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Why human skills and expertise, not technical tools, are what make projects succeed. The project is the basic unit of work in many industries. Software applications, antiviral vaccines, launch-ready spacecraft: all were produced by a team and managed as a project. Project management emphasizes control, processes, and tools—but, according to The Smart Mission, that is not the right way to run a project. Human skills and expertise, not technical tools, are what make projects successful. Projects run on knowledge. This paradigm-shifting book—by three project management experts, all of whom have decades of experience at NASA and elsewhere—challenges the conventional wisdom on project management, focusing on the human dimension: learning, collaboration, teaming, communication, and culture. The authors emphasize three themes: projects are fundamentally about how teams work and learn together to get things done; the local level—not an organization's upper levels—is where the action happens; and projects don't operate in a vacuum but exist within organizations that are responsible to stakeholders. Drawing on examples and case studies from NASA and other organizations, the authors identify three project models—micro, macro, and global—and their different knowledge needs. Successful organizations have a knowledge-based culture. Successful project management guides the interplay of knowledge, projects, and people.

Coping with frequent unexpected events requires an organizational culture that allows the project manager to exercise a great amount of flexibility. The authors' studies, employing experiential data collected from more than 150 successful project managers affiliated with more than 20 organizations, indicate that today's successful project managers cope with unexpected events by a combination of the traditional and agile approaches, assuming four roles. Developing collaboration requires project managers to be people-oriented. Integrating planning and review with learning requires them to be information-oriented. Preventing major disruptions requires them to be action-oriented. Finally, maintaining forward momentum, which is pursued throughout a project, requires them to adopt all three orientations

This chapter covers a descriptive project practitioner-centered knowledge model derived from experience in developing knowledge services (KS) at the National Aeronautics and Space Administration (NASA). It involves an organizational knowledge systems perspective that better negotiates rapid change and accelerated learning in data-rich complex project environments. The authors discuss the modern project environment and cover the strategic imperatives that guide the design and development of a knowledge services (KS) model for individuals, teams, and organizations. The operational components of knowledge capture and retention, sharing and application, and discovery and creation are specified as core processes with individual and organizational inputs of capabilities and expectations. Considerations for the design of KS at NASA are discussed, including governance, federated approach, mapping, biases and heuristics, and practitioner capabilities. A case study covering the James Webb telescope is then used to illuminate the presented strategic imperatives and model, and the article closes with a summary and recommendations for future model application. This model allows organizations to validate and add imperatives within their context to better design knowledge services for diverse challenges and opportunities.

Coronary artery disease (CAD) is the leading cause of death in most developed countries. A large portion of CAD is caused by rupture of unstable plaque, which is not detected by current diagnostic methods. By labeling the unstable plaque with beta emitting radioisotopes, it is possible to detect the plaques with a very narrow in situ detector system that is placed with a coronary angiography guidewire. Our current detection system uses fiber optics, with a single detector at the tip of each fiber, connected to a position sensitive photomultiplier tube (PSPMT). The small diameter of the coronary artery limits our current system to six optical fibers. To increase the number of pixels, we are studying a multi-detector concept for each fiber. By having multiple scintillating fibers coupled in series, with each scintillating fiber emitting a unique wavelength spectrum, the number of pixels can be doubled or tripled. The light is separated before reaching the PSPMT with a single diffraction grating. The PSPMT separates the detectors in one direction by fiber, and in the other direction by the spectral splitting of the different wavelengths of the scintillators.

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.

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/.

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.

Cerium-doped lutetium oxyorthosilicate (LSO) is amongst the most promising new scintillators discovered in almost five decades, with a unique combination of important properties for X and gamma-ray spectroscopy, namely: high density, fast decay, and large light yield. LSO seems to be a prime candidate to replace BGO in PET systems. However, the practical utilization of LSO is hindered by difficulties related to crystal growth (Czochralski method) due to the high temperatures employed. A new approach has been developed using a low-temperature crystal growth technology to produce scintillating LSO crystals. Light transparent polycrystalline LSO samples of a few mm/sup 3/ in volume were grown and characterized by XRD, optical absorption, light decay measurement and gamma-ray spectral response. The properties of the new crystals compared well with high-quality crystals grown by the Czochralski method.

We are developing a detector module consisting of an array of Lutetium Orthosilicate (LSO) scintillation crystals coupled to a new type of silicon photodetector array specifically designed to provide the excellent energy and timing resolution needed for positron emission tomography. We have designed, fabricated and evaluated prototypical detector structures including single 2/spl times/2 mm/sup 2/ pixel test structures and eight-by-eight arrays of the 2/spl times/2 mm/sup 2/ pixels. Leakage current of /spl sim/1 pA per pixel and electronic noise of /spl sim/26 e/sup -/ RMS were obtained with thermoelectric cooling. Quantum efficiency was 40% at 460 nm and 70% at 560 nm. Variation in photoresponse (/spl plusmn//spl sigma/%) over the pixel active area was less than 0.3%. The coincidence resolving time with a LSO crystal was 16 ns. Energy resolutions of 6%, 10% and 13% FWHM for Cs-137 (662 keV) were obtained at the pixel level with 2/spl times/2/spl times/4 mm/sup 3/ CsI(Tl), 2/spl times/2/spl times/10 mm/sup 3/ BGO, and 2/spl times/2/spl times/10 mm/sup 3/ LSO crystals, respectively.

Cerium-doped lutecium orthosilicate (LSO) is the most promising scintillator discovered in almost five decades. It exhibits a unique combination of important properties for x and gamma-ray spectroscopy: high density, fast decay, and large light yield. However, the practical use of LSO is hindered by difficulties related to its fabrication as a single crystal by the Czochralski method. We report on the usefulness of the sol-gel process in obtaining lutecium silicate scintillators. Upon appropriate drying and firing, lutetium silicate crystals can be grown in a silica matrix. The bulk, polycrystalline transparent scintillators are characterized by XRD, optical absorption, light decay measurement and gamma-ray spectral response. Their properties are comparable to that of traditional LSO single crystals.

Intraoperative probes have been employed to assist in the detection and removal of tumors for more than 50 years. For a period of about 40 years, essentially every detector type that could be miniaturized had been tested or at least suggested for use as an intraoperative probe. These detectors included basic Geiger-Muller (GM) tubes, scintillation detectors, and even state-of-the-art solid state detectors. The radiopharmaceuticals have progressed from 32PO4- injections for brain tumors to sophisticated monoclonal antibodies labeled with iodine-125 for colorectal cancers. The early work was mostly anecdotal, primarily interdisciplinary collaborations between surgeons and physical scientists. These collaborations produced a few publications, but never seemed to result in an ongoing clinical practice. In the mid 1980s, several companies offered basic gamma-detecting intraoperative probes as products. This led to the rapid development of radioimmunoguided surgery (RIGS) and sentinel node detection as regularly practiced procedures to assist in the diagnosis and treatment of cancer. In recent years intraoperative imaging probes have been developed. These devices add the ability to see the details of the detected activity, giving the potential of using the technique in a low-contrast environment. Intraoperative probes are now established as clinical devices, they have a commercial infrastructure to support their continued use, and there is ongoing research, both commercial and academic, that would seem to ensure continued progress and renewed interest in this slowly developing field.

A physical phantom that can be used to evaluate the ability of a Positron Emission Tomography (PET) system to perform quantitative dynamic measurements by simulating realistic tissue kinetics was investigated. In this work, the use of ion exchange resin to imitate tracer kinetics in tissue was investigated as a means to construct a realistic dynamic phantom for PET studies. A unit cell perfusion system that demonstrates the feasibility of such an approach was fabricated. The system utilizes competitive binding of /sup 13/NH/sub 4//sup +/ and Na/sup +/ for cation exchange resins. Due to the high binding affinity of /sup 13/NH/sub 4//sup +/ to cation exchange resin, efficient tracer uptake can be completed, followed by gradual washout of tracer from binding sites by an aqueous solution of NaCl. The clearance rates of tracer can be modified to achieve the desired temporal characteristics by changing either the concentration of NaCl eluent or its flow rate. This unit cell perfusion system was also incorporated into a simple cylinder phantom for dynamic PET imaging. Results from the PET phantom images demonstrated the feasibility of using ion-exchange resin to create dynamic phantoms for PET with easily controlled kinetic characteristics.

An adaptive, nonlinear image domain filtering strategy is described which improves positron emission tomography (PET) images. The method was formulated to improve on the linear, low-pass filtering typically applied to each projection in the filtered back-projection (FBP) reconstruction algorithm. The algorithm is a potential alternative to linear smoothing which reduces noise but degrades resolution; this method uses the FBP algorithm for reconstruction, but aims to incorporate some of the statistical information and nonlinear smoothing utilized in iterative reconstruction algorithms. The approach uses sinogram segmentation to separate the sinogram elements with higher and lower signal-to-noise ratios, and then reconstruct each with FBP using a more appropriate choice of filter and cut-off frequency. Also, this algorithm addresses the radial streak artifacts introduced by FBP. The algorithm was evaluated using simulations clinical data of cardiac PET studies on an ECAT 931 PET scanner. The initial results suggest that this technique has advantages over the current clinical protocol. Images processed with the method show generally improved visual image quality and reduced radial streaks without the introduction of artifacts. In simulations, increased contrast recovery and resolution are realized without an increase in the background noise of the reconstructed images.

A key parameter in the practical application of filtered back-projection (FBP), the standard clinical image reconstruction algorithm for positron emission tomography (PET), is the choice of a low-pass filter window function and its cut-off frequency. However, the filter windows and cut-off frequencies for clinical reconstruction are usually chosen empirically, based on a small sample of images and filters. By considering the features of the signal and noise spectra in a sinogram, the desired image resolution, and the signal-to-noise ratio (SNR) of the filtered sinogram, the authors investigated the possibility of establishing a methodology for informed selection of a filter function and cut-off frequency for use with FBP. Simulations of sinogram data similar to whole body or cardiac studies provided information on the signal and noise frequency-domain spectrum of noisy projection data. The improvements in the SNR with different filter windows and cut-off frequencies were evaluated and compared. The projection spectrum SNR measure did not prove to be an accurate indicator of subjective image quality or lesion detectability with variations in Poisson noise and image resolution.

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.

The basic concepts for a dual detector b‘ imaging probe incorporating a method of c-ray background suppression were evaluated. The test devices consisted of a 1.2 cm2 b‘ imaging CaF 2 (Eu) scintillator disk optically coupled to a 1.7 mm thick, area-matched transparent di⁄user disk, which in turn was coupled to sets of 2] 2o r 4 ] 4m m 2by 10 mm long BGO, GSO or LSO scintillators, which acted as both c-detectors as well as light guides for the CaF 2 (Eu) scintillation light, which in turn are attached to a fiber optic bundle coupled to a multichannel photomultiplier. Pulse shape discrimination allowed identification of the detector of interaction and coincidence between the two detectors allowed selection of b‘-only events. This type imaging detector system was shown to be capable of producing high-resolution images of b‘ distributions while suppressing the events due to the 511 keV annihilation radiation background. GSO was found to be the detector of choice for this application. ( 1998 Elsevier Science B.V. All rights reserved.

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.

A 256-element mercuric iodide (HgI/sub 2/) detector array has been developed which is intended for use as an intraoperative gamma camera (IOGC). The camera is specifically designed for use in imaging gamma-emitting radiopharmaceuticals (such as 99m-Tc labeled Sestamibi) incorporated into brain tumors in the intraoperative surgical environment. The system is intended to improve the success of tumor removal surgeries by allowing more complete removal of subclinical tumor cells without removal of excessive normal tissue. The use of HgI/sub 2/ detector arrays in this application facilitates construction of an imaging head that is very compact and has a high SNR. The detector is configured as a cross-grid array. Pixel dimensions are 1.25 mm squares separated by 0.25 mm. The overall dimension of the detector is 23.75 mm on a side. The detector thickness is 1 mm which corresponds to over 60% stopping at 140 keV. The array has good uniformity with average energy resolution of 5.2/spl plusmn/2.9% FWHM at 140 keV (best resolution was 1.9% FWHM). Response uniformity (/spl plusmn//spl sigma/) was 7.9%. A study utilizing realistic tumor phantoms (uptake ratio varied from 2:1 to 100:1) in background (1 mCi/l) was conducted. SNRs for the reasonably achievable uptake ratio of 50:1 were 5.61 /spl sigma/ with 1 cm of background depth ("normal tissue") and 2.74 /spl sigma/ with 4 cm of background for a 6.3 /spl mu/l tumor phantom (/spl sim/270 nCi at the time of the measurement).

In this work, gamma sensitive intra-operative imaging detectors using CsI(Na), CsI(Tl) and NaI(Tl) scintillators coupled through optical fibers were investigated with simulations and measurements. Three coupling methods were employed between the detectors and optical fiber imaging array: discrete crystals coupled one-to-one with the fibers; discrete crystals coupled through a light diffuser to an unmatched fiber array; and continuous, light sharing crystals coupled directly to the fiber optic array. The discrete crystals were arranged in 4/spl times/4 element arrays of sizes ranging from 1/spl times/1/spl times/2.5 mm/sup 3/ up to 2/spl times/2/spl times/6 mm/sup 3/. CsI(Na) had the best pulse height and single crystal energy resolution (44% at 140 keV) when coupled through fibers to the MC-PMT photodetector. All discrete arrays were completely resolved, indicating that the intrinsic spatial resolution corresponded to the crystal size. The spatial resolution for the 12.5 mm /spl phi//spl times/3.5 mm continuous crystals was /spl sim/1.2 mm FWHM, with relative pulse heights of 1.5 to 5.5 times lower than for the discrete array detectors.

We are exploring the possibility of using PIN photodiodes to readout the scintillation crystals used in positron emission tomography (PET) detector designs. Semiconductor photodetectors typically have a lower signal to noise ratio than photomultiplier tubes (PMTs). However, they have the advantage of compactness, and, thus, scintillation crystal readout schemes not available to PMTs because of their size and geometry limitations, are readily available to photodiodes. With current PET detector designs, only a small fraction of the available scintillation light, created from 511 keV gamma ray interactions within the crystal, is collected. Scintillation light collection studies were performed for several crystal geometries and surface treatments using both simulations and measurements. In this report, we present a feasible photodiode readout scheme that allows greater than 90% of the available scintillation light created in either BGO or LSO scintillation crystals to be collected by the photodetector. This improvement in light collection with the photodiode readout somewhat compensates for its lower inherent signal to noise ratios and makes it feasible for use in PET detectors. A coincident timing spectrum resolution of 9.4 ns FWHM was measured for 511 keV interactions with one LSO crystal coupled to a photodiode, the other to a PMT.

Multi-element (4/spl times/4) imaging arrays with high resolution collimators, size matched to discrete CsI(Tl) scintillator arrays and mercuric iodide photodetector arrays (HgI/sub 2/ PDA) were developed as prototypes for larger 16/spl times/16 element arrays for breast imaging. The compact nature of the arrays allows detector positioning in close proximity to the breast to eliminate activity not in the line-of-sight of the collimator, thus reducing image background. Short collimators, size matched to 11.5/spl times/1.5 mm/sup 2/ scintillators show a factor of 2 and 3.4 improvement in spatial resolution and efficiency, respectively, compared to high resolution collimated gamma cameras for the anticipated compressed breast geometries. Monte Carlo simulations, confirmed by measurements, demonstrated that scintillator length played a greater role in efficiency and photofraction for 140 keV gammas than cross sectional area, which affects intrinsic spatial resolution. Simulations also demonstrated that an increase in the ratio of scintillator area to length corresponds to an improvement in light collection. Electronic noise was below 40 e/sup -/ RMS indicating that detector resolution was not noise limited. The high quantum efficiency and spectral match of prototype unity gain HgI/sub 2/ PDAs coupled to 1/spl times/1/spl times/2.5 mm/sup 3/ and 2/spl times/2/spl times/4 mm/sup 3/ CsI(Tl) scintillators demonstrated energy resolutions of 9.4% and 8.8% FWHM at 140 keV, respectively, without the spectral tailing observed in standard high-Z, compound semiconductor detectors. Line spread function measurements matched the scintillator size and pitch, and small, complex phantoms were easily imaged.

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.

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

To improve spatial resolution, positron emission tomography (PET) systems are being developed with finer detector elements. Unfortunately, using a smaller crystal size increases intercrystal Compton scatter and X-ray escape crosstalk, causing positioning errors that can lead to degradation of image contrast. We investigated the use of extremely thin (

We have developed a novel detector device, namely: a mercuric iodide photodetector array (HgI 2 PDA) to be used with a scintillating crystal for X-ray and gamma-ray imaging. A HgI 2 PDA prototype with 16-pixels has been evaluated with segmented scintillators (CsI(Tl)) coupled to the HgI 2 PDA. Pixel energy resolution of 9.4% FWHM has been obtained at 122 keV for a 4 mm thick scintillator corresponding to 72% stopping at 140 keV, and pixel energy resolution of 11% FWHM has been obtained at 122 keV with the 10 mm thick scintillator.

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.

The authors are developing a 1-2 cm/sup 2/ area camera for imaging the distribution of beta-emitting radiopharmaceuticals at the surface of tissue exposed during surgery. The front end consists of a very thin continuous or segmented scintillator sensitive to betas (positrons or electrons) of a few hundred keV, yet insensitive to gamma rays. The light from the scintillator is piped through clear fibers to the photon detector (PD). This approach requires that a sufficient number of scintillation photons be transported from the scintillator, through the fibers to the PD. The scintillator, reflector, surface treatments, geometry, fiber light guides, and optical couplings must be optimized. The authors report here on efforts made to optimize the light collection from

A hybrid attenuation correction technique (ACT) is under development for /sup 18/F-FDG total body positron emission tomography (PET). With a short transmission scan of the thorax, any time within a few days of the imaging session, this technique can correct for attenuation in the entire body. Segmentation, registration, and active contour finding techniques are applied to both emission and short transmission images to locate and map the major attenuating structures in the body. This technique eliminates the need for the patient to remain still from the start of the transmission scan to the end of the emission scan without the added noise of simultaneous or post emission transmission scan measurements. The results of volunteer studies are comparable to standard measured ACT, both visually and quantitatively. The efficient use of scanner time and improved patient comfort make this technique particularly suitable for clinical imaging.

Development of a small area (1-2 cm/sup 2/) in situ beta imaging device includes optimization of the front end scintillation detector which is fiber optically coupled to a remote photon detector. Thin plastic scintillation detectors, which are sensitive to charged particles, are the ideal detectors due to the low sensitivity to ambient gamma backgrounds. The light output of a new binary plastic scintillator was investigated with respect to increasing concentrations of the fluor ([0.5%] to [2.0%] by weight), and varying thickness cylindrical configurations of the intended imaging detector. The new fluor had an emission maximum increasing from 431 to 436 nm with increasing fluor concentration The decay time(s) had two components (0.38, 1.74 nsec). There was an /spl sim/20% increase in light output with increasing fluor concentration, measured with both /sup 204/Tl betas and conversion electrons from /sup 207/Bi. The highest light output of this new scintillator was measured to be /spl sim/30% lower than BC404. Simulations predicted the 1.5 mm scintillator thickness at which light output and energy absorption for /spl sim/700 keV electrons (e.g. from /sup 204/Tl, /sup 18/F) were maximized, which corresponded with measurements As beta continua are relatively featureless, energy calibration for the thin scintillators was investigated using Landau distributions, which appear as distinct peaks in the spectra. As the scintillators were made thinner, gamma backgrounds were shown to linearly decrease.

A nineteen element mercuric iodide (HgI 2 ) detector array has been developed as a prototype for a larger (169 element) array, which is intended for use as an intra-operative gamma camera (IOGC). This work is motivated by the need for identifying and removing residual tumor cells after the removal of bulk tumor, while sparing normal tissue. Prior to surgery, a tumor seeking radiopharmaceutical is injected into the patient, and the IOGC is used to locate and map out the radioactivity. The IOGC can be used with commercially available radioisotopes such as 201 Tl, 99m Tc, and 123 I which have low energy X- and gamma-rays. The use of HgI 2 detector arrays in this application facilitates construction of an imaging head that is very compact and has a high signal-to-noise ratio. The prototype detectors were configured as discrete pixel elements joined by fine wires into novel pseudo crossed-grid arrays to promote improved electric field distribution compared with previous designs, and to maximize the fill factor for the expected circular probe shape. Pixel dimensions are hexagonal with 1.5 mm and 1.9 mm diameters separated by 0.2 mm thick lead septa. The overall detectors are hexagonal with a diameter of ∼1 cm. The sensitive detector thickness is 1.2 mm, which corresponds to >99% efficiency at 59 keV and 67% efficiency at 140 keV. Row, column, and pixel spectra have been measured on the prototypical detector array. Energy resolution was found to vary with the width of the row/column coincidence window that was applied. With the low edge of the coincidence window at 30% below the photopeak, pixel energy resolutions of 2.98% and 3.88% FWHM were obtained on the best individual pixels at 59 keV ( 241 Am) and 140 keV ( 99m Tc), respectively. To characterize this array as an imaging device, the spatial response of the pixels was measured with stepped point sources. The spatial response corresponded well with the pixel geometry, indicating that the spatial resolution was determined by the pixel geometry.

We have compared the high resolution BGO detector blocks from the EXACT HR PET system which use two dual-cathode rectangular photomultiplier (PM) tubes with a new block design, the EXACT HR PLUS, which uses four round PM tubes. Despite the lower coupling area between photocathode and scintillator, the HR PLUS block compares favorably with the HR block. The energy resolution averages 20% for the HR PLUS block and 23% for the HR block, with efficiency variations of 17% in both blocks. Additional measurements were carried out on the HR PLUS block to characterize depth of interaction effects and cross-talk between elements. Coincidence line spread function measurements had a FWHM of 3.0 mm in the axial direction and 2.9 mm in the transaxial direction. In light of these results, limitations of the BGO block design are discussed and some solutions proposed. >

A Monte Carlo simulation has been developed to simulate and correct for the effect of Compton scatter in 3-D acquired PET brain scans. The method utilizes the 3-D reconstructed image volume as the source intensity distribution for a photon-tracking Monte Carlo simulation. It is assumed that the number of events in each pixel of the image represents the isotope concentration at that location in the brain. The history of each annihilation photon's interactions in the scattering medium is followed, and the sinograms for the scattered and unscattered photon pairs are generated in a simulated 3-D PET acquisition. The calculated scatter contribution is used to correct the original data set. The method is general and can be applied to any scanner configuration or geometry. In its current form the simulation requires 25 hours on a single Sparc10 CPU when every pixel in a 15-plane, 128/spl times/128 pixel image volume is sampled, and less than 2 hours when 16 pixels (4/spl times/4) are grouped as a single pixel. Results of the correction applied to 3-D human and phantom studies are presented. >

A spilldown correction method is proposed for the thallium-201 window image in simultaneous dual-isotope technetium-99m/thallium-201 single-photon emission tomographic (SPET) imaging based on a single acquisition into three energy windows. In this method, images are simultaneously acquired in two standard energy windows over the99mTc and201Tl photopeak regions and a third spilldown window adjacent to the201Tl window. Using a Monte Carlo simulation of SPET, the fractional amount of99mTc and201Tl spilldown in the201Tl window with respect to the total counts from the spilldown window, k12, was calculated for simulated images of point sources at varying depths within a water-filled elliptical tub phantom. When applied to experimental acquisitions, k12, multiplied by the total counts from the spilldown window, is then subtracted from the201Tl window image to produce the corrected image. However, for successful applications in SPET, k12 must be determined on a projection-by-projection basis since k12 is depth dependent. Thus, a regression relation was obtained between k12 and the total count ratio of the spilldown to99mTc windows, k23. The spilldown correction method was applied to201T1 photopeak images of an extended source distribution in uniform and nonuniform attenuating media with dual-isotope99mTc/201Tl and single-isotope201Tl. A marked improvement in image contrast was observed between the corrected and uncorrected201Tl window images. The average count ratio of uncorrected dual-isotope201Tl/single-isotope201Tl was 3.08 for uniform and 2.99 for non-uniform attenuating media. When corrected, this value approached unity (0.97 and 1.08) and the corrected dual-isotope images were visually similar The presented method in which spilldown correction for the201Tl photopeak window is performed in a single acquisition on a projection-by-projection basis in simultaneous dual-isotope SPET imaging could form the basis for a rapid and accurate clinical dual-isotope protocol independent of radiotracer administration.

We report the development of high resolution detector modules for use in PET imaging of small laboratory animals. The detectors are based on thin blocks of NaI coupled to a light guide which serves to share the scintillation light between four photomultiplier tubes. Monte Carlo simulations of photon transport within the scintillator and light guide, together with initial experimental results, suggest that a detector resolution of approximately 2 mm is obtainable using 18 F or 11 C labeled tracers.

Two state-of-the-art modular PET block detectors using either discrete or pseudo-discrete BGO crystals coupled to two dual PMTs, utilizing different light sharing schemes, were evaluated. Both detectors were approximately 30 mm thick, while each element of the GE 6/spl times/6 block detector was 4 mm/spl times/8.4 mm, and each element of the CTI 8/spl times/7 block detector was 2.8 mm/spl times/5.8 mm. In addition to measurements with gamma sources, the detectors were also irradiated with a Sr/Y-90 beta source to evaluate the performance of the block detectors without inter-element scatter of annihilation photons. The best and worst energy resolutions of individual elements at 511 keV were 21.8% and 44.2% for the GE detector and 22.7% and 42.5% for the CTI detector. The peak to valley ratios in the detector identification ratio histograms were generally better than 3 to 1. Measurements with beta sources indicate that light sharing in both detectors is a large component of event mispositioning along with inter-detector scatter of the annihilation photons. >

An intraoperative beta probe was designed, built, and tested for detection of radio-labeled malignant tissues that has the advantage of being selectively sensitive to beta while insensitive to gamma radiation. Since beta radiation (electrons or positrons) has a short range in tissue, this probe is ideal for detecting tracers in tumors at the surface of the surgical field. This probe contains a plastic scintillation detector sensitive to beta rays and to a lesser degree some background gamma rays. A second detector counts spurious gamma rays and allows for their subtraction from the activity measured by the first detector. Sensitivity of the dual probe for I-131 and F-18 was measured to be 108 counts/s/kBq (4000 counts/s/microCi). The dual-detector probe faithfully measured the 10:1 "tumor" to background ratio of radioactivity concentrations in a simulated environment of a tumor in the presence of intense background 511 keV photons. In another phantom experiment, simulating abdominal tumor deposits with various realistic I-131 radioactive concentrations, the probe was able to accurately identify tumors of approximately 50 mg with a tumor/normal radioactivity concentration of 3/1 in 10 s.

A nineteen element mercuric iodide (HgI2) detector array has been developed in order to investigate the potential of using this technology for in-vivo x-ray and gamma-ray imaging. A prototype cross-grid detector array was constructed with hexagonal pixels of 1.9 mm diameter (active area = 3.28 mm2) and 0.2 mm thick septa. The overall detector active area is roughly 65 mm2. A detector thickness of 1.2 mm was used to achieve about 100% efficiency at 60 keV and 67% efficiency at 140 keV The detector fabrication, geometry and structure were optimized for charge collection and to minimize crosstalk between elements. A section of a standard high resolution cast-lead gamma-camera collimator was incorporated into the detector to provide collimation matching the discrete pixel geometry. Measurements of spectral and spatial performance of the array were made using 241-Am and 99m-Tc sources. These measurements were compared with similar measurements made using an optimized single HgI2 x-ray detector with active area of about 3 mm2 and thickness of 500 μm.

The authors discuss methods for reducing the computation time for 3-D PET (positron emission tomography) reconstruction through the use of fast computer hardware, vector and parallel programming techniques, and algorithm optimization. The strengths and weaknesses of i860 microprocessor based workstation accelerator boards are investigated in implementations of 3-D PET reconstruction. One key issue in achieving speed on the SuperCard 2 has been minimizing memory accesses, and eliminating redundant calculations through approximations and look-up tables. This has allowed the calculation time to decrease from 20 minutes on the MicroVax 3200 to 146 seconds on a dual i860 system. >

Removal of the interplane septa and configuration of a typical multislice PET scanner to accept all possible coincidence lines of response leads to a fivefold increase in sensitivity. This can be of value in regional CBF studies using bolus 15O-labeled water injections, allowing the injected dose to be reduced by a factor of 4, while maintaining the same number of noise equivalent counts. Thus, for a given cumulative dose limit, four times as many studies can be performed in a single subject. Data from the three-dimensional Hoffman brain phantom, closely matched to count rates seen in human studies, show that for an identical cumulative dose, the noise in subtraction (stimulus minus baseline) images can be reduced by a factor of 2 using three-dimensional data acquisition, with appropriate fractionation of the dose. This improvement is dependent on axial position due to the sensitivity characteristics of three-dimensional scans; however, there is a significant gain in the signal-to-noise ratio (S/N) in all image planes. Studies performed in a human subject demonstrate how the factor of 2 gain in S/N leads to improved detectability of activation sites in PET subtraction images.