Scintillation Detector for Three-Dimensionally Measuring the Gamma-Ray Absorption Position and a Positron CT Apparatus Utilizing the Scintillation Detector

This invention relates to a scintillation detector for three-dimensionally measuring the position of gamma-ray absorption in the detector, and a positron computed tomography (CT) apparatus including plural scintillator detectors arranged in the form of a ring for uniformly obtaining high resolution over the region from the center to the periphery of a visual field without degradation of the spatial resolution in the peripheral region of the visual field.

In nuclear medical diagnosis, a drug labeled with a radioactive isotope (hereinafter referred to as an "RI") is ingested by or injected into a patient, and gamma-rays emitted from the RI in the patient's body are detected outside of the body, thereby providing the distribution of the RI within the body and enabling diagnosis of metabolic functions of internal organs of the body. It is important that the drug labeled with the RI accurately represents the metabolic functions of particular internal organs in the body. For example, where a particular RI is not an inherent element in a living body, a drug labeled with that RI causes metabolic activity and metabolic processes to be significantly changed in comparison with a drug which is not labeled, so that subsequent search for or investigation of the metabolic function is frequently difficult.

A living body is composed of elements including H, C, N and O. These elements also contribute to the metabolism of the living body. A drug labeled with a positron emitter such as .sup.11 C, .sup.13 N, .sup.15 O, or the like, can accurately represent the metabolic function because substances participating in the metabolism, such as C, N, O and the like, are merely replaced by their isotopes and as a result no change is made in chemical structure. Therefore, positron CT for detecting positron nuclides is useful in the field of nuclear medical diagnosis.

FIG. 9 shows a conventional gamma-ray detector employed in a conventional positron CT apparatus. FIG. 9(A) shows a gamma-ray detector comprising a number of scintillators 1 arranged in the form of a ring and photodetectors 2 connected to the scintillators at a ratio of one to one. FIG. 9(B) shows a detector in which scintillators 1 are connected to photodetectors 2 at a ratio of N scintillators to one photodetector, where N is an integer. FIG. 9(C) shows a gamma-ray detector comprising cylindrical scintillators 1 and photodetectors 2 connected to both sides or ends of the scintillators.

In a positron CT apparatus having a gamma-ray detector as described above, two of the scintillators arranged in the form of a ring simultaneously detect two gamma-rays propagating in opposite directions which are radiated when positrons emitted from the RI combine with electrons and are thereby annihilated. A distribution of RIs on a circular plane (hereinafter referred to as "as visual field") defined by the scintillators arranged in the form of a ring can then be observed.

The positron CT apparatus as described above has disadvantage in that spatial resolution is degraded in a peripheral region of the visual field of the apparatus. Moreover, this disadvantage becomes more pronounced as individual scintillators are designed to be narrower to increase their resolving power and longer to increase gamma-ray detection efficiency and sensitivity. Such a problem is caused by the structure and arrangement of the scintillators.

In order to overcome this disadvantage, the present invention provides a positron CT apparatus with a new type of scintillation detector for detecting gamma-rays and three-dimensionally measuring the position of gamma-ray absorption therein, to thereby reduce or prevent the degradation of the spatial resolution in the peripheral region of the visual field and display a positron distribution uniformly over the entire visual field with high resolution.

The scintillation detector according to the present invention comprises a scintillator bundle comprising plural scintillator elements made up into a bundle for emitting scintillation light incident to absorption of gamma-rays in the scintillator elements, each of the scintillator elements having two connecting end surfaces for outputting the scintillation light and side surfaces for leading the scintillation light to the connecting end surfaces and a position detecting photodetector coupled to the connecting end surfaces of each of the scintillator elements, for receiving the scintillation light from the connecting end surfaces and outputting the electrical signals corresponding to the scintillation light, and a positron CT apparatus according to this invention comprises a plurality of scintillation detectors arranged in the form of a ring, each of the scintillation detectors comprising a bundle of plural pillar-shaped scintillator elements and a position detecting photodetector connected to each end of the scintillator bundle.

The accompanying drawings, which are incorporated in and constitute a part of this specification, illustrate a preferred embodiment of the invention and, together with the description, serve to explain the principles of the invention. In the drawings:

FIG. 1 shows an embodiment of a scintillation detector according to the present invention;

FIG. 2 shows a scintillation detector in which plural incisions are made in the direction parallel to a connecting surface of each scintillator element which contacts a photodetector;

FIG. 3 is a graph showing an effect obtained by providing a roughly-ground portion on each scintillator element;

FIG. 4 shows an embodiment of a positron CT apparatus using the scintillation detector according to the present invention;

FIG. 5 illustrates a procedure for determining a gamma-ray absorption position in each coordinate of a scintillator element, in which FIG. 5(A) is a perspective view of scintillator bundle, FIG. 5(B) is a plane view of the bundle, and FIG. 5(C) is a sectional view of the bundle;

FIG. 6 is an explanatory diagram for determining a gamma ray absorption position in the Z-coordinate;

FIG. 7 shows an embodiment of a signal read-out circuit in the detector according to the present invention;

FIG. 8 shows a circuit arrangement of a positron CT apparatus according to the present invention;

FIG. 9 shows a conventional positron CT apparatus, in which FIG. 9(A) shows a scintillation detector comprising scintillators and photodetectors connected to the scintillators in the ratio of one to one, FIG. 9(B) shows a scintillation detector in which scintillators are connected to photodetectors in the ratio of N scintillators to one photodetector, and FIG. 9(C) shows a ring-shaped scintillation detector comprising cylindrical scintillators and photodetectors connected to both sides or ends of the respective cylindrical scintillators;

FIG. 10 shows a coincidence sensitivity for each pair of detectors on a line X--X'; distribution for and

FIG. 11 shows the relationship of resolving power with the distance from the center of the ring-shaped scintillation detector.

FIG. 10 shows a coincidence sensitivity distribution of each pair of detectors on a line X--X' which passes through the center of the detector ring. In FIG. 10, a line connecting a pair of detectors (A and A') passes through the center of the visual field and a line connecting another pair of detectors (B and B') passes through the peripheral region of the visual field. C.sub.1 and C.sub.2 represent coincidence response functions in the center and the peripheral regions of the visual field respectively.

As shown in FIG. 10, when a point-type gamma-ray source is shifted along the line X--X' and detector pairs (A--A' and B--B') are used to coincidentally count the gamma-rays, response functions, that is, coincidence sensitivity distributions, of respective detector pairs which are designated by C.sub.1 for detector pair A--A' and C.sub.2 for detector pair B--B', are obtained.

The width of the response function approximately represents a positional or spatial resolution in the radial direction. The response function C.sub.1 in the center of the visual field which is detected by detector pair A--A' has a symmetrical triangular distribution and its spatial resolution approximately corresponds to a half width of a scintillator.

The response function C.sub.1 is not dependent on the position (in the depth-direction of a scintillator) where gamma-rays are absorbed in the scintillator. On the other hand, the response function C.sub.2 in the peripheral region of the visual field has a surprisingly broad distribution, and consequently the spatial resolution is degraded.

FIG. 11 shows the relationship between positional or spatial resolution in the radial direction and the distance from the center of the ring-shaped detector where the diameter of the ring, and the width and the length of each scintillator are 540 mm, 5 mm and 30 mm, respectively. As is apparent in FIG. 11, the spatial resolution is degraded more significantly as the detecting position approaches the peripheral region of the visual field.

A positron CT apparatus according to the present invention includes a plurality of scintillation detectors arranged in the form of a ring, each of which comprises a scintillator bundle formed by binding pillar-shaped scintillator elements into a bundle. A position detecting photodetector is connected to each end of the scintillator. Accordingly, three-dimensional position information of the gamma-ray absorption in the scintillator can be obtained, and a positron emitter distribution with high resolution in the peripheral region of the visual field as well as in the center region thereof can be obtained. Additionally, one can simultaneously obtain a number of sliced or cross-sectional images.

FIG. 1 shows an embodiment of a scintillation detector according to the present invention. In FIG. 1, reference numeral 11 represents a position detecting photodetector, 12 represents a scintillator bundle, 12.sub.1, 12.sub.2 and 12.sub.3 represent scintillator elements, and 13 represents incident gamma-rays.

In FIG. 1, each of the scintillator elements 12.sub.1, 12.sub.3 and 12.sub.3 is in the form of a pillar and comprises, for example, bismuth germanium oxide(BGO). Each of the pillar-shaped scintillator elements has two end surfaces, each of which can be mated with or connected to a receiving surface of photodetector. Each scintillator also has at least one side wall or surface typically substantially perpendicular to the end surface, through which gamma-rays may enter, as shown in FIG. 1. Scintillator bundle 12 is formed by binding the scintillator elements into a bundle. The end surfaces of each of the scintillator elements are mirror polished. It is preferable to totally reflect scintillation light at the inner surface of the side walls of each scintillation element and to reflect the light to one or both end surfaces thereof. This may be accomplished by coating any surfaces other than the mirror polished end surfaces with a reflecting material such as barium sulphate(BaSO.sub.4) or the like, or by providing spacing or an air gap between respective neighboring scintillator elements, or by combining both the coating and the inter-element spacing. Both ends of the scintillator bundle 12 thus formed are connected to a two-dimensional position detecting type of photomultiplier having, for example, a crossed multiwire-anode. The coupling of the scintillator bundle to the two-dimensional photodetector may be formed directly or through a light guide.

To improve the accuracy in position detection, roughly-ground portions may be provided on any surface (side surface) of each scintillator element other than connecting surfaces (end surfaces). Preferably, plural incisions 14 are formed on the side surfaces of each scintillator element parallel to the connecting end surfaces. FIG. 2 shows a detector having plural incisions 14 formed on each scintillator element in the direction parallel to the connecting surface of the scintillator element.

FIG. 3 is a graphical representation which illustrates an effect obtained by forming a roughly-ground portion on each scintillator element. In FIG. 3, the abscissa of the graph represents a gamma-ray absorbing position and the ordinate represents an output intensity from an end of a scintillator element (3.times.5.times.50 mm).

The curve (a) represents an output intensity of a scintillator element having mirror polished side surfaces coated with a reflecting material. Curve (b) represents a scintillator element having roughly ground side surfaces coated with the reflecting material.

As is apparent in FIG. 3, the inclination of the output intensity distribution of a scintillator element can be made steeper by roughly grinding the side surfaces of the scintillator element, thus the accuracy of detecting position can be improved. The scintillation detector thus formed is arranged in the form of a ring, as shown in FIG. 4, and employed as a gamma-ray detector in a positron CT apparatus.

FIGS. 5 and 6 are explanatory diagrams for explaining the principle of the detector according to this invention. Scintillation light is emitted from both ends of a scintillator element when gamma-rays are absorbed in the scintillator element. The scintillation light thus emitted is detected by position detecting photodetectors connected to each scintillator bundle. On the basis of the output of the photodetectors, an incident position of the gamma-ray is calculated to determine the coordinate position (X, Y).

FIG. 6 shows the relationship between the output intensity of an output A,B with the position of the gamma-ray absorption (along the z-axis). The position of gamma-ray absorption along the Z-coordinate is obtained by comparing the output intensity of the outputs A with that of output B, as shown in FIG. 6.

As the position of gamma-ray absorption in the y-direction of a scintillator can be determined as described above, high spatial resolution can be obtained in a peripheral region of a visual field as well as in the center region of a visual field. Further, detecting units can be densely arranged to obtain high sensitivity, and spatial information in the z-direction can be obtained, thus, simultaneously obtaining multi-sliced tomographic images.

FIG. 7 shows an embodiment of a signal read-out circuit in a position detecting photodetector according to the present invention. In FIG. 7, reference numerals 21 and 22 represent wire anodes of two position detecting photomultipliers connected to each end of a scintillator bundle; 23 and 24 are respective dynodes of the photomultipliers; 25, 26 and 27 are buffer amplifiers; 28, 29 and 30 are x-, y- and z-position calculating circuits, respectively; and 31 is an adder.

As shown in FIG. 7, respective outputs detected by the wire anodes are applied through the buffer amplifiers 25 and 26 to the x- and y-position calculating circuits 28 and 29 and subjected to respective positioning calculations therein. X-address and y-address signals are the outputted from the x-position and y-position calculating circuits, respectively. Furthermore, signals from the dynodes 23 and 24 are independently applied to the buffer amplifiers 27 and the outputs thus amplified are applied to z-position calculating circuit 30 and adder 31 so that a z-address signal is outputted from z-position calculating circuit 30, and an energy signal and a timing signal are outputted from adder 31.

FIG. 8 shows a circuit arrangement of a positron CT apparatus employing the scintillation detector according to the present invention. Reference numerals 41 and 42 represent detecting units for detecting a pair of gamma-rays propagating in opposite directions from each other; 43 and 44 are data selectors; 45 is a coincidence counting circuit; 46 is data bus; 47 is a main memory unit; 48 is a data processing unit; and 49 is an auxiliary memory unit.

When the two detecting units 41 and 42 of the plurality of detecting units arranged in the form of a ring detect a pair of gamma-rays generated by the annihilation of positrons, the outputs from the two detecting units are analyzed by the coincidence counting circuit 45 (using a timing signal) to determine whether or not the outputs were received simultaneously with each other within a given period of time. If the outputs are found to be simultaneously counted, respective address signals (x, y, z) are selected by data selectors 43 and 44 and provided to data bus 46. At this time, unit addresses pointing to the scintillation detectors which detect the gamma rays also are provided to data bus 46. The data included in the address signals and the unit addresses are stored in main memory unit 47 or auxiliary memory unit 49, and are then subjected to ordering and pre-processing operations in data processing unit 48 to reconstruct a corresponding image.

As described above, the scintillation detector of the present invention which comprises a bundle of pillar-shaped scintillator elements generates three-dimensional information on the gamma-ray absorption position. By providing a positron CT apparatus with plural scintillation detectors as described above and utilizing the three-dimensional information obtained by each scintillation detector, spatial resolution in the peripheral region of the visual field can be improved and high sensitivity can be obtained uniformly, so that a position CT apparatus having simple construction and high spatial resolution is obtained.

Furthermore, a positron CT apparatus employing the scintillation detector according to this invention has the advantages that multi sliced tomographic images can be simultaneously obtained and high sensitivity can be obtained by the dense arrangement of the detectors.

Additional advantages and modifications will readily occur to those skilled in the art. Therefore, the invention in its broader aspects is not limited to the specific details, representative devices, and illustrative examples shown and described. Accordingly, departures may be made from such details without departing from the spirit or scope of the general inventive concept as defined by the appended claims and their equivalents.