U.S. patent application number 10/945541 was filed with the patent office on 2005-05-26 for tomography scanner with axially discontinuous detector array.
Invention is credited to Green, Michael V., Seidel, Jurgen, Vaquero, Juan Jose.
Application Number | 20050109943 10/945541 |
Document ID | / |
Family ID | 34375475 |
Filed Date | 2005-05-26 |
United States Patent
Application |
20050109943 |
Kind Code |
A1 |
Vaquero, Juan Jose ; et
al. |
May 26, 2005 |
Tomography scanner with axially discontinuous detector array
Abstract
A tomography scanner has intentionally designed, well defined
gaps between detector rings with image reconstruction obtained with
the use of conventional tomography data processing. The scanner is
particularly advantageous as a small animal PET scanner.
Inventors: |
Vaquero, Juan Jose; (Madrid,
ES) ; Seidel, Jurgen; (Bethesda, MD) ; Green,
Michael V.; (Kensington, MD) |
Correspondence
Address: |
EPSTEIN & GERKEN
1901 RESEARCH BOULEVARD
SUITE 340
ROCKVILLE
MD
20850
US
|
Family ID: |
34375475 |
Appl. No.: |
10/945541 |
Filed: |
September 20, 2004 |
Related U.S. Patent Documents
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Application
Number |
Filing Date |
Patent Number |
|
|
60504321 |
Sep 18, 2003 |
|
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Current U.S.
Class: |
250/363.04 |
Current CPC
Class: |
G01T 1/1644 20130101;
G01T 1/2008 20130101; G01T 1/2985 20130101 |
Class at
Publication: |
250/363.04 |
International
Class: |
G01T 001/166 |
Claims
What is claimed is:
1. A tomography scanner comprising a plurality of axially aligned
detector rings forming a cylindrical detector array around a body
to be imaged with well defined gaps between said detector rings,
said detector rings producing signals in response to radiation
emanating from the body in three dimensions; and means responsive
to said signals to produce tomographic image reconstruction
throughout the axial field-of-view including said well defined
gaps.
2. A tomography scanner as recited in claim 1 wherein said image
reconstruction means includes means performing an iterative
statistical or a classical image reconstruction process.
3. A tomography scanner as recited in claim 1 wherein said image
reconstruction means uses 3D OSEM and 3D MAP algorithms.
4. A tomography scanner as recited in claim 1 wherein said image
reconstruction means includes an iterative statistical
reconstruction method that models some or all of the physics and
geometry of the detector array.
5. A tomography scanner as recited in claim 1 wherein said image
reconstruction means utilizes 3D Fourier re-binning followed by 2D
filtered back projection.
6. A tomography scanner as recited in claim 1 wherein said image
reconstruction means utilizes a 3D Fourier re-binning followed by a
2D iterative statistical reconstruction method.
7. A tomography scanner as recited in claim 1 wherein said image
reconstruction means utilizes a 3D reprojection algorithm.
8. A tomography scanner as recited in claim 1 wherein said detector
rings include a plurality of detector modules angled in an axial
direction.
9. A tomography scanner as recited in claim 1 and further
comprising means to provide relative reciprocating or
unidirectional axial movement between the body to be imaged and
said detector rings to reduce the effects of said gaps.
10. A tomography scanner as recited in claim 1 wherein said
detector rings include a plurality of detector modules carrying a
plurality of detector elements, each detector element having an
axial dimension and said well defined gaps have an axial dimension
greater than or equal to said axial dimensions of said detector
elements.
11. A tomography scanner as recited in claim 10 wherein said well
defined gaps each have an axial dimension at least three times
greater than said axial dimension of said detector elements.
12. A tomography scanner as recited in claim1 wherein said detector
rings each carry a plurality of detector elements having points of
axial resolution with the spacing between said points defining
pitch and said gaps have an axial dimension greater than said
pitch.
13. A positron emission tomography scanner comprising a pair of
detector rings forming a cylindrical detector array around a body
to be imaged with a well defined gap between said detector rings; a
plurality of detector modules carried by said detector rings
producing signals corresponding to the three dimensional points of
interaction in each of a pair of said detector modules in response
to detection of time coincident photons from positron annihilations
in a body containing a positron emitting compound; and means
responsive to said signals to produce tomographic image
reconstruction of the spatial distribution of the amount of
positron emitting compound in the body along the axial
field-of-view of said cylindrical detector array.
14. A positron emission tomography scanner as recited in claim 13
wherein said signal responsive means includes an iterative
process.
15. A positron emission tomography scanner as recited in claim 13
wherein said signal responsive means provides classical
reconstruction.
16. A positron emission tomography scanner as recited in claim 13
wherein said detector modules are angled in an axial direction to
reduce the target area resulting from said gap.
17. A positron emission tomography scanner as recited in claim 13
and further comprising means to provide relative reciprocating
axial movement between the animal and said detector array.
18. A tomography scanner as recited in claim 13 wherein said
detector modules each contain at least one detector element having
an axial dimension and said well defined gap has an axial dimension
greater than said axial dimensions of said detector elements.
19. A tomography scanner as recited in claim 18 wherein said well
defined gap has an axial dimension at least three times greater
than said axial dimensions of said detector elements.
Description
CROSS REFERENCE TO RELATED PATENT APPLICATION
[0001] This application claims priority from prior provisional
patent application Ser. No. 60/504,321, filed Sep. 18, 2003, the
entire disclosure of which is incorporated herein by reference.
BACKGROUND OF THE INVENTION
[0002] 1. Field of the Invention
[0003] The present invention pertains to tomography scanners and,
more particularly, to positron emission tomography (PET) scanners
designed for imaging small animals or humans.
[0004] 2. Brief Discussion of the Related Art
[0005] Small animal PET scanners are commonly used in research
facilities and, desirably, have high spatial resolution and
uniformity and high sensitivity as described in "Molecular Imaging
of Small Animals with Dedicated PET Tomographs," Chatziioannou,
Arion F., European Journal of Nuclear Medicine, Vol. 29, No. 1,
January 2002. Commercial examples of such small animal PET scanners
are the Concorde R4 and P4 "microPET" small animal PET scanners
described in "Performance Evaluation of the MicroPET R4 PET Scanner
for Rodents," Knoess, Christof; Seigel, Stefan; Smith, Anne;
Newport, Danny; Richerzhagen, Norbert; Winkeler, Alexandra; Jacobs,
Andreas; Goble, Rhonda N.; Graf, Rudolph; Wienhard, Klaus; and
Heiss, Wolf-Dieter, European Journal of Nuclear Medicine and
Molecular Imaging, Vol. 30, No. 5, May 2003 and "Performance
Evaluation of the MicroPET P4: A PET System Dedicated to Animal
Imaging," Tai, Y. C.; Chatziioannou, A.; Seigel, S.; Young, J.;
Newport, D., Goble, R. N.; Nutt, R. E.; and Cherry, S. R., Physics
in Medicine and Biology, 46 (2001) 1845-1862, and the Philips
"Mosaic" small animal PET scanner described in "Design Evaluation
of A-PET: A High Sensitivity Animal PET Camera," Surti, S.; Karp,
J. S.; Perkins, A. E.; Freifelder, R.; and Muehllehner, G., IEEE
Transactions on Nuclear Science, Vol. 50, No. 5, October 2003.
These scanners utilize dense arrays of small, individual detector
elements, such as scintillation crystals ultimately viewed by
photomultiplier tubes that encode the location of a scintillation
event. The terms "detector elements," "crystals" and "scintillating
materials" are used interchangeably herein; however, it should be
understood that the term "detector elements" includes any elements
capable of detecting any type of radiation. The arrays are
cylindrically arranged around a small diameter circle or polygon to
form a mass of scintillating material that is nearly continuous in
both the axial and circumferential directions. "Continuous" in this
sense means that the individual scintillation crystals are as close
together as possible such that any space between crystals is
minimal and small compared to the crystal width and that the
crystal positioning is replicated along the entire axial length of
the scanning volume without appreciable or well defined gaps
between rings of scintillation crystals. "Cylindrical detector
array" as used herein means any geometric arrangement in which
scintillation crystals or other detector elements circumferentially
surround an imaging volume, e.g. in a circle, a polygon, an oval or
the like, and have some axial extent. Prior art instrumentation for
3D PET imaging has focused on creating continuous axial and
circumferential arrays of scintillator or other materials able to
detect positron annihilation radiation emanating from a stationary
imaging target or body, e.g. a small laboratory animal or a human.
From these detected events, transverse section images of the
radioactivity distribution from the body can be reconstructed that
span the axial field of view of the device. The perceived need for
continuity in detector arrays has been sufficiently compelling that
prior art scanners have been specifically designed to avoid axially
discontinuous arrays and have attempted to exploit novel array
assembly methods, primarily optical/mechanical, that allow
fabrication of continuous arrays of scintillation crystals, e.g.
use of light guide coupling between crystal arrays and
photodetectors, that allow close packing of crystals in both the
axial and circumferential directions.
[0006] Dedicated PET scanners now on the market for either human or
animal imaging targets use continuous cylindrical arrays of small
scintillation crystals to define the imaging volume of the scanner.
Discontinuous detector arrays e.g. paired, opposed flat panel
detectors in time coincidence, are either mechanically rotated
around the imaging target or the imaging target is rotated between
fixed detectors to achieve the same result. See "A Rotating PET
Scanner Using BGO Block Detectors: Design, Performance and
Applications," Townsend, David W.; Wensveen, Martin; Byars, Larry
G.; Geissbuhler, Antoine; Tochon-Danguy, Henri J.; Christin, Anne;
Defrise, Michael; Bailey, Dale L.; Grootoonk, Sylke; Donath,
Alfred; and Nutt, Ronald, Journal of Nuclear Medicine, 1993;
34:1367-1376, "Design and Characterization of IndyPET-II: A High
Resolution, High Sensitivity Dedicated Research Scanner," Rouze,
Ned C. and Hutchins, Gary D., IEEE Transactions on Nuclear Science,
Vol. 50, No. 5, October 2003, and "ECAT ART--A Continuously
Rotating PET Camera: Performance Characteristics, Initial Clinical
Studies, and Installation Considerations in a Nuclear Medicine
Department," Bailey, Dale L.; Young, Helen; Bloomfield, Peter M.;
Meikle, Steven R.; Glass, Daphne; Meyers, Melvyn J.; Spinks,
Terence J.; Watson, Charles C.; Luk, Paul; Peters, A. Michael; and
Jones, Terry, European Journal of Nuclear Medicine, Vol. 24, No. 1,
January 1997. The design of such scanners is commonly driven by the
perceived need to create continuous, or virtually continuous,
crystal arrays in the sense defined previously. For example, some
of such scanners use individual light guides to connect crystals in
an array to a phototube to eliminate the effect of "dead space" at
the edges of phototubes. In other scanners, a continuous annulus of
glass serves as a light guide to connect the cylindrical array of
closely spaced small crystals to a bank of phototubes. In other
scanners, such as the scanner disclosed in U.S. Pat. No. 6,288,399
to Andreaco et al, a large, axially continuous polygonal crystal
array is created by centering, and packaging together, many small
crystal arrays on clusters of four phototubes, a geometry that
allows large arrays to be made by replication of this pattern as
described in "The ECAT HRRT: Performance and First Clinical
Application of the New High Resolution Research Tomograph,"
Weinhard, K.; Schmand, M.; Casey, M. E.; Baker, K.; Bao, J.;
Eriksson, L.; Jones, W. F.; Knoess, C.; Lenox, M.; Lercher, M.;
Luk, P.; Michel, C.; Reed, J. H.; Richerzhagen, N.; Treffert, J.;
Vollmar, S.; Young, J. W.; Heiss, W. D.; and Nutt, R., IEEE
Transactions on Nuclear Science, Vol. 49, No. 1, February 2002. In
each of these cases, technical innovations of one kind or the other
are applied to allow small scintillation crystals to be packed
closely together and to create detector arrays that are essentially
continuous in both the circumferential and axial directions.
[0007] There are two primary reasons why the use of continuous
arrays is deemed important. First, the idea that continuous arrays
will intercept the largest fraction of annihilation radiation
emanating from the target subject and, hence, will yield the
maximum sensitivity for a particular ring diameter and axial
length. Second, "classical" information theory has been thought to
require continuous, regular and dense sampling of an imaging volume
if imaging performance is to be as good as the system geometry
permits. It has been generally believed that images reconstructed
without dense and uniform sampling, i.e. without a continuous axial
and circumferential distribution of scintillating material, would
be of inferior quality, would contain artifacts or both.
Degradation has been expected to increase if there were actual gaps
in the detector array in either the circumferential or axial
directions. In particular, while the effect on image quality of
small circumferential gaps in detector arrays has been studied in
some detail, see "Statistical Image Reconstruction in PET with
Compensation for Missing Data," Kinahan, P. E.; Fessler, J. A.; and
Karp, J. S., IEEE Transactions on Nuclear Science, Vol. 44, No. 4,
August 1997, "Correction Methods for Missing Data in Sinograms of
the HRRT PET Scanner," de Jong, Hugo W. A. M.; Boellaard, Ronald;
Knoess, Christof; Lenox, Mark; Michel, Christiaan; Casey, Michael;
and Lammertsma, Adriaan A., IEEE Transactions on Nuclear Science,
Vol. 50, No. 5, October 2003, and "A Study of Image Errors Due to
Detector Gaps Using OS-EM Reconstructions," Yu, D.-C. and Chang,
W., IEEE 1998, the literature contains little information about
changes in image quality if a detector array possesses axial gaps,
see "Design Optimization of the PMT-ClearPET Prototypes Based on
Simulation Studies with GEANT3," Heinrichs, U.; Pietrzyk, U.; and
Ziemons, K., IEEE Transactions on Nuclear Science, Vol. 50, No. 5,
October 2003.
[0008] While a continuous cylindrical array of scintillation
crystals surrounding an imaging target is an effective way to
intercept annihilation radiation from an imaging target, prior art
methods possess significant practical disadvantages. For example,
the need to connect individual or small groups of scintillation
crystals to a phototube with a light guide adds complexity to the
manufacturing process. More importantly, there is a demonstrable
loss of scintillation light when the light passes into and through
a light guide, thus, potentially reducing imaging performance. A
similar loss occurs when a bulk light guide is used for the same
purpose. That is, it has been believed that the "dead" regions at
the edges of most photonic devices cannot be tolerated.
[0009] A number of three-dimensional (3D) image reconstruction
methods have been proposed in recent years including the Fourier
re-binning method (FORE) combined with some form of 2D image
reconstruction, e.g. filtered backprojection (FBP) and the 3D
re-projection method (3DRP), as described in "Exact and Approximate
Rebinning Algorithms for 3-D PET Data," Defrise, Michael; Kinahan,
P. E.; Townsend, D. W.; Michel, C.; Sibomana, M.; and Newport, D.
F., IEEE Transactions on Medical Imaging, Vol. 16, No. 2, April
1997 and "Performance of the Fourier Rebinning Algorithm for PET
with Large Acceptance Angles," Matej, Samuel; Karp, Joel S.;
Lewitt, Robert M.; and Becher, Amir, Phys. Med. Biol. 43 (1998)
787-795, which are incorporated herein by reference. Iterative,
statistical methods, such as 3D ordered subset expectation
maximization (3D OSEM), as described in "Accelerated Image
Reconstruction Using Ordered Subsets of Projection Data," Hudson,
H. Malcolm and Larkin, Richard S., IEEE Transactions on Medical
Imaging, Vol. 13, No. 4, December 1994, which is incorporated
herein by reference and 3D maximum a posteriori image
reconstruction (3D MAP) as described in "High Resolution 3D
Bayesian Image Reconstruction Using the MicroPET Small Animal
Scanner," Qi, Jinyi; Leahy, Richard M.; Cherry, Simon R.;
Chatziioannou, Arion; and Farquhart, Thomas H., Phys. Med. Biol.,
43, (1998) 1001-1013, both with system modeling, have also been
introduced. A number of other algorithms that exploit the
expectation maximization-maximum likelihood (EM-ML) approach with
system modeling have also been studied as described in "Fast
EM-Like Methods for Maximum "A Posteriori" Estimates in Emission
Tomography," De Pierro, Alvaro R. and Yamagishi, Michel Eduardo
Beleza, IEEE Transactions on Medical Imaging, Vol. 30, No. 4, April
2001. Each of these methods allows the 3D information potentially
available in cylindrical PET scanners without collimators to be
reconstructed into 2D slices that fully exploit the increased
sensitivity associated with 3D data collections compared to purely
2D collections. To date, these methods have been used only for
image reconstruction from scanners having continuous cylindrical
arrays of scintillation crystals.
SUMMARY OF THE INVENTION
[0010] The present invention avoids the need for specially designed
array assemblies having axially continuous detector arrays by
adapting existing image reconstruction methods to the presence of
axial gaps in a detector array, by mechanical movement of the
imaging target relative to an axially discontinuous detector array
such that lines-of-response from parts of the object that might
otherwise always lie in a gap are translated into locations where a
detector array is continuous, and by arranging the detector modules
in the detector array such that they are tilted with respect to one
another in the axial direction.
[0011] In one aspect, the present invention permits the
construction of a tomography scanner with spaced detector rings
allowing direct coupling of scintillators with photon detectors,
particularly position-sensitive photomultiplier tubes, and uses
conventional image reconstruction methods with tomography scanners
having axially discontinuous arrays of detector
rings/scintillators. The effect of gaps in the detector arrays can
be further minimized, if desired, by appropriate movement of the
imaging target during imaging or by geometric arrangement of the
detector elements in the cylindrical detector array. In one mode
the tomography scanner of the present invention compensates for
missing data introduced by discontinuities in detector arrays of
tomography scanners by using three dimensional (3D) re-binning
and/or reconstruction methods as discussed above.
[0012] A tomography scanner according to the present invention
includes a plurality of axially aligned detector rings forming a
cylindrical detector array that surrounds the imaging target. This
cylindrical array contains one or more well defined gaps between
the detector rings, which produce signals corresponding to points
of interaction in three dimensions of radiation from the body
within the detector array and means responsive to the signals to
produce image reconstructions of some aspect of the imaging target
throughout the axial field-of-view including the well defined
gap.
[0013] In accordance with the present invention, a tomography
scanner includes a plurality of axially aligned detector rings
forming a cylindrical detector array around a body to be imaged
with well defined gaps between the detector rings, the detector
rings producing signals in response to radiation emanating from the
body in three dimensions and means responsive to the signals to
produce tomographic image reconstruction throughout the axial
field-of-view including the well defined gaps.
[0014] Also, in accordance with the present invention, a positron
emission tomography scanner includes a pair of detector rings
forming a cylindrical detector array around a body to be imaged
with a well defined gap between the detector rings, a plurality of
detector modules carried by the detector rings producing signals
corresponding to the three dimensional points of interaction in
each of a pair of the detector modules in response to detection of
time coincident photons from positron annihilations in a body
containing a positron emitting compound and means responsive to the
signals to produce tomographic image reconstruction of the spatial
distribution of the amount of positron emitting compound in the
body along the axial field-of-view of the cylindrical detector
array.
[0015] In a further aspect, a small animal positron emission
tomography scanner according to the present invention includes a
plurality of spaced detector rings forming a cylindrical detector
array for receiving the animal with a well defined gap between at
least two of the detector rings which produce signals corresponding
to positron annihilation radiation from the animal and means
responsive to the signals to produce image reconstruction of the
distribution of a positron emitting radiopharmaceutical in the
animal.
[0016] The tomography scanner of the present invention compensates
for missing axial data by using 3D iterative, statistical,
reconstruction methods that do not specifically require complete
and uniform spatial sampling, e.g. 3D OSEM, 3D MAP algorithms, and
the like as discussed above, and that incorporate a model of the
physics and geometry of the radiation detection/emission process
during image reconstruction.
[0017] Some of the advantages of the present invention over the
prior art are that the tomography scanner of the present invention
can use less detector materials to span the same axial length, is
less expensive and easier to manufacture, permits direct coupling
of scintillators to photomultiplier tubes with substantially less
light loss as occurs with light guides and uses conventional
techniques or methods for reconstructing useful images along the
full axial length of the scanner including the gap regions.
[0018] The above and still further features and advantages of the
present invention will become apparent from the following
description of preferred embodiments of the invention, particularly
when taken in conjunction with the accompanying drawings, wherein
like reference numerals are used to designate like or similar
components thereof.
BRIEF DESCRIPTION OF THE DRAWINGS
[0019] FIGS. 1A and 1B are diagrammatic front and side views,
respectively, of a tomography scanner according to the present
invention.
[0020] 1C is a broken perspective of a detector ring including
modules carrying detector elements.
[0021] FIG. 2 is a diagrammatic side view taken along line 2-2 of
FIG. 1 excluding the body O.
[0022] FIG. 3 is a diagrammatic side view taken along line 2-2 of
FIG. 1 with the body O undergoing reciprocating movement.
[0023] FIG. 4 is a diagrammatic side view of a modification of a
tomography scanner according to the present invention where crystal
arrays are tilted with respect to one another.
[0024] FIGS. 5A and 5B are 3D FORE/2D filtered backprojection
images of point sources distributed in a coronal plane containing a
ring diameter and imaged with a simulated small animal PET scanner
with continuous detector rings and with a gap between detector
rings, respectively.
[0025] FIGS. 6A and 6B are 3D FORE/2D filtered backprojection
images of a coronal plane through a simulated Defrise phantom
containing the axis of the scanner and imaged with a simulated
small animal PET scanner with continuous detector rings and with a
gap between detector rings, respectively.
[0026] FIG. 6C is an image from a scanner with a gap between
detector rings with data reconstructed with a 3D OSEM algorithm
tailored to the gap geometry.
DESCRIPTION OF THE PREFERRED EMBODIMENTS
[0027] Tomography and tomographic images refer to images that
together portray in three dimensions some property of an object
being imaged. Commonly, such images may be in the form of a
sequence of consecutive two dimensional transverse sections closely
spaced along the axis of a tomography scanner to span the entire
axial field-of-view of the scanner and the object therein. A
"property" portray in such images can be, but it not limited to,
the spatial distribution and frequency of occurrence of positron
annihilation sites in the object, and a "property" may also refer
to the distribution of attenuation coefficients, the location and
amount of a light emitting compound distributed within the object,
the amount and location of contrast material introduced into the
object, and other such processes and phenomenon.
[0028] The term "detector ring" as used herein means an annular
structure surrounding an imaging target or body formed of detector
material that is responsive to incident radiation, such as x-rays,
gamma rays, photon pairs from positron annihilation and the like. A
detector ring may be formed of various materials and may be
assembled in various geometries. Detector rings may be formed of
scintillation material, e.g. lutetium oxyorthosilicate, or may be
formed of solid state radiation detection material, e.g. intrinsic
germanium. Detector rings can include material that is continuous
in the circumferential direction, e.g. an annulus of scintillation
material or solid state detector material, or can include
independent segments of such materials tiled around the
circumference of a detector ring to form a polygonal, rather than
circular, ring geometry. Such discrete detector segments are
referred to herein as "detector modules" to distinguish such
segments from continuous rings of material. Detector modules may be
further subdivided into smaller parts referred to herein as
"detector elements". Commonly, many small detector elements would
be packed together to form a detector module. A detector module
would be coupled to a device sensitive to emissions from the
material when struck by incident radiation, and signals from the
device would serve to identify the location of the point of
interaction of the radiation in the detector module ultimately
necessary for creation of tomographic images. Solid state or
continuous detector modules also contain detector elements having
points of axial resolution with the spacing therebetween referred
to as pitch.
[0029] In the case of scintillation detector elements, a number of
readout devices are available for receiving the signals, including
avalanche photodiodes (APDs) that may be coupled to individual
detector elements in a detector module. Position-sensitive APDs can
be coupled to a detector module made up of closely packed arrays of
individual, optically isolated detector elements, and arrays of
detector elements can be coupled to position-sensitive
photomultiplier tubes. Each of these "readout" methods serves the
purpose of producing electrical signals in response to the signals
generated in the detector material by incident radiation. The
electrical signals encode the position of the interaction within
the detector module in either two or three dimensions (three if the
detector module is capable of generating signals that also depend
on the radial depth of penetration of the radiation into the module
before interaction, i.e. a depth-of-interaction detector
module).
[0030] In a preferred embodiment, each detector ring is formed of
modules containing arrays of depth-of-interaction (DOI) capable
phoswich scintillator detector elements, each formed of two or more
scintillators with differing light decay times optically connected
to one another end-on as described in the above mentioned U.S. Pat.
No. 6,288,399 to Andreaco et al, the above mentioned Weinhard et al
article and in "Effect of Depth of Interaction Decoding on
Resolution in PET: A Simulation Study," Astakhov, V.; Gumplinger,
P.; Moisan, C.; Ruth, T. J.; and Sossi, V., IEEE Transactions on
Nuclear Science, Vol. 50, No. 5, October 2003; U.S. Pat. No.
4,843,245 to Lecomte, UK Patent No. GB 2 378 112 to Lecoq, and U.S.
Pat. No. 6,362,479 B1 to Andreaco et al. Depth-of-interaction
information can also be obtained by other means including
measurement of the differential output of light from each end of a
scintillation crystal of a single type or by the means described in
UK Patent No. GB 2 198 620 to Yamashita et al, U.S. Pat. No.
4,831,263 to Yamashita, U.S. Pat. No. 6,087,663 to Moisan et al,
U.S. Pat. No. 5,349,191 to Rogers and U.S. Pat. No. 5,122,667 to
Thompson. Depth-of-interaction detectors are preferred because the
number of lines-of-response penetrating the imaging volume is
increased substantially compared to a non-DOI system with the same
general ring and detector geometry and helps reduce the DOI
parallax effect in the axial direction, as well as the transaxial
direction. The present invention will be described hereinafter
relative to a small animal PET scanner; however, it is understood
that the concept of the present invention can be used with any
tomographic scanner acquiring data in three dimensions, can be used
for human or animal imaging and can utilize a variety of ring
geometries, detector materials and readout methods.
[0031] As shown in FIGS. 1A, 1B, 1C and 2, a PET scanner 10
according to the present invention includes a cylindrical detector
array 11 formed of two or more detector rings 12 of detector
modules 13, the detector rings having a diameter D surrounding body
O (small animal) forming an imaging target or object having a
positron annihilation site P and having an angle of obliquitye. The
detector modules 13 are made of a matrix (13.times.13) of detector
elements 15. The detector rings 12 are spaced from each other by a
well defined axial gap G which may or may not be constant from ring
to ring. Dashed lines 14 indicate possible lines-of-flight of
annihilation photons, and arrowheads 16 indicate points of
interaction within the detector rings. The axial direction is
indicated at A. In the preferred embodiment, the detector rings are
stationary and the body O or imaging target is also stationary.
[0032] Data collection is supplied to a processor means 17 for
image reconstruction. The preferred methods of image reconstruction
for PET scanners according to the present invention are iterative
statistical methods, such as 3D OSEM and 3D MAP, in which a model
of the physical and geometrical response of the imaging system can
be included in the image reconstruction process to correct for
deleterious physical and/or geometrical effects including gaps in
the detector array. These reconstruction methods could also be
applied to the variations described below.
[0033] The body O, such as a small animal, can, during imaging, be
axially translated within the scanner as shown in FIG. 3, or the
scanner can be translated with respect to the animal, as a means of
reducing axial image degradation caused by gaps or spaces between
detector rings in accordance with the present invention.
[0034] Another modification of the PET scanner according to the
present invention is shown in FIG. 4 wherein detector modules
within two detector rings are tilted or angled toward one another
to change the shape of the gap. In this modification, the ring
diameter varies along the axial direction of the scanner and is not
constant. In addition, tilting the detector modules opens small
transaxial gaps (not shown in the drawings) between detector
modules, the magnitude of which also varies with axial position.
Additional detector rings can be added by tiling the perimeter of
an axially oriented polygon or an ellipse centered on the geometric
center of the scanner.
[0035] It has been observed that a prior art PET scanner with no
gaps between detector rings will exhibit certain inherent
transverse and axial imaging properties when image data are
reconstructed with conventional 3D methods. In accordance with the
present invention, the gap has a dimension where axial performance
is acceptable compared to other distortions introduced by the
geometry and/or physical performance of a normal scanner. The term
"well defined gap" as used herein refers to intentionally designed
gaps between detector rings, as opposed to spacing occurring from
designs intended to produce axially continuous detector rings. The
difference between a well defined gap and the very small
inadvertent spaces that occur in prior art scanners can be
distinguished in the following way. Detector modules formed of
arrays of detector elements can be characterized in part by their
"pitch", the center-to-center spacing between adjacent detector
elements. In prior art scanners, the intention has been to span the
entire axial field-of-view of the device with detector elements of
constant pitch and to minimize any space between elements such as
might be needed to optically isolate one detector element from
adjacent detector elements. If pitch is the sum of the width W of a
detector element and T, the thickness of any material between
elements, the pitch is P=W+T. In the prior art, the intent has been
to make the pitch P as close to W as possible by minimizing the
thickness of intervening material. That is, the width of the spaces
is small compared to the axial pitch of the detector elements. In
contrast, gaps according to the present invention are purposefully
chosen such that the gap width between detector rings is as great
as or greater than, the pitch of the detector elements in the
detector modules that form the detector rings.
[0036] The gaps G preferably have an axial dimension two to four
times greater than the axial dimension or pitch of a detector
element.
[0037] The point and distributed source responses of a Monte
Carlo-simulated cylindrical PET scanner with two detector rings of
detector modules without a gap and with a well defined gap equal to
four times the pitch of the detector elements in each detector
module are shown in FIGS. 5A, 6A and 5B, 6B, respectively. In FIGS.
5A and 5B, the simulated data were reconstructed by applying the
FORE algorithm followed by 2D filtered back projection (FBP).
Images in the gap region are obtained simply by applying the FORE
algorithm to the acquired 3D image data to create the "virtual"
sinograms that would have been obtained had the gap region actually
been filled with "real" detector rings and then applying any of
several 2D reconstruction methods to these sinograms to obtain
transverse section images of the object. No other computational
methods were used. The FORE+FBP method is an extreme test of the
ability of an algorithm to compensate for missing axial data. If
such methods yield modest reductions in axial imaging performance,
iterative statistical methods that incorporate accurate models of
system geometry and other characteristic of the radiation detection
process, exhibit even less, or no, reduction in axial imaging
performance. With these latter methods the activity distribution in
the object within the gap region is estimated using
lines-of-response that cross the gap region. In accordance with the
present invention, once the gap distribution is known or estimated,
transverse section images through the distribution can be created
as if there had been actual detector rings that span the gap
region.
[0038] FIGS. 5A and 5B show point sources spread over a sagittal
plane that contains the central axis of simulated scanners with no
axial gap between detector rings and with an axial gap between
detector rings equal to four times the pitch (or axial dimension)
of the detector elements in each module of the detector rings,
respectively. The spacing between sources is 5 mm in the horizontal
and vertical directions. In these images, the axial direction is
vertical and the radial direction is left and right. As can be
seen, the apparent size of the point sources generally increases as
one moves to the left and right away from the central axis. This
variation is due entirely to the radial parallax effect that occurs
in all ring-type PET scanners and is present with, or without, a
gap. There is only a slight apparent difference between the row of
points in the 6 mm-wide gap region (middle row of points) of the
two images, and the overall axial variation is very similar between
the images. The measured transaxial FWHM values in the gap region
are nearly identical to those in the gap-free image while the axial
widths of the spots differ only slightly.
[0039] FIGS. 6A and 6B illustrate axial effects revealed by the
Defrise phantom with no axial detector ring gap and an axial gap
between detector rings equal to four times the pitch of the
detector elements in each ring, respectively. The phantom is formed
of parallel, coaxial cylinders 4.3 mm thick, every other one of
which filled with a pure positron-emitter and the intervening solid
cylinders filled with nothing. The phantom detects defects that
might exist in axial imaging performance and is the most commonly
used phantom to reveal axial imaging flaws. Images shown in FIGS.
6A and 6B were reconstructed with the FORE+FBP method. As noted,
FIG. 6A has no gap and FIG. 6B has an axial gap G equal to four
times the detector element pitch in each detector ring. The images
of the Defrise phantom show a sagittal plane that contains the
central axis of the scanner. It would be expected that any
distortions due to the "four pitch" gap would be most evident in
the central plane (middle row) and would diminish as one moves away
from the center along the axis of the scanner. In a perfect system,
all of the horizontal bands would be the same width (in the
vertical direction). FIGS. 6A and 6B show that some degradation
occurs in the central plane within the gap region as predicted by
"classical theory". The central band in the gap image is somewhat
wider and fainter than the central band in the no gap image while,
away from the central gap in the axial direction, the images are
nearly identical. The maximum amplitude of the band at the center
of the field of view in the four pitch-wide gap image is
approximately 20% less than the same band in the no gap image. A
maximum degradation in axial response of this magnitude is entirely
acceptable in practice given the practical advantages of being able
to use discontinuous detector arrays. Similar experiments using the
preferred 3D OSEM algorithm with system model to reconstruct images
from a dual ring phoswich-based scanner with a four pitch gap
between rings is shown in FIG. 6C and reveals no significant
degradation in axial resolution (vertical thickness of slabs).
There is no measurable degradation in spatial resolution in the
transverse plane anywhere along the axial field-of-view, outside or
inside the gap region, with any of these methods, including the 3D
FORE/2D filtered backprojection method, i.e. radial and tangential
spatial resolution in the transverse plane at given radius do not
vary with axial position for any of these reconstruction
methods.
[0040] The overall performance of a tomography scanner with a small
axial gap according to the present invention, aside from a modest
degradation in axial response in the gap region with the most
sensitive method, is not significantly different from a tomography
scanner without a gap. The FORE method followed by 2D filtered
backprojection were the only computational methods needed to
achieve this result, these algorithms being proposed in the past
only for use with continuous detector arrays and considered to
absolutely require continuous detector arrays. All of the 3D
reconstruction methods noted above possess the "gap-filling"
property to a greater or lesser extent. In particular, the most
resilient of these methods appear to be those based, not on
classical reconstruction methods, but rather on iterative,
approximating reconstruction methods that incorporate a
mathematical model of a system's geometrical and physical response
to annihilation radiation from an imaging volume surrounded by
depth-of-interaction capable detector modules. The 3D OSEM
algorithm with system modeling, for example, is capable of making
nearly distortion-free image reconstructions including nearly
perfect compensation for gaps according to the present
invention.
[0041] As will be appreciated from the above, in accordance with
the present invention, three-dimensional re-binning and/or
reconstruction methods are able to compensate for missing data
introduced by axial gaps or discontinuities. Accordingly, a
tomography scanner 10 may be designed such that an axial gap is
purposefully incorporated into the scanner. Such a gap reduces the
number of detector elements needed to span a given axial
field-of-view and permits light sources, such as
phoswich/scintillator elements, to be optically coupled to sensors,
such as positron-sensitive photomultiplier tubes, with no light
guides. The present invention can be implemented with iterative, or
statistical, reconstruction methods as well as with classical
algorithms to provide the gap-filling function albeit with
different amounts of axial image degradation.
[0042] Tomography scanners with axial gaps between detector rings
in accordance with the present invention can also include
oscillatory or linear translation of the imaging bed or imaging
gantry in the axial direction as shown in FIG. 3. For a scanner
with an axial gap, the imaging bed or the detector gantry can be
driven to and fro in the axial direction during imaging or
translated uni-directionally, and transaxial lines-of-response
acquired for parts of the object that would otherwise always remain
in the gap region. A translation of the bed or the gantry can move
a point in the object out of the gap region and into the region
directly sampled by the detector arrays. If the animal is
translated through the imaging field or the bed or gantry
oscillates axially at a reasonable frequency without moving the
subject appreciably, an additional benefit accrues, namely that the
usual triangular-like sensitivity profile along the scanner's axis
will appear to become flatter and the statistical properties of the
acquired data will be more uniform over the axial imaging
field-of-view. It is understood that such mechanical movements of
the bed or gantry require the position of the bed or gantry
relative to the target to be known at all times so that lines of
response can be transformed into a coordinate system fixed in the
subject rather than the gantry before image reconstruction.
[0043] Tomography scanners with axial gaps between detector rings
in accordance with the present invention can also use axially
tilted detector modules. Conventional human and animal tomography
scanners are usually designed such that the transverse dimensions
of the detector array are constant in the axial direction. With the
detector arrays tilted in the manner shown in FIG. 4, the effective
axial gap width can be reduced. This manner of covering the imaging
volume results in a variable ring diameter as one moves along the
scanner axis and a slightly increasing transaxial gap width between
modules as one moves toward the axial center of the system from
each end. The effect of the gap is "smeared out" over the
circumference of the array so its effect on the acquired data is
reduced. Accordingly, axial resolution degradation will be less
with detector geometry of FIG. 4, and the 3D
re-binning/reconstruction methods noted earlier, or other post
processing methods are more effective in restoring axial resolution
in the gap region.
[0044] Typically, a PET scanner according to the present invention
will be a stationary ring-type scanner with an aperture appropriate
for small animals. Such a scanner typically includes a gantry
containing the detector rings and an aperture into which the animal
on an imaging bed can be accurately inserted into the imaging field
of view. Such scanners might contain several computers for data
acquisition and for data processing. A particular embodiment for a
small animal PET scanner has a cylindrical detector array diameter
between 10 and 40 cm, an aperture within two detector rings between
7 and 30 cm, a useful transverse field-of-view between 5 and 25 cm,
an axial field of view between 4 and 20 cm and an axial gap between
detector rings of between 2 and 20 mm. In a preferred embodiment,
the animal would be surrounded by depth-of-interaction capable
detector modules to increase spatial sampling within the imaging
volume, each formed of arrays of small cross-section, e.g. 1-2 mm
square, optically isolated scintillation crystals or detector
elements, numbering on the order of 100s of detection elements per
module and 10s of thousands of elements for an entire scanner. As
is understood, positron emission tomography involves the sensing of
signals corresponding to three dimensional points of interaction in
each of a pair of the detector modules in response to detection of
time coincident photons from positron annihilations in a body
containing a positron emitting compound.
[0045] Inasmuch as the present invention is subject to various
modifications and changes in detail, it is intended that all
subject matter discussed above and shown in the accompanying
drawings not be taken in a limiting sense.
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