Gamma ray collimator

Details Gamma ray collimator Gamma ray collimator

1998-08-05

2001-08-07

Hannaher, Constantine (Department: 2878)

Radiant energy

Invisible radiant energy responsive electric signalling

With or including a luminophor

C250S363040, C250S363100, C250S503100, C250S505100

Reexamination Certificate

active

06271524

ABSTRACT:

FIELD OF INVENTION
The present invention relates generally to gamma ray detectors and specifically to gamma ray collimation in nuclear medicine.
BACKGROUND OF THE INVENTION
Gamma ray imaging is currently used in medicine to obtain 3D images of patients' internal organs. Positron Emission Tomography (PET) is a medical gamma ray imaging technique frequently used for this purpose.
FIGS. 2A
,
2
B and
3
show representative prior art systems. Prior to an imaging procedure, a patient is given a radiopharmaceutical, which contains a positron emitting substance and which is selectively accumulated in a region of interest. When a positron emitted by the radiopharmaceutical encounters an electron, the electron-positron pair annihilates, emitting two gamma photons of 511 keV each, flying in opposite directions. The simultaneous detection of these two 511 keV gamma photons by two gamma detectors
40
positioned opposite to each other (as shown in FIG.
1
), indicates that a positron has been emitted and annihilated inside an organ of a patient
500
. The simultaneous attribution of 2D coordinates to each one of the photons allows for the determination of the photon's line of flight. The position of the annihilation is along this line. When a multitude of gamma photon pairs are detected and the information processed using appropriate algorithms, electronic circuitry, software, etc., a 3D image of the organ under examination is reconstructed.
Further and more detailed descriptions and analysis of PET will be found in “Performance Parameters of a Positron Imaging Camera”, by Gerd Muehllehner et al., IEEE Transactions of Nuclear Science, Volume NS-23, No. 1, February 1976 and in “Performance Parameters of a Longitudinal Tomographic Positron Imaging System,” by Paans et al., Nuclear Instruments and Methods, Volume 192, Nos. 2, 3, pages 491 -500, Feb. 1, 1982, the disclosures of which are incorporated herein by reference.
Low energy, stray, gamma photons, resulting from 511 keV gamma photons scattered within patient's body, are also present during coincidence measurements and may also reach one or both detectors. These scattered low energy gamma photons do not contain any usable and/or valid information. If these stray gamma photons are allowed to reach the detectors, they increase the count rate at the detector while not adding any usable information. These additional counts, while they may be rejected later, reduce the ability of the detector to detect “real” events, at a high rate.
Another problem encountered in coincidence gamma imaging concerns attenuation artifacts caused by absorption by the patient body and scattering. In order to correct for these effects, a 3D distribution of patient's absorption is preferably previously measured.
Attenuation may be measured (see
FIG. 2A
) by scanning patient
500
using a collimated line source
98
situated opposite a collimated detector
40
or two collimated line sources
98
, opposite two collimated perpendicular detectors
40
and
40
′. When attenuation and coincidence measurements are to be performed consecutively, the configuration of the apparatus has to be changed. This procedure is very time consuming and cumbersome for the following reasons:
a) During coincidence measurements, detectors
40
are positioned parallel to each other (see FIG.
2
B).
b) In order to improve the resolution of the attenuation measurements, a collimator
54
is used on the detector side (see FIG.
2
A). Coincidence measurements use no collimators on detectors (see FIG.
2
B);
c) A “Filter”
56
(see FIGS.
2
B), used in coincidence measurements contains a graded absorber
58
that selectively absorbs, and thus, protects the detectors from large flux of low energy, scattered, stray gamma photons. A line source
98
used in attenuation measurements is, for practical reasons, a source emitting low energy gamma photons (e.g., 100 keV Gd 153). These gamma photons cannot penetrate graded absorber
58
.
Another class of problems concerns the conditioning (collimation) of the radiation from a line source in transmission attenuation measurements. Reference is now made to
FIGS. 4A-4C
. Collimation of line source radiation is performed in one of the following ways:
if a collimation width
62
is larger than source
98
diameter
66
, no substantial collimation exists (see FIG.
4
A);
if collimation width
62
is substantially the same as line source diameter
66
, sensitivity related to manufacturing tolerances is maximal and non-uniform radiation is generated (see FIG.
4
B); and
if line source diameter
66
is larger than collimation width
62
, and a loss of potential radioactivity
68
results. The smaller the ratio of the width of the slit to its length in the direction of the rays, the better the collimation and the greater the loss.
Yet another problem present in coincidence measurements concerns the lack of depth discrimination due to the finite thickness of the scintillation crystal.
Reference is now made to
FIGS. 5A and 5B
. In coincidence measurements, a true
68
or a calculated
68
′ line of flight of gamma photons
70
is determined by the location of a pair of interaction points
72
, of both photons in a pair
76
, with detectors scintillation crystals
42
. The resolution of a detector in coincidence measurements depends on:
a) The detector intrinsic resolution, i.e., the ability of the detector to accurately determine location
72
of interaction of a gamma photon
70
, with scintillation crystal
42
. The thicker the crystal, the higher the probability that the photon interacts with the crystal. However, as is evident from
FIGS. 5A and 5B
, the detector intrinsic resolution is reduced with increasing thickness. In the absence of accuracy in depth discrimination, gamma photons
70
are assumed to interact with the scintillation crystal at its median
84
;
b) The accuracy with which the gantry position is determined; and
c) Loss of resolution due to reconstruction algorithms.
Of these three causes, the loss of resolution in depth discrimination, (shown as X on FIG.
5
B), which strongly depends on incident angle
86
and is a function of crystal's thickness
80
, is most important. In order to increase depth discrimination in coincidence measurements, either the crystal thickness is reduced, or only those photons that have an angle of incidence
86
, under a certain limit are counted, for example, by reducing the flux of photons with large angle of incidence.
“Septa” or “Filter” shields
56
(see
FIGS. 2B
,
3
and
6
) have no substantial localization function per se. They only remove scattered gamma photons
60
, with large axial incidence angle
86
, most of which are not useful for PET and are rejected by the software. To provide this function, the septa are generally about 1 cm apart and have an acceptance angle of about 10 degrees. Prior art septa are placed parallel to the slices of the reconstructed 3D image. The limited collimation of the Septa indirectly improves resolution by reducing the effect of the lack of depth discrimination on location accuracy
It is desirable, in PET, to improve gamma detectors efficiency by reducing the number of stray photons detected relative to the number of non-stray photons detected and to improve the depth discrimination in coincidence measurements without having to reduce the scintillation crystals thickness. It is also desirable to perform attenuation and coincidence measurements in sequence without moving or replacing parts of the imaging system and, in attenuation measurements, to reduce radioactivity losses due to line source diameter while using a large diameter source to improve statistics by increasing the total radiation while keeping the source strictly collimated.
SUMMARY OF THE INVENTION
It is an object of some preferred embodiments of the invention to improve Nuclear Medicine (NM) image quality in coincidence measurements such as in PET imaging.
It is an object of some preferred embodiments of the invention to improve NM detector efficiency in coincidence me

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