This invention relates to a radiation detector for X-rays, .gamma.-rays, etc., and particularly to a radiation detector for use in X-ray CT (computed tomography), a position camera, etc.
X-ray CT includes various types and requires an array of 20 to 1,000 X-ray detecting elements.
Heretofore, a gas chamber filled with xenon, a combination of BGO single crystal (bismuth germanate) and a photomultiplier, and a combination of CsI : T1 single crystal or CdWO.sub.4 single crystal and a photodiode have been used for the radiation detector for use in X-ray CT, etc. The xenon chamber requries a thick window owing to the sealing of a high pressure xenon gas in the detector, and its X-ray absorbance is as low as about 50% because two collimators must be provided per element. In the case of BGO, the luminescence efficiency is low, e.g. about 1%, and thus must be used together with a photomultiplier. That is, the photomultiplier and its accessory high voltage power source are requries and consequently the number of elements in an array is limited.
The CsI : T1 single crystal has a high efficiency, but is deliquescent and has an after-glow (phenomenon of luminescence after the X-ray has been turned off). Thus, it still has a practical problem.
The CdWO.sub.4 crystal has a low luminescence efficiency, and has such problems as easy cleavage at cutting and a toxicity.
The common drawback of the foregoing single crystal scintillator is a fluctuation of luminescence properties in the single crystal. Generally, a single crystal is made to grow from a melting solution, and lattice defects are liable to develop in the crystal during the growth process, and an after-glow will often appear. An activator is often added to the scintillator, but it is difficult to distribute it uniformly in the crystal, and thus the crystal has uneven luminescence. These problems mean that it is very difficult to make the characteristics of the individual detectors uniform.
To solve these problems, some of the present inventors have already proposed a radiation detector using phosphor particles as scintillator [Japanese patent applications Kokai (Laid-open) Nos. 56-151,376 and 57-70,476]. To obtain a sectional image in a radiation detector for X-ray CT, the radiation detector is usually 1 to 3 mm wide and about 20 mm long. Thus, the number of phosphor particles in one radiation detector is, for example, about 300,000, though dependent on their particle sizes. The characteristics of the individual phosphor particles may by slightly different from one another, but a fluctuation of the characteristics as a scintillator can be reduced to 1/square root of number of phosphor particles, that is, about 0.01% by using thoroughly mixed phosphor particles as one scintillator. A satisfactory result can be obtained thereby.
This type of radiation will be described, referring to FIG. 1. After transmission through a cover 5 composed of an aluminum film and a light-scattering layer 4, incoming X-ray 1 makes a scintillator particle layer 2 (phosphor particles solidified by polystyrene) emit light. The emitted light passes through a space 7 and a secondary radiation-preventing layer 8 (Pb glass) and reaches a photodiode 3, where the emitted light is converted to an electric current. To efficiently lead the light emitted from the scintillator particle layer to the photodiode, a vessel 6 is entirely coated with a reflecting film. This type of radiation detector can have about 2-fold output power, as compared with a detector based on a combination of a single crystal scintillator and a photodiode, and is very desirable as a detector for a head scanner, but still has some problem, when applied as a detector for a whole body scanner, because the image processing system for the head scanner is different from that for the whole body scanner.
In FIG. 1, the luminescence of the scintillator layer 2 is uniform, but the emitted light undergoes reflections on the reflecting film surrounding the space 7 many times until the light reaches the photodiode 3. Therefore, if the reflecting film is uneven in the reflectivity, no accurate information of the incoming X-ray can be obtained. When many such detectors are used as elements in an array, there are fluctuations in the characteristis among the individual elements. As a result, the image produced by such an array will have a ring-pattern unevenness or artifact.
To solve these problems, a radiation detector of such a structure as shown in FIG. 2 is preferable. That is, a silicon photodiode 3 is provided at the bottom of a vessel 6 made from brass, and a light-reflecting aluminum layer is formed on the inside surface of the container 6. Pb glass 8 is laid on the silicon photodiode 3 in the vessel 6, and a scintillator layer 2 is formed thereon. The Pb glass is provided to cut the fluorescent X-ray emitted from the scintillator.
In the case of a scintillator particle layer in the structure as shown in FIG. 2, the particle layer itself is optically opaque, and it is difficult to efficiently lead the emitted light to the silicon photodiode. Thus, even if a highly efficient Gd.sub.2 O.sub.2 S: Pr, Ce, F scintillator is used, a signal output of a level only substantially equal to that of a detector based on a combination of single crystal scintillator CdWO.sub.4 and silicon photodiode can be obtained. This is another problem. To increase the output level, a transparent signal crystal scintillator of rare earth oxysulfide should be used, but only single crystal having a size of a few square millimeter can be obtained according to the process disclosed in J. Appl. Phys. 42, 3049 (1971).