Abstract:
A phoswich device for determining depth of interaction (DOI) includes a first scintillator having a first scintillation decay time characteristic, a second scintillator having a second scintillation decay time characteristic substantially equal to the first scintillation decay time, a photodetector coupled to the second scintillator, and a wavelength shifting layer coupled between the first scintillator and the second scintillator, wherein the wavelength shifting layer modifies the first scintillation decay time characteristic of the first scintillator to enable the photodetector to differentiate between the first decay time characteristic and the second decay time characteristic. The phoswich device is particularly applicable to positron emission tomography (PET) applications.

Description:
CROSS-REFERENCE TO RELATED APPLICATION AND CLAIM OF PRIORITY 
       [0001]    This application is a non-provisional of and claims priority to under 35 U.S.C. §119(e) copending Application Ser. No. 61/100,916 filed Sep. 29, 2008. 
         [0002]    Cross reference to related material can also be found in an application having U.S. application Ser. No.: 12/110,544, titled “Implementation of Wavelength Shifters in PHOSWICH Detectors”, which was filed on Apr. 28, 2008, the entire contents of which are incorporated herein by reference. 
     
    
     TECHNICAL FIELD 
       [0003]    The present invention is in the field of nuclear medical imaging. In particular, the present invention relates to techniques for accurate detection of emission radiation in nuclear medical imaging processes such as positron emission tomography (PET). 
       BACKGROUND 
       [0004]    Medical imaging is one of the most useful diagnostic tools available in modern medicine. Medical imaging allows medical personnel to non-intrusively look into a living body in order to detect and assess many types of injuries, diseases, conditions, etc. Medical imaging allows doctors and technicians to more easily and correctly make a diagnosis, decide on a treatment, prescribe medication, perform surgery or other treatments, etc. 
         [0005]    There are medical imaging processes of many types and for many different purposes, situations, or uses. They commonly share the ability to create an image of a region of the body of a patient, and can do so non-invasively. Examples of some common medical imaging types are nuclear medical (NM) imaging such as positron emission tomography (PET) and single photon emission computed tomography (SPECT), electron-beam X-ray computed tomography (CT), magnetic resonance imaging (MRI), and ultrasound (US). Using these or other imaging types and associated machines, an image or series of images may be captured. Other devices may then be used to process the image in some fashion. Finally, a doctor or technician may read the image in order to provide a diagnosis. 
         [0006]    In traditional PET imaging, a patient is injected with a radioactive substance with a short decay time. As the substance undergoes positron emission decay, it emits positrons which, when they collide with electrons in the patient&#39;s tissue emit two simultaneous gamma rays. The gamma rays emerge from the patient&#39;s body at substantially opposite directions. These rays eventually reach a scintillation device positioned around the patient. There is often a ring of scintillation devices surrounding the patient. When the gamma rays interact with oppositely positioned scintillation devices, light is emitted and detected. The light is usually transmitted through a lightguide to a photodetector. The light detected by the photodetector is then interpreted by a processor to enable an image of a slice of the region of interest to be reconstructed. 
         [0007]    In PET (as well as SPECT) it is important to match the scintillator emission wavelength to the photodetector&#39;s optimal wavelength quantum efficiency (QE). For example, a typical photomultiplier tube (PMT) used in PET applications has a peak wavelength sensitivity at 420 nm while a typical LSO scintillator used in PET emits at 420 nm. Therefore, PMTs and LSO are very well matched in terms of wavelength matching. LSO is a very good scintillator for a PMT and is reasonably matched also for other silicon-based photodetectors such as avalanche photodiodes (APDs) and silicon photomultipliers (SiPMs). Scintillators for PET may be made from crystal materials such as, but not limited to, LSO, YSO, LYSO, LuAP (i.e., LuAlO 3 :Ce), LuYAP, or LaBr3. 
         [0008]    The phoswich approach has been used to improve the detection in PET applications by determining the depth-of-interaction (DOI) in the detector. PET scanners are typically made of long, thin detectors with high stopping power to meet high sensitivity requirements. In the absence of DOI information, however, the thickness of the scintillator reduces the spatial resolution due to parallax error. To compensate for reduced spatial resolution, detectors with DOI capability have been used. DOI capability can determine the location of the gamma interaction in the direction of the incident gamma (i.e., depth from the surface of the detector). 
         [0009]    One way to implement DOI capability is to use a multi-layer detector, in which the layers are made of material with different scintillation properties. Because the layers have different characteristics, when a gamma event is detected it is possible to identify which layer absorbed the gamma photon and so to determine more accurately the spatial interaction location in three dimensions. 
         [0010]    A conventional “phoswich” thus is a detector with two or more layers of different scintillators. Phoswich detectors comprising two or more scintillator layers offer a means to simultaneously achieve both high sensitivity and high spatial resolution in nuclear imaging. Each scintillator layer typically has a distinct decay time that allows the DOI of a gamma ray to be determined via pulse shape determination techniques. That is, layer identification is done by using differences in scintillation decay time inherent in the scintillators and pulse shape discrimination techniques. 
         [0011]    The use of different types of scintillators in a phoswich may result in different light yields, emission spectra, densities, effective atomic numbers, and indices of refraction, which can often result in compromises in performance of the phoswich. 
       SUMMARY 
       [0012]    The present invention solves the existing need in the art to determine the depth of interaction for a PET detector using the same or similar scintillation materials in a phoswich detector. An embodiment of the present invention uses a phoswich device for determining depth of interaction (DOI). The phoswich device includes a first scintillator having a first scintillation decay time characteristic and a second scintillator having a second scintillation decay time characteristic substantially equal to the first scintillation decay time. The phoswich device further includes a photodetector coupled to the second scintillator, and a wavelength shifting layer coupled between the first scintillator and the second scintillator, wherein the wavelength shifting layer modifies the first scintillation decay time characteristic of the first scintillator to enable the photodetector to differentiate between the first decay time characteristic (front scintillator) and the second decay time characteristic (back scintillator). 
         [0013]    The phoswich device is particularly applicable to positron emission tomography (PET) applications. 
     
    
     
       BRIEF DESCRIPTION OF THE DRAWINGS: 
         [0014]    The invention will now be described in greater detail in the following by way of example only and with reference to the attached drawings, in which: 
           [0015]      FIG. 1  is a depiction of a conventional phoswich detector configuration. 
           [0016]      FIG. 2  is a depiction of a phoswich detector where a wavelength shifting layer is sandwiched between two scintillators, in accordance with an embodiment of the present invention. 
           [0017]      FIG. 3  a depiction of an exemplary phoswich detector where a wavelength shifting layer is sandwiched between two LSO scintillators, in accordance with an embodiment of the present invention. 
           [0018]      FIG. 4  is a depiction of a PSD configuration, in accordance with a second embodiment of the present invention. 
           [0019]      FIGS. 5A and 5B  are graphs showing the pulse shape discrimination peaks ( 5 A) obtained with the set-up in  FIG. 4  and the energy spectra of the two scintillators ( 5 B). 
           [0020]      FIG. 6  is a schematic of a PET scanner using a phoswich device in accordance with an embodiment of the present invention. 
       
    
    
     DETAILED DESCRIPTION 
       [0021]    As required, disclosures herein provide detailed embodiments of the present invention; however, the disclosed embodiments are merely exemplary of the invention that may be embodied in various and alternative forms. Therefore, there is no intent that specific structural and functional details should be limiting, but rather the intention is that they provide a basis for the claims and as a representative basis for teaching one skilled in the art to variously employ the present invention. 
         [0022]      FIG. 1  depicts a conventional phoswich combination  100  that has been considered for Positron Emission Tomography (PET) imaging. The phoswich combination  100  has a first type scintillator crystal  110  made of LuAP. When a 511 keV gamma photon is absorbed in the LuAP crystal  110 , it emits light scintillations around 360 nm, with a scintillation time of 30 ns. The light photons travel out of the LuAP scintillator array  110 , and is absorbed by LSO scintillator array  120 , and re-emitted at 420 nm with a scintillation time of 40 ns. The 420 nm enters into photodetector  130 , where the light photons are converted into an electrical signal. An exemplary photodetector  130 , such as a photomultiplier tube, detects the 420 nm light with decay time characteristics given as convolution between the decay time of the LuAP crystal and the decay time of the YSO crystal. The following equation shows how the decay time characteristics detected by the PMT  130  are obtained. 
         [0000]        T (combined)= T ( LuAP )           T ( LSO )  Equation 1 
         [0023]    The presence of T(combined) signifies an event in the front scintillator 
         [0024]    As previously discussed, there may be problems with using different types of scintillator crystals in the same phoswich. 
         [0025]      FIG. 2  is a depiction of a phoswich detector utilizing a wavelength shifting layer sandwiched between two scintillators, in accordance with an embodiment of the present invention. The phoswich device  200  includes a scintillator  202 , a wavelength shifting layer  206 , a filter  208 , a scintillator  204  and a photodetecter  210 . The photodetector can comprise a photomultiplier tube, a solid state photodetector such as an APD, MRS-PD and SiPM. 
         [0026]    Wavelength shifting layer  206  may include, but is not limited to a plastic light guide, a crystal, and liquid coatings made from wavelength shifting material. An exemplary wavelength shifting material used was a green wavelength shifter. Specifically, an Eljen 280 foil (WLS) that has an excitation band around 430 nm, suitable to interact with light emission from LSO, and an emission band around 500 nm. The WLS foil has a thickness around 0.1 mm. 
         [0027]    It should be appreciated by those skilled in the art that although a green wavelength shifter is used, the invention may be modified to use other wave length shifting layers as long as interactions occur between the front scintillator and the shifting layer and a decay time modification takes place. Other scintillator types may fall within the scope of the present invention. 
         [0028]    In a first embodiment of the invention, scintillator  202  and scintillator  204  are from the same crystal group which results in scintillator  202  and scintillator  204  having substantially equal decay time characteristics. For example, scintillator  202  comprises a LSO crystal and scintillator  204  comprises a LSO crystal. Scintillator  202  and scintillator  204  may comprise LSO, YSO, LUAP, LUYap, LFS, LYSO, LaBr3, and the like crystals. 
         [0029]    In a second embodiment of the invention, scintillator  202  and scintillator  204  are from different crystal groups but both crystals for light transmission purposes are substantially alike, e.g. have substantially the same decay time characteristics and do not interact. 
         [0030]    In an embodiment of the invention, a 511 keV gamma photon is absorbed by scintillator  202 . The gamma photon emits light scintillations around 420 nm, with a scintillation time of 40 ns. The light photons exit scintillator  202  and are absorbed by wavelength shifting material  206 . Wavelength shifting material  206  can affect the light photons in a number of ways depending on the type of wavelength absorbing material used. For instance, in one embodiment of the invention, the wavelength shifting material  206  increases the decay time characteristic of the light photons entering scintillator  204 . In another embodiment of the invention, the decay time remains substantially the same, however, the rise time of the light photons entering scintillator  204  is affected. 
         [0031]    The light photons exit the wavelength shifting material  206  and enter scintillator  204 . The light photons then exit scintillator  204  with a decay time characteristic of 40 ns but a modified rise time and enter the photodetector  210  where the light photon is converted to an electrical signal. 
         [0032]    The filter  208 , which is preferably a long pass filter prevents light from scintillator  204  from being reflected into scintillator  202 . 
         [0033]    Conversely, a 511 keV photon travels through scintillator  202 , and is absorbed by the scintillator  204 . Scintillator  204  now emits 420 nm light with a decay time of 40 ns. Thus, identification of the location of the gamma interaction in either the front scintillator or the back scintillator can be easily made by analyzing the signals from photodetector  210 . 
         [0034]      FIG. 3  a depiction of an exemplary phoswich detector  300  where a wavelength shifting layer  306  is sandwiched between two LSO scintillators  302 ,  304  in accordance with an embodiment of the present invention. In an embodiment of the invention, a 511 keV gamma photon is absorbed by LSO scintillator  302 . The gamma photon emits light scintillations of 420 nm, with a scintillation time of 40 ns. The light photons exit LSO scintillator  302  and are absorbed by wavelength shifting material  306 . The LSO decay time characteristics is changed due to the interaction and the light emission is shifted upwards to around 500 nm. The shifted light has, thus, the decay time characteristics of a convolution between LSO light response and the response of the wavelength shifter. Specifically, the rise time of the light signal is affected by wavelength shifting layer  306 . Specifically, the light signal has a 12 ns rise time. The decay time remains substantially the same. 
         [0035]    The light photons generated in  302  exit the wavelength shifting material  306  and enter the photodetector  310  via the LSO scintillator  304 . 
         [0036]    Conversely, a 511 keV photon travels through LSO scintillator  302 , and is absorbed in LSO scintillator  304  which emits 420 nm light with a decay time of 40 ns which is detected by photodetector  310 . Filter  308 , prevents light from LSO scintillator  304  from being transmitted in the wave length shifter and into LSO scintillator  302 . One light signal arrives with a light scintillation of 500 nm and a rise time of 12 ns and a 40 ns decay time signifying an event in the front scintillators ( 302 ), another signal arrives with a light scintillation of 420 nm and a rise time of 1 ns and a 40 ns decay time, signifying a 511 keV event has been registered in the back scintillator ( 304 ). With the rise time sensitive pulse shape discriminator circuit, the two light signals can be differentiated. 
         [0037]      FIG. 4  is a depiction of a Pulse Shape Discrimination (PSD) circuit  400  in accordance with a second embodiment of the present invention. Signal  402  also known as signal A comprises an exemplary 500 nm light scintillation with a 12 ns rise time. Signal  404  also known as signal B comprises a 420 nm light scintillation with a 1 ns rise time. Signal A and signal B are detected by a photodetector  406  depicted as a PMT. Two constant fraction discriminator (CFD) circuits are provided. The anode signal goes to discriminator circuit  410 , timing circuit  412  and discriminator circuit  414  and time activity curve circuit (TAC)  416 . The result is illustrated in graph  420  which depicts the time channel of signal A and Signal B. The dynode signal if available is connected to a spectroscopy amplifier to provide the phoswich energy spectra. Graph  418  illustrates the energy channel of signal A and signal B 
         [0038]      FIGS. 5A and 5B  are graphs illustrating a time spectrum and an energy spectrum for the phoswich of  FIG. 3  in accordance with an embodiment of the present invention. For  FIG. 5A , which depicts the PSD time spectrum all events below channel  88  are defined as fast events (e.g. scintillator  304  being closest to the photodetector  310 ). All events greater than channel  88  are defined as slow events (longer rise time, scintillator  302 ). Based on these two gates, the two energy spectra depicted in  FIG. 5B  are acquired. The source  68 Ge, providing the 511 keV photons, was positioned just above scintillator  302 . Based on the two Gaussians from  FIG. 5A , the cross-talk between scintillator  302  and scintillator  304  can be calculated. For the fast setting (e.g., channels below 88) there are 93% true fast events and 7% slow events coming in via scintillator  304 . For the slow setting, the numbers are 98% true slow events and 2% fast events. 
         [0039]      FIG. 6  is a diagram of a PET scanning system  600  using a wavelength shifting material in the phoswich device in accordance with another embodiment of the invention. PET scanning system  600  consists of a number of phoswich detectors  620 . The phoswich detectors may be arranged in a ring configuration. The ring of phoswich detectors  620  forms a space large enough for an adult human body to pass. Each phoswich detector may consist of a first scintillator material, a wavelength shifting material, a second scintillator material and a photodetector. The ring of phoswich detectors  620  may be connected to a processor  630 . The processor  630  is capable of analyzing the data received from the ring of phoswich detectors  620 , reconstructing an image from the acquired data, and outputting tomographic images of the object or patient scanned. The PET scanning system  600  may further include a table or other support structure  610  capable of holding the object or patient to be scanned. The table or other support structure  610  may be adapted to pass through the bore formed by the ring of block detectors  620 . 
         [0040]    The invention having been thus described, it will be apparent to those skilled in the art that the same may be varied in many ways without departing from the spirit and scope of the invention. Any and all such modifications are intended to be covered within the scope of the following claims.