Computerized tomographic scanners

Abstract: A computerized tomographic scanner (10) includes a reconstruction processor (82) for reconstructing an image from digital signals from detector arrays (20). Each detector array includes scintillation crystals (22) arranged in an array for converting x-ray radiation into visible light. An array of photodetectors (24) is mounted beneath the scintillation crystal array for converting light emitted from the scintillation crystals into electrical charge. The combination scintillation crystal, photodetector array is mounted to and preferably integral with an integrated circuit (26) having charge storage devices (52) and downstream circuitry (54). Electrical charge generated by each photodetector is integrated and stored alternately by a corresponding pair of charge storage devices. The charge storage devices operate as a double buffer in which one charge storage device accumulates charge while the other holds an accumulated charge until read out by downstream circuitry. The downstream circuitry includes a plurality of amplifiers (66) and an analog digital converter (62) integrally formed on the integrated circuit substrate (50). The analog to digital converter converts analog signals corresponding to the stored charges into digital values. The digital values created by the detector arrays are transferred in digital form to the reconstruction processor.

Description: This invention relates to computerized tomographic (CT) scanners. Heretofore, CT scanners have included a plurality of discrete radiation detectors which converted X-ray radiation which traversed a patient examination area into electronic signals. Each radiation detector included a radiation sensitive face, such as a scintillation crystal, which converted the received radiation into a corresponding quantity of light. A solid state photodiode was provided to convert the light emitted by the scintillation crystal into analog electrical signals indicative of the intensity of the crystal emitted light, hence the intensity of the received radiation. The radiation detectors were separately arranged on a circuit board. Each circuit board supported a linear array of photodiodes and attached scintillation crystals. In addition, a preamplifier was attached to the circuit board and connected to each photodetector output to convert the photodiode current to an appropriate voltage within the dynamic range of the analog-to-digital conversion system. The analog signals from the circuit board were carried to a central processing area where they were converted from their analog state into a corresponding digital signal. The analog signals were carried to the central processing area via a long bus system which extended around the scanner. One problem relates to degradation of the analog signals as they travel over the long bus system between the radiation detectors and the central processing area. CT scanners operate in a sea of extraneous radio frequency electromagnetic signals, the frequencies of which vary over a wide band. Sources of extraneous signals include nearby operating electrical components, equipment, signals from other detectors, and the like. The long bus systems include long lead wires which inadvertently act as antennas in picking up extraneous electromagnetic signals and converting them into analog signals. The extraneous analog signals are superimposed on and mix with the analog signals from the detectors. The superimposed extraneous signals appear as noise and fictitious data when reconstructed into images. The resulting images are degraded by noise, ghosting, and other artifacts. Prior art scanners commonly collected data from a single slice or a small plurality of.slices. To scan a volume, the patient was shifted longitudinally either incrementally after each slice to generate a multiplicity of parallel slices or continuously for a helical scan. The multiple rotation of the x-ray beam around the subject was time-consuming and stressful on the x-ray tube and generator. Such a device is for example disclosed in US-A-4 719 100. In particular, this device discloses a tomographic scanner as defined in the preamble of claim 1. It is an object of the present invention to provide a CT scanner wherein the above described problems are alleviated. The object of the invention is achieved by a computerized tomographic scanner in accordance with claim 1. One advantage of the present invention is that signals outputted by the radiation detector arrays are in digital form and are immune from the effects of surrounding extraneous radiation. Another advantage of the present invention is that it facilitates acquiring a good quality signal at a very high rate. Another advantage of the present invention is that the scintillation crystal array, the photodetector array, the amplifier, and the analog-to-digital converter can be formed integrally with the substrate to provide a fully integrated hybrid circuit using conventional manufacturing techniques. One CT scanner in 'accordance with the invention will now be described, by way of example, with reference to the accompanying drawings in which:- FIGURE 1 illustrates the CT scanner; FIGURE 2 illustrates an array of detector arrays of the CT scanner of FIGURE 1 arranged for use in receiving cone-shaped x-ray beams; FIGURE 3 illustrates a scintillation crystal and photodetector array of a detector array of Figure 2 in combination with an integrated circuit; FIGURE 4 is a block diagram illustrating one embodiment of the integrated circuit of FIGURE 3; and, FIGURE 5 is a block diagram illustrating a second embodiment of the integrated circuit of Figure 3. Referring to FIGURE 1, the CT scanner 10, which is intended for use in medical diagnostic imaging,selectively images regions of a patient (not shown) supported by a patient table 12 in a scan circle or examination area 14. The patient table 12 is positionable longitudinally. The CT scanner has an x-ray tube 16 which emits a beam of x-rays toward the part of the patient being examined. A collimator 18 collimates the x-rays into a cone beam. When the examination area 14 is irradiated with the x-ray energy, a percentage of the x-rays reaching the examination area 14 is absorbed by the patient's body. The amount of absorption depends on the density of bone or tissue upon which the x-rays are incident. The x-ray energy of the x-rays exiting the patient's body along each ray represents a linear integration of the radiation absorption of each incremental patient element along the ray. The absorption is indicative of relative tissue and skeletal densities. With continued reference to FIGURE 1 and with further reference to FIGURE 2, a plurality of detector arrays 20 receive radiation which has traversed the examination area 14. In the illustrated embodiment, thirteen arrays are arranged to receive a circular or rectangular cross-sectioned cone of radiation. Each array 20 is disposed along a plane normal to the center line of the cone shaped x-ray beam and is mounted for rotation with the x-ray beam. Cone shaped beams allow a larger volumetric region of the patient to be imaged per unit of time when compared with a patient region imaged using conventional fan shaped beams. With continued reference to FIGURE 2 and with further reference to FIGURE 3, each detector array 20 includes an n x m array of scintillation crystals 22 and an n x m array of photodetectors 24 connected or, preferably, integrated into an integrated circuit 26. The scintillation crystal/photodetector array 20 is covered with a light opaque, x-ray transmissive covering (not shown) to prevent extraneous light from reaching the crystals 22 or photodetectors 24. The scintillation crystal and photodetector arrays 22, 24 are preferably fully integrated with the integrated circuit 26 using conventional solid-state component manufacturing techniques. However, as further described below the scintillation crystals 22 and photodetectors 24 may alternatively be connected to the integrated circuit 26 using conventional bonding techniques. The scintillation crystals 22 are formed from cadmium tungstate or other similar materials capable of generating visible light upon being excited with x-ray radiation. A preferred embodiment uses 16 x 16 arrays of scintillation crystals 22 of 1 mm square area. It should be understood that a greater or lesser number of scintillation crystals 22 of varying widths can also be employed without departing from the inventive concept of the present application. Each scintillation crystal 22 of a 16 x 16 array is optically coupled to a respective one of the photodetectors of a 16 x 16 array of photodetectors 24 using conventional optical coupling cement, or the like. During a typical x-ray examination,. the patient is placed between the x-ray source 16 and the detector arrays 20. The scintillation crystals 22 convert x-rays that are transmitted through the examination area 14 to corresponding photons of light. The photodetectors 24 convert the light photons to electrical signals indicative of the integrated x-ray absorption along the corresponding ray. X-rays that have traversed the examination area 14 are received through the front face of a scintillation crystal 22. The scintillation crystal 22 converts these x-rays into visible light of a characteristic wavelength. The visible light exits each scintillation crystal 22 via a surface that is optically coupled to one of the photodetectors 24. The scintillation crystals 22 are covered on all surfaces except the surface which is optically coupled to a photodetector 24 with an x-ray transmissive, optical light opaque coating (not shown), e.g., an aluminum coating. Preferably, the coating is reflective such that substantially all light is reflected to the photodetector 24. Spatial intensity patterns of light emitted from the scintillation crystal arrays are proportional to spatial patterns of the x-ray radiation having traversed the examination area 14. Each of the scintillation crystals 22 is mounted on a photodetector of a 16 x 16 array of photodetectors 24 to convert the light emitted from the scintillation crystal 22 into discrete electrical charge. The preferred embodiment uses a 16 x 16 array of photodetectors 24 of one square millimeter area each. It should be understood that other numbers of photodetectors 24 and scintillation crystals 22 can be used. The technology of forming a 16 x 16 array of photodetectors is readily available. For example, photodetectors in the form of photodiodes are generally manufactured in a continuous sheet which forms one very large photodiode. By etching or cutting the surface layer, but not the substrate, the large photodiode sheet is subdivided into an array of small photodiodes. It should be noted that the photodetectors 24 can be constituted by devices other than photodiodes. Photocapacitors, photoresistors, charge coupled devices (CCD), and the like, are possible alternatives. Light from each scintillation crystal 22 is converted by the associated photodetector 24 into a corresponding charge value. In a photodiode, photons of light create charge pairs in an active surface layer. The charge is channelled from the active layer through detection circuitry to the substrate to create an output current that is proportional to the number of light photons received. In a charge coupled device, each cell is charged to a preselected level. Each photon of received light discharges a corresponding unit charge. The difference between the initial charge and the remaining charge is proportional to the number of photons of light received. With reference to FIGURE 4, aligned beneath the photodetector array and bonded thereto is the integrated circuit 26 which processes both analog and digital signals. The integrated circuit 26 is formed on a substrate 50 and includes charge storage devices 52, several cells of which are shown, and signal processing circuitry 54. In the FIGURE 4 embodiment, the charge storage devices 52 accumulate and store charge from the photodetectors' output currents until read out by the downstream processing circuitry. Each photodetector 24 is connected with charge storage device cells of two distinct groups of charge storage devices 56, 58. The charge storage devices "accumulate" photodetector charge in that they store charge during an x-ray exposure interval. The first and second groups 56, 58 of charge storage devices 52 operate in a dual buffer mode in which one group, while connected, accumulates charge from respective photodetectors as the other group is accessed by downstream circuitry of the integrated circuit which will be more fully explained below..It is to be appreciated that the dual buffer mode provides a faster data output from the radiation detector array. The first and second charge storage device groups 56, 58 preferably take the form of first and second groupings of capacitors integrally formed on the substrate. To facilitate accumulated charge output, rows of capacitors in one grouping are sequentially gated to a common bus (not shown) whereby stored accumulated charges in each row are serially accessed by downstream circuitry while capacitors in the other grouping continue to accumulate charge. In an alternative embodiment, the first and second groups 56, 58 of charge storage devices 52 take the form of charge injected devices (CID) or charge coupled devices (CCD). The CCDs or CIDs are interconnected in rows to form an analog shift register. Accumulated charges in each row are serially shifted until read by downstream circuitry. Again, one of the first or second groups 56, 58 accumulates photodetector generated charges while the other group is shifting accumulated charge to downstream circuitry. The foregoing radiation detector arrays 20 have been described as containing an array of photodetectors 24, each one of which is alternately connected to respective storage devices 52 in first and second groups 56, 58. However, in other embodiments (not shown)each photodetector could selectively be connected to a single charge storage device without departing from the inventive concept of the present invention. To maintain the dual buffer mode described above, the photodetectors in this embodiment function as charge storage devices in addition to functioning as a device for converting light into electrical charge. For example, photodetectors in the form of photocapacitors accumulate charge generated thereby during an x-ray examination interval. At the end of the x-ray examination interval, the accumulated charges in each photocapacitor are transferred to a respective charge storage device. Thereafter, charges stored on the charge storage device are sequentially read out by downstream circuitry while the photocapacitors accumulate further light converted electrical charge. The integrated circuit 26 further includes one or more analog to digital converters 62 integrally formed on the substrate 50 and interconnected with the downstream circuitry wherein analog signals corresponding to accumulated charges are converted into digital signals. Digital signals outputted by a detector array 20 are transferred from the rotating gantry of the CT scanner 10 to the stationary gantry by one or more slip rings (not shown). Optionally, a data compression circuit is used to reduce the volume of digital data transferred between the rotating and stationary gantries. The integrated circuit 26 further includes one or more multiplexers 64 and one or more amplifiers 66 which are integrally formed on the substrate 50. The amplifiers 66 are respectively connected to the charge storage devices 52 and boost the analog signals produced by the-accumulated charges thereon. The amplifiers 66 are interconnected to the charge storage devices 52 such that accumulated charges from distinct charge storage devices can be boosted by the same amplifiers 66. For example, charge storage devices 52 are arranged on the substrate in arrays of rows and columns. Each row of charge storage devices 52 is selectively connected to a respective amplifier. In particular, each common bus to which a row of charge storage capacitors are sequentially gated are connected to an input of a respective amplifier 66. In this arrangement, accumulated charges in the row are serially inputted to the connected amplifier 66. Similarly, each analog shift register formed by interconnecting a row of charge coupled devices or charge injected devices, can be connected to the input of a respective amplifier66. In this arrangement, charges accumulated may be shifted between adjacent charge storage devices until being inputted to an amplifier 66..In either arrangement, it is to be appreciated that the amplifiers 66 output analog signals in parallel. With reference to FIGURE 5, in an alternative embodiment, each amplifier 66 is alternately connected to two rows 56, 58 of charge storage devices 52 which are connected with a common row of photodetectors 24 in the double buffer configuration. In this arrangement, accumulated charges from one row 56 or 58 of charge storage devices 52 are serially inputted to a respective amplifier 66 while the other row 58 or 56 is accumulating charge generated by the photodetectors 24. As can be appreciated, this arrangement reduces the number of amplifiers 66 on the integrated circuit 26 resulting in on board power consumption savings since amplifiers typically require relatively large currents to operate effectively. Optionally, multiplexers 64 are interposed between the amplifiers 66 and the analog to digital converter 62. Each multiplexer 64 is integrally formed on the substrate 50. In this arrangement, analog signals outputted by the amplifiers 66 are multiplexed onto a single line prior to being converted into digital signals. signals. As set forth above, the analog to digital converter 62 converts analog signals to a corresponding digital output. These digital outputs are transferred via the slip ring to the stationary gantry. Bonding photodetectors to an integrated circuit often results in photodetectors of unequal characteristics. The photodetectors might be unequal in that they produce bias offsets which are included in charge accumulated within the detector itself or a connected charge storage device. The offset bias can be slight or extreme. However, the offset bias typically varies from detector to detector. In the extreme, offset bias causes analog signals to be biased outside the dynamic range of the amplifiers. In this situation, AC components of the analog signal which correspond to the x-ray radiation incident upon the scintillation crystals, could be filtered out inadvertently by the amplifier. The substrate 50 further includes gates (not shown) for connecting charge storage devices 52 to the amplifiers 66. In the CCD embodiment, the circuitry for sequentially clocking the charge values across the array to the amplifiers 62 is mounted on the substrate 50. The digital data from the detectors 20 is accumulated in a data memory 80. A cone beam reconstruction processor 82 reconstructs the data into a volumetric image representation which is stored in an image memory 84. A suitable cone beam reconstruction technique is show in "An Inversion Formula for Cone-Beam Reconstruction." SIAM Journal of Applied Mathematics, Tuy, H.K., pp. 546 - 552 (1983). A video processor 86 selectively accesses data from the image memory to produced slice images, 3-D renderings, and other- user requested displays for display on a video monitor 88 as are customary in the art. The present invention is also applicable to fourth generation scanners. By mounting a ring of the above-described 16 x 16 arrays around the stationary gantry, data for sixteen slices is collected concurrently. Of course, arrays of other sizes, such as a 4 x 6 array, may also be used.

Claims: A computerized tomographic scanner comprising an x-ray radiation source (16) for selectively emitting penetrating x-ray radiation which traverses an examination area (14) and an image reconstruction means (82) for reconstructing images representative of a subject in the examination area (14) using digital signals, a plurality of radiation detector arrays (20) positioned opposite the radiation source (16), each detector array (20) comprising (i) a plurality of scintillation crystals (22) arranged in an n x m array, where m and n are integers, each of said scintillation crystals converting x-ray radiation into visible light, (ii) a substrate (50), and (iii) a plurality of photodetectors (24) connected to the substrate (50), arranged in an n x m array, and optically coupled with the scintillation crystal array, each photodetector (24) converting visible light into electrical charge; a plurality of charge storage devices (52) formed on the substrate (50) for accumulating and storing charge, each photodetector (24) being selectively connected to a respective charge storage device (52) for altering the analog charge stored thereby in accordance with the charge generated in response to received light; a plurality of amplifiers (66) formed on the substrate (50) and connected with the charge storage devices (52) for selectively amplifying the stored analog charges; the scanner being characterized by : an analog to digital converter (62) formed on the substrate (50) and connected with the amplifiers (66) for converting the amplified analog charges into the digital signals for reconstruction; and the substrate aligned beneath the photodetector array. A scanner as set forth in claim 1 wherein the charge storage devices (52) include first and second groups (56, 58) of charge storage devices (52) and whereby each of the photodetectors (24) is provided for alternate connection to a respective one of the charge storage devices (52) of the first group (56) and to a respective one of the second group (58) such that while the amplifiers (66) create an analog signal from the charge storage devices (52) of one of the two groups (56,58) the photodetectors (24) alter the charge on the charge storage devices (52) of the other of the two groups (56, 58). A scanner as set forth in one of the claims claim 1 or 2 wherein each radiation detector array (20) further includes: a multiplexer (64) formed on the substrate (50) being selectively connected to the amplifiers (66) and being connected to the analog to digital converter (62). A scanner as set forth in claim 1 wherein each radiation detector array (20) includes first and second groups (56, 58) of charge storage devices (52), each photodetector (24) provided for an alternate connection to a respective one of the charge storage devices (52) from the first group (56) and the second group (58); a second plurality of amplifiers (66) connected with the second group (58) of charge storage devices (52); a multiplexer (64) formed on the substrate (50) and selectively connected to the amplifiers (66); and at least a second analog to digital converter (62) formed on the substrate (50) and connected to an output of the multiplexer (64), the first and second analog to digital converters (62) provided for operating in parallel. A scanner as set forth in any one of claims 1 to 4 wherein the charge storage devices (52) are integrally formed in columns and rows on the substrate (50), and charge storage devices (52) of each row are selectively interconnected to allow accumulated charge to be shifted from adjacent charge storage device (52) to adjacent charge storage device (52) to serialize stored charges supplied to the amplifiers (66). A scanner as set forth in any one of claims 1 to 5 including bias offset correction means. A scanner as set forth in any one of claims 1 to 6 wherein the photodetectors (24), charge storage devices (52), and amplifiers (66) are integrally formed on the substrate (50). A scanner as set forth in claim 9 wherein the scintillation crystals (22) are also integrally formed on the substrate. A scanner as set forth in claim 1 wherein the detector arrays (20) are themselves arranged in an array. A scanner as set forth in claim 9 wherein penetrating x-ray radiation being selectively emitted from the x-ray source (16) forms a cone-shaped beam and the array of radiation detector arrays (20) conforms to a transverse cross-section of the cone beam.

Inventors: BENDULA DALE J, TUY HEANG K, BENDULA, DALE J., TUY, HEANG K.
Applicants: KONINKL PHILIPS ELECTRONICS NV, KONINKLIJKE PHILIPS ELECTRONICS N.V.
Application year: 1995
Main classification: A61B6
IPC codes: A61B6, G01T1