COMPUTERIZED TOMOGRAPHY APPARATUS

Abstract: A computerized tomography apparatus, in a normal scanning mode, senses dose of radiation projected from a radiation source by a plurality of radiation detecting cells (3a to 3d), integrates the output signals from said detecting cells (5a to 5d) by said integrating circuits (6a to 6d), selects sequentially the integrated dose signals from said integrat­ ing circuits (6a to 6d), and supplies the selected dose signals to a computer system. In a high speed scanning mode, the computerized tomography apparatus closes an analog switch (10a) and opens another analog switch (11) to compose the radiation doses sensed by a plurality of radiation detecting cells, and integrates and multiplexes the composite ones as radiation dose outputs from combina­ tions each including a plurality of channels.

Description: Computerized tomography apparatus The present invention relates to a computerized tomography apparatus improved in that reduction of detected radiation data scanned at a high speed is realized by reducing both the number of views and rays. A conventional computerized tomography apparatus (referred to as a CT scanner) is made using radiations such as X-rays, an X-ray source 1 for projecting X-ray Fx in the form of a flat fan beam and a multichannel X-rays detector 2 comprising a plurality of radiation detecting elements 3 for detecting said X-rays are oppositely disposed with an object under diagnosis interposed therebetween, as shown in Figs. 1 and 2. For acquiring X-ray data of the cross sections of the object in every direction, the X-ray source 1 and the X-ray detector 2 are rotated about the object 4 in the same direction and at the same angular speed. After a satisfactory amount of the data is acquired, a digital computer calculates absorption of the penetrating X-rays at individual positions on the cross sections of the object 40. The absorptions of the cross section of the object 4 are used for providing a gradation of a tomogram to be formed. In this way, a clear tomogram is formed of every tissue from soft tissue to hard tissue. The X-ray detector 2 is made up of a number of radiation detecting cells forming an ionization chamber filled with high pressure gas such as Xe gas or photodiode plus scintillator like a cadmium tungstate scintillation (CdWO4) crystal. The X-ray detector 2 thus constructed detects the energy of the X-ray transmitted through the cross section of the object 4 to provide ionization current as X-ray projection data. In particular, in acquiring the X-ray projection data, X-ray photons passing through paths connecting the radiation detecting cells 3 of the ionization chamber to the X-ray source 1 (this path is called an "X-ray path") hit the high pressure gas filled in the ionization chamber to ionize the gas. The ionized charge thus produced is extracted. In this way, the energy of the X-ray is detected in the form of ionization current. The ionization current is integrated for a given time, the result of the integration is subjected to A/D conversion. The result of the A/D conversion is the X-ray projection data on the X-ray path. After the data on all the X-ray paths formed at an angular position when the X-ray radiated from the X-ray source transmits through the object, is acquired, the data acquisition on the X-ray paths formed at the next angular position is performed. Similarly, the data acquisition is performed at the subsequent angular positions. For signal processing on the data acquisition, as shown in Fig. 3, a number of radiation detecting cells 3a to 3n are respectively coupled with amplifiers Sa to 5n for amplifying the output signals from the detecting cells 3a to 3n and for applying them to integrators 6a to 6n. X-ray photons arriving at the detecting cells 3a to 3n are detected by these cells and amplified by the amplifiers 5a to 5n to proper values, and then integrated by integrators 6a to 6n, respectively. The output signals from the integrators 6a to 6n are selectively extracted in a sequential manner and are applied to a dynamic range adaptor 8 for dynamic range adjustment. The dynamic range adjusted signal is then applied to an A/D converter 9 where it is converted into a digital signal. Thus digitized X-ray projection data for each channel are applied to a host system such as digital computer. Incidentally, the dynamic range adaptor is not essential to this signal processing system of the CT scanner under discussion. This CT scanner is called a third generation CT scanner. In addition to the third generation CT scanner, there are first, second and fourth generation CT scanners, although not shown. In the first generation CT scanner, an X ray source for radiating X rays in pencil beam fashion and an X-ray detector using a single radiation sensor are oppositely disposed. In operation, the combination. of the detecting cell and the X-ray source are translated. After one translation is completed, the X-ray projection angle is changed with respect to the object, or rotated about the object, and then the translation of the cell and X-ray combination is repeated. The first generation CT scanner is of translate/rotate type. In the second generation CT scanner, an X-ray source for projecting a relatively narrow X-ray beam in a fan shape is used in place of the pencil beam radiation X-ray source of the first generation CT scanner. The X-ray detector uses a plurality of radiation sensing cells for covering the width of the radiated fan beam. The scanning operation by the CT scanner is similar to that of the first generation one. In the fourth generation CT scanner, a number of radiation sensing cells are disposed in a circle. In taking tomograms of the object or acquiring X-ray projection data, the X-ray source is rotatably moved radiating the fan X-ray beam, while the detector is stationary. The resolution of a reproduced picture depends on how fine the picture elements are. However, improvement of the resolution of the reproduced picture can not be expected unless the pitch for sampling the X-ray projection data is also fine. To obtain a CT scanner with high resolution, the number of channels of the X-ray detector and of the data acquiring system must be increased. The same thing is not applied for the first and second generation CT scanners because these scanners are accompanied by a translation movement. In these CT scanners, however, the data acquiring speed is quite slow and the CT scanners have a problem of long data acquiring time. It is safe to say that the third and fourth CT scanners with high resolution have a large number of channels. The reason for this is as follows. In the third generation CT scanner, the scanning depends only on the rotational movement. Further, a relationship between the number of the radiation detecting cells 3 and an angular scope of the fan beam is fixed and it is impossible to change the sampling pitch for the object. Therefore, for making the sampling pitch finer, the pitch of the channels must be made finer, or the number of the detecting cells must be increased. In the fourth generation CT scanner, the X-ray detecting elements 3 are circularly arranged and only the X-ray source is rotated. The sampling pitch is controllable by controlling the rotation angle for each sampling. Even if the sampling pitch is made fine to increase the number of views, a spatial frequency response can be extended up to high frequencies if the width of an opening of the radiation detecting cell is wide. In this respect, the size of the opening of the detecting cell must be small. For covering the expanded width of the fan beam due to the discrete arrangement of the tiny radiation detecting cells, the number of the channels is inevitably increased. In a high resolution CT scanner, the number channels and the views (projections) as well are large. As is well known, to reproduce exactly the high spatial frequencies of the object, the number of the views must be large. Typical examples of the relationship of the number of channels to the number of views at the present time are: Third generation : 515 channels, 400 views/4.5 sec. scan or 600 views/9 sec. scan Fourth generation: 1088 channels, 1088 views/3 sec. scan, 124 rays/view The amount of data acquired by the fourth generation CT scanner is generally larger than that of the third generation. In the above example, the total number of data is 1.3 x 106, with the line for each data (one ray) as 2.2 ps. In the f#ourth generation CT scanner, a cutoff frequency of MFT (modulation transfer function) which means a high spatial frequency of the reproduced picture to an original picture is approximately 8 lp (line pair)/cm which means that a spatial frequency is eight pairs of black and white bars per 1 cm. When it is desired to expand the cutoff frequency to 40 lp/cm, for example, which is five times the above, that is, the cutoff frequency of the general tomographing, it is necessary that the number of channels of the X-ray detector must be about five times, or 5440, and the number of rays is also about five times, or 6,205 rays/view. As a result, a total number of data is 3.38 x 107. With such an amount of data, it is very difficult to apply high speed scanning for that data acquisition, that is, to rapidly construct an image. Nevertheless, there is still strong demand for shortening the scanning time and improving artifacts of image. The prior art CT scanner has a plural#ity of opera tion modes for dealing with this problem. When high speed scanning is required, the number of data is reduced to speed up the acquisition of data. When no high speed scanning is required, a sufficient number of data is acquired to improve the picture quality. For reducing the data in a high speed scanning mode, the third generation CT scanner has reduced the number of views or the number of projections, and the fourth generation CT scanner has reduced the number of rays per view, that is, the number of X-ray paths per view. Both the third and fourth generation CT scanners sample the object under diagnosis every time that the X-ray source rotates at a fixed angle. Accordingly, the reduction of the data may be attained by controlling the rotating angle pitch for projecting. The reduction of the number of views or rays leads to a deterioration in picture quality. When the number of the views is reduced, streak artifacts occurs. For avoiding these artifacts, a band of the convolution filtering process is narrowed to reduce a spatial resolution. The reduction of the rays results in an increase of angular between the sampling pitches. This is clearly related to a deterioration of the spatial resolution. A moving portion or a portion difficult to keep in a motionless state for a long time, must be tomographed at a high speed, even though the spatial resolution is somewhat impaired. The views and the rays are not parameters designed independently. These parameters are given depending on the desired spatial frequency band. For an optimum design, if one of the parameters is reduced, the other must also be reduced; otherwise, data larger than the data necessary for securing a given picture quality is collected, with the result that time is consumed in unnecessary data acquiring and processing time. Although such is the case, the prior CT scanner controls only one of the parameters, as disclosed in U.S. Patent No. 4,075,492. Let us consider at what speed the prior art CT scanner may operate when the data of 3.38 x 107 is necessary for attaining the cutoff frequency 40 lp/cm, as mentioned above. Even if the sampling pitch is made two times larger than the normal sampling pitch, the assumption is that sacrifice in resolution in high speed scanning is unavoidable and the data will reach even -1.69 x 107. In an acquisition time system two bytes of memory capacity for each data are required. For storing such an amount of data, 12 megabytes (MB) of memory capacity is necessary. For a scanning speed of 0.5 sec. at the ordinary sampling pitch, a storing medium which allows data to be written therein at 67.6 MB/sec. is needed. At the present stage of technology, a magnetic disc is most suitable for the data storing medium for storing a large amount of data. A magnetic disc operable at such a high speed is very expensive and also requires an A/D converter and a data transfer system, which are operable at a corresponding high speed. Therefore, time taken for data acquisition and processing must be shortened such that the number of both the views and the rays are reduced and the data are reduced in a well balanced manner. With a shortening in the data acquisition and processing time, the data storing requirement is eased reducing the cost. The third and the fourth generation CT scanners will be further described. Fig. 4 illustrates a third generation CT scanner. As shown, an X-ray source 1 and an X-ray detector 2 are concurrently rotated about a center 0 in a field or data acquiring area. An X-ray radiation of one pulse is performed at a position a of the X-ray source 1 to first acquire the data and then data acquisition of the first view is completed. In this case, the number of the rays is equal to that of the channels. The data acquisition of the second view is performed at an X-ray source position b. In a high speed scanning mode, the rotation angle /aob of one angular shift is larger than the usual one. That is, the number of views in one rotation of the X-ray source is reduced. In this way, the CT scanner copes with the high speed scanning operation. Fig. 5 shows a fourth generation CT scanner. As shown, an X-ray source 1 rotates about the center 0 of a tomograph view, radiating the X-ray fan beam toward the center, a circular detector is motionlessly mounted with just the source rotating. When the X-ray source 1 is located at a position c, the data of the first ray is acquired by a group of radiation detecting cells located within a range of the fan X-ray beam radiated therefrom. Then, the X-ray source 1 moves to a position s. Here, the data on the second ray is acquired by a group of radiation detecting cells located within a range of the fan beam of X-ray radiated therefrom. This data acquisition is repeated up to a position n. When the data acquisition is completed at the position n, one view as seen from the D1 radiation detecting cell is completed. Views of the other radiation cells are similarly obtained. In a high speed scanning mode, the rotation angle /cDld is made large. In this case, only the number of rays is reduced, while the number of cells is fixed. Therefore, the total number of views for all the cells is unchanged. For reducing the number of both rays and views, the prior scanner needs an additional circuit. Accordingly, the prior scanner reduces either the number of rays or views in the high scanning speed mode not only from a cost/performance standpoint but also because the extremely high-speed scanning and extremely high resolution are not intended. The rays and the views must be reduced in a well balanced manner. If only one of the views or the rays is reduced, the data amount is not reduced much, and more adversely superfluous data is included. Accordingly, the prior art CT scanner involves many problems in data write speed required for the data storing medium, the resulting operation speed of the A/D converter and the data transfer system, the data processing speed, etc. Accordingly, an object of the present invention is to provide a computerized tomography apparatus which secures an ordinary high resolution and realizes a reduction of the detected radiation data by reducing the number of both the views and rays. There is provided a computerized tomography apparatus comprising: radiation source means for projecting a radiation beam spreading in a fan shape toward an object; means for relative motion, between said object and source-detector means combination about an axis of rotation; radiation detector means including a plurality of radiation detecting cells for detecting the radiation beam radiated from the radiation source means and transmitted through the object; radiation dose calculating means respectively connected to the plurality of radiation detecting cells in the radiation detector means and for calculating the dose of radiation detected by the detecting cells to provide the calculated values, the integrating means being provided corresponding to a plurality of channels; multiplexer means connected to the plurality of radiation dose calculating means and for sequentially selecting the calculated radiation dose; means connected to the multiplexer means for analog-to-digital converting the detected radiation data selected by and outputted from the multiplexer means and for supplying the converted digital data to a computer system; and switching means provided for composing the doses of the radiation detected by at least two radiation detector elements and supplying a composite radiation dose as a radiation dose output from a combination of a plurality of channels to the multiplier means. With such an arrangement, the computerized tomography apparatus according to the present invention provides a high-resolution reproduced picture at a low speed or normal scanning mode and reduces the amount of the radiation detected data by reducing the number of both the views and the rays. Accordingly, the requirement for the data transfer time is relaxed and the requirement for a data writing speed to a medium for storing data is remarkably relaxed. A necessary number of A/D converters for A/D converting a radiation dose output from each channel, which is determined by the operation speed of the scanner in a high operation- mode, and is reduced to a value corresponding to the reduced number of the channels, resulting in reduction of cost. With the data fetching capability of the system fixed, a high scanning speed can be expected in comparison with the case where the system has no means for composing the outputs of a plurality of detectors. This invention can be more fully understood from the following detailed description when taken in conjunction with the accompanying drawings, in which: Fig. 1 is a schematic illustration of a third generation CT scanner; Fig. 2 is a schematic side view of an object under diagnosis and the radiation detector shown in Fig. 1; Fig. 3 is a circuit diagram of a prior art data acquiring system for acquiring detected radiation dose data with a plurality of channels; Fig. 4 are schematic illustrations for explaining the data acquiring operation by a third generation CT scanner; Fig. 5 shows a schematic illustration of an operation for acquiring detected radiation dose data by a fourth generation CT scanner; ; Figs. 6 to 8 illustrate circuit arrangements of a data acquiring system for acquiring detected radiation dose data with a plurality of channels, which are different embodiments of the invention; Fig. 9 is an equivalent circuit diagram of two channels coupled with each other in the circuit shown in Fig. 8; Figs. 10 and 11 show circuit arrangements of data acquiring system for acquiring detected radiation dose data with a plurality of channels which are other embodiments of the present invention. Preferred embodiments of the present invention will be described referring to Figs. 6 to 11. Fig. 6 shows a circuit diagram of a data collecting system which is an embodiment of the present invention. In Fig. 2, like reference numerals are used for designating like portions and no explanation thereof will be given here. Basically, the circuit arrangement of Fig. 6 is the same as that of Fig. 3. For ease of illustration, the 1st to 4th channels are illustrated and since the output impedance of the radiation sensing cells 3a to 3d are high, these may be expressed as a constant current signal source. Analog switches lOa and lOb are provided which respectively connect the outputs of two pairs of the adjacent radiation sensing cells 3a and 3b of the first and second channels and the radiation sensing cells 3c and 3d of the third and fourth channels in the data sensing system. Analog switches lla and llb are connected in series in the input circuits of amplifiers Sb and 5d, respectively. When the switches lla and llb are closed but the switches 10a and lOb are opened, the detected signals of the two adjacent channels are composited. Namely, in a low speed or normal scanning mode in which a sufficient number of data are obtained to improve a resolution, a controller (not shown) opens the analog switches lla and llb but closes the analog switches 10a and lOb. Accordingly, since the channels are independently operable, the data of all the channels may be collected as in the prior art. In a high speed scanning mode for reducing the number of the data, the controller closes the analog switches lOa and lOb and opens the analog switches lla and llb. As a result, the two pairs of output signals from the radiation detecting cells 3a and 3b and 3c and 3d, each pair of the output signals being produced from two adjacent channels, are respectively applied to the amplifiers 5a to Sc, respectively. Thus, the output signals from the two adjacent channels are combined together. In this case, the controller controls the multiplexer 7 by incrementing it every two channels. That is, the output signals from the integrators are progressively selected in a skipping way in the order of the output signal from the integrator in the first channel, the output of the integrator 6c in the third channel, the output of the integrator in the fifth channel, ... Since the outputs of the radiation detecting cells 3a to 3d have sufficiently high impedance, the output signals from the two radiation detecting cells are composed and the composite one is applied to the single amplifier 6a or 6b. In this way, the number of the radiation detecting cells are reduced nearly to half by controlling only the analog switches 10a, lla, lOb, llb, and the multiplexer 7. Accordingly, the rotation angle the X-ray source 1 for each sampling is allowed to be large. Further, in the third generation the views and the rays as well may be reduced and in the fourth generation the number of the views and the rays as well may be reduced. In this way, the number of data may be held down to a proper value. A circuit arrangement of Fig. 7 is also so designed that the detected output signals of the two channels are composed together, as in the previous case. In this circuit, the two adjacent channels are treated as a pair unit. In the first pair of channels, a series circuit including an analog switch 10a and a resistor 13a is connected between the input of an integrator capacitor 12a of an integrator circuit 6a in one channel and the output of an amplifier 5b in the other channel. In the second pair of channels, a series circuit including an analog switch lOb and a resistor 13c is likewise connected between the input of an integration capacitor 12c of an integration circuit 6c in one channel and the output of an amplifier Sd in the other channel. Resistance of the resistors 13a and 13b Is equal to that of the integration resistors of the integration circuits 6a and 6c. When the analog switch 10a is closed, the output signal from the amplifier 5b in the other channel is applied through the resistor 13a to the input of an operational amplifier of the integrator 6a of one channel. Similarly, when the analog switch lOb is closed, the output signal from the amplifier 5d in the other channel is applied through the resistor 13c to the input of an operational amplifier in one channel. Accordingly, if the output impedance of the amplifiers is sufficiently low, the integration capacitor 12a or 12c of the integration circuit 6a or 6c is charged proportional to the sum of the detected output signals from the radiation sensing cells 3a and 3b or 3c and 3d in the first of second pair of channels, to produce an integrated value proportional to the sum of the output signals for the corresponding pair of channels. Accordingly, if the multiplexer 7 is incremented every two channels by the controller, much the same operation as that of Fig. 6 may be attained. Thus, the normal scanning mode and the high speed scanning mode may be performed by properly controlling the analog switches 10a and lOb and the multiplexer 7. Since an amplifier with a sufficiently low output impedance is available, the present embodiment may easily be realized. Turning now to Fig. 8, there is shown another embodiment of the present invention, which has a somewhat different construction from that of th above two embodiments. In the channels, amplifiers 5a to 5d have at their poststage series circuits which include resistors 14a to 14d of resistance Ra and capacitors 15a to 15d of capacitance Ca, respectively. The capacitors 15 a to 15d are each grounded at on end. The output signals at the junction points between the resistors 14a to 14d and the capacitors 15a and 15d are applied to a multiplexer 7. Junction points between the resistors 14 (14a to 14d) and the capacitors 15 (15a to 15d) in the two adjacent channels are connected by a switch 10 (10a to lOd). For example, in the first pair of the two adjacent channels including the sensing cells 3a and 3b, junction points between the resistor 14a and the capacitor 15a and between the resistor 14b and the capacitor 15b are connected by an analog switch 10a. In the second pair of the channels which includes the sensors 3c and 3d, an analog switch lOb connects the junction points between the resistor 14c and the capacitor 15c and between the resistor 14d and the capacitor lSd. If the output signal of the CR circuit which includes the resistor 14 and the capacitor 15 is read out at sampling periods approximate to (from a fraction to a multiple) a time constant T as the product of the resistance Rb of the resistor 14 and the capacitance Ca of the capacitor C15, the CR circuit functions like an integrator circuit. The electrical performance of the data collecting system of the present embodiment is inferior to that of the above-mentioned embodiments, but its economical feature is superior. In a normal scanning mode, the analog switches lOa and lOb are open and the multiplexer 7 is incremented in a normal manner. In a high speed scanning mode, the analog switches lOa and lOb are closed and a controller (not shown) controls the multiplexer 7 so as to sequencially select the output signals from the channels every two channels. The detected output signals are applied to the corresponding time constant circuits each made up of the resistor 14 (14a to 14d) of the resistance Rb and the capacitor 15 (15a to 15d) of the capacitance Ca, thereby to charge the capacitor 15 up to a potential corresponding to the detected output level. The output signals from the time constant circuits of the two pairs of channels are applied to the multiplexer 7 when the corresponding analog switches lOa and lOb are closed. An equivalent circuit of the time constant circuits of the first pair of the channels is as shown in Fig. 9. Also in this case, the time constant T is given by T = (Rb/2) x 2Ca = Ca x Rb, and is equal to that when the analog switch is open. Accordingly, when in a normal scanning mode, the charged level of each channel is read out every 1 ms, the additive value of the charged levels of the two channels may be obtained every 1 ms also in a high speed scanning mode if the output signals from the time constant circuits are selected every two channels and read out at 1 ms periods. Let us return to the CT scanner with the abovementioned data collecting system for collecting the detected data of radiation dose shown in Fig. 6. In the fourth generation CT scanner, for obtaining a resolution of 40 lp/cm, for example, data of 3.38 x 107 is required. To obtain this data for three seconds, a writing speed required for storing the data into a memory with a large memory capacity is 33.8 Mega words/sec or 67.6 Mega bytes/sec. Normally, one word of the data treated in the CT scanner is formed of two bytes. The CT scanner involves problems also of the writing speed of data into a data storing medium and the A/D converter for converting the analog detected data into digital data. The A/D converter with a sufficient number of bits for the CT scanning (16 bits because one word is formed of 2 bytes) is generally expensive and does not have high operation speed. In a typical example, if about 20 iis is taken for processing one data, for processing the data of 3.38 x 107 for one second, the necessary number of A/D converters is 676 (3.38 x 107 x 2 x 10 5 x 1 = 676). In this case, even if the resolution is reduced by half unwillingly, the number of channels is unchanged and therefore the number of the A/D converters required is also unchanged. The writing speed of the data storing medium is 33.8 Mega bytes, as mentioned above. In the data acquisition system designed such that the sampling time is doubled and the data of two channels are combined together into one data in a high speed scanning mode, since the number of channels used is halved, the necessary number of the A/D converters is halved to only 338, and the total number of data is 1/4 compared with the acquisition system that the channel data are not combined and the sampling pitch is not large. Therefore, the data writing speed is only 16.7 Mega bytes. In the case of an apparatus capable of obtaining 40 lp/cm for one second scanning, the scanning time is half, 0.5 sec. if it is merely employed to double the sampling pitch. In this case, however, if two channel data are combined together, the scanning time is halved again to 0.25 sec. According to the present invention, the following useful effects are attained; (1) In a high speed scanning mode, the number of views (or rays) is reduced with some sacrifice of the picture quality as given by the resolution and artifacts, and the number of rays (or views) is correspondingly reduced. With this feature, the scanning time is remarkably reduced with little deterioration of the picture quality. (2) Since the above feature reduces the number of data, the memory capacity of the data storing medium is also saved, leading to shortening of picture reproducing time. While the above description has been made on the fourth generation CT scanner, the same thing applies to the third generation CT scanner. Hence, no further description of it will be given to avoid duplication. The number of the combined channels is not limited to two, but may be any other suitable number such as three or four. Further, the radiation used may be Y rays or any other rays, in addition to the X-rays. Circuit arrangements of the data collecting system for combining the three channels data together are illustrated in Figs. 10 and 11. In the embodiment shown in Fig. 10, an analog switch lOa is connected between the output terminals of the radiation detecting cells 3a and 3b, and another analog switch lOb is inserted between the sensing cells 3b and 3c. Analog switches lla and llb are connected in series between the output terminals of the detecting cells 3a and 3c and the input terminals of amplifiers 5a and 5c, respectively, as shown. In a normal scanning mode, the analog switches lOa and lOb are opened while the analog switches Ila and llb are closed. In a high speed scanning mode, the analog switches lOa and lOb are closed while the analog switches lia and llb are opened. With the circuitry thus formed by the switches, the output signals from the three detecting cells 3a to 3c are applied to the single amplifier 5b, thereby obtaining the composite value of these three detected output signals. Although not shown, the multiplexer 7 is of course so controlled as to select the output signals every three channels. In the case of Fig. 11, the analog switches lla and llb in Fig. 10 are used in place of the analog switches Ila and llb in Fig. 10. This circuit arrangement allows the detected output signals from the detecting cells 3a to 3c to be combined and applied to the input of a center amplifier 5b with no resistor.

Claims: Claims: 1. A computerized tomography apparatus comprising: radiation source for radiating a radiation beam spreading in a fan shape toward the center of a region where an object under diagnosis is located; a radiation detector including a plurality of radiation detector elements for sensing the radiation beam radiated from said radiation source and transmitted through the object, said plurality of radiation detector elements being rotatable about the center of a tomographing field irradiated from said radiation source in synchronism with the rotation of said radiation source or said plurality of radiation detector elements being; fixed and circularly arranged about the center of said tomographing field irradiated by said radiation source; a radiation dose calculating circuit respectively connected to said plurality of radiation detector elements in said radiation detector and for calculating the dose of radiation detected by said detector elements to provide the calculated values, said calculating circuit being provided corresponding to a plurality of channels; a multiplexer connected to said plurality of radiation dose calculating circuit and for sequentially selecting said integrated radiation dose; a circuit connected to said multiplexer for analog-to-digital converting the detected radiation data selected by and outputted from said multiplexer and for supplying the converted digital data to a computer system; characterized in that there is further provided with switching means which composes the doses of the radiations detected by at least said two radiation detector elements and supplying the composited dose as a radiation dose output from a combination of a plurality of channels to said multiplier. 2. A computerized tomography apparatus according to claim 1, wherein said switching means are each an analog switch connected between a channel including one of said radiation detector elements and one of said radiation dose calculating circuit and at least another channel with such a construction, and closes to compose the radiation dose sensed by at least two said radiation detect elements. 3. A computerized tomography apparatus according to claim 1, wherein said switching means are each an analog switch connected between said channels which is closed in order to compose the radiation dose calculated by said radiation dose calculating circuit in one of said channels and the one integrated by the radiation dose calculating circuit in another channel.

Inventors: MORI, ISSEI
Applicants: KABUSHIKI KAISHA TOSHIBA
Application year: 1982
Main classification: A61B6
IPC codes: G01N23, A61B6