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Computer Acquisition of Nuclear Medicine Images

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Computer Acquisition of Nuclear Medicine Images CONTINUING EDUCATION Computer Acquisition of Nuclear Medicine Images Mark T. Madsen Department of Radiology, University of Iowa Hospitals and Clinics, Iowa City, Iowa ther 8 bits (1 byte) or 16 bits (2 bytes) to each pixel. A 1-byte Objective: The purpose of this article is to provide a com­ pixel requires half the storage space of a 2-byte pixel but has prehensive description of computer acquisition modes in nu­ a limited dynamic range. The maximum number of counts 8 clear medicine. that can be stored in a 1-byte pixel is 255 (2 - 1). In many Methods: The paper discusses each acquisition mode in instances, the likelihood of acquiring more than 255 counts is detail, explaining when use of each mode is justified. small, especially if the physical dimensions of the pixel are Results: The effects of each acquisition mode on the result­ small (i.e., either a large matrix or a high zoom factor). ing images are discussed. However, there are circumstances where 255 counts per Conclusion: Knowledge of acquisition modes and compe­ tence in acquiring digital images is a vital skill for nuclear pixel is too limiting. The maximum number of counts which 16 medicine technologists. can be stored in a 2-byte pixel is 65,535 (2 - 1). Key Words: Computers, nuclear medicine studies, digital images. IMAGE FORMATION J Nucl Med Techno/1994; 22:3-11 When a gamma ray is absorbed by the sodium iodide crystal, a visible light flash known as a scintillation is cre­ ated. This scintillation is sampled by an array of photomul­ This is the first article in a four-part series on computers in tiplier tubes that create a set of electronic pulses: an x and y nuclear medicine. Upon completion, the technologist should position signal that corresponds to the location of the scin­ be able to list the types of acquisition modes for nuclear tillation on the crystal, and an energy signal whose pulse medicine studies, choose the best acquisition mode for any height is proportional to the energy absorbed in the interac­ given study, and perform studies using these acquisition tion. If the energy signal falls within a selected energy win­ modes. dow, the x andy signals are digitized by a pair of analog-to­ digital convertors. This position information can be stored in Most nuclear medicine imaging systems present their infor­ two ways: (1) as a sequential list of position entries referred mation as digital images. A digital image is stored in the to as list mode, or (2) by incrementing a matrix element computer as an array or matrix of count values and is dis­ (pixel) corresponding to the location of the position, known played by assigning a gray or color scale that depends on the as frame or matrix mode (see Fig. 1). number of counts in each element. Typically (although not Images in both frame mode and list mode can be acquired exclusively), the arrays are square matrices that have dimen­ in a magnification or zoom mode. In zoom mode, the sam­ sions that range from 32 x 32 up to 1024 x 1024, although pling is increased by digitizing over a smaller range of the most nuclear medicine images have dimensions of either position signals. This increases the sampling frequency (de­ 64 X 64, 128 X 128 or 256 X 256 (1,2). creases the pixel size) and also decreases the field of view of Each matrix element (commonly referred to as a pixel) is the computer image. Typical zoom factors range from 1 to 4 a location in computer memory. A 64 x 64 matrix has 4096 in steps of about 0.25. pixels, while a 128 x 128 matrix is four times larger (16,384 Frame Mode pixels), and a 256 x 256 matrix is 16 times larger (65,536 Frame mode is the most common mode of image acquisi­ pixels). The number of counts which can be stored in a pixel tion for nuclear medicine studies. Static, dynamic, gated, depends on how many bits are allocated. Because of the way whole-body, and single-photon emission computed tomogra­ computers are designed, it is most convenient to assign ei- phy (SPECf) studies are acquired in frame mode. With frame mode, a matrix of computer locations is cleared in For reprints contact: Mark T. Madsen, PhD, Dept. of Radiology, memory prior to the start of acquisition. For each detected University of Iowa Hospitals and Clinics, 200 Hawkins Drive, Iowa City, lA 52242. event, the appropriate matrix element is incremented. This VOLUME 22, NUMBER 1, MARCH 1994 3 FIG. 1. Computer acquisition modes. Digi­ tized signals from the gamma camera can be stored in list mode or can point to a pixel which is incremented, frame mode. continues until a preselected time interval or total count topeaks is added together. This is how multi-energy gamma value is reached. The memory or storage space required for ray emitters such as indium-111 (mIn) or gallium-67 (67Ga) a frame mode acquisition is determined solely by the matrix are frequently recorded; the detected counts from all the size and the number of frames acquired. energy windows are combined into a single image. In the dual isotope mode, a separate frame (or list) is reserved for List Mode each selected photopeak (Fig. 2). This yields two distinct List mode acquisition is less commonly used and some images that are spatially registered with one another, each of nuclear medicine computer systems no longer offer it as an which has unique information. option. Since information collected in list mode is a series of One application of dual isotope imaging is technetium-99m 99 111 x andy locations, it cannot be viewed directly. It must be ( mTc) MDP bone scanning in conjunction with In white reconstructed into image matrices in the manner described cell labeling for localizing sites of infection. Dual isotope above. The computer goes through the list of locations and image acquisition is currently available for all types of nu­ increments a matrix element corresponding to that set of clear medicine imaging including SPECT and whole-body coordinates. The advantage of list mode studies is that the studies. All nuclear medicine computer systems permit dual collected data can be framed in a variety of ways. One can isotope imaging and some can acquire image data from up to alter the matrix size and the frame rate to suit a wide range four separate energy windows. of purposes. Physiologic gate signals can also be included in the list, allowing the reconfiguration of image data based on SAMPLING this information. The big disadvantage of list mode studies is that they require large amounts of computer space since each As previously described, the surface area of the gamma event is recorded separately. camera crystal is mapped into a matrix of computer elements (pixels). The physical size of a pixel is found by dividing the Dual Isotope Imaging field of view by the matrix size, as shown in Figure 3. For Most gamma cameras allow the simultaneous selection of example, if a 380 x 380-mm field is coincident with a 128 x multiple photopeaks for gamma rays of different energies. 128 matrix, each pixel corresponds to a 3 x 3-mm area. If the This information can be stored in several ways. In the single matrix is changed to 64 x 64, the pixel size is increased to 6 isotope mode, the information from all of the selected pho- x 6 mm. If an acquisition zoom factor of 2 is used, the field 4 .JOURNAL OF NUCLEAR MI:DICINI: TI:CHNOLOGY FIG. 2. Dual isotope imaging. Two separate images are acquired simultaneously with per­ fect spatial registration for each energy win­ dow. (A) ~c-MDP bone image; (B) 111 ln white blood cell image. of view will be decreased to 200 x 200 mm and the pixel size system response or spread function. If we are concerned will now again be 3 x 3 mm. Since the spatial sampling of the about the smallest objects that can be seen on the image, image is determined by pixel size, increasing the acquisition then the limiting factor is related to the full-width-at-half­ zoom is one way to achieve sufficient sampling without using maximum (FWHM) of the spread function. In order to crit­ the additional storage space required for a larger matrix. ically sample at this limit, the pixel size should be smaller However, this will only work for imaging smaller organs than FWHM/3 (3). With a high resolution collimator, the which can fit inside the reduced field of view. FWHM for a source at 10 em is about 9 mm. This implies By its very nature, a digital image is made up of discrete that the pixel size should be no larger than 3 mm. If the samples. The sampling interval is determined by the pixel gamma camera field of view is 380 mm, this would require a size, which, as discussed above, depends on the matrix size 128 x 128 matrix (3 mm pixel) or a 64 x 64 matrix with a and the image zoom factor. What should the pixel size be? zoom factor of at least 2. The factors that determine the answer to that question are If the pixel size is coarser than this, the spatial resolution the following. of the system is compromised and information is lost. This is illustrated in Figure 4, which shows images of a resolution 1. The spatial resolution of the imaging system. phantom collected at several different pixel sizes with the 2.
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