U.S. patent application number 11/199262 was filed with the patent office on 2006-03-30 for radiation imaging apparatus and nuclear medicine diagnosis apparatus using the same.
Invention is credited to Takafumi Ishitsu, Tsuneaki Kawaguchi, Hiroshi Kitaguchi, Shinya Kominami, Yuichi Morimoto, Masatoshi Tanaka, Katsutoshi Tsuchiya, Kazuma Yokoi.
Application Number | 20060065836 11/199262 |
Document ID | / |
Family ID | 35511292 |
Filed Date | 2006-03-30 |
United States Patent
Application |
20060065836 |
Kind Code |
A1 |
Tsuchiya; Katsutoshi ; et
al. |
March 30, 2006 |
Radiation imaging apparatus and nuclear medicine diagnosis
apparatus using the same
Abstract
Disclosed herein is a radiation imaging apparatus and
radiation-imaging-apparatus-based nuclear medicine diagnosis
apparatus having a collimator in which a plurality of rectangular
through-holes are arranged in a grid pattern and separated by septa
is rotated through a predetermined angle as viewed from above in
relation to the layout of a plurality of rectangular detectors that
are arranged in a grid pattern. The predetermined angle ranges from
20 to 70 degree and more preferably from 30 deg to 60 deg. With
this configuration, the influence of sensitivity variations (moire
patterns) that are included in an image picked up due to
interference with a collimator when pixel type detectors are used
is eliminated.
Inventors: |
Tsuchiya; Katsutoshi;
(Hitachi, JP) ; Kitaguchi; Hiroshi; (Naka, JP)
; Morimoto; Yuichi; (Hitachinaka, JP) ; Kominami;
Shinya; (Mito, JP) ; Yokoi; Kazuma; (Hitachi,
JP) ; Kawaguchi; Tsuneaki; (Kashiwa, JP) ;
Tanaka; Masatoshi; (Kashiwa, JP) ; Ishitsu;
Takafumi; (Hitachi, JP) |
Correspondence
Address: |
DICKSTEIN SHAPIRO MORIN & OSHINSKY LLP
2101 L Street, NW
Washington
DC
20037
US
|
Family ID: |
35511292 |
Appl. No.: |
11/199262 |
Filed: |
August 9, 2005 |
Current U.S.
Class: |
250/363.1 |
Current CPC
Class: |
A61B 6/037 20130101;
G01T 1/1648 20130101 |
Class at
Publication: |
250/363.1 |
International
Class: |
G21K 1/02 20060101
G21K001/02 |
Foreign Application Data
Date |
Code |
Application Number |
Sep 24, 2004 |
JP |
2004-277351 |
Mar 30, 2005 |
JP |
2005-96672 |
Claims
1. A radiation imaging apparatus comprising: a plurality of
rectangular detectors that are arranged in a grid pattern; a
radiation measurement circuit for reading detection signals of the
detectors; and a collimator in which a plurality of rectangular
through-holes are arranged in a grid pattern and separated by
septa; wherein the collimator is rotated through a predetermined
angle in relation to the layout of the detectors as viewed from
above, and used to acquire radiation incidence position information
from each of the detectors and image the acquired information.
2. The radiation imaging apparatus according to claim 1, wherein
the predetermined angle ranges from 20 deg to 70 deg.
3. The radiation imaging apparatus according to claim 2, wherein
the predetermined angle ranges from 30 deg to 60 deg.
4. A nuclear medicine diagnosis apparatus that uses the radiation
imaging apparatus according to claim 1.
5. A radiation imaging apparatus comprising: pixel type detectors
that acquire the position information about radiation incident on
image pixels that are arranged in a grid pattern; a radiation
measurement circuit for reading detection signal of the detector;
and a collimator in which a plurality of rectangular through-holes
are arranged in a grid pattern and separated by septa; wherein the
collimator is rotated through a predetermined angle in relation to
the layout of the detectors as viewed from above.
6. The radiation imaging apparatus according to claim 5, wherein
the predetermined angle ranges from 20 deg to 70 deg.
7. The radiation imaging apparatus according to claim 5, wherein
the predetermined angle ranges from 30 deg to 60 deg.
8. A nuclear medicine diagnosis apparatus, comprising: the
radiation imaging apparatus according to claim 5; a gantry for
supporting the radiation imaging apparatus; a tomographic
information creation device for acquiring detection signals from
the detectors and generating image information; and a display
device for displaying the image information.
9. The radiation imaging apparatus according to claim 6, wherein
the intervals at which the collimator through-holes are positioned
are {square root over (2)} to 1.8 times the intervals at which the
detectors' image pixels that are arranged in a grid pattern are
positioned.
10. The radiation imaging apparatus according to claim 9, wherein
the detectors' image pixels that are arranged in a grid pattern are
positioned at intervals of 1 mm or longer but shorter than 2
mm.
11. The nuclear medicine diagnosis apparatus according to claim 8,
wherein the tomographic information creation device performs a
3.times.3 matrix smoothing filter process.
Description
BACKGROUND OF THE INVENTION
[0001] 1. Technical Field
[0002] The present invention relates to a radiation imaging
apparatus that includes a pixel-type measurement system and images
an incident radiation distribution, and to a nuclear medicine
diagnosis apparatus that uses the radiation imaging apparatus.
[0003] 2. Background Art
[0004] A gamma camera, single photon emission computed tomography
(SPECT) apparatus that uses a gamma camera, or other nuclear
medicine diagnosis apparatus is used as an apparatus that uses a
radiation measurement device for medical purposes. Radiation
detectors (hereinafter may be referred to as detectors) for use in
such a nuclear medicine diagnosis apparatus are mostly a
combination of a scintillator and a photomultiplier tube. For such
a nuclear medicine diagnosis apparatus, a single, large crystal
plate is generally used. A NaI (T1) scintillator is widely used for
the gamma camera and SPECT apparatus.
[0005] FIG. 13 schematically illustrates the configuration of a
scintillator-based gamma camera. A single plate of scintillator
201, which comprises a relatively large single crystal, is a
detector that makes use of a phenomenon in which fluorescence is
generated when radiation energy is absorbed subsequently to
radiation incidence on a particular substance. The generated feeble
light is amplified by a plurality of photomultiplier tubes 203 to
achieve radiation detection. For radiation position measurement
purposes, the output signals generated from the plurality of
photomultiplier tubes 203 are subjected to gravity center
computation to determine a radiation reaction position.
[0006] To project a gamma ray generation position onto an image
pickup surface of the detector, a collimator 206 for controlling
the angle of radiation incidence is positioned in front of the
scintillator 201. At present, the collimator 206 is generally made
of lead that has an infinite number of hexagonal through-holes. The
through-hole diameter approximately ranges from 1 mm to 3 mm. The
through-hole length approximately ranges from 40 mm to 60 mm. The
septa among the through-holes approximately range from 0.2 mm to 3
mm. Hexagonal through-holes are used because they provide the
highest aperture ratio, are easy to fabricate, and exhibit high
strength. In FIG. 13, the reference numeral 202 denotes a light
guide; 204, a measurement circuit; and 205, a measurement circuit
retention board.
[0007] In recent years, individual pixel type detectors, which
acquire position signals in the unit of a small detector, that is,
on an individual pixel basis, have been developed, including a
gamma camera in which a CsI (T1)-based pixel type scintillator and
photodiode are used (Nuclear Medicine Examination Technology,
Japanese Society of Radiological Technology, Ohmsha, pp. 79-80) and
a semiconductor detector for directly converting radiation into
electrical signals (Nuclear Medicine Examination Technology,
Japanese Society of Radiological Technology, Ohmsha, pp. 76-77).
Detectors that determine the radiation reaction position by means
of aforementioned gravity center computation measure one gamma ray
by using a plurality of photomultiplier tubes to capture
scintillator-generated light as a spread of light. Therefore, it
can be said that the detectors make spatially continuous
measurements, that is, analog measurements. On the other hand, it
can be said that pixel-type detectors, which make measurements on
an individual pixel basis, measure one gamma ray by making
spatially discrete measurements, that is, spatially digital
measurements.
[0008] One measurement unit, that is, the radiation incidence cross
section of a pixel, of the above apparatuses is generally
rectangular. The collimator having hexagonal through-holes is not
suitable for the above apparatuses. The reason is that moire
patterns arise although they do not arise with the use of a
conventional scintillator, which comprises a single crystal. The
generation of moire patterns is a problem in which a plurality of
periodical sensitivity variations occur on an image when the
periodical shade changes of septa interfere with each pixel due to
the difference between the detector pitch and through-hole pitch
and anisotropy.
[0009] One solution to avoid moire patterns is to use a collimator
whose hole diameter is smaller than half the pixel size. When the
collimator through-hole is small, the following advantages are
provided. When the collimator is shifted horizontally in relation
to the detectors, a septum positioned over one detector is partly
positioned outside the detector. However, another septum, which has
virtually the same area and was positioned outside the detector, is
now positioned over the detector. As a result, the septum area over
the detector does not significantly change even when the collimator
is shifted. Consequently, the detector sensitivity does not
significantly change. In other words, the resulting image remains
almost unchanged because pixels are almost uniform in sensitivity
even when the collimator is moved forward, rearward, leftward, or
rightward, rotated, or otherwise shifted. The smaller the
collimator through-holes in relation to the detectors, the greater
the produced effect.
[0010] However, when the pixel size is 1 mm or larger, the above
solution does not work due to the manufacturing limitation imposed
on the collimator hole diameter. As a result, moire patterns cannot
be avoided.
[0011] Another solution is to use a matched collimator, which has
rectangular holes that match the pixel size. Since the sensitivity
loss due to the septa 28 is minimized for the pixel-type detectors,
it is said that the use of a matched collimator is ideal. However,
when the current lead-based collimator is used, it is difficult to
maintain the manufacturing accuracy in order to provide the
advantages of the matched collimator. The reason is that lead is
relatively soft and likely to deform. Further, if, for instance,
the collimator mounting position is slightly shifted from normal, a
great sensitivity variation may arise. The collimator can be made
of relatively hard tungsten in order to maintain the required
manufacturing accuracy. Such a solution may work with collimators
for use in a small-size gamma camera, but does not provide a
practical solution for collimators for use in a normal gamma
camera, SPECT, or the like in terms of cost.
[0012] Further, the gamma camera rotates or moves in a complicated
manner during an image pickup operation. During such a movement,
the collimator may deviate from a specified position.
[0013] Even while the gamma camera is at a standstill for a long
period of time, the collimator may gradually deviate from a
specified position due to its weight.
[0014] When displaced, the collimator incurs moire patterns no
matter whether a matched collimator is used.
SUMMARY OF THE INVENTION
[0015] It is an object of the present invention to provide a
radiation imaging apparatus and nuclear medicine diagnosis
apparatus that have the aforementioned pixel type measurement
system and are capable of avoiding moire patterns, which may be
generated during the use of the aforementioned hexagonal collimator
or matched collimator.
[0016] In accomplishing the above object, according to one aspect
of the present invention, there is provided a radiation imaging
apparatus comprising a plurality of rectangular detectors that are
arranged in a grid pattern; a radiation measurement circuit for
reading detector signals; and a collimator in which a plurality of
rectangular through-holes are arranged in a grid pattern and
separated by septa. The radiation imaging apparatus uses the
collimator to control the angle of radiation incidence and images
radiation incidence position information on an individual
rectangular detector basis. The collimator is rotated through a
predetermined angle in relation to the layout of the detectors as
viewed from above.
[0017] According to another aspect of the present invention, there
is provided the radiation imaging apparatus, wherein the
predetermined angle ranges from 20 deg to 70 deg and more
preferably from 30 deg to 60 deg.
[0018] According to another aspect of the present invention, there
is provided a nuclear medicine diagnosis apparatus that uses the
radiation imaging apparatus.
[0019] According to still another aspect of the present invention,
there is provided a radiation imaging apparatus comprising a
plurality of pixel type detectors for acquiring radiation incidence
position information in accordance with image pixels that are
arranged in a grid pattern; a radiation measurement circuit for
reading detection signals from the detectors, and a collimator in
which a plurality of rectangular through-holes are arranged in a
grid pattern and separated by septa. The collimator is rotated
through a predetermined angle in relation to the grid layout of the
detectors as viewed from above.
[0020] When the above configuration is employed, it is possible to
avoid moire patterns that are fatal to the operation performed to
image a radiation distribution (radiation source position) with
pixel type detectors. Further, low-cost lead may be used for
collimator manufacture because the required manufacturing accuracy
and mounting accuracy are not high. As a result, the apparatus cost
can be minimized.
BRIEF DESCRIPTION OF THE DRAWINGS
[0021] FIG. 1 illustrates a SPECT apparatus (gamma camera)
according to one embodiment of the present invention;
[0022] FIG. 2 illustrates a set of pixel type detectors according
to one embodiment of the present invention;
[0023] FIG. 3 illustrates another set of pixel type detectors;
[0024] FIGS. 4A and 4B illustrate another set of pixel type
detectors. FIG. 4A shows the top surface of the set of pixel type
detectors. FIG. 4B shows the bottom surface of the set of pixel
type detectors;
[0025] FIG. 5 illustrates a collimator and detectors according to
one embodiment of the present invention;
[0026] FIG. 6 shows a collimator (hexagonal holes) as a first
comparative example;
[0027] FIG. 7 relates to the first comparative example and
schematically shows a moire pattern that is generated when a
combination of a hexagonal collimator and pixel detectors is viewed
from above;
[0028] FIG. 8 shows a matched collimator as a second comparative
example;
[0029] FIG. 9 relates to the second comparative example and
schematically shows a moire pattern that is generated when a
combination of a matched collimator and pixel detectors is viewed
from above. The upper drawing illustrates a situation where the
collimator is horizontally shifted in relation to the detectors.
The lower drawing illustrates a situation where the collimator is
horizontally shifted in relation to the detectors and slightly
rotated as viewed from above;
[0030] FIGS. 10A to 10D relate to a first embodiment and
schematically shows a moire pattern that is generated when a
combination of a collimator and pixel detectors is rotated through
a predetermined angle as viewed from above;
[0031] FIGS. 11A to 11D relate to a second embodiment and
schematically shows a moire pattern that is generated when a
combination of a collimator and pixel detectors is rotated through
a predetermined angle as viewed from above;
[0032] FIG. 12 illustrates the relationship between a collimator
rotation angle and moire cycle;
[0033] FIG. 13 schematically shows the configuration of a
scintillator-based gamma camera;
[0034] FIG. 14 illustrates the relationship between a collimator
rotation angle and moire cycle;
[0035] FIGS. 15A and 15B show the image pickup simulation results
that are obtained when a conventional hexagonal collimator is used.
FIG. 15A illustrates a case where raw data is used. FIG. 15B shows
a case where a 3.times.3 weighted smoothing filter is used;
[0036] FIGS. 16A and 16B show the image pickup simulation results
that are obtained when a rectangular collimator, which is rotated
through 45 deg in accordance with one embodiment of the present
invention, is used. FIG. 16A illustrates a case where raw data is
used. FIG. 16B shows a case where a 3.times.3 weighted smoothing
filter is used; and
[0037] FIG. 17 illustrates still another set of pixel type
detectors.
DETAILED DESCRIPTION OF THE PREFERRED EMBODIMENTS
[0038] An embodiment of the present invention will now be described
with reference to the accompanying drawings. In the following
description, the terms "detector" and "detector array" are used.
The detector corresponds to a rectangular pixel, whereas the
detector array denotes a set of detectors that are arranged in a
grid pattern.
[0039] As shown in FIG. 1, a SPECT apparatus 1 includes a gantry
10, cameras (image pickup devices) 11A, 11B, a data processing
device 12, and a display device 13. A radiopharmaceutical, such as
a pharmaceutical containing .sup.99mTc having a half-life of 6
hours, is administered to an examinee 15. Gamma rays emitted from
.sup.99mTC in the body of the examinee on a bed 14 are detected
with the cameras 11, which are supported by the gantry 10, to
obtain a tomogram.
[0040] The cameras 11 include a collimator 26 and a large number of
detectors 21, which comprise a semiconductor device. The collimator
26 selects gamma rays emitted from the body of the examinee 15 so
that only gamma rays oriented in a certain direction pass through.
Gamma rays passing through the collimator 26 are detected by the
detectors 21. The cameras 11 include an ASIC (application-specific
integrated circuit) 25 for measuring a gamma ray detection signal.
The gamma ray detection signal is delivered to the ASIC 25 via a
detector circuit board 23 and an ASIC circuit board 24 for the
purpose of inputting the ID of a detector 21 that has detected a
gamma ray, the pulse height value of the detected gamma ray, and
the gamma ray detection time. The components are enclosed within a
light/gamma ray/electromagnetic shield 29, which is a part of the
cameras 11, made of iron, lead, and the like, and used to block
light, gamma rays, and electromagnetic waves. The data processing
device 12 includes a storage device and a tomographic information
creation device (not shown). The data processing device 12 acquires
packet data, which includes the pulse height value of a measured
gamma ray, detection time data, and detector (channel) ID,
generates a planar image or generates tomogram information by
converting the packet data into sinogram data, and displays the
resultant image on the display device 13.
[0041] The cameras 11 can be moved in the radial direction or
circumferential direction of the gantry 10. The cameras 11 pick up
an image while moving along the contour of the examinee 15. The
cameras 11 can also rotate around a gantry mount. When the two
cameras 11A, 11B are fixed side by side, it is possible to obtain a
STATIC image. In this manner, the radiopharmaceutical accumulated,
for instance, on a tumor in the body of the examinee 15 is imaged
to determine the location of the tumor.
[Detector and Collimator]
[0042] Characteristic portions of the present embodiment of the
present invention will now be described.
[0043] As indicated in FIG. 2, the detectors 21 for use in a camera
11 are individually provided for pixels as a rectangular
parallelepiped having upper and lower rectangular surfaces. A large
number of detectors 21 are arranged in a grid pattern to form a
detector array 21A. In marked contrast to the use of a
scintillator, which comprises one big crystal as indicated in FIG.
13, detection signals are collected by individual detectors 21,
that is, on an individual pixel basis. The structure of the
detector array 21A need not be divisible into individual pixels.
Alternatively, the structure may be such that electrodes are
provided for individual pixels as indicated in FIG. 3. Another
alternative structure may also be employed so that a detector array
21C comprises detectors 21 that are separated by dicing as
indicated in FIGS. 4A and 4B. It goes without saying that a
scintillator may be separated into individual pixels to form a
number of detectors 21. In the present embodiment, the term "grid
pattern" means that vertically arranged detectors 21 and
horizontally arranged detectors 21, which constitute detector array
21A, 21B, or 21C, are orthogonal with each other as shown in FIGS.
2, 3, 4A, and 4B.
[0044] The collimator 26A used in the present embodiment is made of
lead. As shown in FIG. 5, the collimator 26A has rectangular
through-holes 27A, which are arranged in a grid pattern. The
through-holes 27A are separated from each other by septa. As is
obvious from FIG. 5, there is a predetermined rotation angle (angle
of cut) between the through-holes 27A arranged within the
collimator 26A and the detectors 21 arranged within the detector
array 21A when they are viewed from above. Typically, the
predetermined rotation angle is 45 deg. As is the case with the
detector array 21A, the term "grid pattern" means that crossing
lines of through-holes 27A are orthogonal with each other.
[0045] The operations and advantages of the present embodiment, in
which the collimator is rotated through a predetermined angle in
relation to the layout of the detectors as viewed from above, will
now be described with reference to comparative examples.
COMPARATIVE EXAMPLE 1
[0046] FIG. 6 is furnished as a first comparative example. It is a
perspective view illustrating a collimator that has hexagonal,
honeycomb through-holes. This collimator 26C is made of lead. The
hexagonal through-holes 27C are separated from each other by septa
28C. In the SPECT apparatus (gamma camera) 1, a collimator 26
having hexagonal through-holes as shown in FIG. 6 were formerly
used (see FIG. 13). In recent years, however, a new problem has
occurred due to the development of a pixel type scintillator and
semiconductor detector, which detect gamma rays in the unit of a
pixel. More specifically, sensitivity variations are caused by the
shades of the collimator septa 28 due to the use of a spatially
digitized detector system so that moire patterns appear on an image
picked up.
[0047] When a photomultiplier tube based conventional technology
was used, no moire-related problem occurred. The through-holes 27
of the collimator 26 are far smaller than the photomultiplier tubes
203 (see FIG. 13). Even if the collimator 26 is displaced in a
situation where the aforementioned collimator hole diameter is
smaller than the pixel size, the septa 28 having virtually the same
area as the septa 28 placed outside a photomultiplier tube 203 is
placed over a photomultiplier tube 203. Therefore, the
photomultiplier tube sensitivity does not significantly change. In
other words, the septa 28 do not significantly affect the
photomultiplier tube sensitivity. Further, since one gamma ray is
dispersed by a plurality of photomultiplier tubes 203 for
measurement purposes, the septa exercise averaged influence.
Consequently, the moire phenomenon does not become obvious.
However, when a pixel type detector is used so that the diameter of
the through-holes 27 of the collimator 26 is close to the size of
the detector 21, great influence is exercised by moire patterns.
Further, detection counting is performed for each pixel. Therefore,
the sensitivity differences among the detectors 21 directly affect
the image. As described later, the shades of the septa 28 of the
collimator 26 vary from one location to another and produce great
sensitivity differences.
[0048] FIG. 7 is furnished as the first comparative example to
illustrate moire patterns that are generated when the collimator
shown in FIG. 6 is used with the detectors shown in FIG. 2. As
indicated in FIG. 7, when the collimator 26C having hexagonal
through-holes 27C as shown in FIG. 6 is used with the detectors 21
(detector array 21A) shown in FIG. 2, moire patterns are generated
if uniform gamma rays are emitted from a planar radiation source.
In the subsequent figures, moire patterns are represented in terms
of image density variation. More specifically, a dark area is a
place where there are many pixels (or detectors 21) whose
sensitivities are significantly lowered, whereas a light area is a
place where there are many pixels whose sensitivities have
insignificantly varied from normal. In FIG. 7, moire patterns are
represented by lines (lines corresponding to the septum thickness
and lines corresponding to a gap between the detectors) that have a
certain thickness and overlap. However, the same density
differences are found in an actual image that is indicated by the
densities of individual pixels (the number of gamma rays detected
by each pixel). If moire patterns constantly appear at the same
locations, it is possible to make corrections to obtain an image
that is close to a correct image.
[0049] However, (see FIG. 1), the moire patterns also depend on the
depth of a diseased part that is to be imaged. The through-holes 27
of the collimator 26 have a certain length. Therefore, when a
radiation source is positioned near the collimator 26, only gamma
rays emitted from a narrow region can pass through the
through-holes 27. If, on the other hand, the radiation source is
positioned far from the collimator 26, gamma rays emitted from a
wide region can pass through the through-holes 27. In reality,
there is a certain gap between the detectors 21 and collimator 26.
Therefore, the gamma rays emitted from the wide region are not only
incident on the detectors 21 that are directly below the
through-holes 27, but also incident on the neighboring detectors
21. It means that the shades of the septa 28 not only affect the
detectors 21 positioned directly below the through-holes but also
affect the neighboring detectors 21. Therefore, individual detector
sensitivities vary depending on the depth of the diseased part so
that the moire patterns vary. Further, the patterns vary when a
slight positional change occurs due, for instance, to collimator
replacement or the cameras 11 are rotated to slightly change the
position of the collimator 26. Since these factors affect in a
complicated manner, it is very difficult to make proper corrections
at an image processing stage.
COMPARATIVE EXAMPLE 2
[0050] FIG. 8 is furnished as a second comparative example. It
illustrates a matched collimator, which matches the layout of
detectors. It is said that a matched collimator 26B should be used
to avoid moire patterns. The matched collimator 26B is such that
the positions of the through-holes 27B of the collimator match the
positions of the detectors 21 as indicated in FIG. 8. As described
earlier, each of the detectors 21 according to the present
embodiment is a rectangular parallelepiped having a rectangular
upper surface. These detectors 21 are arranged in a grid pattern to
form the detector array 21A. The matched collimator 26B includes
rectangular through-holes 27B, which are separated by septa 28B.
The through-holes 27B are arranged in a grid pattern to match the
layout of the detectors 21 within the detector array 21A. The
through-holes 27B are oriented in the same direction as the
detectors 21 as viewed from above (rotation angle=0 deg). The size
of each through-hole 27B is nearly the same as that of the top
surface of a detector 21. However, when the pixel size becomes
smaller, it is difficult to accurately align the detectors 21 with
the through-holes 27B of the matched collimator 26B. Further, it is
difficult to maintain the manufacturing accuracy required for the
matched collimator 26B.
[0051] Further, if the matched collimator 26B is displaced, great
sensitivity differences arise. FIG. 9 illustrates moire patterns
that are generated when the matched collimator is displaced. The
upper drawing in FIG. 9 illustrates a situation where the matched
collimator 26B is displaced in parallel with the detector array
21A. The lower drawing in FIG. 9 illustrates a situation where the
matched collimator 26B is displaced in parallel with the detector
array 21A and then rotated. When the matched collimator 26B is
displaced in parallel with the detector array 21A, the septa 28B of
the matched collimator 26B are uniformly placed over the detectors
in the detector array 21A. Therefore, the overall sensitivity is
greatly lowered while the sensitivity provided by a perfect match
is 100%. When parallel displacement and rotary displacement both
occur, great moire patterns are generated. In the SPECT apparatus 1
(see FIG. 1), which includes the lead collimator 26 that is 50 cm
square and approximately 100 kg in weight, the cameras 11 rotate
and the collimator 26 bends or becomes displaced. Therefore, it is
extremely difficult to achieve proper positioning in any state.
These problems can be solved by using a collimator 26 that is made
of tungsten instead of lead. However, the use of tungsten raises
the production cost and is not practical except for collimator use
with a small camera. As described above, the use of pixel type
detectors 21 and the moire-related problem are unavoidable.
EMBODIMENTS
[0052] The practical collimator 26 should be manufactured at a low
cost and constantly achieve the same image quality without
significantly varying the sensitivity from one pixel to another,
and is acceptable even if does not exhibit high manufacturing
accuracy and mounting accuracy. Alternatively, the employed image
pickup system, which includes the detectors 21, should meet the
above requirements.
[0053] In the embodiments, therefore, the collimator is provided
with rectangular through-holes. The layout of the through-holes is
rotated through a predetermined angle in relation to the layout of
the detectors within the detector array as viewed from above.
[0054] FIGS. 10A to 10D (first embodiment) and FIGS. 11A to 11D
(second embodiment) indicate moire patterns that are generated
while the rotation angle varies. These figures are obtained when
the detectors and collimator are viewed from above. FIGS. 10A and
11A present views prevailing when the rotation angle is 0 deg.
FIGS. 10B and 11B present views prevailing when the rotation angle
is 15 deg. FIGS. 10C and 11C present views prevailing when the
rotation angle is 30 deg. FIGS. 10D and 11D present views
prevailing when the rotation angle is 45 deg. The ratio between the
collimator through-hole pitch and detector pitch is 1.0 for the
matched collimator, 1.1 in FIGS. 11A to 11D, and 1.8 in FIG.
12.
[0055] When the collimator is rotated through 30 deg or more,
almost no moire patterns are visible in FIGS. 10A to 10D, which
indicate the first embodiment, and in FIGS. 11A to 11D, which
indicate the second embodiment. This moire reduction effect is
similarly produced when the position of the collimator 26A is
slightly changed by rotation or parallel displacement. When the
through-holes 27A of the collimator 16A are rectangular with the
through-hole layout direction rotationally displaced from the pixel
layout direction as described above, it is possible to avoid moire
patterns. The rotation angle should range from 20 to 70 degree and
more preferably from 30 deg to 60 deg.
[0056] FIG. 12 (see FIGS. 10A to 10D and 11A to 11D as needed)
indicates that a rotation angle of 30 deg is an angle at which the
moire cycle/detector pitch is 2.0 or smaller for all collimator
pitches according to the present embodiment. In other words, when
the rotation angle is 30 deg or greater, moire patterns disappear
at all collimator pitches according to the present embodiment.
Since symmetry is achieved at a rotation angle of 45 deg, moire
patterns disappear at all collimator pitches according to the
present embodiment when the rotation angle is 60 deg or smaller.
The reason why moire patterns disappear when the moire
cycle/detector pitch is 2.0 or smaller will be described later in
detail.
[0057] When the rotation angle is 20 deg or greater, it can be
expected that the moire reduction effect is produced. Satisfactory
results are obtained particularly when the collimator pitch is 1.5
mm or 1.8 mm. When the rotation angle is smaller than 20 deg, the
influence of moire patterns greatly increases. As regards the
relationship between the moire effect and rotation angle, symmetry
is achieved at a rotation angle of 45 deg. Therefore, when the
rotation angle is 70 deg or smaller, it can be expected that the
moire reduction effect is produced. Satisfactory results are
obtained particularly when the collimator pitch is 1.5 mm or 1.8
mm. When the rotation angle is 70 deg or greater, the influence of
moire patterns greatly increases.
[0058] It is possible to sufficiently avoid moire patterns at a
rotation angle of 20 deg by selecting an appropriate collimator
pitch no matter whether it is indicated in FIG. 12.
[0059] FIG. 12 uses the ratio between the detector pitch and
collimator hole pitch as a parameter and shows the measured ratio
between the detector pitch and the moire cycle T.sub.M for the
rotation angle .theta. of the collimator 26. The moire cycle
T.sub.M represents the distance between the darkest spot of a moire
pattern and the darkest spot of a neighboring moire pattern (or the
distance between the lightest spot of a moire pattern and the
lightest spot of a neighboring moire pattern). Since moire patterns
at rotation angles of 45 deg or greater are the same as for
(90-.theta.), the maximum rotation angle shown in FIG. 12 is 45
deg. Further, when the collimator pitch is more than twice the
detector pitch, many pixels are not shaded by the septa 28.
Therefore, the shades of the septa 28 are directly projected onto
the pixels so that no conspicuous moire patterns are generated.
Moire patterns are conspicuous only when the detector pitch is
close to the collimator pitch. That is why the maximum collimator
pitch shown in FIG. 12 is two times the detector pitch. As
indicated in FIGS. 10A to 10D and 11A to 11D, virtually no moire
patterns are visible when the moire cycle T.sub.M (see FIG. 10B) is
not more than two times the detector pitch, that is, when the
rotation angle is 30 deg or greater. When the rotation angle is 35
deg or greater, no measured data is indicated in FIG. 12 because
the moire cycle T.sub.M could not be measured. To be precise,
almost no moire patterns are generated when T.sub.M sin .phi. and
T.sub.M cos .phi. are not more than two times the detector pitch
p.sub.D, where T.sub.M is the moire cycle and .phi. is a moire
angle (slightly different from the actual rotation angle .theta.),
which is formed between a moire pattern and the detector 21. The
reason is that the minimum cycle for digital imaging is 2 pixels
(that is, when white and black are adjacent to each other). Moire
patterns cannot be recognized when the moire cycle T.sub.M for
projection onto to the pixel layout is 2 pixels, that is, not more
than two times the detector pitch p.sub.D.
[0060] However, moire disappearance is one thing and detector
sensitivity uniforming is another. When the collimator pitch is
relatively great, the whole area of a pixel is positioned within a
through-hole of the collimator 26A' as indicated in FIGS. 11A to
11D. Therefore, a pixel having the maximum sensitivity exists.
Meanwhile, an intersection of septa 28A may overlap a detector 21
for one pixel so that the sensitivity is minimized. In such an
instance, however, there are no great periodic sensitivity
differences but local sensitivity differences. In an actual image
pickup operation, there are up to several hundred counts per pixel.
Local sensitivity differences caused on an individual pixel basis
are not greater than statistical errors, and do not incur any
serious problem.
[0061] When pixel type detectors and a collimator having
rectangular holes are positioned with their layout orientations
displaced from each other as described above, it is possible to
provide a collimator that is capable of avoiding moire patterns and
not dependent on positioning accuracy or manufacturing accuracy. In
other words, the collimator can be made of low-cost lead so that
the manufacturing cost is maintained at a previous level. Further,
various collimator hole diameters and depths are selectable in the
same manner as before. Consequently, a high degree of versatility
results.
[0062] To avoid moire patterns, a matched collimator having
rectangular holes that match the pixel size may be used. Although
the use of such a matched collimator was described earlier, it will
now be described in detail. It is said that a matched collimator
whose hole positions match the pixel positions (in other words, the
septum positions match the detector pixel gap positions) is ideal
for use with pixel type detectors, because the sensitivity loss by
septa 28 is minimized. However, with the current lead-based
collimator, it is difficult to maintain the required manufacturing
accuracy in order to make the most of the features of a matched
collimator. The reason is that lead is relatively soft and likely
to deform. When, for instance, through-holes 27 having a depth of
more than 40 mm are manufactured with 0.2 mm thick septa 28
arranged at a pitch of 1.4 mm, it is extremely difficult to
position the septa and through-holes within a large collimator
measuring 400 mm by 500 mm to an accuracy of 0.05 to 0.1 mm. Even
if the mounting positions are slightly displaced, great sensitivity
variations may result. Hard tungsten or relatively hard tungsten
alloy can be used to manufacture a collimator in order to maintain
the required manufacturing accuracy. However, tungsten is an
expensive metal and its machining cost is extremely high. Such a
solution may work with collimators for use in a small-size gamma
camera, but does not provide a practical solution for collimators
for use in a normal gamma camera, SPECT, or the like in terms of
cost.
[0063] Further, the gamma camera rotates or moves in a complicated
manner during an image pickup operation. During such a movement,
the collimator may deviate from a specified position. Even while
the gamma camera is at a standstill for a long period of time, the
collimator may gradually deviate from a specified position due to
its weight. When displaced, the collimator incurs moire patterns no
matter whether a matched collimator is used. This problem can be
solved by the embodiment described above.
[0064] As mentioned earlier, FIG. 12 uses the ratio between a
detector pitch of 1 mm and the collimator hole pitch as a parameter
and shows the measured ratio between the detector pitch and the
moire cycle T.sub.M for the rotation angle .theta. of the
collimator 26. In other words, the geometric similarity rule can be
applied to this relationship. When the parameter is p.sub.C/p.sub.D
(the ratio of the collimator hole pitch p.sub.C to the detector
pitch p.sub.D) and rendered dimensionless by the detector pitch
p.sub.D, the same graphed relationship as indicated in FIG. 12 is
obtained as shown in FIG. 14. The moire cycle T.sub.M is the
distance between the darkest spot of a moire pattern and the
darkest spot of a neighboring moire pattern (or the distance
between the lightest spot of a moire pattern and the lightest spot
of a neighboring moire pattern). Since moire patterns at rotation
angles of 45 deg or greater are the same as for a rotation angle of
(90-.theta.), the maximum rotation angle shown in FIGS. 12 and 14
is 45 deg. In other words, rotation angle symmetry occurs beginning
with an angle of 45 deg. Therefore, the same moire patterns are
generated at angles of 45+.+-.(45.degree.-.theta.). Further, when
the collimator pitch is more than twice the detector pitch, many
pixels are not shaded by the septa 28. Therefore, the shades of the
septa 28 are directly projected onto the pixels so that no
conspicuous moire patterns are generated. Moire patterns are
conspicuous only when the detector pitch is close to the collimator
pitch. That is why the maximum collimator pitch shown in FIGS. 12
and 14 is two times the detector pitch. FIGS. 12 and 14 (see FIGS.
10A to 10D and 11A to 11D as needed) indicate that a rotation angle
of 30 deg is an angle at which the moire cycle T.sub.M/detector
pitch p.sub.D is 2.0 or smaller for all collimator pitches
according to the present embodiment. In other words, when the
rotation angle is 30 deg or greater, moire patterns, which are
sensitivity variations occurring at long intervals, disappear at
all collimator pitches according to the present embodiment for the
reason described later. As regards the relationship between the
moire effect and rotation angle, symmetry is achieved at a rotation
angle of 45 deg. Therefore, when the rotation angle is between 30
deg (45.degree.-15.degree.) and 60 deg (45.degree.+15.degree.) ,
moire patterns disappear at all collimator pitches according to the
present embodiment.
[0065] When the rotation angle is 20 deg or greater, it can be
expected that a moire reduction effect is produced. A good moire
reduction effect is produced at collimator pitches of 1.5 mm and
1.8 mm (while the ratio to the detector pitch is 1.5 or 1.8)
particularly when the detector pitch is 1 mm. A good moire
reduction effect is produced at collimator pitches of 2.1 mm and
2.52 mm when the detector pitch is 1.4 mm, at collimator pitches of
2.4 mm and 2.88 mm when the detector pitch is 1.6 mm, or at
collimator pitches of 3.0 mm and 3.6 mm when the detector pitch is
2.0 mm. When the rotation angle is smaller than 20 deg, the
influence of moire patterns remarkably increases. As regards the
relationship between the moire effect and rotation angle, symmetry
is achieved at a rotation angle of 45 deg. Therefore, the result
obtained when the rotation angle is 70 deg or smaller is the same
as that is obtained when the rotation angle is 20 deg or greater.
Consequently, the above statement holds true when the rotation
angle is between 20 deg (45.degree.-25.degree.) and 70 deg
(45.degree.+25.degree.).
[0066] Particularly if the rotation angle is 45 deg, a sensitivity
difference arises at intervals of 2 pixels in a situation where the
collimator pitch p.sub.C is {square root over (2)} times the
detector pitch p.sub.D (the collimator pitch is 1.98 mm when the
detector pitch is 1.4 mm, the collimator pitch is 2.26 mm when the
detector pitch is 1.6 mm, or the collimator pitch is 2.83 mm when
the detector pitch is 2.0 mm). When the resulting periodical
position is slightly shifted from the detector center, moire
patterns can be mostly removed. Consequently, it is possible to
reduce the influence of moire patterns with extreme effectiveness.
Even when the cycle is 2 pixels, periodical sensitivity differences
can be eliminated with a 3.times.3 smoothing filter.
[0067] It is possible to sufficiently avoid moire patterns at a
rotation angle of 20 deg or 70 deg by selecting an appropriate
collimator pitch no matter whether it is indicated in FIGS. 12 and
14.
[0068] The reason why moire patterns disappear when the moire cycle
T.sub.M/detector pitch p.sub.D is 2.0 or smaller has already been
described with reference to FIG. 12. However, it will now be
described in detail. As indicated in FIGS. 10A to 10D and 11A to
11D, moire patterns almost disappear when the moire cycle T.sub.M
(see FIG. 10B) is smaller than two times the detector pitch p.sub.D
as indicated in FIGS. 12 and 14, that is, at a rotation angle of 30
deg or greater (FIGS. 10C, 10D, 11C, and 11D). When the rotation
angle is 35 deg or greater, no measured data is indicated in FIGS.
12 and 14 because the moire cycle T.sub.M could not be measured. To
be precise, moire patterns disappear when T.sub.M sin .phi. and
T.sub.M cos .phi., which are moire cycles prevailing when oblique
moire patterns are projected onto a vertical/horizontal layout, are
not more than two times the detector pitch p.sub.D, where T.sub.M
is the moire cycle and .phi. is a moire angle (slightly different
from the actual rotation angle .theta.), which is formed between a
moire pattern and the detector 21. (T.sub.M/p.sub.D)sin
.phi..ltoreq.2, (T.sub.M/p.sub.D)cos .phi..ltoreq.2 (Equation
1)
[0069] The minimum cycle for digital imaging is 2 pixels (that is,
when pixels having different densities such as black and white
pixels are positioned adjacent to each other). Therefore, moire
patterns, which are periodical density variations, cannot be
recognized on image pixels when the moire cycle T.sub.M for
projection onto to the vertical or horizontal pixel layout is not
longer than 2 pixels, that is, not more than two times the detector
pitch p.sub.D. If, for instance, white and black moire patterns are
alternately positioned obliquely on screen, white and black moire
patterns are also alternately positioned when viewed in the
vertical or horizontal direction. The minimum cycle for expressing
such density variations is 2 pixels. If the cycle is shorter than 2
pixels, it cannot be expressed on screen (cannot be measured).
Since sin .phi. and cos .phi. are smaller than 1, moire patterns
are inevitably unrecognizable if T.sub.M/p.sub.D is smaller than 2.
This corresponds to a situation where the rotation angle is between
30 deg and 60 deg as described with reference to FIGS. 12 and 14.
FIGS. 12 and 14 include values measured in a situation where
T.sub.M/p.sub.D is smaller than 2. The reason is that analog
overlays shown in FIGS. 10A to 10D and 11A to 11D are used for
visual measurements. When actual pixel-based digital images are
used, moire patterns are unrecognizable.
[0070] Even if T.sub.M/p.sub.D is not smaller than 2, Equation 1
may be satisfied depending on the moire angle .phi.. The maximum
condition for T.sub.M/p.sub.D is .phi.=45.degree.. In other words,
moire patterns may be rendered unrecognizable until
T.sub.M/p.sub.D=2.83. This is another critical point. In FIGS. 12
and 14, an angle of approximately 20 deg corresponds to the
critical point. Near the critical point, Equation 1 may be
satisfied depending on the combination of collimator and detector
pitches.
[0071] However, the sensitivity cannot completely be uniformed by
controlling moire patterns. When the collimator pitch is relatively
great, the whole area of one pixel is positioned within a
through-hole of the collimator 26A' as indicated in FIGS. 11A to
11D so that pixels having the maximum sensitivity locally exist.
Meanwhile, the intersection of septa 28A may be positioned over a
1-pixel detector 21 so as to minimize the sensitivity. However, the
resulting sensitivity difference is not so periodical. Only a local
sensitivity difference arises. In an actual image pickup operation,
there are several hundred counts per pixel. Local sensitivity
differences caused on an individual pixel basis do not incur any
serious problem because they are not greater than statistical
errors. Images having increased uniformity can be obtained when a
3.times.3 smoothing filter is added for post-processing
purposes.
[0072] As described earlier, two easy-to-use hardware solutions can
be applied to control moire patterns. One is the use of a
collimator whose hole diameter is less than half the pixel size.
The other is the use of a matched collimator having rectangular
holes that match the pixel size. Another method is to control moire
patterns by performing a post-measurement process, that is, a
software process. In this method, a smoothing filter or other
diffusion filter is used to make moire patterns inconspicuous.
However, moire patterns are sensitivity variations having a long
cycle. To make such moire patterns inconspicuous, it is necessary
to provide an increased degree of smoothing with a 5.times.5 matrix
smoothing filter or the like. However, the use of a smoothing
filter having a great matrix considerably deteriorates the spatial
resolution. In practice, therefore, a weighted smoothing filter
having a 3.times.3 matrix, which is the minimum unit of a filter,
is used. As described with reference to an after-mentioned
comparison example (FIGS. 15A and 15B), the influence of moire
patterns remains in a final image when the above-mentioned
3.times.3 matrix weighted smoothing filter is used.
[0073] Typical digital images that are actually displayed are shown
in FIGS. 15A, 15B, 16A, and 16B. FIGS. 15A and 15B show simulation
results that are obtained when a conventional collimator having
hexagonal holes is used. FIGS. 16A and 16B show simulation results
that are obtained when a collimator having rectangular holes, which
are rotated through 45 deg, is used. FIGS. 15A and 15B show
calculation results that are obtained in a situation where a
hexagonal hole collimator having a side-to-side distance of 1.8 mm,
a septum thickness of 0.18 mm, and a length of 39.5 mm is used in
relation to a detector pitch of 1.4 mm with the collimator
positioned at a distance of 6.8 mm from the detector and a .phi.40
mm Co-57 planar radiation source positioned at a distance of 100 mm
from the collimator. FIGS. 16A and 16B show calculation results
that are obtained in a situation where a rectangular hole
collimator having a side-to-side distance of 1.8 mm, a septum
thickness of 0.18 mm, and a length of 40.0 mm is used in relation
to a detector pitch of 1.4 mm with the collimator positioned at a
distance of 4.0 mm from the detector and a .phi.60 mm Co-57 planar
radiation source positioned at a distance of 100 mm from the
collimator. In FIG. 15A, sensitivity variations having a great
periodical structure, that is, moire patterns, are conspicuous. As
indicated in FIG. 15B, the moire patterns do not disappear even
when a common, 3.times.3 weighted smoothing filter is used. In FIG.
16A, slight sensitivity variations remain, but sensitivity
variations having a long cycle are greatly reduced when compared to
moire patterns shown in FIGS. 15A and 15B. This situation can be
sufficiently applied to actual use. However, when a 3.times.3
weighted smoothing filter or the like is used, sensitivity
variations are mostly rendered unrecognizable as shown in FIG.
16B.
[0074] When a 4.times.4 smoothing filter, 5.times.5 smoothing
filter, or other large filter is used, it is possible to control
moire patterns without resort to the present invention. However,
such a large smoothing filter cannot be used under normal
conditions because it lowers the spatial resolution.
[0075] The present invention uses a 3.times.3 smoothing filter to
shorten the cycle of sensitivity variations having a long cycle,
which are recognized as moire patterns, until they cannot be
sufficiently recognized. As a result, the present invention can
control and uniform moire patterns without significantly impairing
the spatial resolution. It goes without saying that the present
invention can be used without causing any practical problem even if
no filter is used.
[0076] As described above, when pixel type detectors and a
collimator having rectangular holes are positioned with their
layout orientations displaced from each other, it is possible to
provide a collimator that fixes a moire problem unique to the pixel
type detectors and does not require high positioning accuracy or
manufacturing accuracy unlike a matched collimator. In other words,
the collimator can be made of low-cost lead so that the
manufacturing cost is maintained at a previous level. Further,
various collimator hole diameters and depths are selectable in the
same manner as before. Consequently, a high degree of versatility
results.
[0077] FIGS. 2, 3, 4A, and 4B are used to illustrate detectors and
a collimator in conjunction with the foregoing embodiment. FIG. 17
illustrates another example of a pixel type detector. A cross strip
type detector may also be used. As indicated in FIG. 17, the cross
strip type detector is such that the front and back surfaces of a
semiconductor detector are provided with lines of electrodes, which
are orthogonal with each other, and that the intersections of
electrode lines are regarded as pixels. In the present embodiment,
the term "grid pattern" means that vertically arranged detectors 21
and horizontally arranged detectors 21 are orthogonal with each
other as shown in FIG. 17. When the example shown in FIG. 17 is
taken into consideration, a set of a plurality of rectangular
pixels is also counted as a detector.
[0078] As described above, the radiation imaging apparatus
according to the present invention comprises one or more detectors
that read, as a detection signal, the position information about
radiation incident on a detector surface that corresponds to the
pixel positions and areas of individual image pixels arranged in a
grid pattern; a radiation measurement circuit for reading incident
radiation information; and a collimator in which a plurality of
rectangular through-holes are arranged in a grid pattern and
separated by septa. The radiation imaging apparatus acquires the
radiation incidence position information about the one or more
detectors and generates an image from the acquired information. The
collimator, which is included in the radiation imaging apparatus,
is rotated through a predetermined angle in relation to the layout
of the one or more detectors, which correspond to image pixels, as
viewed from above. The radiation imaging apparatus according to the
present invention makes it possible to avoid a moire problem, which
is unique to pixel type detectors, without requiring high
positioning accuracy or manufacturing accuracy unlike a matched
collimator.
[0079] In the development of a semiconductor-detector-based nuclear
medicine diagnosis apparatus, moire patterns make it difficult to
obtain practical images with small pixels on the order of 1 mm. The
present invention clears a moire problem while considering actual
collimator manufacture, and provides a practical radiation imaging
apparatus and nuclear medicine diagnosis apparatus that use a
semiconductor or pixel type scintillator.
* * * * *