KR102286445B1 - X-ray Detector and X-ray apparatus - Google Patents

Abstract

The present invention provides an X-ray detector configured to include a pixel generating a pixel signal value according to an irradiated X-ray dose, and configured to have two or more different linear relationships between the X-ray dose and the pixel signal value.

Description

X-ray detector and X-ray apparatus {X-ray Detector and X-ray apparatus}

The present invention relates to an X-ray detector and an X-ray imaging apparatus, and more particularly, to an X-ray detector capable of obtaining an X-ray image of a wide dynamic range and high contrast, and X-ray imaging using the same It's about the device.

A mammography apparatus, which is an example of an X-ray imaging apparatus, is for early diagnosis of breast cancer or microcalcification, etc., and irradiates a certain amount of X-rays to the breast, which is a subject, and detects X-rays that have passed through the breast. 2D X-ray image of the breast (hereinafter referred to as a two-dimensional image) or a three-dimensional X-ray image (hereinafter referred to as a three-dimensional image) of the breast by receiving light.

The X-ray detector, which is a sensor used in a conventional digital mammography apparatus, has a characteristic that a signal value generated according to the amount of X-rays linearly increases with a certain slope and then becomes saturated.

When a conventional X-ray detector having such characteristics is used, there is a problem in that the generated X-ray image has a narrow dynamic area and low contrast.

An object of the present invention is to provide an X-ray detector and an X-ray imaging apparatus having more improved characteristics.

Specifically, an object of the present invention is to provide an X-ray detector capable of obtaining an X-ray image of a wide dynamic range and high contrast, and an X-ray imaging apparatus using the same.

In order to achieve the above object, the present invention includes a pixel generating a pixel signal value according to an irradiated X-ray dose, and the X-ray dose and the pixel signal value are configured to have two or more different linear relationships. A line detector is provided.

Here, when the amount of X-rays is in the first range, the pixel signal values are configured to have a linear relationship of a first slope, and when the amount of X-rays is in a second range greater than the first range, the pixel signal values are It may be configured to have a linear relationship of the second slope that is smaller than the first slope.

The pixel includes a photodiode; It may include a capacitor connected in parallel with the photodiode.

The photodiode may be a PN junction diode, and may have a first capacitance value in the first range and a second capacitance value greater than the first capacitance value in the second range.

The photodiode may have a deep depletion state in the first range and an accumulation state in the second range.

The pixel includes a reset transistor connected to the cathode of the photodiode; a source-follow transistor connected to a node between the photodiode and the reset transistor; It may further include a selection transistor connected between the source follow transistor and the signal output line.

It may be configured to generate a pixel signal value of 75% to 85% with respect to a saturation value of the pixel signal value at a first conversion point between the first range and the second range.

The X-ray dose of the first conversion point may be 150mR ~ 250mR.

The X-ray dose at the second conversion point between the second range and the range of the X-ray dose in which the saturation value is generated may be 300mR to 400mR.

In another aspect, the X-ray detector described above; It provides an X-ray imaging apparatus including a generator for irradiating X-rays to the X-ray detector.

Here, it may include a compression pad for pressing the breast on the X-ray detector.

In the present invention, by connecting a capacitor in parallel to a photodiode having a PN junction structure, the pixel signal value increases in two different linear relationships depending on the amount of X-rays to realize saturation.

Accordingly, the X-ray detector of the X-ray imaging apparatus can implement a wide dynamic area and high-contrast X-ray image.

1 is a perspective view of a mammography apparatus according to an embodiment of the present invention.
Figure 2 is a graph showing the saturation characteristics of the X-ray detector of the mammography apparatus according to an embodiment of the present invention.
Figure 3 is a graph showing another example of the saturation characteristics of the X-ray detector of the mammography apparatus according to an embodiment of the present invention.
Figure 4 is a diagram schematically showing the configuration of the X-ray detector according to an embodiment of the present invention.
5 is a diagram schematically illustrating a pixel structure according to an embodiment of the present invention;
6 is a view showing a CV curve of a photodiode according to an embodiment of the present invention.

Hereinafter, embodiments of the present invention will be described in detail with reference to the drawings.

1 is a perspective view of a mammography apparatus according to an embodiment of the present invention. Hereinafter, a mammography apparatus will be described as an example of an X-ray imaging apparatus according to the present invention for convenience, but the present invention is not limited thereto.

As shown, the mammography apparatus according to an embodiment of the present invention includes a body 100 for X-ray imaging, and a column 200 supporting the body 100 .

The column 200 has a vertical column shape with a lower end fixed to the floor, and the body 100 moves up and down along the longitudinal direction of the column 200 .

The body 100 has a C-shape or a similar shape in which both upper and lower ends face each other. The main body 100 includes a column connection part 110 connected to the column 200 so as to be able to ascend and descend along the column 200, and a vertical connection part 120 connected to the column connection part 110 so as to be rotatable with respect to the column connection part 110. may include.

One end of the main body 100, for example, the upper end is equipped with a generator 130 to irradiate X-rays toward the lower end, the other end, and the other end of the main body 100, for example, the lower end, is equipped with an X-ray detector 140 and its X-rays passing through the interposed subject are received. In addition, the body 100 may include a compression pad 150 that presses the breast placed on the X-ray detector 140 or the support panel 160 between the generator 130 and the X-ray detector 140 .

At this time, the mammography apparatus according to an embodiment of the present invention can photograph the FFDM (full-field digital mammography) mode by rotating the entire body 100 at a certain angle, and the upper end of the main body 100 relative to the lower end at a certain angle. It can be rotated to shoot digital breast tomosynthesis (DBT) or breast tomosynthesis (BTS) mode.

On the other hand, the X-ray detector 140 acquires image information according to the intensity for each position of the X-ray that has passed through the breast. The X-ray detector 140 for this purpose has a form in which a plurality of sensors are two-dimensionally arranged in units of pixels, and a form in which unit pixel sensors form a plurality of matrices and are arranged in a matrix form.

Hereinafter, according to an embodiment of the present invention, the X-ray detector 140 of the mammography apparatus capable of obtaining an X-ray image of a wide dynamic area and high contrast will be described in more detail.

2 is a graph showing the saturation characteristics of the X-ray detector of the mammography apparatus according to an embodiment of the present invention, along with a graph of the saturation characteristics of the conventional X-ray detector. In FIG. 2 , the horizontal axis represents the X-ray dose (mR), and the vertical axis represents a signal value detected by the X-ray detector. Here, the signal value is an offset signal value having a value of 0 when the amount of X-rays is 0, and is a digital value.

As shown in Figure 2, the conventional X-ray detector has a characteristic that the signal value increases linearly as the amount of X-rays increases, and the signal value is saturated at a specific X-ray dose point (Ptr) or more. For example, in the figure, for an X-ray dose in the range of about 130 mR from the origin, the signal value linearly increases with a single slope Sp (ie, a single gain) up to about 13000, and then saturates.

On the other hand, in the X-ray detector according to the present embodiment, the linear relationship between the amount of X-rays and the signal value may have two or more different linear relationships in saturation characteristics.

In this regard, in this embodiment, the case where the saturation characteristic of the X-ray detector is configured to have two linear relationships is taken as an example. In this case, the range of the X-ray dose before the saturation state may be divided into two ranges, a first range that is a relatively low dose range, and a second range that is a relatively high dose range.

As such, when saturation characteristics of two different linear relationships are defined, a transformation point, which is a point at which the slope changes, exists in the saturation characteristic graph, and the first transformation point Pt1 between the first range and the second range , there is a second point of change (or saturation point) Pt2 between the first range and the saturation value range.

In this case, the first conversion point Pt1 is a point corresponding to a signal value of about 75% to 85%, preferably about 80% of the saturation value (about 14500), and the amount of X-rays is about 150mR to 250mR, preferably about 200mR point can be set. In addition, the second conversion point Pt2 is a point at which the saturation value starts, and the amount of X-rays may be set at about 300 mR to 400 mR, preferably at 350 mR.

In this case, it may be configured to have a gain characteristic of the first slope S1 having a relatively high slope in the first range, and to have a gain characteristic of the second slope S2 having a relatively low slope in the second range.

Here, the first range may be defined to include an area QA capable of securing the quality of an X-ray image, but is not limited thereto.

The first range is a relatively low X-ray dose range, and since it has a high first slope (S1), that is, a high gain in this range, contrast can be improved.

The second range is a relatively high X-ray dose range, and since it has a low second slope S2, that is, a low gain in this range, the saturation point P2 can be increased, so that the dynamic area can be increased, i.e., widened.

As such, as in this embodiment, an X-ray detector having a saturation characteristic of a linear relationship in which the slope becomes smaller toward the saturation value as two or more different linear relationships between the amount of X-rays and the signal value is used.

Accordingly, it is possible to implement a wide dynamic area and high contrast X-ray image.

Meanwhile, in FIG. 2 , a case in which the first slope S1 is smaller than the conventional slope Sp is illustrated as an example, but the present invention is not limited thereto, and the first slope S1 may be set to be greater than or equal to the conventional slope. In this regard, as shown in FIG. 3 , the first slope S1 may be set to be larger than the conventional slope Sp.

As described above, the saturation characteristic of the X-ray detector of this embodiment is due to the saturation characteristic of each pixel formed in the X-ray detector, and for this purpose, a linear-logarithmic response pixel or a high-dynamic-range pixel A pixel design technique called (HDR) pixel may be applied.

This will be described with reference to FIGS. 4 and 5 .

4 is a diagram schematically illustrating the configuration of an X-ray detector according to an embodiment of the present invention, and FIG. 5 is a diagram schematically illustrating a pixel structure according to an embodiment of the present invention.

4 and 5 , the X-ray detector 140 of this embodiment includes a pixel array panel 210 , a row control circuit 220 and a read out circuit 230 . may include.

A pixel array composed of a plurality of pixels P is disposed on the pixel array panel 210 , and is arranged in a matrix form along a plurality of row lines and a plurality of column lines.

The row control circuit 220 controls the pixel array panel 210 in a row line unit. The row control circuit 220 may sequentially output the selection signal SEL in a row line unit, and in response, the pixel P of the corresponding row line may be selected.

The readout circuit 230 is connected to the signal output wiring RL extending along each column line, and reads out the pixel signal output from the pixel P of the selected row line through the signal output wiring RL. do. The readout circuit 230 may convert the input pixel signal into a digital signal.

The pixel P is a configuration corresponding to a unit sensor that detects X-rays as electrical signals, and may include a photodiode PD.

In addition, a plurality of transistors may be provided in the pixel P. In this embodiment, a pixel P having a 3T structure including three transistors will be described as an example.

In this case, the pixel P may include a reset transistor Tr, a source follow transistor Tsf, and a selection transistor Tse.

The reset transistor Tr performs a reset function for the pixel P. The gate electrode of the reset transistor Tr receives the reset signal RS, the source electrode receives the high potential power voltage VDD as the first power voltage VDD, and the drain electrode is connected to the photodiode PD. can

The source follow transistor Tsf performs a buffer function for a pixel signal, that is, a voltage generated in the node N according to X-ray irradiation. The gate electrode of the source follow transistor Tsf is connected to the node N between the reset transistor Tr and the photodiode PD, the source electrode receives the first power voltage VDD, and the drain electrode of the select transistor (Tse) can be connected.

The selection transistor Tse performs a function of outputting a pixel signal according to the selection signal SEL. The gate electrode of the selection transistor Tse may receive the selection signal SEL, the source electrode may be connected to the drain electrode of the source follow transistor Tsf, and the drain electrode may be connected to the signal output line RL.

The photodiode PD may be configured as a PN junction diode. The anode, which is the first electrode of the photodiode PD, receives a ground voltage as the second power voltage GND, for example, a low potential power voltage, and the second electrode, the cathode, has a reset transistor Tr. may be connected to the drain electrode of

Meanwhile, in the present embodiment, a capacitor Ca connected in parallel to the photodiode PD may be formed in the pixel P. That is, the first electrode of the capacitor Ca is connected to the anode of the photodiode PD to receive the second power voltage GND, and the second electrode of the capacitor Ca is connected to the cathode of the photodiode PD. can

As described above, in the present embodiment, by connecting the capacitor Ca to the photodiode PD of the PN junction structure in parallel, the pixel signal value increases in two different linear relationships depending on the amount of X-rays to realize saturation. .

In this regard, the gain with respect to the pixel signal value, that is, the voltage value at the cathode-side node N of the photodiode PD, of the photodiode PD and the capacitor Ca connected to the node N in parallel It is inversely proportional to the capacity value.

As such, using the PN junction photodiode PD, the capacitor of the photodiode PD in the first range of low dose, that is, the diode capacitor, has a small capacitance, and in the second range of high dose, the diode capacitor is relatively It is configured to have a large capacity value.

This will be described with reference to FIG. 6 showing the C-V curve of the PN junction photodiode PD of the present embodiment.

As shown in FIG. 6 , the photodiode PD of this embodiment is a diode of a PN junction structure, and its operating states are a deep depletion state and a depletion state depending on a voltage value. And, it may have an accumulation state.

Here, the deep-depletion state corresponds to a state of the photodiode PD when an X-ray dose in the first range, which is a relatively low dose, is irradiated, and in this state, the diode capacitor of the photodiode PD has a first capacitance. As a value, it has a minimum capacity value (Cd_min).

The accumulation state corresponds to the state of the photodiode PD when a relatively high dose of X-rays in the second range is irradiated, and in this state, the diode capacitor of the photodiode PD has the maximum capacitance as the second capacitance value. It has a value (Cd_max).

The depletion state is a critical state between the deep depletion state and the accumulation state, and the minimum capacitance value (Cd_min) and the maximum capacitance value (Cd_max) of the diode capacitor can be switched with the depletion state as a boundary. there is. The photodiode PD voltage in such a depletion state is also called a pinning voltage.

As such, in the photodiode PD of this embodiment, the diode capacitor has a minimum capacitance value Cd_min in the first range of low dose, and has a maximum capacitance value Cd_max when switched to the second range of relatively high dose.

Here, since the minimum capacitance value (Cd_min) of the diode capacitor is negligibly small compared to the maximum capacitance value (Cd_max), it is very difficult to realize the first slope, which is a gain required in the first range of low dose.

Accordingly, in order to realize a desired gain in the first range, that is, the first slope, the capacitor Ca having a very large capacitance value compared to the minimum capacitance value Cd_min of the diode capacitor is connected in parallel.

In this way, as the capacitor Ca is connected in parallel to the photodiode PD configured to have two capacitance values separated according to the amount of X-rays, the first slope, which is a high gain in the first range of low dose, and the second range of high dose The second slope, which is a low gain, can be realized.

In this regard, in the first range, a relatively small capacitance value (Ca + Cd_min) is formed to define a relatively large first slope, and in the second range, a relatively large capacitance value (Ca + Cd_max) is formed. formed so that a relatively small second gradient can be defined.

At this time, as mentioned above, since the minimum dose value (Cd_min) is a very small value that is substantially negligible, in the first range of low dose, the capacitance value becomes Ca, and thus a large gain is defined. will be.

As described above, in the present embodiment, by connecting a capacitor in parallel to the photodiode having the PN junction structure, the pixel signal value increases in two different linear relationships depending on the amount of X-rays, thereby realizing a saturation characteristic.

Accordingly, the X-ray detector can implement a wide dynamic area and high-contrast X-ray image, particularly, a wide dynamic area in high illuminance and high-contrast X-ray image in low illuminance.

On the other hand, the embodiment of the present invention, as well as the mammography apparatus can be applied to the X-ray imaging apparatus for other purposes.

The above-described embodiment of the present invention is an example of the present invention, and free modifications are possible within the scope included in the spirit of the present invention. Accordingly, the present invention includes modifications of the present invention provided they come within the scope of the appended claims and their equivalents.

100: body 110: column connection part
120: vertical connection 130: generator
140: X-ray detector 150: compression pad
160: support panel 200: column
210: pixel array panel 220: row control circuit
230: readout circuit
S1, S2: 1st and 2nd slope
P: pixel
PD: photodiode
Ca: capacitor
Tr: reset transistor
Tsf: source follow transistor
Tse: Select transistor

Claims (11)

Including a pixel generating a pixel signal value according to the irradiated X-ray dose,
The pixel comprises a photodiode,
The photodiode exhibits a deep depletion state at an X-ray dose of a first range, and exhibits an accumulation state at an X-ray dose of a second range greater than the first range, having a first dose value for the X-ray dose, and a second dose value different from the first dose value for the X-ray dose in the second range;
The pixel signal value has a linear relationship of a first inclination with respect to the amount of X-rays in the first range, and a linear relationship of a second inclination different from the first inclination with respect to the amount of X-rays in the second range
X-ray detector.
The method of claim 1,
The second capacitance value is greater than the first capacitance value, and the second slope is smaller than the first slope.
X-ray detector.
The method of claim 1,
The photodiode is a PN junction photodiode,
The pixel includes a capacitor connected in parallel with the PN junction photodiode
X-ray detector.
The method of claim 1,
The photodiode has a depletion state between the first and second ranges.
X-ray detector.
delete The method of claim 1,
The pixel is
a reset transistor connected to the cathode of the photodiode;
a source-follow transistor connected to a node between the photodiode and the reset transistor;
Further comprising a selection transistor connected between the source follow transistor and the signal output wiring
X-ray detector.
The method of claim 1,
configured to generate a pixel signal value of 75% to 85% with respect to a saturation value of the pixel signal value at a first conversion point between the first range and the second range.
X-ray detector.
8. The method of claim 7,
The X-ray dose of the first conversion point is 150mR ~ 250mR
X-ray detector.
8. The method of claim 7,
The X-ray dose at the second conversion point between the second range and the range of the X-ray dose at which the saturation value is generated is 300mR to 400mR
X-ray detector.
The X-ray detector of any one of claims 1 to 4, 6 to 9;
A generator that irradiates X-rays to the X-ray detector
An X-ray imaging device comprising a.
11. The method of claim 10,
Compression pad that presses the breast on the X-ray detector
An X-ray imaging device comprising a.

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