U.S. patent application number 10/779596 was filed with the patent office on 2004-08-19 for scintillation detector array for encoding the energy, position and time coordinates of gamma ray interactions.
This patent application is currently assigned to CTI PET Systems, Inc.. Invention is credited to Andreaco, Mark S., Casey, Michael E., Nutt, Ronald, Williams, Charles W..
Application Number | 20040159792 10/779596 |
Document ID | / |
Family ID | 26761812 |
Filed Date | 2004-08-19 |
United States Patent
Application |
20040159792 |
Kind Code |
A1 |
Andreaco, Mark S. ; et
al. |
August 19, 2004 |
Scintillation detector array for encoding the energy, position and
time coordinates of gamma ray interactions
Abstract
A scintillation detector which includes a plurality of discrete
scintillators composed of one or more scintillator materials. The
discrete scintillators interact with incident radiation to produce
a quantifiable number of photons with characteristic emission
wavelength and decay time. A light guide is operatively associated
with the scintillation crystals and may be either active or
non-active and segmented or non-segmented depending upon the
embodiment of the design. Photodetectors are provided to sense and
quantify the scintillation light emissions. The process and system
embodying various features of the present invention can be utilized
in various applications such as SPECT, PET imaging and simultaneous
PET systems. In accordance with the present invention, the detector
array of the present invention incorporates either a single
scintillator layer of discrete scintillators or discrete
scintillators composed of two stacked different layers that can be
the same scintillator material or of two different scintillator
materials. In either case the different layers are composed of
materials that have distinctly different decay times. The variants
in these figures are the types of optical detectors which are used,
i.e. photomultipliers and/or photodiodes, whether or not a
segmented optical planar light guide is used, and whether the
planar light guide is active or non-active. If a segmented optical
planar light guide is used then the variant is whether the
configuration is inverted or non-inverted.
Inventors: |
Andreaco, Mark S.;
(Knoxville, TN) ; Williams, Charles W.; (Powell,
TN) ; Nutt, Ronald; (Knoxville, TN) ; Casey,
Michael E.; (Knoxville, TN) |
Correspondence
Address: |
PITTS AND BRITTIAN P C
P O BOX 51295
KNOXVILLE
TN
37950-1295
US
|
Assignee: |
CTI PET Systems, Inc.
Knoxville
TN
|
Family ID: |
26761812 |
Appl. No.: |
10/779596 |
Filed: |
February 13, 2004 |
Related U.S. Patent Documents
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Application
Number |
Filing Date |
Patent Number |
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10779596 |
Feb 13, 2004 |
|
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09882101 |
Jun 15, 2001 |
|
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|
09882101 |
Jun 15, 2001 |
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09271770 |
Mar 18, 1999 |
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60079279 |
Mar 25, 1998 |
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Current U.S.
Class: |
250/363.03 |
Current CPC
Class: |
A61B 6/037 20130101;
G01T 1/1648 20130101; G01T 1/1644 20130101 |
Class at
Publication: |
250/363.03 |
International
Class: |
G01T 001/164 |
Claims
Having thus described the aforementioned invention, we claim:
1. A scintillation detector array for encoding energy, position and
time coordinates of gamma ray interactions for use in Positron
Emission Tomography imaging, said scintillation detector array
comprising: a plurality of discrete scintillator elements which
interact with incident gamma-rays to produce a quantifiable number
of scintillation photons, wherein each of said plurality of
discrete scintillators is composed of a first layer having a first
selected decay time and a second layer having a second selected
decay time, wherein said first selected decay time is not equal to
said second selected decay time, and further wherein said first
layer is composed of a first selected scintillator material and
said second layer is composed of a second selected scintillator
material and wherein said first and second selected scintillator
materials are stacked one upon the other, whereby a pulse shape
discrimination technique is used to determine which said layer the
gamma ray interacts; an optical detector associated with each of
said plurality of discrete scintillator elements and positioned for
sensing and quantifying said scintillation photons exiting each of
said plurality of discrete scintillator elements; a continuous
light guide having first and second planar surfaces disposed
between said plurality of discrete scintillator elements and said
associated optical detectors for distributing scintillation photons
exiting said plurality of discrete scintillators to said associated
optical detectors; and a means operatively associated with said
scintillation detector array for determining time, energy, depth
and transverse and longitudinal position coordinates of gamma ray
interactions in said plurality of discrete scintillator
elements.
2. The scintillator detector array of claim 1 wherein said first
and said second layers are composed of High-Z scintillator
materials.
3. The scintillation detector array of claim 1 wherein said
plurality of discrete scintillator elements, which interact with
incident gamma-rays to produce a quantifiable number of
scintillation photons, is arranged in an (m).times.(n) array, and
said plurality of optical detectors is arranged in an (q).times.(p)
array, wherein said plurality of optical detectors is for sensing
and quantifying said scintillation photons exiting each of said
plurality of discrete scintillator elements.
4. The scintillator detector array of claim 3 wherein said
(m).times.(n) array equals said (q).times.(p) array.
5. The scintillator detector array of claim 3 wherein said
(m).times.(n) array does not equal said (q).times.(p) array.
6. The scintillator detector array of claim 2 wherein said first
and said second layer of each of said plurality of discrete
scintillator elements is composed of LSO.
7. The scintillator detector array of claim 2 wherein said High-Z
scintillator material is selected from a group consisting of LSO,
LYSO, LGSO, GSO, LuAP, and YAP.
8. The scintillator detector array of claim 2 wherein said first
layer is composed of a first selected scintillator material and
said second layer is composed of a second selected scintillator
material.
9. The scintillator detector array of claim 8 wherein said first
selected scintillator material and said second selected
scintillator material are selected for use in techniques for
separating low and high energies.
10. The scintillator detector array of claim 8 wherein said first
selected scintillator material and said second selected
scintillator material are selected for use in techniques for
determining depth of interaction of the gamma rays with said
plurality of discrete scintillator elements.
11. The scintillator detector array of claim 8 wherein said first
selected scintillator material and said second selected
scintillator material are selected for use in techniques for
distinguishing pulse heights of gamma ray interactions.
12. The scintillator detector array of claim 1 wherein said first
selected scintillator material is YSO and said second selected
scintillator material is a High-Z scintillator material.
13. The scintillator detector array of claim 1 wherein said first
selected scintillator material is LSO and said second selected
scintillator material is GSO.
14. The scintillator detector array of claim 1 wherein said first
selected scintillator material is YSO and said second selected
scintillation material is LSO.
15. The scintillator detector array of claim 1 wherein said light
guide is active.
16. The scintillation detector array of claim 1 wherein said light
guide is non-active.
17. A scintillation detector array for encoding energy, position
and time coordinates of gamma ray interactions for use in Positron
Emission Tomography imaging, said scintillation detector array
comprising: a plurality of discrete scintillator elements which
interact with incident gamma-rays to produce a quantifiable number
of scintillation photons, wherein each of said plurality of
discrete scintillators is composed of a first layer having a first
selected decay time and a second layer having a second selected
decay time, wherein said first selected decay time is not equal to
said second selected decay time, and further wherein said first and
said second layers are composed of High-Z scintillator materials,
and further wherein said first layer is composed of a first
selected scintillator material and said second layer is composed of
a second selected scintillator material and wherein said first and
second selected scintillator materials are stacked one upon the
other, whereby a pulse shape discrimination technique is used to
determine which said layer the gamma ray interacts; an optical
detector associated with each of said plurality of discrete
scintillator elements and positioned for sensing and quantifying
said scintillation photons exiting each of said plurality of
discrete scintillator elements; a continuous light guide having
first and second planar surfaces disposed between said plurality of
discrete scintillator elements and said associated optical
detectors for distributing scintillation photons exiting said
plurality of discrete scintillators to said associated optical
detectors; and a means operatively associated with said
scintillation detector array for determining time, energy, depth
and transverse and longitudinal position coordinates of gamma ray
interactions in said plurality of discrete scintillator
elements.
18. The scintillation detector array of claim 17 wherein said
plurality of discrete scintillator elements, which interact with
incident gamma-rays to produce a quantifiable number of
scintillation photons, is arranged in an (m).times.(n) array, and
said plurality of optical detectors is arranged in an (q).times.(p)
array, wherein said plurality of optical detectors is for sensing
and quantifying said scintillation photons exiting each of said
plurality of discrete scintillator elements.
19. The scintillator detector array of claim 18 wherein said
(m).times.(n) array equals said (q).times.(p) array.
20. The scintillator detector array of claim 18 wherein said
(m).times.(n) array does not equal said (q).times.(p) array.
21. The scintillator detector array of claim 17 wherein said light
guide is active.
22. The scintillation detector array of claim 17 wherein said light
guide is non-active.
23. A scintillation detector array for encoding energy, position
and time coordinates of gamma ray interactions for use in Positron
Emission Tomography imaging, said scintillation detector array
comprising: a plurality of discrete scintillator elements which
interact with incident gamma-rays to produce a quantifiable number
of scintillation photons, wherein each of said plurality of
discrete scintillators is composed of a first layer having a first
selected decay time and a second layer having a second selected
decay time, wherein said first selected decay time is not equal to
said second selected decay time, and further wherein said first and
said second layers are composed of High-Z scintillator materials,
and further wherein said first layer is composed of a first
selected scintillator material and said second layer is composed of
a second selected scintillator material and wherein said first and
second selected scintillator materials are stacked one upon the
other, whereby a pulse shape discrimination technique is used to
determine which said layer the gamma ray interacts; an optical
detector associated with each of said plurality of discrete
scintillator elements and positioned for sensing and quantifying
said scintillation photons exiting each of said plurality of
discrete scintillator elements; a continuous light guide having
first and second planar surfaces optically bonded to said plurality
of discrete scintillator elements, whereby said plurality of
discrete scintillator elements is disposed between said light guide
and said optical detectors, wherein said plurality of discrete
scintillator elements distribute scintillation photons exiting said
plurality of discrete scintillators to said associated optical
detectors; and a means operatively associated with said
scintillation detector array for determining time, energy, depth
and transverse and longitudinal position coordinates of gamma ray
interactions in said plurality of discrete scintillator
elements.
24. The scintillation detector array of claim 23 wherein said
plurality of discrete scintillator elements, which interact with
incident gamma-rays to produce a quantifiable number of
scintillation photons, is arranged in an (m).times.(n) array, and
said plurality of optical detectors is arranged in an (q).times.(p)
array, wherein said plurality of optical detectors is for sensing
and quantifying said scintillation photons exiting each of said
plurality of discrete scintillator elements.
25. The scintillator detector array of claim 24 wherein said
(m).times.(n) array equals said (q).times.(p) array.
26. The scintillator detector array of claim 24 wherein said
(m).times.(n) array does not equal said (q).times.(p) array.
27. The scintillator detector array of claim 23 wherein said first
and said second layer of each of said plurality of discrete
scintillator elements is composed of LSO.
28. The scintillator detector array of claim 23 wherein said High-Z
scintillator material is selected from a group consisting of LSO,
LYSO, LGSO, GSO, LuAP, and YAP.
29. The scintillator detector array of claim 23 wherein said first
layer is composed of a first selected scintillator material and
said second layer is composed of a second selected scintillator
material.
30. The scintillator detector array of claim 29 wherein said first
selected scintillator material and said second selected
scintillator material are selected for use in techniques for
separating low and high energies.
31. The scintillator detector array of claim 29 wherein said first
selected scintillator material and said second selected
scintillator material are selected for use in techniques for
determining depth of interaction of the gamma rays with said
plurality of discrete scintillator elements.
32. The scintillator detector array of claim 29 wherein said first
selected scintillator material and said second selected
scintillator material are selected for use in techniques for
distinguishing pulse heights of gamma ray interactions.
33. The scintillator detector array of claim 29 wherein said first
selected scintillator material is YSO and said second selected
scintillator material is a High-Z scintillator material.
34. The scintillator detector array of claim 29 wherein said first
selected scintillator material is LSO and said second selected
scintillator material is GSO.
35. The scintillator detector array of claim 29 wherein said first
selected scintillator material is YSO and said second selected
scintillation material is LSO.
36. The scintillator detector array of claim 23 wherein said light
guide is active.
37. The scintillation detector array of claim 23 wherein said light
guide is non-active.
38. A scintillation detector array for encoding energy, position
and time coordinates of gamma ray interactions for use in Positron
Emission Tomography imaging, said scintillation detector array
comprising: a plurality of discrete scintillator elements which
interact with incident gamma rays to produce a quantifiable number
of scintillation photons, wherein each of said plurality of
discrete scintillators is composed of a first layer having a first
selected decay time and a second layer having a second selected
decay time, wherein said first selected decay time is not equal to
said second selected decay time, and further wherein said first
layer is composed of a first selected scintillator material and
said second layer is composed of a second selected scintillator
material and wherein said first and second selected scintillator
materials are stacked one upon the other, whereby a pulse shape
discrimination technique is used to determine which said layer the
gamma ray interacts; an optical detector associated with each of
said plurality of discrete scintillator elements and positioned for
sensing and quantifying said scintillation photons exiting each of
said plurality of discrete scintillator elements wherein said
plurality of discrete scintillator elements, which interact with
incident gamma rays to produce a quantifiable number of
scintillation photons, is arranged in an (m).times.(n) array, and
said plurality of optical detectors is arranged in an (q).times.(p)
array, wherein said (m).times.(n) array does not equal said
(q).times.(p) array and further wherein said plurality of optical
detectors is for sensing and quantifying said scintillation photons
exiting each of said plurality of discrete scintillator elements; a
continuous light guide having first and second planar surfaces
disposed between said plurality of discrete scintillator elements
and said associated optical detectors for distributing
scintillation photons exiting said plurality of discrete
scintillators to said associated optical detectors; and a means
operatively associated with said scintillation detector array for
determining time, energy, depth and transverse and longitudinal
position coordinates of gamma ray interactions in said plurality of
discrete scintillator elements.
39. The scintillator detector array of claim 38 wherein said first
and said second layers are composed of High Z scintillator
materials.
40. The scintillator detector array of claim 39 wherein said first
and said second layer of each of said plurality of discrete
scintillator elements is composed of LSO.
41. The scintillator detector array of claim 39 wherein said High-Z
scintillator material is selected from a group consisting of LSO,
LYSO, LGSO, GSO, LuAP, and YAP.
42. The scintillator detector array of claim 39 wherein said first
layer is composed of a first selected scintillator material and
said second layer is composed of a second selected scintillator
material.
43. The scintillator detector array of claim 42 wherein said first
selected scintillator material and said second selected
scintillator material are selected for use in techniques for
separating low and high energies.
44. The scintillator detector array of claim 42 wherein said first
selected scintillator material and said second selected
scintillator material are selected for use in techniques for
determining depth of interaction of the gamma rays with said
plurality of discrete scintillator elements.
45. The scintillator detector array of claim 42 wherein said first
selected scintillator material and said second selected
scintillator material are selected for use in techniques for
distinguishing pulse heights of gamma ray interactions.
46. The scintillator detector array of claim 38 wherein said first
selected scintillator material is YSO and said second selected
scintillator material is a High Z scintillator material.
47. The scintillator detector array of claim 38 wherein said first
selected scintillator material is LSO and said second selected
scintillator material is GSO.
48. The scintillator detector array of claim 38 wherein said first
selected scintillator material is YSO and said second selected
scintillation material is LSO.
49. The scintillator detector array of claim 38 wherein said light
guide is active.
50. The scintillation detector array of claim 38 wherein said light
guide is non-active.
Description
[0001] This is a continuation of application Ser. No. 09/882,101,
filed on Jun. 15, 2001, which was a continuation-in-part of
application Ser. No. 09/271,770, filed on Mar. 18, 1999, which was
a non-provisional application that claimed the priority benefit of
U.S. Provisional Application No. 60/079,279, filed Mar. 25, 1998.
Accordingly, this non-provisional application also claims the
priority benefit of U.S. Provisional Application No. 60/079,279,
filed Mar. 25, 1998.
TECHNICAL FIELD
[0002] The present invention relates to an apparatus capable of
determining the energy, position and time coordinates of light
emission induced by interactions of gamma-rays in a planar array of
discrete scintillator detectors having either a segmented or
non-segmented light guide. The features of the present invention
find particular application in the field of medical imaging whereby
a single device can be used for Single Photon Imaging which
includes traditional Gamma Cameras, Planar Imaging, Single Photon
Emission Computed Tomography (SPECT) with or without Coincidence
Photon Imaging and Positron Emission Tomography (PET). When
operated in the SPECT mode, the present invention is comparable to
existing high resolution SPECT systems. When operated in the PET
mode, the present invention is an improvement over existing PET
systems in that the device may be operated either in Pulse Height
Discrimination mode or in Pulse Shape Discrimination mode thereby
enabling depth of interaction encoding resulting in improved
spatial resolution. Emission Computed Tomography (ECT) systems
provide a means for sensing, and quantitatively measuring
biochemical and/or physiological changes in the human body or other
living organism. However, the use of the invention is not limited
to such application.
BACKGROUND ART
[0003] Devices for detecting the distribution of gamma rays
transmitted or emitted through objects to study the compositions or
functions of the objects are well known to the art, e.g. the
techniques referred to as Emission Computed Tomography can be
divided into two specific classes; Single Photon Emission Computed
Tomography (SPECT) uses radiotracers which emit gamma rays but do
not emit positrons and Positron Emission Tomography (PET) which
uses radiotracers that emit positrons. Therefore, the fundamental
physical difference between the two techniques is that PET uses
annihilation coincidence detection. The PET technique can
determine, in-vivo, biochemical functions, on the injection of
biochemical analog radiotracer molecules that emit positrons in a
living body. The positrons annihilate with surrounding electrons in
the subject body to produce a pair of gamma-rays, each having 511
keV of photon energy; traveling in nearly opposite directions. The
detection of a pair of annihilation gamma-rays by two opposed
detectors allows for the determination of the location and
direction in space of a trajectory line defined by the opposite
trajectories of the gamma-rays. Tomographic reconstruction is then
used to superpose the numerous trajectory lines obtained by
surveying the subject with an array of detectors to image the
distribution of radiotracer molecules in the living body.
[0004] Emission Computed Tomography systems employ a variety of
geometric configurations for the gamma-ray detectors. The choice of
configuration is typically dictated by the manufacturer's desired
system performance and cost. The detector design must be capable of
providing accurate estimates of gamma-ray energy, position
coordinates, and in addition in the case of PET, coincidence time
interval to reconstruct an image of the distribution of the
radiotracer for in vivo studies. An example of such a device is
disclosed in U.S. Pat. No. 4,750,972 to Casey et al., the
disclosure of which is incorporated herein by reference and relied
upon.
[0005] The position encoder and detector system disclosed by Casey
et. al., is a two dimensional photon counting position encoder
detector system, i.e., the array of scintillation crystals provides
only the transverse coordinates of the photon interaction; the
longitudinal photon interaction position of the excited
scintillation crystal is undetermined. Photons impinging upon such
detector systems at angles other than normal may traverse the path
of several scintillation crystals resulting in uncertainty of their
trajectory lines thereby degrading the image resolution due to
parallax error.
[0006] In U.S. Pat. No. 3,919,556 by Berninger, Berninger discloses
a gamma camera having a light pipe member in which the output
member has a plurality of concave depressions conforming to the
outer surfaces of the convexly curved phototube glass faceplates.
Berninger teaches that the primary function of the light pipe is
simply that of providing a refractive index match between the glass
backing of the scintillator and the glass envelope of the
phototubes.
[0007] A detector system capable of providing both the transverse
and longitudinal position of photon interactions in scintillation
crystals was disclosed in U.S. Pat. No. 4,843,245 by Lecomte. The
approach involves the use of two scintillation crystals of
different decay times which are stacked one upon the other. The
position of photon interaction is determined by the Pulse Shape
Discrimination technique. This method though capable of providing
the transverse and longitudinal position coordinates of photon
interactions in scintillation crystal detector systems will result
in reduced system efficiency if the overall scintillator depth is
constant for two different scintillator materials. If the
scintillators are increased in length to compensate for the
efficiency loss then the system resolution will be degraded.
[0008] Another approach to determine the transverse and
longitudinal positions of photon interactions in scintillation
crystal detector systems was disclosed in U.S. Pat. No. 5,122,667
by Thompson. The approach differs from that of Lecomte in that a
single scintillator is used, further the method does not depend on
decay time differences. The method employs the use of a
scintillation light absorbing band located at the median
interaction coordinate for a specific energy along the longitudinal
axis of the scintillation crystal. The net effect is to divide the
scintillation crystal into two regions whereby the photon is
equally likely to interact. Pulse Height Discrimination is used to
determine which of the two regions of the scintillator the photon
interacted. This approach has the undesired effect of reducing the
total collected scintillation light and of causing the Compton
continuum of the high light yield scintillator to overlap the
photopeak region of the low light yield scintillator. The result is
inherent uncertainty in the contribution of scatter to the full
energy photopeak.
[0009] In U.S. Pat. No. 5,349,191 Rogers discloses a method for
determining the transverse and longitudinal position coordinates
for interactions in scintillation crystal arrays which depends on
the continuous variation of the total collected light with the
longitudinal photon interaction coordinate of the light emission.
The continuous variation in collected light requires a complex
calibration of each detector as a function of longitudinal photon
interaction coordinate from a collimated beam of photons directed
at known positions along the length of the scintillator. This
calibration method is difficult to implement for large arrays of
scintillators.
[0010] In U.S. Provisional Application Serial No. 60/037,519, filed
on Feb. 10, 1997, and U.S. Provisional Application Serial No.
60/042,002, filed on Apr. 16,1997, Moisan and Andreaco et. al.
disclosed a device capable of determining the transverse and
longitudinal coordinates of light emission induced by the
interaction of photons in an array of photon detectors having a
plurality of scintillation light guides. The device uses two or
more layers of stacked scintillators all composed of the same
scintillator material. Pulse Height Discrimination is used to
determine which scintillator layer the photon interaction occurs.
The device requires a difference in the light output from the two
stacked scintillator layers of at least a factor of 1.5 times for
the pulse height discrimination technique to be practicable. The
approach has the undesired effect of causing the Compton continuum
of the high light yield scintillator (which is nearest to the
subject under study) to overlap the photopeak region of the low
light yield scintillator. The result is inherent uncertainty in the
contribution of scatter to the full energy photopeak.
[0011] The detector systems described in the above stated US
Patents when applied to medical imaging are specific to usage in
PET. The predominant scintillator material is Bismuth Germanate
(BGO), though other materials have been proposed or used (see Table
1). The SPECT detector systems are different in that Thallium doped
Sodium-Iodide (NaI(Tl)) is used exclusively as the scintillator
material. Further these systems use large continuous slabs of
NaI(Tl) optically coupled to a continuous light guide. Anger logic
is used for scintillation event localization. The exception to
continuous NaI(Tl) slab detector systems for SPECT imaging was
disclosed by Govaert in U.S. Pat. No. 4,267,452. This detector
system is unique as a SPECT detector in that it is segmented. The
segmentation of the NaI(Tl) is similar to PET block detector
designs which use an active light guide. (For clarification
detector light guides are of two general types: non-active light
guides are composed of optical materials other than the
scintillator; active light guides are composed of scintillator
materials). The detector system disclosed by Govaert does not
result in discrete scintillator elements whereby each element is a
separate detector. Instead the segmentation process results in a
block of NaI(Tl) that is subdivided into elements that share a
common light guide of active scintillator material, i.e. the
NaI(Tl) is not cut all the way through.
[0012] The unique differences in SPECT and PET imaging modalities
have resulted in detector designs which are suitable for their
intended use in either SPECT or PET, but not both. However, the use
of Fluorodeoxyglucose (FDG) with SPECT imaging systems has resulted
in the application of SPECT detector designs in PET imaging. One
problem in the application of SPECT detector designs in PET is that
relatively thin scintillation crystals are preferred in Anger
cameras to provide better intrinsic resolution and image detail.
This results in poor detection efficiency in PET since the
effective-Z and density of NaI(Tl) provides lower stopping power at
511 keV relative to PET scintillators (see Table 1). The efficiency
of SPECT detector systems is further reduced by the use of
absorptive collimation. The continuous slab of NaI(Tl) precludes
the elimination of absorptive collimation.
[0013] SPECT detector system designs which are intended to bridge
both SPECT and PET imaging modalities are known as hybrid devices.
These systems have increased the NaI(Tl) scintillator thickness for
higher efficiency and have added coincidence detection circuitry
and attenuation corrections. Despite these changes the continuous
slab of NaI(Tl) scintillator detector designs are inferior to PET
specific detector designs in terms of system performance.
[0014] The hybrid SPECT detector designs have compromised their
SPECT performance while providing inferior PET performance. A need
has arisen for a hybrid PET/SPECT detector system which provides
state of the art SPECT and PET system performance which does not
suffer from the heretofore stated disadvantages.
DISCLOSURE OF THE INVENTION
[0015] In accordance with the various features of this invention, a
scintillation detector is provided which includes a plurality of
discrete scintillators composed of one or more scintillator
materials. The discrete scintillators interact with incident
radiation to produce a quantifiable number of photons with
characteristic emission wavelength and decay time. A light guide is
operatively associated with the scintillation crystals and may be
either active or non-active and segmented or non-segmented
depending upon the embodiment of the design. Photodetectors are
provided to sense and quantify the scintillation light emissions.
The process and system embodying various features of the present
invention can be utilized in various applications such as SPECT and
PET imaging systems. In accordance with the present invention, the
detector array of the present invention incorporates either a
single layer of discrete scintillators or discrete scintillators
composed of two stacked different layers that can be the same
scintillator material or of two different scintillator materials.
In either case the different layers are composed of materials that
have distinctly different decay times. The variants in these
figures are the types of optical detectors which are used, i.e.
photomultipliers and/or photodiodes, whether or not a segmented
optical light guide is used, and whether the light guide is active
or non-active. If a segmented optical light guide is used then the
variant is whether the configuration is inverted or
non-inverted.
BRIEF DESCRIPTION OF THE DRAWINGS
[0016] The above mentioned features of the invention will become
more apparent from consideration of the following description when
read together with the accompanying drawings, in which:
[0017] FIG. 1 is a perspective view of a medical imaging scanner
embodying scintillation detector arrays for encoding the energy,
position and time coordinates of gamma-ray interactions.
[0018] FIG. 2 is comprised of FIG. 2a through and inclusive of FIG.
2e. These figures present various views of the detector head of the
system illustrated in FIG. 1 and the detector blocks as mounted in
the detector head. FIG. 2a depicts an 8.times.10 array of detector
blocks optically coupled to a 9.times.11 array of photomultiplier
tubes. FIGS. 2b and 2c illustrate cross-sectional views of the
detector head in which the array of detector blocks are optically
coupled to an array of photomultiplier tubes. FIGS. 2d and 2e
illustrate cross-sectional views of the detector head in which the
array of detector blocks is optically coupled to a photodiode
array.
[0019] FIG. 3 is comprised of FIG. 3(a) through and inclusive of
FIG. 3(l). These figures present a perspective view of the types of
detector blocks which could be incorporated in the detector head
illustrated in FIG. 2. The common feature in these figures is the
embodiment of the design incorporating discrete scintillators
composed of two stacked different scintillator materials of
different decay times. The variants in these figures are the types
of optical detectors which are used, i.e. photomultipliers and/or
photodiodes. The other variant is whether or not a segmented
optical light guide is used. If a segmented optical light guide is
used then one variant is whether the configuration is inverted or
non-inverted. Another variant is whether the light guide is active
or non-active.
[0020] FIG. 3(a) is a perspective view of one detector block from
the detector head illustrated in FIG. 2 for one preferred
embodiment of the design incorporating discrete scintillators
composed of two stacked different scintillator materials of
different decay times optically coupled to a non-active segmented
inverted light guide which is optically coupled to photomultiplier
tubes (PMTs). The PMTs are the optical detectors.
[0021] FIG. 3(b) is a perspective view of one detector block from
the detector head illustrated in FIG. 2 for another embodiment of
the design incorporating discrete scintillators composed of two
stacked different scintillator materials of different decay times
optically coupled to a non-active segmented inverted light guide.
The light guide is optically coupled to PMTs in addition a
photodiode array is optically coupled to one of the scintillator
arrays. The PMTs and photodiode arrays are the optical
detectors.
[0022] FIG. 3(c) is a perspective view of one detector block from
the detector head illustrated in FIG. 2 for another embodiment of
the design incorporating discrete scintillators composed of two
stacked different scintillator materials of different decay times
optically coupled to a non-active segmented inverted light guide.
The light guide is optically coupled to a photodiode array. The
photodiode array is the optical detector.
[0023] FIG. 3(d) is a perspective view of one detector block from
the detector head illustrated in FIG. 2 for another embodiment of
the design incorporating discrete scintillators composed of two
stacked different scintillator materials of different decay times
optically coupled to a non-active segmented inverted light guide.
The light guide is optically coupled to a photodiode array in
addition a second photodiode array is optically coupled to one of
the scintillator arrays. The two photodiode arrays are the optical
detectors.
[0024] FIG. 3(e) is a perspective view of one detector block from
the detector head illustrated in FIG. 2 for another embodiment of
the design incorporating discrete scintillators composed of two
stacked different scintillator materials of different decay times.
One of the two scintillator arrays is optically coupled to a
photodiode array. The photodiode array is the optical detector.
[0025] FIG. 3(f) is a perspective view of one detector block from
the detector head illustrated in FIG. 2 for another embodiment of
the design incorporating discrete scintillators composed of two
stacked different scintillator materials of different decay times.
One photodiode array is optically coupled to one of the
scintillator arrays. A second photodiode array is optically coupled
to the other scintillator array. The photodiode arrays are the
optical detectors.
[0026] FIG. 3(g) is a perspective view of one detector block from
the detector head illustrated in FIG. 2 for another embodiment of
the design incorporating discrete scintillators composed of two
stacked different scintillator materials of different decay times
optically coupled to a non-active segmented non-inverted light
guide. The light guide is optically coupled to PMTs. The PMTs are
the optical detectors.
[0027] FIG. 3(h) is a perspective view of one detector block from
the detector head illustrated in FIG. 2 for another embodiment of
the design incorporating discrete scintillators composed of two
stacked different scintillator materials of different decay times
optically coupled to a non-active segmented non-inverted light
guide. The light guide is optically coupled to the PMTs in addition
a photodiode array is optically coupled to one of the scintillator
arrays. The PMTs and the photodiode arrays are the optical
detectors.
[0028] FIG. 3(i) is a perspective view of one detector block from
the detector head illustrated in FIG. 2 for another embodiment of
the design incorporating discrete scintillators composed of two
stacked different scintillator materials of different decay times
optically coupled to a non-active segmented non-inverted light
guide. The light guide is optically coupled to a photodiode array.
The photodiode array is the optical detector.
[0029] FIG. 3(j) is a perspective view of one detector block from
the detector head illustrated in FIG. 2 for another embodiment of
the design incorporating discrete scintillators composed of two
stacked different scintillator materials of different decay times
optically coupled to a non-active segmented non-inverted light
guide. The light guide is optically coupled to a photodiode array.
A second photodiode array is optically coupled to one of the
scintillator arrays. The photodiode arrays are the optical
detectors.
[0030] FIG. 3(k) is a perspective view of an alternative detector
block applicable to any of the embodiments discussed herein showing
that the light depth and configuration of the segmentation is
variable depending upon the thickness of the light guide.
[0031] FIG. 3(l) is a perspective view of an alternative detector
block applicable to any of the embodiments discussed herein showing
that the perimetric dimensions of a single detector block, which in
one embodiment is defined by a 12.times.12 discrete element array
can be coextensive with the perimetric dimension of a 2.times.2
array of four optical detectors. The segmented light guide can be
inverted or non-inverted, active or non-active.
[0032] FIG. 4 is inclusive of FIGS. 4, 4a and 4b. FIG. 4 and 4a
illustrate a side elevation and plan, respectively, of one detector
block from the detector head illustrated in FIG. 2 for another
embodiment of the design incorporating discrete scintillators
composed of two stacked different scintillator materials of
different decay times optically coupled to a non-active
non-segmented planar light guide. The planar light guide is
optically coupled to the PMTs. The PMTs are the optical detectors.
FIG. 4b illustrates an alternate embodiment in which the discrete
scintillators are disposed between the light guide and the optical
detectors.
[0033] FIG. 5 is a perspective view of one detector block from the
detector head illustrated in FIG. 2 for another embodiment of the
design incorporating discrete scintillators composed of a single
scintillator material optically coupled to a non-active segmented
light guide. The light guide may be inverted or non-inverted. The
light guide is optically coupled to the PMTs. The PMTs are the
optical detectors.
[0034] FIG. 6 is inclusive of FIGS. 6, 6a and 6b. FIGS. 6 and 6a
illustrate side elevation and plan views, respectively, of one
detector block from the detector head illustrated in FIG. 2 for
another embodiment of the design incorporating discrete
scintillators composed of a single scintillator material optically
coupled to a non-active non-segmented planar light guide. The
planar light guide is optically coupled to the PMTs. The PMTs are
the optical detectors. FIG. 6b illustrates an alternate embodiment
in which the discrete scintillators are disposed between the light
guide and the optical detectors.
[0035] FIG. 7 is a perspective view of one detector block from the
detector head illustrated in FIG. 2 for another embodiment of the
design incorporating discrete scintillators composed of a single
scintillator material of two different decays times optically
coupled to a non-active segmented inverted light guide. The light
guide is optically coupled to the PMTs. The PMTs are the optical
detectors.
[0036] FIG. 8 is a perspective view of one detector block from the
detector head illustrated in FIG. 2 for another embodiment of the
design incorporating discrete scintillators composed of a single
scintillator material of two different decay times optically
coupled to a non-active non-segmented planar light guide. The light
guide is optically coupled to the PMTs. The PMTs are the optical
detectors.
[0037] FIG. 9(a) depicts the uniform irradiation by a Na-22
radioactive source at 511 keV of a single layer of NaI(Tl)
segmented into a 12.times.12 array yielding 144 discrete elements.
The detector design is exhibited in FIG. 5. The position histogram
displayed in FIG. 9(a) is of the overlap of twelve rows each
containing 12 discrete scintillator elements. The figure displays
excellent peak-to-valleys indicating very good separation among the
144 discrete elements.
[0038] FIG. 9(b) is the same as FIG. 9(a) except Co-57 is used as
the radioactive source with a gamma-ray energy of 122 keV.
[0039] FIG. 10 is the same as FIG. 9(a) except Cd-109 is used as
the radioactive source with a gamma-ray energy of 88 keV.
[0040] FIG. 11 is the same as FIG. 9(b) except the peak-to-valleys
are not displayed. The position histogram displays all 144 discrete
NaI(Tl) scintillator elements for the detector design as exhibited
in FIG. 5 when uniformly irradiated with 122 keV gamma-rays from
Co-57.
[0041] FIG. 12 is the position histogram displaying all 144
discrete NaI(Tl) scintillator elements for the detector design as
exhibited in FIG. 3(a). In this case one scintillator layer (slow)
is composed of NaI(Tl) and the other scintillator layer (fast) is
composed of LSO. The detector is uniformly irradiated with 140 keV
gamma-rays from Tc-99m.
[0042] FIG. 13 is the Cross-over time spectra and position
histograms for the detector design as exhibited in FIG. 3(a). In
this case one scintillator layer (slow) is composed of NaI(Tl) and
the other scintillator layer (fast) is composed of LSO. The
detector is uniformly irradiated by 511 keV gamma-rays from Ge-68.
Pulse Shape Discrimination as provided in FIG. 18 is used to
determine in which of the two scintillator layers the gamma-rays
interacted. The cross-over time spectra is shown in FIG. 13 for the
two scintillator layers. The position histograms are labeled `slow`
for the NaI(Tl) layer and `fast` for the LSO layer. The position
histograms exhibit excellent separation of the 144 discrete
elements for each scintillator layer. FIG. 13 illustrates the
effectiveness of the detector design and the pulse shape separation
technique.
[0043] FIG. 14 is a pulse height energy spectrum for an LSO crystal
irradiated by 662 keV gamma-rays from Cs-137. Also exhibited in the
spectrum is the 2.6% abundant Lu-176 background of LSO.
[0044] FIG. 15 exhibits the energy integration for NaI(Tl) and LSO.
The pulse shape discrimination technique utilized with the detector
designs of this disclosure, involves integrating the detector
charge at two points in time then ratio the values. For LSO and
NaI(Tl) separation the first sample is taken at 80 ns from the
start of the integration and the second sample at 256 ns at the end
of signal integration to provide a value to normalize out event
charge (energy).
[0045] FIG. 16 In the two-sample shape discrimination, the
integrated energy signal is sampled using an 8-bit flash converter
at approximately 80 ns (E1) and again at 256 ns (E2). The energies
(E1) and (E2) are used to determine if the event occurred in a
NaI(Tl) or an LSO crystal. The ratios of the energies (E1) and (E2)
are used for the shape discrimination. A 65536.times.1 static RAM
can be used to indicate the crystal type as shown in FIG. 16.
[0046] FIG. 17 The integrated X and Y values are digitized at 256
ns using flash converters with the integrated energy signal as the
reference to produce the (A+B)/Sum and (A+C)/Sum ratios. The X and
Y ratios are used to determine the crystal in which the event
occurred. A 65536.times.8 static RAM can be used to indicate the
crystal as shown in FIG. 17.
[0047] FIG. 18 displays the block diagram of the setup used to
evaluate the detectors.
[0048] FIG. 19 displays the block diagram for the coincidence
setup.
[0049] FIG. 20 displays the geometric arrangement used in
coincidence timing measurements.
[0050] FIG. 21 displays the coincidence time spectra for the
various scintillator layer combinations.
[0051] FIG. 22 displays the contour plots for coincidence time
centroids in nsec for the various scintillator layer combinations.
The zero-line is based on the mean LSO-LSO centroid position. The
time centroids are symmetric, but the centroid position is
dependent upon the location of the discrete scintillator element
with respect to the PMT and reflects the spatial uniformity in the
anode output. The time centroid shifts can be corrected via a
lookup table.
[0052] FIG. 23 similarly displays the surface plots for coincidence
time centroids in nsec for the various scintillator layer
combinations.
[0053] FIG. 24 displays the geometric arrangement used in the line
spread function measurements.
[0054] FIG. 25 displays the line spread functions for the various
scintillator layer combinations.
[0055] FIG. 26 displays the general system architecture of the
medical imaging system.
BEST MODE FOR CARRYING OUT THE INVENTION
[0056] The present invention is a scintillation detector array for
encoding energy, position and time coordinates of gamma ray
interactions for use in Single Photon Emission Computed Tomography,
("SPECT"), with or without coincidence photon imaging, Planar
Imaging, and Positron Emission Tomography, ("PET"), imaging.
Referring now to the drawings, FIG. 1 depicts a perspective view of
a medical imaging scanner 12 embodying scintillation detector
arrays 10 for encoding the energy, position and time coordinates of
gamma-ray interactions. The detector head assembly 15 comprises an
(n).times.(m) planar array of detector blocks 20, optically coupled
to an (q).times.(p) array of photomultiplier tubes (PMTs) 25 in one
embodiment of the design; or to an (y).times.(z) array of optical
detectors such as Avalanche Photodiodes (APDs) or PIN Photodiodes
25' in another embodiment of the design. Note the variables
(n),(m),(q),(p),(y),(z), may or may not equal each other. FIG. 2a
depicts an 8.times.10 planar array of detector blocks 20 optically
coupled to a 9.times.11 planar array of photomultiplier tubes 25
configured so that the center of each PMT 25 resides over the
corner of each detector block 20. While this is the preferred
arrangement, where the variables (n),(m),(q),(p),(y),(z) are equal
the scintillator elements and the optical detectors can be
positioned so as to be co-linear. As seen in FIGS. 2b and 2c, the
PMTs 25 are partitioned into 99 square compartments using 0.25 mm
thick magnetic shielding arranged in a cross pattern to form a grid
that locates the PMTs 25 and provides structural support for the
1.0 mm glass window 30 that separates the PMTs 25 from the detector
blocks 20. An hermetic enclosure is provided by a thin stainless
steel foil 35 membrane mounted to the subject side of the
enclosure. As seen in FIGS. 2d and 2e, a similar arrangement is
utilized when the optical detector is a photodiode array.
[0057] FIG. 3(a) is a perspective view of a detector block 20 from
the detector head 15 of FIG. 2 for one preferred embodiment of the
design incorporating discrete scintillators 40 composed of two
stacked different scintillator materials, i.e. a slow scintillator
material 140 and a fast scintillator material 240, respectively, of
different decay times optically coupled to a segmented inverted
light guide. Those skilled in the art will appreciate that the
light guides discussed herein can be either "active" or
"non-active". The term "non-active" is used when the light guide is
composed of a non-scintillating material, whereas an active light
guide is composed of scintillating material. The use of an active
light guide has the inherent characteristic of mispositioning of
events which occur as a result of interactions in the non-segmented
portion of the light guide. Provided the active light guide is
properly designed the magnitude of mispositioned events can be
minimized, but never completely eliminated. In the preferred
embodiment, light guide 50 is non-active.
[0058] Moreover, the term "segmented", as used herein, describes a
plurality of barriers defining a preselected number of slots, the
number of which and the depth of which are varied to control the
variable statistical distributions of photons, whereas
non-segmented light guides are continuous such as those used in
conventional gamma cameras. And, as used herein, the term
"inverted" is used to describe a light guide in which the slotted
section of the light guide is optically coupled to the optical
detectors, whereas for a traditional (non-inverted) light guide the
section of the light guide that contains the non-slotted continuous
region is optically coupled to the optical detectors.
[0059] The selection of type and orientation of the light guide is
in response to various manufacturing constraints. For example when
the scintillator and light guide are composed of different
materials they may have to be processed separately each using a
unique set of tooling and chemical processing. Whereas when the
scintillator and light guide are composed of the same material then
the tooling and chemical processing are generally the same and no
bonding agents are required to optically bond the scintillator to
the light guide, since under this circumstance the light guide is
cut into the scintillator. However, the detector designer may
choose to put an optical bond between the scintillator and the
light guide even though they are composed of the same material for
the purpose of depth of interaction encoding by either pulse shape
or pulse height discrimination. One reason why the detector
designer may not want to use the scintillator as an active light
guide is due to the mispositioning of events which occur as a
result of interactions in the non-segmented portion of the light
guide.
[0060] For an (n).times.(m) array of discrete scintillators 40
optically coupled to a segmented light guide, such as light guide
50, the cut depths for an inverted and non-inverted light guide are
unique and are not interchangeable. A detector design incorporating
discrete scintillators 40 and a non-inverted light guide, such as
non-inverted light guide 150 in the Figures, cannot be converted to
a functional inverted light guide detector of equivalent
performance simply by flipping the light guide.
[0061] For the preferred embodiment of the design, an inverted
segmented light guide 50, which is non-active, is used as depicted
in FIG.3(a). The composition of the inverted segmented light guide
50 can be of any material that is chemically compatible with the
scintillator material and is optically transmissive to the
wavelength of emission of the scintillator. Other material
properties constraints the detector designer must consider in
selection of the light guide material is the index of refraction,
thermal characteristics, mechanical characteristics and cost. As
stated above, in the preferred embodiment, the inverted segmented
light guide 50 was designed to be non-active so that mispositioning
of events due to gamma-ray interactions in the light guide would
not occur. The light guide was designed to be inverted so that
registrational tolerances could be eased with respect to the
correspondence of the discrete scintillator elements 40 relative to
the partitioned section of the light guide. The reflector may be of
any material that has high reflectance for the emission wavelength
of the scintillator(s). In the case of NaI(Tl) and LSO, 1.53 micron
Silicon dioxide (SiO.sub.2) was selected as the reflector. Other
particle diameters could be used with the general trend as the
particle diameter increases the optical cross-talk among the
discrete scintillator elements as well as among the partitions of
the light guide increases thereby degrading the signal-to-noise
ratio.
[0062] Two different scintillators 140 and 240 having different
decay times are used in one of the preferred embodiments of the
design as illustrated in FIG. 3(a). NaI(Tl) was selected as the
scintillator of choice for the SPECT measurements and is the
current industry standard. The NaI(Tl) block size is 52 mm.times.52
mm.times.10 mm thickness. Block sizes of different cross sections
or thickness can be selected by the detector designer; the
cross-section will be set by the dimensions of the optical
detector(s) and the thickness based on the level of compromise
between efficiency gain and resolution degradation. Prior to sawing
the NaI(Tl) into segments it must be bonded to a substrate for
mechanical integrity during processing and to preserve the discrete
elements positional registration. Selection of the bonding agent
and substrate requires consideration of mechanical, chemical, and
optical properties. The bonding agent and substrate must be
optically transmissive to the emission wavelength of the NaI(Tl)
scintillator, they must be chemically compatible and provide
mechanical strength without thermal expansion detriment. An optical
glass slide of 0.5 mm thickness was selected as the bonding
substrate. Other glass slide thicknesses could be selected based on
the level of compromise among mechanical strength, optical
cross-talk, and optical attenuation. The bonding agent may be
either epoxy or RTV. Selection of the `sawing` method for
segmentation must consider thermal and mechanical stresses as well
as chemical compatibility. NaI(Tl) processing should be performed
in a dry room to prevent hydration of the scintillator. The cutting
lubricant temperature must be controlled to prevent thermal
fracture of the scintillator; further it must be chemically
compatible so that the mechanical dimensions and optical properties
of the scintillator do not change with time. Regardless of which
cutting method or lubricant is selected all residues from the
process must be removed from the scintillator surface otherwise the
total light emitted will be reduced. A low viscosity oil similar in
viscosity to water was selected as the cutting lubricant. Residues
from the cutting process were subsequently removed either
chemically or mechanically.
[0063] The 52 mm.times.52 mm.times.10 mm NaI(Tl) block was bonded
to a 0.5 mm thick glass substrate and sawed into a 12.times.12
discrete element array of 4.37 mm pitch and 4.0 mm.times.4.0
mm.times.10.0 mm crystal size. Silicon dioxide powder of 1.53
micron particle size is used as the reflector in the interstices 55
of the 144 elements of the array and is also used as the reflector
between the blocks and on the scintillator non-interstitial
surfaces. The silicon dioxide reflector in the interstices 55 is
`sealed` with a small layer of Teflon powder. This is to prevent
the permeation of bonding agents into the silicon dioxide reflector
from subsequent bonding processes.
[0064] Referring to FIG. 3(a). LSO was selected as the other
scintillator of the pair due to its high luminosity, high density,
effective-Z and its fast decay time (see Table 1).
1TABLE 1 Properties of Proposed PET Scintillators NaI(Tl) BGO LSO
LOP GSO CeF.sub.3 Density (gm/cm.sup.3) 3.67 7.13 7.4 6.53 6.71
6.16 Effective Z 50.6 74.2 65.5 62.5 58.6 52.7 Mean Free Path (cm)*
2.93 1.05 1.16 1.37 1.43 1.71 Index of Refraction 1.85 2.15 1.82
1.7 1.91 1.68 Hydroscopic? YES NO NO NO NO NO Rugged? NO YES YES
YES NO Decay Time (ns) 230 300 40 24 60 27 Emission Peak (nm) 410
480 420 360 440 340 LightOutput[NaI(Tl) = 100] 100 15 75 32 25 4-5
Energy Resolution* 7.8 10.1 10.1 8.9 20 Photoelectric Fraction*
.175 .411 .324 .288 .247 .188 Incoherent Fraction* .790 .543 .629
.670 .712 .778 Mass Attenuation (cm.sup.2/gm)* .0930 .1332 .1170
.1118 .1040 .0951 Mass Energy Absorp (cm.sup.2/gm)* .0409 .0731
.0601 .0553 .0495 .0424 Linear Attenuation (cm.sup.-1)* .3411 .9496
.8658 .7302 .6978 .5858 Linear Energy Absorp (cm-1)* .1501 .5214
.4447 .3612 .3323 .2614 BaF.sub.2 PbSO.sub.4 PbCO.sub.3 LAP.sup.(1)
LuAG.sup.(1) YSO.sup.(4) Density (gm/cm.sup.3) 4.89 6.2 6.6 8.34
6.9 4.543 Effective Z 52.2 73.1 75.9 63.9 61.7 34.2 Mean Free Path
(cm)* 2.20 1.22 1.10 1.05 1.31 2.58 Index of Refraction 1.56 1.88
1.8 1.8 Hydroscopic? Slight NO NO NO NO NO Rugged? YES YES YES
Decay Time (ns) 0.6 136 8.5 11.sup.(2) 58.sup.(3) 70 Emission Peak
(nm) 220 350 475 390 500 420 LightOutput[NaI(Tl) = 100] 5 9 1.4 17
25 118 Energy Resolution* 11.4 40 42 14.9 8.0 Photoelectric
Fraction* .186 .399 .432 .305 .277 .051 Incoherent Fraction* .779
.556 .520 .651 .683 .927 Mass Attenuation (cm.sup.2/gm)* .0929
.1320 .1382 .1141 .1105 .0853 Mass Energy Absorp (cm.sup.2/gm)*
.0414 .0718 .0775 .0574 .0540 .0315 Linear Attenuation (cm.sup.-1)*
.4545 .8181 .9122 .9515 .7623 .3875 Linear Energy Absorp (cm-1)*
.2024 .4449 .5113 .4790 .3727 .1431 *at 511 Kev .sup.(1)Computed
10-Feb-1995 M. Andreaco, L. Dyers 13-Nov-1992 V2 .sup.(2)60% @ 11
nsec, 26% @ 28 nsec, 13% @ 835 ns .sup.(3)13% @ 58 nsec, 21% @ 310
nsec, 65% @ 2090 nsec .sup.(4)Computed 5-Jan-1996
[0065] LSO is a rugged scintillator and its processing methods are
unique and different from NaI(Tl). Preferably, the 52 mm.times.52
mm.times.10 mm LSO block is either mechanically polished or etched
to transparency prior to bonding or sawing. The reason the
mechanical and/or etch process is used is to maximize the
transmission of the other scintillator's light through the LSO and
the light guide 50. This mechanical polishing or etching to
transparency condition also applies to the light guide. Those
skilled in the art will recognize that while polishing or etching
to optical transparency is preferred, a polish or etch that results
in less than optical transparency may alleviate certain
manufacturing problems associated with etching to optical
transparency. However, polishing or etching to less than optical
transparency reduces the efficiency of the detector and/or light
guide. Pyrophosphoric acid (H.sub.4P.sub.2O.sub.7) is used as the
chemical etchant for LSO. Etching LSO to transparency is a function
of time and acid temperature. Typically a temperature of
300.degree. C. with approximately 15 minutes duration is required
to etch LSO to transparency, other temperatures and durations may
also be used to the same effect. One problem associated with this
etching temperature is thermal stress which can fracture the
scintillator. Thermal stress can be minimized by having the
scintillator and the acid at the same temperature during the etch
process. Upon removal of the LSO from the etch bath it is allowed
to air cool to 100.degree. C. whereupon the LSO is submerged in
boiling water to rinse the residual pyrophosphoric acid from the
LSO surface. Upon removal of the LSO from the boiling water rinse
bath it is allowed to air dry to room temperature. The LSO is then
submerged in a 37% Hydrochloric acid (HCl) bath for approximately 2
minutes to remove residual pyrophosphoric acid from the LSO which
was not removed by the boiling water rinse. The duration of the HCl
rinse is not critical as it will not etch the LSO, it will only
remove surface contaminants. Following the HCl etch a clean water
rinse is used to remove residual HCl from the LSO surface.
[0066] The mechanically polished and/or etched uncut LSO block is
optically bonded to a mechanically polished optical grade glass or
ultraviolet transmissive plastic whose dimensions are 52
mm.times.52 mm.times.12.7 mm thickness. Other thicknesses can be
used depending on the desired statistical distribution of photons
and optical transmission and optical attenuation compromise. The
light guide 50 can be cut or uncut prior to bonding to the LSO, the
choice really depends upon the fabrication processes selected. The
end result should yield the desired discrete element registration
to light guide partition.
[0067] The 52 mm.times.52 mm.times.10 mm LSO block is cut into a
12.times.12 element array of 4.37 mm pitch and 4.0 mm.times.4.0
mm.times.10 mm crystal size. The cut LSO crystal array must be
etched to remove surface contaminants due to the sawing process.
This etch is generally not designed to be optically transparent
since the etch at that temperature would be detrimental to the
optical bond. Cleaning the crystal surface can be accomplished with
an HCl etch as stated above. However enhancing the LSO light output
after the sawing process requires an HCl etch to remove the surface
contaminants from the sawing process followed by a pyrophosphoric
acid etch at a nominal temperature of 170.degree. C. for a duration
of approximately 20-30 seconds. Etching for shorter or longer
duration at this temperature will not provide the optimum light
yield. Other pyrophosphoric acid temperatures and etch times may be
used. The choice depends upon the manufacturing process selected
and the amount of time allocated to the etch process and the
desired light output. The pyrophosphoric acid etch removes the
sawing process induced micro-fracturing of the LSO crystal surface
which would otherwise trap the LSO luminescence light. The
pyrophosphoric acid etch is followed by a 100.degree. C. water
rinse then a second HCI rinse followed by a room temperature water
rinse.
[0068] The interstices 55 of the cut LSO and light guide arrays are
preferably filled with silicon dioxide powder of 1.53 micron
particle size to serve as reflector. It will be recognized that
other reflector materials, such as titanium dioxide, aluminum
oxide, magnesium oxide, barium sulfate, zinc oxide and Teflon
powder, can also be utilized as reflectors. A small layer of Teflon
powder is then used to `seal` the Silicon dioxide reflector in the
interstices of the arrays. This is to prevent permeation of bonding
agents into the reflector from subsequent bonding processes. The
LSO side of the LSO/light guide array is then optically bonded to
the glass substrate of the NaI(Tl) array.
[0069] Two prototype detectors were fabricated using the methods
described above. Testing of the prototype detectors required that
the NaI(Tl) array be hermetically sealed. For the one prototype an
aluminum housing was used to hermetically seal the NaI(Tl) array
only. Glass was used as the hermetic seal for the second prototype,
however, in this case the NaI(Tl), LSO and light guide were all
enclosed.
[0070] Referring to FIGS. 3a-3l, various configurations of detector
arrays are illustrated. It should be understood that the figures
are not drawn to scale. These figures present a perspective view of
the types of detector blocks which could be incorporated in the
detector head 15 illustrated in FIG. 2. In the various embodiments
illustrated herein, the scintillator can either be a single layered
scintillator or can be composed of two stacks of scintillator
material of different decay times either using the same
scintillator material in each layer or different scintillator
materials. Selection of the scintillator material for each
scintillation layer is application dependent. Table 1 herein
provides examples of various scintillator materials, but is not all
inclusive. Other scintillator materials having similar properties
or other application dependent properties could also be
substituted. The embodiments can be further modified by varying the
types of optical detectors which are used, i.e. photomultipliers
and/or photodiodes. An additional variant is whether or not a
segmented optical light guide is used. If a segmented optical light
guide is used then the variant is whether the configuration is
inverted or non-inverted.
[0071] In this regard, in FIG. 3(a) one detector block 20 from the
detector head 15 illustrated in FIG. 2 incorporates discrete
scintillators 40 composed of two stacked different scintillator
materials 140 and 240 of different decay times optically coupled to
a segmented inverted light guide 50, which is preferably
non-active, which in turn is optically coupled to photomultiplier
tubes (PMTs) 25. The PMTs 25 are the optical detectors. This
embodiment can be further varied, as shown in FIG. 3(b)with the
addition a photodiode array 25' which is optically coupled to one
of the scintillator arrays 20. The PMTs 25 and photodiode arrays
25' are the optical detectors.
[0072] Another embodiment is illustrated in FIG. 3(c). This
embodiment incorporates discrete scintillators 40 composed of two
stacked different scintillator materials 140 and 240 of different
decay times optically coupled to a segmented inverted light guide
50, which is preferably non-active. The light guide 50 is optically
coupled to a photodiode array 25' which serves as the optical
detector. This embodiment can be further modified, as shown in FIG.
3(d) which includes a second photodiode array 25' optically coupled
to one of the scintillator arrays. The two photodiode arrays are
the optical detectors. As illustrated in FIG. 3(e), the embodiment
illustrated in FIG. 3(c) can be further modified by integrating the
light guide function in the discrete scintillators 40. In FIG.
3(f), the embodiment illustrated in FIG. 3(e) has been further
modified by optically coupling a second photodiode array 25' to
scintillator material 140.
[0073] FIGS. 3(g.sub.1) and (g.sub.2) yet another embodiment is
illustrated. This embodiment incorporates discrete scintillators 40
composed of two stacked different scintillator materials 140 and
240 of different decay times optically coupled to a segmented
non-inverted light guide 150, which is preferably non-active. The
light guide 150 is optically coupled to PMTs 25.
[0074] FIG. 3(h) illustrates yet another embodiment of the design
which incorporates discrete scintillators 40 composed of two
stacked different scintillator materials 140 and 240 of different
decay times optically coupled to a segmented non-inverted light
guide 150, which is preferably non-active. The light guide 150 is
optically coupled to the PMTs 25 and in addition a photodiode array
25' is optically coupled to one of the scintillator arrays 140.
[0075] FIG. 3(i) illustrates still another embodiment of the design
incorporating discrete scintillators 40 composed of two stacked
different scintillator materials 140 and 240 of different decay
times optically coupled to a segmented non-inverted light guide 150
which is preferably non-active. The light guide 150 is optically
coupled to a photodiode array 25'. This embodiment can be further
modified, by optically bonding layer 140 to a second photodiode
array 25'.
[0076] FIG. 3(k) is a perspective view of an alternative segmented
light guide 150' from the detector block illustrated in FIG. 3(j)
showing that the light depth and configuration of the segmentation
is variable depending upon the thickness of the light guide. FIG. 4
illustrates an embodiment of the design incorporating discrete
scintillators 40 composed of two stacked different scintillator
materials 140 and 240 of different decay times optically coupled to
a non-segmented light guide 250, which is preferably non-active.
The light guide 250 is optically coupled to the PMTs 25. With
respect to each of these embodiments, and the other embodiments
described herein, selection of the scintillator material for each
scintillation layer is application dependent. Table 1 herein
provides examples of various scintillator materials, but is not all
inclusive. Other scintillator materials having similar properties
or other application dependent properties could also be
substituted.
[0077] As illustrated in FIGS. 4, 6 and 8, in various embodiments,
the planar light guide 250 is continuous, i.e. non-segmented. As
illustrated, the continuous, planar light guide 250 has a first
planar surface that extends beyond the interfacing surface of the
discrete scintillator elements and has a second planar surface that
extends beyond the interfacing surface of the optical detectors. An
alternate embodiment is illustrated in FIGS. 4b and 6b. In this
alternate embodiment, the scintillators, which can either be a
single layer 40 of material as illustrated in FIG. 6b, or two
layers 240 and 140 as illustrated in FIG. 4b, are disposed between
the optical detectors 25 and a thin layer continuous, planar light
guide 250'. In this regard, the planar light guide 250', which can
either be active or non-active, is disposed on the patient side of
the scintillators. For use in PET applications, the following
High-Z scintillator materials are preferred: Cerium-doped Lutetium
Oxyorthosilicate, ("LSO"), Cerium-doped Lutetium Yttrium
Oxyorthosilicate, ("LYSO"), Cerium-doped Lutetium Gadolinium
Oxyorthosilicate, ("LGSO"), Cerium-doped Gadolinium
Oxyorthosilicate, ("GSO"), Cerium-doped Lutetium Aluminum
Perovskite, ("LuAP"), and Cerium-doped Yttrium Aluminum Perovskite,
("YAP"). Also, Cerium-doped Yttrium Oxyorthosilicate, ("YSO"), can
be selected as a scintillator material.
[0078] Referring to FIG. 5, the NaI(Tl) array 20' was optically
bonded to the segmented inverted light guide 50, which is
preferably non-active, for evaluation of the detector design.
Testing of the NaI(Tl) array 20' was conducted using standard
nuclear spectroscopy instrumentation. The array was evaluated for
energy and `position` space at 511 keV, 122 keV and 88 keV, using
point sources. The `position` histograms 60, 62, and 64 for these
three energies are provided in FIGS. 9(a), 9(b) and 10,
respectively, with Spectra Identification number 27-18-1.spm. The
position histogram as depicted exhibits the overlap of all twelve
rows. The twelve position peaks represent all 144 discrete
scintillator elements. Any mispositioning of the 144 peaks would
result in a degradation of the peak-to-valley. No such degradation
is exhibited. The pulse height energy resolution measured at 88
keV, 122 keV, and 511 keV are summarized as follows:
2 Mean Pulse Height Energy Resolution per Energy crystal element 88
keV 9.42% 122 keV 9.45% 511 keV 7.57%
[0079] FIG. 11 provides position information without the complexity
of interpreting twelve overlapped rows. FIG. 11 shows the two
dimensional position histogram 66 of the detector design exhibited
in FIG. 5 where the scintillator layer is composed of NaI(Tl) layer
segmented into the 12.times.12 array. The light guide 50 is
non-active, segmented and inverted. The detector is uniformly
irradiated with 122 keV gamma-rays from Co-57. This data indicates
that the NaI(Tl) array performance is comparable to existing high
resolution SPECT scanners while providing outstanding PET energy,
position and time resolution.
[0080] Evaluation of the bonded NaI(Tl)/LSO/light guide arrays (see
histograms 68 and 70 in FIGS. 12 and 13) requires consideration of
the naturally occurring 2.6% abundant Lu-176 element. The Lu- 176
is radioactive, producing energetic particles that interact in the
scintillator to produce a background count rate of approximately 39
counts per second per gram of LSO scintillator. Due to the high
stopping power of LSO very few of the particles originating from
the decay of Lu-176 actually escape from the LSO scintillator.
Lu-176 background events do not present a problem in PET studies
due to their uncorrelated nature, the system labels the events as
randoms and are thus rejected. However, the Lu-176 background
events which fall into the SPECT energy window are a problem for
they are counted as singles and are of the same order of magnitude
as the signal rate in many Planar and SPECT studies. FIG. 14
exhibits two energy spectra, one for an LSO crystal irradiated by
Cs-137 72, the other spectra is for the Lu-176 background events
74. These spectra indicate standard energy discrimination
techniques will not successfully reject Lu-176 background
events.
[0081] NaI(Tl) and LSO have scintillation decay times of 230 and 40
nanosecs (ns) respectively. For SPECT studies, low energy photons
are stopped in the NaI(Tl) crystal(s), producing scintillation
events of 230 ns decay time; whereas the Lu-1 76 background events
of LSO are produced with 40 ns decay. The electronics circuitry
distinguishes between NaI(Tl) and LSO events based on the decay
time signatures of the two scintillators. The technique is known as
pulse shape discrimination PSD. Approximately 99% of the Lu-176
background events must be rejected to prevent significant noise
counts from occurring in the low count rate SPECT data.
[0082] The PSD technique utilized in the PET/SPECT detector
involves integrating the detector charge at two points in time and
then comparing the ratio of the values. For LSO and NaI(Tl)
separation, the first sample is taken at 80 ns from the start of
the integration and the second sample at 256 ns. The first sample
is selected at 80 ns since this is where the maximum difference
occurs for the LSO and NaI(Tl) scintillators. The LSO signal will
be maximally above the NaI(Tl) signal at 80 ns for a given signal
charge. The second sample at 256 ns is at the end of the signal
integration to provide a value to normalize out event charge
(energy). The two-sample shape discrimination circuit used with the
PET/SPECT ASIC does not require additional shape discrimination
circuitry inside the ASIC. The external logic sequencing circuit
requests an intermediate and final energy sample from the external
energy ADC. These two values can then be effectively ratioed and
discriminated through a look-up logic memory device to determine if
the event came from the NaI(Tl) or LSO scintillator. The integrated
signals for the NaI(Tl) and LSO is shown in FIG. 15. The deviation
from a pure exponential behavior for LSO is caused by the filtering
circuitry necessary to cancel the light decay of NaI(Tl) for
improved count rate performance.
[0083] In order to utilize the two-sample shape discrimination, it
is necessary to ensure that the energy integrator is sufficiently
linear at the sample points for the NaI(Tl) and LSO scintillation
detector signals. The integrated energy signal is sampled using an
8-bit flash converter at approximately 80 ns (E1) and again at 256
ns (E2). The energies (E1) and (E2) are used to determine if the
event occurred in a NaI(Tl) or an LSO crystal. The ratios of the
energies (E1) and (E2) are used for the shape discrimination. A
65536.times.1 static RAM 75 can be used to indicate the crystal
type as shown in FIG. 16. The integrated X and Y values are
digitized at 256 ns using flash converters with the integrated
energy signal as the reference to produce the (A+B)/Sum and
(A+C)/Sum ratios. The X and Y ratios are used to determine the
crystal in which the event occurred. A 65536.times.8 static RAM 77
can be used to indicate the crystal as shown in FIG. 17.
[0084] FIG. 18 exhibits the block diagram of the method used to
evaluate the prototype detectors for energy, position and time (see
FIGS. 12 and 13). The module 79 labeled "CTI Design Box"
incorporates a preamp, timing filter amp, sum amp, CFD, and digital
clock. The rest of the Figure illustrates standard nuclear
spectroscopy instrumentation--such as is available from EG&G
Ortec (e.g. 1995 EG&G Ortec catalog "Modular Pulse-Processing
Electronics and Semiconductor Radiation Detectors")--processing for
performing pulse shape discrimination. For example, those skilled
in the art will appreciate that module 81 is a standard personal
computer with analog to digital converters, module 83 is for pulse
shaping and includes a Constant Fraction Discriminator and that
module 85 is a Time-to-Amplitude Converter. Analysis of the two
prototype blocks in terms of energy and crystal identification
demonstrates that when the devices are operated in the SPECT mode
all the individual NaI(Tl) discrete element detectors can be
identified at 140 keV. Further, the average pulse height energy
resolution at 140 keV is 10.3% for prototype detector block #1 and
10.0% for block #2. The median pulse height energy resolution is
below 10.0% in both cases. This performance is comparable to
current SPECT imaging systems. When operated in the PET mode, all
the individual NaI(Tl) and LSO discrete element detectors are
easily identified. The average pulse height energy resolution
measured at 511 keV is below 10.1% for both scintillator layers
(less than 8.0% for the NaI(Tl) layer). This indicates excellent
PET performance comparable to existing PET imaging systems.
[0085] FIG. 19 exhibits the block diagram used in the determination
of coincidence performance. Similarly FIG. 19 illustrates standard
nuclear spectroscopy instrumentation. For example module 87 is a
fast shaping amplifier along with a constant fraction
discriminator, and module 93 is a time-to-amplitude converter with
a single channel analyzer. The output at module 97 is illustrated
in FIG. 21. FIG. 20 exhibits the geometric configuration used in
measuring the coincidence timing. The figure illustrates 2.times.2
center crystals for each scintillator layer of the one crystal
block is used in coincidence with the 144 crystal elements of each
layer of the opposing block. Time spectra are then measured for
each of the 144 crystal elements of each scintillator layer. FIG.
21 displays the coincidence time spectra for each layer
combination. The time centroids are symmetric, but the centroid
position is dependent upon the location of the discrete crystal
element with respect to the PMT, and reflects the spatial
uniformity in the anode output (see FIGS. 22 & 23). The time
centroid shifts can be corrected via a lookup table. Time
resolution is scintillator layer combination dependent. The LSO-LSO
combination exhibits 1.6 nsec time resolution, indicating PET
coincidence timing with 6 nsec time window is feasible.
[0086] The line spread function (LSF) method was used to assess the
spatial resolution of the prototype detector blocks. FIG. 24
exhibits the geometric arrangement used in the LSF measurement.
FIG. 25 exhibits the step wise data of the LSF for the layer
combinations, and Table 2 provides the LSF data in terms of full
width at half maximum (FWHM) of the LSF.
3TABLE 2 Line spread functions for the layer combinations of the
detector design exhibited in FIG. 3(a) where scintillator layer (1)
is composed of Nal [TI] and scintillator layer (2) is composed of
LSO. The detector is irradiated with 511 keV gamma rays from an
F-18 line source. See FIG. (24) for geometric arrangement.
Combination FWHM [mm] StdDev [mm] LSO-LSO (back-back) 3.5 0.2
Nal-LSO (front-back) 3.0 0.4 LSO-Nal (front-back) 3.1 0.3
Nal-Nal(front-front) 2.8 0.5
[0087] The LSF data indicates the depth of interaction (DOI)
effects and that a reconstructed spatial resolution of less than 4
mm is possible.
[0088] FIG.26. displays the general system architecture for
application in PET/SPECT medical imaging.
Absolute Sensitivity of PET/SPECT
[0089] The following calculation was done to predict the absolute
sensitivity of the proposed PET/SPECT medical imaging system.
Absolute sensitivity is defined as the ratio of the events detected
by the system to those emitted from a line source placed at the
center of the tomograph. The length of the source is the same as
the axial length of the scanner. If the sensitivity of the detector
to single events is known then the calculation of absolute
sensitivity is the product of the fraction of solid angle and the
square of the singles sensitivity.
.eta..sub.Ab=.eta..sub.Geom.multidot..eta..sub.Det.sup.2
[0090] For the NaI/LSO detector, both a Monte Carlo simulation and
a measurement were done. These are summarized in the table
below:
4 Coincidence Sensitivity of NaI/LSO Detector Calculated Calculated
Measured Absolute Relative Relative LSO/LSO 12.7% 52.7% 50.1%
LSO/NAI 4.8% 19.9% 20.4% NaI/LSO 4.8% 19.9% 20.4% NaI/NaI 1.8% 7.5%
9.0% Total Absolute 24.10%
[0091] Assuming two 50 cm transaxial by 40 cm axial detectors,
directly opposing and 72 cm face to face, the fraction of the solid
angle is found to be 6.2%. Using the total detector sensitivity
from table above, the absolute sensitivity is found to be 1.5%.
[0092] For whole body scanning, the body is longer than the axial
extent of the scanner. Then, if one arbitrarily picks a length for
the source of 70 cm, it is possible to put the absolute sensitivity
of the scanner in perspective by comparing to existing scanners.
The table below compares the PET/SPECT, Siemens ECAT ART, and
Siemens ECAT HR+ for both a line source of axial extent and one of
70 cm.
5 Absolute Sensitivity Comparison of PET/SPECT 70 cm Axial Length
Source Source Measured Calculated Calculated ART 1.04% 1.10% 0.25%
HR+ 2.94% 2.80% 0.62% PET/SPECT 1.50% 0.86%
[0093] Count-Rate Performance of PET/SPECT
[0094] To assess the performance of the PET/SPECT in an imaging
situation, the proposed machine was compared to an existing Siemens
ECAT ART tomograph. Since the count-rate performance of the machine
is a function of the detector singles rate, this needs to be known
along with the trues sensitivity. The randoms rate can be computed
from the singles rates and the coincidence window.
[0095] A typical whole-body study performed on an ART was used as a
bench mark. The study consisted of five bed steps ranging from the
patients nose to just above the bladder. The variation in the
count-rate with respect to bed step was minimal. Due to this, the
rates were averaged to give the following results:
6 ART Whole-body Study 5 mCi Injection Trues 20666 cps Randoms
10166 cps Singles/Block 27716 cps
[0096] The singles sensitivity of the ART blocks and dead-time are
obtained by fitting the singles rates obtained in a count-rate
study done with a 20 cm uniform phantom filled with .sup.18F. To
model the response of the system to the subject, the trues were set
so that for the singles rate of the subject study, the trues also
matched. The randoms were calculated from the singles and the
coincidence window. The dead-time is based on the fitted dead-time
function. This process lets the performance of the system be
modeled with respect to activity. The NEC rate shown in the right
most column did not include a scatter term and was simply
T.sup.2/(T+2*R).
[0097] For PET/SPECT, results were used from the Monte Carlo model
of absolute sensitivity to obtain ratios for the singles and trues
sensitivity. These values are given in the table below:
7 PET/SPECT Sensitivity (cps/uCi/cc) Trues Sensitivity 816000
Singles/Block/Sensitivity 175000
[0098] The dead-time of the system is modeled based on the
electronic dead-time of 320 nS and the following relation: 1 S
Observed = S Incident exp ( - 8 S Incident dead ) ( 1 + S Incidnent
dead )
[0099] This relation takes into account the overlapped detector
structure that forces eight surrounding detectors to be dead during
the processing of an event in one detector.
[0100] With these values, it is possible to predict the performance
of the system to different injection levels as shown in the
following chart:
8 Whole-body Imaging Performance of PET/SPECT PET/SPECT 40 cm fov,
ART 320 ns Deadtime Singles Trues Randoms NEC Singles Trues Randoms
NEC Activity.sup.1 cps Livetime cps cps cps cps Livetime cps cps
cps 0.500 89286 36.3% 41550 87111 8001 37304 18.2% 74157.6 53437
30378 0.475 86210 38.0% 41306 82270 8289 36975 19.8% 76690.0 52499
32371 0.450 83031 39.8% 40967 77300 8582 36548 21.5% 79092.1 51294
34432 0.425 79742 41.6% 40524 72215 8879 36015 23.4% 81320.9 49809
36549 0.400 76338 43.6% 39965 67030 9178 35368 25.5% 83326.4 48035
38704 0.375 72814 45.7% 39279 61762 9477 34598 27.8% 85051.1 45965
40873 0.350 69163 48.0% 38452 56431 9771 33694 30.3% 86429.1 43596
43025 0.325 65379 50.3% 37470 51063 10058 32648 33.0% 87384.9 40930
45119 0.300 61455 52.9% 36317 45685 10330 31447 35.9% 87831.9 37974
47102 0.275 57383 55.5% 34975 40330 10579 30081 39.1% 87671.7 34746
48906 0.250 53155 58.4% 33424 35038 10794 28537 42.5% 86792.2 31271
50443 0.225 48762 61.4% 31642 29852 10961 26802 46.3% 85065.7 27584
51601 0.200 44195 64.6% 29604 24826 11058 24862 50.5% 82347.7 23736
52235 0.175 39445 68.1% 27283 20020 11056 22703 55.0% 78474.3 19792
52163 0.150 34499 71.8% 24647 15502 10916 20308 59.9% 73260.0 15837
51147 0.125 29347 75.7% 21664 11355 10577 17662 65.2% 66495.2 11979
48883 0.100 23975 79.9% 18294 7671 9950 14747 71.0% 57943.4 8351
44979 0.075 18370 84.4% 14494 4558 8898 11543 77.3% 47337.8 5117
38923 0.050 12517 89.2% 10216 2142 7198 8032 84.3% 34377.8 2477
30047 0.025 6400 94.4% 5405 567 4468 4192 91.8% 18725.3 675 17467
0.000 0 100.0% 0 0 0 0 100.0% 0 0 0 .sup.1arbitrary units Based on
WB scan on Hannover ART with .about.5 mCi injection (Activity = .11
arbitrary units) Exact with septa has NEC of 17500 for 10 mCi
injection. ART Coincidence Sensitivity: 229000 cps/.mu.Ci/cc
PET/SPECT Coincidence Sensitivity: 816000 cps/.mu.Ci/cc NEC does
not include a scatter term for either scanner
[0101] From the foregoing description, it will be recognized by
those skilled in the art that a detector array, having particular
application in Single Photon Imaging which includes traditional
Gamma Cameras, Planar Imaging, Single Photon Emission Computed
Tomography (SPECT) with or without Coincidence Photon Imaging and
Positron Emission Tomography (PET), offering advantages over the
prior art has been described and shown. Specifically, the detector
array of the present invention incorporates either a single
scintillator layer or two, stacked discrete scintillator layers
that can be the same scintillator material or of two different
scintillator materials. In either case the different layers are
composed of materials that have distinctly different decay times.
The variants in these figures are the types of optical detectors
which are used, i.e. photomultipliers and/or photodiodes.
Additionally, the optical light guide can be integral with the
scintillators or optically bonded thereto. Further, the planar
light guide can be active or non-active. In either of these
variants, the planar light guide can be segmented or non-segmented.
And, if segmented, can either be inverted or non-inverted.
[0102] While a preferred embodiment has been shown and described,
it will be understood that it is not intended to limit the
disclosure, but rather it is intended to cover all modifications
and alternate methods falling within the spirit and the scope of
the invention as defined in the appended claims.
* * * * *