KR20200025699A - X-ray Detector and X-ray apparatus - Google Patents

Abstract

Provided is an X-ray detector with more improved characteristics. The X-ray detector includes a pixel that generates a pixel signal value according to the amount of irradiated X-rays, wherein the amount of irradiated X-rays and the pixel signal value are configured to have two or more different linear relationships.

Description

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, an X-ray detector capable of obtaining a wide dynamic range and a high contrast X-ray image, and X-ray imaging using the same Relates to a device.

The Mammography apparatus, which is an example of an X-ray imaging apparatus, is used for early diagnosis of breast cancer or macrocalcification, and irradiates a certain amount of X-rays to a breast of a subject, and detects X-rays transmitted through the breasts. And a two-dimensional X-ray image (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 is implemented.

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 X-ray dose increases linearly with a predetermined slope and then saturates.

When a conventional X-ray detector having such a characteristic is used, the generated X-ray image has a problem of having a narrow dynamic range and low contrast.

It is an object of the present invention 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 a wide dynamic range and high contrast X-ray image, and an X-ray imaging apparatus using the same.

In order to achieve the above object, the present invention includes a pixel for generating a pixel signal value according to the irradiated X-ray dose, wherein the X-ray dose and the pixel signal value is configured to have two or more linear relationships different from each other Provides a line detector.

Here, when the X-ray amount is the first range, the pixel signal value is configured to have a linear relationship of the first slope, and when the X-ray amount is the second range larger than the first range, the pixel signal value is It may be configured to have a linear relationship of the second slope 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 configured of 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 have an accumulation state in the second range.

The pixel may include: a reset transistor connected to the cathode of the photodiode; A source follow transistor coupled to a node between the photodiode and a reset transistor; The transistor may further include a selection transistor connected between the source follower transistor and the signal output line.

The pixel signal value of 75% to 85% may be generated with respect to the saturation value of the pixel signal value at the first conversion point between the first range and the second range.

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

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

In another aspect, the above-described X-ray detector; An X-ray imaging apparatus including a generator for irradiating X-rays to the X-ray detector is provided.

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

In the present invention, by connecting a capacitor to the photodiode of the PN junction structure in parallel, it is possible to implement the characteristic that the pixel signal value is increased and saturated in two different linear relationships according to the X-ray dose.

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

1 is a perspective view of a mammography apparatus according to an embodiment of the present invention.
2 is a graph showing the saturation characteristics of the X-ray detector of the mammography apparatus according to the embodiment of the present invention.
3 is a graph showing another example of the saturation characteristics of the X-ray detector of the mammography apparatus according to the embodiment of the present invention.
4 is a view schematically showing the configuration of an X-ray detector according to an embodiment of the present invention.
5 schematically illustrates a pixel structure according to an embodiment of the invention.
6 illustrates 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, for convenience, the mammography apparatus is described as an example of the X-ray imaging apparatus according to the present invention, but the present invention is not limited thereto.

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

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

The main body 100 has a C or 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 to be able to move up and down 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. It may include.

The generator 130 is mounted at one end of the main body 100, for example, at the upper end, and irradiates X-rays toward the lower end, which is the other end, and the X-ray detector 140 is mounted at the other end of the main body 100, for example, at the lower end. X-rays transmitted through the subject placed in between are received. In addition, the main body 100 may include a compression pad 150 for pressing the breast placed on the X-ray detector 140 or the support panel 160 between the generator 130 and the X-ray detector 140.

In this case, in the mammography apparatus according to the embodiment of the present invention, the entire body 100 may be rotated by a predetermined angle to capture the full-field digital mammography (FFDM) mode, and the upper end of the main body 100 has a predetermined angle relative to the lower end. It can rotate to take a digital breast tomosynthesis (DBT) or breast tomosynthesis (BTS) mode.

On the other hand, the X-ray detector 140 obtains image information according to the intensity of each position of the X-rays transmitted through the breast. The X-ray detector 140 for this purpose is a form in which a plurality of sensors are two-dimensionally arranged in units of pixels, and a unit pixel sensor forms a plurality of matrices and is 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 a wide dynamic range and a high contrast X-ray image 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 the embodiment of the present invention, together with the 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 the signal value detected by the X-ray detector. Here, the signal value is an offset signal value having a value of zero when the X-ray dose is zero, and is a digital value.

As shown in FIG. 2, the conventional X-ray detector has a characteristic that the signal value increases linearly as the X-ray dose increases and the signal value is saturated above a specific X-ray dose point Ptr. For example, in the drawing, for an X-ray dose in the range of about 130 mR from the origin, the signal value is linearly increased to a single slope Sp (that is, a single gain) to about 13000 and then saturated.

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

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

As such, when the saturation characteristics of two different linear relations are defined, a saturation characteristic graph has a transformation point, which is a point at which the slope is changed, and the first transformation point Pt1 between the first range and the second range and , 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 X-ray dose is about 150 mR to 250 mR preferably about 200 mR. Can be set to a point. In addition, the second conversion point Pt2 may be set to a point at which the saturation value starts, and the X-ray dose is approximately 300 mR to 400 mR, preferably 350 mR.

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

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

Since the first range is a relatively low X-ray dose range and has a high first slope S1, that is, a high gain, the contrast can be improved.

Since the second range is a relatively high X-ray dose range and 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 range can be increased or widened.

As described above, as in the present 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 X-ray dose and the signal value is used.

As a result, a wide dynamic range and high contrast X-ray image can be realized.

Meanwhile, FIG. 2 illustrates a case where the first slope S1 is smaller than the conventional slope Sp, but is not limited thereto, and the first slope S1 may be set to be equal to or more than the conventional slope. In this regard, as shown in FIG. 3, the first slope S1 may be set larger than the conventional slope Sp.

The saturation characteristic of the X-ray detector of the present embodiment as described above 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 a 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 the present embodiment includes a pixel array panel 210, a row control circuit 220, and a read out circuit 230. It may include.

The pixel array panel 210 includes a pixel array including a plurality of pixels P, which are arranged in a matrix 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 sequentially outputs the selection signal SEL in units of row lines, and in response thereto, the pixels P of the corresponding row lines may be selected.

The readout circuit 230 is connected to the signal output line 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 line RL. do. The readout circuit 230 may convert the input pixel signal into a digital signal.

The pixel P corresponds to a unit sensor that detects X-rays as an electrical signal, and may include a photodiode PD.

In addition, a plurality of transistors may be provided in the pixel P. In the present embodiment, a pixel P having a 3T structure including three transistors is 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 supply voltage VDD, and the drain electrode is connected to the photodiode PD. Can be.

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

The selection transistor Tse outputs 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 follower transistor Tsf, and the drain electrode may be connected to the signal output line RL.

The photodiode PD may be composed of a PN junction diode. An anode, which is a first electrode of the photodiode PD, receives a ground voltage as a low potential power supply voltage, for example, as a second power supply voltage GND, and a cathode, which is a second electrode, is a reset transistor Tr. It can 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 supply voltage GND, and the second electrode of the capacitor Ca is connected to the cathode of the photodiode PD. Can be.

As described above, in the present embodiment, the capacitor Ca is connected to the photodiode PD of the PN junction structure in parallel, thereby realizing a characteristic that the pixel signal value is increased and saturated in two different linear relations according to the X-ray dose. .

In this regard, the gain of the pixel signal value, that is, the voltage value, at the cathode side node N of the photodiode PD is determined by the photodiode PD and the capacitor Ca of the parallel relationship connected to the node N. It is inversely proportional to the dose value.

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

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

As shown in FIG. 6, the photodiode PD of the present embodiment is a diode having a PN junction structure, and its operating state is a deep depletion state and a depletion state according to a voltage value. And, it may have an accumulation state (accumulation).

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

The accumulation state corresponds to the photodiode (PD) state in the case where the X ray dose of the second range, which is a relatively high high dose, is irradiated. In this state, the diode capacitor of the photodiode PD is the maximum capacitance as the second capacitance value. It has a value Cd_max.

The depletion state is a threshold 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 may be switched on the basis of the depletion state. have. The photodiode (PD) voltage in this depletion state is also called a pinning voltage.

As described above, the photodiode PD of the present embodiment has a maximum capacitance value Cd_max when the diode capacitor has a minimum capacitance value Cd_min in the first range of low dose and is converted to a second range of relatively high dose.

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, a 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.

As such, when the capacitor Ca is connected in parallel to the photodiode PD configured to separate two capacitance values according to the X-ray dose, the first slope, which is high gain in the first range of low dose, and the second range of high dose, The low gain second slope can be realized.

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

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

As described above, in this embodiment, by connecting a capacitor to a photodiode having a PN junction structure in parallel, a pixel signal value increases and saturates in two different linear relationships according to X-ray dose.

Accordingly, the X-ray detector can realize a wide dynamic range and high contrast X-ray image, particularly a high dynamic range at high illumination and a high contrast X-ray image at low light.

On the other hand, the embodiment of the present invention can be applied not only to the mammography apparatus but also to the X-ray imaging apparatus for other purposes.

Embodiment of the present invention described above is an example of the present invention, it is possible to change freely within the scope included in the spirit of the present invention. Accordingly, the invention includes modifications of the invention within the scope of the appended claims and their equivalents.

100: main body 110: column connection portion
120: vertical connection portion 130: generator
140: X-ray detector 150: pressure pad
160: support panel 200: column
210: pixel array panel 220: row control circuit
230: lead-out circuit
S1, S2: 1st, 2nd Slope
P: pixels
PD: photodiode
Ca: Capacitor
Tr: Reset Transistor
Tsf: SourceFollowersTransistors
Tse: Select transistor

Claims (11)

A pixel generating a pixel signal value according to the irradiated X-ray dose,
Configured to have two or more linear relationships between the X-ray dose and the pixel signal value
X-ray detector.
The method of claim 1,
When the X-ray dose is in the first range, the pixel signal value is configured to have a linear relationship of the first slope,
The pixel signal value having a linear relationship of a second slope smaller than the first slope, when the X-ray amount is a second range greater than the first range
X-ray detector.
The method of claim 2,
The pixel,
A photodiode;
A capacitor connected in parallel with the photodiode
X-ray detector.
The method of claim 3, wherein
The photodiode is composed of a PN junction diode, and has a first capacitance value in the first range and a second capacitance value greater than the first capacitance value in the second range.
X-ray detector.
The method of claim 4, wherein
The photodiode has a deep depletion state in the first range, and has an accumulation state in the second range.
X-ray detector.
The method of claim 3, wherein
The pixel,
A reset transistor coupled to the cathode of the photodiode;
A source follow transistor coupled to a node between the photodiode and a 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 2,
A pixel signal value of 75% to 85% is generated 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.
The method of claim 7, wherein
X-ray dose of the first conversion point is 150mR ~ 250mR
X-ray detector.
The method of claim 7, wherein
The X-ray dose of the second conversion point between the second range and the range of X-ray dose at which the saturation value is generated is 300 mR to 400 mR.
X-ray detector.
An X-ray detector of any one of claims 1 to 9;
Generator for irradiating X-rays to the X-ray detector
X-ray imaging apparatus comprising a.
The method of claim 10,
Compression pads for pressing the breast on the X-ray detector
X-ray imaging apparatus comprising a.

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