RADIATION IMAGING SYSTEMS

Abstract: A radiation imaging apparatus and method e.g. for com­ puted tomography x-ray imaging employing a radiation detector which enhances the detector output due to lower energy compo­ nents of the incident radiation. One embodiment includes first and second layers (190, 192) of crystalline scintillation material mutually aligned in a path of x-rays to be detected, to receive the x-rays in sequence. The layer (190) upstream in the x-ray path comprises a scintillation material having a relatively high effi­ ciency for converting x-ray energy to light. The downstream layer (192) comprises a scintillation material having a relatively lower efficiency for x-ray/light conversion. A photodiode (198) is positioned to view both scintillation layers (190, 192) simultane­ ously and to respond to scintillations in either or both. Scintillation crystal material surfaces can be coated with reflective material (194) to enhance the effects of their scintillations. The photodiode (198) thus combines x-ray indicating scintillations from both crystals while in analog form. Another embodiment comprises a photodiode and an optically coupled scintillation crystal, with the photodiode upstream in the x-ray beam path relative to the crystal.

Description: The present invention relates to radiation imaging systems e.g. for computed tomography imaging. In both emission and transmission computed tomography, information concerning the internal structure of a subject of interest is obtained without invasive procedures. In computed tomography a two dimensional image is generated from multiple one dimensional projections. A source of radiation, e.g. x-radiation, transmits x-rays through a subject of interest and the intensity of that radiation is monitored on an opposite side by one or more detectors, which are often single scintillation crystals each optically coupled to a different photodiode. If the source is moved about the patient and x-ray attenuation data is obtained from a number of directions, a computed tomography reconstruction process can be utilized to generate an image of the subject cross-section. The x-ray tube radiation source used in computed tomography emits x-ray photons over polychromatic or continuous energy spectrum which extends from near zero in energy to a maximum which is determined by the voltage (kv) applied to the source. The performance and energy spectra of x-ray tubes is well understood and documented by those of ordinary skill in the relevant art. The radiation detectors utilized in computed tomography each produce a signal proportional to the total x-ray energy absorbed by the detector. The lower energy x-ray photons striking a detector contribute a relatively small amount to the total output signal of the detector, whereas the higher energy photons incident on the detector produce correspondingly larger contributions to the output signal. The low energy x-ray photons are more readily absorbed in the patient, while the high energy photons have a greater tendency to pass through the patient's body and to undergo less attenuation than do the low energy x-ray photons. Hence, the low energy radiation contains more contrast information defining the patient's internal body structure than does the high energy radiation, which more uniformly passes through the patient's body, as a result of less attenuation. Therefore, the detector output signal is more a function of the high energy radiation, which carries relatively less contrast information, than it is a function of the lower energy radiation which is relatively rich in image information. Accordingly, the present detectors that employ a single energy sensitive element (e.g., a single scintillation crystal or a gas filled ion chamber) weights its output signal in favor of the less informative high energy radiation than in favor of the more informative low energy radiation. The generation of x-ray photons by an x-ray tube anode is a random process. Even where the x-ray tube kV and mA (voltage and current) are held constant, the number of photons emitted from the anode fluctuates statistically in time about an average value with a Poisson probability distribution. The absorption, or attenuation, of the x-ray photons in matter, such as in a patient's body, is also a random process following similar laws of probability theory. Therefore, the number of photons detected during a fixed time period, with all other conditions held constant, will vary from one measurement to the next. This statistical fluctuation in the measurement gives rise to an uncertainty or ambiguity in the true value of the attenuation. This ambiguity is sometimes referred to as "quantum noise" or "quantum statistics". While this quantum noise sets an absolute, or fundamental limit on the quality of an image that can ultimately be obtained, unequal response by the detector to photons of different energies magnifies this noise problem. Noise in the detector output signal which results from the quantum statistics associated with the detection of N x-ray photons is given by the following relation: wbere: σ = the standard deviation (or RMS noise) of the output signal, expressed as a fraction of the signal (i.e., noise to signal ratio); W n = Weight (or signal contribution) of the n th photon, and N = Total number of photons detected. In most present detection systems used in computed tomography, W is proportional to the absorbed energy of the photon. However, less noise (or σ) is obtained if W n is constant (i.e., every photon is weighted equally). In this optimum case, the standard deviation is given by the familiar formula: To illustrate with an example, consider two cases: Case I. (Unequal Weighting) Let N = 100 and assume that half of the detected photons have a relative weighting of W, and the other half are higher in energy and have a relative weighting of 4W, then Case II. (Equal Weighting) Let N = 100 and assume that all photons are detected with relative weighting of W; then As can be seen in this set of examples, Case I is 17% noisier than Case II. In order for Case I to achieve the reduced level of noise of Case II, the number of detected photons, i.e., N, would have to be increased to 136. This would correspond to an increase in patient dose of 36%. Thus, the characteristic of present detectors to weight, their response to high energy photons more heavily than to low energy photons results in an exaggerated effect of noise in the output signal, and, in order to compensate for this noise, requires a larger x-ray dose to be applied to the subject. It is an object of this invention to provide an apparatus wherein this difficulty is alleviated. According to a first aspect of the present invention there is provided a radiation imaging system comprising: a radiation source for propagating radiation along a path; a detector assembly in said radiation path for producing an analog output signal in response to said radiation, said detector assembly comprising first and second scintillator components preferentially responsive to lower and higher energy components of said radiation respectively and a detector component for producing said output signal in response to the responses of said first and second scintillator components to said radiation; and imaging circuitry responsive to said output signal, said first and second scintillator detector assembly components comprising scintillation material and said detector component comprising a detector for scintillation produced by said material, said first and second scintillator detector assembly components being arranged serially along said radiation path with said first scintillator detector assembly component nearer said source and having a higher photon to light conversion efficiency than said second scintillator detector assembly component, characterised in that said detector component is disposed along said radiation path between said first scintillator component and said source thereby to enhance a component of said output signal produced in response to a lower energy component of said radiation. According to a second aspect of the present invention there is provided a radiation imaging system comprising: a radiation source for propagating radiation along a path; a detector assembly in said radiation path for producing an analog output signal in response to said radiation, said detector assembly comprising a scintillator component and a detector component for producing said output signal in response to the response of said scintillator component to said radiation; and imaging circuitry responsive to said output signal, said scintillator detector assembly component comprising scintillation material and said detector component comprising a detector for scintillation produced by said material, characterised in that: said scintillator component comprises a single layer of scintillation material; and said detector component is disposed along said radiation path between said source and said scintillator component thereby to enhance a component of said output signal produced in response to a lower energy component of said radiation. W0-A-8100457 discloses a radiation imaging system including a detector assembly comprising first and second components preferentially responsive to lower and higher energy components of radiation respectively and a third component for producing an output signal in response to the responses of the first and second components to said radiation thereby to enhance a component of the output signal produced in response to a lower energy component of the radiation. One radiation system in accordance with the invention will now be described by way of example with reference to the accompanying drawings in which: Figure 1 is a perspective view, partially broken away, illustrating the general form of the system; Figure 2 is a flow diagram illustrating functions of operation of the system of Figure 1; and Figures 3-6 are drawings of respective radiation detectors for use in the system of Figure 1, the detectors of Figures 4 and 5 only being in accordance with the present invention. Turning now to the drawings, Figure 1 illustrates the general form of a computed tomography scanning system 10 used in imaging cross-sectional slices of interest in a patient. The computed tomography system 10 comprises a scanner 12, a viewing console 14, a computer system 16, and specialized electronics 17 needed by the scanner 12 for control and data handling. While the invention can be employed in any CT geometry, the scanner as shown for example at 12 comprises a fourth generation computed tomography scanner having a movable x-ray source (not shown) and a fixed array of detectors such as 19 surrounding a patient aperture 18. During imaging, a patient is positioned on a couch 20 and then moved into and through the patient aperture 18 until a cross-sectional slice to be imaged is appropriately positioned. A series of electronic subsystems 30 shown to the side of the computed tomography scanner 12 provide power suitable for creating x-radiation. In computed tomography scanning, the special electronics 17 analyzes intensity values detected by the scanner 12. This specialized electronics 17 measures the output from the scanner detectors as well as controls movement of an X-ray tube source and coordinates this movement with the analysis of the output signals. High speed computed tomography imaging is possible though use of a high speed data processing computer system 16. The computer system 16 performs the data processing for reconstructing an image of attenuation variations inside the patient slice from intensity readings taken from a plurality of detectors surrounding the patient aperture. The computer system is responsible for analyzing and reconstructing cross-sectional image densities and for displaying this information on the console 14. The console depicted in Figure 1 includes a first CRT 37 for a technician operating the computed tomography apparatus and a second CRT 38 for a person responsible for diagnosing from the images produced. The system employs new and improved detector structure. Each preferred detector comprises at least one scintillation crystal coupled to at least one photodiode, as explained in more detail below. In operation, when x-radiation impinges upon a scintillation crystal, it is converted to visible light which in turn causes current flow in a photodiode. The analog current produced by the x-radiation is converted to a digital signal by an analog to digital converter 51. Electronics for generating these digital signals are known in the art. The steps (See Figure 2) of detecting the radiation 50 and generating the digital output 51 of the converter are depicted in a flow chart schematically describing the computed tomography process. The two steps 50, 51 are followed by taking the logarithm of the data and a storing 54 of that data in the computer system. The logarithm of the intensity data yields attenuation information. The remaining steps in the computed tomography process are performed by the computer system 16. The computer system first performs a series of calibration calculations 56 on the data. These calculations are based upon data obtained during a CT set-up phase. These calculations take into account variations in detector sensitivity, gain, and offsets in the electronics. Once these calibration steps have been completed, a digital filtering step 58 is performed where all data from each detector is filtered in accordance with a filter function. The process consists of performing a forward fast fourier transform (FFT) of the data, multiplying the transformed data by a frequency filter, and then performing an inverse FFT to produce the filtered data. Finally, the data are back projected at 66 into a memory to produce an image of a particular patient slice under examination. Once this back projection process has been completed, the image can be viewed at the console 14 (figure 1). An important purpose of this invention is the provision of improved embodiments for the detectors, having enhanced response to lower energy radiation and which reduce the noise resulting from the statistical variations of the radiation photons. An important feature of the invention involves increasing the weighting, or detector output signal contribution of the lower energy x-ray photons relative to the weighting or signal contribution of the higher energy photons. Such a radiation detector will produce an output signal which is lower in noise and higher in information content than could be produced by prior art detectors. One such detector comprises a scintillator including a sandwich of two or more layers of scintillation material having different energy conversion efficiency characteristics. Referring to Figure 3, the geometry of such a detector 100 is such that radiation from the source first impinges upon a first layer 102 of scintillation material, such as cesium iodide, of thickness 0.2 mm, having a relatively high efficiency in converting x-ray photons to light photons. In the layer 102, the lower energy x-ray photons tend to be preferentially absorbed, while the higher energy photons tend to pass through the layer 102 and to be absorbed in a layer 104 which comprises a scintillation material, such as cadmium tungstate, of thickness 2.0 mm, having a relatively lower efficiency in converting x-ray photons to light photons. The interface between the layers 102, 104 is at least partially light transmissive. The light scintillations from both the layers 102, 104 are detected by a photodiode 106, which views both layers simultaneously, and produces an analog electrical output signal on a lead 108 which is a function of the detected analog scintillations from both layers. The net effect of this phenomenon is that the component of the analog output signal, appearing at the lead 108, which is due to the lower energy photons, has enhanced weighting relative to the contribution attributable to the higher energy x-ray photons striking the detector. That is, the signal at the lead 108 has a relatively enhanced component which comprises a contribution resultant upon the conversion to light energy of the lower energy x-ray photons, compared to the results obtained in prior art detectors having only a single crystal upstream of the photodiode in the x-ray path. Additionally, the enhancement of the response of the output 108 due to lower energy x-ray photons renders the signal appearing at the output 108 more independent of random photon fluctuations of the incident x-ray energy than would be the case if only the lower efficiency crystal 104 were used. This is because the high and low energy components of the analog output signal at 108 are more nearly equal, for a polyenergetic spectrum, when the weighting of the lower energy contribution is enhanced. Figure 4 shows a first detector in accordance with the present invention which achieves enhanced weighting of the detector response to the lower energy x-ray photons and includes a single relatively thick scintillator layer 110 and a single photodiode 112. The photodiode is positioned upstream in the x-ray path with respect to the crystal 110. Thus, x-rays transmitted through the subject first pass through the relatively thin photodiode 112, and are then absorbed in the crystal 110. The lower energy x-ray photons are preferentially absorbed, and converted to light photons, in the region of the layer 110 near its upstream surface 116. The higher energy x-ray photons are converted to light photons generally more near the lower, or downstream, surface 118 of the layer 110. Thus, the light photons resulting from the lower energy x-ray photons are preferentially converted to light scintillations nearer the photodiode 112 than are the scintillations resulting from the higher energy x-ray photons. This is true because the higher energy photons tend to penetrate further through the layer 110, before being converted to light scintillations, than do the lower energy x-ray photons. Therefore, the photodiode 112 will respond more strongly to the scintillations near the surface 116 than to those near the surface 118, due to the fact that the scintillations near the surface 116 are closer to the photodiode than those further away near the surface 118. The photodiode 112 comprises an epitaxially grown layer of silicon. The scintillation layer 110 comprises a portion of semi-transparent crystal, for example cadmium tungstate, and has a thickness of about 2 mm. The semi-transparency of the layer 110 helps to exaggerate the attenuation of the response of the photodiode 112 to more distant scintillations occurring near the surface 118, to assist in further enhancing the weighting of the low energy response in the output signal at the lead 114. Figure 5 shows a second detector in accordance with the present invention incorporating a sandwich crystalline layer assembly 120, 122, in combination with a photodiode 124. As in the first detector in accordance with the present invention of Figure 4 the photodiode 124 is positioned relatively upstream in the x-ray path, and receives incident x-rays prior to the crystalline layers. The incident x-rays first penetrate the photodiode 124, which is substantially transparent to x-ray, and are absorbed in the layers 120, 122. In the detector of Figure 5 the layer 120 comprises a high efficiency absorber, such as cesium iodide, of thickness 0.2 mm, which preferentially absorbs and converts to light lower energy x-ray photons. The layer 122 comprises a material having lower efficiency for converting x-ray photons to light photons, such as cadmium tungstate, with a thickness of 2 mm. As in the case of the detectors of Figures 3 and 4, the photodiode 124 produces an output signal 126 having an enhanced response to lower energy photons, and which combines lower and higher energy representing signals in analog form. In order to maintain the statistical noise at a minimum, it is important that only a very small fraction of the x-ray photons be absorbed in the diffusion layer of the silicon in the photodiode. A photodiode of epitaxially grown silicon, as described above, satisfies this criterion. In the detectors described above, additional measures can be taken to further influence the respective weighting of response to higher and lower energy x-rays. For example, different scintillation phosphors can be chosen based on their respective preferential absorption of lower and higher x-ray energy. Choices of such phosphor materials for their preferential absorption characteristics are generally described in published European Patent Application No. 0115125, published on August 8, 1984, by Gary L. Barnes, which is expressly incorporated by reference. Suitable low efficiency crystal materials include bismuth germanate, and higher efficiency materials include cadmium tungstate. Preferably, scintillator materials are chosen which fluoresce in the green to red portion of the spectrum, to which photodiodes tend to be more responsive than to other colors. The scintillators preferably have k-absorption edges below about 40 kev. In the Figure 3 detector the crystal 102 has a higher photon to light conversion efficiency than does the crystal 104. The crystal 104, however, is more radioabsorptive than the crystal 102. Generally, the crystal 102 can have a thickness of about 0.1 mm to about 0.4 mm, and the crystal 104 can have a thickness of about 2 mm. The detector of Figure 6 can be employed to alter or enhance the relative weighting of the x-ray energy components of a detector output signal. The detector of Figure 6 includes crystalline layers 190, 192, and a photodiode 198 having an output 199. For example, as shown in Figure 6, increasing the thickness of the layer 190 with respect to the layer 192 further alters the response of the layer 190 to x-rays and therefore enhances the response of the photodiode 198 to scintillations occurring in the layer 190. Furthermore, the relative response of the photodiode 198 to scintillations in the layers 190, 192, respectively, can be altered by choosing the crystals 190, 192 to have differing geometry and/or area size, as also shown in Figure 6. The response of the photodiode 198 to scintillations occurring in a given layer, such as 190, can also be enhanced by providing a layer or coating 194 of reflective material about some or all of the outer surface of the scintillator. Selection of the type of reflective material and means of application to create such reflective surfaces is within the ordinary skill in the art. All the foregoing aspects of the detector of Figure 6 can be employed in order to desirably tailor the response of the photodiode 198 to scintillations within the respective layers 190, 192, in order to enhance the weighting of the response to low energy x-ray photons in the output signal 199 from the photodiode 198. Considering also the second detector in accordance with the present invention of Figure 5, it is seen that a layer 121 which is only partially transmissive of visible light can further enhance the weighting of the response of the photodiode 124 to scintillations occurring in the layer 120, by inhibiting the response of the photodiode to scintillations occurring in the crystal 122.

Claims: 1. A radiation imaging system comprising: a radiation source for propagating radiation along a path; a detector assembly (19) in said radiation path for producing an analog output signal in response to said radiation, said detector assembly comprising first and second scintillator components (120, 122) preferentially responsive to lower and higher energy components of said radiation respectively and a detector component (124) for producing said output signal in response to the responses of said first and second scintillator components (120, 122) to said radiation; and imaging circuitry (14, 16, 17) responsive to said output signal, said first and second scintillator detector assembly components (120, 122) comprising scintillation material and said detector component (124) comprising a detector for scintillation produced by said material, said first and second scintillator detector assembly components (120, 122) being arranged serially along said radiation path with said first scintillator detector assembly component (120) nearer said source and having a higher photon to light conversion efficiency than said second scintillator detector assembly component (122), characterised in that said detector component (124) is disposed along said radiation path between said first scintillator component (120) and said source thereby to enhance a component of said output signal produced in response to a lower energy component of said radiation. 2. A system according to Claim 1 wherein said detector assembly includes a further component (121) between said first and second scintillator components (120, 122) to reduce the response of said detector component (124) to scintillation produced in said second scintillator component (122). 3. A system according to Claim 1 or Claim 2 wherein said first scintillator component (120) comprises cesium iodide. 4. A system according to Claim 1 or Claim 2 or Claim 3 wherein said second scintillator component comprises cadmium tungstate. 5. A system according to Claim 1 or Claim 2 wherein said first (120) and second (122) scintillator components comprise cadmium tungstate and bismuth germanate respectively. 6. A system according to any one of Claims 1 to 5 wherein said first (120) and second (122) scintillator components are in the form of layers of scintillation material having thicknesses in the range 0.1 to 0.3 mm and of substantially 2.0 mms respectively. 7. A radiation imaging system comprising: a radiation source for propagating radiation along a path; a detector assembly (19) in said radiation path for producing an analog output signal in response to said radiation, said detector assembly comprising a scintillator component (110) and a detector component (112) for producing said output signal in response to the response of said scintillator component (110) to said radiation; and imaging circuitry (14, 16, 17) responsive to said output signal, said scintillator detector assembly component (110) comprising scintillation material and said detector component (112) comprising a detector for scintillation produced by said material, characterised in that: said scintillator component comprises a single layer of scintillation material (110); and said detector component (112) is disposed along said radiation path between said source and said scintillator component thereby to enhance a component of said output signal produced in response to a lower energy component of said radiation. 8. A system according to Claim 7 wherein said scintillation material comprises a material which is semitransparent to the scintillations produced by said material. 9. A system according to Claim 8 wherein said scintillation material comprises cadmium tungstate. 10. A system according to any one of Claims 1 to 9 wherein said detector component (112 or 124) is a photodiode. 11. A system according to any one of Claims 1 to 6 wherein said first and second scintillator detector assembly components (120, 122) have different geometries. 12. A system according to any one of Claims 1 to 6 or Claim 11 wherein at least one of said first and second scintillator detector assembly components (120, 122) includes a surface layer (194) of reflective material to assist in said enhancement. 13. A system according to any one of Claims 7 to 9 wherein said scintillator component (110) includes a surface layer (194) of reflective material to assist in said enhancement. 14. A system according to any one of Claims 1 to 6 or Claim 11 or Claim 12 wherein said first and second scintillator components (120, 122) and said detector component (124) are arranged in stacked closely spaced configuration. 15. A system according to any one of Claims 7 to 9 or Claim 13 wherein said scintillator component (110) and said detector component (112) are arranged in stacked closely spaced configuration. 16. A system according to any one of the preceding claims wherein the system is a computed tomography x-ray image system.

Inventors: BRUNNETT, CARL J., MATTSON, RODNEY A.
Applicants: PICKER INTERNATIONAL, INC.
Application year: 1986
Main classification: G01T1
IPC codes: A61B6, G01T1, G01N23