Radiant energy – Invisible radiant energy responsive electric signalling – Semiconductor system
Reexamination Certificate
2001-04-16
2003-02-04
Hannaher, Constantine (Department: 2878)
Radiant energy
Invisible radiant energy responsive electric signalling
Semiconductor system
C250S370060
Reexamination Certificate
active
06515286
ABSTRACT:
BACKGROUND OF THE INVENTION
1. Field of the Invention
The present invention relates to a radiation detector in which radiographic image information is recorded by converting radiation to fluorescent light in a scintillator, and detecting the fluorescent light by photoelectric converters so that the recorded radiographic image information can be read out.
2. Description of the Related Art
Currently, various radiographic image record-and-readout apparatuses using a solid-state radiation detector are proposed and practically used in radiography in the field of medical diagnosis or the like. In the solid-state radiation detector, charges are generated and temporarily stored in charge storing portions of solid-state detector elements when radiation is detected. Thereafter, the stored charges are converted to an electric signal representing radiographic image information, and then the electric signal is output. In addition, various types of solid-state radiation detectors have been proposed for use in the above image record-and-readout apparatuses. When the solid-state radiation detectors are classified by charge generation process, the so-called optical-conversion type solid-state radiation detectors are known, for example, as disclosed in Japanese Unexamined Patent Publication Nos. 59(1984)-211263, 2(1990)-164067, PCT International Publication No. WO92/06501, and Larry E. Antonuk et al., “Signal, noise, and readout considerations in the development of amorphous silicon photodiode arrays for radiotherapy and diagnostic x-ray imaging,” SPIE Proceedings Vol.1443 (“Medical Imaging V: Image Physics”) 1991, pp. 108-119. In the optical-conversion type solid-state radiation detectors, a scintillator is exposed to radiation, and converts the radiation to fluorescent light, photoelectric conversion elements (photodiodes) detect the fluorescent light, and generate signal charges the amounts of which corresponds to the intensities of the fluorescent light (i.e., the intensities of the radiation) at the locations of the photoelectric conversion elements, and capacitor elements respectively connected to the photoelectric conversion elements store the signal charges, where the photoelectric conversion elements are made of a semiconductor such as silicon and selenium. In order to obtain an electric signal (image signal) representing the amounts of the signal charges stored in the capacitor elements, the so-called TFT readout method is used. According to the TFT readout method, switches, e.g., thin-film transistors (TFTs), are arranged at midpoints of signal lines respectively connected to the above capacitor elements, and the switches are sequentially driven in a scanning order.
When radiation is detected by a solid-state radiation detector which is formed with a combination of a scintillator and photoelectric conversion elements as described above, the scintillator is required to have high emission efficiency in order to reduce an exposure dose in a patient. In addition, the radiation absorption in the scintillator is required to be great in order to reduce quantization noise and achieve emission of a great amount of fluorescent light, although generally, transmittance of radiation through material is high, and radiation produces quantization noise. Further, the scintillator is required to have a wavelength-light emission characteristic which matches well to the wavelength-sensitivity (spectroscopic sensitivity) characteristic of the photoelectric conversion elements.
When the above solid-state radiation detector of the optical conversion type is used in applications in which a highly sharp image is required, e.g., in medical X ray imaging, usually, the photoelectric conversion elements are arranged at a pixel pitch of 50 to 200 micrometers, where the pixel pitch corresponds to the pixel size.
On the other hand, since layers which constitute each photoelectric conversion element have a dielectric property, the capacitor elements are usually realized by the photoelectric conversion elements per se. Therefore, the capacitance of each capacitor element is determined by the pixel pitch, and as small as 0.5 to 2 pF. Further, the maximum storable charge amount of each capacitor element is also small.
In addition, in order to increase the radiation absorption in the scintillator, usually, the thickness of the scintillator is increased, and scintillators having a thickness of about 500 micrometers are widely used.
However, when the thickness of the scintillator is increased, the amount of the fluorescent light emitted from the scintillator increases, and the amounts of charges generated in the photoelectric conversion elements also increase. As a result, it is probable that the amount of charges generated in each photoelectric conversion element exceeds the maximum storable charge amount of each capacitor element, i.e., the solid-state radiation detector is saturated. Therefore, X ray imaging can be performed only with a low radiation dose. In other words, it is impossible to secure a sufficient dynamic range of the radiation dose.
For example, even when solid-state radiation detectors include a CsI:Tl scintillator and a photoelectric conversion element containing Si as a main component and having a wavelength-sensitivity characteristic matched with the wavelength-light emission characteristic of the CsI:Tl scintillator, or a CsI:Na scintillator and a photoelectric conversion element containing Se as a main component and having a wavelength-sensitivity characteristic matched with the wavelength-light emission characteristic of the CsI:Na scintillator, the solid-state radiation detectors have a dynamic range as narrow as 7 to 10 mR, which corresponds to the saturation limit.
SUMMARY OF THE INVENTION
An object of the present invention is to provide a radiation detector which realizes a sufficiently large dynamic range of a radiation dose even when the amount of radiation absorption in a scintillator is increased in order to reduce quantization noise.
In order to achieve the above object, according to the present invention, the amounts of charges generated by photoelectric conversion elements are reduced while maintaining great radiation absorption in the scintillator.
Specifically, according to the first aspect of the present invention, there is provided a radiation detector comprising a scintillator, a plurality of photoelectric conversion elements, and a plurality of capacitor elements. The scintillator is made of CsI:Tl (CsI doped with Tl) receives radiation corresponding to a number X of radiation quantums for each of a plurality of pixels, and converts the radiation to fluorescent light so that a number L of photons which constitute the fluorescent light are emitted in response to each radiation quantum. Each of the plurality of photoelectric conversion elements is provided for one of the plurality of pixels, contains Si as a main component, detects the fluorescent light, and generates charges when the fluorescent light is detected, where each of the plurality of photoelectric conversion elements has a fill factor F and a photoelectric conversion efficiency &eegr;, and is arranged so that the fluorescent light enters each of the plurality of photoelectric conversion elements with an entrance efficiency T. Each of the plurality of capacitor elements is connected to one of the plurality of photoelectric conversion elements, stores the charges generated by the one of the plurality of photoelectric conversion elements, and has a maximum storable charge amount Q. When the radiation detector receives a 10 to 300 mR dose of the radiation, the number X of radiation quantums, the number L of photons of the fluorescent light, the entrance efficiency T, the fill factor F, and the photoelectric conversion efficiency &eegr; satisfy a relationship X·L·T·F·&eegr;≦Q.
The left side of the above inequality corresponds to the amount of the charges generated by each photoelectric conversion element when the radiation detector is exposed to a radiation dose determined by the above number X. Therefore, the above relat
Fuji Photo Film Co. , Ltd.
Gabor Otilia
Hannaher Constantine
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