JP3569526B2 - X-ray detector for low dose scanning beam type digital X-ray imaging system - Google Patents

Description

AP_X×AP_YIMAGEデータ列を有用な画像に結合させる第２の現在好ましい方法は、マルチアウトプット旋回法と呼ばれる。この場合、DET_X×DET_Yセンサのセンサアレイとともに、DET_X×DET_Yのデジタイザ（または同等なものや、多重のもの）および同数の画素合算回路が有る。各センサからデジタル化された値は、SENSOR（j,i）と呼ばれる。最後のOUTIMAGE列は、以下のように計算される。出力イメージ列OUTIMAGE（y,x）［ｙ＝１〜AP_Y、ｘ＝１〜AP_X］の各画素について、DET_X×DET_YソースイメージSENSOR（j,i）の各々からの１つの画素は、合計されて［ｊ＝１〜DET_Y、ｉ＝１〜DET_X］、目的のイメージ画素OUTIMAGE（ｙ−ｊ＊FOCUS,x−ｉ＊FOCUS）となる。それから、各素子がDET_X＊DET_Yで割られることにより、OUTIMAGE列の規格化が行なわれる。
これらの技術のさらなる改良が、FOCUS要素の分数部分にもとずく線形補間を行うことにより得ることができる。
マルチイメージ旋回法のマルチアウトプット旋回法に対する利点は、前者では、データを取った後にソフトウエアで最適焦点の面を選択できるのに対し、後者ではできないことがある。しかし、後者は時間が限界となる場合に、より素早い操作が可能である。
SDBXデータからの３次元像の再構築
ここで述べられたSBDXシステムは、物体80の断層写真３次元像を形成するために用いられる１組の連続的平面像を作るのに用いることができる。種々のFOCUS値でセットされたデータを再分析することによる、種々の深さの一連の像からなる３次元で像を作るために、１組の像は解析できる。用いられる、自然のFOCUS値は、n/DET_Xまたはn/DET_Yであり、ここに、ｎはそれぞれ０からDET_XまたはDET_Yまでの整数である。通常、焦点の値は、物体80内の関心のある平面に一致するように分析される。例えば、下記の表１に現されたSBDXシステムにおいて、焦点の平面は、22.86cm（９インチ）（最適焦点面）の普通の焦点面の近くで、約2.54cm（１インチ）の間隔で置かれる。
次の数式は、一連の平面像が、アノード50からの距離に換算して位置されることを示す。

サブ（sub）−サンプリング技術を用いる時も、計算は変わらず、「はねない（not skipped）」コリメーショングリッドの開口からのデータが扱われる。しかし、たとえ間にはいるコリメーショングリッドの穴がなくても、λ_ｔは同じである。
ネガティブフィードバックｘ線束制御
図13に示されるBDX撮影システムは、ｘ線ビーム100のｘ線束を制御するネガティブフィードバック経路305を用いる。好ましくは、センサアレイからのネガティブフィードバックは、センサアレイが常にほぼ同じｘ線束レベルを見るようにｘ線束を制御する。このように、（ｘ線に対し比較的透明な）軟らかい組織がスキャンされるとき、ｘ線束が落ち、患者（対象物）に対する全体の照射量を低下する。ネガティブフィードバック線束制御を用いることにより、コントラストとダイナミックレンジが改善される。本実施形態により、差動増幅器310は、ユーザにより設定可能な、調節可能参照レベル320を有する。ネガティブフィードバックループ305は、ｘ線管に対しｘ線束のネガティブフィードバックをする。
時間ドメインスキャンモード
また、時間ドメイン撮影システムは、また、ここに説明される原理を用いて、実行できる。そのようなシステムにおいて、種々の画素からあらかじめ測定されたｘ線束に達する時間は、計算でき、マップにできる。次に、ネガティブフィードバック制御は、取り扱うスキャン期間のためのあらかじめ決定した線束レベルに達した画素に対応する開口からｘ線束を除きまたは減少するために使用できる。この場合、集められた情報は、線束レベルへの時間であり、マップされ撮影された情報は、強度よりはむしろ時間に対応する。そのようなシステムの能力は、ずっと大きな信号雑音比を与え、コントラストを改善し、検査中の目的物へのｘ線照射量を劇的に減少し、ダイナミックレンジを改善することを可能にする。
多重エネルギーｘ線映像モード
本発明の１実施形態によれば、２以上のグループのｘ線ビーム100が、１以上の検出器アレイの方に進む。第１グループのｘ線ビームは、第１の特性ｘ線エネルギースペクトルを有する。第２グループのｘ線ビームは、異なった第２の特性ｘ線エネルギースペクトルを有する。第１グループと第２グループのｘ線ビームの測定された透過率を比較することにより、検査中の対象物における物体の存在が検出できる。微分ｘ線撮影の基本的概念は、従来より知られていて、米国特許第5,185,773号（「金属の非破壊的選択的決定のための方法と装置」）に開示され、この引用により本明細書に組み込まれる。
この２グループのｘ線は、多数の方法で発生できる。１つの方法では、第１グループの開口の近くの材料の第１材料または第１厚さ、および、第２グループの開口の近くの材料の第２材料または第２厚さを備える特別なアノードの製造をおこなう。このように、第１グループに関連する開口は、第１特性エネルギースペクトルを有するｘ線を放出し、第２グループに関連する開口は、第２特性エネルギースペクトルを有するｘ線を放出する。他の方法では、Ｋ−フィルタリング（Ｋ端フィルタリング）は、同様な効果を生じる開口140の一部の中に、フィルタ材料（モリブデンなど）を配置することにより使用できる。この場合、開口の第１グループは、その中に挿入された第１フィルタを備え、開口の第２グループは、その中に挿入された第２フィルタを備える。第２フィルタは、フィルタでなくてもよい。前の場合におけるような、異なった特性エネルギースペクトルを有する２グループのｘ線は、２グループの開口と関連する。
少なくとも２グループの開口が異なった特性エネルギースペクトルと関連づけられると、広帯域ｘ線で通常見ることができない微細石灰化（乳癌）や他の異常を検知することが可能になる。たとえば、第１像を形成する第１グループの開口のスキャンを行ない、第２像を形成する第２グループの開口のスキャンを行ない、次に、像の除算によりそれらの比を強調することにより、微小石灰化や他の異常を低照射量スキャンｘ線撮影システムを用いてリアルタイムに検出できる。同様に、多重検出器アレイ装置は、第１検出器アレイの方に向けた第１グループの開口と第２検出器アレイの方に向けた第２グループの開口と共に使用できる。
多重エネルギー撮影の他の実施形態が次に説明される。検出されたｘ線光子からの電気パルスの大きさが、光子のエネルギー（kV）に比例するので、２以上のエネルギーバンドにおける光子からくるパルスを別々に計数することが可能である。パルスは、強度により分離され、次に計数され、別々に処理され、２以上の別々の像を作る。これらの像は、比として表示できる。
また、対象物における異なった密度領域を区別するため、フライ（fly）における選択されたエネルギーレベルを変えることが可能である。本実施形態の効果は、上述の実施形態よりも柔軟であることであり、特別なコリメーショングリッド、アノード材料または２重検出器を必要としないことである。
多数の好ましい実施形態が本発明の種々の構成についてさきに説明されたが、以下の説明は、本発明による好ましいSBDX撮影システムを説明するものである。

したがって、セグメントに分割された検出器アレイを利用するこのSBDX撮影システムは、高分解能で高感度であると同時に、検査中の対象物への照射線量が低い。また、このシステムは、線源50と検出器アレイ110との間の任意の点に最適焦点を設定することができ、視野の有効作用深さを与えることができる。
ビームサブスキャン技法
以下の説明は、本発明の特に好ましい実施形態に関連し、この実施形態は、計算機の処理のオーバーヘッドを減少するために、ビームのサブサンプリングの技法を用いる。
標準のビデオイメージは、640×480の画素を用い、30Hzで更新される。これは、約12MHzの画素サンプル速度を必要とする。この速度でｘ線管の高電圧電子ビームを正確に250,000個の連続的な異なる開口の背後に位置することは、高精度と比較的大きな電力消費を必要とする。12MHzの速度でのｘ線検出器の大きなアレイからの信号のデジタル化は、同様に、効果的であり、電力を消費する。したがって、SBDXの空間的分解能または時間的分解能の大きな減少なしに、画素サンプル速度の12MHzより下への減少は、初期のユニットのコスト、電力消費による操業コスト、および、ｘ線管によるむだな熱のための冷却要求を減少するのに有用である。
したがって、画素サンプル速度を減少しつつ実質的に同じ空間的分解能と時間的分解能を与える機構が発展された。この機構は、サブサンプリングといわれ、SBDXの他の構成でも明らかに適用できるが、ここに説明するSBDXの実施形態において最もよく実行される。本実施形態の効果は、電力消費の減少、ｘ線管内での電子ビームの偏向のより簡単な回路、平行化グリット90の作成コストの低下、対象物80の像を分解するのに必要な計算の複雑さの減少、および、当業者にとって明らかなその他の効果である。
本実施形態において、コリメーショングリッド90は、500個よりは少なくした数の開口、好ましくはAP_x＝AP_y＝166（他の数も明らかに用いることができるが）を備えるように作成される。計算の観点からのこの減少の効果は、以下で明らかになる。しかし、製造の観点からの効果は、製造に必要な開口の数の約1/9の、ずっと簡単な構造である。開口の数が減少するので、より高い偏向角（すなわち、開口がコリメーショングリッドの前面260に対してなす角）でグリッドを製造することは、近接する開口と交差する開口を有するという問題を生じることなく、より容易である。このことは、立体グリッドが製造されるとき、立体グリットにおける近接する開口が異なった検出器アレイの方に向き、したがって、開口交差を防ぐため非立体グリッドよりも広い物理的分離を必要とするので、特に有用である。
コリメーショングリッドの開口（複数）は、最大寸法の円の中にAP_x行とAP_y列の円で配置される。計算の目的のために、これは、情報に寄与しない円の外側にある、すなわち、常に「暗い」すなわちｘ線により照射されない素子（複数）を備えたAP_x行とAP_y列の寸法の矩形として取り扱える。
ｘ線検出器アレイ110のセンサ160は、図15に示すように、最大寸法でDET_x行とDET_y列の円状アレイの中に配置される。画素のサンプル速度は、コリメーショングリッドの全部より少ない開口の照射、すなわち、サブサンプリングにより低下できる。好ましくは、照射されない開口がないコリメーショングリッドが使用される。この検出器アレイを用いて像を作るために、各行におけるてすべてのDET_x番目のコリメータの穴と各行におけるDET_y番目のコリメータの穴とのみが照射される必要があり、したがって、像は、各々DET_x画素とDET_y画素の寸法の、画素の像タイルから組み立てできる。これは、DET_x×DET_yのサブサンプリング比に対応するが、サブサンプリングは、１×１のサンプリング速度には対応しない。したがって、このサブサンプリング比は、ｘ方向（行）に１からDET_xまで、ｙ方向（列）に１からDET_yまで調節できる。この好ましい実施形態によれば、図15に示すように、DET_x＝DET_y＝12である。
12×12の検出器が用いられ、サブサンプリング比は12であるとき、像は、デイビッド・ホックニー（David Hockney）の光モザイクにずっと似て、実質的に共に貼り付けられる複数の重複しない像から作成される。実際のシンチレータと検出器とは、全く完全というわけでなく、また、全く同一に応答するというわけでもないので、ｘ線鉛筆ビームは、完全には一様ではなく、コリメーショングリッドの開口は、理想的な面積と全く正確に同一というわけではない。そして、正方形の検出器でなく円の検出器が使用されるので、ある程度の重なりは、検出器の非線形性と雑音とを平均することを可能にするために、非常に好ましい。
もしサブサンプリング比が、画素中の検出器の寸法より小さければ（すなわち、この好ましい実施形態では12より小さければ）、像は、重複した「タイル」から組み立てられ、これらは合計されまたは平均されねばならない。もしサブサンプリング比（複数）が（画素中の）検出器の寸法の複数倍でないとき、または、もし検出器のアレイが直方体でないとき、異なった数のサンプルが各画素に追加され、異なった除算器が各画素を平均化するために必要である。理想の環境より少ない数を扱う技法は当業者によく知られていて、説明を複雑にしすぎないためにここでは説明しない。
以下の計算では、SSXは、Ｘ方向（行）でのサブサンプリング寸法を表し、SSYは、Ｙ方向（列）でのサブサンプリング寸法を表す。たとえば、もしSSX＝SSY＝１であるならば、サブサンプリングはなく、処理は、上述の本発明の他の実施形態におけるのと全く同じである。同様に、もし本実施形態におけるようにSSX＝SSY＝１であるならば、画素平均を用いない「光モザイク」に戻る。もしSSXとSSYが３つあり、円のアクティブ領域が500×500であるならば、166×166の開口がスキャンされ、ｘ方向に1/3、ｙ方向に1/3であり、得られるデータの数が因子９だけ減少する。もし1/9の開口を全時間使用するならば、これらの計算の必要はなく、コリメーショングリッドが含まれる必要がない。
したがって、像を作成するために、元のコリメーショングリッドにおける1/（SSX＊SSY）のみの開口（500×500の開口）が使用される必要があり、すなわち、ｘ線照射のための電子ビームにより照射される必要がある。もしフレーム速度が一定に、すなわち、30Hzに保たれるならば、電子ビームを駆動する回路の周波数応答のように、電子ビームの運動の数は、SSX＊SSYだけ減少される。電子ビームが動く全体の距離（およびスキャン線の数）は、1/SSYだけ減少され、このため、ターゲットアノードを通る平均ビーム速度は、1/SSYだけ減少される。像を再構成する画素速度は、コリメーション開口速度（開口がスキャンまたは照射される速度）と同じであり、同様に、1/（SSX＊SSY）により減少される。
この方式によれば、各表示画素へと平均化されるサンプルの数は、（DDT_x/SSX）＊（DDT_y/SSY）である。最大のサンプリング速度を用いると（SSX＝DDT_xおよびSSY＝DDT_y）、１つのみのデジタイザサンプルが各表示画素に平均される（「光モザイク」モード）。サンプルの平均化は、ビーム、シンチレータ、検出器および増幅器における非一様性を滑らかにするために重要である。サブサンプリングの大きさ（SSXとSSY）は、許容可能な像の質を確保するために提示される条件のために、適当なレベルに設定されねばならない。これは、像の質と特定の１組の環境により提出される条件とについてのユーザの好みによりフライについてユーザにより設定できる。
図15に示す検出器アレイ110は、好ましくは、約１インチの直径を有する円状面積においてほぼ配置される96個の個々の検出器素子160のアレイである。アレイの中心で、縦の列に12個の検出器（DDT_x）があり、アレイの中心で、水平の行の列に12個の検出器（DDT_y）がある。シンチレータ結晶は、好ましくは、４角の水平断面に切断され、0.005インチの厚さのステンレス鋼のストリップからなる「玉子クレート」構造により支持される。全シンチレータ結晶（ハッチされている）が中に位置された図15の円400は、好ましくは、約0.800インチの直径である。
検出器アレイ110におけるシンチレーション結晶の長さは、好ましくは、0.10cmであり、その前の入力面は、好ましくは、0.135cm×0.135cmである。シンチレーション結晶は、好ましくは、YSO、LSOまたはBGOであるが、上述のように、他の材料も使用できる。BGOは、この用途におけるその光出力のための適当に減少したデケイ時間（50nS）のために、約100℃に加熱されねばならない。したがって、抵抗加熱素子が備えていてもよい。
図17は、本発明の好ましい実施形態による検出器アッセンブリ402を示す。ｘ線は、上から、ｘ線窓404を通り、鉛シールド406の中に入る。散乱ｘ線を減衰しつつｘ線がコリメーショングリッド90の開口から出て検出器アレイ110に当たることを可能にするために、ｘ線窓404は、好ましくは、円であり、直径が1.91cmである。光シールド408は、検出器を迷った光から遮断するために備えられる。光シールド408は、実質的にｘ線を減衰することなく光を減衰するように選択されたアルミニウムまたはベリリウムの薄板から製造できる。
検出器アレイ110は、BGOシンチレータとともに使用するため、適当な加熱素子410の近くに位置される。加熱素子410は、100℃の作動温度で検出器アレイ110を維持するように設計された抵抗加熱素子であってもよい。ファイバ光学的撮影テーパー412を検出器アレイ110の底414から現れる光子を、96チャンネルの光電子倍増管（PMT）416に向ける。検出器アッセンブリ402は、迷う光が雑音を発生するのを防止するために、光を漏らさない外側ハウジング418の中に囲まれている。３個の肩ねじ420と３個のセンタリングねじ422は、当業者に周知であるように、面状と線状の配列のために備えられる。回転配列は、外側ハウジング418をPMT取付台426に対して回転することにより達成される。ファイバ光学的撮影テーパー412は、米国カリフォルニア州キャンベルのコリメーテッド・ホールズ社から市販されていて、直径2.03cmの円状入力開口と、直径3.38cmの円状出力開口を備える。テーパー412は、PMT416の寸法（0.10インチ）に対し、各シンチレータ結晶のピッチ寸法（0.06インチ）を適合する。すなわち、テーパーは、1.667倍の倍率を有する。ダウコーニング社から市販されている高粘性の光結合液体（型200）は、ガラスの屈折率にほぼ等しい屈折率を有し、シンチレータ結晶160からテーパー412へ、テーパー412からPMT入力面424への光移送効率を最大にするために、光結合媒体としてテーパーの２面で使用される。
光電子倍増管416は、フィリップス社からの型XP1724Aとして市販されている96チャンネルの管（１チャンネルが各シンチレーション結晶に対応する）である。この光電子倍増管416は、シンチレーションアレイの空間的配置が、光学面板の他方の面上でのPMTに位置されるPMT光カソードへ正確に実行されるように、ファイバ光学面板を備える。１つのシンチレーション160に当たるｘ線光子は、PMT光カソードに結合される光パルスを生じる。これは、光カソード上で、対応する電子パルスを生じ、このパルスは、PMTダイオード構造の１チャンネルで、1,000,000倍まで増幅される。
このPMT出力パルスは、30MHz帯域の増幅器の入力に結合され、増幅器の出力は、パルスが微分されるように、0.5〜5.0ボルトの範囲にあり、約30nsecの長さである。これにより、パルス速度が変わるにつれベースライン参照電圧を保つDC復元（restorer）回路の必要性をなくす。
増幅器の出力は、比較器に入力され、比較器は、その入力の大きさにかかわらず、一定の大きさの出力パルスを出力する。比較器のための参照電圧は、増幅器が雑音出力レベルでトリガーされないように、雑音出力レベルよりわずかに大きな値に設定される。増幅器のチェーンは、検出器アレイにおける各シンチレーション結晶について一度で、96回反復される。比較器の出力パルスは、データ収集・像再構成システムについての新しいデータとなる。試験が示したように、プロトタイプシステムは、約10MHzの速度までｘ線光子をランダムに計数できる。
本発明の実施形態と用途が以上に説明されたが、当業者には、上に説明したよりも多くの変形が、本発明の概念から離れることなく可能であることが明らかである。したがって、本発明は、添付した請求の範囲の精神の中を除いて、制限されない。 1.Technical field of invention
The present invention relates to a diagnostic X-ray imaging apparatus. More particularly, the present invention relates to a real-time scanning beam digital x-ray system with improved resolution and reduced x-ray emission by incorporating multiple apertured collimation grids and a segmented x-ray detector array. .
2. Background technology
With the development of therapeutic techniques, real-time radiography in medical procedures is increasingly desired. For example, many electrophysiological, peripheral vascular, urological, and orthopedic procedures in cardiology rely on real-time radiography.
Unfortunately, current clinical real-time x-ray equipment emits high levels of x-rays to both patients and participating staff. The U.S. Food and Drug Administration (FDA) reports anecdotal evidence of acute radiation sickness in patients and problems with occupationally overexposed physicians (Radiological Health Bulletin, Vol. 26, No. 8, August 1992) .
Conventionally, many real-time X-ray imaging systems are known. These are equipped with a system based on X-ray fluoroscopy, in which X-rays are directed at the subject and the shadow created by the X-ray opaque body in the object is opposite to the object by the X-ray source. Is displayed on the X-ray fluoroscope located on the side. Scanning X-ray tubes have been known in connection with fluoroscopy techniques since at least the early 1950s. See Moon, "Amplification and enhancement of fluoroscopy by scanning x-ray tubes", Science, October 6, 1950, pages 389-395.
Scanning beam digital radiography systems are also well known in the art. In this system, X-rays are generated using an X-ray tube. An X-ray is generated from this spot by generating an electron beam inside the X-ray tube and condensing it on a small spot on a relatively large anode (transmission target) of the tube. The electron beam is deflected (electromagnetically or electrostatically) into a raster scan pattern across the anode. A small X-ray detector is disposed at a predetermined distance from the anode of the X-ray tube. The detector converts the impinging X-rays into an electrical signal proportional to the detected X-ray flux. When the object is placed between the X-ray tube and the detector, the X-rays are attenuated and dispersed by the object according to the X-ray density of the object. If the x-ray tube is in scan mode, the signal from the detector is modulated in proportion to the x-ray density of the object.
Examples of conventional scanning beam digital X-ray systems include Albert US Pat. No. 3,949,229, Albert US Pat. No. 4,032,787, Albert US Pat. No. 4,057,745, Albert US Pat. No. 4,144,457, Albert US Pat. U.S. Pat.No. 4,149,076, Albert U.S. Pat.No. 4,196,351, Albert U.S. Pat. No. 4,323,779, Albert U.S. Pat. No. 4,465,540, Albert U.S. Pat. No. 4,519,092, and Albert U.S. Pat. No. 4,730,350.
In a typical prior art scanning beam digital X-ray system, the output signal from the detector is applied to the Z-axis (luminance) input of a video monitor. This signal adjusts the brightness of the screen. The x and y inputs to the video monitor are derived from the same signal that deflects the x-ray signal of the x-ray tube. Therefore, the brightness of a point on the screen is inversely proportional to the respiration of the X-rays from the X-ray source through the object to the detector.
Medical X-ray systems are operated at the lowest level of X-ray dose that meets the resolution requirements appropriate for the procedure. Therefore, radiation dose and resolution are limited by the signal-to-noise ratio.
As used herein, the term "low radiation dose" means an X-ray exposure level of about 2.0 R / min or less that reaches the patient during operation.
The time and region distributions of X-ray photons follow the Poisson distribution and have unavoidable randomness, respectively. Randomness is expressed as the standard deviation of the average flux and is equal to its square root. The signal-to-noise ratio of the X-ray image under these conditions is therefore equal to the average radiation flux divided by the square root of the average radiation flux, and for an average radiation flux of 100 photons, the noise is ± 10 photons and the signal The noise-to-noise ratio is 10.
Therefore, the spatial resolution and the signal-to-noise ratio of the X-ray image formed by the scanning X-ray imaging system are significantly dependent on the size of the sensitive area of the detector. As the aperture area of the detector is increased, more divergent light is detected, the effective sensitivity is increased, and the signal-to-noise ratio is improved. However, at the same time, as the "pixel" size (measured in the plane of the object being imaged) increases, the larger detector aperture reduces the adjustable spatial resolution. This is necessary because most objects that are imaged in the medical field (eg, the internal structure of the human body) are some distance from the X-ray source. Therefore, in the prior art, the aperture size of the detector had to be selected to balance resolution and sensitivity, and it was not possible to maximize both resolution and sensitivity at the same time.
In medical imaging technology, the irradiation amount to the patient, the frame rate (the number of times per second for scanning the object and updating the image), and the resolution of the image of the object are key parameters. It is. High x-ray flux readily provides high resolution and high frame rate, but results in unacceptably high x-ray exposure to patients and participating staff. Similarly, even if the image is not visible and the refresh rate is insufficient, the dose can be reduced if the refresh rate is allowed. A preferred medical imaging system should simultaneously achieve low dose, high resolution, and a sufficient update rate of at least about 15 images per second. Therefore, systems such as the scanning beam digital radiography systems described above have a relatively long irradiation time, and the same is always true for a real patient, but the X-ray dose received by the patient is kept to a minimum. Not suitable for diagnostic medical procedures that must be performed.
Accordingly, it is an object of the present invention to provide a scanning beam digital radiography system that can be used in medical diagnostic procedures in which patients can be safely received.
It is another object of the present invention to provide a scanning beam type digital X-ray imaging system which provides a high resolution image at a sufficient frame rate and minimizes the amount of X-ray irradiation on an inspection object. .
Still another object of the present invention is to provide a scanning beam type digital X-ray imaging system in which the resolution at an area away from the surface of the X-ray source is improved while the level of the X-ray flux is reduced.
The above and many other objects and advantages of the present invention will be apparent to those skilled in the art from a consideration of the drawings and the following description of the invention.
Disclosure of the invention
A scanning beam digital X-ray imaging system (SBDX) according to the present invention includes an X-ray tube having an electron beam source and an anode serving as a target. A circuit is provided which focuses the beam and directs or scans the beam in a predetermined pattern via the target anode. For example, the predetermined pattern may be a raster scan pattern, a curved or S-shaped pattern, a spiral pattern, a random pattern, a Gaussian distribution pattern centered at a predetermined point on the target anode, or a short work pattern. Other patterns that are effective for
A collimating element (preferably formed in a grid) may be interposed between the X-ray source and the object to be irradiated with X-rays. The collimating element comprises, for example, a circular metal plate having a diameter of about 25.4 centimeters (10 inches) and an aperture array of 500 rows and 500 columns in the center of the collimating element. The collimating element is preferably arranged immediately before the light emitting surface of the X-ray tube. Other shapes of collimating elements can also be used. In one preferred embodiment of the present invention, each of the apertures of the collimating element is configured to be directed toward (or point to) a detection point on a surface located at any distance from the collimating element. ing. The distance is selected such that the object to be irradiated with X-rays is located between the collimating element and the detection point. The function of the collimating element is to form thin beams of X-rays (arranged like pixels), all directed from one point on the anode of the X-ray tube towards the detector.
Array of detector elements (preferably DET_X× DET_YThe center of the divided detector array having a rectangular or square-like area array, and more preferably a relatively round array, is located at the detection point. The detector array preferably comprises a plurality of closely packed X-ray detectors. Such an array is designed, positioned and used in accordance with the present invention to provide high sensitivity without loss of resolution, and therefore has a resolution equal to or better than conventional X-ray equipment. However, an X-ray system having an irradiation amount at least one order of magnitude smaller than that of the conventional X-ray system can be realized. This feature of the invention has important implications in medicine and other fields. Irradiation to patients or participating medical staff during current procedures is reduced. The danger of assault allows for measures that are not currently possible.
The output of the detector array is the intensity value for each element of the detector array when the x-ray beam is emitted through an aperture in the grid. Since each aperture is located at a spatially different point with respect to the test object and the detector array, a different output is obtained from the detector array for each aperture through which the X-ray beam passes. The output of the detector array can be converted to an image in various ways. One method is to simply combine the array output, i.e., sum the intensity values of the array elements corresponding to each scanned aperture and further normalize. The output array can then be used to drive a video or other display. Even more preferred are the multiple image synthesis method and the multiple output synthesis method described below, which provide improved visible output.
The SBDX imaging system can also perform stereoscopic imaging using a collimation grid having two groups of apertures. In this case, the first group of apertures is directed to the first detection point where the split first detector is located, while the second group of apertures is the first detection point where the split second detector is located. 2 detection points. By forming the two images with the two split detectors and using conventional stereoscopic display methods, a stereoscopic image is created.
SBDX imaging systems can also produce enhanced imaging of materials that exhibit different X-ray transmission at different X-ray photon energies. Thus, for example, microcalcification, a precursor of breast cancer, can be imaged. By configuring the grid and / or anode to emit two or more groups of X-ray beams, each having a different X-ray energy spectrum, and directing each group to a detector array (multiple detector arrays can also be used) Since it is possible to image that the transmittance of the inspection object is different at various X-ray photon energies, it is possible to emphasize only a portion showing a different X-ray transmittance inside the inspection object. For example, if optimized for calcium detection, this imaging system is a powerful tool that can detect breast cancer and other abnormalities early.
Maximum sensitivity without sacrificing resolution obtained using small surface area detectors by utilizing a split array that blocks the entire collimated x-ray beam and image processing the array output. Can be provided. Non-split detectors of the same size as the split array have the same sensitivity but lower resolution.
In addition, subsampling techniques may be used to process data from the array detector, reducing system complexity, required processing speed, and energy consumption while providing substantially the same image quality.
The system described herein is used in conjunction with a "catheter with an X-ray sensitive optical sensor detection device" described in U.S. patent application Ser. No. 08 / 008,455 filed Jan. 25, 1993 (CAM-003). Can be used, the contents of which are incorporated herein by reference. This US patent application Ser. No. 08 / 008,455 is owned by the assignee of the present application.
[Brief description of the drawings]
FIG. 1 is a diagram showing the basic components of a low-dose scanning beam digital X-ray imaging system.
FIG. 2 shows the distribution of X-rays from an SBDX system without a collimation grid.
FIG. 3 is an enlarged view of the grid and anode of an X-ray tube for a low dose scanning beam digital X-ray imaging system.
3A, 3B, and 3C are partial cross-sectional views of a collimation grid effective for the device of the present invention.
FIG. 4 is a diagram of an X-ray tube for a low dose scanning beam digital X-ray imaging system.
FIG. 5 is a sectional view showing the structure of an X-ray tube for a low-dose scanning beam type digital X-ray imaging system.
FIG. 6 is a diagram of a stereoscopic scanning beam type digital X-ray imaging system.
FIG. 7A is a diagram of an apertured X-ray source that cooperates with an undivided simple detector.
FIG. 7B is a diagram of X-rays from one aperture of an apertured X-ray source cooperating with a split detector array.
FIG. 7C is an illustration of X-rays from multiple apertures of an apertured X source cooperating with a split detector array.
FIG. 7D is a diagram of X-rays from two apertures of an X-ray collimation grid cooperating with an inspection object and a split detector array.
FIG. 8 is a view of the exposed surface of a 5 × 5 detector array for a low dose scanning beam digital radiography system.
FIG. 9 is a diagram of a 5 × 5 detector array for a low dose scanning beam digital radiography system.
FIG. 9A is a diagram showing a scintillator element according to one preferred embodiment of the present invention.
FIG. 10 is a diagram of a detector element for a low dose scanning beam digital radiography system.
FIG. 11 shows an array of pencil-shaped detector elements of a non-planar detector array.
FIG. 12 is a diagram of a 3 × 3 detector array for a low dose scanning beam digital radiography system.
FIG. 13 is a diagram showing the basic elements of a low-dose scanning beam digital X-ray imaging system using negative feedback to control the X-ray flux.
FIG. 14 is a perspective view showing a grid sealing device.
FIG. 15 is a diagram of an arrangement of a 96-element detector array according to a preferred embodiment of the present invention.
FIG. 16 is a diagram showing the interaction between the collimation grid and the detector array.
FIG. 17 shows a preferred embodiment of the detector device.
Embodiment of the present invention
It will be understood by those skilled in the art that the following description of the present invention is by way of example and not by way of limitation. Other embodiments of the invention will be readily apparent to those skilled in the art.
Overview of the device
FIG. 1 shows a scanning beam digital X-ray apparatus according to a preferred embodiment of the present invention. The scanning X-ray tube 10 is used as an X-ray source. As is well known in the art, a power supply of approximately 100 KV to 120 KV is used to operate the X-ray tube 10. A 100 KV power supply provides an X-ray spectrum that extends to 100 KeV. Here, 100 KV X-ray refers to this spectrum. X-ray tube 10 includes a deflection coil 20 that is controlled by a scan generator 30, as is well known in the art. The electron beam 40 generated in the X-ray tube 10 scans from one side of the grounded anode 50 to the other in a predetermined pattern in the X-ray tube 10. For example, the predetermined pattern is a raster scan pattern, a meandering or S-shaped pattern, a spiral pattern, a random pattern, a Gaussian distribution pattern centered on the above predetermined point of the target, or a task in the near future Other patterns that are useful for Presently preferred are meandering or S-shaped patterns that eliminate the need for "retrace" in raster scans.
The electron beam 40 strikes the anode 50 at a point 60, a cascade of X-rays 70 is emitted, and travels outside the X-ray tube 10 toward an object 80 to be examined by X-rays. To make the best use of the operation of the device, conical X-ray photons are generated that radiate out just to cover the detector array 110. This is preferably achieved by placing a collimation grid between the anode of the scanning X-ray tube and the detector. Therefore, the collimation grid 90 is disposed between the object 80 and the X-ray tube 10. The collimation grid 90 is designed such that only X-rays 100 directed to the detector 110 pass through it. The collimation grid 90 does not move with respect to the detector array 110 while the device is operating. Thus, whenever the electron beam 40 scans the anode 50, there is only a single x-ray beam 100 passing from the anode 50 to the detector array 110.
FIG. 2 shows the distribution of X-rays without a collimation grid. The output of the detector array 110 is processed and displayed on the monitor 120 as a luminance value at the x, y position on the monitor 120 corresponding to the x, y position of the anode 50. This is achieved by using the same scan generator that drives the x, y position of the electron beam 40 and the electron beam in the video monitor 120. Also, image processing techniques can be used to produce computer operated images on a suitable display or photographic media.
The inventive apparatus disclosed herein is typically measured at the point of incidence on the patient and ranges from about 0.15 R / min at an update rate of 15 frames / sec to about 0.33 R / min at an update rate of 30 frames / sec. It is a low-dose device that gives patients doses up to minutes. With this device, an update rate of 30 frames / second would be about 0.50 R / min for the whole body. The actual range of radiation dose at the patient surface for operation of the invention disclosed herein is from 0.15 R / min to 2.00 R / min.
X-ray tube
FIG. 3 shows an enlarged structure of the grid and the anode. The anode 50 is formed on a beryllium anode support 130 and is preferably assembled with a target layer of a material having good vacuum properties and the ability to resist high temperature and electron bombardment. Additionally, aluminum or other materials that are relatively transparent to X-rays can be used for the anode support 130. A presently preferred target layer comprises, in order of preference: (1) a second layer of tantalum sputter deposited approximately 5 microns thick onto a first layer of niobium sputter deposited approximately 1 micron thick on the anode support (this The structure is such that niobium has a coefficient of thermal expansion intermediate between that of beryllium (anode support 130) and tantalum, thus suppressing or reducing microcracks due to thermal cycling of the target anode with the tube on and off. (2) a sputter deposited layer of tantalum approximately 5 microns thick, (3) a sputter deposited layer of tungsten-rhenium approximately 5 microns thick, and (4) a sputter deposited layer of approximately 5-7 microns thick. This is a sputter-deposited layer of tungsten. Tantalum, tungsten and tungsten-rhenium are preferred for anode 50 because they have relatively high atomic numbers and densities and readily emit X-rays when irradiated with an electron beam. The high melting point of 3370 ° C. and good vacuum characteristics of tungsten are suitable for the high temperature and ultra-high vacuum conditions in the X-ray tube. Tantalum and tungsten-rhenium have similar properties as known to those skilled in the art. The thickness of the anode layer is selected to be approximately equal to the distance required for the anode layer to effectively convert 100 KV electrons to x-rays. Beryllium is preferred for the anode support 130 because it is robust and does not significantly reduce or dissipate X-rays emitted from the anode 50. The thickness of the beryllium anode support 130 is preferably about 0.5 cm. The anode support 130 is subject to the physical constraint that it must be strong enough to resist a pressure gradient with one atmosphere outside, but must be as thin as possible.
The collimation grid 90, according to one preferred form of the invention, comprises an array of apertures 140, each oriented or oriented in the direction of the detector array 110. That is, the apertures in the collimation grid 90 are not parallel to each other and, for example, in a use such as a chest x-ray application, relative to the front face 260 of the collimation grid 90, the 0 ° in the center of the collimation grid 90 to the grid 90. An angle of about 20 ° at the end can be made. In use of the present invention in a breast examination application, the grid 90 can be assembled such that the opening is angled with the front surface at the end of the grid in the range of 45 °. The number of openings 140 in the collimation grid 90 can correspond to, for example, 500 × 500 to 1024 × 1024 pixels at the center of the round collimation grid 90, and can determine the resolution of the device. Also, apertures smaller than the number of pixels can be used in connection with the sub-sampling technique described below. The thickness of the grid 90 and the dimensions of the apertures 140 depend on the distance from the X-ray tube 10 to the detector array 110 (preferably 36 inches), and any X-rays not directed to the detector. The size of the detector element 160 (not shown in this figure) of the detector array 110 is determined by the need to subtract the line. Although not critical to the present invention, the openings 140 as viewed from the front surface 260 are preferably designed in a rectangular column and row pattern having a circular boundary of 25.4 cm (10 inches) in diameter. The aperture array can be a convenient interrelated layout that uses the detection and convolution techniques outlined below to analyze the image of the object 80. This aperture array is called a "circular active area." At the center of the circular active area, the total number of openings is preferably 500 × 500 according to a preferred embodiment of the present invention. The non-opening portion 150 of the collimation grid 90 is designed to absorb the shifted X-rays so that the shifted X-rays do not illuminate the object 80. This means that the X-rays that impinge on the non-aperture 150 are "1/2" (the amount of material needed to reduce the X-rays that impinge on it at half the energy of this device, here at 100 KeV). This is done by assembling the grid to show at least a factor of ten. The shifted X-rays irradiate the object and staff with X-rays, but do not give useful information to the image. As shown in FIGS. 3A and 3B, the collimation grid 90 may be formed from multiple sheets of X-ray absorbing material 143, 144 having openings 140 therethrough to pass the X-ray beam 100 to the detector. it can. The collimation grid 90 is preferably assembled by stacking 50 thin sheets of molybdenum 0.00.04 inch (0.010 inch) thick and holding them together. Molybdenum is preferred because it stops x-rays generated by the x-ray tube 10 that are not directed to the detector 110 from futile and detrimental impact on objects 80, including, of course, human patients. Lead or similar X-ray dense materials can also be used.
The aperture 140 of the collimation grid 90 is preferably square in cross-section to obtain maximum packing density and to match the preferred square shape of the detector array element 160. Other shapes, especially hexagons, can also be used. The square opening 140 is preferably 0.0381 cm (0.015 inch) by 0.0381 cm, providing a cross-sectional area of about 1/100 of that of a conventional collimator used in fluoroscopy. Due to this tighter collimation, a smaller beam width for the X-ray beam 100 is achieved. This means that the cross-sectional area of the detector surface can be correspondingly smaller than in conventional devices. As a result, the X-rays scattered by the object are not detected by the detector, and the image is disturbed as if the scattered X-rays acted in a conventional device using a detector having a relatively large surface area. Absent.
The preferred method of assembling the collimation grid 90 is by photochemical milling or etching. Photochemical milling is preferred because of its cost efficiency and accuracy. According to this method, a set of 50 photomasks is made to etch holes and gaps in 50 sheets of 0.0254 cm (0.010 inch) thin sheet material. The etched sheets are stacked, arranged, and held together to form a grid assembly having a number of stepped openings, each having a predetermined angle with respect to each sheet. FIG. 3A shows a variation of the inventive collimation grid 90. This variation includes a number of X-ray absorbing sheets 143 having individual openings of a constant cross section (but the cross section need not be constant). The resulting aperture 14 is stepped, as shown, but passes the x-ray beam 100 to a detector array 110. The variation shown in FIG. 3B is quite similar to that shown in FIG. 3A, except that the individual openings formed in the X-ray absorbing sheet 144 are themselves stepped. These stepped apertures can be made by milling or chemical etching from each side of the sheet 144 with slight staggering to be the shape shown, as will be apparent to those skilled in the art. The shape of FIG. 3B results in less x-ray energy absorption within the stepped aperture 140 of the collimation grid 90, so that the x-ray flux at the end of the x-ray beam 100 is the same as the variation shown in FIG. 3a. It is more preferable because it is not reduced to the extent.
One preferred method of holding the etched sheet (s) forming the grid 90 is shown in FIG. Each of the etched sheets 91 (preferably 50) is provided with an alignment hole or alignment opening 94. Alignment pegs 95 are positioned within respective alignment openings 94 to align the etched sheet 91. The assembly of seat 91 and peg 95 is located within aluminum ring 359. The aluminum ring 359 is provided with a vacuum port 370 that can be sealed with a pinch-off 375. Then, the aluminum sheet 365 having a thickness of 0.1 cm is bonded and sealed to the upper surface 380 of the ring 359 with a vacuum adhesive. Aluminum sheet 360 is similarly adhered to lower surface 385 of ring 359. Port 370 is partially evacuated via port 370, and port 370 is sealed with a pinch-off 375 as is well known in the art. In this manner, the relatively x-ray permeable aluminum sheets 360, 365 provide a clamping action that helps to hold and orient the etched sheets 91 as a grid assembly 90 together.
The aperture 140 furthest from the center of the grid 90 has a stepped surface and is preferably square in cross section. X-rays are generally unaffected by the asperities of the passages due to the stepped surfaces, and even if they are scattered, they will not measurably affect the combined beam. As described above, the material used for the collimation grid 90 is currently preferably molybdenum, brass, lead, or copper with molybdenum. The preferred error for the hole locations is ± 0.00127 cm (0.0005 inches) from center to center, excluding the cumulative error, and the hole size error is ± 0.00254 cm (0.001 inches).
Other methods of making usable collimation grid 90 include e-beam machining, drilling, mini-machining, and laser drilling. Drilling and laser drilling have the disadvantage of creating rounder holes than square holes. Circular openings work well, but are currently undesirable.
More details of the preferred scanning X-ray tube 10 are shown in FIGS. The electron gun 161 is located on the side opposite to the front of the X-ray tube 10 and operates at a potential of about -100 KV to -120 KV. The grounded anode 50 is located in front of the tube, and the electron beam 40 moves between the electron gun 161 and the anode 50. The grounded electron aperture plate 162 is located near the electron gun 161 and has an opening 163 at the center where the electron beam 40 passes. The magnetic focusing lens 164 and the deflection coil 20 position the beam spot on the anode 50 using self-focusing as is well known in the art. The tube is assembled such that the electron beam 40 has a 25.4 cm (10 inch) diameter circular active area that intersects the anode 50 with the electron beam deflected by about 30 ° at the end of the circular active area. . When the beam is not "fired" through a certain aperture, it is preferable to stop producing the beam, resulting in power savings of about 25%.
FIG. 5 shows a cross-sectional view of the front portion of a suitable X-ray tube 10. The interior of the X-ray tube 340, which is maintained in a vacuum, is behind the anode 50. Anode 50 is a coating of an anode material as described above. In front of the anode 50 is a 0.5 cm thick beryllium anode support 130. The front of the beryllium anode support 130 is a cooling jacket 350 that is preferably 0.4 cm thick and can be adapted to flow water or forced air. Aluminum grid supports 360, 365 are etched 0.1 cm thick and help support a collimation grid 90, which is preferably 0.5 inches thick.
When an x-ray tube 340 is used, only one aperture 140 in the collimation grid 90 will pass a significant amount of x-ray at any given moment. According to a preferred embodiment, the electron beam 40 can be stopped when the electron beam 40 is not directly in front of the aperture 140. Thus, the x-ray tube is effectively operable in scan pulse mode to reduce power consumption and reduce wear and tear on the target anode 50.
Stereoscopic X-ray photography
In FIG. 6, according to another preferred embodiment of the present invention, it is possible to provide a grid having a plurality of focal points so that stereoscopic radiography can be obtained. For example, if all other rows of apertures in grid 90 were directed to focus F1 (92) and the remaining apertures were directed to focus F2 (93), place the first sensor array at F1 (92). By placing the second sensor array in F2 (93), it is possible to scan the aperture in a raster or serpentine pattern, thereby providing a "line" of data for the first sensor array and a second sensor array. A "line" of data about the array can be created. This can be repeated to assemble two completed images as seen from two different points F1 and F2 in space, thereby providing a conventional stereoscopic display that provides a stereoscopic X-ray image. You can display them. FIG. 3C shows how to construct a stereoscopic collimation grid from within the layer 144 of X-ray absorbing material. In this case, the apertures 140A, 140B, as shown, provide a separate path along the "V" "leg" for the x-ray beams 100A, 100B, and are actually a "V" shaped It can be. However, although it is not necessary that the apertures 140A, 140B be connected as shown, the advantage of a "V" shaped aperture where X-rays are incident at the apex of the "V" is that both detections The detectors are illuminated simultaneously and "V" is to act as a signal separator between the X-rays going to F1 and the X-rays going to F2. This halves the power required for beam and deflection current. The price is lower, but the scattering and consequent image haze increase.
Array detector
In order to achieve a resolution of a few lines / mm at the object plane, as required in some medical applications, the spatial resolution limit is mainly determined by the size of the detector. This can produce very high power levels, which would be required in today's X-ray tube technology to obtain very well-directed X-ray radiation, as well as the combined X-ray direction. Neither is it possible to develop an attachment.
If the detector is made smaller than the area that intersects the cone of emitted X-rays, a large percentage of the X-rays emitted by source 50 will be detected by detector 250, as shown in FIG. 7A. Not done. In practice, such an industrial beam scanning digital X-ray inspection apparatus is designed in this way, where the irradiation dose is normal and not a problem. As a result, the irradiation dose is increased to maintain the desired resolution.
Thus, the resolution is improved with the use of a smaller detector, but when the area of the detector is equal to or greater than the area defined by the cone of radiated X-rays intersecting the detector plane 270, the X irradiation Dose is minimized.
The resolution of the scanning X-ray apparatus is determined by the cross-sectional area of the detector element projected on the object plane 280 (a plane perpendicular to the line between the center of the anode 50 and the center of the detector 110 where the object 80 is located). Is determined by Thus, as shown in the front view of the detector array in FIG. 8, if the large area detector is divided into smaller array elements, the large capture area of the detector assembly is maintained, while at the same time. , Maintaining an image resolution proportional to the size of each small detector element 160.
The resolution defined by the individual detector elements 160 is maintained by distributing and summing the readings from the individual elements to each address or pixel to a memory buffer corresponding to a particular location on the object plane 280. Is done. As the x-ray beam 100 moves individually across the collimation grid 90 located before the x-rays radiating to the anode 50, the address at which the output of a given detector element is added changes. The imaging geometry is shown in FIGS. 7B and 7C. FIG. 7B shows how a single beam position is divided into five pixels. FIG. 7C shows how successive beam positions add beams to each other within a single pixel.
In other words, the signal for each detector element is stored in the image buffer at a very small specific area on the object surface 280, ie, at a memory address corresponding to a single pixel. Accordingly, the memory storage address for each detector element changes to the position of the scanned x-ray beam in a manner in which each pixel of the memory is arranged to include the sum of the radiation that has passed through a particular spot on the object surface 280. In this manner, the resolution of the device is determined by the size of a single detector element, while the sensitivity of the device is highest because substantially all of the X-rays reaching detector surface 270 are recorded.
A further advantage of the imaging geometry of this array detector is that the object plane 280 is narrowly defined. Structures in front of or behind it will be hazy (out of focus). The X-rays from the first opening 141 and the second opening 142 pass through the object surface 280 at a distance SO from the openings 141 and 142, and pass through a surface 281 twice the distance SO from the openings 141 and 142 in FIG. 7D. Is shown in As can be readily seen, the resolution obtained at twice the SO drops to about 1/2 of the resolution obtained at the SO. This feature provides improved placement and viewing of detailed structures at the plane of interest 280, while providing sufficient depth of the field that can be changed by the geometry of the device.
The array of the presently preferred embodiment is a 96-element pseudo-circular array of 0.135 cm-sided square detector elements arranged in a circle of about 1.93 cm (0.72 inch) diameter. It is not necessary to be so large, but it is also possible to have three or more detectors arranged such that they do not all line up within a circle of radius equal to the length of one detector, here 0.135 cm.
X-ray detector
Conventional image intensifier technology basically has the constraint of limiting the sensitivity of the system. The thickness of this usable scintillator material is limited by its light transmission properties. Typically, it is made thick enough to capture about 50 percent of the incident X-ray photons. Only half of the emitted photons reach the photocathode. At the photocathode, only about 10 percent of the incident photons produce photoelectrons. In this way, only 2.5 percent (0.5 × 0.5 × 0.1) of the incident X-ray photon energy is maintained in the image intensifier tube system. In addition to this limited conversion factor, photons are scattered longitudinally by the scintillator material, resulting in blurring that reduces the resolution of the system at a given dose level.
One of the underlying objects of the present invention is to provide an SBDX imaging system that ensures that the object to be inspected is exposed to the lowest possible X-ray levels while achieving the appropriate image quality required for the procedure to be performed. It is to provide. This means that the system used to detect the X-ray photons emerging from the object must have the highest photon-to-electrical signal conversion efficiency. To achieve this, the material used for the detector must be long enough to ensure that the photon travels in a direction in which it does not enter from far away from the incident x-ray. No. For example, to maximize the output of the detector, photon energy must be properly wasted in the material. There are many types of detectors that can be used in the SBDX system described here. Preferred are scintillators in which the X-ray photon energy is converted into visible energy and the light intensity is converted into an electrical signal by means of a photomultiplier, photodiode, CCD or such device means. Since each pixel of the SBDX image must occur in a very short time of about 140 nanoseconds, the scintillator material must have high responsiveness and minimal afterglow time. Afterglow refers to a phenomenon in which the scintillator continues to emit light after the excitation incident X-ray radiation ends. Plastic scintillators such as polystyrene mixed with organics are suitable in that they have the required high responsiveness, but have relatively small X-ray photon interaction cross-sections and low linear X-ray absorption coefficients. Value. As a result, a significant thickness is required to stop X-ray photons. For 100 kV x-rays, a typical plastic scintillator requires a 28 cm (11 inch) thickness to capture 99% of the input x-rays. More preferably (in a preferred choice) herein, (1) cerium-doped YSO (available from Yttrium Oxy-Orthosilicate, Airtron (Litton), Charlotte, NY); (2) cerium-doped LSO ( Lithium oxy-orthosilicate, available from Schlumberger, and (3) BGO (bismuth germanate, available from Lexington Component, Beachwood, Ohio). YSO and LSO are excellent in that they can be used at room temperature. BGO must be heated to about 100 ° C. to achieve a reasonable light output decay time on the order of 50 naniseconds. These scintillator materials do not need to be as long as plastic scintillators and are effective at 0.10 cm in length.
According to the presently preferred embodiment of the present invention, the SBDX array detector 110 comprises 96 closely bundled individual, spaced 9 inches (36 inches) from the X-ray source 50. X-ray detector 160 comprises an array of 12 × 12 pseudo-circles. (A 5x5, 3x3 array is also considered as a non-square array (filling around the X-ray target) with a square detector. See, for example, the lower part of Table 1.) Aspect ratio shape of the collimation grid The distance from the x-ray source 50 and the size of the x-ray source 50 provide a square pyramid at the entrance of the detector array 110 of approximately 0.9 inches in diameter with a total included angle of 1.46 °. Each scintillator 170 Must therefore be about 0.152 cm (0.06 inches) from center to center in the plane of the detector.If the scintillator 170 has parallel sides, X-rays incident near the edge Would not be able to travel the required distance without hitting the scintillator wall, so that these x-rays would pass to a nearby scintillator and if they were not shielded, Apparent emergence from a spatial location results in a consequent degradation of the image quality of the object.As shown in FIG. 9A, to avoid this effect, each scintillator according to a preferred embodiment of the present invention is , Preferably tapered, and the interface 173 has an included angle equal to the extreme angle of the incident x-ray 100 ', which is particularly useful with long plastic scintillators. Each scintillator 170 is a preferred 28 cm long truncated pyramid having an entrance surface 172 of 0.285 cm diameter and a photodetector end surface 174 of 0.37 cm diameter.Each of the 81 detector bundles is therefore a polyhedron And each face is in contact with the surface of a sphere located at the center of the X-ray source 50.
Further improvement in scintillator detection efficiency is achieved with a scintillator taper at an angle greater than the angle of the incident x-ray beam 100. Near the ends of the scintillator, photoelectrons and scattered x-rays generated by the interaction between incident x-rays and scintillator atoms can be lost to the shielding material, which preferably separates adjacent scintillators. Such lost photoelectrons do not produce any light and do not behave as light exiting the scintillator. Therefore, such extinction reduces the efficiency of the scintillator. The maximum distance a photoelectron travels depends on its energy and the material it travels. In the interaction of electrons with 100 kV X-rays in a plastic scintillator, no photoelectrons can travel more than about 0.01 cm. If the truncated pyramid of the scintillator is made with an included angle larger than the X-ray beam 100, its size will be a short distance 2 × 0.01 cm compared to the length of the detector (28 cm). Being larger than the enclosed beam, loss of efficiency due to lost photoelectrons is minimized. In this case, the center of the sphere touching the surface no longer coincides with the x-ray source 50 and is closer to the detector array 110.
Scattered photons travel a longer distance than photoelectrons. To prevent them from escaping to adjacent scintillators, the value of the taper of the scintillator pyramid is set to be greater than that required for full photoelectron capture to maximize the capture of scattered photons.
Referring now to FIG. 9, according to a preferred embodiment of the present invention, a light pipe that contacts each scintillator element 170 and optically couples each scintillator element 170 to a corresponding photomultiplier tube 190 or solid state detector. Alternatively, a fiber optic cable 180 is provided. Alternatively, the scintillator 170 may be located in physical proximity to a particular photodetector.
FIG. 10 shows a preferred shape of the detector element 160. An X-ray opaque sheet 200 having an aperture 210 corresponding to each detector 160 is placed in front of the detector array 110. Each detector element 160 is surrounded by an enclosure 220 that is light tight and opaque to X-rays. A light blocking window 230, preferably made of an aluminum plate, is placed in front of the light tight enclosure 220. The light blocking window 230 transmits X-rays. Within the light tight enclosure 220, the scintillator element 170 is located close to a photomultiplier tube 190 electrically connected to the preamplifier 240. Preferably, the analog signal from preamplifier 240 is converted to a digital signal in a conventional manner by further processing.
Alternatively, the scintillator can be placed directly on or near an array of photodiodes, phototransistors, or solid-state imaging devices (CCD) to obtain a coarser, more compact detector. Especially when a solid state device such as a CCD is used, cooling using a Peltier cooler or the like can be used to increase the signal to noise ratio of the device.
Alternatively, the scintillator array can be located directly or close to the photomultiplier tube, which, like its size, identifies one or more locations that provide an output signal, identifying locations equivalent to the light source. .
In another preferred embodiment, the sensor array may consist of a collection of aligned pencil detectors 285, for example, as shown in FIG. In FIG. 11, the tapered scintillator 290 is aligned with the path of the x-ray beam 100, and the scintillator corresponding to a particular cross-sectional area of the beam 100 absorbs all x-rays within that cross-sectional area. The optical amplifier tube 300 is provided physically close to the scintillator 290, and an electric signal is generated in response to X-ray absorption by the scintillator 290. A solid state device can also be used in place of the light amplification tube 300.
In accordance with this preferred embodiment of the present invention, the scintillator, such as silicon dioxide, prevents light from escaping (or entering) along its length and at the plane of incidence and aids internal reflection within the scintillator. It is covered by a material that reflects light.
According to another preferred embodiment of the present invention, each scintillator element 179 is separated from adjacent scintillator elements by a thin film 171 of a high radiopaque material such as gold or lead. The membrane 171 is preferably about 0.0102 cm (0.004 inches) to 0.0127 cm (0.005 inches) thick. The location of the membrane 171 between the scintillators 170 is shown in FIG.
As shown here, the circular active area of the collimation grid 90 is larger than the area of 110 of the detector array. In this way, the X-ray pencil beams emitted from the respective apertures 140 of the collimation grid 90 all converge on the detector array 110. On the other hand, each independent X-ray beam 100 branches and spreads like a flash beam.
Image processing
An important aspect of the present invention relates to the application of the image processing system to further reduce the required irradiation dose. In practice, the signal from the detector is not normally used directly for the "Z" or luminance input of the video monitor. Instead, the digitized intensity of each pixel is stored at a separate address in the "frame store buffer". Pixel addresses in the buffer can be accessed randomly and numerical intensity values are treated mathematically. This feature has many image expansion algorithm applications that can be applied, and allows the allocation of pixels of data from independent parts of the detector array.
According to a preferred embodiment of the present invention, the SBDX image is composed of about 250,000 or more pixels of 500 columns and 500 rows (corresponding to 500 columns and 500 rows of the central opening of the collimation grid 90). For the purposes of the following example, it is assumed that the scanning X-ray source is at a certain moment at the center P on a pixel at 100 columns, 100 rows of the collimation grid 90. Further, for this embodiment, detector array 110 comprises a 3 × 3 array including nine segments 179 (FIG. 12), with each segment 179 blocking all x-ray radiation associated with a single pixel. Size. Obviously, other array shapes may be used as detailed herein.
The digitized numbers from the individual segments of the detector array 110 are assigned the following pixel addresses.
Segment 1-99 columns, 99 rows
Segment 2-99 columns, 100 rows
Segment 3-99 columns, 101 rows
Segment 4-100 columns, 99 rows
Segment P-100 columns, 100 rows
Segment 6-100 columns, 101 rows
Segment 7-101 columns, 99 rows
Segment 8-101 columns, 100 rows
Segment 9-101 columns, 101 rows
As the scanning line X-ray beam passes through all pixels, the same data allocation pattern is repeated.
In the displayed image, the value of each pixel is equal to the sum of the "n" parts. Here, “n” is the number of the segments 179 of the array 110 (n = 9 in this example). When configured as shown here, the detector array 110 has a fixed operating distance for optimal focus and is not available with conventional non-segmented (non-segmented) detector array SBDX imaging systems. This has the effect of providing a plane of optimal focus.
The following parameters must be taken into account in the design of the detector.
1. The size and shape of the parallel beam from the X-ray source (anode target) 50.
2. Distance between X-ray source 50 and detector array 110; "SD"
3. Distance between X-ray source 50 and center of object 80 of interest; "SO"
4. Desired resolution or pixel size of object 80 of interest
5. In medical applications, the entire area of the array must be large enough to block X from the collimation grid 90.
In the SBDX system for the preferred embodiment of the present invention, the distance between the x-ray source 50 and the exit 260 of the collimation grid 90 is about 0.894 inches (see FIGS. 3 and 5). The opening 140 is a 0.0381 cm (0.015 inch) × 0.0381 cm square. The spot size of the electron beam 40 on the anode 50 is about 0.0254 cm (0.010 inch) in diameter. The detector array 110 is 36 inches (91.4 cm) from the anode 50. Thus, the beam width of X-ray beam 100 is 2 * ARCTAN ((spot diameter / 2) / ((opening width / 2) + (spot diameter / 2)) * 2.271 cm (0.894 inches), or 1.6 °. At a distance of 91.4 cm (36 inches) from the anode 50, the illuminated X-ray beam diameter is 91.4 * TAN (1.6 °) cm. Therefore, the detector array 110 is of this preferred embodiment. On the side, it should be about 2.54 cm (1 inch), for example, the object to be photographed is 22.86 cm (9 inches) from the anode and the desired pixel size is 0.0508 cm (0.020 inches) on the object And the distance from the X-ray source to the detector is 91.4 cm (36 inches) and the size of the optical detector array is 2.54 cm (1 inch) square, the illuminated at the detector plane 270 The pixel size is simply the (SD / SO) * pixel size of the object or 0.2032 cm (0.080 inch) Dividing 2.54 cm (1 inch) by 0.2032 cm (0.080 inch) yields the desired resolution using a square-sectioned detector with 12 to 13 segments on the surface. Obviously, many other shapes will be used depending on the context in which the SBDX system is used.
Outside the surface SO (280 in FIG. 7D) of the optimum resolution, the resolution is reduced by half at 0.5 × SO and 2 × SO (281 in FIG. 7D). This allows a reasonable depth of focus for many applications. In some applications, such as an image of the human heart, a focus degraded outside this depth range is seen as an advantage. Blurring of details outside the region of interest increases the perception of details within the region of interest.
Many methods can be used to obtain a usable image from the data obtained as described above. As mentioned above, a simple convolution can be used, but in this case the resolution has not been optimized at all. Two additional methods are preferred for obtaining maximum resolution and sensitivity from the obtained data. These are called multi-image convolution and multi-output convolution. In both cases, the following is assumed.
Aperture AP in collimation grid 90_XRow, opening AP_YThere are rows. The intersection of each row and column is a "pixel". Those pixels outside the circular active area of the collimation grid 90 are treated as if there was no measured intensity of the image, for example, as "dark". Pixels that are not illuminated by the X-ray beam during scanning are similarly treated as if they were "dark", as if there were no measured intensity of the image.
Referring to FIG. 15, the pseudo circular sensor array 110 has the DET of the sensor element 160 of the detector array._XColumn maximum and DET of sensor element 160 of detector array 110_YThere is a maximum value for the row.
ZRATIO is a real number between 0 and 1. If ZRATIO = 1, the focus is set on the sensor plane. If ZRATI = 0.5, the focal point is set at the midpoint between the X-ray source and the sensor plane. PIXELRATIO is the number of pixels for the physical distance between neighboring sensors in a column or row. For example, if the distance between the pixel centers on the object plane 280 is 0.01 cm, the distance between the sensors on the detector plane 270 will be 1.0 cm, PIXELRATIO = 10, FOCUS = ZRATIO * PIXELRATIO.
IMAGE is a DET containing intensity information for special scans and corresponding to special pixels._X× DET_YIs the data string of the dimension. PIXEL is the DET obtained by scanning all (or part) of the aperture_X× DET_YAP containing IMAGE data column_X× AP_Y× DET_X× DET_YIn a four-dimensional array. PIXEL is refreshed after each scan according to one preferred embodiment of the present invention. As the beam is scanned across the anode surface, it is effectively placed before the center of the selected aperture 140, "fired" and then reset. Thus, an IMAGE sequence of data is obtained for each irradiation. While these images are assembled into displayable images with some direct use, high resolution, high intensity is obtained by combining them. The preferred method of combining the first images is called the multi-image rotation method. In this multi-image rotation method, the AP that can be displayed on a CRT or similar display means_X× AP_YAn OUTIMAGE sequence of dimensional intensities is formed by assigning OUTIMAGE (y, x).

AP_X× AP_YA second currently preferred method of combining IMAGE data sequences into useful images is called the multi-output swirl method. In this case, DET_X× DET_YDET with sensor array of sensors_X× DET_YDigitizers (or equivalent or multiplexed) and the same number of pixel summing circuits. The digitized value from each sensor is called SENSOR (j, i). The final OUTIMAGE column is calculated as follows: Output image sequence OUTIMAGE (y, x) [y = 1 to AP_Y, X = 1 to AP_X] For each pixel_X× DET_YOne pixel from each of the source images SENSOR (j, i) is summed [j = 1 to DET_Y, I = 1 to DET_X], And the target image pixel OUTIMAGE (y−j * FOCUS, x−i * FOCUS). Then, each element is DET_X* DET_YBy dividing by, the OUTIMAGE sequence is standardized.
Further refinements of these techniques can be obtained by performing linear interpolation based on the fractional part of the FOCUS element.
The advantage of the multi-image swivel method over the multi-output swivel method is that the former allows the software to select the plane of optimal focus after data acquisition, whereas the latter does not. However, the latter allows faster operation when time is at a limit.
Reconstruction of 3D image from SDBX data
The SBDX system described herein can be used to create a set of continuous planar images that are used to form a tomographic three-dimensional image of the object 80. A set of images can be analyzed to produce images in three dimensions consisting of a series of images at different depths by re-analyzing the data set at different FOCUS values. The natural FOCUS value used is n / DET_XOr n / DET_YWhere n is each 0 to DET_XOr DET_YIs an integer up to. Typically, the focus value is analyzed to match a plane of interest within the object 80. For example, in the SBDX system presented in Table 1 below, the plane of focus is near the normal focal plane of 9 inches (optimal focal plane) and spaced about 2.54 cm (1 inch) apart. I will be.
The following equation shows that a series of planar images is located in terms of distance from the anode 50.

When using the sub-sampling technique, the calculations remain the same and the data from the aperture of the "not skipped" collimation grid is handled. However, even if there is no intervening collimation grid hole, λ_tIs the same.
Negative feedback x-ray flux control
The BDX imaging system shown in FIG. 13 uses a negative feedback path 305 that controls the x-ray flux of the x-ray beam 100. Preferably, the negative feedback from the sensor array controls the x-ray flux so that the sensor array always sees approximately the same x-ray flux level. Thus, when soft tissue (which is relatively transparent to x-rays) is scanned, the x-ray flux drops, reducing the overall dose to the patient (object). The use of negative feedback flux control improves contrast and dynamic range. According to this embodiment, the differential amplifier 310 has an adjustable reference level 320 that can be set by a user. The negative feedback loop 305 provides negative feedback of the x-ray flux to the x-ray tube.
Time domain scan mode
Also, a time domain imaging system can also be implemented using the principles described herein. In such a system, the time to reach a pre-measured x-ray flux from various pixels can be calculated and mapped. Negative feedback control can then be used to remove or reduce x-ray flux from the aperture corresponding to pixels that have reached a predetermined flux level for the scan period being handled. In this case, the information gathered is the time to flux level, and the mapped and captured information corresponds to time rather than intensity. The ability of such a system provides a much higher signal-to-noise ratio, improves contrast, dramatically reduces x-ray exposure to the object under inspection, and allows for improved dynamic range.
Multi-energy x-ray imaging mode
According to one embodiment of the present invention, two or more groups of x-ray beams 100 travel toward one or more detector arrays. The first group of x-ray beams has a first characteristic x-ray energy spectrum. The second group of x-ray beams has a different second characteristic x-ray energy spectrum. By comparing the measured transmittances of the x-ray beams of the first and second groups, the presence of an object in the object under inspection can be detected. The basic concept of differential radiography is previously known and is disclosed in US Patent No. 5,185,773 ("Methods and Apparatus for Non-Destructive Selective Determination of Metals"), which is hereby incorporated by reference. Incorporated in
The two groups of x-rays can be generated in a number of ways. In one method, a special anode comprising a first material or first thickness of material near a first group of openings and a second material or second thickness of material near a second group of openings. Perform manufacturing. Thus, apertures associated with the first group emit x-rays having a first characteristic energy spectrum, and apertures associated with the second group emit x-rays having a second characteristic energy spectrum. Alternatively, K-filtering (K-edge filtering) can be used by placing a filter material (such as molybdenum) within a portion of the aperture 140 that produces a similar effect. In this case, a first group of apertures comprises a first filter inserted therein, and a second group of apertures comprises a second filter inserted therein. The second filter may not be a filter. As in the previous case, two groups of x-rays having different characteristic energy spectra are associated with two groups of apertures.
When at least two groups of apertures are associated with different characteristic energy spectra, it is possible to detect microcalcifications (breast cancer) and other abnormalities not normally visible on broadband x-rays. For example, by scanning a first group of apertures forming a first image, scanning a second group of apertures forming a second image, and then enhancing those ratios by dividing the image, Microcalcifications and other abnormalities can be detected in real time using a low dose scan x-ray imaging system. Similarly, a multiple detector array device can be used with a first group of apertures towards the first detector array and a second group of apertures towards the second detector array.
Another embodiment of multiple energy imaging will now be described. Since the magnitude of the electrical pulse from the detected x-ray photon is proportional to the energy (kV) of the photon, it is possible to separately count the pulses coming from the photon in more than one energy band. The pulses are separated by intensity, then counted and processed separately to create two or more separate images. These images can be displayed as ratios.
Also, it is possible to vary the selected energy level in the fly to distinguish different density regions in the object. The advantage of this embodiment is that it is more flexible than the embodiments described above, and does not require special collimation grids, anode materials or dual detectors.
While a number of preferred embodiments have been described above for various configurations of the present invention, the following description describes a preferred SBDX imaging system according to the present invention.

Thus, this SBDX imaging system, which utilizes a segmented detector array, has high resolution and high sensitivity, and at the same time, low irradiation dose to the object under inspection. The system can also set the optimal focus at any point between the source 50 and the detector array 110, and provide an effective working depth of field.
Beam subscan technique
The following description relates to a particularly preferred embodiment of the present invention, which uses a technique of beam sub-sampling to reduce computer processing overhead.
The standard video image uses 640 × 480 pixels and is updated at 30 Hz. This requires a pixel sample rate of about 12 MHz. Positioning the high-voltage electron beam of the x-ray tube exactly 250,000 consecutive different apertures at this speed requires high precision and relatively high power consumption. Digitization of signals from a large array of x-ray detectors at a rate of 12 MHz is equally effective and power consuming. Thus, without a significant decrease in the spatial or temporal resolution of the SBDX, a decrease in pixel sample rate below 12 MHz would result in an initial unit cost, operating costs due to power consumption, and wasted heat due to x-ray tubes. Useful for reducing cooling requirements for
Accordingly, mechanisms have been developed that provide substantially the same spatial and temporal resolution while reducing the pixel sample rate. This mechanism, referred to as subsampling, is obviously applicable in other configurations of SBDX, but is best implemented in the SBDX embodiments described herein. The advantages of this embodiment include reduced power consumption, a simpler circuit for deflecting the electron beam in the x-ray tube, lowering the cost of creating the collimating grit 90, and the calculations required to resolve the image of the object 80. , And other effects apparent to those skilled in the art.
In this embodiment, the collimation grid 90 has less than 500 openings, preferably AP_x= AP_y= 166 (although other numbers could obviously be used). The effect of this reduction from a computational point of view will become apparent below. However, the advantage from a manufacturing point of view is a much simpler structure, about 1/9 of the number of openings required for manufacturing. Manufacturing the grid at a higher deflection angle (i.e., the angle that the openings make with respect to the front surface 260 of the collimation grid), as the number of openings is reduced, creates the problem of having openings that intersect adjacent openings. Not easier. This is because, when a solid grid is manufactured, adjacent apertures in the solid grid are directed toward different detector arrays, and thus require a wider physical separation than non-solid grids to prevent aperture crossing. Especially useful.
The aperture (s) in the collimation grid are AP_xRows and AP_yThe rows are arranged in circles. For computational purposes, this is the AP with the element (s) outside the circle that does not contribute information, ie, always "dark", ie not illuminated by x-rays_xRows and AP_yCan be treated as a rectangle with column dimensions.
The sensor 160 of the x-ray detector array 110 has a maximum size of DET as shown in FIG._xRows and DET_yIt is arranged in a circular array of rows. The pixel sampling rate can be reduced by illuminating less than all of the apertures in the collimation grid, ie, subsampling. Preferably, a collimation grid with no non-illuminated openings is used. In order to create an image using this detector array, all DET_xDET in the hole of each collimator and each row_yOnly the holes of the second collimator need to be illuminated, so the images are each DET_xPixel and DET_yIt can be assembled from pixel image tiles of pixel dimensions. This is DET_x× DET_y, But subsampling does not correspond to a 1 × 1 sampling rate. Therefore, this sub-sampling ratio is 1 to DET in the x direction (row)._xFrom 1 to DET in the y direction (row)_yCan be adjusted up to. According to this preferred embodiment, as shown in FIG._x= DET_y= 12.
When a 12 × 12 detector is used and the sub-sampling ratio is 12, the image is much more similar to David Hockney's light mosaic, from multiple non-overlapping images that are substantially stuck together. Created. Since the actual scintillator and detector are not completely perfect and do not respond exactly the same, the x-ray pencil beam is not perfectly uniform and the aperture of the collimation grid should be ideal. Is not exactly the same as the typical area. And since circular detectors are used instead of square detectors, some overlap is highly preferred to allow averaging of the detector nonlinearity and noise.
If the sub-sampling ratio is smaller than the size of the detector in the pixel (ie, smaller than 12 in this preferred embodiment), the image is assembled from overlapping "tiles" that must be summed or averaged. No. If the subsampling ratios are not multiples of the detector dimensions (in pixels), or if the detector array is not rectangular, a different number of samples will be added to each pixel and a different division Is needed to average each pixel. Techniques for handling fewer than ideal environments are well known to those skilled in the art and will not be described here to avoid overcomplicating the description.
In the following calculations, SSX represents the sub-sampling dimension in the X direction (row), and SSY represents the sub-sampling dimension in the Y direction (column). For example, if SSX = SSY = 1, there is no subsampling and the process is exactly the same as in the other embodiments of the invention described above. Similarly, if SSX = SSY = 1 as in the present embodiment, the process returns to “light mosaic” without using pixel averaging. If there are three SSXs and SSYs and the active area of the circle is 500 × 500, a 166 × 166 aperture is scanned, 1/3 in the x direction and 1/3 in the y direction, resulting data Are reduced by a factor of nine. If 1/9 aperture is used all the time, these calculations are not necessary and the collimation grid need not be included.
Therefore, only 1 / (SSX * SSY) aperture (500 × 500 aperture) in the original collimation grid needs to be used to create the image, ie, the electron beam for x-ray irradiation Need to be irradiated. If the frame rate is kept constant, ie, 30 Hz, the number of electron beam motions is reduced by SSX * SSY, such as the frequency response of the circuit driving the electron beam. The total distance the electron beam travels (and the number of scan lines) is reduced by 1 / SSY, so that the average beam velocity through the target anode is reduced by 1 / SSY. The pixel speed at which the image is reconstructed is the same as the collimation aperture speed (the speed at which the aperture is scanned or illuminated) and is similarly reduced by 1 / (SSX * SSY).
According to this scheme, the number of samples averaged to each display pixel is (DDT_x/ SSX) * (DDT_y/ SSY). Using the maximum sampling rate (SSX = DDT_xAnd SSY = DDT_y) Only one digitizer sample is averaged for each display pixel ("light mosaic" mode). Sample averaging is important to smooth out non-uniformities in the beam, scintillator, detector and amplifier. The subsampling magnitudes (SSX and SSY) must be set to an appropriate level for the conditions presented to ensure acceptable image quality. This can be set by the user for the fly depending on the user's preference for image quality and the conditions submitted by a particular set of environments.
The detector array 110 shown in FIG. 15 is preferably an array of 96 individual detector elements 160 arranged approximately in a circular area having a diameter of about one inch. At the center of the array, 12 detectors (DDT_x) With 12 detectors (DDTs) in the center of the array, in horizontal rows and columns_y). The scintillator crystals are preferably cut into square horizontal cross sections and supported by an "egg crate" structure consisting of a 0.005 inch thick stainless steel strip. The circle 400 of FIG. 15 with the entire scintillator crystal (hatched) positioned therein is preferably about 0.800 inches in diameter.
The length of the scintillation crystal in detector array 110 is preferably 0.10 cm, and the input surface before it is preferably 0.135 cm × 0.135 cm. The scintillation crystals are preferably YSO, LSO or BGO, but other materials can be used as described above. The BGO must be heated to about 100 ° C. for a suitably reduced decay time (50 nS) for its light output in this application. Therefore, the resistance heating element may be provided.
FIG. 17 shows a detector assembly 402 according to a preferred embodiment of the present invention. The x-ray passes from above through the x-ray window 404 and into the lead shield 406. The x-ray window 404 is preferably circular and 1.91 cm in diameter to allow the x-rays to exit the aperture of the collimation grid 90 and strike the detector array 110 while attenuating the scattered x-rays. . A light shield 408 is provided to shield the detector from stray light. Light shield 408 can be fabricated from a sheet of aluminum or beryllium selected to attenuate light without substantially attenuating x-rays.
The detector array 110 is located near a suitable heating element 410 for use with a BGO scintillator. Heating element 410 may be a resistive heating element designed to maintain detector array 110 at an operating temperature of 100 ° C. The photons emerging from the fiber optic imaging taper 412 from the bottom 414 of the detector array 110 are directed to a 96 channel photomultiplier tube (PMT) 416. The detector assembly 402 is enclosed within a light tight outer housing 418 to prevent stray light from generating noise. Three shoulder screws 420 and three centering screws 422 are provided for planar and linear arrangements, as is well known to those skilled in the art. Rotational alignment is achieved by rotating outer housing 418 relative to PMT mount 426. Fiber optic imaging taper 412 is commercially available from Collimated Halls, Inc. of Campbell, Calif., And has a 2.03 cm diameter circular input aperture and a 3.38 cm diameter circular output aperture. The taper 412 fits the pitch dimension (0.06 inch) of each scintillator crystal to the PMT 416 dimension (0.10 inch). That is, the taper has a magnification of 1.667 times. A highly viscous optically coupled liquid (type 200), commercially available from Dow Corning, has a refractive index approximately equal to the refractive index of the glass, with a scintillator crystal 160 to taper 412 and a taper 412 to PMT input surface 424. To maximize light transfer efficiency, it is used as a light coupling medium with two tapered surfaces.
The photomultiplier tube 416 is a 96-channel tube (one channel corresponding to each scintillation crystal) commercially available as type XP1724A from Philips. The photomultiplier 416 includes a fiber optic faceplate such that the spatial arrangement of the scintillation array is accurately performed on a PMT photocathode located on the PMT on the other face of the optical faceplate. An x-ray photon that hits one scintillation 160 produces a light pulse that is coupled to the PMT photocathode. This results in a corresponding electron pulse on the photocathode, which pulse is amplified up to 1,000,000 times in one channel of the PMT diode structure.
This PMT output pulse is coupled to the input of a 30 MHz band amplifier, and the output of the amplifier is in the range of 0.5-5.0 volts and is about 30 nsec long so that the pulse is differentiated. This eliminates the need for a DC restorer circuit that maintains the baseline reference voltage as the pulse rate changes.
The output of the amplifier is input to a comparator, and the comparator outputs an output pulse of a fixed size regardless of the size of the input. The reference voltage for the comparator is set to a value slightly higher than the noise output level so that the amplifier is not triggered at the noise output level. The amplifier chain is repeated 96 times, once for each scintillation crystal in the detector array. The output pulse of the comparator provides new data for the data acquisition and image reconstruction system. As tests have shown, the prototype system can randomly count x-ray photons up to a speed of about 10 MHz.
While embodiments and applications of the present invention have been described above, it will be apparent to those skilled in the art that many more variations than those described above are possible without departing from the inventive concept. Accordingly, the invention is not to be restricted except in the spirit of the appended claims.

Claims (3)

A scanning X-ray tube having an electron beam generator, an anode, and an output surface adjacent the anode;
A circuit for locating the electron beam generated by the electron beam generator at selected points on the anode, and applying the electron beam to the selected points for a predetermined time, A scan generator that stays in a fixed pattern,
A collimation grid formed of a material that absorbs X-rays and provided in parallel with and adjacent to the output surface, and having a flat input surface and a flat output surface,
The collimation grid has a plurality of openings, the plurality of openings penetrate the output surface in a first region on the output surface of the collimation grid,
An X-ray detector array having at least three separate X-ray detectors,
Each of the X-ray detectors has a detection input surface, and all of the plurality of the detection input surfaces are arranged two-dimensionally , and are located in a second area on the detection surface smaller than the first area. Included
The plurality of openings are not provided in parallel to each other and at an angle which gradually changes toward the end from the center of the collimation grid for the front of the next contact is, and the collimation grid with each other , each aperture provided with a X-ray to direct said X-ray detector array from the anode, the scanning beam digital X-ray imaging system. The scanning beam digital x-ray imaging system of claim 1 , further comprising a feedback loop for limiting the output of the scanning x-ray tube based on a plurality of measurements detected by the x-ray detector array. The scanning beam digital X-ray imaging system according to claim 1 , further comprising a controller configured to emit X-rays from a plurality of selected apertures in a predetermined pattern.

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