JP2007071602A - Radiation detector - Google Patents

Abstract

<P>PROBLEM TO BE SOLVED: To provide a radiation detector which can detect radiation with a high count rate and collect energy information as well. <P>SOLUTION: An X-ray detector 3 comprises a matrix M made of silicon, which is a semiconductor, and first to fourth electrodes 21 to 24 for outputting a current developed in the matrix M. The matrix M is formed in a plate shape extending in the incident direction of filtered X-rays and is provided with a ground side electrode E on its one end. The first to fourth electrodes 21 to 24 are arranged in order along the incident direction of the filtered X-rays on the top side of the matrix M. <P>COPYRIGHT: (C)2007,JPO&INPIT

Description

  The present invention relates to a radiation detector used for radiation transmission imaging and the like.

  X-rays and gamma rays, which are electromagnetic radiation (hereinafter simply referred to as radiation), have a high transmission capability, so they are widely used for non-destructive inspection and crystal structure analysis in the industrial field, including diagnosis in the medical field. Yes. In recent years, as a radiation transmission imaging method using the radiation transmission capability, instead of X-ray imaging using a photosensitive action on a photographic plate, the radiation excitation action on a detection medium is electrically detected, and the digital result is obtained based on the detection result. What gets images has become mainstream. When radiation energy is applied to the inside of the radiation detector, depending on the type of detection medium, for example, an electron / hole pair for a semiconductor, an electron / ion pair for a gas, or an scintillator In the case of fluorescence and superconductors, excitons such as quasi-electrons are generated. In the radiation detector, a voltage proportional to the applied energy is induced by the movement of these excitons to the electrode, and the energy of the radiation transmitted through the human body or the like can be measured.

  However, when the radiation detector described above is used, the induced voltage is measured every time the radiation is incident. Therefore, when the next radiation is incident while the excitons are moving to the electrode, There is a risk of recognizing it as a single radiation. As a result, the number of counts of radiation incident per unit time (that is, the count rate) is limited, and a very high count is achieved in a short time as in a conventional radiographic imaging method (such as an X-ray CT scan). When measuring the rate, there is a problem that the energy information of radiation cannot be used. Therefore, in a conventional radiation transmission imaging apparatus, a method of measuring energy per unit time given from radiation by detecting a current flowing through an electrode has been adopted (see Patent Document 1).

On the other hand, when an image for determining the presence or absence of cancer tissue is obtained using an iodine contrast agent, only a small part of the irradiated X-rays are absorbed by iodine, so it is necessary to irradiate a large amount of X-rays on the human body. There was a problem of increasing the risk of radiation exposure. In view of this, a low-exposure type X-ray transmission imaging apparatus has been developed that employs an energy difference method that utilizes the difference in X-ray absorption between the human body and iodine. In this type of X-ray transmission apparatus, those using white X-rays and those using two types of monochromatic X-rays are generally used. However, in the past proposed by the present applicant, a lantern filter (La filter) is used. ) Irradiate the human body with filter X-rays from which the high-energy portion has been cut, and measure the filter X-rays transmitted through the human body in the upper and lower energy ranges with the K shell absorption edge of iodine in between. By subtracting the two types of energy information, an image in which only the contrast medium is emphasized is obtained (see Patent Document 2).
JP-A-2005-77152 JP 2004-223158 A

  The radiation detector of Patent Document 1 can cope with a high counting rate, but information on energy applied from individual radiation is lost. For example, for conventional X-ray transmission imaging, a certain tissue is used. Only information on whether the amount of X-rays that passed through is large or small can be obtained. Therefore, when such a radiation detector is applied to the energy difference method of Patent Document 2, measurement with a high counting rate cannot be performed with a small amount of radiation, such as an X-ray CT scan, and the like It was extremely difficult to obtain a clear image of a fast-moving organ.

  The present invention has been made in view of such a background, and an object thereof is to provide a radiation detector capable of collecting energy information while detecting radiation with a high count rate.

  The radiation detector according to the invention of claim 1 is a detection medium that generates a charge by energy applied from incident radiation, and the detection medium at a position where the distance from the radiation incident end of the detection medium is different from each other. And a plurality of electrodes installed in the.

  According to a second aspect of the present invention, in the radiation detector according to the first aspect, the detection medium is a solid, and a groove is formed between the electrodes in the detection medium.

  According to a third aspect of the present invention, in the radiation detector according to the first or second aspect, the detection medium is a semiconductor, and the detection medium includes a semiconductor layer having a polarity different from that of the semiconductor between the electrodes. Is formed.

  According to a fourth aspect of the present invention, in the radiation detector according to the first aspect, the detection medium is divided between the electrodes.

  According to the radiation detector of claim 1, since the radiation incident on the detection medium travels in the detection medium while being attenuated by releasing energy, a high count rate is obtained by measuring the current flowing through each electrode. Radiation detection and energy information collection can be performed simultaneously. Moreover, according to the radiation detector of Claims 2-4, the insulation between each electrode improves and a more accurate detection is implement | achieved.

Hereinafter, some embodiments of a radiation detector to which the present invention is applied will be described in detail with reference to the drawings.
[First Embodiment]
FIG. 1 is a schematic configuration diagram of the X-ray inspection apparatus according to the first embodiment, and FIG. 2 is a perspective view of the X-ray detector according to the first embodiment. FIG. 3 is a graph showing an X-ray energy spectrum according to the first embodiment, and FIGS. 4 and 5 are graphs showing current values of the respective electrodes according to the first embodiment.

<< Configuration of First Embodiment >>
<X-ray inspection equipment>
As shown in FIG. 1, an X-ray inspection apparatus 1 according to the first embodiment includes an X-ray tube 2 and an X-ray detector array 4 in which X-ray detectors (radiation detectors) 3 are arranged vertically and horizontally. It comprises a preamplifier 5, a main amplifier 6, two types of integrators 7, 8, a contrast agent thickness calculator 9, an imaging device 10, and the like.

<X-ray detector>
As shown in FIG. 2, the X-ray detector 3 according to the first embodiment includes a base material M made of silicon, which is a semiconductor, and first electrodes 21 to 21 for outputting a current generated in the base material M. And a fourth electrode 24. The base material M has a rectangular plate shape extending along the incident direction of the filter X-ray, and a ground side electrode E is provided at one end thereof. In addition, the first electrode 21 to the fourth electrode 24 are arranged on the upper surface of the base material M in order along the incident direction of the filter X-ray, and the currents I1 to I4 output from the respective electrodes cause the diode D to be arranged. To the preamplifier 5 described above.

<< Operation of First Embodiment >>
When the X-ray inspection apparatus 1 is activated and the subject X is irradiated with the filter X-rays from the X-ray tube 2, the filter X-rays transmitted through the subject S are X-ray detectors 3 in the X-ray detector array 4. Is incident on. In the X-ray tube 2 of the present embodiment, electrons accelerated to 50 kV are collided with a tungsten target, and a high energy portion (38.9 keV or more) is removed from the emitted white X-rays by a La filter. Is generated.

When filtered X-rays enter the X-ray detector 3, exciton electron / hole pairs are generated in the base material M, and current I _{1 is} supplied from the first electrode 21 to the fourth electrode 24 to the preamplifier 5. ~ I ₄ is output. The currents I _{1 to} I ₄ are amplified by the preamplifier 5 and the main amplifier 6, integrated by the integrators 7 and 8 for each energy region, and output to the contrast agent thickness calculator. In the contrast agent thickness calculator 9, the thickness of the iodine contrast agent in the subject S is calculated based on the integration results of the integrators 7 and 8, and the imaging device 10 generates a transmission X-ray image based on the calculation result. .

When the subject S in which iodine contrast medium is injected into a lesion or blood vessel such as cancer is irradiated with filter X-rays, as shown in FIG. 3, the energy level (33.2 keV) around the K shell absorption edge of iodine is discontinuous. An X-ray energy spectrum is obtained. FIG. 3 is a graph of an X-ray energy spectrum obtained by irradiating filter X-rays into an aqueous layer having a thickness of 20 cm imitating a human body and containing iodine imitating an iodine contrast agent. It shows four energy range _E 1 to E ₄ of the lines and the number _Y 1 to Y ₄ of the X-rays in these energy range _E 1 to E _4. In addition, in the patent document 2 cited above, the applicant of the present invention is based on the ratio of the number of X-rays Y ₂ and Y ₃ included in the energy ranges E ₂ and E ₃ in FIG. It has been disclosed that it is possible to determine the thickness.

式１において、ｗは一つの電子・正孔対を生成するために必要なエネルギであり、Ｔは測定時間であるが、以降の説明においては、説明が煩雑になることを防ぐため、ｗおよびＴはともに１とする（考慮しない）。また、Ｅ_ｊ（ｊ＝１，４）は各エネルギ範囲の平均エネルギであり、Ｃ_ｉｊ（ｉ＝１，４，ｊ＝１，４）は平均エネルギＥ_ｊのＸ線が母材Ｍの第ｉ電極（２１〜２４）に対応する部位で吸収される相対確率を示す。 As described above, by measuring the induced voltage due to the incidence of one radiation, it is possible to determine in which energy range the radiation is included. In the case of using the radiation detector which measures by the above method, the energy of individual X-rays cannot be measured correctly by the current output from a single electrode. That is, in FIG. 2, since the X-rays in the entire energy range E _{1 to} E ₄ contribute to the current I ₁ output from the first electrode 21, when the current is measured only with the first electrode 21. Of course, it is not possible to measure how many X-rays in which energy range exist (Equation 1).

In Equation 1, w is the energy required to generate one electron / hole pair, and T is the measurement time. In the following description, in order to prevent the explanation from becoming complicated, w and T is 1 (not considered). E _j (j = 1, 4) is the average energy of each energy range, and C _ij (i = 1, 4, j = 1, 4) is the X-ray of the average energy E _j of the base material M. The relative probability of being absorbed in the site | part corresponding to i electrode (21-24) is shown.

式２から判るように、電流Ｉ_１〜Ｉ_４は、母材Ｍ（シリコン）の各電極２１〜２４に対応する部位で吸収される数に依存している。ここで、Ｅ_ｊ（ｊ＝１，４）は既知の値であるが、Ｙ_ｉ（ｉ＝１，４）およびＣ_ｉｊ（ｉ＝１，４，ｊ＝１，４）は未知数であり、式が４つで未知数が合計２０個となるため、そのままでは式２の連立方程式を解くことができない。 Formula 2 shows currents I _{1 to} I ₄ output from the first electrode 21 to the fourth electrode 24.

As can be seen from Equation 2, the currents I _{1 to} I ₄ depend on the number absorbed at the portions corresponding to the electrodes 21 to 24 of the base material M (silicon). Here, E _j (j = 1, 4) is a known value, but Y _i (i = 1, 4) and C _ij (i = 1, 4, j = 1, 4) are unknowns, Since there are four equations and a total of 20 unknowns, the simultaneous equations of Equation 2 cannot be solved as they are.

ここで、μ（Ｅ_ｊ）はエネルギに依存する減衰係数である。また、Ｘ_１−２は第１電極２１と第２電極２２との間の距離、Ｘ_２−３は第２電極２２と第３電極２３との間の距離、Ｘ_３−４は第３電極２３と第４電極２４との間の距離である。 By the way, the intensity of X-rays decreases exponentially when passing through a substance (Equation 3).

Here, μ (E _j ) is an attenuation coefficient depending on energy. _X1-2 is a distance between the first electrode 21 and the second electrode 22, _X2-3 is a distance between the second electrode 22 and the third electrode 23, and _X3-4 is a third electrode. 23 and the fourth electrode 24.

Therefore, C _ij can be obtained by previously obtaining the absorption probability in each electrode segment when a photon of energy E _j is incident from the simulation code of the interaction between the photon and the substance. As a result, the simultaneous equations of Expression 2 include _four unknowns Y ₁ to Y 4 and can be solved (Expression 4).

Specific calculation examples are given below.
The representative values of the _four energy ranges E _{1 to} E ₄ in FIG. 3 are 26.1 keV, 31.7 keV, 37.5 keV, and 44.5 keV, respectively, and the size of the X-ray detector 3 shown in FIG. Both have a width of 6 mm, a thickness of 2 mm, and a total length of 11 mm (2 mm long electrodes installed at 1 mm intervals) and a total length of 19 mm (4 mm long electrodes installed at 1 mm intervals). The test was performed with the thing (B). As a result, the graph shown in FIG. 4 was obtained in the former (A), and the graph shown in FIG. 5 was obtained in the latter (B). 4 and 5, in both cases, the rate at which 26.1 keV X-rays are absorbed at the portion corresponding to the first electrode 21 is normalized to 1.

Based on the results of FIGS. 4 and 5, when the determinant of 4 rows and 4 columns of Equation 4 is C, Equation 5 corresponding to (A) and Equation 6 corresponding to (B) are obtained.

  This makes it possible to solve Equation 4 (that is, to know how many X-rays in which energy range exist), so that the contrast agent thickness calculation device 9 calculates the iodine contrast agent thickness and the image. The transmission apparatus 10 can generate a transmission X-ray image.

<< Effects of First Embodiment >>
Thus, in the first embodiment, since the energy spectrum of the filter X-ray can be obtained while performing current measurement, the amount of filter X-ray irradiation per unit time can be increased, and clear transmission X can be obtained in a short time. A line image could be obtained. Further, if the number of electrodes provided in the X-ray detector 3 is increased, a more detailed energy spectrum can be obtained (that is, a clearer transmitted X-ray image can be obtained). In addition, after measuring the current with a certain number of electrodes, changing the response function (determinant C) obtained by calculation in advance, the measured energy range can be obtained without replacing the X-ray detector 3 or the like. Can be changed freely.

≪Comparison with conventional products≫
In a conventional X-ray inspection apparatus, an X-ray detector having a single electrode is used to detect a current generated by X-rays in the entire energy region shown in FIG. The energy spectrum of the filter X-ray that has passed through the normal tissue has almost no absorption by the iodine contrast agent, and the energy spectrum of the filter X-ray that has passed through the lesion portion containing a large amount of iodine contrast agent has an iodine as shown in FIG. Since the portion corresponding to the energy level (33.2 keV) at the K-shell absorption edge of is absorbed, a significant decrease in the measured current is observed in the lesion portion. This can be expressed as follows:

となり、病巣部分を通過したフィルタＸ線による電流値Ｉ’は、

となるため、それぞれの電流値Ｉ，Ｉ’の差ΔＩは、

となる。 First, the current value I by the filter X-ray that has passed through the normal tissue is

The current value I ′ by the filter X-ray that has passed through the lesion is

Therefore, the difference ΔI between the current values I and I ′ is

It becomes.

となるため、従来のエネルギ差分法における信号雑音比Ｓ／Ｎは、式１１で与えられる。

Considering the current quantum noise, the signal-to-noise ratio S / N is

Therefore, the signal-to-noise ratio S / N in the conventional energy difference method is given by Equation 11.

On the other hand, when the X-ray detector of the present embodiment is used, the current value and the energy information can be obtained at the same time, and therefore the number of X-rays Y _{1 to} Y in each energy range E _{1 to} E ₄ from the measured current value. ₄ can be identified, and Equation 11 can be an energy representation (Equation 12).

とすることが可能となり、信号雑音比Ｓ／Ｎの分母が小さくなる分だけ信号が雑音に埋もれ難くなる（すなわち、有意な信号を得やすい）。 Here, since the current values by the X-rays in the energy ranges E ₁ and E ₄ where there is almost no absorption by the iodine contrast agent are almost the same between the case of passing through the normal tissue and the case of passing through the lesion part, Y ₁ ′ = Y ₁ and Y ₄ ′ = Y ₄ . Moreover, Y ₁ E ₁ and Y ₄ E _{4 that} are constant amounts regardless of the presence or absence of an iodine contrast agent can be set to 1. As a result, the signal-to-noise ratio S / N is

And the signal is less likely to be buried in the noise (ie, it is easy to obtain a significant signal) as the denominator of the signal / noise ratio S / N becomes smaller.

  Thus, according to the present embodiment, an equivalent transmission X-ray image is obtained by using a smaller amount of filter X-rays compared to a conventional X-ray inspection apparatus using a radiation detector having a single electrode. It was possible to achieve low exposure effectively.

[Second Embodiment]
FIG. 6 is a perspective view of an X-ray detector according to the second embodiment.
As shown in the figure, the X-ray detector 3 of the second embodiment has the same overall configuration as that of the first embodiment described above, but the grooves that partition the electrodes 21 to 24 on the upper surface of the base material M. 31 is engraved. Thus, in this embodiment, it improves the electrical insulation between the electrodes 21 to 24 were improved output accuracy of the current I ₁ ~I _4.

[Third Embodiment]
FIG. 7 is a perspective view of an X-ray detector according to the third embodiment.
As shown in the figure, the X-ray detector 3 of the third embodiment is the same in overall configuration as that of the first embodiment described above, but the base material M and the polarity are formed in the portions that partition the electrodes 21 to 24. Semiconductor layers 32 (or p layers if the base material M is an n layer) are formed by ion implantation or thermal diffusion. Thus, in the present embodiment, the electrical insulation is further improved between the electrodes 21 to 24 by the depletion layer, the output accuracy of the current I ₁ ~I ₄ is improved.

[Fourth Embodiment]
FIG. 8 is a perspective view of an X-ray detector according to the fourth embodiment.
As shown in the figure, the X-ray detector 3 of the fourth embodiment is similar in overall configuration to the first embodiment described above, but the first mother is divided along the incident direction of the filter X-ray. The material M1-the 4th base material M4 are provided, and the 1st electrode 21-the 4th electrode 24 are each installed in the upper surface of these base materials M1-M4. Thus, in the present embodiment, the electrical insulation is further improved between the respective electrodes 21 to 24, the output accuracy of the current I ₁ ~I ₄ is improved.

[Fifth Embodiment]
FIG. 9 is a perspective view of an X-ray detector according to the fifth embodiment.
As shown in the figure, the X-ray detector 3 of the fifth embodiment arranges the first base material M1 to the fourth base material M4 filled with gas in series along the incident direction of the filter X-ray, The first electrode 21 to the fourth electrode 24 are installed on the base materials M1 to M4.

[Sixth Embodiment]
FIG. 10 is a perspective view of an X-ray detector according to the sixth embodiment.
As shown in the figure, the X-ray detector 3 of the sixth embodiment is installed on a plate-like scintillator (insulator) SC extending along the incident direction of the filter X-ray and on the upper surface of the scintillator 41. The first photomultiplier tube 41 to the fourth photomultiplier tube 44 are provided as the formed electrodes.

[Seventh Embodiment]
FIG. 11 is a perspective view of an X-ray detector according to the seventh embodiment.
As shown in the figure, the X-ray detector 3 of the seventh embodiment is similar in overall configuration to the sixth embodiment, but is divided into first scintillators SC1 to SC1 divided along the incident direction of the filter X-ray. A fourth scintillator SC4 is provided, and a first photomultiplier tube 41 to a fourth photomultiplier tube 44 are respectively installed on the upper surfaces of the scintillators SC1 to SC4.

  Although the description of the specific embodiment is finished as described above, the present invention is not limited to the above embodiment and can be widely modified. For example, although the above embodiment is an application of the present invention to a medical X-ray inspection apparatus using the energy difference method, it can also be applied to an industrial X-ray inspection apparatus, a gamma ray inspection apparatus, or the like. In this case, CdTe (cadmium telluride) or the like may be employed as the base material of the radiation detector according to the energy of X-rays or gamma rays to be detected. In the above embodiment, the number of electrodes and photomultiplier tubes is four, but it may be two, three, or five or more. In addition, the specific configuration and the like of the radiation detector can be appropriately changed without departing from the spirit of the present invention.

1 is a schematic configuration diagram of an X-ray inspection apparatus according to a first embodiment. It is a perspective view of the X-ray detector which concerns on 1st Embodiment. It is a graph which shows the X-ray energy spectrum which concerns on 1st Embodiment. It is a graph which shows the electric current value of each electrode which concerns on 1st Embodiment. It is a graph which shows the electric current value of each electrode which concerns on 1st Embodiment. It is a perspective view of the X-ray detector which concerns on 2nd Embodiment. It is a perspective view of the X-ray detector which concerns on 3rd Embodiment. It is a perspective view of the X-ray detector which concerns on 4th Embodiment. It is a perspective view of the X-ray detector which concerns on 5th Embodiment. It is a perspective view of the X-ray detector which concerns on 6th Embodiment. It is a perspective view of the X-ray detector which concerns on 7th Embodiment.

Explanation of symbols

1 X-ray inspection device 2 X-ray tube 3 X-ray detector (radiation detector)
21 1st electrode 22 2nd electrode 23 3rd electrode 24 4th electrode 31 Groove 32 Semiconductor layer 41 1st photomultiplier tube 42 2nd photomultiplier tube 43 3rd photomultiplier tube 44 4th photomultiplier tube M Base material M1 1st 1 base material M2 2nd base material M3 3rd base material M4 4th base material SC scintillator SC1 1st scintillator SC2 2nd scintillator SC3 3rd scintillator SC4 4th scintillator

Claims (4)

A detection medium that generates charge by energy imparted from incident radiation;
A radiation detector comprising: a plurality of electrodes installed on the detection medium at positions where the distances from the radiation incident end of the detection medium are different from each other.   The radiation detector according to claim 1, wherein the detection medium is a solid, and a groove is formed between the electrodes in the detection medium in order to improve insulation between the electrodes.   The detection medium is a semiconductor, and a semiconductor layer having a polarity different from that of the semiconductor is formed between the electrodes in order to improve insulation between the electrodes. Or the radiation detector of Claim 2.   The radiation detector according to claim 1, wherein the detection medium is divided between the electrodes in order to improve insulation between the electrodes.

Images