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WO2000062674A1 - Detecteur de topographie par ordinateur de taille reduite a demi-champ de vue - Google Patents

Detecteur de topographie par ordinateur de taille reduite a demi-champ de vue Download PDF

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Publication number
WO2000062674A1
WO2000062674A1 PCT/US2000/010170 US0010170W WO0062674A1 WO 2000062674 A1 WO2000062674 A1 WO 2000062674A1 US 0010170 W US0010170 W US 0010170W WO 0062674 A1 WO0062674 A1 WO 0062674A1
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WO
WIPO (PCT)
Prior art keywords
detector
projection
data
rays
projection data
Prior art date
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Ceased
Application number
PCT/US2000/010170
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English (en)
Inventor
Wen-Tai Lin
Ishaque N. Ahmad
Peter M. Edic
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General Electric Co
Original Assignee
General Electric Co
Priority date (The priority date is an assumption and is not a legal conclusion. Google has not performed a legal analysis and makes no representation as to the accuracy of the date listed.)
Filing date
Publication date
Application filed by General Electric Co filed Critical General Electric Co
Priority to JP2000611814A priority Critical patent/JP2002541896A/ja
Priority to DE10081367T priority patent/DE10081367T1/de
Priority to AU43524/00A priority patent/AU4352400A/en
Publication of WO2000062674A1 publication Critical patent/WO2000062674A1/fr
Anticipated expiration legal-status Critical
Ceased legal-status Critical Current

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Classifications

    • A—HUMAN NECESSITIES
    • A61—MEDICAL OR VETERINARY SCIENCE; HYGIENE
    • A61B—DIAGNOSIS; SURGERY; IDENTIFICATION
    • A61B6/00—Apparatus or devices for radiation diagnosis; Apparatus or devices for radiation diagnosis combined with radiation therapy equipment
    • A61B6/02—Arrangements for diagnosis sequentially in different planes; Stereoscopic radiation diagnosis
    • A61B6/03—Computed tomography [CT]
    • A61B6/032—Transmission computed tomography [CT]
    • A—HUMAN NECESSITIES
    • A61—MEDICAL OR VETERINARY SCIENCE; HYGIENE
    • A61B—DIAGNOSIS; SURGERY; IDENTIFICATION
    • A61B6/00—Apparatus or devices for radiation diagnosis; Apparatus or devices for radiation diagnosis combined with radiation therapy equipment
    • A61B6/02—Arrangements for diagnosis sequentially in different planes; Stereoscopic radiation diagnosis
    • A61B6/027—Arrangements for diagnosis sequentially in different planes; Stereoscopic radiation diagnosis characterised by the use of a particular data acquisition trajectory, e.g. helical or spiral

Definitions

  • the present invention relates to method and apparatus for use in a volumetric computed tomography (VCT) system that utilizes a reduced-sized area detector that covers only half of the field of view, thereby enabling the size and costs of the area detector to be reduced without increasing, or substantially increasing, artifacts.
  • VCT volumetric computed tomography
  • Computed tomography is a technique that generally involves subjecting a patient to X-rays, acquiring digital X-ray projection data of a portion of the patient's body, and processing and back-projecting the digital X- ray projection data to produce an image that is then displayed on a display monitor of the CT system.
  • CT systems typically comprise a gantry, a table, an X-ray tube, an X-ray detector array, a computer and a display monitor.
  • the computer sends commands to controllers of the gantry to cause the gantry to rotate the X-ray tube and/or the detector array at a particular rotational speed.
  • relative rotational motion is produced between the gantry, partly comprised by the detector array and the X-ray tube, and the patient's body.
  • the computer controls the data acquisition process performed by the X-ray tube and the detector array to acquire digital X-ray radiographs.
  • the computer then processes and back-projects the digital X-ray radiograph data by performing a reconstruction algorithm and displays the reconstructed CT image on the display monitor.
  • CT systems in use today utilize a single row of detectors in the gantry, which is normally referred to as a linear array of detector elements. More advanced CT systems use two to four linear arrays of detectors to construct a multi-row detector. Although both detector arrangements can be used with a helical scanning protocol, the multi-row detector facilitates patient scanning since a specified axial coverage of the patient can be scanned in less time by increasing the helical pitch of the detector array.
  • Helical pitch is typically defined as the ratio of the displacement of the table supporting the patient during one rotation of the gantry to the detector pitch. For example, a helical pitch of one refers to translating the patient table an amount equal to the detector pitch during one revolution of the CT gantry of the CT system.
  • the linear detector or the multi-row detector arrays cover the entire field of view of the X-ray fan beam emitted by the X-ray source.
  • the X-rays that pass through or illuminate the area of the object being scanned, which may or may not be a patient, are absorbed by the detector array.
  • CT imaging systems it is desirable and may, in some instances, be necessary to reduce the size of the detector array.
  • area detector arrays composed of many rows of linear detector arrays, for CT data acquisition.
  • a detector panel that can cover the entire field of view or extent of the patient being imaged is not yet available.
  • some systems that use linear detector arrays support a very large field of view for the patient being scanned. It is desirable to reduce the size and costs of the detector array in this situation as well.
  • One technique that has been utilized to overcome these limitations is to translate a detector array that is smaller in size by half of the width of the detector array.
  • the original size of the detector array should be 80 cm to cover the desired field of view of the patient.
  • a smaller detector that may have a width equal to half of the width of the original detector, 40 cm in this case, may be used.
  • This detector is offset by half of its width, or 20 cm in this example, such that it covers approximately one half of the field of view of the CT imaging system.
  • the same field of view in the patient is obtained by a detector that has a width equal to half of its original value.
  • the field of view of a system that has a detector with a fixed width.
  • the projection of the center of rotation of the CT imaging system is aligned with the center of the detector panel.
  • the center of rotation in a CT imaging is the physical location of the point which the X-ray source and the detector array rotate about.
  • the field of view (FOV) of this system can be increased by offsetting the detector by half of its width with respect to its original position.
  • the projection of the center of rotation of the imaging system is near an edge of the linear or multi-row detector that has been shifted, although the detector still measures projection data from X-rays that pass through the physical center of rotation (i.e., at the ISO) of the imaging system.
  • This arrangement effectively doubles the field of view of the original imaging system configuration, which can result in a large increase in the field of view of the imaging system.
  • the system configuration of shifting the detector by half of its width is typically referred to half-detector shift.
  • a fan-beam CT system a CT system where the X-ray source is a point which emits an angular aperture of X-rays that illuminate only the detector panel and resemble a fan
  • the gantry rotates a angular region about the patient equaling 180° plus the fan angle.
  • the fan angle is the measure of the angular aperture of X-rays that illuminate only the detector array in the axial plane of the imaging system.
  • some of the projection data must be redundant.
  • a technique that is currently being used to mitigate the artifacts that are caused by the discontinuities of projection data within the field of view utilizes a weighting function to smooth the discontinuities in the data within a transition region.
  • This technique requires that the detector has some additional detector elements that extend past the projection of the center of rotation of the imaging system on the detector. As the gantry rotates 360° about the patient, the region of the shifted detector panel that extends slightly past the projection of the center of rotation on the detector in both directions is called the transition region. Actual data is measured by the detector in half of the transition region, and data can be generated from alternate views of the gantry in the second half of the transition regionData within the transition region are multiplied by a weighting factor that smoothes the discontinuities.
  • a computed tomography (CT) system for obtaining projection data of an object, which comprises an X-ray source, and a detector.
  • the detector is shifted by half of its width relative to a center position which corresponds to the projection of a center of rotation of the CT system onto the detector.
  • a detector element value, Va is chosen of a detector element that is closest to the an ISO center of the CT system.
  • a value, Nb for the detector element is estimated from an opposing direction or from a forward projection in the same direction.
  • a smoothing function capable of eliminating a discrepancy between Va and Vb is then selected.
  • the smoothing function is then applied to eliminate the discrepancy between Va and Vb.
  • a weighting function is then applied to pave the step difference when true projection data and estimated projection data are combined to create a smooth transition region.
  • Fig. 1 is a block diagram of the CT system of the present invention.
  • Fig. 2 illustrates the detector offset utilized in accordance with the method of the present invention.
  • Fig. 3 is a block diagram illustrating the method of the present invention in accordance with the preferred embodiment.
  • FIG. 1 is a block diagram of a volumetric CT scanning system that is suitable for implementing the method and apparatus of the present invention.
  • the Volumetric CT scanning system will be discussed with respect to its use in reconstructing an image of an anatomical feature of a patient, although it will be understood that the present invention is not limited to imaging any particular object.
  • the present invention may also be used for industrial processes, as will be understood by those skilled in the art.
  • the present invention is not limited to medical CT equipment, but includes industrial systems where the X-ray source and detector geometry are held fixed while the object is rotated during the scanning time.
  • a volumetric CT scanning system the gantry is rotated about an object, such as a human patient, and projection data are acquired.
  • a computer 1 controls the operations of the volumetric CT scanning system.
  • that phrase is intended to denote rotation of the X-ray tube 2 and/or rotation of the detector 3. which preferably is a high resolution area detector.
  • the X-ray tube 2 and the area detector 3 are comprised by the gantry.
  • the controllers 4A and 4B are controlled by the Volumetric CT scanning system computer 1 and are coupled to the X-ray tube 2 and to the detector 3, respectively.
  • the controllers 4A and 4B cause the appropriate relative rotational motion to be imparted to the X-ray tube 2 and/or to the detector 3. Individual controllers are not necessary.
  • a single controller component may be used to rotate the gantry.
  • the computer 1 controls the variations in image scanning time, image resolution and/or axial coverage in order to implement the methods of the present invention.
  • the computer 1 controls the data acquisition process by instructing the data acquisition system 6 as to when to sample the detector 3 and by controlling the speed of the gantry. Additionally, the computer 1 instructs the data acquisition system 6 to configure the resolution of the radiographs obtained by the area detector 3, thereby allowing the resolution of the system to be varied.
  • the data acquisition system 6 comprises the read out electronics, as shown.
  • the area detector 3 is comprised of an array of detector elements (not shown). Each detector element measures an intensity value associated therewith that is related to the amount of X-ray energy that impinges on the detector element.
  • the present invention is not limited to any particular computer for performing the data acquisition and processing tasks of the present invention.
  • the term "computer”, as that term is used herein, is intended to denote any machine capable of performing the calculations, or computations, necessary to perform the tasks of the present invention. Therefore, the computer utilized to perform the control algorithm 10 of the present invention may be any machine that is capable of performing the necessary tasks
  • an alternate method utilizes an iterative algorithm to estimate the projection data that would have been measured if a large detector array that covered the entire field of view had been used to acquire the data. Both approaches will be discussed within the same context below since errors in the transition region are handle in a similar manner.
  • This technique forms a set of X-ray projections ⁇ P by either forward projecting the reconstructed data obtained from a prior iteration step or by interpolating the redundant detector data obtained from a set of opposing rays, which form another set of projection data in the reversed directions of ⁇ P a ⁇ -
  • the technique of forward projection is the process where rays are emitted from a hypothetical X-ray source; these rays traverse the reconstructed volume toward individual detector elements. Along the ray, the linear attenuation value of the reconstructed values along the rays are summed and denoted as the line integral of the line attenuation coefficient.
  • the technique of forward projecting the reconstructed data (denoted as
  • FPT is generally suitable for generating projection data corresponding to a larger cone angle (i.e., as is used in VCT systems) while the technique of interpolating redundant projection data(denoted as PDT) is more suitable for projection data that is closer to the mid-plane (i.e., that is closer to the ISO center).
  • the cone angle refers to the angular extent of the X-ray emitted from the X-ray source in a direction orthogonal to the fan angle direction.
  • the discrepancy between the estimated detector values obtained using the either FPT or PDT and the original values that would be obtained if the data had actually been measured may cause distortions in images close to the ISO center.
  • a technique has been developed that utilizes a smoothing function, which can be stated as follows with reference to
  • Figure 3 1. For each projection view, choose the known detector element that is closest to the ISO center 21 , which will be referred to hereinafter as Va.
  • Vb an estimated value 22 (i.e., via interpolation projection data from alternate views(PDT) or from forward projecting the reconstructed data(FPT)), which will be referred to hereinafter as Vb.
  • a smoothing function reduces the discrepancy between Va and Vb and gradually smoothes this within a region near the center of the field of view of the imaging system 24.
  • Equationl Equationl
  • x 0 is the absolute value of the distance from the detector location corresponding to the detector element value Va and a is a factor controlling the slope of the curve associated with the smoothing function.
  • the exponential function is added/subtracted to projection values located to one side of the central ray location(the detector location corresponding to the projection of the center of rotation in the imaging system onto the detector) to bring the lower/higher estimated values up/down and is subtracted/added to projection values to the alternate side of the central ray location to reduce/increase the higher/lower original values.
  • VCT Volumetric CT
  • each ray is a shallow tetrahedron originating at the source and ending at the detector element. There are no identical rays traversing exactly the same part of the object unless the object is totally homogeneous and circularly symmetrical.
  • the process of interpolation may introduce errors.
  • d( ⁇ ) is the difference between f ⁇ (N 2 ) and f ⁇ -(N 2 ) at the angular source position ⁇ . If d( ⁇ ) is totally random, then an error introduced in the reconstructed image probably will be obscured by the quantum noise associated with other CT random errors. However, if the error is somewhat systematic, it will introduce obvious artifacts in the reconstructed image. For this reason, a smoothing process must be utilized at the transition region. In other words , the smoothing function was developed to make the step error of f ⁇ and f ⁇ - smaller around the central ray location denoted by detector element N 2 .
  • W and W ' will be used to represent the smoothing function for f ⁇ and f fi -, respectively.
  • certain criteria must be considered, which will be understood by those skilled in the art.
  • smoothing functions are suitable for this purpose, other than those specifically set forth herein, as will be understood by those skilled in the art. For example,
  • ⁇ n sets the extent of smoothness of W and W and ⁇ is a differential operator.
  • the conventional smoothing function for this type of application is also generally known as a feathering function.
  • W and W' are used to find a transition region between the forward and opposing rays.
  • ⁇ n must be an integer larger than zero. In fact, the larger ⁇ n is, the better the smoothing function will work.
  • increasing ⁇ n too much could require that additional detector elements be used to extend beyond N 2 , which is the detector element at the central ray location. Therefore, ⁇ n should be selected so that it is sufficiently large, but not so large as to require that additional detector element be added to the transition region.
  • the actual detector signals being used in the back projection process are Wf ⁇ and Wf ⁇ .
  • the incentive is that when one starts to increase the number or detector rows in a CT system, i.e. in an area detector, the additional detector elements ( ⁇ n times the number of rows) becomes larger.
  • ⁇ n the number of detector elements
  • d( ⁇ ) is viewed as a step error between f ⁇ ( N 2 ) and f ⁇ (N 2 ) where ⁇ is the angular position of the x-ray source.
  • Equation 1 a is a control factor controlling the smoothness of the exponential function.
  • the forward and opposing ray functions, f ⁇ (n) and f ⁇ (n) are converted into two other respective functions.
  • the following procedure is implemented: 1. Obtain the original half FOV projection data calling it f ⁇ (n), with zero- padding according to Equation 2.
  • the present invention has been discussed with respect to certain embodiments. However, the present invention is not limited to these embodiments.
  • the three scenarios discussed are not meant to be all inclusive of the manner in which tradeoffs of the aforementioned parameters can be utilized to obtain the proper mode of operation of the VCT system. These scenarios are discussed in order to illustrate the concepts of the present invention and the manner in which these fundamental parameters can be traded off in order to achieve the proper scanning protocol.
  • the trade-offs are not limited to a scanning protocol, i.e. they apply to both axial scanning (the patient table is not moved during the scanning period) and helical scanning protocols. Those skilled in the art will understand the manner in which these concepts can be utilized and extrapolated to achieve other area detector scanning protocols that are useful for particular applications.

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  • Health & Medical Sciences (AREA)
  • Life Sciences & Earth Sciences (AREA)
  • Engineering & Computer Science (AREA)
  • Medical Informatics (AREA)
  • Radiology & Medical Imaging (AREA)
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  • High Energy & Nuclear Physics (AREA)
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  • Optics & Photonics (AREA)
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  • General Health & Medical Sciences (AREA)
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Abstract

L'invention concerne un système de topographie par ordinateur, dont le détecteur est déplacé de la moitié de sa largeur par rapport au centre ISO, ce qui produit un champ de projection, Va (21). Vb (22) est évalué à partir de la direction opposée ou d'une projection avancée. Lissage (23, 24) et pesage (25) sont utilisés pour éliminer la divergence entre Va et Vb.
PCT/US2000/010170 1999-04-15 2000-04-14 Detecteur de topographie par ordinateur de taille reduite a demi-champ de vue Ceased WO2000062674A1 (fr)

Priority Applications (3)

Application Number Priority Date Filing Date Title
JP2000611814A JP2002541896A (ja) 1999-04-15 2000-04-14 半撮影域のみをカバーする縮小サイズ検出器を利用するコンピュータ断層撮影システムに用いる装置及び方法
DE10081367T DE10081367T1 (de) 1999-04-15 2000-04-14 Verfahren und Gerät zur Verwendung bei einem Computertomographiesystem, das eine verkleinerte Erfassungsvorrichtung für die Abdeckung lediglich des halben Sichtfeldes verwendet
AU43524/00A AU4352400A (en) 1999-04-15 2000-04-14 Half field of view reduced-size ct detector

Applications Claiming Priority (4)

Application Number Priority Date Filing Date Title
US12939899P 1999-04-15 1999-04-15
US60/129,398 1999-04-15
US16650099P 1999-11-19 1999-11-19
US60/166,500 1999-11-19

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AU (1) AU4352400A (fr)
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WO (1) WO2000062674A1 (fr)

Cited By (9)

* Cited by examiner, † Cited by third party
Publication number Priority date Publication date Assignee Title
FR2820045A1 (fr) 2001-01-29 2002-08-02 Chabunda Christophe Mwanza Dispositif bilina-imatron de stereoradiotherapie par acceleration et collision des particules des 2 faisceaux d'irriadiation synchrones associe a un dispositif de verification instantannee de delivrance des radiations
WO2004000123A3 (fr) * 2002-06-25 2004-03-25 Univ Michigan Systeme de tomographie par ordinateur (ct) par rayons x a resolution spatiale elevee
WO2006013325A1 (fr) * 2004-08-05 2006-02-09 Elekta Ab (Publ) Appareil à rayons x
WO2007020318A3 (fr) * 2005-08-17 2007-04-12 Palodex Group Oy Appareil d'imagerie par rayons x et procede d'imagerie par rayons x
WO2010084389A1 (fr) * 2009-01-21 2010-07-29 Koninklijke Philips Electronics N.V. Procédé et appareil d'imagerie à grand champ de vision et détection et compensation d'artéfacts de mouvement
US7945012B2 (en) 2006-08-17 2011-05-17 Koninklijke Philips Electronics N.V. Computed tomography image acquisition
US8908953B2 (en) 2007-06-11 2014-12-09 Koninklijke Philips N.V. Imaging system and imaging method for imaging a region of interest
US8934602B2 (en) 2007-01-24 2015-01-13 Dental Imaging Technologies Corporation Adjustable scanner
US10426424B2 (en) 2017-11-21 2019-10-01 General Electric Company System and method for generating and performing imaging protocol simulations

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JP4820561B2 (ja) * 2005-03-14 2011-11-24 株式会社東芝 核医学診断装置
CN100435733C (zh) * 2005-12-31 2008-11-26 清华大学 X-ct扫描系统
JP5514195B2 (ja) * 2008-05-09 2014-06-04 コーニンクレッカ フィリップス エヌ ヴェ 運動する物体の画像を生成するための装置
CN104545976B (zh) * 2014-12-30 2017-04-19 上海优益基医疗器械有限公司 计算机体层摄影方法和装置
WO2018115160A1 (fr) * 2016-12-21 2018-06-28 Koninklijke Philips N.V. Reconstruction de radiographie mise à l'échelle
CN107714072B (zh) * 2017-11-20 2019-09-20 中国科学院高能物理研究所 缺失数据的补偿方法、计算机断层扫描成像方法及系统
CN111374693A (zh) * 2018-12-27 2020-07-07 有方（合肥）医疗科技有限公司 口腔ct
CN117257340A (zh) * 2023-11-21 2023-12-22 北京朗视仪器股份有限公司 一种异形探测器、医疗影像设备以及图像补全方法

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JP3637074B2 (ja) * 1992-12-15 2005-04-06 株式会社東芝 ヘリカルスキャン方式のコンピュータ断層撮影装置
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JPH10290798A (ja) * 1997-04-17 1998-11-04 Ge Yokogawa Medical Syst Ltd 投影データ測定方法および装置並びにｘ線ｃｔ装置
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Cited By (13)

* Cited by examiner, † Cited by third party
Publication number Priority date Publication date Assignee Title
FR2820045A1 (fr) 2001-01-29 2002-08-02 Chabunda Christophe Mwanza Dispositif bilina-imatron de stereoradiotherapie par acceleration et collision des particules des 2 faisceaux d'irriadiation synchrones associe a un dispositif de verification instantannee de delivrance des radiations
WO2004000123A3 (fr) * 2002-06-25 2004-03-25 Univ Michigan Systeme de tomographie par ordinateur (ct) par rayons x a resolution spatiale elevee
US7099428B2 (en) 2002-06-25 2006-08-29 The Regents Of The University Of Michigan High spatial resolution X-ray computed tomography (CT) system
US7460637B2 (en) 2002-06-25 2008-12-02 The Regents Of The University Of Michigan High spatial resolution X-ray computed tomography (CT) method and system
WO2006013325A1 (fr) * 2004-08-05 2006-02-09 Elekta Ab (Publ) Appareil à rayons x
US7796728B2 (en) 2004-08-05 2010-09-14 Elekta Ab (Publ) X-ray apparatus
WO2007020318A3 (fr) * 2005-08-17 2007-04-12 Palodex Group Oy Appareil d'imagerie par rayons x et procede d'imagerie par rayons x
US7945012B2 (en) 2006-08-17 2011-05-17 Koninklijke Philips Electronics N.V. Computed tomography image acquisition
US8934602B2 (en) 2007-01-24 2015-01-13 Dental Imaging Technologies Corporation Adjustable scanner
US8908953B2 (en) 2007-06-11 2014-12-09 Koninklijke Philips N.V. Imaging system and imaging method for imaging a region of interest
WO2010084389A1 (fr) * 2009-01-21 2010-07-29 Koninklijke Philips Electronics N.V. Procédé et appareil d'imagerie à grand champ de vision et détection et compensation d'artéfacts de mouvement
US9710936B2 (en) 2009-01-21 2017-07-18 Koninklijke Philips N.V. Method and apparatus for large field of view imaging and detection and compensation of motion artifacts
US10426424B2 (en) 2017-11-21 2019-10-01 General Electric Company System and method for generating and performing imaging protocol simulations

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DE10081367T1 (de) 2001-06-28
JP2002541896A (ja) 2002-12-10
JP2011031070A (ja) 2011-02-17
AU4352400A (en) 2000-11-02
CN1300201A (zh) 2001-06-20
JP5194095B2 (ja) 2013-05-08
CN1210000C (zh) 2005-07-13

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