Nuclear Medicine In nuclear medicine the region under study becomes an active source of radiation. The imaging problem is then to determine the three dimensional source. Radioactive materials are administered selectively to create the source. The source of radioactivity must be less intense than in radiography using x-rays. This is because in radiography the patient is only exposed during the time the external source is active, which is a fraction of a second in the case of an x-ray shadow graph to a few seconds in the case of a CT scan. In nuclear medicine the patient is irradiated from the time the dose of radioactive material is administered until it is eliminated from the body. A second reason for keeping the source at a low intensity is due to the collimation necessary and the type of detectors used. While it is the source that is being imaged and not the attenuation coefficients of intervening tissues that are being measured, it would seem that a high energy source would be ideal. However it is then difficult to collimate the beam and to detect it as the photons pass straight through the collimator and the detector. Photons must be absorbed by the detector to be detected. Radioactive materials used in nuclear medicine are gamma ray emitters. I131 was the earliest material used as it was selectively absorbed by the thyroid. Tc99, an isotope of technicium, is now the most widely used source. It can be made easily without requiring a cyclotron. It has a short half life of six hours, so that it does not irradiate the patient for a long period. Its gamma ray emission energy of 140 kev is a good compromise between body attenuation and source imaging. It can be attached to a sufficiently wide variety of compounds that can be selectively taken up by different body organs. The most commonly used radionuclides in nuclear medicine are shown in Table 1.

TABLE 1 Radionuclides Commonly Used in Nuclear Medicine

Radionuclide Half-life Transition Eγ (MeV) Production Carbon-11 20.38 min β+ 0.511 Cyclotron Fluorine-18 109.77 min β+ 0.511 Cyclotron Phosphorus-32 14.29 days β- 0.695 Reactor Chromium-51 27.70 days EC 0.320 Reactor Cobalt-57 270.90 days EC - Cyclotron Gallium-67 78.26 hr EC - Cyclotron Molybdenum-99 66.00 hr β+ - Reactor Technetium-99m 6.02 hr IT 0.141 Generator Indium-111 2.83 days EC - Cyclotron Indium-113 100.00 min IT 0.393 Cyclotron Iodine-123 13.20 hr EC - Cyclotron Iodine-125 60.14 days EC 0.027 Reactor Iodine-131 8.04 days β- 0.364 Reactor Xenon-133 5.24 days β- 0.081 Reactor Thallium-201 3.04 days EC - Cyclotron

Scanning Detectors Earliest detectors consisted of a single crystal of a scintillating material, like sodium iodide, at one end of a long narrow diameter hole drilled in a lead block (Figure 228). The lead block collimated the gamma ray beam. Behind the detecting crystal was a photomultiplier tube which picked up the light generated in the crystal by bombardment with the gamma rays. The detector was scanned over the area of interest to produce an image.

Figure 228 Scanning gamma ray detector.

Depending on the organ being imaged, different degrees of gamma ray intensity have different implications. In the thyroid gland in the neck, areas of high radioactivity also are areas of abnormally high tissue activity and vice versa. Tumours are usually inactive and so form "cold nodules" in the image. Brain tumours tend to take up high levels of radioactive material. Liver will take up material, but tumours inside the liver will not take up radioactive substances. Scanning with a single detector is slow. The image is affected by patient movement and it is not possible to do any dynamic studies where the change in function of an organ is measured with time. Gamma cameras have been developed which can view an entire region simultaneously. The speed with which an image can be formed is governed by the efficiency of the camera and the maximum radioactive dose that can be safely given to the patient. The dose of radiation is the Curie (Ci), which is defined as 3.7 x 1010 disintegrations per second. In most procedures a dose of 5 Ci is given. This is usually much greater than the dose given by an x-ray radiographic procedure. Gamma Ray Cameras Gamma ray cameras allow all points of a two dimensional image to be formed simultaneously without requiring a mechanical scan. They are the most commonly used instruments in nuclear medicine. A basic camera is shown in Figure 229. The camera consists of a collimator of some type to convert the essentially random ray distribution into a two dimensional image. The collimator has a similar function to the lens on a light camera. Behind the collimator is a detector for detecting the position of each gamma ray photon. The output of the detector goes to a recorder. The detector material, as in radiography, needs to be sufficiently thick to stop high energy photons. Typically 10 to 12 mm thick sodium iodide crystal are used. The recorder is often

an array of photomultipliers. More modern machines use diode arrays as recorders or specially doped arrays of diodes which act as a combined detector and recorder. The output of the diode arrays or photomultipliers are digitised and displayed or stored.

Figure 229 Basic gamma camera.

One problem is that many of the gamma ray photons do not come directly from the organ to be imaged but are subject to Compton scattering en route. These lead to distortions in the image. They have lower energy than the direct photons so can be removed by thresholding the output of the recorders. A pulse height analyser, which forms a histogram of the intensities of the photons, can be used to set the threshold level. Anger Camera This camera, named after its inventor, allows a higher resolution than the spacing between detector elements, as applies with the ordinary gamma camera. Light from a gamma ray photon is emitted in many directions and can strike a number of recorders simultaneously (Figure 230). By interpolating between all the photodetectors that are recording simultaneously according to their output amplitudes, the position of the original gamma photon in the detecting crystal can be found. Thus in a system consisting of two photodetectors, the position of an event (gamma ray photon strike) x is: x = I1 x1 + I2 x2 (81) I 1 + I2 If two or more gamma rays strike the detector crystal simultaneously the event is misinterpreted. Fortunately this is a rare occurrence. Resolution is typically 2 mm for an array size of 400 mm (ie about 40,000 pixels) using 25 photomultipliers.

Figure 230 Anger camera principle.

A typical system is shown in Figure 231. A single detecting crystal is used. If the output pulses from the photomultipliers do not fall within the energy range for a direct gamma ray photon for the particular isotope used, the screen is not illuminated at that point. Centroid calculations are used to determine the x and y positions by a generalisation of equation (81).

Figure 231 Components of an Anger camera.

Collimation The simplest collimator is a simple pinhole aperture which works in the same way as a light pinhole camera (Figure 232). The image is magnified by M: M = id (810) z This has the same problems as the equivalent light camera. The intensity of photons available to form the image is small. The magnification varies inversely as the distance of the object from the camera. If the organ imaged is a known distance away there is no problem. If it is at an unknown depth it is difficult to determine the size of lesions. A pinhole system and thyroid image obtained is shown in Figure 233.

Figure 233 Thyroid image and pinhole collimator scanner.

The parallel hole collimator shown in Figure 234 is placed as closely as possible to the organ to be imaged. Each hole in the array allows the portion of the source near it to illuminate the corresponding part of the image detector. This produces a 1:1 source to image magnification regardless of the distance of the source from the camera. There is some distortion of the image due to photons travelling obliquely through the holes. Although the signal to noise ratio is higher than with the pinhole collimator, it is still too low in many applications.

Figure 234 parallel hole collimator imaging system.

Coded aperture systems are used to overcome the low signal to noise of the pinhole and parallel hole collimators. An array of holes, known as a coded aperture plate, allows a greater number of photons to reach the detectors (Figure 235). The resultant image is the convolution of the source with the aperture plate. The original source can be reconstructed by a suitable cross correlation.

Figure 235 Imaging with a coded aperture.

An example of a scanner and the output of a whole body scan is shown in Figure 236.

Figure 236 Commercial scanner and typical whole body image. Positron Emission Tomography (PET) Scanners Some radioactive substances emit positrons which immediately interact with electrons, annihilating both particles and generating two gamma rays, each having energies of 510 kev and travelling in almost opposite directions. A simple imaging system can be constructed as shown in Figure 237. If the image depth z is known, the x and y coordinates (at right angles to the plane of the paper and also in the plane of the paper at right angles to the z direction) can be calculated: x = x1 z2 + x2 z1 z1 + z2 z1 + z2 y = y1 z2 + y2 z1 z1 + z2 z1 + z2 The position indicating detectors only record rays if both are picked up simultaneously. This avoids errors due to external radiation or Compton scattering. The detectors can be set to selectively respond to photons of around 510 kev. The disadvantage of this approach is that if there is radioactive material in other xy planes (at different z values) there will be blurring of the image.

Figure 237 Positron emission imaging system.

An alternate approach is similar to projection measurements made in x-ray CT scans. A ring of detectors surrounds the patient as shown in Figure 238. Only coincident events are recorded (Figure 239). If the number of coincident events reaching a pair of detector locations x1, y1 and x2, y2 is summed, the result is the line integral of all sources along the line joining the two detectors. When the line integrals are found over all angles and all positions, the source distribution can be reconstructed in the same way as in x-ray CT scans. There are two limitations to positron emission tomography. The positron can move a significant distance from its point of creation to its point of annihilation. The two photons may not be 1800 apart, depending on the momentum of the positron. Some attempt to improve the resolution of the image is being made by measuring the small difference in times of "coincident"

events. As in ultrasonic imaging, this can be used to measure the depth at which the annihilation took place. Due to the pico second accuracy needed in the timing measurements only a window can be placed on the annihilation position. Nevertheless some improvement in image resolution can be gained. Resolution is approximately 1 cm. A practical difficulty is that many of the required isotopes require an on-site cyclotron to produce them. Radionuclides Although technetium is the most commonly used radionuclide, a number of others are used in many common applications (Table 2). The choice of a radionuclide for medical image depends on four main factors:

1. Emission of gamma rays or positrons 2. Chemical compounds must be biologically convenient for administration 3. Chemical compounds must be physiologically selective and relevant 4. Radiation dose should be as short as possible, ie short half-life and/or rapid excretion.

Although short half-life reduces the radiation dose, it can be inconvenient in that it does not allow enough time to transport the radionuclide from a reactor to the site where it is being used. To overcome this problem some specialist hospitals have an on-site cyclotron to produce the radionuclides. Cyclotrons are smaller, cheaper and safer than a reactor.

TABLE 2 Frequently Used Procedures in Nuclear Medicine

Procedure Radionuclide Chemical form Plasma/blood volume 125I, 131I Iodinated serum albumin Red cell mass & life 51Cr Haemoglobin Vitamin B12 absorption 57Co Cyanocobalamin (B12) Thyroid function 125I, 131I Sodium iodide Tumour & abscess 67Ga Citrated complex Thrombus 111In Oxine labelled platelets Cerebro-spinal fluid 111In DTPA chelated complex Myocardium 201Tl Thallous chloride Cardiac blood pool 99Tc Blood Brain 99Tc Pertechnetate Kidney 99Tc DTPA chelated complex Bone 99Tc Methylene diphosphonate Liver/spleen 99Tc Colloid Liver/gall bladder 99Tc Iminodiacetic acid Lung perfusion 99Tc, 18O Albumin, gas Lung ventilation 133Xe Gas Thyroid therapy 131I Sodium iodide Polycythaemia rubra vera tmt 32P Sodium phosphate

Figure 238 Positron ring detector.

Figure 239 Coincidence counter in a positron imaging system.