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The Silicon Photomultiplier (SiPM) is a compact detector which shows great promise for use in MR compatible PET systems. SiPMs are insensitive to magnetic fields, have a high intrinsic gain (about 106) and operate at low voltages (40 V), eliminating the need for sophisticated shielding and low noise preamplifiers. A prototype 1×1 mm2 detector currently under investigation at the University of Toronto consists of an array of 556 independent microcells operating in Geiger mode. Each microcell thus produces a pulse of constant amplitude with the detection of one or more electrons. All microcells are connected in parallel through integrated quenching resistors, and thus the output of the SiPM is the sum of the individual standardized signals and is proportional to energy provided the number of detected light photons is less than the number of microcells. The detector was operated without a preamplifier, and using a 2×2×6 mm2 LSO crystal and 22Na source, the energy resolution was measured to be 25.6±0.4% FWHM and timing resolution for a detector pair was 1.9±0.1 ns FWHM. LED measurements and measurements using 133Ba (356 keV) and 137Cs (662 keV) verify that the detector is operating linearly at 511 keV. The major noise contribution is the dark current, which at −44 V results in a peak dark count rate of about 16 kHz that reduces to negligible levels above a 90 keV lower energy threshold. In the next generation of compact multi-element SiPM detectors with improved optical coupling the energy resolution should improve to below 15% FWHM, making them well suited for use in a high performance MR compatible depth of interaction PET scanner.

Pinhole collimation of gamma rays to image distributions of radiolabeled tracers is considered promising for use in small animal imaging. The recent availability of transgenic mice, coupled with the development of /sup 125/I and /sup 99m/Tc labeled tracers, has allowed the study of a range of human disease models while creating demand for ultrahigh resolution imaging devices. We have developed a compact gamma camera that, in combination with pinhole collimation, allows for accessible, ultrahigh resolution in vivo single photon emission computed tomography (SPECT) imaging of small animals. The system is based on a pixilated array of NaI(Tl) crystals coupled to an array of position sensitive photomultiplier tubes. Interchangeable tungsten pinholes with diameters ranging from 0.5 to 3 mm are available, allowing the camera to be optimized for a variety of imaging situations. We use a three dimensional maximum likelihood expectation maximization algorithm to reconstruct the images. Our evaluation indicates that high quality, submillimeter spatial resolution images can be achieved in living mice. Reconstructed axial spatial resolution was measured to be 0.53, 0.74, and 0.96 mm full width at half maximum (FWHM) for rotation radii of 1, 2, and 3 cm, respectively, using the 0.5-mm pinhole. In this configuration, sensitivity is comparable to that of a high-resolution parallel hole collimator. SPECT images of hot- and cold-rod phantoms and a highly structured monkey brain phantom illustrate that high quality images can be obtained with the system. Images of living mice demonstrate the ability of the system to obtain high-resolution images in vivo. The effect of object size on the quantitative assessment of isotope distributions in an image was also studied.

As part of the development of dedicated scintillation cameras for breast imaging, we compared the performances of two dedicated cameras with a standard clinical camera (Siemens Orbiter). One dedicated camera was based on a single position-sensitive photomultiplier tube (PSPMT) coupled to a 6 cm, by 6 cm by 6 mm NaI(Tl) crystal. The other (LumaGem) was based on multiple-PSPMTs coupled to a 2 mm /spl times/ 2 mm by 6 mm matrix of NaI(Tl) crystals. Spatial resolution was measured with all cameras as a function of depth. The ability of the cameras to image small superficial tumors was tested with a phantom consisting of six hot cylindrical tumors of height 3 mm and varying diameters against a warm background. The tumors were imaged at various depths within the background using tumor-to-background activity concentration ratios of 10:1 and 5:1. Tumor contrast and signal-to-noise ratio were measured in all cases for all cameras. The dedicated cameras show improved shallow depth of field spatial resolution, contrast, and signal-to-noise ratio relative to the clinical camera when imaging the breast tumor phantom., This, with the additional advantage of being able to position the cameras close to the region of interest, suggests that dedicated cameras will play a strong role in scintimammography and other shallow depth of field applications.

Position-sensitive gamma-radiation detectors equipped with collimators have been used for in vivo im- aging of the distribution of radiolabelled molecules in laboratory animals and humans for several decades. To date, the best image resolution achieved in a rodent is on the order of 1 mm. Here we demonstrate how a basic and compact gamma camera can be constructed for in vivo radionuclide imaging in small animals, at much higher spatial resolution. Resolution improvements were ob- tained by combining dense, shaped, micro-pinhole aper- tures with iodine-125, an isotope with low energy emis- sions, ease of incorporation into a wide range of mole- cules, and straightforward translation into the clinic via other isotopes of iodine that are suitable for nuclear med- icine imaging. 125 I images of test distributions and a mouse thyroid have been obtained at a resolution of as high as 200 µm using this simple bench-top camera. Pos- sible future applications and extension to ultra-high-res- olution emission tomography are discussed.

We are investigating the use of a retro-reflective tape (3M Industries), which reflects light along a line parallel to the incident light, as a substitute for Teflon tape or other reflective material used with scintillation crystals. We expect this retro-reflective property to improve the spatial resolution in large continuous crystals relative to that seen with Teflon, or other reflective coatings, when the reflective tape is coupled to the front (radiation incident) side of the crystal. Three CsI(TI) crystals of dimensions 5.8 cm /spl times/ 5.6 cm /spl times/ 5 mm, 6 cm /spl times/ 6 cm /spl times/ 12 mm, and 6 cm /spl times/ 6 cm /spl times/ 20 mm were evaluated by stepping a collimated /sup 99m/Tc source across the crystal face. Images were taken with the crystal coupled to a position-sensitive photomultiplier tube. The energy resolution and signal amplitude were similar for both the reflector and Teflon. The average spatial resolution in the central region of the 5-mm-thick crystal with black sides was measured to be 3.6 /spl plusmn/ 0.3 mm and 5.0 /spl plusmn/ 0.3 mm for retro-reflective tape and Teflon tape, respectively.

To evaluate the effect of ultrasonographic (US) contrast agents on measurements of peak velocity with spectral Doppler US in stenotic and nonstenotic flow states.Nonpulsatile flow was established in a flow phantom with 0%, 50%, 75%, and 90% stenoses. SH U 508A, perflenapent emulsion, and perfluorohexane emulsion were the contrast agents evaluated. Before and after administration of each contrast agent, two peak velocity measurements obtained proximal to, at the site of, and distal to the stenosis in each vessel model were averaged. The percentage difference in peak velocity after contrast agent administration was calculated for each site interrogated. The mean, SD, and coefficient of variation of the percentage difference in peak velocity were calculated.Percentage differences in peak velocity after contrast agent administration at different sample volume sites were not significantly different irrespective of the degree of stenosis or the contrast agent evaluated.The contrast agents evaluated do not produce a statistically significant increase in peak velocity. If this result is corroborated in clinical practice, contrast agents can be used without reevaluating existing Doppler US thresholds for stenosis.

Pinhole collimation of gamma rays to image distributions of radiolabeled tracers is considered promising for use in small animal imaging. The recent availability of transgenic mice, coupled with the development of /sup 125/I and /sup 99m/Tc labeled tracers, has allowed the study of a range of human disease models while creating demand for ultra-high resolution imaging devices. We have developed a compact gamma camera that, in combination with pinhole collimation, allows for accessible, ultra-high resolution in vivo SPECT imaging of small animals. The system is based on a pixilated array of NaI(Tl) crystals coupled to an array of position sensitive photomultiplier tubes (PSPMTs). Interchangeable tungsten pinholes with diameters ranging from 0.5 mm to 3 mm are available, allowing the camera to be optimized for a variety of imaging situations. We use a 3D MLEM algorithm to reconstruct projection images, allowing for 3D positioning of isotope distributions. Our evaluation indicates that high quality, sub-millimeter spatial resolution images can be achieved in living mice. Reconstructed axial spatial resolution was measured to be 0.53 mm and 0.74 mm FWHM for rotation radii of 1 cm and 2 cm, respectively, using the 0.5 mm pinhole. In this configuration, sensitivity is comparable to that of a standard high resolution parallel hole collimator. SPECT images of hot and cold rod phantoms and a highly structured monkey brain phantom illustrate that high quality images can be obtained with the system. Images of living mice demonstrate the ability of the system to obtain high resolution images in vivo.

A hand-held scintillation camera has been developed for localizing Sentinel Lymph Nodes (SLN) during or prior to surgery. Use of gamma proves has increased in recent years to identify SLNs as these are increasingly implicated in the spread of cancerous cells. The functionality of this new imaging probe will go beyond that of current non-imaging devices by displaying activity distributions in addition to counting rates. Using different collimators changes the spatial resolution, sensitivity and field of view (FOV) of the camera. In addition, the camera produces audible tones proportional to the count rate. The tone is modulated in real time to indicate the presence of complex distributions of activity within the FOV. The modulated tone alerts the user to localized concentrations of activity that may be nodes. The 22 mm FOV camera has 1 mm intrinsic spatial resolution. System FWHM in air is 2.1 mm and 2.6 mm at 0 cm and 1 cm from a high-resolution parallel hole collimator, respectively. FWHM of 3.8 mm is measured 2 cm from a 3 mm pinhole. Pinhole sensitivity is 600 cps/MBq above a 125 cps/MBq background for a 1 cm lesion 1 cm below a water surface. Nodes are identified in images even when overall count rate is not above background.

As part of the development of dedicated scintillation cameras, we compared the performances of 2 dedicated cameras with a standard clinical camera. The dedicated cameras were based on Position Sensitive Photo multipliers (PSPMTs): first, a single PSPMT coupled to a 6 by 6 by 0.6 cm NaI(Tl) crystal and second, multiple-PSPMTs coupled to a matrix of 2 by 2 by 6 mm NaI(Tl) crystal with hexagonal hole collimator. Image resolution was measured with all cameras as a function of depth. The ability of the cameras to measure small superficial tumors was tested with a phantom consisting of 6 hot cylindrical tumors of height 3 mm and varying diameters against a warm background of 4,1 cm cylinders. Tumors were stepped through the background and imaged at each level, starting at the collimator face, using a tumor to background activity concentration ratio of 10:1 and adjusting the imaging time for each to compensate for decay. Images were made of an anthropomorphic human thorax phantom with simulated breast lesions and detectability between the multi-PSPMT camera and the clinical camera was compared. The improve performance of dedicated cameras in specific tasks suggests that these devices will have a role in scintimammography and assisting in O.R. procedures such as sentinel node dissection and other shallow depth of field applications.

A high resolution, hand-held scintillation camera has been designed and built for specific Nuclear Medicine applications. Primary intended applications are pre-surgical and intra-operative lymphoscintigraphy. The detector head is highly compact with a 1-inch by 1-inch physical field of view. A variety of easily interchangeable collimators including parallel hole, diverging hole, and pinhole allow several choices of image parameters including variable spatial resolution, sensitivity and field of view. The camera can be operated in imaging mode or as a probe in a non-imaging mode. Surgeons performing sentinel node surgeries have the option of using the device asa standard audio-guided counting probe or as an imaging device to improve surgical management. The 20 mm FOV camera has 1 mm intrinsic spatial resolution. System FWHM in air is 2.1 mm and 2.6 mm at 0 cm from a high-resolution parallel hole collimator, respectively. FWHM of 3.8 mm is measured 2 cm from a 3 mm pinhole. Pinhole sensitivity is 600 cps/MBq above a 125 cps/MBq background for a 1 cm lesion 1 cm below a water surface. Nodes are identified in images even when overall count rate is not above the background from a nearby injection site.

Screen film mammography is currently accepted as being the best method available for the early detection of breast cancer. Digital mammography promises to replace screen film mammography as the primary screening tool of the future. The primary advantage it offers is the ability to perform image processing for optimisation of image contrast, which may be especially useful for lesion detection in patients with dense breasts [1] and may, despite its lower spatial resolution, make it more effective than conventional mammography at detecting lesions. Screen film systems are limited because there is no possibility for image contrast enhancement and hence when imaging thick or dense breasts, which require a higher energy spectrum for penetration and dose reduction, there is a loss of subject contrast resulting in a loss of calcification detection [2].