JP2004325183A - Radiation detection method, radiation detector, and radiation imaging system with this detector loaded thereon - Google Patents

Abstract

<P>PROBLEM TO BE SOLVED: To provide a radiation detector of quantum counting type which can attain high contrast, high detection sensitivity and small-sizing. <P>SOLUTION: The radiation detector 1 is provided with a sensor part S and a processing circuit C for every collection pixel. The sensor part S has an incident window of a size which the occurrence of superposition phenomenon between pulse signals responding to each introduction when radiation particles continuously come in a plurality of times at the same position or the vicinity, can be practically negligible. Responding to each of the radiation particles introducing in the incident window, electric signal is generated. The processing circuit C is provided with a charge amplifier 41 individually detecting electric signal from each of a plurality of pixels, a waveform shaping circuit 42 individually adjusting gain of the electric signal and offset detected by the charge amplifier 41, circuits 43, 44 splitting energy of radiation particles that electric signal output from the waveform shaping circuit 42 indicates, and a counter 45 for counting each number of the split radiation particles. <P>COPYRIGHT: (C)2005,JPO&NCIPI

Description

【００６０】
カウンタ１〜３の総ビット数は４ビット、１５ｃｏｕｎｔまで計測できれば十分で、少し各カウンタ構成に自由度を持たせ、実際は３ビットで、計９ビットあれば十分である。
【００６１】
このため、上述の▲６▼の条件より９ビットのカウント値は、１ビット当り、１３０ｎｓｅｃ／９≒１４．５ｎｓｅｃの転送時間でシリアル信号として出力することができる。１収集画素につき、１ｍｓｅｃでカウントリセットを繰り返して収集される。この収集において、全カウント値の出力に要する時間は１３０ｎｓｅｃと短く、この１３０ｎｓｅｃの期間は、デジタルノイズがチャージアンプなどのアナログ系に飛び込むことを防ぐために、カウント計測を停止したとしても、その影響は無視することができる。
【００６２】
また、上述した▲７▼の基準値ＴＨ１〜３の情報は同一ラインを使ってデータ収集前にプリセットできる構成にしておけば回路構成は簡素にできる。さらに、Ｘ線検出器１２は、検出器モジュールを稠密に並べることでＣＴスキャナ用検出器を構成する場合は、シリアル信号の読出す単位を、各検出器モジュールとする構成にすることができる。
【００６３】
（設計例２）
次に、Ｘ線粒子の入射に応答した電気パルス信号の数を「正確に」（すなわち、入射粒子間の重畳現象の発生を実質的に無視可能な状態で）計測するための収集画素のサイズの設計例を説明する。
【００６４】
【外２】

【００６５】
Ｘ線ＣＴスキャナにおいて、パルス計測モードによる再構成の事例は報告が無いので、ＳＰＥＣＴ（Ｓｉｎｇｌｅ Ｐｈｏｔｏｎ Ｅｍｉｓｓｉｏｎ Ｃｏｍｐｕｔｅｄ Ｔｏｍｏｇｒａｐｈｙ）の場合の収集画素サイズを先に計算する。通常、ＳＰＥＣＴでは４ｍｍ×４ｍｍを収集画素とした場合、２０ｋｃｐｓ程度の情報で３０分程度収集を行わないと、良好な再構成画像が得られない。このために必要なトータルカウントは、２００００×３０×６０＝３６Ｍｃｏｕｎｔｓである。
【００６６】
このときの収集画素サイズは１６ｍｍ^２なので、Ｘ線ＣＴスキャナの１．５ｍｍ^２に換算すると、
【数１】

となり、３８４Ｍｃｏｕｎｔｓのカウントが必要になる。
【００６７】
また、Ｘ線ＣＴスキャナの場合、図１１に模式的に示すように、人体を通過しないＸ線も検出器に入射するため、その部分は１００倍程度のＸ線強度になると仮定できる。この仮定を考慮すると、最低でも３８４００ＭｃｏｕｎｔｓのカウントがＸ線粒子が各収集画素に入ってくることを想定する必要がある。
【００６８】
いまの条件によれば、収集時間は０．５ｓｅｃなので、Ｘ線検出器全体としては、
【数２】
３８４００×２＝７６８００Ｍｃｐｓ
のカウントが必要になる。これを１収集画素に換算すると、
【数３】

となる。このことから、パルス計測を１０ｋｃｐｓ程度に抑えるためには、
【数４】

の面積に抑える必要がある。したがって、各収集画素のサイズは、
【数５】

となる。ゆえに、約２００μｍの収集画素サイズに設定しないと、いわゆる「正確な」パルス数の計測は難しいと結論できる。
【００６９】
以上の理由により、本発明では、Ｘ線の入射条件が最も厳しく、収集画素サイズの大小によってのみ「正確」なパルス計測を行おうとするときの、収集画素サイズの上限は約２００μｍである。この上限値は、Ｘ線の入射条件によって変わるもので、かかる条件がもっと緩やかなときには、収集画素サイズの上限ももっと大きくても構わない。
【００７０】
（作用効果）
以上のように、本実施形態に係るＸ線検出器１２によれば、従来の単なるパルス計測モードとは異なり、Ｘ線粒子それぞれが有するエネルギ量で考慮したパルス計測モードでＸ線を検出できる。このパルス計測モードでの検出動作を可能にする主な要因としては、各収集画素に時間的に近接して入射するＸ線粒子のパルス信号が相互に重畳する現象の発生を無視し得る収集画素サイズを採用していること、Ｘ線粒子の入射情報を一度、そのエネルギ量に置き換えて弁別することが挙げられる。
【００７１】
したがって、本実施形態に係るＸ線検出器１２及びこれを搭載したＸ線ＣＴスキャナ１によれば、散乱線の影響が少なくコントラストの明瞭なＣＴ画像を得ることができ、Ｘ線検出器１２及びその後段のチャージアンプ４１から発生する暗電流の影響を回避して、Ｓ／Ｎを大幅に改善した高感度なＸ線検出を行うことができる。さらに、Ｘ線検出感度が高いため、Ｘ線管の管電流を下げて同等又はそれ以上の精度の検出を行うことができるため、被検者のＸ線被曝線量を減らすことができる。また、このようにＸ線管の管電流を下げて使うことで、Ｘ線管の故障も少なくなり、Ｘ線ＣＴスキャナの長寿命化を図ることができる。さらに、従来の積分モードで動作させるＸ線検出器のように大形化及び高コスト化の要因であったＡ／Ｄ変換器が不要になるため、Ｘ線検出器１２の全体として、その構成が簡素化されるとともに、トータルの製造コストも下げることができる。このため、Ｘ線検出器の更なる大形化も可能になる。
【００７２】
さらに、本実施形態に係るＸ線検出器１２は、その半導体セル（センサ部）ＳとしてＣｄＴｅ半導体を用いているので、これにより高精度な検出能も確保される。この検出能について説明する。
【００７３】
現状、Ｘ線を粒子とみなし、それを高速に波形整形した上で、暗電流のレベルを考慮し、最低の閾値レベルを１０ｋｅＶ前後に設定するには、擬似Ｓｃｈｏｔｔｋｙ型のＣｄＴｅが最適である。特に検出器をモノリシック（Ｍｏｎｏｌｉｔｈｉｃ）構造で構成し、ＣｄＴｅの厚みを０．５ｍｍ〜１．０ｍｍで構成すれば従来のＣｓＩ、Ｇｅ、Ｓｉ、Ｘ線フィルムなどと比較しても、優れたＸ線の検出感度を得ることができる。また、この厚みの場合、例えばガンマ線源である１２２Ｃｏでのエネルギ分解能は８０％以下であり、きわめて良好にエネルギの閾値検出を行うことができる。また、ＣｄＴｅは直接Ｘ線を電荷に変換するため、ＣｓＩのようにフォトダイオードで光電変換する必要がないため、変換ロスが少ない。
【００７４】
これに対し、従来のようにシンチレータを用いる場合、収集画素を形成するには、シンチレータを画素単位に切り、しかもその収集画素間に反射剤を塗り、シンチレーション光を相互にアイソレーションする必要がある。このため、このアイソレーションを無視することができず、デットスペースができる一方で、フォトダイオードに到達する光に一定のロスが生じて、感度低下する原因になる。
【００７５】
これらを総合すると、ＣｄＴｅは他のＸ線検出器に比べ、検出感度の優れた素材といえる。しかも、閾値検出により、設定した最低レベル以下のエネルギは検出しないようにできるため、暗電流が混入したり、より低エネルギ側にシフトする散乱線が自動的に混入したりするという事態も殆ど確実に排除でき、したがって、コントラストの極めて高い画質を得ることができる。
【００７６】
以下、上述したＸ線検出器１２を採用した各種の装置を別の実施形態として説明する。
【００７７】
（第２の実施形態）
次に、本発明の放射線撮像システムに係る第２の実施形態として、本発明に係るＸ線検出器を採用した別のＸ線診断装置を、図１２、１３を参照して説明する。
【００７８】
前述したＸ線検出器１２によれば、その収集画素毎に、入射したＸ線の粒子数のカウント値に応じたデータが、しかもそのエネルギ領域別に検出されるので、このカウントデータをエネルギ領域に応じて選択的に画像化することができる。そこで、この実施形態では、このエネルギ領域に応じた選択的な画像再構成を行う機能を有し、薄い造影剤を静脈から投与した場合でも高感度に血管造影を行うことができるＸ線診断装置を説明する。
【００７９】
周知のように、ヨード糸造影剤を用いた冠状動脈描出は心血管の閉塞を検出するには有効であるとして実施されている。しかし、現状の積分モード型のＸ線検出の場合、動脈に造影剤を投入して造影剤濃度の濃い状態で、画像化しないと十分なコントラストが得られないという欠点がある。また、患者の動脈からの注入により、非常に侵襲度の高い検査方法であると言われている。
【００８０】
この状況を改善するために、単色Ｘ線発生装置を利用し、ヨード系造影剤のＫ吸収端エネルギ（３３．１７ｋｅＶ：図１２参照）近辺に単色Ｘ線を絞り、これを患者に照射する撮影法が盛んに研究されている。つまり、この単色Ｘ線を用いて、そのＫ吸収端エネルギよりも大きい及び小さい値のエネルギで画像化し、この画像間でサブトラクションをすることにより、静脈投与で造影剤が薄まった状態でも、鮮明に冠状動脈を描出しようとするものである。しかしながら、この撮影を行うには、単色Ｘ線発生装置のような大掛かりな設備が必要であり、この撮影法を広く普及させるには限界がある。
【００８１】
そこで、本実施形態に係るＸ線診断装置１Ａでは、図１２に示すように、少なくとも３３．１７ＫｅＶのエネルギ位置及びこれを挟んだエネルギ準位の上下２つの位置に３つのエネルギ弁別用の基準値ＴＨ１〜ＴＨ３を設定し、これにより３３．１７ｋｅＶの前後にエネルギ領域１、２を設定する。
【００８２】
このエネルギ領域１，２のカウントデータを選択的に収集し、その２種類のカウントデータを互いにサブトラクションする。これにより、単色Ｘ線で撮影したときとほぼ等価な生データを得ることができるので、この生データを再構成すればよい。なお、２種類のカウントデータを先に各別に再構成した後で、相互にサブトラクションするようにしてもよい。
【００８３】
これを実現するためのＸ線診断装置の構成例を図１３に示す。同図は、前述した図４の構成のうち、異なる部分の概要のみを示す。Ｘ線検出器１２の各検出器ブロック１２Ｊには、エネルギ領域１，２に入るＸ線粒子のカウントデータＫ_ｉ１，Ｋ_ｉ２（ｉ＝１〜Ｎ）をそれぞれ出力するカウンタ４６_１，４６_２を設ける。このカウントデータＫ_ｉ１，Ｋ_ｉ２は収集画素毎に検出されてコンソール５に送られる。コンソール５に入力したカウントデータＫ_ｉ１，Ｋ_ｉ２は、そのインターフェース２１を通って画素演算回路３１に送られる。この画素演算回路３１は、カウントデータＫ_ｉ１，Ｋ_ｉ２の相互の差分を演算する差分回路３１Ａ、それらを相互に加算する加算回路３１Ｂ、及び、カウントデータＫ_ｉ１、Ｋ_ｉ２それぞれを選択的に通過させる選択回路３１Ｃ、３１Ｄを備える。また、この画素演算回路３１は、コントローラ２０から与えられる切換制御信号ＳＣに応答して、この回路３１Ａ〜３１Ｄのうちの、１つ又はそれ以上の回路を選択的に作動させることができるようになっている。
【００８４】
この画素演算回路３１で演算又は選択されたカウントデータ「Ｋ_ｉ１−Ｋ_ｉ２」、「Ｋ_ｉ１＋Ｋ_ｉ２」、「Ｋ_ｉ１」、及び／又は、「Ｋ_ｉ２」は図４に示す記憶ユニット２２に保管される。再構成ユニット２３は、記憶ユニット２２に保管されているカウントデータのうち、差分演算に係る「Ｋ_ｉ１−Ｋ_ｉ２」を読み出して画像再構成する。これにより、エネルギ領域１，２間のサブトラクション画像が得られるので、造影剤の静脈投与による高感度な血管造影を行なうことができる。
【００８５】
このように、本発明に係るＸ線検出器１２を用いることで、各収集画素の入射Ｘ線粒子のエネルギ領域別の高精度なカウント値を選択的に利用して、冠動脈のみの画像を鮮明に描出することができ、通常の強度のＸ線源を搭載したＸ線診断装置であっても、造影剤を静脈から注入して、冠動脈を画像化できる。患者への負担を通常のルーチン検査並みに減らすことができ、冠動脈撮影を高度医療の領域から開放できるという大きな画期的なメリットが得られる。
【００８６】
なお、本実施形態に係るＸ線診断装置１によれば、上述した差分カウントデータ「Ｋ_ｉ１−Ｋ_ｉ２」に加えて、加算カウントデータ「Ｋ_ｉ１＋Ｋ_ｉ２」、各エネルギ領域別のカウントデータ「Ｋ_ｉ１」及び／又は「Ｋ_ｉ２」をも選択的に収集できる。このため、加算カウントデータ「Ｋ_ｉ１＋Ｋ_ｉ２」によるデータを再構成することで、一般のＸ線診断装置に相当するＸ線画像（すなわち、全スペクトラムを包括的に含むエネルギ量のデータから再構成したＸ線画像）を得るこができる。さらに、エネルギ領域別に選択したカウントデータ「Ｋ_ｉ１」及び／又は「Ｋ_ｉ２」を用いた画像再構成を行うことができる。このように、Ｘ線粒子が有する特定エネルギ成分の画像を選択的に提供できるので、画像情報を先鋭化させて診断に寄与可能になる。
【００８７】
（第３の実施形態）
次に、本発明の放射線撮像システムに係る第３の実施形態として、本発明に係るＸ線検出器を採用したＸ線ＣＴスキャナを、図１４、１５を参照して説明する。
【００８８】
このＸ線ＣＴスキャナは、ＩＶＲ−ＣＴと呼ばれている複合装置を、１台のＸ線ＣＴスキャナで担おうとするものである。
【００８９】
近年、ＩＶＲ（Ｉｎｔｅｒｖｅｎｔｉｏｎａｌ Ｒａｄｉｏｌｏｇｙ）技術の急速な進歩により、血管造影アプリケーションも多様化している。このため、血管造影装置と高性能Ｘ線ＣＴスキャナとを組み合わせることにより、従来の患者移送により施行していたＣＴアンギオ検査を一つの撮影室で施行できる、いわゆるＩＶＲ−ＣＴが盛んに使用されるようになってきている。しかしながら、この従来のＩＶＲ−ＣＴには、血管造影装置とＸ線ＣＴスキャナという２種類の装置が必要であるため、全体の装置が非常に大掛かりになり、装置の操作が複雑である、広い設置スペースが必要である、設置コストや運用コストが高いなどの問題があった。
【００９０】
そこで、本実施形態に係るＸ線ＣＴスキャナ１は、図１４に示すように、図１３の構成を採用するとともに、コンソール５の画素データ演算回路３１の出力側に、画素束ね回路３２と切換回路３３とを備える。画素束ね回路３２はＮ個のデータ（ここでは、収集画素毎のカウントデータ「Ｋ_ｉ１−Ｋ_ｉ２」、「Ｋ_ｉ１＋Ｋ_ｉ２」、「Ｋ_ｉ１」、及び／又は、「Ｋ_ｉ２」）を所定数ずつ相互に加算してｊ個のデータを作成する。
【００９１】
この加算により、図１５に示すように、Ｘ線検出器１２における全収集画素が所定数（領域）の画素群毎に束ねられて１つの画素になったことと同等になる。これを詳述すると、本実施形態に係るＸ線検出器１２の収集画素は、前述したように、従来のＣＴスキャナ用Ｘ線検出器のそれに比べて非常に精細に形成されており、例えば２００μｍ×２００μｍのサイズになっている。このため、通常のＸ線平面画像を得るときには、かかる精細な画素をそのまま用いて高精細な画像を得るようにする。一方で、ＣＴ画像を得るときには、この画素サイズは再構成に時間が掛かりすぎる。つまり、画素が精細過ぎるので、幾つかの隣接する画素同士を束ねて１つのデータを出力できるようにする（図のクロスハッチングの部分を参照）。図１５の例では、５×５の収集画素を信号加算により束ねて１つの画素とみなす。これにより、１ｍｍ×１ｍｍの大きさの画素を見かけ上、形成できる。
【００９２】
図１４に戻って、この画素束ね回路３２には、これを迂回する出力ラインＯＬが設けられており、この出力ラインＯＬと画素束ね回路３２の出力ラインとを、切換回路３２により、そのうちの何れか一方を選択して記憶ユニット２２にデータ保管できるようになっている。なお、画素データ演算回路３１及び切換回路３２は、コントローラ２０から与えられる切換制御信号ＳＣ１，ＳＣ２により、それぞれのモード（ＣＴ撮影モード、通常Ｘ撮影モード）に応じて回路切換又はスイッチ切換がなされる。
【００９３】
すなわち、血管描出をＣＴ撮影モードで行うときには、画素データ演算回路３１が差分回路３１Ａに切り換えられ、切換回路３２が画素束ね回路３２の出力ライン側に切り換えられる。一方、通常Ｘ線撮影モードのときには、画素データ演算回路３１が差分回路３１Ａに切り換えられ、切換回路３２が迂回出力ラインＯＬ側に切り換えられる。画素データ演算回路３１の加算回路３１Ｂ又は選択回路３１Ｃ，３１Ｄへの切換は、オペレータが所望するときにコントローラ２０を介して指令される。
【００９４】
このため、本実施形態に係るＩＶＲ−ＣＴ用のＸ線ＣＴスキャナ１によれば、被検体にヨード糸造影剤を静脈に注入し、通常のＸ線撮影を行って平面画像による血管描出を行うことができる。この血管描出は通常Ｘ線撮影モードで実行されるため、収集画素サイズが例えば２００μｍ×２００μｍであり、非常に高精細な画像になる。この段階で、画素データ演算回路３１の回路を差分回路３１Ａに切り換えることで、エネルギ領域を考慮したサブトラクション画像も得ることができる。
【００９５】
これらの適宜に且つ多様な観点から収集した高精細な画像を観察して血管の閉塞状況を把握の後、ターゲットとする血管に対して寝台４を位置決めをした後、すぐにその場でＣＴ撮影モードに切り換えてＣＴ撮影を行うことができる。このＣＴ撮影のデータ収集により、前述したエネルギ領域１、２に分離した再構成画像をサブトラクションしたＣＴ画像を得ることができる。
【００９６】
このように本実施形態に係るＸ線ＣＴスキャナ１によれば、一台の装置で、従来の血管造影装置と高性能Ｘ線ＣＴスキャナの機能を十分に賄うことができる。このため、従来よりも格段に高画質の平面画像を位置決めなどに用いることができるので、位置決め精度が高くなり、その場でＣＴ撮影に移行できるので、患者及び術者双方の負担も著しく軽減される。加えて、設置スペースが少なくてすみ、また設置や運用のためのコストも少なくなるので、病院の設備投資の面でも有利になる。さらに、このＸ線ＣＴスキャナによれば、必要に応じて、通常のＸ線平面画像の収集機能を省くことができ、その代わりに、サブトラクションＣＴ画像により血管のみを精密に３次元表示するようにしてもよい。
【００９７】
（第４の実施形態）
次に、本発明の放射線撮像システムに係る第４の実施形態として、本発明に係るＸ線検出器を採用した別のＸ線ＣＴスキャナを、図１６を参照して説明する。
【００９８】
このＸ線ＣＴスキャナは、肺がんの早期発見を目的とする集団検診に適したＸ線ＣＴスキャナに関する。
【００９９】
図１６に、このＸ線ＣＴスキャ１ナの構成例を示す。Ｘ線検出器１２は、その各収集画素に対応して３個のカウンタ４７_１〜４７_３を備える。このカウンタ４７_１〜４７_３は、それぞれ、Ｘ線エネルギのスペクトラム上に設定した所定の４個の異なるエネルギ閾値ＴＨ１〜ＴＨ４（図示せず）で分割されるエネルギ領域１〜３のＸ線粒子入射数をカウントし、カウントデータＫ_ｉ１、Ｋ_ｉ２，Ｋ_ｉ３（ｉ＝１〜３）して出力する。これに対して、コンソール５には、インターフェース４８、束ね回路網４９、及びＮ個の収集画素チャンネルに対応した３個の画素束ね回路５０_１〜５０_３を備える。
【０１００】
各収集画素チャンネルからのカウントデータＫ_ｉ１、Ｋ_ｉ２，Ｋ_ｉ３はインターフェースを介して束ね回路網４９に入力する。束ね回路網４９は、各画素収集チャンネルの３つのカウント出力経路それぞれを３個の画素束ね回路５０_１〜５０_３にそれぞれ振り分けて入力させるようになっている。
【０１０１】
画素束ね回路５０_１〜５０_３は、コントローラ２０から与えられる切換制御信号ＳＣ３に応答してＮ個の入力信号を、所望数の相互に隣接する画素群毎に加算する。これにより、収集画素を所望画素数毎に束ねることができる。このとき、画素束ね回路５０_１〜５０_３は、それぞれ、切換制御信号Ｓ３の指令内容に応じて、束ねる画素数（画素群）を複数種にわたって変更できるようになっている。例えば、基本の収集画素を２００μｍとした場合、束ねない状態で出力する２００μｍのほか、束ねることによる４００μｍ、８００μｍ、および１６００μｍのうちの４種類のうちの所望の値に変更できる。この束ね法はオペレータが選択して指令することができる。
【０１０２】
この画素束ね回路５０_１〜５０_３は、エネルギ領域１〜３に対する画素束ねを担っている。このため、各画素束ね回路５０_１（〜５０_３）は、切換制御信号ＳＣ３を介して、例えば４種類の収集画素サイズ２００μｍ、４００μｍ、８００μｍ、および１６００μｍの中から所望サイズのカウントデータを出力できることになる。
【０１０３】
そこで、所望のエネルギ領域別に且つ所望の収集画素サイズ別にカウントデータ蓄積して画像再構成することで、肺がんの特徴描出に適したエネルギ範囲と分解能でＣＴ画像を得ることができる。
【０１０４】
このため、現状の再構成画像の分解能０．５ｍｍ程度が一般的であるが、再構成時間の短縮により、分解能０．１ｍｍ以下の描出も可能となる。また、関心エネルギーウインドウに入るＸ線を使って画像再構成を行うことができるので、肺がんに絞った描出を行うことができ、肺がんに対する検出能が飛躍的に高まる。したがって、真に初期段階の肺ガンの検診が短時間に可能となる。このようなＸ線ＣＴスキャナが普及することで、肺がんの早期発見を目指して集団検診を行うことができる。
【０１０５】
（第５の実施形態）
次に、本発明の放射線撮像システムに係る第５の実施形態として、本発明に係るＸ線検出器を採用したマンモグラフィ装置（乳房用Ｘ線撮影装置）を、図１７，１８を参照して説明する。
【０１０６】
Ｘ線を用いて乳ガンを検出する場合、ガンの初期段階で発見したとする観点から、石灰化した微小組織の検出が重要である。石灰化した組織は２０ｋｅＶ前後のエネルギのＸ線に対し、他の正常組織との際立った吸収係数の差を呈するが、ガン化（石灰化）してしまった組織は正常組織との吸収差が少ない。この微小石灰化は、比較的エネルギの低いＸ線に検出感度を有するので、低エネルギのみに絞った平面画像を高いコントラストで撮影する必要がある。
【０１０７】
本実施形態では、図１７に示すように、フィルターを用いた線質を低エネルギ側にシフトさせ、患者被曝線量を、２０ＫｅＶを中心とするエネルギ領域１と、その低エネルギ側のエネルギ領域２に分けた状態でカウントデータをエネルギ領域別に収集するようにする。
【０１０８】
具体的なマンモグラフィ装置の構成は、血管造影のシステムと同じである。エネルギ閾値は例えば、
【外３】

に設定する。
【０１０９】
前述した血管造影撮影のときと同様に、エネルギ領域１及び２それぞれから収集したカウントデータを相互にサブトラクションすることで、早期の石灰化した組織のみの高コントラストのＸ線画像を得ることができる。これにより、乳がんも早期に発見可能となる。
【０１１０】
このＸ線画像の場合、１収集画素あたりのＸ線粒子のカウント数をＸ線ＣＴスキャナの場合よりも多くする必要がある。これについては、各収集画素に内蔵するカウンタがオーバーフローする時間を線量強度に応じて算定し、それよりも短い時間でカウンタを定期的にリセットしながら、収集カウントをＸ線爆射時間の間、積分すれば、高カウント値のマンモグラフィ画像を得ることができる。
【０１１１】
従来のマンモグラフィの場合、石灰化した組織を検出するためには、収集画素を細かくし、検出能を上げてやる必要があったが、本実施形態によれば、検出感度が高いので、２００μｍ程度の収集画素サイズでも十分、微小石灰化を描出できる。
【０１１２】
また、本実施形態では、Ｘ線検出器１２をＣｄＴｅ半導体で構成しているので、その検出面（Ｘ線入射面）を検出器端面ぎりぎり（例えば５ｍｍ以内）まで形成することができる。このため、図１８に模式的に示すように、圧迫板６０を用いることにより、極めて広い乳房領域を一度に撮影できる。また、本実施形態の事例の場合、１０ｋｅＶ以下の散乱成分は検出しないため、散乱成分が極めて少なく、コントラストの良いＸ線画像になる。
【０１１３】
（第６の実施形態）
次に、本発明の放射線撮像システムに係る第６の実施形態として、本発明に係るＸ線検出器を採用した手荷物検査装置を、図１９を参照して説明する。
【０１１４】
現在、空港などでは、手荷物検査のために、産業用Ｘ線ＣＴスキャナが導入されているが、低感度であるため、Ｘ線強度を上げてやる必要がある。その結果、通常の未現像の写真フィルムが感光してしまうなどの不都合がある。加えて、現状のＸ線ＣＴスキャナは空港の大勢の人の手荷物全部の検査を賄うだけの高速動作はできないので、手荷物検査に時間がかかり過ぎるという問題がある。このため、手荷物の全数を一旦、通常のエックス線で検査し、疑わしい手荷物だけをＸ線ＣＴスキャナで精密に検査しているという状況にあり、手間と時間がかかっている。
【０１１５】
そこで、本実施形態に係る手荷物検査装置７０は、図１９に示すように、架台７１にＸ線管１１と本発明に係るＸ線検出器１２との対を回転可能に設置し、架台７０を貫通する開口部ＯＰをベルトコンベア７２が通過できるようにする。ベルトコンベア７２により一定速度で運搬されてきた手荷物は架台７１の開口部ＯＰを通過するので、Ｘ線管１１とＸ線検出器７２とが開口部ＯＰの周りを回転することで、Ｘ線ＣＴのヘリカルスキャンが可能になる。このスキャンを介して、Ｘ線検出器１２により前述の如く検出されたＸ線粒子数に応じたカウントデータ（エネルギ領域別のカウントデータを含む）が検出される。
【０１１６】
このカウントデータは再構成装置７２に送られる。再構成装置７２は、荷物がガントリ７１を通過し終わる時点までには、そのＣＴ画像データを生成する。このＣＴ画像はモニタ７３に表示され、監視員に提示される。
【０１１７】
また、ＣＴ画像のデータは危険物自動認識装置７４に送られる。これにより、危険物自動認識装置７４は、予め設定した危険物（例えば金属、麻薬）にエネルギ領域を絞ったカウントデータに基づくＣＴ画像のデータを受ける。そこで、危険物自動認識装置７４は、このＣＴ画像に危険物自動認識処理（特徴抽出プログラム）を施して、該当する危険物の有無を自動的に判断し、その結果を異常情報検出提示装置７５に送る。
【０１１８】
本実施形態によれば、かかるＸ線ＣＴスキャナの機能によりリアルタイムに且つ連続的に複数の手荷物を検査することができる。つまり、エネルギ領域別にカウントデータを検出可能なＸ線検出器１２を用いているので、高感度さが要求される危険物検出に適する高エネルギに絞った画像再構成から麻薬などの低エネルギに絞った画像再構成まで幅広く対応可能である。このため、全ての手荷物をＣＴスキャンに連続的にかけることができる。さらに、効率よく且つ無人化して危険物を検出することができる。さらに、現場におけるリアルタイムなＣＴ画像の取得に加え、危険物描出が容易な関心エネルギのみのＣＴ画像を事後に編集することが可能である。これにより、より確度の高い検査を短時間に行なうことができ、目視検査を併用した半自動化の手荷物検査又は完全自動化した手荷物検査が可能にもなる。勿論、検出感度が優れるため、Ｘ線強度をその分下げることができ、未現像の写真フィルムに対する感光の問題も解消される。
【０１１９】
（第７の実施形態）
次に、本発明の放射線撮像システムに係る第７の実施形態として、本発明に係るＸ線検出器を採用したパイプライン検査システム７９を、図２０を参照して説明する。
【０１２０】
通常アラスカ、サハリンなど寒冷地に設置された石油パイプラインの亀裂などの検査は困難を極める。この石油パイプラインに沿ってＸ線非破壊検査装置を自動走行させてＸ線非破壊検査が行なえれば一番良いが、しかし、このパイプラインの敷設された環境下でＸ線を発生させることは現実的でない。
【０１２１】
そこで、図２０に模式的に示すように、本実施形態に係るパイプライン検査システム７９は、密封型のガンマ線源を用いる。これは、本発明に係る放射線検出器がエネルギ領域での画像描出が可能なことに拠る。
【０１２２】
具体的には、前述してきたＸ線検出器１２をガンマ線検出に用いるもので、このライン状のガンマ線検出器１２とライン状の密封型ガンマ線源８０とをパイプライン８１を介して対向配置させる。ガンマ線源としては、^１２５Ｃｏ，^７５Ｓｅ，^１３７Ｃｓなどが用いられる。ガンマ線源８０のガンマ線出射側には、ガンマ線を平行に導くコリメータ８２が一体に設けられている。このガンマ線検出器１２及びガンマ線源８０（コリメータ８２）の組を、パイプライン８１の周りに適宜に回転させながら、その長手方向に沿って移動させる走行機構（図示せず）が設けられている。
【０１２３】
これにより、ガンマ線源８０からコリメータ８２を通して平行に照射されたガンマ線がパイプを透過し、トランスミッション像としてガンマ線検出器１２に検出される。この結果、ガンマ線検出器１２からガンマ線の粒子数に応じたカウントデータが得られる。このとき、前述のように、ガンマ線のスペクトラムに対してエネルギ領域を弁別する閾値（例えば２つの閾値ＴＨ１，ＴＨ２とし、そのエネルギ範囲は密封型ガンマ線エネルギの±１０％に設定）が設定される。このため、カウントデータもその閾値で設定されるエネルギ領域に対応したデータになる。
【０１２４】
ガンマ線検出器１２で検出されたカウントデータはメモリ８３を介して電波送信機８４に送られ、位置信号器８５からの位置信号とともに別の場所のアンテナ８６に電波で送信される。アンテナ８６で受けた電波は受信機８７で処理され、この処理された受信データがパイプライン異常検出装置８８に送られる。この異常検出装置８８により、受信データに基づいてＣＴ画像を再構成するなど、適宜な処理を行って、パイプラインの損傷や内部歪などの異常の有無及び内容情報を提供する。
【０１２５】
この結果、パイプに入った亀裂などの画像データを遠隔地で得ることができ、厳しい自然環境の中に設置されたパイプランであっても、それを無人で自動的に検査することができる。
【０１２６】
なお、このような検査は、外部からのガンマ線に対する遮断構造がとれるならば、工場のパイプラインの自動検査にも用いることができる。
【０１２７】
また、本発明に係るＸ線又はガンマ線による検出器を採用したＣＴスキャナを用いて内部構造を検査する例は、パイプラインに限られず、
【外４】

に対しても適用可能である。
【０１２８】
例えば、現状のリチュームイオン検査では、電極部の不良検査をＸ線の透視撮影で行なっているが、製品を流すためのピッチタイムが短いため、現状のＸ線撮影時間では長すぎる。そこで、同じＸ線非破壊検査装置を３台併設し、検査工程のみ３つの工程に平行分離して行なっていたが、本願発明に係る放射線検出器を用いることにより、１台の非破壊検査装置で検査をこなせるようになる。加えて、より検査に適したエネルギのみに基づいてコントラストの高い画像を得るから、不良品を高精度に抽出することができる。従って、検査にかけるコストを大幅に削減することが可能となる。
【０１２９】
また、金属製品の内部構造をＣＴで分析する場合、Ｘ線のエネルギが高くないとＸ線は金属を通過できないので、従来では、非常にコントラストの低いＸ線画像しか得られず、ＣＴをかかる分析に使用することはあまり行われていなかった。しかし、本発明の放射線検出器によれば、関心エネルギを峻別できるので、例えば高エネルギ成分だけの再構成画像を得ることができ、Ｘ線ＣＴであっても、内部構造をより鮮明に描出できる。
【０１３０】
さらに、切り出した木材を、前述した手荷物検査システムと同様の構成のシステムで検査することで、木材の内部の空洞の有無を知ることができる。これにより、木材内部の構造を確実に知ることができ、木材加工の能率アップや材木製品の品質安定に寄与する。
【０１３１】
このように、本発明に係るＸ線検出器は、あらゆるＸ線非破壊検査の分野において、著しく高感度であること、及び、エネルギ領域の峻別が可能であることによって、その優れた有用性を発揮することができる。
【０１３２】
（第８の実施形態）
次に、本発明の放射線撮像システムに係る第８の実施形態として、本発明に係るＸ線検出器を採用したＸ線天文学用の２次元画像システムを、図２１を参照して説明する。
【０１３３】
最近のＸ線天文学の進歩の中で、エネルギに分けたＸ線画像の撮影が宇宙の構造解析（成分分析）には不可欠である。
【０１３４】
そこで、図２１に示すように、宇宙空間の静止軌道上に打ち上げられた人工衛星により、２次元画像システム９０のセンサ装置９１を宇宙空間に滞在させる。このセンサ装置９１は、本発明に係るＸ線検出器１２を用いる。このＸ線検出器１２は、０．５ｍｍ厚の擬似ＳｃｈｏｔｔｋｙＣｄＴｅピクセル検出器で成り、これを−２０℃まで冷却して使用するようになっている。これにより、このＸ線検出器１２は、Ｘ線感度が高いことに加え、現状では検出感度は低いがエネルギ分解能としては最も優れているＧｅ検出器並みのエネルギ分解能を有する。したがって、Ｘ線天文学のような厳密性を要する検出器としても前例のない特性を示すことができる。
【０１３５】
このＸ線検出器１２のＸ線入射面にはコリメータ９２が設置される。また、この検出器１２に設ける処理回路では、少なくとも特性Ｘ線を捕捉することができるように、エネルギ領域をＭ段（Ｍは、２以上の正の整数）に設定している。
【０１３６】
そこで、Ｘ線検出器１２からのカウントデータは特性Ｘ線検出回路９３に送られ、この検出回路９３により、そのカウントデータがエネルギ領域別に検出されて、Ｍ個のエネルギ別用のメモリ９４_１〜９４_Ｍに分けて記憶される。この記憶データは切換器９５により、シリアル信号に変換され、送信機９７を介して地球アンテナ９８に向けて送信される。なお、ＣＰＵ９９は、処理回路の各回路の動作を制御する。
【０１３７】
地球アンテナ９８に送られたデータは、受信機１００及びメモリ１０１を経て処理・表示装置１０２に送られ、特性Ｘ線を観測するための所定の処理に付される。
【０１３８】
このようなシステム構成にすることにより、常時観測したい方向にＸ線検出器１２を向け、その天台付近から発する特性Ｘ線を関心エネルギ毎に同時に画像化して捕えることができる。このため、時々刻々と変化する天台の様子を時系列に捕えるような観測を地球上で行うことができる。つまり、このＸ線検出器１２は、一度に、複数のエネルギのＸ線画像を同時に捉えることができるので、瞬時のバースト現象を捉える必要があるＸ線天文学用の２次元センサとしても好適に機能する。
【０１３９】
（作用効果のまとめ）
以上のように、本発明に係るＸ線検出器は、ＣｄＴｅ半導体を用いて、２００μｍ前後の画素をＭｏｎｏｌｉｔｈｉｃ構造で構成し、これに検出器の背面にはみ出さずにチャージアンプを含む処理回路をＡＳＩＣ化して設けて成る。この検出器により、Ｘ線を粒子とみなし、各Ｘ線粒子が入射毎に発生するイオナイズされた電気信号をチャージアンプで検出させる。この検出信号をパルス整形回路に送り、そのピーク値がエネルギに比例する信号を発生させ、これを複数の比較器に入力させる。これらの比較器で、複数のエネルギ閾値（ＤＣ電圧）との比較を行わせ、入射粒子のエネルギが所定エネルギを超えたかどうかについてエネルギ領域毎に判定させる。各エネルギ領域に入ったエネルギの粒子数をそれぞれカウントさせる。したがって、Ａ／Ｄ変換器を設けずに、各エネルギ領域での出力をデジタル値で得ることができる。
【０１４０】
このようなパルス数を計測する方法によれば、エネルギ閾値は暗電流より高いレベルに設定されるため、積分型に比べてＳ／Ｎの極めて良い画像が得られる。また各画素の大きさは２００μｍと小さいため、各画素に入射するＸ線粒子数の数え落とし特性の現れる１（Ｋｃｏｕｎｔ／ｓｅｃ）よりはるかに下回るため、数え落としもなく、計測することができる。
【０１４１】
通常、ＣＴ画像を得るために必要な画素サイズは１ｍｍ前後であるから、収集画素をその検出単位にするには、例えば２００μｍの画素サイズの検出器出力を２５個束ねれば（加算すれば）、等価的に数え落としなく出力を得ることができる。
【０１４２】
また、もともとの画素が２００μｍの場合、その出力をそのまま束ねずに出力することにより、Ｘ線の平面画像を得ることもできる。これはＩＶＲ−ＣＴで代表される医療用Ｘ線ＣＴスキャナと通常のＸ線診断装置を１台のＣＴスキャナでこなせることを意味しており、大幅な設置スペースや装置コストなどの削減を実現することができる。
【０１４３】
同じく手荷物検査用の非破壊検査機器では、まずは通常のＸ線透過画像で疑わしき手荷物を見つけ、これをＣＴスキャナで精密に判定するような手順を踏んでいるが、本発明によれば、１台のＣＴスキャナで最初からＸ線透視画像が得られるので、コスト、設置スペース、スループットなど面で極めて有利である。
【０１４４】
また、そのあるエネルギを超えたＸ線粒子のみを計測しているため、人体などで散乱して検出器に入るＸ線がＸ線管より発せられるＸ線よりエネルギが小さくなっており、散乱線の混入割合の低い画像を得ることができる。
【０１４５】
複数のエネルギ閾値を設けてやることにより、スキャン後に、再構成データを編集することにより、例えば高エネルギ画像、低エネルギ画像、あるいは、ある範囲のエネルギのみの画像を任意に選択的に再構成することもできる。
【０１４６】
Ｘ線検出器はＣｄＴｅ半導体を用いているので、現状のＧＯＳやＣｓＩのフォトダイオードを含めた検出器系と比べて、その検出感度に優れている。このため、Ｘ線管のＸ線強度を落としても、それまでと同程度の感度の画像が得られる、被検者の被曝も抑えることができる。
【０１４７】
空港の手荷物検査などで使用されるＸ線ＣＴスキャナでは、写真のフィルムを手荷物内に入れていると感光してしまうため、手荷物内には未現像のフィルムを入れることは禁止されているが、Ｘ線の量を減らせれば、感光のリスクも低減させることができる。
【０１４８】
ＣｄＴｅ半導体を用いたＸ線検出器は、放射線を直接電気信号に変換し、かつカウント値としてデジタル信号を出力するため、従来存在していたフォトダイオードとＡ／Ｄ変換器が不要になり、かつ、その後段の回路構成が簡素化される。また、このＸ線検出器はＭｏｎｏｌｉｔｈｉｃ構造を採用しているため、電極の形成がすなわち画素の形成となり、シンチレータによる画素の形成に比べて工数が少なくて済む。このため、トータルとして製造コストを大幅に低減できる。
【０１４９】
なお、パルスカウントをベースにシステムを構築する場合、各収集画素に対する波形整形回路からのパルス（エネルギ値に比例）のゲイン、オフセットのばらつきとエネルギ閾値（ＤＣ電圧）の相対的な関係が極めて重要で、これが精度良く設定されていななければ、ＣＴ画像のアーチファクトの大きな原因となり、実用化は難しくなる。そこで本発明では、実際のＸ線源のエネルギ領域に近い密封型のガンマ線源である^２４１Ａｍ（５９．５ＫｅＶ）と^１２２Ｃｏ（１２２ＫｅＶ）をリファレンスとして用い、検出器の出力調整の段階で各画素の回路系のゲイン及びオフセットを計測し、それに見合った精度の良いエネルギ閾値の設定をプログラム化して対処する。このプログラムを装置据え付け時などの適宜なタイミングでキャリブレーションとして実行することで、かかるばらつきを確実に抑制して、アーチファクトを抑制した高品質のＣＴ画像を確保できる。
【０１５０】
なお、本発明に放射線検出器及び放射線撮像システムは、上述した実施形態に記載の構成及びその変形例の構成に限定されるものではなく、当業者においては、特許請求の範囲に記載の要旨を逸脱しない範囲で適宜に変更、変形可能なものである。例えば、Ｘ線検出器は、前述したように集電電極により収集画素の位置及びサイズが決まるモノシリック構造において、前述した各実施形態のように、収集画素が２次元配列（図２参照）になるものみに限定されず、例えば図２２（ａ），（ｂ）に示すように、収集画素を１次元配列の検出器、又は、２次元配列の検出器を積層してなる３次元配列の検出器であってもよい。さらに、同図（ｃ）に示すように、所望のサイズを有する１個の画素（入射ウィンドウ）であっても、放射線粒子がこの入射ウィンドウに複数個連続して入射したときの各入射に応答したパルス信号間の重畳現象の発生を実質的に無視可能であれば、使用可能である。
【０１５１】
【発明の効果】
以上説明したように、本発明に係る射線検出方法、放射線検出器、及び、放射線撮像システムによれば、医療用から産業用まで広範囲の分野にわたって応用可能であって、散乱線の影響を大幅に減らして高コントラストを確保し、且つ、暗電流の影響を排して極めて高いＳ／Ｎを確保することで非常に高い検出感度を有するとともに（放射線の被曝量も減）、センサ部により検出された信号を処理する回路の回路規模を積分モードで動作するＸ線検出器に比べて大幅に減らして小形化図ることができる。
【０１５２】
また、放射線のエネルギに応じた検出を行うことができるなど、１台の装置で多様な種類の放射線の検出に対応可能になる。
【図面の簡単な説明】
【図１】本発明の放射線撮像システムとしてのＸ線ＣＴスキャナの一実施形態の概略を示す説明図。
【図２】Ｘ線ＣＴスキャナに搭載したＸ線検出器の概略を示す斜視図。
【図３】数え落とし特性を説明する入力カウント数、対、出力カウント数のグラフ。
【図４】Ｘ線検出器を中心としたＸ線ＣＴスキャナの電気的な概略構成を示すブロック図。
【図５】収集画素単位で形成するＸ線検出の処理回路を説明する図。
【図６】入射Ｘ線粒子のエネルギ弁別から、そのカウント動作までを説明するタイミングチャート。
【図７】通常のＸ線管球から曝射されるＸ線のエネルギスペクトラムと閾値との関係を示すグラフ。
【図８】収集画素毎に実施するキャリブレーションを説明するブロック図。
【図９】キャリブレーションの処理の手順の一例を説明する概略フローチャート。
【図１０】キャリブレーションのリファレンス設定を説明する図。
【図１１】Ｘ線管から曝射されたＸ線の典型的な経路を説明する図。
【図１２】血管造影のための閾値設定を説明する図。
【図１３】本発明に係る放射線撮像システムの第２の実施形態としての、血管造影を行うＸ線診断装置の部分ブロック図。
【図１４】本発明に係る放射線撮像システムの第３の実施形態としての、ＩＶＲ−ＣＴを行うＸ線ＣＴスキャナの部分ブロック図。
【図１５】複数の収集画素の束ねを説明する図。
【図１６】本発明に係る放射線撮像システムの第４の実施形態に係る、肺がんの集団検診が可能なＸ線ＣＴスキャナの部分ブロック図。
【図１７】本発明に係る放射線撮像システムの第５の実施形態としてのＸ線撮影装置で実行されるマンモグラフィのためのＸ線スペクトラムと閾値設定との関係を説明する図。
【図１８】圧迫板を用いたマンモグラフィの様子を説明する図。
【図１９】本発明に係る放射線撮像システムの第６の実施形態としての手荷物検査システムを示す概略ブロック図。
【図２０】本発明に係る放射線撮像システムの第７の実施形態としてのパイプライン検査システムを示す概略ブロック図。
【図２１】本発明に係る放射線撮像システムの第８の実施形態としてのＸ線天文学用の２次元画像システムを示す概略ブロック図。
【図２２】Ｘ線検出器の収集画素配列の別の例を説明する図。
【符号の説明】
１ Ｘ線ＣＴスキャナ（放射線撮像システム）
１Ａ Ｘ線診断装置（放射線撮像システム）
１１ Ｘ線管（放射線源）
５ コンソール
１２ Ｘ線検出器（放射線検出器）
１２Ｊ 検出器ブロック
２０ コントローラ
２２ 記憶ユニット
２３ 再構成ユニット
３１ 画素データ演算回路
３２ 画素束ね回路
３３ 切換回路
４１ チャージアンプ
４２ 波形整形回路
４３ 比較器
４４ エネルギ領域振分回路
４５ カウンタ
４６ カウンタ
４７ カウンタ
４９ 束ね回路網
５０ 画素束ね回路
５１ 調整器
５２ メモリ
７０ 手荷物検査システム
７９ パイプライン検査システム
９０ Ｘ線天文学用の２次元画像システム
Ｓ 半導体セル
ＡＳ ＡＳＩＣ層
Ｅ１ 荷電電極
Ｅ２ 集電電極
Ｃ 処理回路[0001]
TECHNICAL FIELD OF THE INVENTION
The present invention relates to a radiation detection method for detecting radiation such as X-rays, a radiation detector, and a radiation imaging system equipped with the detector. In particular, the present invention regards incident radiation as particles and counts the number of particles. The present invention relates to a radiation detection method, a radiation detector, and a radiation imaging system equipped with this detector, which are provided with a quantum counting type circuit and exhibit epoch-making detection performance which has not been achieved in the past with respect to detection sensitivity, detection signal processing, and the like.
[0002]
[Prior art]
In recent years, as seen in medical image diagnosis, nondestructive inspection of structures, and the like, the importance of imaging the inside of a human body or an object using X-rays, which is a kind of radiation, has been increasing. In order to meet this demand, it is required to image a target object with a small X-ray dose, a high S / N, and a wide field of view, and research for this purpose is being actively pursued.
[0003]
This will be described with reference to an X-ray detector mounted on an X-ray CT scanner. Conventional X-ray detectors begin with the application of the ionization phenomenon of Xe gas, and recently use a two-dimensional solid-state detector such as GOS, CsI, or CWO (a detector using a combination of a scintillator and a photodiode). Is also being used. At the same time, in recent years, the field of view has been required to be expanded in the body axis direction, and the development of a large aperture has been progressing so that the X-ray detector can cope with a wide field of view capable of performing a volume scan of the heart and lung fields.
[0004]
As the field of view is expanded in the body axis direction in this way, the influence of scattered radiation from the human body increases, and at the same time, the sensitivity of the detector system is poor, and the X-ray exposure dose that does not contribute to reconstruction increases, In addition to increasing the exposure dose of the subject, the lack of sensitivity has been compensated by increasing the X-ray intensity by increasing the current flowing through the X-ray tube. Generally, if the current flowing through the X-ray tube is increased, the life of the X-ray tube is shortened, and it is easy to cause a failure.
[0005]
Further, since the sensor portion of the X-ray detector is formed of a scintillator, processing for pixel separation is necessary to avoid crosstalk of signals between pixels. This process requires enormous man-hours. Further, before converting the detection signal into a digital value, various types of circuit elements such as a scintillator, a photodiode, a preamplifier, and an A / D converter are required. For this reason, the present situation has been that the manufacturing cost of the entire X-ray detector must be increased. In particular, when the size of the X-ray detector is increased to provide a wide field of view, the price of the X-ray CT scanner naturally rebounds, which may be a hindrance to the spread of the X-ray CT scanner. .
[0006]
The X-ray CT scanner including the above-described scintillator, photodiode, preamplifier, and A / D converter for each pixel has a detection principle of an integration mode as disclosed in Patent Document 1 and Non-Patent Document 1, for example. A working device. That is, after converting the incident X-rays into light, this optical signal is converted into an electric signal, and the electric signal is integrated for a predetermined period of time to be output as a detector signal.
[0007]
However, in the case of an X-ray detector that operates in the integration mode, the influence of scattered radiation is likely to occur. However, there is a problem that artifacts are likely to appear. In addition, since reconstruction is performed under the condition including scattered radiation, there is a problem that the image contrast is inferior to an actual tomographic image. Furthermore, since a scintillator is used, it is difficult to increase the definition of pixels, and the current pixel size is at least about 1 mm. For this reason, it is not possible to cover both a CT image and an X-ray image represented by IVR-CT (International CT) with one detector. Therefore, conventionally, IVR-CT has to be implemented by combining two dedicated X-ray CT apparatuses and two X-ray apparatuses.
[0008]
On the other hand, as shown in Patent Literature 2 and Non-Patent Literature 2, a quantum counting type X-ray device that detects an incident X-ray in a pulse measurement mode using a cadmium telluride semiconductor (hereinafter, referred to as a CdTe semiconductor) is also available. Are known.
[0009]
[Patent Document 1]
JP-A-2000-131440
[0010]
[Patent Document 2]
JP 2000-69369 A
[0011]
[Non-patent document 1]
"Present and Future of Room Temperature Semiconductor Radiation Detectors," Atomic Energy Society of Japan Autumn Meeting Planning Session, Presentation Date: September 12, 1999, Presenter: Masaki Misawa, Mechanical Engineering Laboratory, Industrial Technology Institute, Ministry of International Trade and Industry "
[0012]
[Non-patent document 2]
"Report on the results of scientific research at Kobe University School of Medicine" "Study on Image Processing Method of Quantitative Counting X-Ray Imaging in Chest Imaging", Publication date: March 1999, Research representative: Eiichiro Itoi, Call number: Kaken 27. -MID: 000094541 "
[0013]
[Problems to be solved by the invention]
However, in the case of the above-described conventional quantum counting type X-ray apparatus, since the size of each pixel is large in view of the X-ray incidence characteristics, the amount of X-ray incident on one pixel is also large.
[0014]
In general, in the pulse measurement mode, it is desirable that the characteristic of the output count number with respect to the input count number be linear. However, in the case of such a quantum counting type X-ray apparatus, the superposition phenomenon of the detection pulse accompanying the detection of the incident radiation particles. For this reason, as the input count number increases, the output count number saturates or decreases, that is, the so-called count-down characteristic becomes severe, and the device is far from practical use. In particular, since the X-ray CT scanner is required to have sufficient linearity of the output value with respect to the detector input dose, the existence of such a count-down characteristic requires the use of the quantum counting type X-ray apparatus described above in the X-ray CT scanner. It meant that it was difficult.
[0015]
INDUSTRIAL APPLICABILITY The present invention can solve the above-mentioned problems of the conventional integral type and quantum counting type X-ray detectors at once, and can be applied to a wide range of fields from medical use to industrial use. The effect of the line is greatly reduced to ensure high contrast, and the effect of dark current is eliminated to ensure an extremely high S / N. Thus, a very high detection sensitivity can be obtained. A quantum counting type, which can greatly reduce the circuit scale of a circuit for processing a signal detected by the sensor unit as compared with an X-ray detector operating in an integration mode, can be reduced. It is a first object of the present invention to provide a radiation detector and a radiation imaging system equipped with the detector.
[0016]
Further, in addition to the first object described above, the present invention not only simply outputs an electric signal according to the incident radiation but also performs detection according to the type (energy) of the radiation as in the related art. A second object of the present invention is to provide a quantum counting type radiation detector capable of responding to detection of various types of radiation with one device and a radiation imaging system equipped with the detector. I do.
[0017]
[Means for Solving the Problems]
In the present invention, in order to achieve the above object, the radiation detector according to the present invention is a radiation detector configured to detect the number of particles of the radiation by assuming the radiation as particles and the radiation particles are An incident window having a size that can substantially ignore the occurrence of a superposition phenomenon between pulse signals in response to each incident when a plurality of incident light are continuously incident on the same position or in the vicinity thereof, and the incident window is incident on the incident window. A sensor unit that generates an electric signal in response to each of the radiation particles; and an energy detection unit that detects an electric signal generated by the sensor unit and detects an energy value of each of the radiation particles incident on the incident window. The main feature is to have.
[0018]
The radiation is, for example, radiation having a continuous spectrum of energy.
[0019]
Preferably, energy discriminating means for discriminating an energy value detected by the energy value detecting means into a desired energy area, and counting the number of radiation particles incident on the energy area discriminated by the energy discriminating means. Means for outputting a signal representing the count value as the detection signal.
[0020]
As an example, the incident window includes a plurality of sets of pixels each formed to have a size that can substantially ignore the occurrence of the superimposition phenomenon. For example, each of the plurality of pixels has a pixel size of 200 μm × 200 μm or less, or a pixel size capable of detecting the radiation particles with an accuracy of 0.1% or less. The plurality of pixels are arranged in a one-dimensional, two-dimensional, or three-dimensional array.
[0021]
Preferably, a calibration means is provided for calibrating a relative magnitude relationship between the energy value detected by the energy value detection means and the predetermined value based on a reference energy value. For example, the calibration unit is configured to perform the calibration using an energy value of a gamma ray source as the reference energy value. The gamma ray source is composed of at least two types of sources having different energy values, for example, having an energy value of 59.5 KeV.^{2 41}Am and energy value of 122 KeV⁵⁷It is characterized by containing Co.
[0022]
Further preferably, the sensor unit includes a semiconductor layer that directly converts the radiation into an electric signal. This semiconductor layer is preferably made of a compound semiconductor. As an example, the semiconductor layer is made of a cadmium telluride semiconductor (CdTe semiconductor), a cadmium zinc telluride semiconductor (CdZnTe semiconductor), or the like. Further, the semiconductor layer may be formed of a silicon semiconductor (Si semiconductor).
[0023]
Preferably, the sensor section covers the radiation incident surface of the semiconductor layer with a charging electrode for voltage application (corresponding to an incident window), and a current collector electrode on a surface opposite to the radiation incident surface of the semiconductor layer. Has a covering monolithic structure. The current collecting electrode has a structure in which the radiation is divided into a plurality of pixels on the back surface for detecting the radiation for each particle. Further, the sensor unit integrally includes a processing circuit including a charge amplifier for detecting an electric signal generated in the semiconductor layer in response to the incidence of the radiation particles, on a surface of the current collecting electrode of the semiconductor layer using an ASIC. It is desirable to arrange them in layers.
[0024]
On the other hand, according to the radiation imaging system according to the present invention, a radiation detection unit that detects radiation using the above-described radiation detector, and an image generation unit that generates an image using a signal of the radiation detected by the radiation detection unit Means.
[0025]
Further, according to the radiation detection method according to the present invention, a method of detecting the number of particles of the radiation by irradiating the radiation having the energy of the continuous spectrum as particles, the radiation particles at the same position or in the vicinity thereof A plurality of successively incident light, having an incident window of a size that can substantially neglect the occurrence of superposition between pulse signals in response to each incident, and responding to each of the radiation particles incident on the incident window; A sensor unit for generating an electric signal, detecting the electric signal generated by the sensor unit, detecting the energy value of each of the radiation particles incident on the incident window, and discriminating the detected energy value. Counting the number of incident radiation particles exhibiting the discriminated energy value and outputting the detection signal representing the count value. And it features.
[0026]
BEST MODE FOR CARRYING OUT THE INVENTION
Hereinafter, an embodiment of a radiation imaging system according to the present invention will be described with reference to the accompanying drawings. In this radiation imaging system, an X-ray detector as a radiation detector according to the present invention is also implemented. In addition, it goes without saying that only the X-ray detector can be implemented alone.
[0027]
(1st Embodiment)
An X-ray CT scanner as a first embodiment of a radiation imaging system according to the present invention will be described with reference to FIGS.
[0028]
FIG. 1 shows a schematic configuration of the X-ray CT scanner.
[0029]
As shown in FIG. 1, the X-ray CT scanner 1 includes a gantry 2, a high-voltage generator 3, a bed 4, and a console 5. Inside the gantry 2, an X-ray tube 11 and an X-ray detector (radiation detector) 12 as an X-ray source (radiation source) (not shown) are arranged to face each other via a cylindrical opening OP. A pair of the X-ray tube 11 and the X-ray detector 12 is rotatably arranged in the gantry 2 around the opening OP.
[0030]
The X-ray tube 11 is supplied with a high voltage from the high-voltage generator 3, whereby the X-rays are emitted toward the opening OP, for example, in a pulsed manner. The top plate 4P of the bed 4 is inserted into the opening OP so as to freely advance and retreat. Since the subject P is placed on the top plate 4P, the X-rays emitted from the X-ray tube 11 pass through the subject P and travel toward the opposite side of the opening OP.
[0031]
X-rays transmitted through the subject P are counted for each pixel as X-ray particles (ie, X-ray photons) by a two-dimensional X-ray detector 12 provided inside the gantry 2 and output as a digital amount detection signal. It is supposed to be. That is, the signal detected by the X-ray detector 12 corresponds to the count value of the X-ray particles. The X detector 12 forms a part of the main feature of the present invention, and will be described in detail again in different sections.
[0032]
The detection signal of the digital amount for each pixel output from the X-ray detector 12 is original data of image reconstruction generally called raw data. The raw data is sent to a fixed unit via a slip ring or a non-contact type communication connection unit (not shown), and is transmitted from the fixed unit to the console 5 via a communication line. The console 5 includes an interface 21, a storage unit 22, a reconstruction unit 17, a display 18, and an input device 19, in addition to a controller 20 that controls all the components of the X-ray CT scanner 1.
[0033]
The digital raw data transmitted from the X-ray detector 12 for each pixel is temporarily stored in the storage unit 15. The data stored in the storage unit 15 is read out at an appropriate timing and sent to the reconstruction unit 17. The reconstruction unit 17 applies a desired reconstruction algorithm to the data to reconstruct a CT image. This CT image is displayed on the display 18, for example. The console 5 is provided with a post-processing processor for post-processing, and performs appropriate post-processing such as correction processing on raw digital data transmitted from the X-ray detector 12 for each pixel. The processed data may be used for reconstruction.
[0034]
(X-ray detector)
Next, the above-described X-ray detector 12 will be described.
[0035]
As shown in FIG. 2, the X-ray detector 12 mounted on the X-ray CT scanner 1 of the present embodiment is generally perpendicular to the slice direction corresponding to the body axis direction of the subject P and the slice direction. A two-dimensional detector having a channel direction and a plurality of huge pixels arranged in a two-dimensional array along this two-dimensional plane. In addition, the channel direction of the X-ray detector 12 is curved in particular in consideration of the spread angle of the X-ray beam from the X-ray tube 11. The overall shape of the X-ray detector 12 depends on the application, and may be a flat plate.
[0036]
The X-ray detector 12 includes a plurality of detector blocks 12 that divide the two-dimensional surface into a plurality of surfaces.₁~ 12_J(J is a positive integer of 2 or more; hereinafter, represented by 12J), and the detector blocks can be detachably connected to each other. On the front surface of the detector block 12J on the X-ray incident side, an X-ray transmission image from the X-ray tube 11 is obtained with a collimator (not shown) made of molybdenum or tungsten arranged in the slice direction. Has become.
[0037]
Each detector block 12J is made of a compound semiconductor and has a layered semiconductor cell S of a predetermined size (for example, several centimeters × several centimeters), and a radiation incident surface of the semiconductor cell S is covered with a charging electrode E1 for voltage application. The semiconductor cell S has a monolithic structure in which the surface opposite to the radiation incident surface is covered with a plurality of current collecting electrodes E2 divided into a two-dimensional array. As a material of the semiconductor cell S, a cadmium telluride semiconductor (CdTe semiconductor), a cadmium zinc telluride semiconductor (CdZnTe semiconductor), a silicon semiconductor (Si semiconductor), or the like is used. A relatively high voltage of, for example, several tens of volts to several hundred volts is applied to the charging electrode E1. As a result, a pair of an electron and a hole is generated therein due to the X-ray photon incident on the semiconductor cell S, and the electron is collected by the current collecting electrode E2 having a relatively positive potential. The charge is detected as a pulse signal. That is, the X-rays incident on the radiation incident surface are directly converted into pulse signals of electric quantity.
[0038]
The collecting electrode E2 is divided in a grid pattern, and a plurality of electrodes E2₁~ E2_L(L is a positive integer of 2 or more). This collecting electrode E2₁~ E2_LDetermines the size of the collected pixels for X-rays. As a result, L collection pixel channels are formed for each detector block 12J, and N collection pixel channels are formed for the entire X-ray detector 12.
[0039]
Current collecting electrode E2₁~ E2_LIn this embodiment, each size is processed to a size sufficient to detect X-rays as its particles. In the present embodiment, the size at which X-rays can be detected as particles is defined as “between pulse signals responding to each incidence when a plurality of radiation (eg, X-ray) particles are continuously incident at the same position or in the vicinity thereof. Is a size that can substantially ignore the occurrence of the superimposition phenomenon of When this superimposition phenomenon occurs, the characteristic of the count number (input count number) with respect to the incidence of X-ray particles and the characteristic of the actual count number (output count number) with respect to the characteristics of the X-ray particles are reduced. For this reason, the size of the collected pixels formed in the X-ray detector 12 is set to a size that can be regarded as not generating or substantially not generating the counting-down characteristic.
[0040]
FIG. 3 shows the above-described counting-down characteristics. When an X-ray particle is incident on one collection pixel in a temporally close manner, a pulse signal detected with the continuous incidence has a two-modal waveform by superimposition of two pulse signals. For this reason, if this bimodal pulse signal is discriminated by the threshold value TH, the two pulse signals are counted as one pulse signal, resulting in erroneous detection. Generally, in the related art, as shown in FIG. 3, as the input count number increases, the output count number saturates or decreases due to the above-described superimposition of the detection pulse signal (that is, the count-down characteristic occurs). For this reason, in order to use a count region having no count-down characteristics, measures such as weakening the X-ray intensity and reducing the number of X-rays incident by forming each collection pixel small are taken. Usually, an X-ray detector such as an X-ray CT scanner having a collection pixel size of 2 to 3 mm²In the case of (1), it is required that the count rate increase linearly to about 500 to 1000 kcps (that is, the counting characteristic does not occur; see the dotted straight line in FIG. 3). In this case, the collected pixels of the X-ray detector have a value of about 100 to 200 μm. That is, in the case of such an X-ray CT scanner, it is required that the collection pixels be 200 μm or less.
[0041]
As described above, the pixel size required for the X-ray detector 12 relatively changes depending on the apparatus conditions and the like. For example, as described above, the value of 200 μm or less, or 0.1% or less of the radiation particles is used as described above. Pixel size that can be detected with the precision of
[0042]
Further, each detector block 12J of the X-ray detector 12 has various kinds of processing for processing pulse signals for each X-ray particle detected by the semiconductor cell S corresponding to each of the L collected pixel channels. A processing circuit C composed of a circuit is formed by an ASIC (application-specific integrated circuit), and is provided as an ASIC layer: AS. More specifically, the operation of the processing circuit C will be described later, but the charge amplifier 41, the waveform shaping circuit 42, and the multi-stage comparator 43 are arranged for each collection pixel.₁~ 43_n(N is a positive integer of 2 or more) Energy region distribution circuit 44 and multi-stage counter 45₁~ 45_n(See FIG. 4). These circuits 41 to 45 are provided for each collection pixel of the X-ray detector 12 described above (that is, for a total of N (= “L current collection electrodes E2”).₁~ E2_L) × “J detector blocks 12J”) for each of the collection pixel channels (N is a positive integer greater than or equal to 4). That is, the charge amplifier 41, the waveform shaping circuit 42, and the multi-stage comparator 43 for all the collected pixels of each detector block 12J.₁~ 43_n, Energy region distribution circuit 44, and multi-stage counter 45₁~ 45_nIs formed by the ASIC, and the collecting electrode E2 of the semiconductor cell S₁~ E2_LAre laminated and provided integrally.
[0043]
At this time, the ASIC layer AC is manufactured so as not to protrude from the end face of the semiconductor cell S for each detector block 12J. This makes it possible to easily configure the X-ray detector 12 having a wide incidence window by connecting the plurality of detector blocks 12J vertically and horizontally without any gap.
[0044]
FIG. 4 is an electrical block diagram mainly showing the X-ray detector 12 of the X-ray CT scanner 1. FIG. 5 shows an X-ray detector 12 schematically representing collected pixels arranged in a matrix and an outline of a processing circuit formed by an ASIC for each collected pixel.
[0045]
As shown in FIG. 4, the X-ray detector 12 includes a plurality of processing circuits C including an ASIC layer AS provided in a semiconductor cell (sensor unit) S for each collection pixel channel.
[0046]
As described above, each processing circuit C includes a charge amplifier 41, a waveform shaping circuit 42, and a multi-stage comparator 43.₁~ 43_n, Energy region distribution circuit 44, and multi-stage counter 45₁~ 45_nIs provided. The charge amplifier 41 is connected to the current collecting electrode E2 of the semiconductor cell S.₁~ E2_LAnd charges up the electric charge collected in response to the incidence of the X-ray particles, and outputs the electric charge as a pulse signal.
[0047]
The output terminal of the charge amplifier 41 is connected to a waveform shaping circuit 42 capable of adjusting a gain and an offset, and shapes the detected pulse signal by processing the waveform of the detected pulse signal with a gain and an offset that have been adjusted in advance. The gain and offset of the waveform shaping circuit 42 are adjustment parameters in consideration of non-uniformity with respect to the charge-up characteristics of each collection pixel of the semiconductor cell S. By adjusting the gain and offset of the waveform shaping circuit 42 for each collection pixel channel in advance by a calibration operation described later, it is possible to shape the waveform without such non-uniformity. As a result, the waveform-shaped pulse signal output from the waveform shaping circuit 42 of each acquisition channel has a characteristic substantially reflecting the energy amount of the incident X-ray particles, and this variation between the acquisition pixel channels. Is almost completely eliminated.
[0048]
The output terminal of the waveform shaping circuit 42 is connected to a plurality of comparators 43.₁~ 43_nAre respectively connected to the comparison input terminals. The plurality of comparators 43₁~ 43_nAre applied with reference values TH1 (ｎTHn) having different values. As described above, the reason why one pulse signal is compared with the different reference TH1 (、 THn) is that the energy amount of the incident X-ray particle falls into any one of the energy regions divided and set in advance. To do that information. As illustrated in FIG. 6A, which of the reference values TH1 to TH3 (when n = 3) exceeds the peak value of the pulse signal (that is, the energy amount of each incident X-ray particle). , The energy region to be discriminated is different. When such a peak value is between the reference values TH1 and TH2, the energy amount of the X-ray particle currently being measured is discriminated from those falling in the energy region 1, and when it is between the reference values TH2 and TH3, The energy amount of the X-ray particles being measured is discriminated from those falling into the energy region 2, and if the X-ray particles are equal to or more than the reference value TH3, the energy amount of the X-ray particles being currently measured is discriminated from those falling into the energy region 3. . Note that the lowest reference value TH1 is set as a threshold value that does not detect disturbance, white noise from the semiconductor cell S or the charge amplifier 41. Further, the number of reference values, that is, the number of comparators is not necessarily limited to three (the number of discriminable energy regions = 3), and may be two or four, and in some cases, one. It may be. In the case of one, information as to whether or not X-ray particles have entered can be obtained.
[0049]
FIG. 7 shows a setting example of the reference values TH1 to THn. The waveform shown in the figure shows the energy spectrum of X-rays emitted from a normally used X-ray tube. The count intensity on the vertical axis is an amount dependent on the X-ray tube current, and the energy spectrum on the horizontal axis is an amount dependent on the tube voltage of the X-ray tube. For this spectrum, the first reference value TH1 is set to a value that allows the number of X-ray particles to be discriminated from the uncountable region and the lower energy region 1, and the second reference value TH2 is set to the lower energy region. 1 and the medium energy region 2 are set to values that can be distinguished from each other, and the third reference value TH3 is set to a value that allows the medium energy region 2 and the higher energy region 3 to be distinguished from each other. ing. In the present embodiment, for example, by setting the three reference values TH1 to TH3 as described above, the incident X-rays are detected as being regarded as particles, and then classified into the energy regions 1 to 3, respectively. The set number and set values of the reference values TH1 to THn are arbitrary.
[0050]
Multiple comparators 43₁~ 43_nIs connected to an energy region distribution circuit 44 as shown in FIG. The energy area distribution circuit 44 includes a plurality of comparators 43₁~ 43_nHas a function of reading which comparison result is turned on (off) from among the outputs, and allocating to which region the energy of the X-ray particles currently being measured falls. Specifically, the circuit 44 is formed by a combination of logic circuits. For example, in the case of the pulse signal shown in FIG. 6A, the peak value (energy amount) of the pulse signal indicates the energy region 2, and therefore, as shown in FIG.₁, 43₂Is turned on, and the comparator 43_nIs turned off. Therefore, the energy area distribution circuit 44 reads and decodes the on / off information, and outputs an energy distribution signal from its output terminal. As an example, in the case of FIG. 6B, this energy distribution signal is a signal in which only the signal indicating “energy region 2” is turned on. As a result, distribution information that the energy of the X-ray particle currently measured enters the “energy region 2” is obtained.
[0051]
A plurality of output terminals of the energy area distribution circuit 44 are connected to a plurality of counters 45.₁~ 45_nAre connected to each other, and the energy distribution signal is separately supplied to a plurality of counters 45.₁~ 45_nGiven to each. Therefore, the counter 45₁~ 45_nCan increment (count up) the count value only when the input energy distribution signal is ON, and can measure the number of X-ray particles entering each assigned energy region over a certain period of time. As an example, in the situations shown in FIGS. 6C and 6D, the energy of the X-ray particles enters the energy region 2, and therefore, the counter 45 that measures the energy region 2 is used.₂Is incremented.
[0052]
In this manner, the plurality of counters 45 during a certain time until resetting are performed.₁~ 45_nThus, the number of X-ray particles incident on the X-ray detector 12 is measured for each collection pixel and for each energy region. This count value, that is, the count value of the number of X-ray particles,₁~ 45_nIs read out as digital amount detection data (raw data). This data reading is performed, as shown in FIG. 5, by a counter 45 for each collection pixel in the ASIC layer AS.₁~ 45_nAre read out via a shift register 46 integrally formed with the data. The shift registers 46 are arranged in a matrix in a circuit in the ASIC layer AS, and N ′ (N × n) are connected to control lines CL connected to the plurality of shift registers 46 in the respective detector channel directions. ) Series count data (particle number measurement data; that is, raw data)) 1 to N ′ (K₁₁~ K_Nn) Is externally given a read pulse for serially reading). As a result, the count data 1 to N ′ (K) of each collected pixel are read from the read lines RD arranged in the slice direction (body axis direction).₁₁~ K_Nn) Is read out serially.
[0053]
The count data 1 to N ′ (K₁₁~ K_Nn) Is sent to the console 5 and temporarily stored in the storage unit 22 via the interface 21. According to a predetermined scanning method, the scanning is performed by moving the top of the bed or by fixing the position of the top to the pair of the X-ray tube 11 and the X-ray detector 12 around the subject. Then, X-rays are emitted at each rotation position moved by a predetermined angle during the rotation, and an X-ray scan is performed on the subject. As a result, count data (raw data) of the number of X-ray particles incident on each collection pixel at each rotation position is detected as view data and sequentially stored in the storage unit 22. Therefore, the reconstruction unit 17 reads the data of each view necessary for image reconstruction from the storage unit 22 at an appropriate timing, reconstructs this data by a desired reconstruction algorithm, and performs the reconstruction based on the number of X-ray particles. Obtain a CT image.
[0054]
(Calibration)
Here, the above-described calibration will be described.
[0055]
In order to accurately count the number of X-ray particles in each energy region, it is desirable that each reference value TH can be set with an accuracy of about 0.25 keV. For this reason, each reference value TH is calculated by allocating an energy range of 0 to 150 keV to a 10-bit digital value, converting the reference value TH into an analog voltage by a D / A converter built in each module, and₁~ 43_nRespectively. However, on the other hand, naturally, the pulse signal (the peak value indicates energy) detected via the semiconductor cell S or the charge amplifier 41 varies among the collected pixels in terms of the gain and the offset. I have. If there is this variation, the above-described highly accurate setting of each reference value TH loses its meaning, and it becomes almost impossible to count the number of X-ray particles with high accuracy for each energy region.
[0056]
For this reason, the X-ray CT scanner 1 according to the present embodiment is provided with a calibration device that externally adjusts the offset and gain of each waveform shaping circuit 42 for each collected pixel in order to correct the above-described variation. . FIG. 8 shows an outline of this calibration device for the detector block 12J for one collected pixel. This calibration device includes a counter 45₁~ 45_n9 and a memory 52 for storing a reference value measured and set in advance, and the adjuster 51 operates according to the calibration algorithm shown in steps S1 to S5 in FIG. The offset and gain of the waveform shaping circuit 42 are adjusted using the reference value.
[0057]
As a specific adjustment method, as a different reference source,²⁴¹Am and⁵⁷A gamma ray source F sealed with Co is placed on the X-ray detector 12. The adjuster 51 sets a predetermined range of energy energy regions 1 and 2 around 59.5 keV and 122 keV corresponding to the energy of 241 Am and 57 Co, respectively, on the algorithm. For example, as shown in FIG. 10, one calibration energy region 1 is selected in a narrow range of 56.5 to 62.5 keV, and the other calibration energy region 2 is selected in a narrow range of 119 to 125 keV. Is set. Therefore, the adjuster 51 repeatedly adjusts the gain and the offset of the waveform shaping circuit 42 in the procedure shown in FIG. 9 and finally converges so that the count value becomes maximum in each of the energy regions 1 and 2. In general, the influence of the offset with respect to the above-described variation is larger than that of the gain, so it is preferable to perform the offset adjustment first. The adjuster 51 causes this adjustment to be performed for all collected pixels. Note that this calibration work is performed in advance, for example, at the time of installation of the X-ray CT scanner, at the time of regular service / maintenance, or before scanning a patient before activation. By this calibration, the energy discrimination result based on the reference value TH can be held with extremely high accuracy.
[0058]
(Design example 1)
Here, when the above-described X-ray CT scanner 1 is assumed to be used under the most severe X-ray incidence conditions at present, a design example of a counter in the X-ray detector 12 and output of the detection information to the outside are described. An example will be described.
[0059]
[Outside 1]

[0060]
It is sufficient if the total number of bits of the counters 1 to 3 can be measured up to 4 bits and up to 15 counts. Each counter configuration has a certain degree of freedom. Actually, 3 bits are required, and a total of 9 bits is sufficient.
[0061]
Therefore, the 9-bit count value can be output as a serial signal with a transfer time of 130 nsec / 9 ≒ 14.5 nsec per bit based on the above condition (6). Each pixel is collected by repeating count reset at 1 msec. In this collection, the time required to output all count values is as short as 130 nsec. During this 130 nsec period, even if the count measurement is stopped in order to prevent digital noise from jumping into an analog system such as a charge amplifier, the effect is not affected. Can be ignored.
[0062]
Also, if the information of the reference values TH1 to TH3 described in (7) is configured to be preset before data collection using the same line, the circuit configuration can be simplified. Further, in the case where the X-ray detector 12 configures a detector for a CT scanner by densely arranging the detector modules, a unit for reading out a serial signal can be configured as each detector module.
[0063]
(Design example 2)
Next, the size of the collection pixels for measuring the number of electrical pulse signals in response to the incidence of the X-ray particles “accurately” (ie, with substantially negligible occurrence of superposition between the incident particles) A design example will be described.
[0064]
[Outside 2]

[0065]
In the X-ray CT scanner, since there is no report of a case of reconstruction in the pulse measurement mode, the collection pixel size in the case of SPECT (Single Photon Emission Computed Tomography) is calculated first. Normally, in SPECT, when 4 mm × 4 mm is set as a collection pixel, a good reconstructed image cannot be obtained unless information of about 20 kcps is collected for about 30 minutes. The total count required for this is 20,000 × 30 × 60 = 36 Mcounts.
[0066]
The collection pixel size at this time is 16 mm²So, 1.5mm of X-ray CT scanner²When converted to
(Equation 1)

384 Mcounts is required.
[0067]
In the case of an X-ray CT scanner, as schematically shown in FIG. 11, X-rays that do not pass through the human body also enter the detector, so that it can be assumed that the X-ray intensity is about 100 times as high. With this assumption in mind, a minimum of 38400 Mcounts counts must be assumed to allow X-ray particles to enter each acquisition pixel.
[0068]
According to the current conditions, the acquisition time is 0.5 sec, so that the X-ray detector as a whole is:
(Equation 2)
38400 × 2 = 76800 Mcps
Count is required. Converting this to one collection pixel,
(Equation 3)

Becomes From this, in order to suppress pulse measurement to about 10 kcps,
(Equation 4)

Area. Therefore, the size of each collection pixel is
(Equation 5)

Becomes Therefore, it can be concluded that unless an acquisition pixel size of about 200 μm is set, it is difficult to measure the so-called “accurate” pulse number.
[0069]
For the above reasons, in the present invention, the X-ray incidence condition is the strictest, and the upper limit of the collection pixel size is about 200 μm when performing “accurate” pulse measurement only depending on the size of the collection pixel size. This upper limit varies depending on the X-ray incidence conditions. When such conditions are more moderate, the upper limit of the collection pixel size may be larger.
[0070]
(Effect)
As described above, according to the X-ray detector 12 according to the present embodiment, unlike the conventional simple pulse measurement mode, X-rays can be detected in the pulse measurement mode that takes into account the energy amount of each X-ray particle. The main factor that enables the detection operation in this pulse measurement mode is a collection pixel that can ignore the occurrence of a phenomenon in which pulse signals of X-ray particles incident on each collection pixel in a temporally close proximity are mutually superimposed. The size is adopted, and the incident information of the X-ray particles is once replaced with the energy amount for discrimination.
[0071]
Therefore, according to the X-ray detector 12 and the X-ray CT scanner 1 equipped with the same according to the present embodiment, it is possible to obtain a CT image with less influence of scattered radiation and a clear contrast. By avoiding the influence of the dark current generated from the charge amplifier 41 in the subsequent stage, high-sensitivity X-ray detection with greatly improved S / N can be performed. Further, since the X-ray detection sensitivity is high, the tube current of the X-ray tube can be reduced to perform detection with the same or higher accuracy, so that the X-ray exposure dose to the subject can be reduced. Further, by using the tube current of the X-ray tube at a low level, the failure of the X-ray tube is reduced, and the life of the X-ray CT scanner can be extended. Further, since an A / D converter, which is a factor of increasing the size and increasing the cost as in the conventional X-ray detector operated in the integration mode, is not required, the configuration of the X-ray detector 12 as a whole is Can be simplified, and the total manufacturing cost can be reduced. Therefore, the size of the X-ray detector can be further increased.
[0072]
Furthermore, since the X-ray detector 12 according to the present embodiment uses a CdTe semiconductor as its semiconductor cell (sensor unit) S, a highly accurate detection capability is also ensured. This detection ability will be described.
[0073]
At present, a pseudo-Schottky type CdTe is optimal for considering X-rays as particles, shaping the waveforms at high speed, and setting the minimum threshold level at about 10 keV in consideration of the dark current level. In particular, if the detector is constituted by a monolithic (Monolithic) structure and the thickness of CdTe is constituted by 0.5 mm to 1.0 mm, excellent X-rays can be obtained as compared with conventional CsI, Ge, Si, X-ray films and the like. Can be obtained. In addition, in the case of this thickness, the energy resolution of, for example, 122Co, which is a gamma ray source, is 80% or less, and the threshold value of energy can be detected extremely well. In addition, since CdTe directly converts X-rays into electric charges, there is no need to perform photoelectric conversion using a photodiode unlike CsI, so that conversion loss is small.
[0074]
On the other hand, when a scintillator is used as in the prior art, in order to form a collection pixel, it is necessary to cut the scintillator into pixel units, apply a reflective agent between the collection pixels, and isolate the scintillation light from each other. . For this reason, this isolation cannot be neglected and a dead space is created, but a certain loss occurs in light reaching the photodiode, which causes a decrease in sensitivity.
[0075]
Taken together, it can be said that CdTe is a material with excellent detection sensitivity compared to other X-ray detectors. In addition, since the energy below the set minimum level can be prevented from being detected by the threshold value detection, it is almost certain that a dark current is mixed in or a scattered ray that shifts to a lower energy side is automatically mixed in. Therefore, an image with extremely high contrast can be obtained.
[0076]
Hereinafter, various devices employing the above-described X-ray detector 12 will be described as another embodiment.
[0077]
(Second embodiment)
Next, as a second embodiment of the radiation imaging system of the present invention, another X-ray diagnostic apparatus employing the X-ray detector according to the present invention will be described with reference to FIGS.
[0078]
According to the X-ray detector 12 described above, data corresponding to the count value of the number of incident X-ray particles is detected for each of the collected pixels and for each energy region. An image can be selectively formed in response to the request. Therefore, in this embodiment, an X-ray diagnostic apparatus having a function of performing selective image reconstruction in accordance with the energy region and capable of performing angiography with high sensitivity even when a thin contrast agent is administered from a vein Will be described.
[0079]
As is well known, coronary artery imaging using an iodine thread contrast agent has been practiced as being effective for detecting cardiovascular occlusion. However, in the case of the current integration mode type X-ray detection, there is a disadvantage that a sufficient contrast cannot be obtained unless a contrast agent is injected into an artery and an image is formed in a state where the concentration of the contrast agent is high. In addition, it is said to be an extremely invasive examination method by injection from a patient's artery.
[0080]
In order to improve this situation, a monochromatic X-ray generator is used to focus monochromatic X-rays near the K-absorption edge energy (33.17 keV: see FIG. 12) of the iodine-based contrast agent, and irradiate the patient with the monochromatic X-rays. The law is being actively studied. In other words, by using this monochromatic X-ray, an image is formed with an energy having a value larger or smaller than the K absorption edge energy, and subtraction is performed between the images. It is intended to depict a coronary artery. However, large-scale equipment such as a monochromatic X-ray generator is required to perform this imaging, and there is a limit to widespread use of this imaging method.
[0081]
Therefore, in the X-ray diagnostic apparatus 1A according to the present embodiment, as shown in FIG. 12, three energy discrimination reference values are set at an energy position of at least 33.17 KeV and two positions above and below the energy level sandwiching the energy position. TH1 to TH3 are set, whereby the energy regions 1 and 2 are set around 33.17 keV.
[0082]
The count data of the energy regions 1 and 2 are selectively collected, and the two types of count data are subtracted from each other. As a result, it is possible to obtain raw data that is substantially equivalent to that obtained when imaging with monochromatic X-rays, and it is sufficient to reconstruct this raw data. The two types of count data may be reconstructed separately and then subtracted from each other.
[0083]
FIG. 13 shows a configuration example of an X-ray diagnostic apparatus for realizing this. This figure shows only an outline of a different part of the configuration of FIG. 4 described above. Each detector block 12J of the X-ray detector 12 has count data K of X-ray particles entering the energy regions 1 and 2._i1, K_i2(I = 1 to N) counter 46₁, 46₂Is provided. This count data K_i1, K_i2Are detected for each collected pixel and sent to the console 5. Count data K input to console 5_i1, K_i2Is sent to the pixel operation circuit 31 through the interface 21. This pixel operation circuit 31 calculates the count data K_i1, K_i2, A difference circuit 31A for calculating the difference between them, an addition circuit 31B for adding them to each other, and count data K_i1, K_i2Selection circuits 31C and 31D for selectively passing each of them are provided. In addition, the pixel operation circuit 31 can selectively operate one or more of the circuits 31A to 31D in response to a switching control signal SC provided from the controller 20. Has become.
[0084]
The count data “K” calculated or selected by the pixel calculation circuit 31_i1-K_i2”,“ K_i1+ K_i2”,“ K_i1And / or "K_i2Are stored in the storage unit 22 shown in FIG. The reconstructing unit 23 outputs “K” related to the difference operation among the count data stored in the storage unit 22._i1-K_i2Is read out and the image is reconstructed. As a result, a subtraction image between the energy regions 1 and 2 is obtained, so that highly sensitive angiography by intravenous administration of a contrast agent can be performed.
[0085]
As described above, by using the X-ray detector 12 according to the present invention, an image of only the coronary artery is sharpened by selectively using a high-accuracy count value for each energy region of incident X-ray particles of each collected pixel. Even in an X-ray diagnostic apparatus equipped with an X-ray source of normal intensity, a contrast agent can be injected from a vein to image a coronary artery. The burden on the patient can be reduced to the level of a normal routine examination, and a great breakthrough advantage can be obtained in that coronary angiography can be released from the field of advanced medical care.
[0086]
According to the X-ray diagnostic apparatus 1 of the present embodiment, the difference count data “K_i1-K_i2”And the additional count data“ K_i1+ K_i2”, The count data“ K_i1And / or "K_i2"Can also be collected selectively. Therefore, the addition count data “K_i1+ K_i2, It is possible to obtain an X-ray image corresponding to a general X-ray diagnostic apparatus (that is, an X-ray image reconstructed from data of an energy amount comprehensively including the entire spectrum). Furthermore, the count data “K_i1And / or "K_i2Can be used for image reconstruction. As described above, since the image of the specific energy component of the X-ray particle can be selectively provided, it is possible to sharpen the image information and contribute to the diagnosis.
[0087]
(Third embodiment)
Next, as a third embodiment according to the radiation imaging system of the present invention, an X-ray CT scanner employing the X-ray detector according to the present invention will be described with reference to FIGS.
[0088]
In this X-ray CT scanner, a composite device called IVR-CT is intended to be carried by one X-ray CT scanner.
[0089]
2. Description of the Related Art In recent years, angiographic applications have been diversified with the rapid progress of IVR (International Radiology) technology. For this reason, by combining an angiography apparatus and a high-performance X-ray CT scanner, a so-called IVR-CT, which can perform a CT angiography performed by a conventional patient transfer in a single imaging room, is widely used. It is becoming. However, since this conventional IVR-CT requires two types of devices, an angiography device and an X-ray CT scanner, the entire device becomes very large and the operation of the device is complicated, and the wide installation is required. There were problems such as the need for space and the high installation and operation costs.
[0090]
Therefore, the X-ray CT scanner 1 according to the present embodiment adopts the configuration shown in FIG. 13 as shown in FIG. 14 and also includes a pixel bundling circuit 32 and a switching circuit on the output side of the pixel data operation circuit 31 of the console 5. 33. The pixel bundling circuit 32 has N pieces of data (here, the count data “K_i1-K_i2”,“ K_i1+ K_i2”,“ K_i1And / or "K_i2)) Are added to each other by a predetermined number to create j data.
[0091]
By this addition, as shown in FIG. 15, it is equivalent to a case where all the collected pixels in the X-ray detector 12 are bundled into a predetermined number (region) of pixel groups to form one pixel. More specifically, as described above, the collection pixels of the X-ray detector 12 according to the present embodiment are formed much more finely than those of the conventional X-ray detector for a CT scanner. It has a size of × 200 μm. Therefore, when obtaining a normal X-ray plane image, a high-definition image is obtained by using such fine pixels as they are. On the other hand, when obtaining a CT image, this pixel size takes too much time for reconstruction. In other words, since the pixels are too fine, some adjacent pixels are bundled together to output one data (see the cross-hatched portion in the figure). In the example of FIG. 15, 5 × 5 collected pixels are bundled by signal addition and regarded as one pixel. Thereby, a pixel having a size of 1 mm × 1 mm can be apparently formed.
[0092]
Returning to FIG. 14, the pixel bundling circuit 32 is provided with an output line OL that bypasses the pixel bundling circuit 32. The output line OL and the output line of the pixel bundling circuit 32 are connected to each other by the switching circuit 32. Either one can be selected to store data in the storage unit 22. The pixel data arithmetic circuit 31 and the switching circuit 32 are switched or switched according to the respective modes (CT imaging mode, normal X imaging mode) by the switching control signals SC1 and SC2 given from the controller 20. .
[0093]
That is, when performing the blood vessel rendering in the CT imaging mode, the pixel data calculation circuit 31 is switched to the difference circuit 31A, and the switching circuit 32 is switched to the output line side of the pixel bundling circuit 32. On the other hand, in the normal X-ray imaging mode, the pixel data calculation circuit 31 is switched to the difference circuit 31A, and the switching circuit 32 is switched to the bypass output line OL. Switching of the pixel data operation circuit 31 to the addition circuit 31B or the selection circuits 31C and 31D is instructed via the controller 20 when the operator desires.
[0094]
For this reason, according to the X-ray CT scanner 1 for IVR-CT according to the present embodiment, an iodine thread contrast agent is injected into a subject into a vein, normal X-ray imaging is performed, and a blood vessel is drawn by a planar image. be able to. Since this blood vessel delineation is usually performed in the X-ray imaging mode, the acquired pixel size is, for example, 200 μm × 200 μm, and an extremely high-definition image is obtained. At this stage, by switching the circuit of the pixel data operation circuit 31 to the difference circuit 31A, it is possible to obtain a subtraction image in consideration of the energy region.
[0095]
After observing the occlusion state of the blood vessel by observing the high-definition images collected from these appropriate and various viewpoints, the couch 4 is positioned with respect to the target blood vessel, and immediately thereafter, CT imaging is performed on the spot. The mode can be switched to CT imaging. By collecting the data of the CT imaging, it is possible to obtain a CT image obtained by subtracting the reconstructed image separated into the energy regions 1 and 2 described above.
[0096]
As described above, according to the X-ray CT scanner 1 according to the present embodiment, the functions of the conventional angiography apparatus and the high-performance X-ray CT scanner can be sufficiently covered by one apparatus. For this reason, a much higher-quality planar image can be used for positioning and the like than before, so that the positioning accuracy is increased and CT imaging can be performed on the spot, so that the burden on both the patient and the operator is significantly reduced. You. In addition, less installation space is required, and installation and operation costs are also reduced, which is advantageous in terms of hospital equipment investment. Furthermore, according to this X-ray CT scanner, the function of collecting a normal X-ray plane image can be omitted if necessary, and instead, only a blood vessel can be precisely displayed three-dimensionally by a subtraction CT image. May be.
[0097]
(Fourth embodiment)
Next, as a fourth embodiment of the radiation imaging system of the present invention, another X-ray CT scanner employing the X-ray detector according to the present invention will be described with reference to FIG.
[0098]
This X-ray CT scanner relates to an X-ray CT scanner suitable for mass screening for early detection of lung cancer.
[0099]
FIG. 16 shows a configuration example of the X-ray CT scanner. The X-ray detector 12 has three counters 47 corresponding to each collected pixel.₁~ 47₃Is provided. This counter 47₁~ 47₃Respectively counts the number of incident X-ray particles in energy regions 1 to 3 divided by predetermined four different energy thresholds TH1 to TH4 (not shown) set on the spectrum of X-ray energy, and count data K_i1, K_i2, K_i3(I = 1 to 3) and output. In contrast, the console 5 includes an interface 48, a bundling network 49, and three pixel bundling circuits 50 corresponding to the N collected pixel channels.₁~ 50₃Is provided.
[0100]
Count data K from each collection pixel channel_i1, K_i2, K_i3Are input to the bundle network 49 via the interface. A bundling network 49 connects each of the three count output paths of each pixel collection channel to three pixel bundling circuits 50.₁~ 50₃To be input separately.
[0101]
Pixel bundling circuit 50₁~ 50₃Adds N input signals for each of a desired number of mutually adjacent pixel groups in response to a switching control signal SC3 provided from the controller 20. Thus, the collected pixels can be bundled for each desired number of pixels. At this time, the pixel bundling circuit 50₁~ 50₃Are configured to be able to change the number of pixels to be bundled (pixel group) over a plurality of types according to the instruction content of the switching control signal S3. For example, if the basic collection pixel is 200 μm, the value can be changed to a desired value among four types of 400 μm, 800 μm, and 1600 μm by bundling, in addition to 200 μm which is output in an unbundled state. This binding method can be selected and commanded by the operator.
[0102]
This pixel bundling circuit 50₁~ 50₃Is responsible for bundling pixels for the energy regions 1 to 3. Therefore, each pixel bundling circuit 50₁(~ 50₃) Can output count data of a desired size from, for example, four types of collected pixel sizes of 200 μm, 400 μm, 800 μm, and 1600 μm via the switching control signal SC3.
[0103]
Therefore, by accumulating count data for each desired energy region and for each desired collection pixel size and reconstructing the image, a CT image can be obtained with an energy range and resolution suitable for delineating the characteristics of lung cancer.
[0104]
For this reason, the resolution of the current reconstructed image is generally about 0.5 mm, but by reducing the reconstructing time, it is possible to depict a resolution of 0.1 mm or less. In addition, since the image can be reconstructed using X-rays falling within the energy window of interest, it is possible to perform a depiction focused on lung cancer, and the detection capability for lung cancer is dramatically increased. Therefore, screening for lung cancer in a truly early stage can be performed in a short time. With the widespread use of such X-ray CT scanners, it is possible to perform mass screening for early detection of lung cancer.
[0105]
(Fifth embodiment)
Next, as a fifth embodiment according to the radiation imaging system of the present invention, a mammography apparatus (breast X-ray imaging apparatus) employing the X-ray detector according to the present invention will be described with reference to FIGS. I do.
[0106]
When a breast cancer is detected using X-rays, it is important to detect calcified micro-tissue from the viewpoint that the cancer is discovered at an early stage of the cancer. The calcified tissue exhibits a remarkable difference in absorption coefficient from other normal tissues to X-rays having an energy of about 20 keV, but the cancerous (calcified) tissue has a difference in absorption from normal tissues. Few. Since the microcalcification has a detection sensitivity to X-rays having relatively low energy, it is necessary to photograph a planar image focused only on low energy with high contrast.
[0107]
In the present embodiment, as shown in FIG. 17, the beam quality using the filter is shifted to the low energy side, and the patient exposure dose is reduced to the energy region 1 centered on 20 KeV and the energy region 2 on the low energy side. Count data is collected for each energy region in a divided state.
[0108]
The specific configuration of the mammography apparatus is the same as that of the angiography system. The energy threshold is, for example,
[Outside 3]

Set to.
[0109]
As in the case of the above-described angiography, by subtracting the count data collected from each of the energy regions 1 and 2 from each other, a high-contrast X-ray image of only the calcified tissue at an early stage can be obtained. This allows early detection of breast cancer.
[0110]
In the case of this X-ray image, the number of counts of X-ray particles per collected pixel needs to be larger than in the case of an X-ray CT scanner. In this regard, the time at which the counter built into each collection pixel overflows is calculated according to the dose intensity, and the counter is periodically reset in a shorter time, while the collection count is set during the X-ray exposure time. By integration, a mammography image with a high count value can be obtained.
[0111]
In the case of conventional mammography, in order to detect a calcified tissue, it was necessary to make the collected pixels finer and increase the detection performance. However, according to the present embodiment, since the detection sensitivity is high, about 200 μm Even with the collection pixel size of, microcalcification can be drawn.
[0112]
Further, in the present embodiment, since the X-ray detector 12 is made of a CdTe semiconductor, the detection surface (X-ray incidence surface) can be formed to the very end of the detector end surface (for example, within 5 mm). For this reason, as shown schematically in FIG. 18, by using the compression plate 60, an extremely large breast region can be imaged at a time. Further, in the case of the present embodiment, since a scattered component of 10 keV or less is not detected, an X-ray image with very small scattered component and good contrast is obtained.
[0113]
(Sixth embodiment)
Next, as a sixth embodiment according to the radiation imaging system of the present invention, a baggage inspection apparatus employing the X-ray detector according to the present invention will be described with reference to FIG.
[0114]
At present, at airports and the like, industrial X-ray CT scanners have been introduced for baggage inspection. However, due to low sensitivity, it is necessary to increase the X-ray intensity. As a result, there is a disadvantage that a normal undeveloped photographic film is exposed. In addition, the current X-ray CT scanner cannot operate at a high speed enough to cover all of the baggage of a large number of people at the airport. For this reason, the entire number of baggage is once inspected by ordinary X-rays, and only suspicious baggage is precisely inspected by an X-ray CT scanner, which is troublesome and time-consuming.
[0115]
Therefore, the baggage inspection device 70 according to the present embodiment, as shown in FIG. 19, mounts a pair of the X-ray tube 11 and the X-ray detector 12 according to the present invention on a gantry 71 so as to be rotatable. The belt conveyor 72 is allowed to pass through the opening OP penetrating therethrough. Since the baggage carried at a constant speed by the belt conveyor 72 passes through the opening OP of the gantry 71, the X-ray tube 11 and the X-ray detector 72 rotate around the opening OP, so that the X-ray CT Helical scan is possible. Through this scan, count data (including count data for each energy region) corresponding to the number of X-ray particles detected by the X-ray detector 12 as described above is detected.
[0116]
This count data is sent to the reconstruction device 72. The reconstructing device 72 generates the CT image data by the time the luggage finishes passing through the gantry 71. This CT image is displayed on the monitor 73 and is presented to the observer.
[0117]
In addition, the data of the CT image is sent to the automatic dangerous goods recognition device 74. As a result, the dangerous substance automatic recognition device 74 receives CT image data based on count data in which the energy area is narrowed down to a preset dangerous substance (for example, metal or drug). Therefore, the automatic dangerous substance recognition device 74 performs an automatic dangerous substance recognition process (feature extraction program) on the CT image, automatically determines the presence or absence of the corresponding dangerous substance, and uses the result as an abnormal information detection and presentation device 75. Send to
[0118]
According to the present embodiment, a plurality of pieces of baggage can be inspected in real time and continuously by the function of the X-ray CT scanner. That is, since the X-ray detector 12 capable of detecting the count data for each energy region is used, the image reconstruction is focused on high energy suitable for detecting dangerous substances requiring high sensitivity, and is reduced to low energy such as drugs. A wide range of image reconstruction is possible. For this reason, all the baggage can be continuously applied to the CT scan. Further, it is possible to detect dangerous substances efficiently and unmannedly. Further, in addition to real-time CT image acquisition at the site, it is possible to post-edit a CT image of only the energy of interest, which makes it easy to depict dangerous substances. As a result, a more accurate inspection can be performed in a short time, and a semi-automated baggage inspection or a fully automated baggage inspection using a visual inspection can be performed. Of course, since the detection sensitivity is excellent, the X-ray intensity can be reduced correspondingly, and the problem of exposure to an undeveloped photographic film can be solved.
[0119]
(Seventh embodiment)
Next, as a seventh embodiment according to the radiation imaging system of the present invention, a pipeline inspection system 79 employing an X-ray detector according to the present invention will be described with reference to FIG.
[0120]
Inspection of cracks in oil pipelines usually installed in cold regions such as Alaska and Sakhalin is extremely difficult. It would be best if the X-ray non-destructive inspection equipment could be operated automatically along this oil pipeline to perform X-ray non-destructive inspection, but it is necessary to generate X-rays in the environment where this pipeline is laid. Is not realistic.
[0121]
Therefore, as schematically shown in FIG. 20, the pipeline inspection system 79 according to the present embodiment uses a sealed gamma ray source. This is because the radiation detector according to the present invention can render an image in an energy region.
[0122]
Specifically, the above-mentioned X-ray detector 12 is used for gamma ray detection, and the linear gamma ray detector 12 and the linear sealed gamma ray source 80 are arranged to face each other via a pipeline 81. As a gamma ray source,¹²⁵Co,⁷⁵Se,¹³⁷Cs or the like is used. On the gamma ray emission side of the gamma ray source 80, a collimator 82 for guiding gamma rays in parallel is provided integrally. A traveling mechanism (not shown) is provided for moving the set of the gamma ray detector 12 and the gamma ray source 80 (collimator 82) along the longitudinal direction thereof while appropriately rotating the pipeline around the pipeline 81.
[0123]
Thereby, gamma rays emitted in parallel from the gamma ray source 80 through the collimator 82 pass through the pipe and are detected by the gamma ray detector 12 as a transmission image. As a result, count data corresponding to the number of gamma ray particles is obtained from the gamma ray detector 12. At this time, as described above, a threshold (for example, two thresholds TH1 and TH2, and the energy range is set to ± 10% of the sealed gamma ray energy) is set for discriminating the energy region from the spectrum of the gamma ray. Therefore, the count data also becomes data corresponding to the energy region set by the threshold value.
[0124]
The count data detected by the gamma ray detector 12 is sent to a radio transmitter 84 via a memory 83, and is transmitted by radio to an antenna 86 at another location together with a position signal from a position signal 85. The radio wave received by the antenna 86 is processed by the receiver 87, and the processed received data is sent to the pipeline abnormality detection device 88. The abnormality detection device 88 performs appropriate processing such as reconstructing a CT image based on the received data, and provides information on the presence or absence of abnormality such as damage to the pipeline and internal distortion and content information.
[0125]
As a result, image data such as cracks in a pipe can be obtained at a remote location, and even a pipe run installed in a severe natural environment can be automatically and automatically inspected.
[0126]
In addition, such an inspection can be used for an automatic inspection of a factory pipeline as long as a structure capable of blocking gamma rays from the outside can be taken.
[0127]
Further, an example of inspecting an internal structure using a CT scanner employing a detector by X-rays or gamma rays according to the present invention is not limited to a pipeline,
[Outside 4]

It is also applicable to
[0128]
For example, in the current lithium ion inspection, the defect inspection of the electrode portion is performed by X-ray fluoroscopy, but the current X-ray imaging time is too long because the pitch time for flowing the product is short. Therefore, the same X-ray non-destructive inspection apparatus was provided in parallel, and only the inspection process was performed in parallel with the three steps. However, by using the radiation detector according to the present invention, one non-destructive inspection apparatus was used. In order to be able to do the inspection. In addition, since a high-contrast image is obtained based only on energy more suitable for inspection, defective products can be extracted with high accuracy. Therefore, it is possible to significantly reduce the cost for the inspection.
[0129]
Also, when analyzing the internal structure of a metal product by CT, X-rays cannot pass through the metal unless the energy of the X-rays is high, so that conventionally, only an X-ray image with very low contrast is obtained, and CT is applied. Less commonly used for analysis. However, according to the radiation detector of the present invention, since the energy of interest can be distinguished sharply, a reconstructed image of, for example, only a high energy component can be obtained, and even in the case of X-ray CT, the internal structure can be more clearly depicted. .
[0130]
Furthermore, by inspecting the cut wood with a system having the same configuration as the above-described baggage inspection system, it is possible to know the presence or absence of a hollow inside the wood. As a result, the internal structure of the timber can be surely known, which contributes to improving the efficiency of wood processing and stabilizing the quality of timber products.
[0131]
As described above, the X-ray detector according to the present invention has extremely high sensitivity in all fields of non-destructive X-ray inspection, and has an excellent usefulness due to its ability to distinguish energy regions. Can be demonstrated.
[0132]
(Eighth embodiment)
Next, as an eighth embodiment according to the radiation imaging system of the present invention, a two-dimensional image system for X-ray astronomy employing the X-ray detector according to the present invention will be described with reference to FIG.
[0133]
With recent advances in X-ray astronomy, taking X-ray images separated by energy is indispensable for structural analysis (component analysis) of the universe.
[0134]
Therefore, as shown in FIG. 21, the sensor device 91 of the two-dimensional image system 90 is caused to stay in outer space by an artificial satellite launched into a geosynchronous orbit in outer space. This sensor device 91 uses the X-ray detector 12 according to the present invention. The X-ray detector 12 is a 0.5 mm thick pseudo-Schottky CdTe pixel detector, which is cooled to -20 ° C. for use. Thus, in addition to the high X-ray sensitivity, the X-ray detector 12 has the same energy resolution as the Ge detector, which currently has low detection sensitivity but has the highest energy resolution. Therefore, a detector that requires rigor such as X-ray astronomy can exhibit unprecedented characteristics.
[0135]
A collimator 92 is provided on the X-ray incidence surface of the X-ray detector 12. In the processing circuit provided in the detector 12, the energy region is set to M stages (M is a positive integer of 2 or more) so that at least characteristic X-rays can be captured.
[0136]
Therefore, the count data from the X-ray detector 12 is sent to the characteristic X-ray detection circuit 93, and the count data is detected for each energy region by the detection circuit 93, and M memories 94 for energy use are detected.₁~ 94_MIs stored separately. The stored data is converted into a serial signal by the switch 95 and transmitted to the earth antenna 98 via the transmitter 97. Note that the CPU 99 controls the operation of each circuit of the processing circuit.
[0137]
The data sent to the earth antenna 98 is sent to the processing / display device 102 via the receiver 100 and the memory 101, and is subjected to predetermined processing for observing characteristic X-rays.
[0138]
With such a system configuration, the X-ray detector 12 can be directed in the direction in which observation is always desired, and characteristic X-rays emitted from the vicinity of the table can be simultaneously imaged and captured for each energy of interest. For this reason, it is possible to perform observations on the earth that capture the time-varying state of the platform in chronological order. That is, since the X-ray detector 12 can simultaneously capture X-ray images of a plurality of energies at a time, the X-ray detector 12 suitably functions as a two-dimensional sensor for X-ray astronomy that needs to capture instantaneous burst phenomena. I do.
[0139]
(Summary of effects)
As described above, the X-ray detector according to the present invention uses a CdTe semiconductor to configure a pixel of about 200 μm in a monolithic structure, and further includes a processing circuit including a charge amplifier without protruding to the back of the detector. It is provided as an ASIC. With this detector, X-rays are regarded as particles, and an ionized electric signal generated each time each X-ray particle is incident is detected by a charge amplifier. This detection signal is sent to a pulse shaping circuit, and a signal whose peak value is proportional to the energy is generated and input to a plurality of comparators. These comparators make comparisons with a plurality of energy thresholds (DC voltages), and determine whether the energy of the incident particle exceeds a predetermined energy for each energy region. The number of particles of energy entering each energy region is counted. Therefore, an output in each energy region can be obtained as a digital value without providing an A / D converter.
[0140]
According to such a method of measuring the number of pulses, since the energy threshold is set to a level higher than the dark current, an image having an extremely high S / N compared to the integral type can be obtained. In addition, since the size of each pixel is as small as 200 μm, the number of X-ray particles incident on each pixel is much lower than 1 (Kcount / sec) at which the counting characteristic appears, so that measurement can be performed without counting.
[0141]
Normally, the pixel size required to obtain a CT image is about 1 mm. Therefore, in order to use collected pixels as a detection unit, for example, 25 detector outputs having a pixel size of 200 μm are bundled (by adding). The output can be equivalently obtained without being counted down.
[0142]
Further, when the original pixel is 200 μm, by outputting the output without bundling it as it is, a plane image of X-ray can be obtained. This means that a medical X-ray CT scanner represented by IVR-CT and a normal X-ray diagnostic apparatus can be performed by a single CT scanner, and a large reduction in installation space and apparatus cost is realized. be able to.
[0143]
Similarly, in a non-destructive inspection device for baggage inspection, a procedure is first performed in which a suspected baggage is first detected with a normal X-ray transmission image, and this is precisely determined by a CT scanner. X-ray fluoroscopic images can be obtained from the beginning with the CT scanner, which is extremely advantageous in terms of cost, installation space, throughput, and the like.
[0144]
Also, since only X-ray particles exceeding a certain energy are measured, the energy of X-rays scattered by the human body and entering the detector is smaller than that of X-rays emitted from the X-ray tube. Can be obtained.
[0145]
By providing a plurality of energy thresholds, by editing reconstruction data after scanning, for example, a high-energy image, a low-energy image, or an image having only a certain range of energy can be arbitrarily and selectively reconstructed. You can also.
[0146]
Since the X-ray detector uses a CdTe semiconductor, the X-ray detector is excellent in detection sensitivity as compared with a current detector system including a GOS or CsI photodiode. For this reason, even if the X-ray intensity of the X-ray tube is reduced, it is possible to obtain an image having the same sensitivity as before, and to suppress the exposure of the subject.
[0147]
X-ray CT scanners used in airport baggage inspections, etc. expose the photo film in baggage, so it is forbidden to put undeveloped film in baggage. If the amount of X-rays can be reduced, the risk of exposure can be reduced.
[0148]
An X-ray detector using a CdTe semiconductor directly converts radiation into an electric signal and outputs a digital signal as a count value, so that a photodiode and an A / D converter, which exist conventionally, become unnecessary, and The circuit configuration at the subsequent stage is simplified. In addition, since the X-ray detector employs a monolithic structure, the formation of electrodes, ie, the formation of pixels, requires less man-hours than the formation of pixels using a scintillator. Therefore, the total manufacturing cost can be significantly reduced.
[0149]
When a system is constructed based on the pulse count, the relative relationship between the variation in gain and offset of the pulse (proportional to the energy value) from the waveform shaping circuit for each collected pixel and the energy threshold (DC voltage) is extremely important. If this is not set with high accuracy, it becomes a major cause of CT image artifacts, and practical application becomes difficult. Therefore, in the present invention, a sealed gamma ray source close to the energy range of an actual X-ray source is used.²⁴¹Am (59.5 KeV)¹²²Using Co (122 KeV) as a reference, the gain and offset of the circuit system of each pixel are measured at the stage of the output adjustment of the detector, and the setting of the energy threshold with high accuracy corresponding thereto is programmed and dealt with. By executing this program as calibration at an appropriate timing such as at the time of installation of the apparatus, it is possible to reliably suppress such variations and to secure a high-quality CT image in which artifacts are suppressed.
[0150]
Note that the radiation detector and the radiation imaging system according to the present invention are not limited to the configurations described in the above-described embodiments and the configurations of the modifications, and those skilled in the art may describe the gist described in the claims. It can be appropriately changed and modified without departing from the scope. For example, in the X-ray detector, in a monolithic structure in which the position and the size of the collection pixel are determined by the collection electrode as described above, the collection pixels are two-dimensionally arranged (see FIG. 2) as in each of the above-described embodiments. For example, as shown in FIGS. 22A and 22B, detection of a three-dimensional array obtained by stacking detectors of a one-dimensional array or detectors of a two-dimensional array as shown in FIGS. It may be a vessel. Further, as shown in FIG. 3C, even if a single pixel (incident window) having a desired size is used, it responds to each incident when a plurality of radiation particles continuously enter the incident window. If the occurrence of the superposition phenomenon between the pulse signals described above can be substantially ignored, it can be used.
[0151]
【The invention's effect】
As described above, the radiation detection method, the radiation detector, and the radiation imaging system according to the present invention can be applied to a wide range of fields from medical use to industrial use, and greatly reduce the influence of scattered radiation. It has a very high detection sensitivity by reducing the effect of dark current and ensuring an extremely high S / N by reducing the influence of dark current (the amount of radiation exposure is also reduced) and is detected by the sensor unit. The circuit scale of the circuit that processes the generated signal can be greatly reduced as compared with the X-ray detector that operates in the integration mode, and the size can be reduced.
[0152]
In addition, a single apparatus can detect various types of radiation, such as performing detection in accordance with the energy of the radiation.
[Brief description of the drawings]
FIG. 1 is an explanatory view schematically showing an embodiment of an X-ray CT scanner as a radiation imaging system of the present invention.
FIG. 2 is a perspective view schematically showing an X-ray detector mounted on the X-ray CT scanner.
FIG. 3 is a graph of an input count number versus an output count number for explaining the counting characteristic.
FIG. 4 is a block diagram showing an electrical schematic configuration of an X-ray CT scanner centering on an X-ray detector.
FIG. 5 is a diagram illustrating a processing circuit for X-ray detection formed in units of collection pixels.
FIG. 6 is a timing chart for explaining from energy discrimination of incident X-ray particles to counting operation thereof.
FIG. 7 is a graph showing a relationship between an energy spectrum of X-rays emitted from a normal X-ray tube and a threshold.
FIG. 8 is a block diagram illustrating calibration performed for each collection pixel.
FIG. 9 is a schematic flowchart illustrating an example of a procedure of a calibration process.
FIG. 10 is a view for explaining calibration reference setting.
FIG. 11 is a diagram illustrating a typical path of X-rays emitted from an X-ray tube.
FIG. 12 is a view for explaining threshold setting for angiography.
FIG. 13 is a partial block diagram of an X-ray diagnostic apparatus that performs angiography as a second embodiment of the radiation imaging system according to the present invention.
FIG. 14 is a partial block diagram of an X-ray CT scanner that performs IVR-CT as a third embodiment of the radiation imaging system according to the present invention.
FIG. 15 is a diagram illustrating a bundle of a plurality of collected pixels.
FIG. 16 is a partial block diagram of an X-ray CT scanner capable of performing a mass screening of lung cancer according to a fourth embodiment of the radiation imaging system according to the present invention.
FIG. 17 is a diagram illustrating a relationship between an X-ray spectrum for mammography executed by an X-ray imaging apparatus as a fifth embodiment of the radiation imaging system according to the present invention and threshold setting.
FIG. 18 is a diagram illustrating a state of mammography using a compression plate.
FIG. 19 is a schematic block diagram showing a baggage inspection system as a sixth embodiment of the radiation imaging system according to the present invention.
FIG. 20 is a schematic block diagram showing a pipeline inspection system as a seventh embodiment of the radiation imaging system according to the present invention.
FIG. 21 is a schematic block diagram showing a two-dimensional image system for X-ray astronomy as an eighth embodiment of the radiation imaging system according to the present invention.
FIG. 22 is a view for explaining another example of the collection pixel array of the X-ray detector.
[Explanation of symbols]
1 X-ray CT scanner (radiation imaging system)
1A X-ray diagnostic equipment (radiation imaging system)
11 X-ray tube (radiation source)
5 Console
12 X-ray detector (radiation detector)
12J detector block
20 Controller
22 Storage unit
23 Reconstruction unit
31 Pixel Data Operation Circuit
32 pixel bundling circuit
33 Switching circuit
41 Charge Amplifier
42 Waveform shaping circuit
43 Comparator
44 Energy region distribution circuit
45 counter
46 counter
47 counter
49 Bundling Network
50 pixel bundling circuit
51 Adjuster
52 memory
70 Baggage Inspection System
79 Pipeline Inspection System
90 Two-dimensional imaging system for X-ray astronomy
S semiconductor cell
AS ASIC layer
E1 charging electrode
E2 current collecting electrode
C processing circuit

Claims (28)

In a radiation detector that detects the number of particles of the radiation by letting the radiation be regarded as particles,
An incident window having a size that can substantially ignore the occurrence of a superposition phenomenon between pulse signals in response to each incident when a plurality of the radiation particles are incident continuously at or near the same position, and A sensor unit for generating an electric signal in response to each of the radiation particles incident on the sensor, and an energy for sensing an electric signal generated by the sensor unit and detecting an energy value of each of the radiation particles incident on the incident window A radiation detector, comprising: a detection unit. The radiation detector according to claim 1,
The said sensor part is a sensor part which injects radiation which has energy of a continuous spectrum as the said radiation, The radiation detector characterized by the above-mentioned. The radiation detector according to claim 2,
Energy discriminating means for discriminating the energy value detected by the energy value detecting means into a desired energy region; and counting the number of radiation particles incident on the energy region discriminated by the energy discriminating device to represent the count value. A radiation counting device that outputs a signal as the detection signal. The radiation detector according to any one of claims 1 to 3,
The radiation detector according to claim 1, wherein the incident window is a single-area window formed in a size that can substantially ignore the occurrence of the superimposition phenomenon. The radiation detector according to any one of claims 1 to 3,
The radiation detector according to claim 1, wherein the incident window includes a plurality of sets of pixels each having a size that can substantially ignore the occurrence of the superimposition phenomenon. The radiation detector according to claim 5,
The plurality of pixels each have a pixel size of 200 μm × 200 μm or less, or a pixel size capable of detecting the radiation particles with an accuracy of 0.1% or less. The radiation detector according to claim 5 or 6,
The radiation detector according to claim 1, wherein the incident window is formed by arranging the set of pixels in a one-dimensional, two-dimensional, or three-dimensional array. The radiation detector according to any one of claims 3 to 7,
The energy discriminating unit includes a threshold setting unit that sets a plurality of thresholds, and a comparing unit that compares an energy value detected by the energy value detecting unit with each of the plurality of thresholds,
As a result of the comparison by the comparing means, the number-of-particles counting means counts the number of incident radiation particles exhibiting an energy value larger than each of the plurality of thresholds for each of the thresholds, and uses the count value as the plurality of detection signals. A radiation detector, which is configured to output the radiation. The radiation detector according to any one of claims 3 to 7,
The energy discriminating means includes threshold setting means for setting one or more thresholds, and comparing means for comparing the energy value detected by the energy value detecting means with each of the one or more thresholds,
The particle number counting means includes means for counting the number of incident radiation particles exhibiting an energy value belonging to a desired energy region determined by the one or more threshold values as a result of the comparison by the comparison means. Detector. The radiation detector according to any one of claims 1 to 9,
A radiation detector comprising: a calibration unit configured to calibrate a relative magnitude relationship between an energy value detected by the energy value detection unit and the predetermined value based on a reference energy value. The radiation detector according to claim 10,
The radiation detector according to claim 1, wherein the calibration unit performs the calibration using an energy value of a gamma ray source as the reference energy value. The radiation detector according to claim 11,
The gamma ray source comprises at least two types of radiation sources having different energy values. The radiation detector according to claim 11,
The gamma ray source having at least two kinds of energy values includes ²⁴¹ Am having an energy value of 59.5 KeV and ⁵⁷ Co having an energy value of 122 KeV. The radiation detector according to any one of claims 1 to 13,
The radiation detector according to claim 1, wherein the sensor unit includes a semiconductor layer that directly converts the radiation into an electric signal. The radiation detector according to claim 14,
The radiation detector, wherein the semiconductor layer is made of a compound semiconductor. The radiation detector according to claim 15,
The radiation detector according to claim 1, wherein the semiconductor layer is made of a cadmium telluride semiconductor (CdTe semiconductor). The radiation detector according to claim 15,
The radiation detector, wherein the semiconductor layer is made of a cadmium zinc telluride semiconductor (CdZnTe semiconductor). The radiation detector according to claim 14,
The radiation detector, wherein the semiconductor layer is made of a silicon semiconductor (Si semiconductor). The radiation detector according to claim 14,
The sensor unit has a monolithic structure in which a radiation incident surface of the semiconductor layer is covered with a charging electrode for applying a voltage, and a surface opposite to the radiation incident surface of the semiconductor layer is covered with a current collecting electrode. Radiation detector. The radiation detector according to claim 19,
The radiation detector has a structure in which the current collecting electrode is divided into a plurality of pixels for detecting the radiation on a particle-by-particle basis on the back surface. The radiation detector according to claim 19,
The sensor unit includes a processing circuit including a charge amplifier for detecting an electric signal generated in the semiconductor layer in response to the incidence of the radiation particles, and a processing circuit including a charge amplifier integrally disposed on the surface of the current collecting electrode of the semiconductor layer in an ASIC. A radiation detector characterized in that: The radiation detector according to claim 21,
The sensor unit is configured such that the ASIC is disposed integrally in the plane of the current collecting electrode, and the integrated structure of the semiconductor layer and the ASIC is formed as a block. A radiation detector having a structure capable of coupling in a plane direction. The radiation detector according to any one of claims 1 to 22,
The radiation detector is an X-ray. A radiation detection unit that detects radiation using the radiation detector according to any one of claims 1 to 23, and an image generation unit that generates an image using a signal of the radiation detected by the radiation detection unit. A radiation imaging system comprising: The radiation imaging system according to claim 24,
A radiation imaging system, comprising: signal bundling means for adding the detection signals detected from the pixels arranged in an array of the sensor unit to each other and bundling them into a single detection signal. The radiation imaging system according to claim 25,
The signal bundling unit is a unit that adds together the detection signals detected from adjacent pixels among the pixels arranged in the array and bundles them into a single detection signal. system. The radiation imaging system according to claim 26,
The image generating means includes: calculating means for processing the detection signal output from the particle counting means for calculation of a predetermined process; and means for generating the image from a signal obtained as a result of the calculation by the calculating means. A radiation imaging system, characterized in that: In a radiation detection method for detecting the number of particles of the radiation by letting radiation having energy of continuous spectrum be incident as particles,
An incident window having a size that can substantially ignore the occurrence of a superposition phenomenon between pulse signals in response to each incident when a plurality of the radiation particles are incident continuously at or near the same position, and A sensor unit for generating an electric signal in response to each of the radiation particles incident on the incident window; detecting an electric signal generated by the sensor unit to detect an energy value of each of the radiation particles incident on the incident window; And detecting the detected energy value, counting the number of incident radiation particles exhibiting the determined energy value, and outputting the detection signal representing the count value.

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