe volume in the processing unit
memory (step 304), i.e., Xc+dx, Yc+dy, Zc+dz. Processing unit 212
calculates the number of times that the calculation process has
been done in each voxel from the beginning of the image formation
(step 305), i.e., M(Xc+dx, Yc+dy, Zc+dz). Processing unit 212 then
calculates the new average count rate values in each voxel (step
306), in accordance with the formula:
N(Xc+dx, Yc+dy, Zc+dz)=[N(Xc+dx, Yc+dy, Zc+dz)+N(Xc, Yc, Zc, .rho.,
.theta., .phi.)]/[M (Xc+dx, Yc+dy, Zc+dz)+1]
[0164] Processing unit 212 then corrects the display image that
represents the perceived voxels at N(Xc+dx, Yc+dy, Zc+dz) (step
307). The algorithm then repeats itself for the next probe position
(step 308).
[0165] The resulting graph of the averaging algorithm of FIG. 15,
as applied to the example of FIG. 14, is shown in FIG. 16.
[0166] FIGS. 17 and 18 respectively show examples of a hot cross
phantom image and a hot 4.77 mm bar phantom image, produced by a
gamma radiation probe coupled with position tracking system 206 and
the averaging algorithm of FIG. 15. The probe images were made by
using a probe, EG&G Ortec NaI(Tl) model 905-1 (thickness=1'',
diameter=1'') connected to a ScintiPack model 296. The position
tracking system used was the Ascension miniBIRD, commercially
available from Ascension Technology Corporation, P.O. Box 527,
Burlington, Vt. 05402 USA
(http://www.ascension-tech.com/graphic.htm). The magnetic tracking
and location systems of Ascension Technology Corporation use DC
magnetic fields to overcome blocking and distortion from nearby
conductive metals. Signals pass through the human body without
attenuation.
[0167] In another embodiment of the invention, the imaging software
algorithm 217 may employ a minimizing process to refine the curve
of FIG. 14 as is now described with reference to FIG. 19.
[0168] Probe counter 215 feeds probe count rate information N(Xc,
Yc, Zc, .rho., .theta., .phi.) to processing unit 212 (step 401).
Position sensor 204 feeds probe position information (Xc, Yc, Zc,
.rho., .theta., .phi.) to processing unit 212 (step 402). Probe
parameters (such as its physical size, dx, dy, dz) are also input
into processing unit 212 (step 403).
[0169] Processing unit 212 then finds all the voxels that represent
the probe volume in the processing unit memory (step 404), i.e.,
Xc+dx, Yc+dy, Zc+dz. From the voxels that represent the probe
volume in the processing unit memory, processing unit 212 finds
those that have a higher count rate value N(Xc+dx, Yc+dy, Zc+dz)
than the inputted probe count rate N(Xc, Yc, Zc, .rho., .theta.,
.phi.) (step 405). Processing unit 212 then changes the higher
count rate voxels to that of inputted probe count rate N(Xc, Yc,
Zc, .rho., .theta., .phi.) (step 406), and corrects the display
image at the higher count rate voxels N(Xc+dx, Yc+dy, Zc+dz) (step
407). The algorithm then repeats itself for the next probe position
(step 408).
[0170] The resulting graph of the minimizing algorithm of FIG. 19,
as applied to the example of FIG. 14, is shown in FIG. 20.
[0171] Reference is now made to FIG. 21, which illustrates an image
reconstruction system 450, constructed and operative in accordance
with a preferred embodiment of the present invention. Image
reconstruction system 450 produces a combined image 451 made up of
the images coming from the medical imaging system 208 with the
position of the peak radiation location (and its uncertainty area)
from processing unit 212, together with the location of a
therapeutic instrument 452, such as a biopsy needle. The combined
image 451 allows the physician to better assess the relative
position of therapeutic instrument 452 in relation to the
anatomical image (from medical imaging system 208) and the position
of the radioactive area as inferred by the radiation detection
algorithm.
[0172] In accordance with preferred embodiments of the present
invention, image acquisition and reconstruction algorithms are
provided, for image acquisition with wide-aperture collimation and
image reconstruction which includes deconvolution, for resolution
enhancement. The overall algorithms are herethereto referred to as
wide-aperture collimation-deconvolution algorithms.
[0173] In essence, the wide-aperture collimation-deconvolution
algorithms enable one to obtain a high-efficiency, high resolution
image of a radioactivity emitting source, by scanning the
radioactivity emitting source with a probe of a wide-aperture
collimator, and at the same time, monitoring the position of the
radioactive emission probe, at very fine time intervals, to obtain
the equivalence of fine-aperture collimation. The blurring effect
of the wide aperture is then corrected mathematically.
[0174] The wide-aperture collimation-deconvolution algorithms are
described hereinbelow, in conjunction with FIGS. 27A-27I, and FIG.
22. Experimental results, showing images produced by the
wide-aperture collimation-deconvolution algorithms of the present
invention, in comparison to a conventional gamma camera are seen in
FIGS. 23A-26B, and FIG. 36, hereinbelow.
[0175] Referring further to the drawings, FIGS. 27A-27I describe
wide-aperture collimation-deconvolution algorithms, in accordance
with preferred embodiments of the present invention. The
wide-aperture collimation-deconvolution algorithms are designed for
estimating the distribution of radiation sources in a control
volume, thus constructing an image of the radiation sources in the
control volume. For simplicity, it is assumed that the radiation
sources comprise dot sources that radiate uniformly in all
directions, are localized, and are smoothly distributed in the
control volume.
[0176] Consider a single-pixel, wide-bore-collimator probe 707,
seen in FIG. 27H, formed for example, as a lead tube collimator
708, having a length, L.sub.collimator of about 20 mm and a
radiation detector 706, for example, a solid state CdZnTl crystal,
at its base, with a diameter, D.sub.detector of about 10 mm, so
that the ratio of L.sub.collimator to D.sub.detector is about 2. An
angular opening .delta. of wide-bore-collimator probe 707 may be
about 40 degrees; wide-bore-collimator probe 707 is sensitive to
any incident photons within the confining of about a 40-degree
angular opening. It will be appreciated that single-pixel,
wide-bore-collimator probes of other dimensions and ratios of
L.sub.collimator to D.sub.detector are similarly possible. For
example, the ratio of L.sub.collimator to D.sub.detector may be
about 1, or about 3.
[0177] By contrast, as seen in FIG. 27I, conventional gamma cameras
have a multi-pixel, small-aperture collimation probe, wherein a
collimator-pixel arrangement 709 is about 2 mm in diameter and 40
mm in length, so that the ratio of L.sub.collimator to
D.sub.detector is about 20. An angular opening c of the
small-aperture collimation probe is very small. In consequence,
wide-bore-collimator probe 707 may be about 30 times more sensitive
than the small-aperture collimation probe of collimator--pixel
arrangement 709.
[0178] Yet, there are disadvantages to wide-bore-collimator probe
707: [0179] 1. a single-pixel probe does not lend itself to the
generation of an image; and [0180] 2. a wide aperture blurs the
information regarding the direction from which the radiation
comes.
[0181] With regard to the first point, the generation of an image,
as wide-bore-collimator probe 707 is moved in space, while its
position and angular orientation are accurately tracked, at very
short time intervals, for example, of about 100 or 200 ms, or any
other short time interval, a one-, two-, or three-dimensional image
of counting rate as a function of position can be obtained, by data
processing.
[0182] More specifically, knowing the position and orientation of
probe 707 at each time interval and the photon count rate at that
position and orientation can be used to reconstruct the radiation
density of an unknown object, even in three dimensions.
[0183] With regard to the second point, the blurring effect of the
wide-angle aperture, while it is known that the information is
blurred, or convolved, by the wide-aperture collimator, a
deconvolution process may be used to obtain dependable results.
Moreover, the convolving function for a wide aperture depends only
on the geometry of the collimator and may be expressed as a set of
linear equations that can be readily solved.
[0184] Thus, in accordance with the present invention, a wide-bore
collimator probe, having a position tracking system, operating at
very short time intervals, and connected to a data processor for
wide-aperture collimation-deconvolution algorithms, can be used for
the construction of an image of radiation density in one-, two-, or
three-dimensions, with high degree of accuracy.
[0185] It will be appreciated that similar analyses may be employed
for wide aperture probes having non-circular collimators. For
example, the wide-aperture probe may have a rectangular collimator
of a single cell or of a coarse grid of cells. Additionally, the
collimator may be a wide-angle collimator. A corresponding
deconvolution function may be used by the data processor.
[0186] Reference is now made to FIGS. 27A and 27B, which illustrate
a radiation sensor 600, preferably generally shaped as a tube
collimator. Radiation quanta 602 are registered by the radiation
sensor 600, as described hereinabove, thereby providing the average
number of quanta per unit time. The radiation sensor 600 may be
moved around a volume of interest 604, preferably, with
six-dimensional freedom of motion, and six dimensional monitoring.
However, it will be appreciated that a more limited motion and
corresponding monitoring, for example, only in the X and .rho.
directions is possible. The position of the sensor 600 and its
direction (as well as the position of the investigated volume 604)
are assumed to be known at any given moment (FIG. 27A).
[0187] The tube collimator is preferably provided with a plane
circular detector 606 of radiation quanta. The quanta detector 606
is preferably disposed on a rear end 608 of the tube and radiation
quanta can reach the detector 606 only through an open front end
610 of the tube (FIG. 27B)
[0188] Reference is now made to FIG. 27C, which illustrates a
system of coordinates (x, y, z) with the origin O in the center of
the radiation sensor 600, the (x, y) plane being the plane of the
detector, and the z axis being in the center of the collimator
tube. The geometry parameters of the collimator tube--height h and
radius .rho.--are known.
[0189] From the rotational symmetry of the tube, it is clear that
having a radiation source Q=Q(x, y, z) with the total intensity I
uniformly radiating in all directions, the portion of the intensity
registered by the quanta detector 606 of the radiation sensor 600
is determined only by the distance r from Q to the axis of the
collimator (axis z) and the distance z from Q to the (x, y) plane.
In other words, there is a function (r, z), which is defined only
by the collimator parameters .rho. and h (corresponding expressions
from .rho., h, r and z may be easily written in explicit form),
such that the intensity of the radiation spot Q=Q(x, y, z)=Q(r, z)
registered by detector 606 is proportional to (r, z) and to the
total intensity I of the radiation spot.
[0190] Reference is now made to FIG. 27D. It follows from the
foregoing discussion, that if instead of one radiation spot there
is a radiation distribution I(Q)=I(Q(r, z)) in a volume V, then the
radiation intensity, registered by radiation sensor 600, is
proportional (with some constant not depending on the radiation
distribution and the sensor position) to the following
integral:
.intg. V I ( Q ( r , z ) ) .PHI. ( r , z ) Q ( 1 ) ##EQU00001##
[0191] An algorithm for estimating the intensity distribution I(Q)
from the values obtained in the measurement scheme of Equation (1)
is now discussed. For the sake of simplicity, the first case is
discussed with reference to FIG. 27E for a two-dimensional problem,
wherein intensities I(Q) are distributed in some 2-dimensional
plane. The 3D problem is a direct generalization of the
corresponding 2D problem, as is explained hereinbelow.
[0192] As seen in FIG. 27E, the radiation sources are distributed
in a rectangular region V in a plane. Two systems of coordinates
are considered. The first one is the sensor coordinate system (x,
y, z) corresponding to the sensor 600. The second one is the
radiation source coordinate system (u, v, w) corresponding to the
radiation sources plane (u, v).
[0193] It is assumed that at each discrete time increment, the
position of the origin of (x, y, z) system and the direction of the
z-axis unit vector in (u, v, w) coordinates are known. In other
words, the position and direction of the moving sensor in the (u,
v, w) coordinate system is known, and the (u, v, w) coordinate
system is assumed to be motionless.
[0194] The radiation sources are considered to be distributed in
accordance with the distribution function I(Q)) in some bounded,
given rectangle V on the plane (u, v). I(Q)=I(u, v) is the unknown
and sought-for radiation (or radiation intensity) distribution
function defined in V.
[0195] To regularize the problem of estimation of the radiation
distribution function I(Q), the function I(Q) will be considered to
be given from some finite dimensional space H of functions defined
in V. In other words, the function I(Q) itself will not be
estimated but rather some finite dimensional approximation of the
distribution I(Q).
[0196] The simplest approach to finite dimensional approximation is
to subdivide the rectangle V into sets of equal rectangular cells
and consider the space H of step-functions corresponding to this
subdivision (i.e., the space of functions that are constant in the
cells of subdivision), as shown in FIG. 27F.
[0197] If the subdivision of rectangle V into small rectangles is
sufficiently fine, then this step-function approximation is good
enough for the estimation of radiation distribution I(Q).
[0198] Let each side of rectangle V be divided into n equal parts
(FIG. 27F). Then m=n.sup.2 is the dimension of the space H of
step-functions on the corresponding subdivision.
[0199] The space H is naturally isomorphic to the m-dimensional
space of n.times.n matrixes (with its natural scalar product
<.cndot., .cndot.>).
[0200] Let I=(I.sub.i j).sub.i, j=1, . . . , n be the unknown
element of H that it is desired to estimate. Suppose that element I
is measured on K functionals {.PHI..sub.k}.sub.k=1 . . . K of the
form of the integral (1):
<I,.PHI..sub.k>=.SIGMA..sub.i,j=1 . . .
nI.sub.ij.PHI..sub.ij.sup.(k) (2)
[0201] where .PHI..sub.k=(.PHI..sub.i j.sup.(k)).sub.i, j=1, . . .
, n, k=1, . . . , K (after approximation of function I(Q), by the
corresponding step-function, the integral (1) is transformed to the
sum (2)).
[0202] Functionals .PHI..sub.k, k=1, . . . , K, correspond to K
discrete positions of the sensor (FIG. 27E). Knowing the explicit
expressions for functions (r, z) from (1) and knowing for each time
moment k, the position of the sensor relative to inspection region
V, one can calculate all matrixes .PHI..sub.k=(.PHI..sub.i j
.sup.(k)).sub.i, j=1, . . . , n, k=1, . . . , K.
[0203] Accordingly the following scheme of measurements are
obtained:
.psi..sub.k=<I,.PHI..sub.k>+.epsilon..sub.k,k=1, . . . ,K.
(3)
[0204] Here .psi..sub.k are results of measurements of the unknown
element I of the space H, and .epsilon..sub.k are random errors
(.epsilon..sub.k-- independent random variables,
E.epsilon..sub.k=0, k=1, . . . , K).
[0205] Let M: H.fwdarw.H the operator in the space H of the
form:
M=.SIGMA..sub.k=1 . . . K.PHI..sub.k.PHI..sub.k. (4)
[0206] Then the best non-biased linear estimate I of the element I
is given by the formula:
I=M.sup.-1.PSI., (5)
[0207] where M.sup.-1: H.fwdarw.H the inverse operator to the
operator M of the form (4), and
.PSI.=.SIGMA..sub.k=1 . . . K.psi..sub.k.PSI..sub.k, (6)
[0208] (where .psi..sub.k are the results of measurements of the
form (3)).
[0209] One problem of using estimates (5) (besides computational
problems if the dimension m of the space H is very large) is that
the operator M: H.fwdarw.H of the form (3) is "bad invertible". In
other words, the estimation problem is "ill-posed". It means that
having a noise .epsilon..sub.k in the measurements scheme (3), even
if the noise is small, may sometimes result in a very large
estimation error dist(I, I).
[0210] This means that the estimation problem requires additional
regularization. This is a general problem of solving a large set of
linear equations. There are several methods for solving such
equations. Below is described one of the known methods for solving
such equations but numerous other methods are also possible, theses
include gradient decent methods such as in
(http://www-visl.technion.ac.il/1999/99-03/www/) and other methods
that are generally known in the art. Further, it is possible to
improve the image reconstruction by taking into account the
correlation between measurements as they are done with substantial
overlap. Also, in the following description, a regular step
function is assumed for the representation of the pixels or voxels,
other basis may be used such as wavelet basis, Gaussian basis,
etc., which may be better suited for some applications.
[0211] To obtain regularized estimate I.sub.R instead of the
estimate I, the eigenvector decomposition of the operator M may be
used:
[0212] Let .phi..sub.1, .phi..sub.2, . . . , .phi..sub.m be
eigenvectors of operator M: H.fwdarw.H corresponding to Eigenvalues
.lamda..sub.1.gtoreq..lamda..sub.2.gtoreq. . . .
.gtoreq..lamda..sub.m.gtoreq.0.
[0213] Let R be some natural number, 1<R<m (R is the
"regularization parameter"). Let H.sup.(R) be the subspace of the
space H spanned by the first R eigenvectors (.phi..sub.1, . . .
.phi..sub.R.
H.sup.(R)=sp{.phi..sub.k}.sub.k=1 . . . R. (7)
[0214] Let P.sup.(R): H.fwdarw.H.sup.(R) be the orthogonal
projection on subspace H.sup.(R).
[0215] The regularized estimate I.sub.R may be obtained as
follows:
[0216] Let .PSI..sub.k .sup.(R)=P.sup.(R).PSI..sub.k, k=1, . . . ,
K.
.PSI..sup.(R)=.SIGMA..sub.k=1 . . .
K.psi..sub.k.PHI..sub.k.sup.(R), (8)
[0217] M.sup.(R): H.sup.(R).fwdarw.H.sup.(R) the operator of the
form:
M.sup.(R)=.SIGMA..sub.k=1 . . .
K.PHI..sub.k.sup.(R).PHI..sub.k.sup.(R) (9)
[0218] (operator M.sup.(R) is the restriction of the operator M of
the form (4) to the subspace H.sup.(R) of the form (7)),
then I.sup.(R)=(M.sup.(R)).sup.-1 .PSI..sup.(R). (10)
[0219] When the regularization parameter R is properly chosen (so
that the eigenvalue .lamda..sub.R is not too small), then the
estimate (10) becomes stable.
[0220] There are several possible approaches to choosing the
parameter R. One approach is to leave R as a "program parameter"
and to obtain the reasonable value "in experiment". Another
approach is to choose some "optimal" value. This is possible if the
covariation operators of the random noise .epsilon..sub.k in (3)
are known, and information about the element I of the space H is
known a priori.
[0221] The subdivision of rectangle V into a large number of equal
rectangles has the disadvantage of making the dimension of the
space H too big (especially in the 3D case). If each side of
rectangle V is subdivided into n equal parts, then the dimension of
the space H will be n.sup.2 and the dimension of the matrices used
in solving the corresponding estimation equations would be
n.sup.2.times.n.sup.2=n.sup.4 (in the 3D case,
n.sup.3.times.n.sup.3=n.sup.6). It is clear that for large n, this
situation may cause serious memory and computation time.
[0222] In accordance with a preferred embodiment of the present
invention, an irregular subdivision of the rectangle V is used.
This irregular approach may significantly decrease the dimension of
the problem and facilitate computer calculations.
[0223] More specifically, a drawback of the regular subdivision of
the investigated region V, discussed hereinabove, is that a lot of
cells that actually have no signal may be taken into account (FIG.
27F). It would be much better to have small cells only in regions
with high signal and have big cells in regions with low signal.
[0224] Reference is now made to FIG. 27G, which illustrates an
advantageous irregular cell subdivision of rectangle V, in
accordance with a preferred embodiment of the present
invention.
[0225] In a first stage, regular subdivisions are made in "large"
cells, and measurements and estimations are made as described
hereinabove. In this manner, the intensity distribution is
estimated in the large cells.
[0226] In a second stage, the large cells, which have an intensity
larger than some threshold, are subdivided into 4 equal subcells
(or 8 subcells in the 3D case). A suitable threshold may be
obtained by taking the average intensity (of all large cells) minus
two (or three) sigmas (standard deviation), for example.
Measurements and estimations are made in these subdivisions as
described hereinabove.
[0227] The act of subdivision and subsequent measurements and
estimations are continued until a desired accuracy is reached at
some smaller level of subdivision, typically defined by the
computational and memory capabilities of the computer being
used.
[0228] The 3D problem may be treated in the same way as the 2D
case, the only difference being that instead of rectangle V, there
is a parallelepiped V (FIG. 27D). Accordingly, the cells in each
subdivision are also parallelepipeds.
[0229] In accordance with the present invention, the wide-aperture
collimation-deconvolution algorithms described hereinabove may be
used for a variety of imaging systems. For example, the algorithms
may be used with single radiation detector probe, an array of
radiation detector probes, large gamma cameras of various designs,
such as multi-head cameras, general purpose cameras, and automatic
white balance (AWB) scanners. The algorithms are suitable for SPECT
and planar imaging, and may be used for all types of isotopes of
with any type of photon energy.
[0230] From the foregoing discussion, the skilled artisan will
appreciate that the algorithms described hereinabove may be used to
predict the location of the radiation source and the uncertainty
region (based on the system measurement errors) in the vicinity of
the radiation source. The algorithms also guide the user to perform
additional measurements to minimize the uncertainty region
according to the requirements of the system operator.
[0231] The algorithms thus comprise a feedback system that employs
analysis to determine the bounds of an uncertainty region about the
radiation source, and which guides medical personnel to conduct
additional scans in these uncertainty regions to improve accuracy,
reduce error, and hence minimize the bounds of the uncertainty
regions.
[0232] Continuous sampling with radiation probe 202 may provide
localization of a tumor and a physiological radiation activity map
of the tumor region. Higher safety and accuracy are gained by a
greater number of scans.
[0233] Other deconvolution methods are known and are often used in
image processing procedures. Examples of such deconvolution methods
are described in U.S. Pat. No. 6,166,853 to Sapia et al., the
disclosure of which is incorporated herein by reference. (However,
it is appreciated that these are just examples and the present
invention is not limited to the deconvolution methods mentioned in
U.S. Pat. No. 6,166,853)
[0234] Reference is now made to FIG. 22 which illustrates a flow
chart of a radiation map reconstruction algorithm, by another
system of wide-aperture collimation-deconvolution algorithms, in
accordance with another preferred embodiment of the present
invention.
[0235] In typical image acquisition, light (or other
electromagnetic wave energy) passes through a finite aperture to an
image plane. The acquired image is a result of a convolution of the
source object's light with the aperture of the imaging system. A
system transfer-function may be generally obtained directly by
taking the Fourier transform of the aperture. As is known in the
art, the blurring effects due to convolution generally exist in
two-dimensions only, i.e., the x-y planes. A point-spread-function
(PSF) is an expression used to describe the convolutional blurring
in two-dimensions. The PSF physically results from imaging a point
source. The Fourier transform of the PSF is the system
transfer-function, which is obtained by convolving the system
transfer-function with a Dirac-delta function. A point source is
the physical equivalent of a Dirac-delta function, and, in the
frequency domain, the Dirac-delta function is a unity operator
across the spectrum. Therefore, the Fourier transform of the PSF
should be the Fourier transform of the aperture. However, the PSF
contains noise and blurring due to other effects such as
aberrations.
[0236] The PSF contribution to the overall blurriness may be
diminished or eliminated by deconvolution.
[0237] In the case of the present invention, the transfer function
of the radiation detector may be determined by taking the Fourier
transform of the aperture of the detector, and taking into account
the noise and blurring due to other effects such as aberrations
(step 500). An example of a transfer function may be a normal
distribution. Using known mathematical techniques, the
deconvolution of the transfer function may be determined (step
502).
[0238] The count readings of each spatial location of the detector
constitute the sum of radiation counts from all the voxels (or
pixels in the case of two-dimensional maps, the term "voxel" being
used herein to include both pixels and voxels) within the
detector's field of view. At least one voxel, or preferably each
such voxel, may be assigned a count value based on the
deconvolution of the unique transfer function of the radiation
detector in use (step 504). An additional mathematical procedure
may treat the various values that each voxel receives due to the
multiple readings from viewpoints of different detectors (step
506). This treatment may constitute for example a simple algebraic
average, minimum value or reciprocal of averaged reciprocals in
order to produce a single value of readings in each voxel. The
deconvolution is then used to reconstruct the voxels of the
radiation map with diminished or no blurriness (step 508).
[0239] The algorithms described herein are applicable not only to
the analysis of readings obtained using a directional radioactivity
detector, rather they also apply for spatially sensitive
(pixellated) radioactivity detectors. In this case, the readings of
each pixel are algorithmically treated as described herein like for
a directional radioactivity detector. The motivation behind using a
spatially sensitive detector is to save on measurement time by
receiving readings from a multitude of directions in parallel.
This, in essence, creates a number of overlapping low resolution
images which can then be processed to form a high resolution image.
In addition, the spatially sensitive detector can be scanned to
improve even further the resolution using the algorithms described
hereinabove.
[0240] Thus, the same algorithms that apply for a directional
detector apply for the spatially sensitive detector, only now
instead of one radiation reading at each position, a large set of
desecrate positions are processed in parallel. Each pixel can be
seen as a separate detector with an angle of acceptance dictated by
the geometry of a segmented collimator employed thereby. Each of
the pixels occupies a different position in space and hence can be
seen as a new position of a single directional probe by the
algorithm described herein. It is also possible, like with the
directional detector, to scan the whole set of pixels by scanning
the spatially sensitive detector and to acquire a new set of data
points from the new position. Once obtaining a low resolution image
from each of the pixels of the spatially sensitive detector, a
super resolution algorithm can be employed to generate an image of
higher resolution. Suitable super resolution algorithms are
described in, for example, J. Acoust. Soc. Am., Vol. 77, No. 2,
February 1985 Pages 567-572; Yokota and Sato, IEEE Trans. Acoust.
Speech Signal Process. (April 1984); Yokota and Sato, Acoustical
Imaging (Plenum, N.Y., 1982, Vol. 12; H. Shekarforoush and R.
Chellappa, "Data-Driven Multi-channel Super-resolution with
Application to Video Sequences", Journal of Optical Society of
America-A, vol. 16, no. 3, pp. 481-492, 1999; H. Shekarforoush, J.
Zerubia and M. Berthod, "Extension of Phase Correlation to
Sub-pixel Registration", IEEE Trans. Image Processing, to appear;
P. Cheeseman, B. Kanefsky, R. Kruft, J. Stutz, and R. Hanson,
"Super-Resolved Surface Reconstruction From Multiple Images," NASA
Technical Report FIA-94-12, December, 1994; A. M. Tekalp, M. K.
Ozkan, and M. I. Sezan, "High-Resolution Image Reconstruction for
Lower-Resolution Image Sequences and Space-Varying Image
Restoration," IEEE International Conference on Acoustics, Speech,
and Signal Processing (San Francisco, Calif.), pp. III-169-172,
Mar. 23-26, 1992, http://www-visl.technion.ac.il/1999/99-03/www/,
which are incorporated herein by reference.
[0241] Referring further to the drawings, FIGS. 28A-28F
schematically illustrate a radioactive emission probe 700,
preferably, handheld, for free-hand scanning, in accordance with
preferred embodiments of the present invention.
[0242] As seen in FIG. 28A, handheld probe 700 is a lightweight
nuclear imaging camera, adapted for free-hand movement and having a
positioning device 714 incorporated thereto.
[0243] Handheld probe 700 includes a radiation detector 706,
mounted onto a housing 702, wherein the various electronic
components (not shown) are contained. Housing 702 is preferably
formed of a rigid, lightweight plastic, a composite, or the like,
and includes a handle 704, for easy maneuvering of probe 700.
[0244] A control unit 710, which may be mounted on housing 702, may
include basic control knobs, such as "stop", "start", and "pause."
Additionally or alternatively, control unit 710 may be a computer
unit, such as a microcomputer or the like, having processing and
memory units, contained within housing unit 702. Additionally,
control unit 710 may include a display screen 718, mounted on
housing unit 702, for displaying information such as gamma counts,
gamma energies, device position, and the like. Display screen 718
may be interactive, for control and display. Control unit 710 may
further include a data unit 720, for receiving a diskette, a
minidisk, or the like. Preferably, data unit 720 is a read and
write unit. An eject button 722 may be included with data unit
720.
[0245] A cable 724 may provide signal communication between
handheld probe 700 and a main-data-collection-and-analysis-system
726, such as a PC computer or a server. Additionally or
alternatively, signal communication with
main-data-collection-and-analysis-system 726 may be wireless, for
example, by radio frequency or infrared waves. Alternatively,
hand-held probe 700 may be a stand-alone unit, operative with
built-in computer unit 710. Additionally, data may be transferred
to and from main-data-collection-and-analysis-system 726 via a
diskette or a minidisk, from data unit 720.
[0246] A preferably rechargeable battery 716 may be attached to
housing unit 702. Additionally or alternatively, a cable 712 may be
used to provide power communication between handheld probe 700 and
the grid (not shown), or between handheld probe 700 and
main-data-collection-and-analysis-system 726.
[0247] Positioning device 714 is preferably a Navigation sensor,
for determination of six coordinates X, Y and Z axes and rotational
angles .rho., .theta. and .phi., as shown in FIG. 28F. Positioning
device 714 may be mounted on housing 702, or located within it.
Preferably, positioning device 714 is a magnetic tracking and
location system, known as miniBIRD, commercially available from
Ascension Technology Corporation, P.O. Box 527, Burlington, Vt.
05402 USA (http://www.ascension-tech.com/graphic.htm).
Alternatively, any other known, preferably 6D tracking and
positioning system may be used. Positioning device 714 maintains
signal communication with either one or both of control unit 710
and main-data-collection-and-analysis-system 726.
[0248] As seen in FIGS. 28A and 28B, nuclear imaging detector 706
may be a disk, forming a single pixel, having a diameter D of about
10 mm, and a length L1 of about 5 mm. It will be appreciated that
other dimensions, which may be larger or smaller, are also
possible. Preferably, radiation detector 706 is connected to a
preamplifier (not shown).
[0249] Additionally, as seen in FIGS. 28A and 28B, handheld probe
700 may include a tubular, wide-bore collimator 708, formed for
example, of lead or of tungsten. The collimator may have a length
L2 of about 10 to 30 mm, and preferably, about 15 mm. It will be
appreciated that other dimensions, which may be larger or smaller,
are also possible.
[0250] Alternatively, as seen in FIGS. 28C and 28D, collimator 708
may be a square grid of 4.times.4 cells 705, or 16.times.16 cells
705. Collimator grid 708 may have sides w of about 10 mm, and
length L2 of about 20-40 mm mm. It will be appreciated that other
dimensions, which may be larger or smaller, are also possible.
[0251] As seen in FIG. 28E, each cell 705 may include a plurality
of radiation detector pixels 703, formed for example, as square
pixels of 5.times.5 mm and length L1 also of about 5 mm. The reason
for the division of nuclear detector 708 into pixels 703 is that
while the wide aperture of a coarse collimator increases the
counting efficiency, radiation detector uniformity is better
maintained when the size of the radiation detectors is small. Thus,
it is desired to combine a coarse collimator grid with a fine pixel
size. Lead spacers are placed between the pixels. From spatial
resolution consideration, each cell 705 preferably operates as a
single pixel, but each pixel 703 within cell 705, must be
calibrated individually, to correct for material nonuniformity, as
will be described hereinbelow, in conjunction with FIG. 29B.
[0252] Preferably, radiation detector 706 is equipped with a
positive contact array, for forming contact with each pixel 703,
and a common negative contact. Preferably, each pixel 703 is
connected to a preamplifier.
[0253] Preferably, radiation detector 706 is a single module array,
for example, of 4.times.4, or of 16.times.16 pixels, of room
temperature CdZnTe, obtained, for example, from IMARAD IMAGING
SYSTEMS LTD., of Rehovot, ISRAEL, 76124, www.imarad.com.
[0254] Room temperature solid-state CdZnTe (CZT) is among the more
promising nuclear detectors currently available. It has a better
count-rate capability than other detectors on the market, and its
pixilated structure provides intrinsic spatial resolution.
Furthermore, because of the direct conversion of the gamma photon
to charge-carriers, energy resolution is enhanced and there is
better rejection of scatter events and improved contrast.
[0255] In accordance with the present invention, the detector may
be optimized, in accordance with the teaching of "Electron lifetime
determination in semiconductor gamma detector arrays,"
http://urila.tripod.com/hecht.htm, "GdTe and CdZnTe Crystal Growth
and Production of Gamma Radiation Detectors,"
http://members.tripod.com/-urila/crystal.htm, and "Driving Energy
Resolution to the Noise Limit in Semiconductor Gamma Detector
Arrays," Poster presented at NSS2000 Conference, Lyon France, 15-20
Oct. 2000, http://urila.tripod.com/NSS.htm, all by Uri Lachish, of
Guma Science, P.O. Box 2104, Rehovot 76120, Israel,
urila@internet-zahav.net, all of whose disclosures are incorporated
herein by reference.
[0256] Accordingly, radiation detector 706 may be a monolithic
CdZnTe crystal, doped with a trivalent donor, such as indium.
Alternatively, aluminum may be used as the trivalent donor. When a
trivalent dopant, such as indium, replaces a bivalent cadmium atom
within the crystal lattice, the extra electron falls into a deep
trap, leaving behind an ionized shallow donor. The addition of more
donors shifts the Fermi level from below the trapping band to
somewhere within it. An optimal donor concentration is achieved
when nearly all the deep traps become occupied and the Fermi level
shifts to just above the deep trapping band.
[0257] Optimal spectral resolution may be achieved by adjusting the
gamma charge collection time (i.e., the shape time) with respect to
the electron transition time from contact to contact. Gamma photons
are absorbed at different depth within the detector where they
generate the electrons. As a result, these electrons travel a
different distance to the counter electrode and therefore produce a
different external signal for each gamma absorption event. By
making the shape time shorter than the electron transition time,
from contact to contact, these external signals become more or less
equal leading to a dramatic improvement in resolution.
[0258] Furthermore, for a multi-pixel detector, the electrons move
from the point of photon absorption towards the positive contact of
a specific pixel. The holes, which are far slower, move towards the
negative contact, and their signal contribution is distributed over
a number of pixels. By adjusting the gamma charge collection time
(i.e., the shape time) with respect to the electron transition time
from contact to contact, the detector circuit collects only the
electrons' contribution to the signal, and the spectral response is
not deteriorated by the charge of the holes.
[0259] For an optimal detector, crystal electrical resistively may
be, for example, about 5.times.10 .sup.8 ohm cm. The bias voltage
may be, for example, -200 volts. The shape time may be, for
example, 0.5.times.10.sup.-6 sec. It will be appreciated that other
values, which may be larger or smaller are also possible.
[0260] In accordance with a preferred embodiment of the present
invention, control unit 710 includes data acquisition and control
components, which further include a very small, single-module
Detector Carrier Board (DCB). The DCB preferably includes a
temperature control system, having a temperature sensor, a desired
temperature setting, and a heat removal system, preferably
operative by peltier cooling. Preferably, the DCB further includes
a High Voltage (HV) power supply, a HV setting, a HV and dark
current output-sensing, and power regulators, preferably having a
range of about .+-.2V, a XAIM_MBAIS current driver, and noise
filtering components.
[0261] Preferably, data acquisition and control components are
based on Application Specific Integrated Circuit (ASIC)
architecture for the control, calibration, and readout. ASIC data
acquisition and control includes calibration of ASIC data, I/V
conversion, amplification and shaping. Data identification includes
threshold discrimination (i.e., noise rejection) and energy windows
discrimination, wherein there may be one or several energy windows
per radionuclide. Preferably, there are up to four energy windows
per radionuclide, but it will be appreciated that another number is
also possible.
[0262] Data acquisition relates to counting events per pixel, per
energy window, for an acquisition period, which may be very short,
for example, about 100 msec, in a continuous mode. It will be
appreciated that larger or smaller counting periods are also
possible.
[0263] The events per pixel 703, per energy window, is synchronized
with signals from positioning device 714, preferably via a DIO
card, either within control unit 710, or main-data
collection-and-analysis-system 726, or both. Additionally,
initialization information, including Digital to Analog Convertor
(DAC) information, may be stored in software of
main-data-collection-and-analysis-system 726 or within control
system 710.
[0264] Preferably, an Integrated Drive Electronics (IDE) system
provides the electronic interface.
[0265] Referring further to the drawings, FIGS. 29A-29B
schematically illustrate the manner of calibrating handheld probe
700, in accordance with preferred embodiments of the present
invention. Non-uniformity of the material leads to different
sensitivities for each pixel. A correction is required to eliminate
the sensitivity effect of the different pixels.
[0266] An energy spectrum for each pixel 703 of radiation detector
706 is obtained in a classical fashion. Radiation detector 706 is
irradiated by a flood source of a known photon energy, for example,
24 KeV, and for each pixel 703, such as 703(1), 703(2), 703(3).
703(n), a summation of counts as a function of voltage is obtained,
during a finite time interval. The location of the energy peak,
corresponding to the peak photon energy, for example, 24 KeV is
bounded, by an upper level (UL) and a lower level (LL) for that
pixel 703(n), forming an energy window for the specific pixel. Only
photons which fall within the bounded energy window for each pixel
are counted, to exclude Compton scattering and pair production. A
multi-energy source may be used, and several energy windows may be
bounded for each pixel.
[0267] After the energy windows, representing specific energy
peaks, are determined for each pixel, they are stored in a memory
unit.
[0268] During acquisition, each collimator cell 705(n) is operative
as a single pixel. Thus, the sensitivity correction factor has to
be at a collimator-cell level, rather than at a pixel-level.
[0269] To calculate the sensitivity correction factor, an average
counts per cell 705, is obtained as:
average counts per cell 705 = sum of counts in all pixels 703 of
detector 706 number of collimator cells 705 in detector 706
##EQU00002## correction factor for cell 705 ( n ) = average counts
per cell 705 counts in a cell 705 ( n ) ##EQU00002.2##
[0270] During acquisition, the sum of counts within the bounded
windows of all pixels in cell 705(n) is multiplied by the
correction factor for that cell 705(n) to obtain a sensitivity
corrected count rate.
[0271] Referring further to the drawings, FIG. 30 schematically
illustrates the manner of synchronizing event readings by radiation
detector 706 and position readings by positioning device 714, in
accordance with preferred embodiments of the present invention.
Preferably, event readings, for each cell 705 is tabulated with the
position reading, for each time interval, such as t.sub.1, t.sub.2,
t.sub.3, after correction for sensitivity.
[0272] Referring further to the drawings, FIGS. 31A-31C describe a
spatial resolution bar-phantom test of probe 700, in accordance
with a preferred embodiment of the present invention. As in FIGS.
17 and 18, hereinabove, the aim of the bar phantom test is to
evaluate image quality by measuring image resolution of a bar
phantom. Spatial resolution is expressed in terms of the finest bar
pattern, visible on nuclear images.
[0273] As seen in FIG. 31A, a bar phantom 750 was constructed of
lead strips 752, encased in a plastic holder (not shown). The lead
strips were of equal width and spacing. Thus, a 2-mm bar pattern
consisted of 2-mm wide strips 752, separated by 2-mm spaces 754,
along an x axis. The thickness of lead strips 752 was, sufficient
to ensure that bars 752 are virtually opaque to the .gamma. rays
being imaged.
[0274] A uniform radiation field, formed of a sheet source (not
shown), was placed directly behind the bar phantom 750, and bar
phantom 750 was scanned with probe 700.
[0275] Probe 700 of the present embodiment comprised an eV Pen-type
CdZnTe solid state detector with collimator diameter D, of 5 mm,
and collimator length, L2 of 30 mm. The CdZnTe detector was
attached to a charge sensitive preamp probe and connected to a
Digital and analog V-target alpha system PC cards. The probe energy
window calibration was performed at the beginning of the test by
using EG&G Ortec Maestro 2K Tramp ISA Spectra Processing
connected to the analog and the digital V-target alpha system
cards. The probe position information was provided by
MicroScribe-6DLX arm manufactured by Immersion Corporation.
Acquisition time was 6 minutes.
[0276] FIG. 31B illustrates a bar phantom image 760, produced by
probe 700, after 6 minutes of acquisition time, in accordance with
the present invention. Image 760 shows bar images 756, and space
images 758, of the 2-mm bar pattern, clearly demonstrating a high
degree of spatial resolution.
[0277] FIG. 31C provides a similar observation, based counts as a
function of length axis, X. The well-defined, narrow sinusoidal
peaks of counts as a function of length axis demonstrate a high
degree of spatial resolution for the 2-mm bar pattern.
[0278] Referring further to the drawings, FIGS. 32A-32D describe a
spatial resolution bar-phantom test of a prior-art probe, a High
Resolution Gamma Camera of a competitor. Although for the prior art
probe, acquisition time was 3 hours, a bar pattern is barely
observed for a 3-mm bar pattern of FIG. 32C, and clearly observed
only for a 3.5 mm bar pattern of FIG. 32D.
[0279] Table 1 provides a comparison between the prior art
technology and that of the present invention, illustrating a clear
advantage to the present invention.
TABLE-US-00001 TABLE 1 Detector - Maximum Imaging collimator
Acquisition Spatial Technology arrangement Time Resolution a
multi-pixel 180 min. Worse then High-resolution detector, 3 mm
Gamma-Camera tube collimators of a 1.9 mm in dia. competitor's 40
mm in length Probe 700 single-pixel 6 min. 2 mm of FIG. 28A
detector 10 mm in dia. tube collimator 25 mm in length Summary
V-target V-target's probe resolution is 30 time is 2 time faster
better
[0280] Good energy resolution of the detector is important for
precise identification and separation of .gamma. rays, for
radionuclide identification and for scatter rejection. A major
advantage of CdZnTe detector is better energy resolution, for
example than that of the Nai(Tl) detector.
[0281] Referring further to the drawings, FIGS. 33A-33B illustrate
the energy resolution of a single pixel 703 of probe 700, in
accordance with the present invention.
[0282] FIG. 33A compares a spectrum of a single pixel of a CdZnTe
detector with that obtained by a single pixel of an NaI(Tl)
detector, at .about.120 KeV, showing the difference in energy
resolution between the two detectors. Clearly, the a CdZnTe
detector provides a sharper, more well-defined energy peak, which
is about 1/3 higher than the Nai(Tl) peak, and having a FWHM that
is 2/3 that of the Nai(Tl) peak.
[0283] FIG. 33B provides a spectrum of a a CdZnTe detector, showing
good energy resolution at the low energy level of about 24 KeV.
[0284] Referring further to the drawings, FIGS. 34A-34C
schematically illustrate endoscopic radioactive emission probes
800, adapted to be inserted into the body, on a shaft or a catheter
713, via a trocar valve 802 in a tissue 810.
[0285] As seen in FIG. 34A, endoscopic radioactive emission probe
800 may include, for example, radiation detector 706, for example,
of a single pixel, wide-bore collimator 708, position tracking
system 714, and an extracorporeal control unit 804. Preferably, the
diameter of radiation detector 706, D.sub.detector, is less than 12
mm. In accordance with another embodiment of the present invention,
radioactive emission probe 800 may be used in body lumens, for
example, for scanning the stomach or the uterus.
[0286] As seen in FIG. 34B, endoscopic radioactive emission probe
800 may include several radiation detectors 706, for example, of a
single pixel each, and having a dedicated collimator 708, each.
Unlike the collimator grid structure of FIGS. 28C and 28D, the
collimators of FIG. 34B are not parallel to each other, and each
may point in a different direction. Alternatively, each radiation
detector 706 may include a plurality of pixels.
[0287] As seen in FIG. 34C, endoscopic radioactive emission probe
800 may include several radiation detectors 706, which may include
wide-angle collimators.
[0288] Referring further to the drawings, FIG. 35 schematically
illustrates a manner of calculating a distance, d, between
radioactive emission probe, such as probe 700 (FIGS. 28A-28F) and a
radiation source 766, based on the attenuation of photons of
different energies, which are emitted from the same source. Probe
700 is on an extracorporeal side 762 and radiation source 766 is on
an intracorporeal side 764 of a tissue 768.
[0289] The transmitted intensity of a monochromatic gamma-ray beam
of photon energy E and intensity I.sub.0 crossing an absorbing
layer, such as a tissue layer, of a depth d, is a function of the
photons energy E, the layer thickness d, and the total linear
attenuation coefficient as a function of energy, .mu..sub.total, of
the material constituting the layer. The transmitted intensity is
given by:
I(E)=I.sub.0(E)exp(-.mu..sub.total(E)d) (Eq-1)
[0290] The linear attenuation coefficient is expressed in cm.sup.-1
and is obtained by multiplying the cross section (in cm.sup.2/gr)
with the absorbing medium density. The total linear attenuation
coefficient accounts for absorption due to all the relevant
interaction mechanisms: photoelectric effect, Compton scattering,
and pair production.
[0291] The layer thickness d, which is the depth of the radiation
source, can be measured using a radionuclide that emits two or more
photon energies, since attenuation is a function of the photon
energy.
[0292] For photons of different energies, originating from a same
source, and having the same half-life, we get:
I ( E 1 ) I ( E 2 ) = I 0 ( E 1 ) I 0 ( E 2 ) - .mu. ( E 1 ) d -
.mu. ( E 2 ) d ( Eq - 2 ) A = I ( E 1 ) I ( E 2 ) & B = I 0 ( E
1 ) I 0 ( E 2 ) ( Eq - 3 ) R = A B = [ I ( E 1 ) / I ( E 2 ) ] [ I
0 ( E 1 ) / I 0 ( E 2 ) ] ( Eq - 4 ) A = B - .mu. ( E 1 ) d - .mu.
( E 2 ) d R = - .mu. ( E 1 ) d - .mu. ( E 2 ) d ( Eq - 5 ) R = {
.mu. ( E 2 ) - .mu. ( E 1 ) } d ln ( R ) = [ .mu. ( E 2 ) - .mu. (
E 1 ) ] d ( Eq - 6 ) ##EQU00003##
[0293] The source depth, d, can be calculated as:
d = ln ( R ) .mu. ( E 2 ) - .mu. ( E 1 ) = ln { [ I ( E 1 ) / I ( E
2 ) ] / [ I 0 ( E 1 ) / I 0 ( E 2 ) ] } .mu. ( E 2 ) - .mu. ( E 1 )
( Eq - 7 ) ##EQU00004##
[0294] The linear attenuation coefficients for specific energies
are calculated from the photon mass attenuation and energy
absorption coefficients table.
[0295] The ratio I.sub.0(E.sub.1)/I.sub.0(E.sub.2) for a known
isotope or isotopes is calculated from an isotope table and the
ratio I(E.sub.1)/I(E.sub.2) is measured at the inspected object
edges.
[0296] Additionally, X-ray information, for example, mammography
may be used, for more material linear attenuation coefficient,
since X-ray imaging provides information about the material
density.
[0297] For example, we may consider a lesion inside a soft tissue,
such as a breast, at a distance d that holds an isotope of
.sup.123I, which emits two main photons, as follows:
[0298] E.sub.1=27 keV with 86.5%; and
[0299] E.sub.2=159 keV with 83.4%.
[0300] The soft tissue density is 0.8 g/cm.sup.3.
[0301] The linear attenuation coefficients for E.sub.1 and E.sub.2
are calculated from J. H Hubbell, "Photon Mass Attenuation and
Energy-absorption Coefficients from 1 keV to 20 MeV", Int. J. Appl.
Radiat. Isat. Vol. 33. pp. 1269 to 1290, 1982, as follows:
[0302] .mu.(E.sub.1)=0.474 cm/g.fwdarw..mu.(E.sub.1)=0.474 cm/g0.8
g/cm.sup.3=0.38 1/cm
[0303] .mu.(E.sub.2)=0.1489 cm/g.fwdarw..mu.(E.sub.2)=0.1489
cm/g0.8 g/cm.sup.3=0.119 1/cm
[0304] The measured ratio between I(E.sub.1) to I(E.sub.2) on the
attenuating soft tissue edge is 0.6153.
[0305] Using Eq-7, one gets that the lesion depth is: d=2 cm.
[0306] Alternatively, if the measured ration between I(E.sub.1) to
I(E.sub.2) is 0.473 the extracted lesion depth is: d=3 cm.
[0307] It will be appreciated that when two or more isotopes used
for obtaining the different photons, corrections for their
respective half-lives may be required.
[0308] It will be appreciated that by calculating the depth of a
radiation source at each position, one may obtain a radiation
source depth map, in effect, a three-dimensional image of the
radiation source. This information may be superimposed on the
radiation source image, as produced by other methods of the present
invention, as an independent check of the other methods. For
example, a radiation source image produced by the wide-aperture
collimation-deconvolution algorithms may be superimposed on a
radiation source image produced by depth calculations of based on
the attenuation of photons of different energies.
EXPERIMENTAL RESULTS
[0309] In a series of clinical experiments, some of the basic
concepts of the invention have been tested on patients who were
pre-injected with a suitable radiopharmaceutical for their
particular pathology. Two-dimensional color-coded maps have been
constructed based on a scan of a pre-determined lesion area by a
hand-held radiation detector with a magnetic position-tracking
system. The resulting maps, which represented the radiation count
level, were compared to images of conventional gamma camera. The
list of radiopharmaceuticals tested includes .sup.18FDG,
.sup.99MTc-MDP, .sup.99MTc sodium pertechnetate, .sup.99MTc
erythrocytes. Similar radiolabeled patterns were observed in the
images produced by the system of the invention and in the images
produced by a conventional gamma camera in the following
pathologies:
[0310] FIGS. 23A and 23B illustrate radiolabeled patterns observed
in images produced by the system of the invention and by a
conventional gamma camera, respectively, of an autonomous adenoma
of a thyroid of a 58 year-old male.
[0311] FIGS. 24A and 24B illustrate radiolabeled patterns observed
in images produced by the system of the invention and by a
conventional gamma camera, respectively, of suspected Paget's
disease of a humerus in an 89 year-old female.
[0312] FIGS. 25A and 25B illustrate radiolabeled patterns observed
in images produced by the system of the invention and by a
conventional gamma camera, respectively, of chronic osteomyelitis
in a 19 year-old female.
[0313] FIGS. 26A and 26B illustrate radiolabeled patterns observed
in images produced by the system of the invention and by a
conventional gamma camera, respectively, of skeletal metastasis
from medulloblastoma in an 18 year-old male.
[0314] FIG. 36 illustrates a control volume 701, and a
three-dimensional image 703 of the radioactivity emitting source,
produced by free-hand scanning of a cancerous prostate gland, ex
vivo, with hand-held probe 700 of FIG. 28A. The diameter of the
detector was 10 mm, and the length of the collimator was 25 mm. The
wide-aperture collimation-deconvoluting algorithm was used, for
image reconstruction. A scale 705 illustrates brightness as a
function of count rate density. The brighter the area, the higher
the count density. The Figure shows the capability of hand-held
probe 700 to detect cancerous tissue. The By comparison, the
competitor's High-resolution, Gamma-Camera failed to detect any
cancerous growth.
[0315] The following provides a list of known procedures which can
take advantage of the system and method of the present
invention:
[0316] In cancer diagnosis the system and method of the present
invention can find uses for screening for cancer and (or) directing
invasive diagnosis (biopsies) either from outside the body or by
way of endoscopic approach. Examples include, but are not limited
to, lung cancer biopsy, breast cancer biopsy, prostate cancer
biopsy, cervical cancer biopsy, liver cancer biopsy, lymph node
cancer biopsy, thyroid cancer biopsy, brain cancer biopsy, bone
cancer biopsy, colon cancer biopsy, gastro intestine cancer
endoscopy and biopsy, endoscopic screening for vaginal cancer,
endoscopic screening for prostate cancer (by way of the rectum),
endoscopic screening for ovarian cancer. (by way of the vagina),
endoscopic screening for cervical cancer (by way of the vagina),
endoscopic screening for bladder cancer (by way of the urinary
track), endoscopic screening for bile cancer (by way of the
gastrointestinal track), screening for lung cancer, screening for
breast cancer, screening for melanoma, screening for brain cancer,
screening for lymph cancer, screening for kidney cancer, screening
for gastro intestinal cancer (from the outside).
[0317] In the special case of MRI, the radiation detector can be
combined and packaged together with a small RF coil for the
transmission and reception or reception only of the MRI signals in
a rectal probe configuration for prostate diagnosis and treatment
or any other close confinement position such as the vagina,
airways, the uper portion of the gastrointestinal track, etc)
[0318] Procedures known as directing localized treatment of cancer
can also benefit from the system and method of the present
invention. Examples include, but are not limited to, intra tumoral
chemotherapy, intra tumoral brachytherapy, intra tumoral cryogenic
ablation, intra tumoral radio frequency ablation, intra tumoral
ultrasound ablation, and intra tumoral laser ablation, in cases of,
for example, lung cancer, breast cancer, prostate cancer, cervical
cancer, liver cancer, lymph cancer, thyroid cancer, brain cancer,
bone cancer, colon cancer (by way of endoscopy through the rectum),
gastric cancer (by way of endoscopy through the thorax), thoracic
cancer, small intestine cancer (by way of endoscopy through the
rectum or, by way of endoscopy through the thorax), bladder cancer,
kidney cancer, vaginal cancer and ovarian cancer.
[0319] In interventional cardiology the following procedures can
take advantage of the present invention wherein the method and
system can be used to assess tissue perfusion, tissue viability and
blood flow intra operatively during PTCA procedure (balloon alone
or in conjunction with the placement of a stent), in cases of
cardiogenic shock to asses damage to the heart, following
myocardial infarct to asses damage to the heart, in assessing heart
failure condition tissue in terms of tissue viability and tissue
perfusion, in intra vascular tissue viability and perfusion
assessment prior to CABG operation.
[0320] The radioactivity detector can be mounted on a catheter that
is entered through the blood vessels to the heart to evaluate
ischemia from within the heart in order to guide ablation probes or
another type of treatment to the appropriate location within the
heart. Another application which may benefit from the present
invention is the localization of blood clots. For example, a
radioactivity detector as described herein can be used to asses and
differentiate between new clots and old clots. Thus, for example,
the radioactivity detector can be placed on a very small caliber
wire such as a guide wire that is used during PTCA in order to
image blood-vessel clots. Blood-vessel clots can be searched for in
the aortic arc as clots therein are responsible for about 75% of
stroke cases.
[0321] Using the method and system of the present invention to
assess tissue perfusion, tissue viability and blood flow intra
operatively can also be employed in the following: during CABG
operation to asses tissue viability, to mark infarct areas, during
CABG operations to asses the success of the re vascularization.
[0322] The present invention has many other applications in the
direction of therapeutics, such as, but not limited to, implanting
brachytherapy seeds, ultrasound microwave radio-frequency
cryotherapy and localized radiation ablations.
[0323] It will be appreciated that many other procedures may also
take advantage of the present invention.
[0324] It is appreciated that certain features of the invention,
which are, for clarity, described in the context of separate
embodiments, may also be provided in combination in a single
embodiment. Conversely, various features of the invention, which
are, for brevity, described in the context of a single embodiment,
may also be provided separately or in any suitable
subcombination.
[0325] Although the invention has been described in conjunction
with specific embodiments thereof, it is evident that many
alternatives, modifications and variations will be apparent to
those skilled in the art. Accordingly, it is intended to embrace
all such alternatives, modifications and variations that fall
within the spirit and broad scope of the appended claims. All
publications in printed or electronic form, patents and patent
applications mentioned in this specification are herein
incorporated in their entirety by reference into the specification,
to the same extent as if each individual publication, patent or
patent application was specifically and individually indicated to
be incorporated herein by reference. In addition, citation or
identification of any reference in this application shall not be
construed as an admission that such reference is available as prior
art to the present invention.
* * * * *
References