WO2012034178A1

WO2012034178A1 - Radiation detector method and apparatus

Info

Publication number: WO2012034178A1
Application number: PCT/AU2011/001185
Authority: WO
Inventors: Anatoly Rozenfeld; Michael Lloyd Franz Lerch; Marco Petasecca
Original assignee: University Of Wollongong
Priority date: 2010-09-17
Filing date: 2011-09-16
Publication date: 2012-03-22

Abstract

A radiation detector is disclosed that comprises a scintillator that emits electromagnetic radiation in response to excitation by radiation of interest, and a photodetector with a semiconductor active layer adapted to interact with the electromagnetic radiation and a substrate. Either the scintillator constitutes the substrate or the substrate is located such that the substrate is between said active layer and the scintillator, and the substrate is at least partially transparent to the electromagnetic radiation.

Description

RADIATION DETECTOR METHOD AND APPARATUS

FIELD OF THE INVENTION

This invention is related to a radiation detector method and apparatus, and in particular a scintillator based radiation detector, of particular but by no means

exclusive application in providing a PET detector module for coincidence detection of annihilation photons emitted by a source of radiation and in PET imaging.

BACKGROUND OF THE INVENTION

Scintillator based detectors use scintillator materials to convert energy deposited from radiation to optical photons , followed by detection of the photons by

photodetectors . Generally efficient gamma-ray detectors use high Z scintillators optically coupled to a

photomultiplier (PM) . PMs have the advantage of internal amplification and hence a good signal-to-ratio (SNR) . PMs also offer the possibility of a fast signal outcome for use in timing applications where events are triggered on the rising edge of a pulse received from a photodetector. A disadvantage of PMs is their large size and high voltage requirements (800-1200 V) . Also, this type of detector can be expensive when a large scintillator area is required, as is the case with homeland security portals used by customs services and the like. Additionally, the performance of PMs is affected by magnetic fields and this prevents their application in the proximity of a magnetic resonance imaging (MRI ) scanner .

Another scintillator based radiation detector uses semiconductor photodiodes . Photodiodes have certain advantages over PMs : they are smaller and require lower bias voltages in operation, but they are slow and do not have internal amplification. Silicon photomultipliers (SiPMs) , which are operated in Geiger mode for a similar internal amplification as PMs, have opened new possibilities for fast timing and

coincidence applications, relevant in PET instrumentation. A SiPM has a multi-pixel configuration; each pixel comprises an avalanche diode that is triggered by the absorption of a photon, which generates charge and an increased current through the SiPM producing a voltage drop across a series resistor. The total output pulse current is proportional to the number of activated pixels and the generated current rises extremely quickly until it is stopped by a quenching process .

A timing resolution for 511 keV photons of up to 8-10 ns has been reported, when using a two pixel configuration of

2

SiPMs, each with an active area 3x3 mm and optically coupled to YLSO scintillators and fast readout

electronics . An array of SiPMs with limited pixels up to a 4x4 matrix are available. The aforementioned SiPMs are produced on bulk Si wafers.

The major disadvantage of SiPMs operated in Geiger mode is the strong link of their amplification coefficient to the temperature and to the applied bias voltage. Furthermore, there is a non-linearity of response versus the energy deposited by radiation in a scintillator, owing to pile-up from a number of sub-pixels .

An avalanche photodiode (APD) with internal amplification is a related Si based photodetector , and may be compared to a single pixel SiPM. An APD employs an avalanche process, which is initiated by optical photon absorption. Within the Si region a strong build-up of the electrical field occurs which is produced by a special implantation process for creation of a p+-p region under the surface of the diode . The disadvantage of avalanche photodiodes is in the non-uniform avalanche response along the diode's area of sensitivity , whereby there is a preferential response on the sharp edges of the p-n junction where the electrical field is stronger . This phenomenon leads to non-uniform sensitivity of the active area of the

avalanche photodiode and an overall loss of efficiency.

Single SiPM or avalanche photodiodes are not widely used in medical imaging applications and are limited in use with large area radiation detector applications.

U.S. Patent No. 7,759,650 discloses an improved APD and a method of fabrication for the photodiode to overcome aforementioned problems . The disclosed method utilises a standard CMOS technology for production of an APD on a thick Si layer (s) formed on a substrate. This is said to allow production of an array of APDs using a trench formation within the semiconductor layers , with the trenches filled with material that reduces curvature of the edge of the p-n junction. This is intended to improve uniformity of the avalanche response along the pixel's sensitive area and increase the photodetector optical photon detection efficiency . Another advantage of CMOS technology for production of an APD array is the

possibility of integrating readout electronics for each pixel into the same silicon layer (s) and reducing the package size for an array of photodetectors . This document also describes the possible optical coupling of a pixellated scintillator directly above the APDs pixels . US Patent No. 5,464,984 describes an array of

photodetectors without internal amplification that is similar to a charge-coupled device (CCD) covered from one side by scintillator material and the other side available for bump bonding of sensor arrays to a processing circuit. In such an embodiment, an array is produced on a Silicon- on-Insulator (SOI) substrate that is transparent to the scintillator' s light response and allows placement of scintillator or phosphorous on the back of the SOI while a plurality of processing circuits are bump bonded from the other side. This approach is aimed at X-ray intensity measurements only and where related to 2D diagnostic imaging where spectroscopy of X-ray is not required.

A passive Si p-i-n photodiode optically coupled to a scintillator has advantages for gamma-ray spectroscopy as an alternative to a PM. The advantages of this technique are similar to those of the SiP in being small and hence allowing a compact detector system (especially

advantageous in packaging multiple detector modules within PET and SPECT systems) . Pixellated p-i-n diode arrays, however, are easier to produce than are SiPMs . The disadvantage of p-i-n diodes in comparison to SiPMs and PMs are the absence of internal amplification, so their timing properties are less attractive when used in coincidence timing (owing to reduced SNR) . The lower SNR limits the performance of a passive p-i-n photodiode as noise levels are comparable to or exceed the detection threshold for spectroscopy of low energy X-ray sources . Improvement of the SNR is a key issue to their wider application . Improvements in SNR are possible by reducing the

capacitance of the p-i-n diode and can be achieved by utilizing high resistivity Si and increasing the reverse voltage bias for a full depletion of the p-i-n diode.

Normally this leads to a Si wafer of about 300-400 microns thickness which achieves capacitance for a fully depleted diode of 3-4 pF/pixel (based on a pixel size 3x3 mm typical for PET and SPECT imaging applications) . A disadvantage of this approach is the relatively high voltage bias needed (50-60 V) , not much lower than is required by a SiPM. There is also an increased noise level owing to higher reverse current. Another technical problem arises if integrating readout electronics onto the same high resistivity Si wafer using standard CMOS technology. For a readout array of Si p-i-n diodes optically coupled to a scintillator, wirebonding or flip chip bump connecting technology is necessary for connection to an ASIC readout circuit.

Further complications are found when optical coupling the scintillator directly to the Si surface of the p-i-n diode using optical compounds; degradation of the p-i-n diode surface over time and an increase in leakage of the diode causing diminishing SNR and energy resolution. This problem prohibits the wide adoption of a p-i-n diode based detector modules within PET and SPECT systems where the replacement of detector modules would be associated with maintenance downtimes and serious maintenance costs .

Most high Z scintillators (YLSO, NAI BGO , LSO, etc) used in gamma spectroscopy and imaging applications such as PET and SPECT produce photons with wavelengths in the range

380-420 nm. These photons are absorbed at the surface of the p-i-n diodes within a depth of 0.1-0.2 microns. The high doped p+ or n+ implanted regions operate at a near zero electric field, which prohibits the charge collection created by the absorbed photons. This has limited the application of p-i-n diodes optically coupled to high Z scintillators in gamma spectroscopy and, in particular, in PET. A method of creating the p-n junction on the surface of the p-i-n diode to improve charge collection has reportedly achieved an energy resolution for 511 keV gamma photons of at best 12-13% (Lerch et al . , Spectral

Characterization of blue-enhanced silicon photodiode, IEEE Trans, on Nucl . Sci . , 48, N-4 (2001) 1220-1224). More detail on the advantages , and the limitations , of PET modules based on PM and Si photodiodes is provided by Humm et al. in From PET detectors to PET scanners, Eur. J.

Nucl. Med. Mol . Imaging, 30 (2003) 1574-1597. SUMMARY OF THE INVENTION

According to a first broad aspect, the present invention provides a radiation detector, comprising:

a scintillator that emits electromagnetic radiation in response to excitation by radiation of interest (such as gamma-ray radiation) ; and

a photodetector with a semiconductor active layer (such as of silicon) adapted to interact with said

electromagnetic radiation and a substrate;

wherein either said scintillator constitutes said substrate or said substrate is located such that said substrate is between said active layer and said

scintillator, and said substrate is at least partially transparent to said electromagnetic radiation.

The radiation detector of the present invention can provide, such as for PET or gamma spectroscopy (for use in, say, security applications) a spectroscopic grade performance. For example, it can provide an energy threshold in PET of approximately 350 keV for events acquisition . Also , it can provide a gamma camera mode with an energy window for the different isotopes of interest dependent on application . In security

applications, it may be desired to identify isotopes from NORM (Naturally Occurring Radiation Materials) and, in particular, weapons grade isotopes such as U-235 and Pu- 239 with lines at energies of, respectively, approximately 190 keV and 400 keV. In some security customs

applications, count mode alone is sufficient to determine increased radiation compared to the background, that is, by measuring gamma intensity without energy resolution .

It will also be appreciated that the present invention has application in the detection of other particles, such as heavy ions (when using scintillator materials such as Csl doped with thallium to emit wavelengths of approximately 550 run) , neutrons and electrons .

In one embodiment, the substrate is substantially

transparent to a principal, average, significant or representative electromagnetic radiation received from said scintillator following emission by said scintillator in response to excitation of said scintillator by

radiation of interest.

In a particular embodiment, the substrate comprises sapphire .

In one embodiment, the scintillator comprises a high Z scintillator material (such as LSO, YLSO or LaBr) . For example, the scintillator may emit photon spectra with an average wavelength of 420 micron.

In another embodiment, the scintillator constitutes said substrate and the active layer is optically coupled to the scintillator (such as composed of inorganic material) with or without an optical interface layer between the active layer and the scintillator. For example, this may be achieved by growing an active layer of silicon or another suitable semiconductor material onto the scintillator (whether or not provided with an optical interface layer) . Such a structure may be achieved by epitaxial grown of the active layer above , for example, an inorganic scintillator with high Z and large volume suitable for effective detection of gamma photons . The advantages of this approach may include improved quantum efficiency (QE) of the photodiode and better SNR of detector-scintillator module and/or cost of

fabrication. The photodiode may be produced by

evaporation of Shottky metal contact on the front surface of the active layer thereby avoiding the need for high- temperature processing, which could damage the inorganic scintillator. For example, the detector in such an embodiment may comprise an InGaAsP pin diode epitaxially grown on the surface of a thin (e.g. 0.5 mm thick) InP scintillator wafer; lacking a large volume scintillator, this arrangement does not have high efficiency and is unlikely to be suitable for gamma-ray spectroscopy, but should be suitable for short range charged particles or low energy X-rays that can be detected by bulk

semiconductor photodiodes directly.

In a certain embodiment, the scintillator is polished on the side optically coupled to the substrate and has (such as by mechanical treatment) diffusely reflective other surfaces leading to diffuse reflection of photons emitted by the scintillator. The scintillator may be covered with a diffuse reflector (such as T1O₂ paint provided by BICRON (trade mark)) on all other sides. In one embodiment, the front surface of the scintillator (i.e. that through which incident radiation enters the scintillator) is provided with a reflective coating that is reflective to photons emitted by the scintillator. In a particular embodiment, the detector includes a reflective layer or coating located such that said active layer is between said reflective layer or coating and said substrate, adapted to reflect photons that pass undetected through said active layer from said scintillator back into said active layer.

In a certain embodiment, the detector comprises a semi- reflective layer or coating between the substrate and the scintillator, adapted to reflect into said active layer photons reflected back into said active layer by said reflective layer or coating (that is, if received by the semi-reflective layer or coating after passing undetected through the active layer) .

In a particular embodiment, the semi-reflective layer or coating is substantially more transparent to

electromagnetic radiation received from said scintillator than to the photons reflected back into said active layer by said reflective layer or coating. In one embodiment, the detector comprises an optical interface layer between said photodetector and said scintillator that optically couples said photodetector and said scintillator. The optical interface layer may comprise a coating or optical paste. In a particular embodiment, the optical interface layer provides

refractive index matching between the scintillator and the substrate .

In a particular embodiment, the detector comprises electronics located on the active layer. The electronics may comprise, for example, readout electronics, processing electronics or combined readout and processing

electronics . Thus, a radiation detector can be constructed according to the invention with low noise and improved SNR, of

particular value in coincidence timing application in - for example - PET . The geometry of certain embodiments avoids the degradation of the photodetector from direct attachment to its surface of the scintillator (thus avoiding mechanical damage, the effects of optical paste on the surface properties of some photodiodes , and degradation over time) , can provide improved functionality by including built-in readout/processing electronics on each pixel, can allow a reduction in the voltage bias necessary for the detector to as low as 3-5 V, and can increase packaging flexibility (such as for deployment in a range of imaging applications) .

In one embodiment, the radiation detector comprises a plurality of photodetectors each with a semiconductor active layer, wherein the substrate is common to said plurality of photodetectors .

In this embodiment, the substrate may have grooves or trenches, optionally containing optically isolating material, adapted to minimize or avoid crosstalk between pairs of said photodiodes . The groove or trench may comprise a material (such as T1O₂ paint) that is optically reflective to photons emitted by the scintillator, so that such photons - if not yet detected - are reflected back into the photodiode .

In one embodiment, the radiation detector comprises a plurality of photodetectors each with a semiconductor active layer, wherein the substrate comprises a discrete substrate portion for each respective of said

photodetectors .

In this embodiment, the detector may include optically reflecting material between the discrete substrate portions, adapted to minimize or avoid crosstalk between pairs of said photodiodes .

In one embodiment, the photodiode is a planar photodiode.

In a certain embodiment, the photodiode includes doped p+ and n+ regions of the photodiode configured such that a portion of the active layer separates the p+ and n+ regions and the substrate.

In some embodiments , the detector is adapted for mounting such that radiation of interest is received by the scintillator from a source located opposite the active layer (and hence propagating generally, and disregarding scattering, perpendicularly to a plane of the substrate) . In other embodiments , however , the detector is adapted for mounting such that radiation of interest is received by the scintillator from a source located beside the

scintillator (and hence propagating generally parallel to a plane of the substrate) . Other configurations and orientations in use are, as will be appreciated, entirely possible.

According to this aspect, the invention also provides a method of detecting radiation, comprising:

locating a radiation detector as described above so as to receive radiation of interest from a source of radiation (such as gamma-ray radiation) .

According to a second broad aspect, the present invention provides a radiation detector module, comprising one or more radiation detectors as described above.

In one embodiment, the module comprises a regular array (such as an 8x8 array) of said radiation detectors. In one embodiment, the module is adapted to be mounted such that radiation of interest is received by the scintillators from a source located generally beside the scintillators (and hence propagating generally parallel to a plane of each respective substrate) .

According to a third broad aspect, the present invention provides a radiation detection system, comprising:

one or more radiation detector modules as described above .

In a particular embodiment, the modules are mounted such that radiation of interest is received by the scintillators from a source located generally beside the scintillators (and hence propagating generally parallel to a plane of each respective substrate) . For example , the radiation detection system may be a gamma spectroscopy detector, a PET system or a SPECT system.

In one embodiment, the system is a PET system with a detector ring comprising a plurality of said modules . In this embodiment, the system may have Depth-of-Interaction (DOI) capability.

According to this aspect, the present invention provides a method of detecting radiation, comprising:

locating a radiation detection system as described above so as to receive radiation of interest from a source of radiation (such as gamma-ray radiation) .

This may entail locating a source of radiation (such as of annihilation radiation) within a detector ring of the system.

According to a fourth broad aspect, the present invention provides an imaging system, comprising a radiation detection system as described above.

In a particular embodiment, the imagining system is a PET system. According to this aspect, the present invention provides an imaging method, comprising:

employing an imaging system as claimed above to form an image . It should be noted that any of the various features of each of the above aspects of the invention can be combined as suitable and desired. BRIEF DESCRIPTION OF THE DRAWING

In order that the invention may be more clearly

ascertained, embodiments will now be described, by way of example, with reference to the accompanying drawing, in which :

Figure 1 is a schematic view of a radiation detector module according to an embodiment of the present invention ;

Figure 2 is a schematic rear view of the radiation detector module of figure 1 ;

Figure 3 is a schematic, cross sectional view of the structure of an individual pixel of the module of figure 1 ;

Figure 4 is a schematic view of a detail of the module of figure 1 ;

Figure 5 is a simplified schematic view of a PET detector ring of a PET scanner according to an embodiment of the present invention; and

Figure 6 is a schematic, cross sectional view of the structure of an individual pixel of a radiation detector module according to another embodiment of the present invention. DETAILED DESCRIPTION

A radiation detector module according to an embodiment of the present invention is shown schematically at 10 in figure 1. Detector module 10 comprises an array of pixels each comprising a lateral planar photodiode 12 and corresponding scintillator element 14 of a high Z

scintillator material (such as LSO, YLSO or LaBr) for emitting photon spectra with an average wavelength of about 420 micron. In the illustrated example, module 10 is shown as a 5x5 array, but it will be appreciated that other sized arrays are entirely possible according to the present invention (and indeed an 8x8 array may be more typically used) , as well as single pixel detectors . Indeed, module 10 is suitable for use in large radiation detector panels with very low power consumption and high quantum efficiency. Although in the illustrated

embodiment the scintillator is pixellated into individual scintillator elements, this may not be necessary in certain applications in which energy resolution need not be as high. Figure 2 is a schematic rear view of module 10,

illustrating the configuration of lateral photodiodes 12. (Although the rear of module 10 is shown, the visible faces of photodiodes 12 are - in conventional photodiodes - regarded as their front faces . )

Each photodiode 12 is provided with its own readout and processing electronics 22, embedded into each respective photodiode 12 (effectively in the corner of the each pixel) . The charge collected from each pixel is thus processed on that respective pixel, so this configuration may be referred to as Active Pixel Technology . Placing readout and processing electronics 22 on the Si wafer that constitutes the active Si layer of the photodiodes 12 (described further below) affords a cheaper

implementation, as well as a more compact detector (since external electronics are not employed or required) , more flexible packaging options and higher readout speeds .

Figure 3 is a schematic, cross sectional view of the structure of an individual pixel 30, comprising lateral photodiode 12 and scintillator element 14. (Note that the thicknesses of the layers of the structure are not shown to scale.) From front (i.e. scintillator side) to rear, pixel 30 comprises scintillator element 14, an optical interface layer 32, and photodiode 12. Photodiode 12, which is fundamentally of a SOI structure, comprises a semi-transparent substrate 34, an active Si layer 36 (provided with an N+ strip 38 and a P+ strip 40 that form a p-i-n diode) , a reflective coating 42 (deposited on the front surface of photodiode 12 and hence located at the rear of pixel 30) and two aluminium contacts 44a, 44b projecting from the front of photodiode 12 (and hence rear of pixel 30) and electrically coupled, respectively, to N+ strip 38 and P+ strip 40.

Gamma-ray photons to be detected impinge and interact with scintillator element 14, which emits photons λ. At least some of photons λ propagate through optical interface layer 32 and substrate 34 into active Si layer 36 for detection . Substrate 34 is semi-transparent so that photons can enter active Si layer 36 through substrate 34, rather than - as conventionally - from the opposite (or ^Λ front') face of photodiode 12 (which, in a conventional design, would comprise an active layer without a reflective coating) . This configuration means that readout and processing electronics 22 are behind (from the perspective of the photons λ) the active volume of photodiode 12, so do not block those photons from detection . Optical interface layer 32 is composed of a coating or optical paste, and is used to match the refractive index between scintillator element 14 and substrate 34. This is to maximize the coupling of photons λ into active Si layer 36 which, having a thickness of the order of 0.1 micron, absorbs only some of the photons λ (of wavelength around 420 nm in the example of a YLSO scintillator) . If substrate 34 is selected so as to be close to 100% transparent to photons λ , the quantum efficiency of detector module 10 will be limited essentially only by photon absorption efficiency of active Si layer 36, despite its thinness . For example, at the wavelength of blue light (420 micron) sapphire is essentially 100% transparent and active Si layer 36 is about 70% efficient in absorbing photons, so - with a substrate 34 of sapphire - about 70% of photons λ will be absorbed over the thickness of active Si layer 36. About ~30% will pass through active Si layer 36

unabsorbed .

The absorption efficiency of the thin active Si layer 36 is augmented, however, by reflective coating 42, which reflects unabsorbed photons back into active Si layer 36 to increase the SNR by ~30% (though it will be appreciated that some photons will still escape detection in active Si layer 36 even on this second pass) .

Thus, reflective coating 42 directs photons that fail to interact with active Si layer 36 on first pass back (as shown at X_R) into active Si layer 36, increasing the probability of their detection and hence quantum

efficiency and SNR, owing to their again entering active

Si layer 36. Reflective coating 42 is selected or adapted to reflect the particular wavelength (s) of interest after emission by scintillator element 14 (e.g. 420 nm for YLSO, as used for PET) and passage through active Si layer 36. Reflective coating 42 may comprise Al or Si0₂ with a thickness λ/4η (assuming that the refractive index of reflective coating 42 is higher than that of active Si layer 36) , or λ/2η (where the refractive index of

reflective coating 42 is lower than that of active Si layer 36) , where n is the reflectivity index of reflective coating 42. In any event, the reflective coating has an optical length such that photons reflected from the reflective coating-air interface and from the active layer-reflective coating interface form a constructive interference in the active layer. S1O₂ , for example , may be deposited by known techniques (such as are used in microelectronics) . A coating 42 of S1O₂ would act as a passive layer only in respect of readout and processing electronics 22.

Various materials are satisfactory for substrate 34 , but - whatever material is used - substrate 34 should be thick enough to have sufficient mechanical rigidity to function as a substrate while, advantageously, minimizing the attenuation of photons of interest from scintillator element 14. Consequently, scintillator material will be chosen according to application and substrate material . In this embodiment, substrate 34 comprises sapphire

(which, at the wavelength of blue light - 420 micron - is essentially 100% transparent) , with a thickness of 400-500 micron (though other thicknesses meeting the above criteria would also be satisfactory) . The combination of Si layer 36 and substrate 34 may thus be described as a ^Λ silicon-on-sapphire' (SOS) configuration. In other embodiments, however, with different scintillator

materials , different semi-transparent substrate materials may be more appropriate .

This arrangement, in which scintillator element 14 is mounted - in effect - behind photodiode 12, avoids problems associated with mounting a scintillator above photodiode circuitry (which can damage the photodiode' s electronics) . It also avoids the need to introduce a protective layer above the photodiodes .

The depth of active Si layer 36 is in the range 0.1 to 0.15 microns (and in this embodiment 0.1 micron) . This structure operates with a very low capacitance owing to the very thin active Si layer 36, which is advantageous for the envisaged timing applications. Indeed, such a thin layer of Si yields a very low parasitic capacitance (effectively zero) relative to substrate 34, and low p-i-n photodiode 12 capacitance (approximately 0.06 pF) owing to the thin active Si layer 36 and use of lateral p-i-n photodiode 12. (This is in contrast to high resistivity Si photodiodes with Si layer thickness around 380 micron and capacitances of 3 pF/3x3 mm pixel when fully

depleted. ) The low capacitance p-i-n photodiodes 12 render readout and processing electronics 22 low in noise allowing high signal-to-noise ratio (SNR) detection and outstanding timing.

The low thickness of active Si layer 36 nonetheless provides sufficient absorption of the majority of photons: as mentioned above, for photons of 420 nm wavelength, the 0.1 micron active Si layer 36 is approximately 70% efficient in absorbing photons. Indeed, increasing the thickness of active Si layer 36 to 0.15 or 0.2 micron is not beneficial for the technology used in fabrication of readout electronics 22 and requires a higher reverse voltage to be applied on photodiode 12.

It will be appreciated by those skilled in the art that the thickness of the active Si layer may vary in

accordance with the wavelength of photons emitted by the scintillator material , while still retaining the key advantage of a thin active Si layer with significantly reduced capacitance .

It will also be noted that the implantation of the N+ and P+ strips 38, 40 to a depth less than that of active Si layer 36 allows use of 100% of the sensitive volume of lateral photodiode 12, as photons λ from scintillator element 14 have a significant likelihood of detection in active Si layer 36 before reaching the N+ strip 38 or the P+ strip 40. Similarly, owing to the reversed geometry of photodiode 12, these photons reach the active Si layer 36 before reaching contacts 44a, 44b. Thus, the efficiency of photodiode 12 is increased compared with designs in which scintillator element 14 is located on the front of photodiode 12 , whereby the aluminium covered electrodes can reflect light and hence not reach the scintillator. In use, a small reverse voltage (3-4 V) is applied to achieve full p-i-n diode depletion, and in some

embodiments an operating voltage of only 1-2 V is

possible. Active Si layer 36 of photodiode 12 absorbs photons of wavelength 300-450 nm and responds by

generating electrical charges that are transmitted via aluminium contacts 44a, 44b to readout and processing electronics 22.

Figure 4 is a schematic view of a detail 50 of detector module 10. Illustrated are exemplary lateral photodiodes

2

52a, 52b, 52c (each 3x3 mm ) that are optically coupled at their substrate sides to respective scintillator elements 54a, 54b, 54c; scintillator elements 54a, 54b, 54c are mutually optically isolated and matched in size to respective, corresponding photodiodes 52a, 52b, 52c.

Thus , gamma-ray photon interaction in a particular scintillator element 54a, 54b, 54c produces photons for subsequent detection by the correspondingly adjacent photodiode 52a , 52b , 52c .

As alluded to above, module 10 comprises a single Si wafer on sapphire (SOS) 56, respective portions of which

constitute the respective photodiodes 52a, 52b, 52c.

Trenches 58 are provided in the substrate portion of SOS 56; these trenches 58 delineate the respective photodiodes 52a, 52b, 52c and are filled with optically isolating material to reduce or avoid crosstalk between the

respective photodiodes 52a, 52b, 52c. Optionally, trenches 58 may be filled with material that is optically reflective to photons (such as T1O₂ paint) that are not yet detected by the Si active layer 36, to reflect such photons back into photodiode 12. Trenches 58 can be produced by laser, and optimally are as deep as possible while not unduly compromising the mechanical integrity of the structure. It should be noted that, in this

embodiment, no trench is provided within active Si layer 36, and photodiodes 52a, 52b, 52c are not operated as an avalanche diode driven in Geiger mode .

Scintillator elements 54a, 54b, 54c are polished on the side optically coupled to the substrate and are mechanical treated on their other surfaces 59 leading to diffuse reflection of the photons inside of the scintillator crystal . The scintillator is covered with a diffuse reflector (such as T1O₂ paint provided by BICRON) on all other sides 59.

A coating reflective to photons emitted by scintillator elements 54a, 54b, 54c is provided on the front surface (not shown) of scintillator elements 54a, 54b, 54c, that is , the surface through which the gamma rays enter the scintillator elements 54a, 54b, 54c. This coating reflects at least some of these photons back towards respective photodiodes 52a, 52b, 52c for detection.

Module 10 is suitable for use in many applications, including for example PET imaging. Figure 5 is a

simplified schematic view of a PET detector ring 60 of a PET scanner according to an embodiment of the present invention, with a plurality of detector modules 10 arranged edge-on. (In reality, an essentially complete ring of detector modules 10 would be provided.) This arrangement allows Depth-of-Interaction (DOI) measurements with accurate tracking of the line of response of

coincident events in pairs of detector modules (such as 10' and 10") ; this allows the accurate determination of the point of interaction of pairs of gamma-ray photons emitted in opposite directions (following an annihilation event) from the sample under investigation. It should be noted that the pair of detector modules that detect a coincident event need not be diametrically opposite each other in ring 60 , but rather are on opposite sides of the event. Consequently, accurate image reconstruction is possible independently of the positioning of the object within the field of view of the scanner, so image quality is uniform across the field of view.

A radiation detector module according to another

embodiment of the present invention is shown schematically at 70 in figure 6. Detector module 70 is identical in most respects to detector module 10 of figure 1, and like reference numerals have been used to indicate like

features .

However, detector module 70 includes an additional layer in the form of a semi-reflective coating 72 between optical interface layer 32 and substrate 34. If photons X_R reflected by reflective coating 42 again pass through active Si layer 36 unabsorbed, some photons \_R> will be reflected by semi-reflective coating 72 back into active Si layer 36, further increasing their likelihood of absorption and hence detection. It will be appreciated by those skilled in this art that the inclusion of semi- reflective coating 72 may require the selection of a different optical interface layer 32 (owing to the changed optical coupling requirements between scintillator element 14 and active Si layer 36) . Also, it is expected that semi-reflective coating 72 would be employed only in cases where either so doing augments the total number of

detected photons or increases the SNR.

The above embodiments thus provide good timing resolution with a passive and inexpensive Si photodiode , and permit edge on packaging of detector modules for Depth-of- Interaction capability. Improvements in SNR are made possible by employing a thin (0.1-0.2 micron) active Si layer leading to a) essentially zero parasitic capacitance between substrate and active elements on the pixel (in the above embodiments, the PIN photodiodes) , and b) an extremely small capacitance of the photodetector (such as a lateral PIN photodiode) , that is approximately 100 times less than would conventionally be the case from a vertical pin photodiode of the same active area located on bulk high resistivity silicon (as

currently used for PET.

Such a low pixel capacitance and parasitic capacitance provided under a small bias voltage (e.g. 2 V) results in a low input capacitance presented to the readout

electronics for each pixel. Moreover a low dark current of the lateral pin diode is achieved. This has the effect of reducing noise generated, both serial and parallel, at the pin diode and allows for its application to

spectroscopy with scintillator without the need for internal amplification (as is the case of a SiPM) and for good timing resolution that is important for PET

applications .

Modifications within the scope of the invention may be readily effected by those skilled in the art. It is to be understood, therefore, that this invention is not limited to the particular embodiments described by way of example hereinabove . In the claims that follow and in the preceding description of the invention, except where the context requires otherwise owing to express language or necessary

implication, the word "comprise" or variations such as "comprises" or "comprising" is used in an inclusive sense, that is, to specify the presence of the stated features but not to preclude the presence or addition of further features in various embodiments of the invention. Further, any reference herein to prior art is not intended to imply that such prior art forms or formed a part of the common general knowledge in Australia or any other country .

Claims

CLAIMS :

1. A radiation detector, comprising:

a scintillator that emits electromagnetic radiation in response to excitation by radiation of interest; and

a photodetector with a semiconductor active layer adapted to interact with said electromagnetic radiation and a substrate;

scintillator, and said substrate is at least partially transparent to said electromagnetic radiation .

2. A detector as claimed in claim 1 , wherein the

substrate is substantially transparent to a principal, average, significant or representative electromagnetic radiation received from said scintillator following emission by said scintillator in response to excitation of said scintillator by radiation of interest.

3. A detector as claimed in claim 1 , wherein the

substrate comprises sapphire.

4. A detector as claimed in claim 1 , wherein the

scintillator comprises a high Z scintillator material .

5. A detector as claimed in claim 1 , wherein said

scintillator constitutes said substrate and the active layer is optically coupled to the scintillator.

6. A detector as claimed in claim 5 , wherein said active layer is grown onto said scintillator .

7. A detector as claimed in claim 1 , wherein the

scintillator comprises LSO, YLSO or LaBr.

8. A detector as claimed in claim 1 , wherein one or more sides of the scintillator are covered with a reflective paint or other reflective coating.

9. A detector as claimed in claim 1 , wherein a front or incident surface of the scintillator is covered with a reflective paint or other reflective coating.

10. A detector as claimed in claim 1, comprising a reflective layer or coating located such that said active layer is between said reflective layer or coating and said substrate, adapted to reflect photons that pass undetected through said active layer from said scintillator back into said active layer.

11. A detector as claimed in claim 1, comprising a semi- reflective layer or coating between the substrate and the scintillator, adapted to reflect into said active layer photons reflected back into said active layer by said reflective layer or coating.

12. A detector as claimed in claim 11, wherein the semi- reflective layer or coating is substantially more

transparent to electromagnetic radiation received from said scintillator than to the photons reflected back into said active layer by said reflective layer or coating.

13. A detector as claimed in claim 1 , comprising an optical interface layer between said photodetector and said scintillator that optically couples said

photodetector and said scintillator.

14. A detector as claimed in claim 13, wherein the optical interface layer comprises a coating or optical paste .

15. A detector as claimed in claim 13, wherein the optical interface layer provides refractive index matching between the scintillator and the substrate.

16. A detector as claimed in claim 1, comprising

electronics located on the active layer.

17. A detector as claimed in claim 16, wherein said electronics comprise readout electronics, processing electronics or combined readout and processing

electronics .

18. A detector as claimed in claim 1, comprising a plurality of photodetectors each with a semiconductor active layer, wherein the substrate is common to said plurality of photodetectors .

19. A detector as claimed in claim 18, wherein the substrate comprises grooves or trenches adapted to minimize or avoid crosstalk between pairs of said

photodiodes .

20. A detector as claimed in claim 19, wherein the grooves or trenches contain optically isolating or reflective material.

21. A detector as claimed in claim 1, comprising a plurality of photodetectors each with a semiconductor active layer, wherein the substrate comprises a discrete substrate portion for each respective of said

photodetectors .

22. A detector as claimed in claim 19, comprising optically isolating material between the discrete

substrate portions, adapted to minimize or avoid crosstalk between pairs of said photodiodes .

23. A detector as claimed in claim 1 , wherein the

photodiode is a planar photodiode .

24. A detector as claimed in claim 1, wherein the

photodiode includes doped p+ and n+ regions of the

photodiode configured such that a portion of the active layer separates the p+ and n+ regions and the substrate .

25. A detector as claimed in claim 1 , adapted for

mounting such that radiation of interest is received by the scintillator from a source located opposite the active layer .

26. A detector as claimed in claim 1, adapted for

mounting such that radiation of interest is received by the scintillator from a source located beside the

scintillator .

27. A method of detecting radiation, comprising:

locating a radiation detector as claimed in any one of the preceding claims , so as to receive radiation of interest from a source of radiation.

28. A radiation detector module, comprising one or more radiation detectors as claimed in any one of claims 1 to

26.

29. A module as claimed in claim 26, adapted to be mounted such that radiation of interest is received by the scintillators from a source located generally beside the scintillators and hence propagating generally parallel to a plane of each respective substrate .

30. A radiation detection system, comprising:

one or more radiation detector modules as claimed in either claim 28 or 29.

31. A radiation detection system as claimed in claim 30, wherein the modules are mounted so that radiation of interest is received by the scintillators from a source located generally beside the scintillators and hence propagating generally parallel to a plane of each

respective substrate .

32. A radiation detection system as claimed in either claim 30 or 31 , wherein the radiation detection system comprises a gamma spectroscopy detector, a PET system or a SPECT system.

33. A radiation detection system as claimed in any one of claims 30 to 32 , wherein the system is a PET system with a detector ring comprising a plurality of said modules.

34. A radiation detection system as claimed in any one of claims 30 to 33, wherein the system has Depth-of- Interaction (DOI) capability.

35. A method of detecting radiation, comprising:

locating a radiation detection system as claimed in any one of claims 30 to 34 so as to receive radiation of interest from a source of radiation .

36. An imaging system, comprising a radiation detection system as claimed in any one of 30 to 34.

37. An imaging system as claimed in claim 36, wherein the imagining system is a PET system.

38. An imaging method, comprising:

employing an imaging system according to claim 37 to form an image .