JP2002017717A - X-ray ct apparatus - Google Patents

Abstract

PROBLEM TO BE SOLVED: To obtain required S/N on projection data even if a subject is substantially photographed in a low dose mode in an X-ray CT apparatus. SOLUTION: This X-ray CT apparatus is provided with an X-ray tube 40 and an X-ray detector array 70 opposed to each other with the subject 100 placed between, and recomposes a CT tomographic image of the subject on the basis of projection data collected from the X-ray detector array. The X-ray CT apparatus is further provided with an integrator 1 integrating the detection output of an X-ray detector XDi and capable of changing the integrating time constant, an X-ray dose estimating means 2 for estimating an X-ray dose per picture element inputted to the X-ray detector XDi, on the basis of specified set information related to a scanning plan, and a control means 3 for changing the integrating time constant of the integrator in a diminishing direction when the estimated X-ray dose is less than a specified threshold TH.

Description

DETAILED DESCRIPTION OF THE INVENTION

[0001]

BACKGROUND OF THE INVENTION 1. Field of the Invention The present invention relates to an X-ray CT apparatus, and more particularly, to an X-ray CT apparatus having an X-ray tube and an X-ray detector array opposed to each other with a subject interposed therebetween. The present invention relates to an X-ray CT apparatus that reconstructs a CT tomographic image of a subject based on the X-ray CT apparatus.

In this type of X-ray CT apparatus, when the X-ray dose input to the X-ray detector decreases, electrical noise generated in electronic circuits (X-ray detector, integrator, etc.) cannot be ignored. Therefore, the X-ray detection signal, and further the integration
The S / N ratio of the / D converted projection data deteriorates. Therefore, it is desired that the S / N ratio of the X-ray detection signal (projection data) is not deteriorated even if the X-ray dose becomes small.

[0003]

2. Description of the Related Art FIG. 11 is a block diagram of a main part of a conventional X-ray CT apparatus, mainly showing a configuration of a data acquisition system (DAS). In the figure, 40 is a rotating anode type X-ray tube, 50
Is a collimator for limiting the X-ray irradiation range (mainly in the slice thickness direction), 100 is the subject, 20 is an imaging table on which the subject is placed and moved in the body axis direction, and 70 is a large number (for example, n = 1000). XD detector array in which the X-ray detectors are arranged in a single row in an arc shape, XD1 to XDn are X-ray detectors composed of, for example, a scintillator and a photodiode, and 81 _{1 to} 81 _n are X-ray detectors. A1 to An are amplifiers, SH1 to SHn are sample and hold circuits, and 82 is a signal multiplexer (MP
X), 83 are A / D converters (A / D), 15 is a data acquisition buffer, 11 is a central processing unit that performs main control and processing (scan control, CT tomographic image reconstruction processing, etc.) of the X-ray CT apparatus , 11a is its CPU, 11b is a main memory (MEM) storing a control program executed by the CPU, etc., 1
Reference numeral 4 denotes a control interface of the CPU 11a, and reference numeral 84 denotes a timing generator (TG) that generates various timing signals related to X-ray data detection / collection control.

In operation, the fan beam from the X-ray tube 40 passes through the subject 100 through the X-ray detector array 70.
All at once. Now, focusing on the signal detection processing of the X-ray beam XB1, the X-ray detector XD1
And outputs a current signal IB1 corresponding to the intensity of the integrator 81 ₁
Integrates the input current signal IB1 with a constant time constant (capacitance C).

[0005] This integration operation can be described by a general equation:
Equation (1) is applied between the input current i _S and the output voltage V _{O of the} integrator,

[0006]

(Equation 1)

[0007] There is a relationship. Here, the output voltage V _O of the integrator has a relationship of integrating the input current i _S at a ratio of (1 / C). Therefore, in this specification, the capacitance C is also referred to as a time constant of integration. The resistor R in this case is a so-called circuit protection resistor having a relatively small resistance value, and does not contribute to the integration operation (time constant).

Furthermore, the amplifier A1 amplifies the output of the integrator 81 _1, the sample-hold circuit SH1 samples and holds the output of the amplifier A1 at a predetermined timing. Other X
The same applies to each signal detection processing of the line beams XB2 to XBn. The signal multiplexer 82 scans each sample output of the sample hold circuits SH1 to SHn at high speed, and the A / D converter 83
2 is A / D converted at high speed. A series of main signal data (projection data) thus obtained is stored in the data collection buffer 15 and processed by the CPU 11a.

In the integrators 81 _{1 to} 81 _n , the block “RESET” is a reset circuit for the capacitor C. When the reset signal RS of the timing generator 84 is activated, the analog switch S5 is closed, The charge Q of the capacitor C is discharged via the reset circuit. The block “AUTO ZERO” is an input offset voltage calibration circuit of the operational amplifier OA, and when the calibration signal CAL of the timing generation unit 84 is activated, the analog switches S6 and S7 are switched to the opposite sides shown in the drawing. The bias state of the calibration circuit is updated so that the output voltage V _O of the operational amplifier OA at that time becomes 0 V.

[0010] By the way, noises which are problematic in this type of X-ray CT apparatus are roughly classified into quantum noise due to X-rays and amplifier noise due to an electronic circuit at the first stage of X-ray detection. Quantum noise is caused by temporal and spatial fluctuations in X-ray photon density. On the other hand, amplifier noise is caused by thermal noise of circuit resistance in the first-stage amplifier, shot noise accompanying current of transistors and the like. These noises are X
This is not a problem when a sufficient X-ray dose is input to the X-ray detector XD (in the high dose mode), but when the X-ray dose is reduced (in the low dose mode), these noises are reduced. This cannot be ignored and degrades the S / N ratio of the projection data. In particular, the latter amplifier noise occurs regardless of the presence or absence of an input signal and is superimposed on the signal component in the form of addition, thus having a considerable adverse effect on the S / N ratio of the projection data.

[0011]

However, in the above-mentioned conventional X-ray CT apparatus, no countermeasure against such amplifier noise is taken, and therefore, when the subject is photographed in the low dose mode. However, there is a problem that a scout image (two-dimensional X-ray image) obtained by the scout scan and a CT tomographic image obtained by the axial / helical scan are significantly deteriorated.

SUMMARY OF THE INVENTION The present invention has been made in view of the above-mentioned problems of the prior art, and has as its object the purpose of providing projection data required even when an object is to be photographed in a substantially low-dose mode. An object of the present invention is to provide an X-ray CT apparatus capable of obtaining an S / N ratio.

[0013]

The above-mentioned problem is solved, for example, by referring to FIG.
Is solved. That is, the X-ray C of the present invention (1)
The T apparatus includes an X-ray tube 40 facing each other across the subject 100.
And an X-ray detector array for reconstructing a CT tomographic image of the subject based on the projection data collected from the X-ray detector array, and integrating the detection output of the X-ray detector XDi. X-ray dose estimator for estimating the X-ray dose per pixel input to the X-ray detector XDi based on predetermined setting information relating to a scan plan, the detector 1 being capable of changing its integration time constant. Means 2 and control means 3 for changing the integration time constant of the integrator 1 to a direction in which the integration time constant of the integrator 1 is reduced when the estimated X-ray dose falls below a predetermined threshold TH.
Is provided.

In the present invention (1), the integrator 1 can improve the S / N ratio of the X-ray detection signal (ie, the integrator 1) in the low dose mode by changing its integration time constant. . The principle configuration and operation will be apparent in the description of the embodiment below. And
The X-ray dose estimating means 2 estimates the X-ray dose per pixel input to the X-ray detector XDi based on predetermined setting information relating to the scan plan, and the control means 3 determines whether the estimated X-ray dose is When the value falls below the predetermined threshold value TH, the integration time constant of the integrator 1 is changed to a smaller value.

Therefore, even when the subject is to be imaged in the substantially low-dose mode, such a state can be accurately estimated, and at that time, the X-ray detection signal (integration By improving the S / N ratio of the device 1), high-quality projection data of the subject can always be obtained.

The predetermined threshold value TH can be set in a plurality of steps, and the integration time constant can be reduced in a plurality of steps accordingly.

Preferably, in the present invention (2), in the above-mentioned present invention (1), the X-ray dose estimating means 2 comprises: an irradiation energy of an X-ray tube relating to a scan plan; This is for estimating an X-ray dose per pixel input to the X-ray detector based on setting information of one or more of the scan time of the subject and the body type of the subject.

In the present invention (2), the X-ray dose per pixel input to the X-ray detector is determined by the irradiation energy of the X-ray tube, the slice thickness of the subject, the scan time of the subject, and / or the subject. Since the subject changes depending on the body type (obesity, etc.) of the subject, the subject may be substantially imaged in the low-dose mode based on one or more of these setting information. It can be accurately estimated. The information on the body type of the subject can be manually set based on, for example, the feeling of the operator visually observing the subject.

Preferably, in the present invention (3),
In the present invention (2), the body type estimating means 4 for estimating the body type of the subject based on the scout image data of the subject obtained by the scout scan performed by previously fixing the rotation angle around the body axis of the X-ray imaging system X-ray dose estimating means 2
Is for setting the estimated output of the body type estimating means 4 as setting information relating to the body type of the subject. Therefore, according to the present invention (3), the body type of the subject can be objectively and accurately estimated.

[0020]

DETAILED DESCRIPTION OF THE PREFERRED EMBODIMENTS Preferred embodiments of the present invention will be described below in detail with reference to the accompanying drawings. Note that the same reference numerals indicate the same or corresponding parts throughout the drawings.

FIG. 2 is a block diagram of the main part of the X-ray CT apparatus according to the embodiment. In FIG. 2, reference numeral 10 denotes an operation console operated by a user, reference numeral 20 denotes an imaging table on which a subject is placed and moved in the body axis direction, Numeral denotes a scanning gantry for performing Axial / Herical scanning / reading of a subject by an X-ray fan beam.

In the scanning gantry 30, reference numeral 40 denotes a rotating anode type X-ray tube, 41 denotes X-ray irradiation timing (corresponding to the X-ray irradiation time Sec to the subject) and X-ray intensity (tube voltage k).
V, tube current mA), 50 is a collimator that limits the X-ray irradiation range (mainly in the slice thickness direction), 51 is the X-ray transmission slit width (in the body axis direction of the subject). A collimator control unit for adjusting the detection width Thic) and position; a rotation control unit 60 for rotating an X-ray imaging system including the X-ray tube 40 and the X-ray detector array 70 around the body axis of the subject; 70 is a large number (for example, about n = 1000) of X
An X-ray detector array in which the line detectors are arranged in, for example, one line in an arc shape, and 80 is a data acquisition unit (DAS) that acquires detection data (projection data) of the X-ray detector array.

In the operation console 10, reference numeral 11 denotes a central processing unit which performs main control and processing (scan control, CT image reconstruction processing, time constant control of the integrator according to the present invention) of the X-ray CT apparatus, and 12 denotes a keyboard. And an input device 13 including a mouse and the like, and 13 is a scan plan (imaging parameters kV, m
A, Sec, Thic, etc.) and a display device (CRT) 14 for displaying a CT reconstructed image and the like, the CPU 11a is provided with various control signals C between the scanning gantry 30 and the imaging table 20.
And a control interface for exchanging the monitor signal SD, a data collection buffer 15 for temporarily storing main signal data from the data collection unit 80, and 16 various data and application programs necessary for operating the X-ray CT apparatus. Storage device (disk device, etc.) that stores data
It is.

With this configuration, the fan beam from the X-ray tube 40 simultaneously enters the X-ray detector array 70 via the subject 100. The data collection unit 80 scans and collects the detection data (projection data) of the X-ray detector array 70 and stores it in the data collection buffer 15. Further scanning gantry 30
Performs projection similar to the above in each view slightly rotated, collects and accumulates projection data for one rotation of the scanning gantry (for example, about 1000 views), and moves the imaging table 20 in the body axis direction according to the axial / helical scan method. Intermittently / continuously to the subject 1
All projection data for the required imaging area of 00 is collected and stored. Then, the CPU 11a reconstructs a CT tomographic image of the subject 100 based on the obtained total projection data and displays the CT tomographic image on the display device 13.

FIG. 3 shows a data collection unit (D) according to the embodiment.
AS) 80 is a block diagram. The integrators 81 _{1 to} 81 _{n according to the} present embodiment include, for example, two capacitors C1 and C2 in parallel in order to be able to change the integration time constant. A switch ASW is provided in series and the CPU 11a
ON / OFF the contact SW according to the control signal SL from the
F is provided. Now, assuming that the capacitance value of the integration in the normal state (that is, when the mode is not the low-dose mode) is the same as the conventional capacitance value C, for example, there is a relationship of C = C1 + C2. Although the ratio C1 / C2 of these capacitance values C1 and C2 can be selected as desired, for example, it is assumed that C1 = C2 = C / 2. Other configurations may be the same as those described with reference to FIG.

Next, the S / N of the X-ray detection signal in the present invention
The principle configuration and operation of the ratio improvement will be described. 9 and 11
0 is a diagram (1), (2) for explaining the S / N ratio of the integrator according to the embodiment, and FIG.
X comprising a line detector 71 and a mirror type integrator 81
3 shows a basic configuration of a line detection circuit. The photodiode 71 has a detection current i corresponding to the input X-ray beam intensity.
_S , which is denoted by current source i _S. Cd is a junction capacitance of the photodiode and exists in parallel with the current source i _S. The integrator 81 converts the input signal current i _S ′ into a capacitance C (= C1 + C
Integrate in 2). Now, when a noise voltage converted to the input side of the operational amplifier OA and v _n, the ratio V _S / V of the S / N ratio of the X-ray detection circuit of the output voltage V _N of the output voltage V _S and a noise component of the signal component Given by _N. Hereinafter, each component is obtained.

FIG. 10A is a diagram for obtaining the output voltage V _S of the signal component. Now, the output impedance of the operational amplifier OA is sufficiently small, the signal current of the current source i _S i _S
Is divided into the junction capacitance Cd and the integration capacitance C. At this time, the signal current i _s ′ flowing to the integration capacitance C is given by the equation (2).

[0028]

(Equation 2)

Therefore, the output voltage V _S of the signal component becomes (3)
It is obtained by the equation.

[0030]

(Equation 3)

FIG. 10B is a diagram for obtaining the output voltage V _N of the noise component. In the figure, the operational amplifier OA - voltage side input terminal V _- is given by equation (4).

[0032]

(Equation 4)

The output voltage V _N of the noise component is obtained from the equation (5) based on the voltage V ₊ = v _{N at the +} input terminal of the operational amplifier OA.

[0034]

(Equation 5)

When this is arranged for the output voltage V _N , the equation (6) is obtained.

[0036]

(Equation 6)

Further, since the open gain A of the operational amplifier OA is very large, the output voltage V _N of the noise component is finally given by the equation (7).

[0038]

(Equation 7)

Returning to FIG. 9, FIG. 10 (A) and FIG. 10 (B) are superimposed, and the output signal V
_O includes the signal component V _S and the noise component V _N obtained above. The S / N ratio of this X-ray detection circuit is finally given by equation (8) as the ratio V _S / V _N of these signal components.

[0040]

(Equation 8)

Note that CdＣC.

According to the above equation (8), S of the X-ray detection circuit
/ N ratio is proportional to the signal current i _S, and the combined capacitance (C +
2Cd) and the noise voltage v _N. Among them, the signal current i _S is determined by the input X-ray beam intensity, and is good in the high dose mode, but the S / N ratio becomes a problem in the low dose mode. The junction capacitance Cd is determined by the element structure (characteristics) of the photodiode, and the noise voltage v _N is determined by the circuit structure (characteristics) of the operational amplifier OA. That is, these elements cannot be easily changed from the outside.

Therefore, in the present embodiment, the capacitance value of the integrating capacitor C is changed. That is, the above (8)
According to the formula, when the signal current i _S of the X-ray detection is reduced (i.e., if that would subject 100 is substantially captured in the low-dose mode) be the capacitance value of the integrating capacitor C Is reduced, the S / N ratio of the X-ray detection circuit can be improved.

Next, scan control of the X-ray CT apparatus based on the above configuration will be described. FIG. 4 is a flowchart of the scan control according to the embodiment.
When performing a helical scan or the like, this process is input. The following step S1 executed by the CPU 11a
1 and S12 correspond to the X-ray dose estimation means 2 in FIG. 1 described above, and the processing in steps S13 to S15 corresponds to the control means 3 in FIG.

In step S11, scan plan (protocol) setting information set in advance by the operator is input. FIG. 5 shows an image diagram of scan plan information setting input in the embodiment. For example, after the end of the scout scan performed in advance, a scan setting screen 13a for the subsequent axial / helical scan is displayed on the screen 13A of the display device 13, and the operator clicks or inputs the necessary scan parameters with a mouse or key. input. An example scan plan for acquiring image A is as follows.

Scan type [Scan Type] = Axial scan Scan start position on body (Z) axis [Start Loc] = Z1 Scan end position on body (Z) axis [End Loc] = Z2 Number of captured images [NO.of Images] = 13 slice thickness of subject [Thick] = 2 mm Scan time [Sec] = 13 seconds Tube voltage of X-ray tube [kV] = 120 kV Tube current of X-ray tube [mA] = 280 mA For Sec, the axial scan represents the X-ray irradiation time for one revolution of the gantry,
In the scout scan, it indicates the total X-ray irradiation time. This scan time affects the X-ray dose per pixel input to the X-ray detector 71 (that is, the integration time).

Further, on this scan setting screen, the operator
When the [Show Localizer] icon is clicked, a scout image 100A of the subject as shown in the figure is displayed in the image display area 13b of the display screen 13A, and a line indicating the axial slice position is displayed thereon. The bold line in the figure represents the start position and the end position, and the dotted line represents each intermediate slice position (corresponding to the center line of the slice width).

Returning to FIG. 4, in step S12, the CPU 1
1a is input to the X-ray detector 71 at the time of the subsequent axial scan based on the setting information of the scan plan.
The X-ray dose per pixel (one integration time) is estimated. Various methods are conceivable for this estimation, and an example of a simple method will be described below.

When the standard body 100 is photographed for a predetermined medical purpose, certain reference photographing parameters are tube current = mA _r , slice thickness = Thic _r , and scan time = Se.
c _r, when the tube voltage = kV _r, tubes imaging parameters based on the scan plan current = mA _a, slice thickness = Th
Assuming that ic _a , scan time = Sec _a , and tube voltage = kV _a , one pixel (one integration time) input to the X-ray detector 71
An example X-ray dose per is estimated by the product R _ATiO of any one or more of the ratio of these parameters. For example, the equation for estimating the X-ray dose when all of the above four parameters are used is shown in equation (9).

[0050]

(Equation 9)

Here, for example, β = 3. The above (9)
According to the equation, the X-ray dose R _atio per pixel input to the X-ray detector 71 is such that the smaller the irradiation energy of the X-ray tube 40 and the thinner the slice thickness of the subject 100,
Also, the shorter the scan time of the subject, the smaller the relationship. Hereinafter, this relationship will be specifically described with reference to FIGS.

FIG. 6A shows the tube voltage-tube current characteristics of the X-ray tube 40. The relationship between the irradiation energy of the X-ray tube 40 and the X-ray dose per pixel input to the X-ray detector 71 is shown. Shows the relationship. When the filament of the X-ray tube 40 is heated, a tube current mA flows from the anode to the cathode. At this time, if the DC filament current If is kept constant, a substantially constant tube current mA flows at a certain tube voltage kV or higher irrespective of the tube voltage kV as shown in the figure, and this region is also called a temperature limit region. Therefore, in the present embodiment, in low-voltage imaging in which the tube voltage kV falls below, for example, 100 kV, the irradiation energy of the X-ray tube 40 (that is, the tube current mA) is changed by controlling the tube voltage kV. The tube voltage is 100k
In a high-voltage shooting that exceeds V, the filament current I
By controlling f, the irradiation energy (tube current mA) of the X-ray tube 40 is changed.

Therefore, the X-ray tube 40 based on the scan plan
Tube voltage kV _a, the X-ray dose per pixel to be input to the X-ray detector 71 based on the tube current mA _a can be estimated. That is, the smaller the tube voltage kV _a and / or the tube current mA _a (filament current If), the smaller the X-ray dose.

FIG. 6B shows the relationship between the slice thickness Thic of the subject and the X-ray dose per pixel input to the X-ray detector 71. Although the X-ray detector 71 has a certain width W in the body axis direction of the subject 100, the X-ray dose per pixel actually input to the X-ray detector 71 is the slice of the subject 100. It changes according to the opening width of the collimator 50 corresponding to the thickness Thic1, Thic2, and the like. That is, the smaller the slice thickness Thic, the smaller the X-ray dose.

FIG. 7 shows the relationship between the scan time Sec of the subject and the X-ray dose per pixel input to the X-ray detector 71. If the scan time Sec is in the normal range, the integration time per pixel = t2, and even with the normal integration capacitance C (= C1 + C2), the output voltage V _O can be integrated to the full range LIM within the time t2. on the other hand,
When the scan time Sec becomes relatively short, the integration time t1 per pixel becomes relatively short. With the normal integration capacitance C (= C1 + C2), the integration to the full range LIM of the output voltage V _{O is completed} within the integration time t1. Can not do,
As a result, the S / N ratio of the projection data deteriorates.

The fact that the integration time t1 is short means that the X-rays per pixel input to the X-ray
This is equivalent to a small radiation dose. Therefore, the shorter the scan time Sec (that is, the higher the rotation speed of the scanning operation gantry 30), the smaller the X-ray dose.

Although the above equation (9) does not include information on the body type (obesity etc.) of the subject 100,
In this embodiment, the operator sets the object 100 in advance.
Is visually determined to determine its body type. For example, if it is determined that the subject is obese, when inputting the setting of the scan plan information, for example, X
This is because it is assumed that sets the tube voltage kV _a and or tube current mA _a line pipe 40 slightly larger. Note that a method of automatically determining the body type of the subject 100 separately and evaluating the obtained body shape information by adding it to the above-described equation (9) will be described later with reference to FIG.

Returning to FIG. 4, in step S13, the CPU 1
1a compares the X-ray dose Ratio estimated by the above equation (9) with a predetermined threshold TH, and determines whether Ratio <TH. If Ratio <TH is not satisfied, the time constant of the integrator 81 is set to C (= C1 + C2) in step S14, and
If io <TH, the time constant of the integrator 81 is set to C1 (<C) in step S15. In step S16, an axial scan based on the scan plan is performed.

At this time, integrators 81 ₁ to 81 of all channels
An appropriate integration time constant of 1n is selected in accordance with the X-ray dose per pixel input to the X-ray detectors XD1 to XDn. Even in the case of being photographed, the S / N ratio of the projection data can be secured as required.

FIG. 8 is a diagram for explaining the body shape estimation processing of the subject based on the scout image in the embodiment. FIG.
In (A), a scout scan is performed on the subject 100 in advance at a view angle of 0 ° (FIG. A) and / or a view angle of 90 ° (FIG. B) to obtain a scout elephant for a predetermined imaging region. The scout scan is normally performed before the original axial / helical scan. In the present embodiment, the two-dimensional projection data of the subject obtained by the scout scan is used to estimate the body shape of the subject. Available to

FIG. 8B shows a scout image 100a when a subject 100 having a non-obese figure is scout-scanned at a view angle of 90 °. This scout image 1
00a is obtained from each projection data g (X, 90) having a fixed view angle. The CPU 11a (corresponding to the body type estimating means 4 in FIG. 1) compares each projection data g (X, 90) with a predetermined threshold value Th for separating the air and the subject 100, thereby obtaining the contour of the subject 100. A binary image whose shape can be recognized is generated, and the subject width of each binary image (corresponding to the subject's chest thickness) h
Estimating the type Vol _a of the subject 100 based on the average value Ha or maximum value H _max.

FIG. 8C shows a subject 10 of a certain obese body type.
0 shows a scout image 100b when a scout scan is performed at a view angle of 90 °. Type Vol above similarly-obtained specimen width of each binary image (corresponding to the chest thickness of the object) average Ha or the subject 100 based on the maximum value H _max of h
to estimate the _a.

Now, the body type of the subject having the standard body shape is V
Assuming that ol _r , the obtained estimated body type Vol _a is added to the evaluation of the estimated X-ray amount _{Ratio in} the form of the following equation (10), for example.

[0064]

(Equation 10)

According to the above equation (10), the X-ray dose _Ratio per pixel input to the X-ray detector 71 decreases as the body type of the subject becomes obese. Although not shown, the same applies to a case where the subject 100 is scout-scanned at a view angle of 0 °.
If the scout scan is performed at 90 ° and 90 °, the body shape of the subject 100 can be more accurately estimated. According to the present embodiment,
Since the body shape of the subject 100 can be automatically and objectively estimated without depending on the operator's feeling, the change control of the integration time constant becomes more reliable.

[0066] In the above embodiment has been described the case of generating the integrated voltage signal V1 to integrator 81 ₁ corresponding to integrated directly into the detection current output IB1 of X-ray detectors XD1, not limited to this . Depending on the X-ray CT apparatus, the detection current output IB of the X-ray detector XD1, as shown by the arrow in FIG.
1 is once converted into a voltage signal VB1 (including logarithmic conversion) by a preamplifier PA1 and then input to an integrator 81 ₁ . In the input circuit of the integrator 81 ₁ (the operational amplifier OA1) in this case, the relationship of IB1 = VB1 / R holds. That is, generally speaking, the relationship between the input current Ii of the integrator 81 ₁ and the input voltage Vi is obtained. Since the relationship of Ii = Vi / R is established between them, when this is substituted into the above-mentioned equation (1) of the current integration, the relationship of the equation (11) between the output voltage Vo of the integrator and the input current Ii is obtained. There is.

[0067]

[Equation 11]

This is a case where the input (signal source) of the integrator is a voltage source. According to the above equation (11), the output voltage Vo of the integrator is obtained by subtracting the input voltage Vi from (1 / C · R). There is a relationship of integrating by ratio. Therefore, in this case, the product (C · R) of the capacitance C and the resistance R is called a time constant of integration. The present invention is also applicable to this type of X-ray CT apparatus. In this case, the integration time constant can be improved by reducing the resistance R and / or the capacitance C to improve the S / N ratio when the subject is to be photographed substantially in the low dose mode.

In the above embodiment, the case where the integration time constant can be changed in two steps has been described. However, the integration time constant may be changed in three or more steps. In this case, the X-ray estimation amount R
_A plurality of threshold _values are provided for the _atio , and the X-ray dose per pixel input to the X-ray detector 71 is estimated in a plurality of stages.

According to the present embodiment, there is also an advantage that the dynamic range of the output signal Vo in the low dose mode can be increased by reducing the time constant of integration.

Although the preferred embodiments of the present invention have been described, it goes without saying that various changes in the configuration, control, processing, and combinations thereof can be made without departing from the spirit of the present invention. There is no.

[0072]

As described above, according to the present invention, a required S / N ratio can be ensured for the projection data even when the subject is to be imaged substantially in the low dose mode. This greatly contributes to the improvement of X-ray CT imaging performance and image reliability.

[Brief description of the drawings]

FIG. 1 is a diagram illustrating the principle of the present invention.

FIG. 2 is a main part configuration diagram of an X-ray CT apparatus according to an embodiment.

FIG. 3 shows a data collection unit (DAS) 80 according to the embodiment.
It is a block diagram of.

FIG. 4 is a flowchart of scan control according to the embodiment.

FIG. 5 is an image diagram of scan plan information setting input in the embodiment.

FIG. 6 is a diagram (1) illustrating an X-ray dose estimation process in the embodiment.

FIG. 7 is a diagram (2) illustrating an X-ray dose estimation process according to the embodiment.

FIG. 8 is a diagram illustrating a process of estimating a body type of a subject based on a scout image according to the embodiment.

FIG. 9 is a diagram (1) illustrating an S / N ratio of the integrator according to the embodiment.

FIG. 10 is a diagram (2) illustrating an S / N ratio of the integrator according to the embodiment.

FIG. 11 is a configuration diagram of a main part of a conventional X-ray CT apparatus.

[Explanation of symbols]

Reference Signs List 40 X-ray tube 50 Collimator 20 Imaging table 70 X-ray detector array 81 _{1 to} 81 _n Integrator 100 Subject ASW Analog switch OA Operational amplifier XD1 to XDn X-ray detector

 ────────────────────────────────────────────────── ─── Continued on the front page (72) Inventor Yoshi Ichinoseki 4-7, Asahigaoka, Hino-shi, Tokyo 127 G F Yokogawa Medical System Co., Ltd. F-term (reference) 4C093 AA22 BA03 CA06 EA02 FA33 FC04

Claims (3)

[Claims] An X-ray tube and an X-ray tube facing each other across a subject
An X-ray detector for reconstructing a CT tomographic image of a subject based on projection data collected from the X-ray detector array
In the X-ray CT apparatus, an integrator that integrates a detection output of an X-ray detector, the integration time constant of which can be changed, and an integrator that is input to the X-ray detector based on predetermined setting information related to a scan plan X-ray dose estimating means for estimating the X-ray dose per pixel; and control means for changing the integrated time constant of the integrator to a smaller direction when the estimated X-ray dose falls below a predetermined threshold. An X-ray CT apparatus characterized by the above-mentioned. 2. An X-ray dose estimating means, comprising: one of an irradiation energy of an X-ray tube according to a scan plan, a slice thickness in an axial direction of a subject, a scan time of the subject, and a body type of the subject. The X-ray CT apparatus according to claim 1, wherein an X-ray dose per pixel input to the X-ray detector is estimated based on two or more setting information. 3. A body shape estimating means for estimating a body shape of the subject based on scout image data of the subject obtained by a scout scan performed beforehand by fixing a rotation angle around a body axis of the X-ray imaging system, 3. The X-ray CT apparatus according to claim 2, wherein the X-ray dose estimating unit uses the estimated output of the body type estimating unit as setting information relating to the body type of the subject. 4.