CN1300201A - Half field of view reduced CT detector - Google Patents

Half field of view reduced CT detector Download PDF


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CN1300201A CN 00800601 CN00800601A CN1300201A CN 1300201 A CN1300201 A CN 1300201A CN 00800601 CN00800601 CN 00800601 CN 00800601 A CN00800601 A CN 00800601A CN 1300201 A CN1300201 A CN 1300201A
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CN1210000C (en
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    • A61B6/00Apparatus for radiation diagnosis, e.g. combined with radiation therapy equipment
    • A61B6/02Devices for diagnosis sequentially in different planes; Stereoscopic radiation diagnosis
    • A61B6/03Computerised tomographs
    • A61B6/032Transmission computed tomography [CT]
    • A61B6/00Apparatus for radiation diagnosis, e.g. combined with radiation therapy equipment
    • A61B6/02Devices for diagnosis sequentially in different planes; Stereoscopic radiation diagnosis
    • A61B6/027Devices for diagnosis sequentially in different planes; Stereoscopic radiation diagnosis characterised by the use of a particular data acquisition trajectory, e.g. helical or spiral


一种CT系统,具有相对于ISO中心移动了其宽度的一半的检测器,产生投影视图Va(21),从反方向中估计或从正向投影中估计Vb(22)。 A CT system, with respect to the ISO center movement detector half its width, to produce projection views Va (21), estimation or estimation Vb (22) from the forward projecting from the opposite direction. 应用平滑步骤(23,24)和加权步骤(25)来消除在Va和Vb之间的差。 Applying a smoothing step and the weighting step (25) (23, 24) to eliminate the difference between Va and Vb.


尺寸减小的半视场CT检测器 Size reduction of the half field detector CT

本申请要求分别以在1999年4月15日和1999年11月19日申请的申请号为No.60/129,398和60/166,500的临时申请的申请日为优先权,在此以引用的方式将它们全部结合在本申请中。 This application claims to Application No. 1999 April 15, 1999 and November 19 to apply for the No.60 / 129,398 and filing date of provisional application 60 / 166,500 is a priority in this by reference to all of which are incorporated in the present application.

本发明涉及一种应用在体积型计算机X-射线断层成像(VCT)系统中应用的方法和装置,该系统应用一种尺寸减小的面积检测器,这种面积检测器仅覆盖一半的视场,由此降低这种面积检测器的尺寸和成本但不增加或基本不增加假象。 The present invention relates to a method and apparatus for field application of a volume application X- ray computed tomography (VCT) system, the system having a reduced size of the application area detector, this detector area covers only half of the , this area thereby reducing the size and cost of the detector without increasing or without substantially increasing the artifacts.

计算机X-射线断层成像(CT)是这样的一种技术:一般包含对患者进行X-射线辐射、采集患者的部分身体的数字X-射线投影数据以及处理和背式投影数字X-射线投影数据以产生图象,然后将该图象显示在CT系统的监视器上。 X- ray computer tomography (CT) is a technique: the patient generally comprise X- ray radiation, collecting part of the body of the digital X- ray projection data, and processing and back-projection digital X- ray projection data of the patient to produce an image, then the image is displayed on the monitor of the CT system. CT系统通常包括台架、工作台、X-射线管、X-射线检测器阵列、计算机和显示检测器。 CT systems typically comprise a gantry, a table, X- ray tube, X- ray detector array, a computer display and detector. 计算机给台架的控制器发送指令以使台架使X-射线管和/或检测器阵列以特定的转速转动。 The computer sends commands to controllers of the gantry so that the gantry X- ray tube and / or the detector array at a particular rotational speed.

在第三代CT系统中,在检测器阵列和X-射线管部分地包括的台架和患者的身体之间产生相对转动。 In third generation CT systems, the patient is generated between the carriage body and X- ray detector array comprises a tube portion and a relative rotation. 由于产生这种相对转动,计算机控制通过X-射线管和检测器阵列执行的数据采集过程以采集数字X-射线照相。 Because of this relative rotation is generated, the computer controls the data acquisition process performed by the X- ray tube and the detector array to acquire digital X- ray radiography. 然后计算机进行处理并通过执行重构算法背向投影数字X-射线照相数据,并在显示监视器上显示所重构的CT图像。 The computer then processes and back-X- digital radiographic projection data by performing a reconstruction algorithm and displays the reconstructed CT image on the display monitor.

如今所应用的许多CT系统都利用在台架中的单行检测器,这种单行的检测器通常称为检测器元件的线性阵列。 Many CT systems today utilize a single row of the applied detector gantry, which is generally referred to as a single row of detector elements is a linear array detector. 更先进的CT系统应用两至四个线性检测器阵列以构成多行检测器。 More advanced CT systems use two to four linear detector arrays to form a multi-row detector. 虽然这两者检测器结构都可以用于螺旋扫描方案,但是由于通过增加检测器阵列的螺旋间距多行检测器能够在更少的时间中扫描患者特定的轴线区域,所以它有利于患者扫描。 Although both the detector structure may be used for a helical scanning scheme, but because by increasing the helical pitch of the detector array can be a multi-row detector scan area of ​​the patient in a particular axis in less time, so it is beneficial to the patient scan. 螺旋间距通常定义为在台架旋转一圈中支撑患者的工作台的位移与检测器间距之比。 The helical pitch is generally defined as the ratio of one rotation of the gantry supporting a patient table with a displacement of the detector spacing. 例如,一个螺旋间距是指在CT系统的CT台架的旋转一圈中将患者工作台移动等于检测器间距的量。 For example, a helical pitch is the amount of rotation of the table moves the patient in the CT gantry in a CT system is equal to the detector pitch.

通常,线性检测器或多行检测器阵列覆盖由X-射线源发射的X-射线扇形束的整个视场。 Typically, one or more linear detector arrays cover the entire row detector field of view X- ray fan beam emitted from the X- ray source. 换句话说,通过检测器阵列吸收穿过或照射所扫描的对象的面积的X-射线,该对象可能是或不是患者。 In other words, the scanned area illuminated through or absorbed by the detector array X- ray object, the object may be a patient or not.

在CT成像系统中,比较理想的是并且在某些情况下也是必需的是减小检测器阵列的尺寸。 In CT imaging systems, and is ideal in some cases it is also necessary to reduce the size of the detector array. 例如,在新近发展的CT技术中应用包括许多行线性检测器阵列的面积检测器阵列进行CT数据采集。 For example, application area including a plurality of rows of the detector array is a linear array of detectors recent development in CT technology in the CT data acquisition. 当前,仍然还没有能够覆盖整个成像的视场或患者范围的检测器面板。 Current, yet still be able to cover the entire field of view of the imaging or patients with a range of detector panel. 此外,一些应用线性检测器阵列的系统支持对于所扫描的患者的很大的视场。 Further, some applications of the system of linear detector arrays support a large field of view for the patient being scanned. 理想的是在这种情况下也减小检测器阵列的大小和成本。 It is desirable in this case also to reduce the size and cost of the detector array.

用于克服这些局限性的一种方法是将更小的检测器阵列平移其宽度的一半。 For overcoming these limitations is a process will be smaller detector array translate half of its width. 例如,假设为覆盖患者的所需的视场的检测器阵列的最初的尺寸应该是80厘米。 For example, assuming a desired initial size of the detector array to cover the patient's field of view should be 80 cm. 可以应用等于最初检测器的宽度一半的更小的检测器,即在这种情况为40厘米。 It may be applied equal to half the width of the first detector smaller detector, i.e. in this case 40 cm. 这种检测器偏移它的一半的宽度(在这种情况为20厘米)以使它大致覆盖CT成像系统的视场的一半。 This detection offset half its width (20 cm in this case) so that it covers substantially half of the field of view of a CT imaging system. 在本实例中,通过宽度等于它的最初宽度值的一半的检测器获得了在患者上的相同的视场。 In the present example, it is obtained in the same field of view on the patient by a width equal to half the width of the detector to its original value.

还能够增加具有固定宽度的检测器的系统的视场。 Also possible to increase the field of view of the detector system having a fixed width. 通常,CT成像系统的中心旋转的投影与检测器面板的中心对准。 Typically, the center of rotation of a CT imaging system and the detection center of the projection panel is aligned. 在CT成像中的旋转中心是X-射线源和检测器阵列绕其旋转的点的物理位置。 Center of rotation in a CT imaging is the physical location of the X- ray source and the detector array about the point of rotation. 然而,通过使检测器相对于它的原始位置偏移它的一半宽度能够增加这种系统的视场(FOV)。 However, by the detector with respect to its original position shifted half the width of its field of view can be increased such systems (FOV). 虽然检测器仍然测量穿过成像系统的物理旋转中心(即,在ISO)的X-射线的投影数据,但是成像系统的旋转中心的投影是在已经平移的线性或多行检测器的边沿附近。 While still detector for measuring physical center of rotation passing through the imaging system (i.e., ISO) X- ray projection data, but the projection of the center of rotation in the imaging system is near an edge has a linear translational or more rows of detectors. 这种结构反过来又有效地使原始成像系统结构的视场加倍,这就能够较大地增加成像系统的视场。 This structure, in turn, effectively making the field of view of the original structure of the imaging system doubled, which can greatly increase the field of view of the imaging system. 将检测器移动它的一半宽度的系统结构通常称为半检测器移动。 Which move the detector system configuration is commonly referred to as half the width of the half-detector shift.

在扇形束CT系统中,在该CT系统中的X射线源是辐射具有孔径张角的X-射线的点,该具有孔径张角的X-射线仅辐射检测器面板并类似于扇形,需要采集CT台架的整个旋转的一部分旋转的投影数据。 In the fan-beam CT systems, X-ray source in the CT system is a point X- ray radiation with the angular aperture, the angular aperture of having only the X- ray detector panel and is similar to the radiation fan, to be collected part of the rotatable CT gantry rotation of projection data. 具体地说,需要在台架绕患者旋转180°加上该扇形角的角度区的同时采集投影数据。 Specifically, patient rotation angle of 180 ° plus fan angle region of the projection data are acquired while the gantry about the need. 在再一次测量中,扇形角X-射线的孔径张角的度量,具有该孔径张角的X-射线仅辐射在成像系统的轴向平面中的检测器阵列。 Again the measurements, the angular aperture of the fan angle measure X- rays, the angular aperture of having only the X- ray radiation detector array in the axial plane of the imaging system. 可以清楚地看到,由于不需要在台架绕患者旋转的整个360°的过程中测量投影,一些投影数据必定是冗余的。 Can be clearly seen, it is not necessary during the rotation of the gantry about the patient's entire 360 ​​° of projection measurements, some of the projection data must be redundant.

在CT系统的半检测器移动结构中,在台架的整个360°旋转中采集数据。 In the half-detector shift configuration of the CT system, data collection throughout the 360 ​​° rotation of the gantry. 在每个视角或台架上,仅测量一半的投影数据。 In each view angle or the gantry, only one half of the measured projection data. 应用来自台架的其它视图的数据来完成在给定的视角上的投影数据。 Application Data from other views of the gantry to complete the projection data on a given viewing angle. 在本领域中完成这种处理的方法是公知的。 The method to accomplish this in the present process are well known in the art. 然而,当将所测量的覆盖成像系统的视场的一半的投影数据与从其它的台架的视图中产生的数据相结合时,所得到的投影数据并不能与投影数据的中心附近相匹配。 However, when the measured half of the field of view of projection data covering the imaging system is combined with the data generated from other views of the gantry, the projection data is obtained does not match the vicinity of the center of projection data. 如果没有降低或消除这些不匹配的话,它将在所重构的图像中产生不希望的假象。 If not reduce or eliminate these do not match, it will produce undesirable artifacts in the reconstructed image.

当前应用减小由在视场中的投影数据的不连续性引起的假象的一种技术是利用加权函数来平滑在过渡区中的数据的不连续性。 Current application technique of reducing artifacts discontinuity in the projection data caused by the field of view is to use a weighting function to smooth the transition region in the data discontinuity. 这种技术要求检测器具有额外的检测器元件,这些额外的检测器元件延伸通过成像系统的旋转中心在检测器上的投影。 This technique requires an additional detector having detector elements, these additional detector elements of the imaging system extends through the center of rotation projected onto the detector. 由于台架绕患者旋转360°,在两方向上稍稍移动延伸通过旋转中心在检测器上的投影的检测器面板区称为过渡区。 Since the 360 ​​° rotation of the gantry about the patient, extending slightly shifted detector panel zone projected on the detector is called a transition zone in the center of rotation by the two directions. 实际的数据是通过检测器在一半过渡区中测量的,以及从台架的可替换视图中产生第二半过渡区中的数据。 The actual data is measured by the detector in half of the transition region, and generating data of a second half in the transition region from an alternative view of the gantry. 对在过渡区中的数据乘以用于平滑不连续性的加权系数。 In the transition region of the data is multiplied by a weighting coefficient smoothing discontinuities. 通常,较大的过渡区产生更好的图像质量,但是由于这种系统的结构的视场梢小于比应用半检测器移动结构所产生的视场,所以它还导致了更高的系统成本。 In general, larger transition region results in better image quality, but since the field of view of the tip is less than half-detector shift configuration than the application of such a system configuration of a generated field of view, it also leads to higher system cost.

需要改善对所测量的数据和在过渡区内所产生的数据的完整性,以便能够实现检测器阵列的完整视场。 A need to improve data integrity, and the measured data generated by the transition region in order to achieve full field of view of the detector array.

在应用半检测器移动结构的体积型CT系统中,在成像系统的一半视场中测量投影数据,同时另一半投影照相数据必需从反向的射线中产生。 In a volumetric CT system configuration of a semi-moving vessel detected, measured projection data in a half field of the imaging system, while the other half of the data necessary to create a photographic projection of rays from the reverse. 不幸的是,如果检测器是它的最初的宽度的两倍并且没有偏移,则在CT台架的其它的投影角上测量的投影数据与已经测量的射线方向不具有相同的方向。 Unfortunately, if the detector is twice its original width and not offset, then the projection data measured at other projection angles of the CT gantry of the radiation direction has been measured does not have the same direction. 因此,人们需要一种VCT系统,这种VCT系统应用在半检测器移动结构中的面积检测器并实现其优点,由此克服前述的困难。 Thus, the need exists for a VCT system, this VCT system application area detector in a half-detector shift configuration and achieve the advantage, thereby overcoming the aforementioned difficulties.

一种获得对象的投影数据的计算机X-射线断层成像(CT)系统,包括X-射线源和检测器。 X- ray computer tomography for obtaining projection data of an object (CT) system, comprising an X- ray source and detector. 检测器相对于中心位置移动它的宽度的一半,该中心位置对应于CT系统的旋转中心在检测器上的投影。 A detector moved relative to half the width of its central position, which corresponds to the center position of the rotation center of the CT system is projected on the detector. 依据本发明的方法,对于每个投影视图,检测器元件值Va选择最靠近CT系统的ISO中心的检测器元件。 The method according to the present invention, for each projection view, a detector element value Va ISO selective detector element closest to the center of the CT system. 然后,对于所选择的检测器元件,检测器元件值Vb是从相同方向的正向投影或相反的方向中估计的。 Then, for the chosen detector element, the detector element is estimated from the value Vb forward projected in the same direction or opposite directions in. 然后选择能够消除Va和Vb之差的平滑函数。 Then select the smoothing function capable of eliminating a difference of Va and Vb. 然后应用该平滑函数来消除a和Vb之差。 The smoothing function is then applied to eliminate the difference between a and Vb. 然后在将真实的投影数据和所估计的投影数据结合时应用加权函数来消除幅值差以产生平滑的过渡区。 Then the weighting function is applied when the actual projection data and estimated projection data are combined to eliminate amplitude difference to create a smooth transition region.

附图1所示为本发明的CT系统方块图。 BRIEF shown a block diagram of the CT system of the present invention.

附图2所示为依据本发明的方法应用的检测器偏移。 2 shown in the drawings is an offset method according to the invention is applied to the detector.

附图3所示为依据优选的实施例说明本发明的方法的方块图。 BRIEF is a block diagram illustrating the method of the present invention according to a preferred embodiment shown in FIG.

在描述本发明的方法和装置之前,参考附图1总体上讨论一下本发明的VCT系统。 Before describing the method and apparatus of the present invention, with reference to the drawings 1 generally discuss the VCT system of the present invention. 附图1所示为适合于实施本发明的方法和装置的体积型CT扫描系统的方块图。 BRIEF suitable for volumetric CT scanning system embodiment of the method and apparatus of the present invention is shown in a block diagram. 虽然可以理解的是本发明并不限于对任何特定对象的成像,但仍然结合体积型CT扫描系统在重构患者的解剖学特征的图像的应用中进行讨论。 While it is understood that the present invention is not limited to imaging any particular object, but still binding volumetric CT scanning system is discussed in the application of the reconstructed image of the anatomical features of the patient. 正如本领域的熟练技术人员将理解的是,本发明还可以用于工业过程。 As the person skilled in the art will appreciate that the present invention may also be used in industrial processes. 此外,本发明并不限于医用CT设备,而是包括工业系统,在这些工业系统中X-射线源和检测器结构都保持固定同时在扫描时间中绕对象旋转。 Further, the present invention is not limited to medical CT equipment, but includes industrial systems, industrial systems in which X- ray source and detector structures are held stationary while rotating about the object in the scan time.

在体积型CT扫描系统中,台架绕对象比如病人旋转,并采集投影数据。 In a volumetric CT scanning system, the gantry rotating around an object such as a patient, and projection data is acquired. 计算机1控制体积型CT扫描系统的运行。 The computer 1 controls the operation of a volumetric CT scanning system. 当在此称台架的旋转时,该术语是指X-射线管2的旋转和/或检测器3的旋转,可取的是该检测器3是一种较高分辨率的面积检测器。 When this known rotating gantry, the term refers to rotating the rotary X- ray tube 2 and / or the detector 3, it is desirable that the detector 3 is a high resolution area detector. 台架包括X-射线管2和面积检测器3。 Stage including X- ray tube 2 and the area detector 3. 控制器4A和4B受体积型CT扫描系统计算机1控制并分别连接到X-射线管2和检测器3。 The controllers 4A and 4B 1 controlled by volumetric CT scanning system computer and connected to the X- ray tube 2 and the detector 3. 控制器4A和4B使适当的旋转运动传递到X-射线管2和/或检测器3。 Controllers 4A and 4B cause the appropriate rotational movement to the X- ray tube 2 and / or the detector 3. 并不是每个控制器都需要。 Each controller is not required. 可以使用单一控制器部件使台架旋转。 Member so that the controller can use a single gantry rotation. 还应该指出的是,为实施本发明的方法计算机1控制图像扫描时间、图像分辨率和/或轴向覆盖区的变化。 It should also be noted that the present invention is a computer-implemented method of controlling an image scanning time, image resolution change region and / or axial coverage.

当对检测器3进行采集时计算机1通过给数据采集系统6发指令并控制台架的速度来控制数据采集过程。 When the detector 1 is collected by the computer 3 to the data acquisition system 6 and the speed of the gantry send commands to control the data acquisition process. 此外,计算机1指令数据采集系统6构造通过面积检测器3获得的射线照相的分辨率,由此能够改变系统的分辨率。 Further, a computer data acquisition system 6 instructions configured obtained by the area detector 3 radiographic resolution, it is possible to change the resolution of the system. 如图所示数据采集系统6包括读出电子系统。 Data acquisition system shown in Figure 6 comprises an electronic readout system.

面积检测器3包括检测器元件阵列(未示)。 Area 3 comprises a detector element array (not shown) detector. 每个检测器元件测量与其相关的强度值,该强度值与辐照到检测器元件上的X-射线能量的大小相关。 Each detector element measures an intensity value associated therewith, the intensity value associated with the size of the irradiation energy of the X- ray detector element. 当本发明的装置和方法并入到体积型CT扫描系统中,产生了一种新的体积型CT扫描系统。 When the apparatus and method of the present invention is incorporated into a volumetric CT scanning system to produce a new volumetric CT scanning system. 因此,本发明还提供一种新颖的体积型CT扫描系统。 Accordingly, the present invention also provides a new volumetric CT scanning system.

还应该指出的是本发明并不限于任何特定的计算机来执行本发明的数据采集和处理的任务。 It should also be noted that the present invention is not limited to any particular computer of the present invention to perform data acquisition and processing tasks. 如这里所使用的术语“计算机”是指任何能够执行计算并需要完成本发明的任务的机器。 The term "computer" as used herein refers to any machine capable of executing calculation and complete the object of the present invention. 因此,用于实现本发明的控制算法10的计算机可以是能够执行所需的任务的任何计算机。 Thus, a computer for implementing the control algorithm 10 of the present invention may be any computer capable of performing the required task.

关于本发明,已经确定通过数据平滑的替换方案来消除了需要应用额外的检测器元件来覆盖过渡区的需要。 About the present invention, it has been determined to eliminate the need for application of additional detector elements to cover the transition zone required by data smoothing alternative. 此外,如果已经应用覆盖整个视场的较大的检测器阵列来采集数据,则一种变型的方法应用迭代算法来估计已经测量的投影数据。 Further, if the application has to cover a larger detector array to collect the entire field of view data, a variant of the method of applying an iterative algorithm to estimate the projection data has been measured. 由于在过渡区中的误差可以以类似的方式处理,下文将在相同的条件中讨论这两种方法。 Due to an error in the transition zone may be processed in a similar manner, the two methods will be discussed below in the same conditions.

这种技术通过正向投影从前面的重复步骤中所获得的重构的数据或通过对从一组反向射线中获得的冗余检测器数据进行插值来形成X-射线投影数组{Pa},该反向射线在{Pa}的相反的方向中形成了另一投影数据组。 This X- ray projection techniques to form an array {Pa} by the redundant detector data obtained from a set of reverse ray in the forward projection data by interpolation obtained from the foregoing step is repeated or reconstructed, the reverse ray projection data forms another in opposite directions in the {Pa}. 这种正向投影技术是这样的一种方法,即射线从假想的X-射线源发射;这些射线对着每个检测器元件穿过所重构的体积。 Such forward projection technique is a method in which rays emitted from an imaginary X- ray source; these rays toward each detector element through the reconstructed volume. 沿着该射线,对所重构的值沿着射线的线性衰减值求和并表示为线性衰减系数的线积分。 Along the ray attenuation value of the reconstructed values ​​are summed and expressed as a line integral along a linear attenuation coefficient of linear rays.

正向投影所重构的数据的技术(表示为FPT)通常适合于产生与较大的锥形角相对应的投影数据(即,当用于VCT系统中时),而对冗余的投影数据进行插值的技术(表示为PDT)更适合于对更靠近中平面(即,更靠近ISO中心的平面)的投影数据。 Technical forward projecting the reconstructed data (denoted as FPT) is generally suitable for generating projection data corresponding to a large taper angle (i.e., when used in VCT systems) while the projection data redundancy interpolation technique (denoted as PDT) is more suitable for the projection closer to the mid-plane (i.e., closer to the center plane of the ISO) data. 与扇形角类似,锥形角是指在与扇形角方向正交的方向上从X-射线源发射的X-射线的角度范围。 Similarly to the fan angle, the cone angle refers to an angle range in the direction orthogonal to the fan angle direction X- ray emitted from the X- ray source. 在应用FPT或PDT所获得的估计检测器值和原始值(如果实际已经测量了该数据则将获得该值)之间的差可能使在靠近ISO中心的图像产生畸变。 The difference between the estimated detector values ​​and original values ​​in the application FPT or PDT obtained (actually having been measured if the data of the value would be obtained) may cause distortion in an image near the center of the ISO.

为降低这种畸变,已经研究出了一种应用平滑函数的方法,参考附图3该平滑函数表述如下:1. To reduce this distortion, we have developed a method of applying a smoothing function, with reference to the accompanying drawings 3 smoothing function expressed as follows: 1. 对于每个投影视图,选择最接近ISO中心21的已知的检测器元件,在此以后称为Va。 For each projection view, choose the known detector element closest to the ISO center 21, after this called Va.

2. 2. 对于相同的检测器元件,得出一种估计值22(即,对来自替换的视图中的投影数据的插值(PDT)或对所重构的数据进行正向投影(FPT)),在此以后称为Vb。 For the same detector element, obtain an estimate and 22 (i.e., the view from the interpolation alternative projection data (PDT) or the reconstructed forward projected data (FPT of)), after this called Vb.

3. 3. 产生适当的平滑函数23。 23 generate appropriate smoothing function.

4. 4. 平滑函数降低在Va和Vb之间的差并逐渐平滑在成像系统24的视场的中心附近的区域中的这种差。 Smoothing function to reduce the difference between Va and Vb is gradually and smoothly such a region near the center of the field of view 24 of the difference of the imaging system.

这可以从下面的讨论中得出,使d=Va-Vb,其中d是在视场的中心的投影数据中的不连续量。 This can be derived from the following discussion, so that d = Va-Vb, where d is the discrete quantities of projection data center in the field of view. 作为实例,可以用来逐步平滑该差的可能的平滑函数是一种如下定义的指数函数:V=0.5de-axo(等式1)这里Xo是距离与检测器元件值Va相对应的检测器位置的距离的绝对值,a是控制与平滑函数相关的曲线的斜率的系数。 As an example, a possible smoothing function can be used to smooth the phase difference is an exponential function defined as follows: V = 0.5de-axo (Equation 1) where Xo is the distance from the detector element value Va corresponding to the detector from the absolute value of a position, a is a smoothing function coefficient associated with the control curve slope. 将指数函数加到位于中心射线位置(对应于成像系统中的旋转中心在检测器上的投影的检测器位置)的一侧的投影值中/从位于中心射线位置(对应于成像系统中的旋转中心在检测器上的投影的检测器位置)的一侧的投影值中减去该指数函数以降低/提高所估计的值,以及从在中心射线位置的替换侧的投影值中减去/加到在中心射线位置的替换侧的投影值中以降低/增加更高/更低的原始值。 The exponential function added to projection values ​​located at the central ray location (the detector location corresponding to the projected center of rotation in the imaging system on the detector) side in / from the central ray position located (corresponding to the rotation of the imaging system side of the central projection value in the projection on the detector of the position detector) subtracting the exponential function to decrease / increase the estimated value, and the replacement value is subtracted from the projection side of the central ray location / plus to replace the value in the projection side of the central ray location to reduce / increase the higher / lower original value. 换句话说,它提供了一种方法,当将真实投影数据和估计投影数据结合在一起时这种方法降低了在中心射线的投影数据的不一致性,因此该方法在数据中提供了一种光滑的过渡区,该过渡区降低或消除了假象25。 In other words, it provides a method for, when the true projection data and estimated projection data are joined together in this way reduces the inconsistency of the central ray projection data, this method provides a smooth data in the transition zone, reducing or eliminating the false impression that the transition zone 25.

目前,在已有的技术中如何在半检测器移动结构的体积型CT(VCT)系统中应用面积检测器并不清楚。 Currently, how the system is applied in the area of ​​the detector is not clear volumetric CT detected the mobile structure is a half (VCT) in the conventional art. 前文已经描述了VCT系统和通过产生过渡区消除在成像系统的视场中的投影数据的不连续性的不同的公知的技术,下文将描述本发明的另一方面。 The foregoing has described a VCT system and the different discontinuities of known techniques in the field of view of projection data in the imaging system is eliminated by generating a transition zone, it will be described another aspect of the present invention.

应用变量fθ和fθ'来分别表示在源角度θ处所得到的正向和反向射线(相同的角度方位但以相反的方向通过的射线)的信号强度,这里 Variables and Applications f [theta] fθ 'respectively represent the source at an angle θ obtained spaces forward and reverse ray (same azimuth angle, but in opposite directions by radiation) signal intensity, where

fθ(n)=0 对于N2<n<N (等式2)fθ'(n)=0 对于1<n<N2(等式3)理想的fθ(N2)应该精确地等于fθ'(N2),因为两者都穿过对象的相同的部分。 fθ (n) = 0 for N2 <n <N (Equation 2) fθ '(n) = 0 for 1 <n <N2 (Equation 3) preferably fθ (N2) should be exactly equal to fθ' (N2) , as both pass through the same portion of the object. 但是由于下面的原因这永远不可能:(a)每个射线的实际形状是从源发出并在检测器终止的空的四面体。 However, due to the following reason never: (a) The actual shape of each ray is emitted tetrahedral empty and terminates the detector from the source. 没有完全相同的射线通过对象的相同的部分,除非对象是完全均匀的并圆形对称。 Not identical to any portion of the same ray through the object unless the object is totally homogeneous and circularly symmetrical.

(b)在扫描周期中对象/患者的运动都可能在每个正向/反向射线对中引入附加的误差。 (B) the object in the scan period / patient motion may introduce additional errors in each forward / reverse ray pairs.

(c)迄今为止还没有研制出完美的有效的插值方案。 (C) so far has not developed a perfectly valid interpolation schemes. 这就是说插值过程可能引入误差。 This means that the interpolation process may introduce errors.

假设d(θ)是在角度源位置θ的fθ(N2)和fθ'(N2)之差,如果d(θ)是完全随机的,则可能由所重构的图像感应的误差会被与其它的CT随机误差相连的量子噪声所掩盖。 Suppose d (θ) is the difference of the angle [theta] source location fθ (N2) and fθ '(N2) of, if d (θ) is completely random, then the error may be provided by the reconstructed image is induced with other CT is connected to the random error quantum noise masked. 然而,如果误差是某种系统的误差,它将在所重构的图像中引入明显的假象。 However, if the error is an error of some kind of system, it will introduce significant artifacts in the reconstructed image. 由于这个原因,必需在过渡区应用平滑处理。 For this reason, it is necessary in the transition region smoothing processing application. 换句话说,研究出一种平滑函数来使在检测器元件N2所表示的中心射线位置周围的fθ和fθ'的幅值误差更小。 In other words, developed a smoothing function to enable the detector element N2 indicated position around the central ray and f [theta] fθ 'of smaller amplitude error.

应用W和W'分别表示fθ和fθ'的的平滑函数。 Application of W and W 'represent the f [theta] and fθ' of the smoothing function. 当对W和W'求导时,必需考虑一定的规则,对于本领域的熟练技术人员来说这些规则都是可以理解的。 When W and W 'derivative, certain rules must be considered, the skilled person in the art that these rules are understandable. 此外,正如本领域的熟练技术人员将会理解的是,除了在此所特别说明的函数外,不同的平滑函数可以用于此目的。 Moreover, as a person skilled in the art will appreciate that, in addition to the special function described herein, different smoothing function can be used for this purpose. 例如W(n)+W'(n)=1对于所有的n (等式4)&delta;W&delta;n=&delta;W&prime;&delta;n=0]]>在n=N2±Δn (等式5)这里Δn设定W和W'的平滑范围,δ是微分算子。 For example, W (n) + W '(n) = 1 for all n (Equation 4) & delta; W & delta; n = & delta; W & prime; & delta; n = 0]]> at n = N2 ± Δn (Equation 5 ) where Δn is set W and W 'in the smoothing range, δ is the differential operator. 应该指出的是,对于这种类型的应用常规的平滑函数还通常称作尾翼函数(feathering function)。 It should be noted that, for applications of this type of conventional smoothing function is also commonly referred to as function tail (feathering function). 应用W和W'在正向和反向射线之间寻找过渡区。 Application of W and W 'between looking forward and reverse ray transition zone. 应该指出的是,为使平滑函数有效,Δn必需是比零大的整数。 It should be noted that the smoothing function is effective, Δn must be an integer greater than zero. 实际上,Δn越大,平滑函数的效果越好。 In fact, the greater the [Delta] n, the better the smoothing function. 然而,Δn增加得太多可能要求附加的检测器元件来延伸到在中心射线位置的检测器元件N2之外。 However, Δn is possible to increase too much the detector element extends into the central ray at the N2 position requires additional elements other than the detector. 因此,应该选择Δn足够大但是又不大到要求对过渡区增加额外的检测器元件。 Accordingly, Δn should be chosen large enough but not so large as in claim additional detector elements of the transition zone. 这样,对fθ和fθ'作如下的限制: Thus, for f [theta] and fθ 'limits as follows:

fθ(n)=0 对于N2+Δn<n<N (等式6)fθ'(n)=0 对于1<n<N2-Δn (等式7)在背面投影方法中所应用的实际检测器信号是Wfθ和W'fθ'。 fθ (n) = 0 for N2 + Δn <n <N (Equation 6) fθ '(n) = 0 for 1 <n <N2-Δn (Equation 7) In the rear projection method applied in the actual detector and the signal is Wfθ W'fθ '. 还应该指出的是,由于更宽的过渡区易于消除在正向和反向射线之间的许多不匹配的误差,所以对于每个半视场(FOV)的投影数据不需要叠加反向射线。 It should also be noted that, due to the wider transition zone easily eliminated many mismatch errors between the forward and opposing rays, so the projection data for each half field of view (FOV) need not be superimposed on opposing rays. 换句话说,每半个FOV数据(加上附加的Δn检测器值)填充以零以获得长度为N的检测器数据,在此之后进行常规的过滤投影程序。 In other words, each half FOV data (plus additional detector values ​​Δn) filled with a zero length N to obtain detector data, conventional filtration procedure after this projection. 不包含插值过程。 It does not include the interpolation process.

起因是当开始增加在CT系统中即在面积检测器中的检测器的行数时附加检测器元件(Δn乘以行数)变得更大。 When the cause is a CT system begins to increase the number of rows in that area detector an additional detector detector elements ([Delta] n times the number of rows) becomes larger. 因此,需要通过设计一种使Δn最小的方法和装置来改善常规的方法。 Therefore, by making Δn minimum design a method and apparatus to improve the conventional methods.

在本发明中这种方法可以将Δn减小到1,而计算机模拟表明常规的平滑方法要求Δn大约20才能实现相当的假象水平。 In this method of the present invention, Δn can be reduced to 1, the computer simulation showed that the conventional smoothing method requires Δn of about 20 to achieve comparable artifact level. 在VCT应用中这种优点更有意义,在VCT中在面积检测器的过渡区中所需的检测器元件的数目可能比在线性阵列中所需的元件还多三次幂的数量级。 This advantage in VCT applications more significant, the number of elements needed in the transition area of ​​the detector in the VCT detector elements that may be needed than in a linear array of more than three orders of magnitude power.

在正向和反向射线中可能有系统误差,通过插值或正向投影获得在反向射线中系统误差。 Forward and reverse ray system may have errors, systematic errors obtained by interpolation in the reverse rays or forward projected. 通过求fθ(N2)和fθ'(N2)之差测量幅值误差。 Measuring the amplitude error by calculating fθ (N2) and fθ '(N2) of the difference. 使d(θ)=fθ(N2)-fθ'(N2)。 So that d (θ) = fθ (N2) -fθ '(N2). 将d(θ)看作fθ(N2)和fθ'(N2)之间的幅值误差,在此θ是X-射线源的角度位置。 The d (θ) considered fθ (N2) and fθ 'amplitude error between the (N2), [theta] is the angular position of this X- ray source.

我们的方法是应用如在等式1中所述的指数函数来消除幅值误差,在等式1中a是控制指数函数的平滑性的控制系数。 Our approach is to use an exponential function as in Equation 1. The amplitude error to eliminate, in Equation 1 is a control factor controlling the smoothness of the exponential function. 将正向和反向射线函数fθ(n)和fθ'(n)分别转换成如下的两个其它的函数gθ(n)和gθ'(n):gθ(n)=fθ(n)-pθ(N2-n) (等式8)gθ'(n)=fθ'(n)+Pθ(n-N2) (等式9)使等式8和等式9的gθ(N2)=gθ'(N2)。 The forward and reverse ray function f [theta] (n) and the fθ 'conversion, respectively (n) into two other functions gθ (n) and the following gθ' (n): gθ (n) = fθ (n) -pθ (N2-n) (equation 8) gθ '(n) = fθ' (n) + Pθ (n-N2) (equation 9) of equation 9 and equation gθ 8 (N2) = gθ '( N2).

对于每个投影图像,实施下面的过程:1. For each projection image, the implementation of the following procedure: 1. 获得原始的半FOV投影数据,称其为fθ(n),并依据等式2补零。 Obtain the original half FOV projection data, called fθ (n), according to Equation 2 and zeros.

2. 2. 获得反向射线fθ'(n)的数组,并依据等式3进行补零。 Reverse ray obtained fθ '(n) of the array, and zero padding according to Equation 3.

3. 3. 依据等式8和等式9基于幅值误差d(θ)(这里d(θ)=fθ(N2)-fθ'(N2))应用平滑函数。 Based on Equation 8 and Equation 9 based on the amplitude error d (θ) (where d (θ) = fθ (N2) -fθ '(N2)) applying a smoothing function.

4. 4. 对gθ(n)和gθ'(n)进行积分以形成N-检测器数组,称为hθ(n),这里当N2<n<N时,hθ(n)=gθ(n)和当1<n<N2时,hθ(n)=gθ'(n)。 Of gθ (n) and gθ '(n) is integrated to form N- detector array, called hθ (n), where when N2 <n <N when, hθ (n) = gθ (n), and when 1 < n <N2 when, hθ (n) = gθ '(n).

5. 5. 对hθ(n)应用常规的滤波背向投影。 Of hθ (n) using conventional filtered back projection.

6. 6. 对所有的投影角度重复步骤1至5。 Repeat steps 1 through 5 for all projection angles.

对于在那些可以通过对在不同的角度之外的其它投影数据进行插值来获得反向射线的任何CT扫描器来说上面的过程都是有效的。 For the CT scanner can be any of those obtained by reverse ray projection data other than the different angles are interpolated for the above process is effective. 在如下的2D扇形束中比较理想:另一半FOV数据总是可以大致地从冗余的扇形束投影数据中计算出。 In the following ideal 2D fan-beam: the other half FOV of data can always be approximately calculated beam projection data from a redundant sector. 当将这种方法扩展到3D的VCT中时,当使用圆形轨道时仅在中平面上能够精确地插值。 When this method is extended to 3D VCT, when a circular orbit can be used only in the plane interpolation accuracy.

我们的模拟表明当将相同的方法应用到应用圆形轨道的VCT中时,对于在±1.5度的锥形角度这种方法比常规的平滑方法(应用超过中心射线的位置的20个附加的检测器)更优越。 Our simulation shows that when the same method is applied to a circular orbit in VCT applications, for 20 additional detection than the conventional smoothing method (applied over the location of the central ray ± 1.5 ° taper angle of this method device) superior. 如果应用完整的检测器,由于反向射线的角度方位与已经测量的数据差别极大,所以对于较大的锥形角度这种结果就不正确。 If the application is complete the detector, due to the reverse-rays has been measured and the angular orientation data vary widely, it is not proper for a larger taper angle such results. 为补救这种情况,可以应用迭代的方法来提高图像的质量。 To remedy this situation, an iterative method can be applied to improve the quality of the image. 该过程如下:1. The process is as follows: 1. 依据上述的步骤1至6获得初始的3D图像。 According to the above Step 1-6 to obtain an initial 3D image.

2. 2. 在该方法的第二次迭代中应用正向投影方法获得每个半FOV投影数据组的“反向射线”并依据上文所述的依据步骤3至步骤6将它们结合成一个完整的投影数据组。 In the second iteration of the method to obtain each half FOV projection data set through forward projection method "reverse ray" according to the above steps and according to claim 3 to 6 are combined into the step of a complete projection data group.

3. 3. 继续步骤2直到过程收敛,即不能再提高图像的质量。 Continue step 2 until the process converges, i.e., can not improve the image quality.

应该注意的是,结合一定的实施例已经讨论了本发明。 It should be noted that, in conjunction with certain embodiments of the present invention have been discussed. 然而,本发明并不限于这些实施例。 However, the present invention is not limited to these embodiments. 例如,所讨论的三种方案并不意味着都包括应用前述的参数的折衷来获得VCT系统的正确的操作模式的所有方案。 For example, three options discussed include trade-off does not mean that all the aforementioned parameters of the application program to get all the correct operating mode of the VCT system. 讨论这些方案是为了说明本发明的概念和方法,在这些方法中对这些基本参数进行折衷以实现正确的扫描方案。 These schemes are discussed to illustrate the concepts and methods of the present invention, these basic parameters of compromise in these processes to achieve the proper scanning protocol. 此外,这些折衷方案并非限于一种扫描方案,即它们还可以应用到轴向扫描(在扫描周期中患者的工作台并不移动)和螺旋扫描方案中。 In addition, these trade-offs are not limited to one scanning scheme, i.e., they may also be applied to an axial scanning (the patient table is not moved in the scanning period) and helical scanning scheme. 在本领域的熟练技术人员会理解这些方式,在这些方式中应用这些概念并外推以实现对特定领域的应用很有用的其它面积检测器扫描方案。 Person skilled in the art will appreciate that these methods, the application of these concepts and extrapolating these ways other area detector scanning scheme to enable application to specific fields useful.

Claims (2)

1. 1. 一种获得对象的投影数据的体积型计算机X-射线断层成像(VCT)系统,该VCT系统包括:X-射线源,该X-射线源对对象投影X-射线;检测器,该检测器相对于与CT系统的旋转中心在检测器上的投影相对应的中心位置移动了其宽度的一半,该检测器接收从X-射线源投影的X-射线并产生响应辐照在其上的X-射线的电信号;从检测器读取电信号并将该电信号转换为数字信号的数据采集系统;以及能够执行重构算法的计算机,该计算机从数据采集系统部分接收电信号,其中当计算机运行数据采集部分以处理所说的数字信号时,计算机重构图像。 A method of obtaining projection data of an object volume X- ray computed tomography (VCT) system, the VCT system comprising: X- ray source, the X- ray source X- ray projection of an object; a detector, the detector relative and the rotational center of the CT system to a projection on the detector corresponding to the center position of the half width of the X- ray detector receives from X- ray projection source and generates a response irradiated thereon X- rays electrical signal; reading the electric signal from the detector and the data acquisition system converts the electrical signal into a digital signal; and a computer capable of executing the reconstruction algorithm, the computer data acquisition system receives an electrical signal from the section, wherein when the computer is running data acquisition to said digital signal processing, image reconstruction computer.
2. 2. 一种应用计算机X-射线断层成像(CT)系统获得对象的投影数据的方法,该方法包括如下的步骤:从X-射线源向对象投影X-射线;在检测器的投影视图上接收CT系统在检测器投影的X-射线,该检测器包括许多检测器元件,该检测器元件产生响应在辐射在其上的X-射线的电信号I;数字化该电信号;对于每个投影视图,选择最靠近CT系统的中心的检测器元件的检测器元件值Va;对于所选择的检测器元件,经过从反方向或在相同的方向中的正向投影中插值估计一个检测器元件值Vb;选择能够消除在值Va和Vb之差的平滑函数;应用该平滑函数以消除Va和Vb之差;以及当结合正确的投影数据和估计的投影数据时应用加权函数来消除幅值差以产生平滑的过渡区。 X- method of application of the computer tomography (CT) system to obtain projection data of the object, the method comprising the steps of: X- ray from the X- ray source projected to the subject; receiving projection view on a CT system detector in the projector of the X- ray detector, which detector comprises a plurality of detector elements, the detector elements generate a response in an X- ray radiation on the electrical signal I; digitizing the electrical signals; for each projection view, choose the detector element value Va detector elements closest to the center of the CT system; selected for the detector element, through a detector element estimation value Vb in the backward direction or the forward projection interpolation in the same direction; selecting possible to eliminate the difference value Va and Vb of the smoothing function; applying the smoothing function to eliminate the difference between Va and Vb; and applying a weighting function when combined with the proper projection data and estimated projection data to eliminate the difference in amplitude to produce a smooth Transition zone.
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