Lawrence R. MacDonald

Parallax correction in PET using depth of interaction information

The authors have investigated the restoration of radial resolution for transaxially off-center sources in a PET ring detector configuration using directly measured depth of interaction (DOI) information. Lutetium oxyorthosilicate (LSO) crystals are coupled to a photodetector along the radial dimension of the PET ring configuration. Four millimeter wide crystals are segmented into 4 mm and 10 mm lengths in the radial dimension and optically isolated so that scintillation light is localized to the radial extent of the segment, giving 4 mm and 10 mm DOI resolution. The authors also segmented 2 mm wide crystals into 10 mm lengths. This DOI technique is unambiguous but not practical for a multi-ring PET scanner. The objective here is to quantify the intrinsic radial resolution as a function of position within the FOV and DOI resolution. The authors found the degradation in radial resolution across 75% of the field of view was reduced by 45% (50%) with DOI resolution equal to one-half (one-third) the crystal length. Improved resolution via narrower crystals and worse resolution due to non-colinearity were both overshadowed by DOI effects at 75% off-center positions in 30 mm deep crystals with 10 mm of DOI resolution.

Design considerations and initial performance of a 1.2 cm/sup 2/ beta imaging intra-operative probe

A novel small area beta (/spl beta//sup /spl plusmn//) detector is under development for nuclear emission imaging of surgically exposed, radiolabeled tumor beds. The imaging device front-end consists of a 0.5 mm thick by 1.25 cm diameter CaF/sub 2/(Eu) scintillator disk coupled to a rigid bundle of 2 mm diameter double clad optical fibers through a polystyrene light diffuser. The detector area (1.2 cm/sup 2/) was determined by the requirement of introducing the probe into small cavities, e.g. during neuro-surgical lesion resection, but large enough to produce images of clinical significance. Flexible back-end optical fibers (1.9 m long) were coupled to the front-end components allowing /spl sim/75 photo-electrons to be detected for mean beta energies of 250 keV, indicating that sufficient signal can be obtained with clinical beta emitters (e.g. /sup 18/F, /sup 131/I). The long flexible fibers guide the scintillation light to a Philips XP1700 series, fiber optic faceplate, Multi-Channel PMT. The parallel MC-PMT outputs are fed into a variable gain, charge divider network and an i-V pre-amplifier/line driver network, whose resulting four outputs are digitized and histogrammed with standard Anger positioning logic. The various components in the imaging chain were evaluated and optimized by both simulations and measurements. Line spread functions measured in the 10.8 mm FOV were 0.50 mm /spl plusmn/0.038 mm and 0.55 mm /spl plusmn/0.065 mm FWHM in X and Y, respectively. A 20% variation in pulse height and minimal variation in spatial resolution was observed. The differential image uniformity was measured to be /spl plusmn/15.6% with /spl sim/150 cts/pixel. Preliminary images show excellent reproduction of phantom activity distributions.

A miniature phoswich detector for gamma-ray localization and beta imaging

A combined /spl gamma/-ray probe//spl beta//sup /spl plusmn// imaging detector was created by modification of a beta imaging intra-operative probe. This phoswich detector consists of a thin CaF/sub 2/(Eu) /spl beta//sup /spl plusmn// imaging disk, coupled through a light diffuser to 7 or 19 parallelepiped high-Z scintillation crystals arranged hexagonally in columns. These elements are either 2/spl times/2/spl times/10 mm/sup 3/ or 4/spl times/4/spl times/10 mm/sup 3/ LSO, GSO or EGO crystals, and act as both light guides for the imaging light and as intrinsic /spl gamma/ detectors. The coincidence between the CaF/sub 2/(Eu) and phoswich scintillators identifies /spl beta//sup +/ or true /spl beta//sup -/ events and suppresses accompanying annihilation or background /spl gamma/ events. The phoswich crystals are in turn coupled through optical fibers to a multi-channel PMT with positioning and discrimination electronics. Various characteristics of this novel imaging phoswich detector are investigated and presented for each of the phoswich combinations including: detector anisotropy, effects of time blocking and energy windows on phoswich imaging, intrinsic singles and phoswich spectral and spatial resolution characteristics, and phoswich imaging ability within /spl gamma/ background environments. The performance characteristics of the assembled prototype devices demonstrate that in surgery, this device can be used three ways: (1) rapid localization of /spl gamma/ emitting radionuclides with the modestly directional, self-collimated phoswich crystals; (2) high resolution /spl beta//sup m/aging or coarse (annihilation background corrupted) /spl beta//sup +/ imaging in singles mode; or (3) /spl beta//sup +/ imaging in coincidence mode with reduced /spl gamma/ background contamination.

A YSO/LSO phoswich array detector for single and coincidence photon imaging

The performance of a phoswich array detector module for possible use in a combined single and coincidence photon imaging system has been evaluated. The detector consist of a linear array of discrete 4/spl times/4/spl times/15 mm/sup 3/ YSO elements coupled to a combined detector array/light guide of LSO, 10 mm thick. Since the scintillation light decay time is different in YSO and LSO (70 and 40 ns, respectively), events originating from the two detector materials can be separated by pulse shape discrimination. The front layer of YSO could then be used for detection of low energy, single photon events and the LSO layer for coincidence detection of annihilation radiation. The light collection of the PMTs coupled to the detector was found to be adequate to accurately identify each detector element in the array using the same positioning logic used in conventional EGO block detectors. The average energy resolution of the YSO elements at 140 keV for the block detector was found to be 14.5% FWHM, ranging from 13.8 to 15.4%. Spatial resolution of the detector block in single photon mode, using a high resolution collimator (geometric resolution 6.5 mm at 10 cm) was measured by scanning a /sup 99m/Tc line source. The resolution at 5 and 10 cm from the collimator face was found to be 5.9 and 8.5 mm FWHM, respectively.

Depth of interaction for PET using segmented crystals

The authors investigate the use of segmented scintillation crystals to obtain depth of interaction (DOI) information in positron emission tomography (PET) detectors. Coupling segmented crystals to a photodetector at one end yields distinct pulse heights from the different segments and the depth of interaction is determined to within the segment length. The method is simple to implement because no additional instrumentation is required. Since the technique uses light attenuation, a scintillator with sufficient light yield is required. The authors found that 4/spl times/4/spl times/10 mm/sup 3/ segments of lutetium oxyorthosilicate (LSO) provide enough light for this application. The energy spectra of segmented crystals show clearly distinct photopeaks for each segment and the photopeak separation can be adjusted by altering the the optical couple between segments. An ambiguity exists between photopeak events from an attenuated crystal and scattered events from the crystal without attenuation. This leads to the same error as a system with no depth information. Using 10 mm segments the authors find that the fraction of the events in the lower energy window due to scatter contamination in 2-segment (3-segment) crystals increases from 8% (13%) to 32% (60%) for a source in air vs. behind 16 cm of water. In imaging situations with minimal scatter medium (e.g. rodents) DOI is accurately provided by the segmented crystals, while in the presence of scatter (e.g. WB, cardiac) the method is only partially effective, and potential improvements in scanner cost, sensitivity and/or image quality must be carefully evaluated.

PSPMT and photodiode designs of a small scintillation camera for imaging malignant breast tumors

The authors are investigating the concept of utilizing a small gamma ray scintillation camera to help identify and localize malignant breast tumors after a in a mammogram. Excellent sensitivity and specificity for malignant breast tumors has been achieved using conventional nuclear medicine scintillation cameras with certain /sup 99m/Tc-labeled radiopharmaceuticals. However, because of the large size, low image resolution and high cost of these devices, they are not ideal for use in breast imaging in a mammography suite. A dedicated miniaturized camera would allow imaging at angles that are physically impossible with the standard camera. These lateral views would not include the background activity from the heart and liver. In addition, with a potentially higher intrinsic resolution, shorter collimator and if breast compression is applied, a small camera could significantly improve the sensitivity and signal to noise ratio for the scintillation imaging method. The authors are exploring 2 different photodetector technologies for a small prototype camera development. The first uses a position sensitive photomultiplier (PSPMT) as the photodetector, the second, an array of silicon PIN photodiodes (PD). In this report, the authors present imaging results obtained with a NaI(Tl)-PSPMT design, and the design features, expected performance and relevant energy and position measurements obtained for a test CsI(TI)-PD device.

HgI/sub 2/ polycrystalline films for digital X-ray imagers

This paper describes recent results obtained with mercuric iodide (HgI/sub 2/) polycrystalline films that we have produced. The ultimate goal of this effort is to develop a new detector technology for digital X-ray imaging based on HgI/sub 2/ polycrystalline films coupled to large-area flat-panel amorphous silicon thin-film transistor-addressed readout arrays. We have employed two approaches for producing the polycrystalline films: 1) thermal evaporation (sublimation) and 2) deposition of films from various solutions. The 50- to 150-/spl mu/m-thick films were characterized with respect to their electrical properties and in response to ionizing radiation. The leakage current was about 40 pA/cm/sup 2/ at an operating bias voltage of /spl sim/50 V. Signals from the HgI/sub 2/ polycrystalline detectors, in response to ionizing radiation, compare favorably to the best published results for all high Z polycrystalline films grown elsewhere, including TlBr, PbI/sub 2/, and HgI/sub 2/.

Annihilation /spl gamma/-ray background characterization and rejection for a small beta camera used for tumor localization during surgery

The authors have developed a miniature (1.2 cm/sup 2/) beta-ray camera prototype to assist a surgeon in locating the margins of a resected tumor. With this technique, one directly detects betas emitted from exposed radio-labeled tissue. When imaging positron emitting radiopharmaceuticals, annihilation gamma ray interactions in the detector can mimic those of the betas. The extent of the background contamination depends on the detector, geometry and the tumor specificity of the radiopharmaceutical. The authors have characterized the effects that annihilation gamma rays have on positron imaging with their small camera. They studied beta and gamma ray detection rates and imaging using small positron or electron sources directly exposed to the detector to simulate hot tumor remnants, and a cylinder filled with /sup 18/F to simulate annihilation background from the brain. For various ratios of phantom head/tumor activity, a background gamma rate of 2.0 cts/sec//spl mu/Ci was measured in the CaF/sub 2/(Eu) detector. The authors present two gamma-ray background rejection schemes that require a /spl beta/-/spl gamma/ coincidence. The first configuration uses a high efficiency scintillator coincidence shield, the second, a phoswich. Results show that these coincidence methods work with /spl sim/99% gamma ray rejection efficiency.

GAMMA AND BETA INTRA-OPERATIVE IMAGING PROBES

Small area ( -1.5 cm2) scintillation cameras for imaging gammas and betas using inter-changeable detector front ends were built and characterized. Components common to both emission imaging cameras include: (1) fiber optic bundles 2-3m long, comprised of multi-clad fibers which connect the scintillation detector to (2) an MC-PMT; (3) parallel MC-PMT outputs feed a resistive positioning network and i-V converter/line driver network which produce balanced +X, -X, + Y, and - Y outputs; and (4) four ADCs and a Macintosh PC for system control and image display. The beta and gamma devices used distinct scintillation detectors which were characterized by both simulation and measurement. The beta camera utilized a 0.5 mm by 1.25 cm 4 CaF,(Eu) scintillation crystal coupled, through a diffusing light guide, to 19 2-mm 4 optical fibers. These front-end fibers are in turn coupled by a more flexible fiber bundle to the MC-PMT. CaF,(Eu) has high light output, high beta sensitivity, and low gamma sensitivity. Image signals are histogrammed and displayed after Anger logic computations are performed on digitized signals. The beta camera has <0.6mm FWHM intrinsic resolution. The gamma camera concept was tested with matrices of discrete 1 x 1 mm2 and 2 x 2 mm’ CsI(T1) and NaI(T1) crystals of various lengths, and 3 mm thick continuous crystals. Configurations using 4 x 4 element matrices with one-to-one coupling between crystals and fiber channels. and light diffusers between each crystal matrix and fibers were evaluated. The continuous crystals were coupled directly to the fiber optics with signal and data processing analogous to the beta camera. Coupling of discrete crystals to fiber optics by both methods gave essentially perfect identification of the crystal of interaction, allowing spatial resolution to be defined by the crystal size and collimator. The continuous crystal gamma camera gave intrinsic resolution of - 1.4mm FWHM.

Evaluation of breast tumor detectability with two dedicated, compact scintillation cameras

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.