IMPROVED SOLID STATE DETECTOR

Abstract: A detector for high energy radiation such as X-ray radiation or gamma radiation comprises a layer (1) of scintillator material, an intermediate Mylar layer (2) and a solid state detector array (3) located behind them. The intermediate layer has an atomic number which is less than that of the scintillator material or the radiation sensitive regions of the detector. Hence, secondary electrons produced at the detector region move to the intermediate layer, reducing the dose delivered by X-rays to the sensitive regions and prolonging the life of the detector.

Description: IMPROVED SOLID STATE DETECTOR This invention relates to detectors for high energy radiation and more particularly, but not exclusively, to detectors employed in medical imaging applications. In dental X-ray imaging and for other medical applications, for example, the use of solid state detectors is now being advocated to replace previously used X-ray sensitive film. In a typical device, a layer of scintillator material is deposited on a fibre optic taper which is in turn bonded to a detector array, for example a charged coupled device (CCD) array. The object to be imaged is irradiated by a beam of high energy radiation, such as X-rays, which, after passing through the object, are incident on the scintillator material. The X- rays are converted into optical radiation at the scintillator material for detection by the solid state detector array, being directed by the fibre optic taper. In some scanning arrangements such as in a computed tomography (CT) scanner, high doses of X-ray radiation are required to produce a satisfactory image. However, the X-rays may be of sufficiently high energy that significant numbers of X-rays avoid annihilation within the scintillator and deposit their energy within the detector array itself, causing damage to sensitive structures at its surface. In a CCD array, ionising radiation may damage the dielectric material by causing charge to be trapped within the dielectric which leads to a voltage shift, thus requiring a change in operating conditions with time. Also, damage at the surface tends to cause an increase in dark current. Thus, compensation must be made to take into account the change in the characteristics of the device with time and also the lifetime of the CCD is reduced. High energy radiation also causes damage to other types of solid state detectors. For example, in a photodiode array, switching components are highly susceptible to damage by radiation. To avoid radiation damage it has previously been proposed to use an optical taper of a thickness which prevents X-rays which leave the scintillator material from reaching the detector array. It may be necessary for particularly high energy radiation to use a fibre optic taper of several centimetres thickness. Alternatively, the thickness of the scintillator material is increased such that it absorbs substantially all incident X-rays. The present invention seeks to provide an improved detector for high energy radiation which is suitable for medical and non-medical applications such as CT systems and for dental use. According to a first aspect of the invention, there is provided a CCD arrangement for detecting high energy radiation comprising: a layer of scintillator material for converting incident high energy radiation into optical radiation; a CCD for detecting optical radiation and having regions which are sensitive to damage by ionizing radiation; and an intermediate layer located between the scintillator and the detector regions and adjacent the detector regions, the intermediate layer being substantially transmissive to optical radiation and having an atomic number which is lower than that of the detector regions. The atomic number for a compound is taken to be the effective atomic number. This is a number calculated from the composition and atomic numbers of a compound or mixture. An element of this atomic number would interact with photons in the same way as the compound or mixture. Various formulae for this number have been developed. For example, Spier's is based on theoretical considerations involving absorption and scattering coefficients, the effective atomic number, Z, being given by Z = (a 1 Z 1 2 - 94 +a^ 2 294 +....) " 294 where Z,, Z 2 etc., are atomic numbers of individual constituents, and a,, a j etc., the fractional electron contents of elements Z,, Z 2 etc. in the compound. According to Fricke and Glasser, based on theoretical considerations of photoelectron production, Where Z,, Z j etc. are the atomic numbers of the constituents, and al, a2 their fractions by weight. By "optical radiation" it is meant radiation falling within the visible part of the spectrum, ultra violet and/or infrared radiation. By "substantially transmissive" it is meant that sufficient optical radiation may be transmitted through the intermediate layer for the solid state detector region to adequately image the high energy radiation directed onto the scintillator material. By employing the invention, it is possible to use the detector with high energy radiation such as X-rays gamma radiation and electrons with satisfactory results without the need for a long fibre optic taper, say, to prevent X-rays from reaching the detector region. When high energy radiation irradiates a material, secondary electrons are produced by processes such as the photoelectric effect and the Compton effect. Where the X-rays are directed onto a material of higher atomic number located in front and adjacent to a material of a lower atomic number, the secondary electrons produced in the front high atomic number layer are more numerous than those in the layer located behind it. At the boundary between the two layers, therefore, secondary electrons from the high atomic number (Z) layer tend to move across the boundary into the low Z region, increasing the number of secondary electrons in the low Z layer at the boundary. This enhanced dose effect is employed in the present invention. The intermediate layer in accordance with the first aspect of the invention has a lower atomic number Z than that of the adjacent CCD detector regions. When the CCD arrangement is irradiated with high energy radiation, secondary electrons from the scintillator material will travel into the intermediate layer causing the number of electrons to be higher at the boundary in the intermediate layer than would be the case without the enhanced dose effect. However, any dose deposited at the surface of the detector regions adjacent the intermediate layer causes energy to be transferred from the layers of the detector regions sensitive to damage by ionizing radiation to the intermediate layer as secondary electrons migrate into the intermediate layer. This reduces the dose at the surface of the detector region and hence prolongs the life of the CCD and reduces variations in operating characteristics with time. By using the invention, therefore, although the bulk material of the detector region is subject to an X-ray dose similar to that which would be delivered in the absence of the use of the intermediate material, that deposited within the sensitive layers has been found to be reduced by a factor which in some embodiments is of the order of 70. Figure 1 schematically illustrates the relative dose deposited in the sensitive regions of the detector for different thicknesses of the intermediate layer for incident radiation at several energy levels. As can be seen, even though a thickness of, say, 5μm would not stop any X-rays use of the invention significantly reduces the relative dose at the sensiti ve regions . Use of the invention avoids the need to place significant thicknesses of material between the scintillator material and the CCD regions in order to stop high energy X-rays from causing damage at the sensitive layer of the detector regions whilst permitting optical radiation to also be incident at those regions for detection. The invention may be particularly advantageously used in computed tomography scanner arrangements, for example, where it is desirable that the size and mass of the detector be minimized because of the high scanning speeds involved. Another field where the invention can provide significant benefit is intra-oral dental imaging. A particularly advantageous material for the intermediate layer is Mylar (Trade Name). This has an effective atomic number of approximately 10. Silicon has an atomic number of 14. A preferred scintillator material is gadolinium oxysulphide which has an effective atomic number of approximately 64. Other scintillator materials may be used instead. Another material which may be used for the intermediate layer is polyimide. According to a second aspect of the invention there is provided an arrangement for detecting high energy radiation comprising: a layer of scintillator material for converting incident high energy radiation into optical radiation; a solid state detector having regions at which charge is generated which is representative of incident optical radiation and having areas which are sensitive to damage by ionizing radiation; and an intermediate layer located between the scintillator and the damage sensitive areas and adjacent the sensitive areas, the intermediate layer having an atomic number which is lower than that of adjacent sensitive areas, and in which materials located in front of the areas are such that a sufficient proportion of incident high energy radiation reaches said areas that a dose of 10 krads or more would be deposited at the areas in the absence of the intermediate layer. The invention may be used where a dose of 200 krads or more, that is two orders of magnitude greater than what would normally be considered acceptable for a non- radiation hardened device, would be deposited in the sensitive areas if the intermediate layer were absent. Although the same proportion of ionizing radiation reaches those areas its damaging effect is substantially reduced by the enhanced dose effect provided by use of the intermediate layer. Doses of 300 krads or more may be handled using the invention where the detector is a standard production device with no radiation hardening. For devices which are radiation hardened, use of the invention will increase the amount of radiation exposure they can tolerate. The solid state detector may be a CCD or a photodiode array, for example. In the latter case, the areas vulnerable to damage by ionising radiation are laterally positioned with respect to regions where optical radiation is received. Hence, they may be shielded in some manner from high energy radiation whilst leaving the detecting regions exposed. However, by employing the invention, separate shielding, which may require several processing steps to fabricate, may be eliminated. By employing the invention, significant doses of radiation may be deposited in the sensitive areas without adversely affecting performance. Thus the thickness of the scintillator layer, and any other materials included between it and the intermediate layer, need not be constrained to absorb all high energy radiation before it reaches the sensitive areas. This permits the scintillator thickness to be optimised for satisfactory light conversion efficiency. It also may allow one particular detector arrangement to be used with a greater range of high energy radiation than might otherwise be the case. Some ways in which the invention may be performed are now described with reference to the accompanying drawings, in which: Figure 2 schematically illustrates a detector in accordance with the invention; and Figures 3 and 4 are explanatory diagrams relating to the operation of the detector shown in Figure 2. With reference to Figure 2 which is not drawn to scale, a photodiode array suitable for detecting high energy radiation, in this case X-rays, has a layer of scintillator material 1 , gadolinium oxysulphide, of about 1 mm thickness laid down on an intermediate layer 2 of Mylar. The Mylar is substantially optically transmissive and in this embodiment of the invention, which is intended to be used with an X-ray source having a spectrum of X-rays with a maximum energy of the order of 120 keV, is approximately 150 μm thick. A photodiode detector array 3 is located adjacent to and behind the intermediate layer 2. The detector region is located at the surface of a silicon substrate at its boundary with the intermediate layer 2 and includes switching FETs and other vulnerable structures at its surface 3 A. The detector is intended to be used in a CT scanner and has pixel sizes of the order of 1 mm. The scintillator material has a high effective atomic number, being approximately 64, the Mylar has an effective atomic number of approximately 10, and the silicon and silicon dioxide at the boundary with the intermediate layer 2 have an atomic number of approximately 14. In Figure 3, the three regions 1, 2 and 3 are indicated with the abscissa showing the distance through the detector and the ordinate giving the energy per kilogram of the secondary electrons present in each region. The thickness of the Mylar layer 2 is chosen such that it is sufficient to prevent secondary electrons produced at the boundary of the scintillator material 1 and the Mylar from reaching the detector array 3. The effective atomic number of the Mylar is below that of the adjacent damage sensitive areas 3A and energy is therefore transferred from the sensitive layers of the detector array 3 to the Mylar buffer intermediate layer 2. Figure 4 schematically illustrates relative dose deposited in a sensitive silicon dioxide layer vulnerable to damage by radiation and underlying silicon for a case where no intermediate layer is included and an embodiment in accordance with the invention using a Mylar intermediate layer. Ionizing radiation is permitted to reach the damage sensitive areas but use of the invention reduces the relative dose to an acceptable level. In another embodiment of the invention, the arrangement of Figure 2 is used but the detector array is a CCD array and the intermediate layer is polyimide. The polyimide layer 2 is of approximately 3 microns thickness, or greater, and the detector arrangement is for use in a dental imaging system. An X ray source of suitable energy level is directed to irradiate a patient's jaw and the detector is used intra orally to detect radiation after passage through the jaw. The thickness of the intermediate layer may be selected so as to be approximately half the lateral dimension of the pixels to give good spatial resolution.

Claims: CLAIMS 1. A CCD arrangement for detecting high energy radiation comprising: a layer of scintillator material for converting incident high energy radiation into optical radiation; a CCD for detecting optical radiation and having regions which are sensitive to damage by ionizing radiation; and an intermediate layer located between the scintillator and the detector regions and adjacent the detector regions, the intermediate layer being substantially transmissive to optical radiation and having an atomic number which is lower than that of the detector regions. 2. An arrangement for detecting high energy radiation comprising: a layer of scintillator material for converting incident high energy radiation into optical radiation; a solid state detector having regions at which charge is generated which is representative of incident optical radiation and having areas which are sensitive to damage by ionizing radiation; and an intermediate layer located between the scintillator and the damage sensitive areas and adjacent the sensitive areas, the intermediate layer having an atomic number which is lower than that of adjacent sensitive areas, and in which materials located in front of the areas are such that a sufficient proportion of incident high energy radiation reaches said areas that a dose of 10 krads or more would be deposited at the areas in the absence of the intermediate layer. 3. An arrangement as claimed in claim 2 wherein a sufficient proportion of incident high energy radiation reaches said areas that a dose of 200 krads or more would be deposited at the areas in the absence of the intermediate layer. 4. An arrangement as claimed in claim 2 or 3 wherein the solid state detector is a CCD and said areas are at regions arranged to receive and detect optical radiation. 5. An arrangement as claimed in claim 2 or 3 wherein the solid state detector is a photodiode array and said areas include switching means. 6. An arrangement as claimed in any preceding claim wherein the solid state detector is non-radiation hardened. 7. An arrangement as claimed in any preceding claim wherein the intermediate layer is of polyimide. 8. An arrangement as claimed in any of claims 1 to 6 wherein the intermediate layer is of Mylar (Trade Name). 9. An arrangement as claimed in any preceding claim wherein the scintillator material has an atomic number which is higher than that of the intermediate layer. 10. An arrangement as claimed in any preceding claim wherein the scintillator material is gadolinium oxysulphide. 11. An arrangement as claimed in any preceding claim wherein the intermediate layer has a thickness of lOμm or less. 12. An arrangement as claimed in any preceding claim wherein the detector is pixellated and the thickness of the intermediate layer is approximately half the lateral dimension of the pixels. 13. An arrangement as claimed in any preceding claim adapted for use in dental X-ray imaging. 14. An arrangement comprising a source of high energy radiation and a CCD arrangement as claimed in any preceding claim except when dependent on claim 5 located to receive radiation after it has passed through a body irradiated by the source. 15. A method of imaging high energy radiation including: providing a detecting arrangement which comprises: a layer of scintillator material for converting incident high energy radiation into optical radiation; a solid state detector having regions at which signal charge is generated which is representative of incident optical radiation and having areas which are sensitive to damage by ionizing radiation; and an intermediate layer located between the scintillator and the damage sensitive areas and adjacent said areas, the intermediate layer having an atomic number which is lower than that of adjacent sensitive areas; and arranging for a sufficient proportion of incident high energy radiation to be incident at the sensitive areas of the detecting arrangement such that a dose of 10 krads or more would be deposited at the areas in the absence of the inteπnediate layer. 16. A method as claimed in claim 15 wherein the detecting arrangement is a photodiode array. 17. A method as claimed in claim 15 wherein the detecting arrangement is a CCD array. 18. An arrangement for detecting high energy radiation substantially as illustrated in and described with reference to the accompanying drawings.

Inventors: ROBBINS MARK STANFORD, ROBBINS, MARK, STANFORD
Applicants: EEV LTD, ROBBINS MARK STANFORD, EEV LIMITED, ROBBINS, MARK, STANFORD
Application year: 1997
Main classification: G01T1
IPC codes: G01T1