Classifications

• • H—ELECTRICITY
  • H01—ELECTRIC ELEMENTS
  • H01L—SEMICONDUCTOR DEVICES NOT COVERED BY CLASS H10
  • H01L27/00—Devices consisting of a plurality of semiconductor or other solid-state components formed in or on a common substrate
  • H01L27/14—Devices consisting of a plurality of semiconductor or other solid-state components formed in or on a common substrate including semiconductor components sensitive to infrared radiation, light, electromagnetic radiation of shorter wavelength or corpuscular radiation and specially adapted either for the conversion of the energy of such radiation into electrical energy or for the control of electrical energy by such radiation
  • H01L27/144—Devices controlled by radiation
  • H01L27/1446—Devices controlled by radiation in a repetitive configuration
• • A—HUMAN NECESSITIES
  • A61—MEDICAL OR VETERINARY SCIENCE; HYGIENE
  • A61B—DIAGNOSIS; SURGERY; IDENTIFICATION
  • A61B6/00—Apparatus or devices for radiation diagnosis; Apparatus or devices for radiation diagnosis combined with radiation therapy equipment
  • A61B6/02—Arrangements for diagnosis sequentially in different planes; Stereoscopic radiation diagnosis
  • A61B6/03—Computed tomography [CT]
  • A61B6/037—Emission tomography
• • G—PHYSICS
  • G01—MEASURING; TESTING
  • G01J—MEASUREMENT OF INTENSITY, VELOCITY, SPECTRAL CONTENT, POLARISATION, PHASE OR PULSE CHARACTERISTICS OF INFRARED, VISIBLE OR ULTRAVIOLET LIGHT; COLORIMETRY; RADIATION PYROMETRY
  • G01J1/00—Photometry, e.g. photographic exposure meter
  • G01J1/42—Photometry, e.g. photographic exposure meter using electric radiation detectors
  • G01J1/44—Electric circuits
• • H—ELECTRICITY
  • H01—ELECTRIC ELEMENTS
  • H01L—SEMICONDUCTOR DEVICES NOT COVERED BY CLASS H10
  • H01L31/00—Semiconductor devices sensitive to infrared radiation, light, electromagnetic radiation of shorter wavelength or corpuscular radiation and specially adapted either for the conversion of the energy of such radiation into electrical energy or for the control of electrical energy by such radiation; Processes or apparatus specially adapted for the manufacture or treatment thereof or of parts thereof; Details thereof
  • H01L31/08—Semiconductor devices sensitive to infrared radiation, light, electromagnetic radiation of shorter wavelength or corpuscular radiation and specially adapted either for the conversion of the energy of such radiation into electrical energy or for the control of electrical energy by such radiation; Processes or apparatus specially adapted for the manufacture or treatment thereof or of parts thereof; Details thereof in which radiation controls flow of current through the device, e.g. photoresistors
  • H01L31/10—Semiconductor devices sensitive to infrared radiation, light, electromagnetic radiation of shorter wavelength or corpuscular radiation and specially adapted either for the conversion of the energy of such radiation into electrical energy or for the control of electrical energy by such radiation; Processes or apparatus specially adapted for the manufacture or treatment thereof or of parts thereof; Details thereof in which radiation controls flow of current through the device, e.g. photoresistors characterised by potential barriers, e.g. phototransistors
  • H01L31/101—Devices sensitive to infrared, visible or ultraviolet radiation
  • H01L31/102—Devices sensitive to infrared, visible or ultraviolet radiation characterised by only one potential barrier
  • H01L31/107—Devices sensitive to infrared, visible or ultraviolet radiation characterised by only one potential barrier the potential barrier working in avalanche mode, e.g. avalanche photodiodes

Definitions

• the present invention relates to a detection system for photon sensing and for measuring photon fluxes according to the preamble of claim 1 , a scanner for computed tomography (CT), positron emission tomography (PET), single photon emission computed tomography (SPECT), or portal imaging applications, or any combination of these, according to the preamble of claim 19, use of a photon detection system in PET, SPECT, CT, or radiotherapy portal imaging applications, or any combination of these, and a method for photon sensing and for measuring photon fluxes according to the preamble of claim 21.
• CT computed tomography
• PET positron emission tomography
• SPECT single photon emission computed tomography
• portal imaging applications or any combination of these
• Photon sensing is an important component in many medical and technical applications. Besides optical imaging, there are several types of applications concerned with detecting ionizing radiation where the radiation is converted into optical photons in scintillators and a large dynamic intensity range is needed, such as industrial radiography, nuclear safeguards, environmental radiation monitoring and airport security. Other important ionizing radiation detection applications are diagnostic and therapeutic medical imaging applications.
• Molecular imaging is a rapidly growing modality in medical diagnostics and of particular interest is the recent development and rapid growth of combined computed tomography (CT) and positron emission tomography (PET) systems.
• CT computed tomography
• PET positron emission tomography
• the photomultiplier tubes (PMTs) used together with scintillator crystals to detect X-rays and gamma rays in current state-of-the-art PET scanners and single photon emission computed tomography (SPECT) scanners are about to be replaced with photo detectors based on semiconductor technology, such as avalanche photo diodes (APDs).
• a major rationale behind this ongoing transition to silicon-based devices is that they can operate in a strong magnetic field and, thus, provide the technology for combined CT/SPECT/PET systems and magnetic resonance imaging (MRI) systems.
• the radiation detectors are usually operated in a "photon counting mode" (pulse mode) which means that each signal corresponding to a detected primary photon with the desired properties will generate a pulse that can be used by a data acquisition system, additionally after being processed by fast photon counting electronics, such as preamplifiers, discriminator circuits, etc.
• pulse mode pulse mode
• the radiation detectors consist of APDs coupled to scintillator crystals
• the APDs may be operated in either proportional mode or "Geiger mode".
• proportional mode each APD generates pulses whose magnitudes are proportional to the number of detected photons generated in the scintillation process and hence proportional to the energy deposited by the primary radiation.
• the APDs are usually run at high gain (typically up to ⁇ 1000) in order to optimize the signal to noise ratio and the APD signals are passed on to preamplifier and discriminator circuits designed to select events depending on energy and timing.
• Geiger mode operation is achieved by operating the APDs above their breakdown voltage. When operated in Geiger mode, the internal gain is extremely high (approximately 10 6 - 10 7 ) and the breakdown current caused by a diode avalanche has to be limited to a fixed value by a current limiter ("quenching resistor").
• a current limiter quenching resistor
• each channel of such a system can be regarded as a digital "yes or no” sensor.
• arrays of such detectors sometimes called silicon photomultipliers (SiPMs) or multi-pixel photon counters (MPPCs) have been constructed. These arrays can emulate the linear response of normal photomultiplier tubes and APDs operated in proportional mode by adding the outputs from a large number of individual Geiger mode operated detector pixels.
• APDs operated in single-photoiTcounting mode In applications with high radiation fluxes, such as high-speed X-ray CT imaging or portal imaging systems, utilization of APDs operated in single-photoiTcounting mode is not possible or inefficient due to saturation effects due to dead time and pulse pile -up.
• the maximum count rate per detector element in state-of-the-art single-photon counting systems is well below the mean count rate per unit area required in standard CT imaging.
• APDs can practically be used for rates up to about 10 MHz, mainly limited by the decay time of the scintillator. Therefore, common photodiodes that photovdltaically generate a current that is substantially proportional to the energy flux of the measured radiation are normally used for such high-dose rate applications.
• two separate detector systems has previously been required; one detector system comprising, e.g., common photodiodes for enabling high-rate current readout needed for fast CT scanning, and one high gain detector system consisting of, for example, APDs for enabling high-resolution single-photon readout needed for PET (and SPECT) scanning.
• APDs for enabling high-resolution single-photon readout needed for PET (and SPECT) scanning.
• the requirement of two separate detector systems implies more complex, expensive and cumbersome radiation detection devices.
• WO 2006/034585 Al presents a method and system for acquiring CT images using a photon- counting detector, such as an avalanche photodiode (APD) or an array of APDs (e.g., a SiPM).
• a photon- counting detector such as an avalanche photodiode (APD) or an array of APDs (e.g., a SiPM).
• CT scans often require a high radiation dose and detectors operating in photon-counting mode are unable to measure high photon fluxes due to saturation effects.
• such a dual mode APD When operated in pulse mode and used together with a scintillator in medical X-ray imaging, such a dual mode APD can be used to detect single primary photons up to fluxes of millions of photons per mm 2 per second.
• the dynamic range When operated in current sensing mode the dynamic range can be extended by many orders of magnitude, in principle far beyond what is practically useful for most applications.
• the output signal amplitude carries no information of the primary photon energy incident on the scintillator crystal.
• Such problems could be remedied by using a SiPM device comprising a plurality of APDs connected in parallel.
• a SiPM is used and the number of secondary photons incident on the photo detector per unit area of the device, within the single-pixel recovery time (which for state-of-the-art devices lies in the range ⁇ 100ns - 1 ⁇ s), becomes comparable to the pixel density (currently limited to around 1000 - 2000 pixels per mm 2 , depending on the type and manufacturer), significant nonlinearities in the sensor response will occur.
• the finite single-pixel recovery time of Geiger mode APDs makes them count rate limited and introduces non-linearities in the current response at high photon fluxes. Not only does this preclude quantification of high radiation fluxes in single-photon counting mode due to saturation, the nonlinearities in the sensor response also severely compromise the readout of the continuous mean current in current mode operation.
• a detection system for photon sensing and for measuring photon fluxes comprising at least one photo detector element arranged to detect incident radiation and comprising at least one avalanche photo diode (APD).
• the detection system comprises measuring means for measuring the rate and magnitude of the discrete electric pulses, and the mean current, generated by the detector element as a result of the incident radiation.
• the detection system is characterized in that it comprises a bias regulator comprising means for altering a bias voltage applied to the detector element between at least a first and a second voltage level, said first voltage level being below the breakdown voltage of the detector element, and said second voltage level being above the breakdown voltage of the detector element.
• the detector element can be operated in Geiger mode (bias voltage above the break-down voltage) at low to moderate photon fluxes requiring a high gain for optimizing the readout in pulse mode, and in proportional mode (bias voltage below the breakdown voltage) at high photon fluxes for optimizing the readout in current mode, thus extending the dynamic intensity range of the detector element.
• Geiger mode bias voltage above the break-down voltage
• proportional mode bias voltage below the breakdown voltage
• the detector element preferably comprises a plurality of APDs, constituting what is sometimes called a silicon photomultiplier (SiPM) or a multi-pixel photon counter (MPPC).
• SiPM silicon photomultiplier
• MPPC multi-pixel photon counter
• the bias regulator comprises an operator input device, allowing an operator of the detection system to switch between the two different bias voltage levels.
• an operator can chose to optimize the readout in pulse mode for low to moderate rate applications, such as PET and SPECT, and to optimize the readout in current mode for high rate applications, such as standard CT.
• the bias regulator may also be arranged to automatically switch between the two different bias voltage levels applied to the detector element in dependence of a parameter value indicative of the radiation intensity incident on the detector element.
• the bias regulator is then arranged to apply a bias voltage above the breakdown voltage of the detector element if the parameter value is indicative of a low incident radiation intensity, and a bias voltage below the breakdown voltage of the detector element if the parameter value is indicative of a high incident radiation intensity.
• the parameter value received by the bias regulator and used in the decision of whether to apply a bias voltage above or below the breakdown voltage of the detector element may be a parameter value relating to the rate and magnitude of the discrete pulses and/or the mean current generated by the detector element as a result of the incident radiation. It may also be a parameter value relating to the radiation intensity I R emitted from a radiation source.
• the present invention eliminates the need for two separate detection systems in applications requiring such a large radiation intensity range.
• State-of-the-art technology e.g. for dual-modality CT/PET scanners, is based on the use of separate detection systems combined into one imaging system.
• the present invention improves both the cost efficiency and the performance of such a scanner by providing the means for a single detection system being able to operate in single photon counting Geiger mode for PET, SPECT or low-rate CT applications and in current mode for high-rate CT applications.
• the combined CT/PET/SPECT scanner according to the present invention can provide greatly enhanced medical diagnostics such as in oncology as well as cardiology and neurology.
• the detection system of the present invention utilizes photo detectors based on semiconductor technology, the detection system can be integrated in MRI systems, thus providing multiple-modality CT/PET/MRI or CT/PET/SPECT/MRI systems.
• One particularly interesting field of application of the present invention is for combined diagnostic and therapeutic imaging.
• This concerns imaging in connection with radiation-based cancer therapy where it is important to couple as closely as possible the diagnostic imaging of the patient with the portal imaging that is performed online to verify the dose delivery.
• the invention will enable the design of a single detector system that can perform both diagnostic CT/PET/SPECT as well as the portal imaging tasks during a cancer therapy session. This will greatly increase the accuracy, quality and efficiency of the dose delivery and lead to higher protection of healthy tissues by reducing uncertainties in patient positioning and by providing the possibility of online corrections to the therapeutic beam.
• the object is also achieved by a method for photon sensing and for measuring photon fluxes within a large dynamic radiation range.
• the method preferably comprises the use of a detection system as disclosed above.
• Fig. 1 shows an embodiment of the detector system according to the present invention, realized in a system intended for combined PET/CT or PET/SPECT/CT scanning.
• Fig. 2 shows the photo detector element used in the detector system shown in Fig. 1
• Fig. 3 shows a flowchart illustrating an example of a method for measuring photon fluxes within a large dynamic range according to the present invention.
• CT computed tomography
• PET positron emission tomography
• SPECT single photon emission computed tomography
• Fig. 1 shows part of a detection system for a combined PET/SPECT/CT scanner according to one embodiment of the present invention.
• a radiation source 4 emits ionizing radiation, illustrated as arrows in the figure, towards a target region 8 with an intensity regulated by a radiation control unit (not shown) contained within or connected to the radiation source 4.
• Scintillators are well known in the art and serve to fluoresce photons at characteristic wavelengths in response to absorbed incident ionizing radiation.
• the number of generated fluorescence photons is substantially proportional to the energy of the incident primary X-ray or gamma photon.
• the scintillators 1 may be any known type of scintillators, such as organic crystal scintillators, organic liquid scintillators, organic plastic scintillators or inorganic crystal scintillators.
• the scintillators 1 are optically coupled to a plurality of semiconductor photo detector elements 3.
• the scintillators 1 may be coupled to the photo detector elements 3 via light guides (not shown).
• Fig. 1 there is a one-to-one correspondence between detector elements 3 and scintillators 1 , i.e. each photo detector element 3 is coupled to one scintillator.
• each photo detector element 3 may be coupled to more than one scintillator 1 , or more than one photo detector element 3 may be coupled to each scintillator 1.
• the photo detector elements 3 may comprise a single avalanche photo diode (APD) 2 or, preferably, a plurality (an array) of APDs 2 constituting what is known as a silicon photo multiplier (SiPM) device, multi-pixel photon counter (MPPC), or multi-cell avalanche photo diode (MAPD).
• SiPM silicon photo multiplier
• MPPC multi-pixel photon counter
• MPD multi-cell avalanche photo diode
• the photo detector elements 3 are in turn connected to measuring means for measuring the rate and magnitude of the discrete electric pulses, and the mean current, generated by the at least one APD 2 as a result of the incident radiation. Consequently, each photo detector element 3 is arranged to operate both in single photon counting mode (pulse mode) and current mode.
• the means for measuring the rate and magnitude of the electric pulses and the means for measuring the mean current may comprise any components and circuitries known in the art and need not to be described in detail.
• the detector elements 3 are arranged to operate in pulse mode and current mode simultaneously, thus allowing simultaneous measurements of both the rate and magnitude of the discrete electric pulses, and the mean current, generated by the detector elements 3.
• each photo detector element 3 is coupled to a current processing circuitry 9 via a current-sensitive amplifier 5 for current-mode operation, and to a pulse processing circuitry 11 via an optional amplifier 7 for pulse-mode operation. It is also possible to add the output from several photo detector elements 3 and then pass on the summed signal to the amplifiers 5, 7 and the processing circuitries 9, 11, or to add the signals produced after amplification before passing them to the processing circuitries 9, 11. Appropriate current and pulse processing circuitries 9, 11 for CT, PET and SPECT applications are well known in the art and need not further be disclosed herein.
• the above described components constitute a detector module 6 and a plurality of such detector modules 6 constitutes the detector system in the combined CT/PET/SPECT scanner.
• the total number of detector modules 6 used in the combined CT/PET/SPECT scanner may, of course, vary with the size of the area that needs to be covered by detectors and the position resolution requirements for the particular application.
• the current processing circuitry 9 and the pulse processing circuitry 11 of each detector module 6 are coupled to a data processing unit 15 via two separate channels CH 1 , CH 2.
• the output from the current processing circuitry 9 and/or the output from the pulse processing circuitry 11 of several detector modules 6 may be added before being sent to the data processing unit 15. That is, the data processing unit 15 may have a varying number of data channels.
• the data processing unit 15 registers data corresponding to the energy and timing of the received signals and the mean currents induced in the photo detectors and uses this information in a known manner to create diagnostic images of the target region 8.
• the volume "seen" by the detector system in CT/PET/SPECT applications is often referred to as "the field of view".
• the field of view does not necessarily coincide with the irradiated target region in CT applications and images are, of course, only possible to obtain for the field of view.
• target region used in this description should be interpreted as the field of view in any of the above mentioned medical imaging applications.
• the photo detector elements 3 are also electrically connected to means 13 for providing a controlled bias voltage V B thereto.
• the means 13 for providing a controlled bias voltage V B to the detector assemblies 3 will henceforth be referred to as the bias regulator 13.
• the bias regulator 13 can in its simplest form be a voltage supply arranged to provide two different voltage levels to the photo detector elements 3: one level above the break-down voltage of the photodetectors 3, and one level below the break-down voltage of the photodetectors 3.
• the bias regulator 13 preferably comprises an operator input device for allowing an operator of the equipment to switch between the two levels. For example, an operator is then able to choose a voltage level above the breakdown voltage of the photo detector elements 3 for PET/SPECT or low to moderate rate CT applications, and a voltage level below the breakdown voltage of the photo detector elements 3 for high-rate CT applications.
• the bias regulator 13 comprises a voltage supply and a control unit for regulating the voltage V B supplied to the photo detector elements 3 in dependence of certain variables relating to the photon flux intensity Ii incident on the photo detector elements 3.
• the bias regulator 13 is arranged to receive such radiation related information from the data processing unit 15, via a communication channel 17, and/or from the radiation source 4, via another communication channel 19. These communication channels 17, 19 are illustrated as dotted lines in Fig. 1.
• the radiation source 4 or its control unit may be arranged to provide the bias regulator 13 with information related to the radiation emitted from the radiation source 4, e.g.
• the emitted radiation intensity I R and the data processing unit 15 may be arranged to provide the bias regulator 13 with information related to the pulses or the mean current outputted by the detector modules 6.
• the radiation source 4 may be connected to the data processing unit 15, in which case information related to the radiation source 4 or its control unit can be provided to the bias regulator 13 via communication channel 17.
• the detector system comprising a bias voltage V B regulatory system according to the present invention thus allows the photo detector elements 3 to be operated in Geiger mode by applying a bias voltage V B above the breakdown voltage and in normal mode by lowering the bias voltage V B below the breakdown voltage.
• the detector system according to the invention can be used in applications requiring a large dynamic radiation range and be optimized for the radiation intensity currently used.
• the internal gain of the detector elements 3 is sufficient to detect single photons, thus optimizing the detector system for low to moderate radiation rate applications, such as PET, and by operating the detector elements 3 in proportional mode, the mean current outputted from the detector elements 3 will stay proportional to the incident photon flux even at high fluxes, thus optimizing the detector system for high radiation rate applications, such as CT.
• the bias voltage regulatory system described above can be arranged to stabilize the gain in either mode when the radiation intensity is changing. If, for example, a protective resistor is connected in series with the bias regulator 13, a varying current will cause changes of the voltage drop across the detector elements 3 and hence in the gain. This can be compensated for by the bias regulatory system.
• the bias regulatory system may also be arranged to control the bias voltage V B applied to the detector elements 3 in dependence of parameters not mentioned above. For example, temperature variations of the detector elements 3 may be measured and the bias regulatory system may be arranged to vary the applied bias voltage in dependence of the measured temperature values.
• a photo detector element 3 is schematically illustrated. As described above, the detector elements 3 may comprise a single APD 2 or a plurality of APDs 2, constituting a SiPM 10. According to the preferred embodiment of the present invention the detector elements 3 are SiPMs 10, as shown in Fig. 2.
• the SiPM 10 comprises a matrix of APDs 2 connected in parallel.
• the pixel density and the active surface of each SiPM device may vary but is approximately 100-2000 or more pixels per mm 2 , and 0.1-25 mm 2 , respectively.
• V B When operated in Geiger mode, i.e. when the applied bias voltage V B is above the breakdown voltage of the APDs 2, one or a few secondary photons trigger an avalanche breakdown of the APD 2, resulting in an electric pulse whose magnitude is independent of the energy of the incident photons.
• the pulses from each APD 2 are added and due to the large number of pixels per unit area, the magnitude of the total pulse outputted from the SiPM device can, up to a certain number of incident secondary photons within the single-pixel recovery time, be substantially proportional to the number of secondary photons striking that particular SiPM device 10 and hence proportional to the incident primary photon energy striking the scintillators 1.
• the output from the SiPM device 10 can emulate the linear response of normal photomultiplier tubes and APDs operated in proportional mode.
• SiPM device 10 Another advantage with the SiPM device 10 is that a signal gain similar to that of photomultiplier tubes (10 6 -10 7 ) can be achieved with a bias of only around 20-80 volts.
• the high gain provided by the SiPM 10 reduces the cost of the electronics chain in the detector system. A high cost of the electronics usually limits the number of electronics channels in an imaging system, which results in reduced throughput and lower sensitivity.
• the SiPM device 10 is also rugged, compact (millimeters in size), and inexpensive.
• the bias voltage V B applied to the detector elements 3 may be chosen by an operator of the detection system and be supplied to the detector elements 3 by an ordinary voltage supply 13. But the bias voltage V B may also be automatically controlled by the bias regulator 13 in dependence of certain parameter values relating to the radiation intensity I 1 incident on the detector elements 3.
• a method for measuring photon fluxes within a large dynamic range, utilizing such automatic control of the bias voltage V B is described.
• Fig. 3 a flowchart illustrating a method for measuring photon fluxes within a large dynamic range according to the present invention is shown. The method can be used in, e.g., CT, PET, SPECT, or radiotherapy portal imaging applications, or any combination of these.
• the method is not limited to any particular application, but can be implemented in any detection system utilizing APDs or SiPM devices as detector elements.
• At least one detector element 3 is arranged to detect the radiation passing through a target region 8 emitted from an external radiation source, or radiation arising from positron annihilation or other physical processes within the target region 8.
• the detector element 3 may be arranged to detect this primary radiation directly or being optically coupled to for example scintillators for detecting secondary radiation that is generated by the scintillators and proportional to the primary radiation.
• the detector element 3 is an APD 2 or a matrix of APDs constituting a SiPM device 10.
• step 2 a controlled bias voltage V B is applied to the detector element 3.
• step 3 the magnitude and rate of the electric pulses and the mean current generated by the detector element 3 as response to the incident photon intensity I 1 are measured.
• the output from the detector element 3 is registered by a data processing unit 15, preferably after being processed by preamplifiers 5, 7 and signal processing circuitries 9, 11 adapted to optimize the mean current and pulse readout.
• step 4 at least one parameter value indicative of the photon flux intensity Ii that is to be measured by the detector element 3 is provided to the bias regulator 13 for regulating the bias voltage V B applied to the detector element 3.
• the parameter value may be provided by the radiation source 4, or by a control unit controlling the radiation source 4, and relate to, e.g., the radiation intensity IR emitted by the radiation source 4. It can also be provided by the data processing unit 15 and relate to the registered magnitude and/or rate of the electric pulses or the mean current outputted from the detector element 3.
• the bias regulator 13 comprises means for comparing the received parameter value to a predetermined threshold value and increase or decrease the bias voltage V B applied to the detector element 3 in response to the result of the comparison.
• the comparison is taking place in step 4 and depending on the result of the comparison, the bias voltage V B of the detector element 3 is increased, decreased ('Yes' in step 4) or unchanged ('No' in step 4).
• the bias regulator 13 may be arranged to receive a value from the data processing unit 15 relating to the mean current outputted from the detector element 3.
• the bias regulator 13 applies a bias voltage V B to the detector element 3 that exceeds its breakdown voltage, thus putting the detector element 3 into Geiger mode to achieve a high internal gain. If the mean current is above the predetermined threshold value on the other hand, the bias regulator 13 applies a bias voltage V B below the breakdown voltage of the detector element 3, thus putting it into proportional mode in order to avoid non- linearities and saturation in the current readout. If the detector element 3 already operates in the operational mode that is optimal for the photon flux intensity indicated by the parameter value, no regulation of the applied bias voltage is necessary. However, optionally a SiPM gain stabilization can be achieved by fine tuning the applied bias voltage as a function of the mean current and possibly also as a function of the temperature of the detector unit.
• one single bias regulator 13 is arranged to control the bias voltage V B of all the photo detector elements 3 in the detector system.
• a plurality of bias regulators 13 is utilized in order to apply different bias voltages V B to different detector elements 3.
• the bias voltage V B applied to individual detector elements 3 can be altered in dependence of local radiation intensity variations.
• a detector element 3 located in an area with a high intensity I) of incident radiation may be operated with a bias voltage V B below the breakdown voltage, thus optimizing the readout in current mode, while a detector element 3 located in an area with a low to moderate intensity I 1 of incident radiation may be operated with a bias voltage V B above the breakdown voltage, thereby optimizing the readout in pulse mode.
• the data processing unit 15 utilizes the optimized outputs from all the detector elements 3, representing the energy and/or timing of the detected photons and/or the mean currents induced in the photo detectors, to create diagnostic images (CT/PET/SPECT imaging) and/or therapeutic images (portal imaging) of the target region 8 in a known way.
• CT/PET/SPECT imaging diagnostic images
• therapeutic images therapeutic images
• the detection system and method according to the present invention may be used for any applications concerned with detecting electromagnetic radiation using APDs or SiPM devices.
• the detection system has been described in the context of a combined CT/PET/SPECT scanner for diagnostic and therapeutic imaging, it may as well be employed in, e.g., industrial radiography applications, environmental radiation monitoring, nuclear safeguard applications, airport security applications, and optical imaging applications.
• the radiation may as well originate from a non-controllable radiation source, such as radioactive contamination, as from a controlled or known radiation source. That is, there may not be any radiation source to "control” or “direct” or even a "target region” to examine as in the case with the combined CT/PET/SPECT application described herein.
• a non-controllable radiation source such as radioactive contamination
• the present invention is not limited to the exemplary configuration disclosed and illustrated herein. Different detector configurations and electronic circuitries may also be used to implement the invention.
• the detailed disclosure of the invention only is illustrative and exemplary and merely serves the purpose of providing a full and enabling disclosure thereof. Accordingly, it is intended that the invention should be limited only by the scope of the claims appended hereinafter.

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