European Nuclear Medicine Guide
Nuclear Medicine (NM) images display functional information of various organs and tissues, because each pixel value reflects the uptake of an administered radiopharmaceutical in a location in the patient that matches the location of the pixel in the image. Sometimes, the location and distribution can be difficult to determine, even if the uptake of activity is high, simply because there are no other uptakes aiding as landmarks for the uptake of interest. Thus, the ability to have complementing anatomical information increases the diagnostic value of the NM study. The advantage of multimodality imaging is to be able to acquire both modalities at the same time to create tomographic images that match each other both in size and in location. By using graphics, these images can be shown superimposed using different colour scales. The common method for anatomical information is Computed Tomography (CT)[130].
In the early days, the two studies were acquired separately, and this resulted in two sets of images that were not registered to each other. Different types of mathematical registration methods were suggested, but since the performance of these methods are limited, internal mismatches were expected, even if the patients were positioned on the bed as carefully as possible. Because the patients moved to different places for the two studies, the organs were also expected to misalign internally, even if the external positioning of the patient was accurate. Furthermore, transfer of data between SPECT and CT was, at that time not fully standardized and straight-forward. Hasegawa et al were one of the first researchers that designed a hybrid scintillation camera that combined SPECT with CT[131,132], and the first commercial product was released by GE Healthcare in 1999. It consisted of a low-dose one-slice CT which took about 10 minutes to acquire a full field-of-view study. The resolution was around 3.5 mm, but the possibility to perform dual acquisition aligned SPECT/CT images was a major improvement. Considering the poor resolution of a SPECT image (order 10-15 mm), the 3.5 mm spatial resolution of the low-dose HawkEyeTM CT still provided a good localization of the activity uptake. Today, SPECT/CTs are mostly equipped with high-resolution and fast diagnostic CT systems. The potential advantage is, of course, that these CT images can be used also for diagnostic purposes, and the investigation time for the patient is generally reduced, because these devices have the capability of performing two studies on the same occasion. This also is very important for the healthcare logistic.
Fusion – localization
The combination of SPECT and CT imaging has proven to be clinically useful and superior to only using SPECT by many publications, see for example[133–139]. It improves anatomic localization and adds confidence by increasing both sensitivity and specificity in diagnostic decisions resulting in an improved patient care. For example, in small regions of the spine it can be difficult to determine from a SPECT image only if an abnormality resides in the bone, or if it is located in the adjacent joints. A high-resolution CT image provides such additional information[140]. Figure 1 is an example of this improvement and shows anatomical and functional information in a fused image.
Figure 1. The images show the advantage of a fusion image (right) as compared to a planar posterior image (left) and a SPECT image (middle).
Attenuation and scatter correction
CT information can be also used for correction of attenuation in the patient and scatter in the image. These are, thus, both related to photon interactions in the patients that either absorb the photon, and thereby reducing the acquired counts, or changes the direction of the photon, with a potential mis-positioning of the event in the image. The effects from scatter and attenuation are strongly dependent on the geometry and composition of the patient, a dependence that also varies with the projection angles. These effects are usually treated separately, although they are linked to each other.
As a simple interpretation, one can understand attenuation as the reduction of the expected number of detected photons due to absorption in the patient or scattering in a new direction that makes detection impossible. To correct for this, one needs to calculate the probability for a photon to be transmitted through the patient. For this, we can use the CT images. By calculating the length of the expected line and taking into account the variation in attenuation along the line from the CT information, we can calculate a correction factor that, for each projection view, very accurately corrects the registered events. The principle is indicated in Figure 2, where the correcting factor is the reciprocal of the probability for photon #1 to travel the distance x1, where the average attenuation factor is μ1.
Scatter in the image is the result of detected photons that never should have been measured but appear in the image because of the poor energy resolution of the scintillation camera. Scatter does not produce any artefact in the images, but the locations of the scatter events are wrong. This leads to a loss of image contrast and limits the ability to obtain quantitative images. Therefore, correction should also be made for scatter. Most commercial SPECT reconstruction programs base their scatter correction methods on the use of information acquired in additional energy windows. However, since the data obtained from scatter windows generally do not represent the scatter in the primary energy window regarding amount and spatial distribution, an alternative is to model the scatter by analytical methods or from Monte-Carlo-based simulations[141].
Figure 2. Examples of two pathways for a primary photon (#1) and a scattered photon (#2) and the probabilities of escape. These probabilities are used to compensate for attenuation and scatter by including them in the modelling of attenuation and scatter in the reconstruction program.
Both methods use calculated probabilities for a scattered photon to be detected. In both cases, detailed information about the patient’s size and internal tissue distribution are essential, since photon interaction very much depends on the body composition and shape. Figure 2 shows the principles for how scatter can be modelled from CT images by calculating the probability for a photon to 1) travel the distance x2, to be 2) scattered an angle θ toward the camera, followed by 3) an escape to travel a distance x3.
Segmentation (VOI)
The spatial resolution of a SPECT image is generally much lower than a CT image by an order of a magnitude. Furthermore, image contrast is also low in areas of low uptake, which then leads to difficulties in the segmentation (i.e., the delineation either by hand or by computer algorithms) of region-of-interest (ROI) or volume-of-interest (VOI). However, if a set of CT images is properly registered to a set of SPECT images, the CT images can be used to define volume-of-interest (VOI) that can then be transferred for use on the SPECT images. For example, the geometric volumes of organs and their masses is an essential parameter when calculating the absorbed dose in radionuclide dosimetry. However, it should be recalled that spill-out (or spill-in) of counts over the outline that defines a geometric VOI will occur due to the spatial resolution of a SPECT image. Therefore, when applying a CT-based VOI on a set of SPECT images, this loss of counts due to spill-out needs to be accounted for in order to calculate the correct count density or the activity concentration if translated by a calibration factor. A method to overcome this count loss is to apply so-called recovery factors. These are often determined from experimental measurements or computer simulations of geometries with known activities and volumes and define the ratio between the total measured counts to the counts determined within a geometrically defined VOI of the source. The main limitation is that the VOI of tumours generally have arbitrary shapes, and therefore a recovery correction by a factor, obtained from a spherical source, may not completely account for the spill-out/spill-in.
Dosimetry calculations
An important field of nuclear medicine is the use of radionuclides emitting charge-particles for treatment of malignant systemic diseases. The principle is administering radionuclides or radiopharmaceuticals that accumulate close to or in the target volume, e.g. cancer cells. Damages to nearby cancer cells will occur by the interactions between the emitted charge-particles and the orbital electrons in the molecules that will cause ionizations and chemical breakdowns. Since these radiopharmaceuticals will not completely target just the main volume of interest, the radiation may also damage normal and healthy tissue. Therefore, it is important, as in external radiation therapy, to perform dosimetry to optimize the treatment with respect to the outcome and side-effects. This can be made by quantitative sets of SPECT images that display the radionuclide distribution over time, and from which mean organ absorbed doses can be calculated. In many standard treatments today, a simplified and approximate calculation scheme is used, but having access to a Monte-Carlo-based radiation transport program and a good description of the body composition (available from registered SPECT/CT study) patient-specific absorbed doses can be calculated[142,143].
As stated above, dosimetry calculations require, in principle, a procedure based on sequential SPECT/CT studies in order to determine the total absorbed dose to an organ or tumour. Then, the problem of mis-registration, due to patient movements between the time-points, appear even if each time point is acquired with a hybrid SPECT/CT study. The only way to compensate for this mis-match is to rely on mathematical registration methods, but using the SPECT data for this purpose may not be reliable due to image noise and potentially large redistributions of the radiopharmaceuticals between the time points. However, if a SPECT/CT is conducted at each time point, then a non-rigid transformation can be determined by a CT-CT image registration, and from the result of this the transformation parameters can be applied to the SPECT-SPECT images[144]. The reason why CT-CT registration is preferable, is because the CT-CT images are more similar as compared to the matching SPECT-SPECT images.
In some dosimetry applications, the calculation of the absorbed doses still relies on quantitative planar scintillation camera imaging[145]. These images are also affected by photon attenuation. The most common method to correct for attenuation is the conjugate-view method. In order to correct for photon attenuation when using this method, an image of attenuation factors is used, where each factor corrects a corresponding pixel value. If a SPECT/CT system is available, such an attenuation map can be calculated from a planar scout measurement using the x-ray unit on the SPECT/CT system[146,147]. Figure 3 shows an example of such an image. The value of a pixel in the middle image represents the probability for absorption along a line at the pixel position.
An extensive collection of SPECT/CT artefacts can be found in ref xx (SPECT/CT Atlas of quality control and image artefacts. IAEA Human Health Series no. 36, IAEA, Vienna 2019)
https://www.iaea.org/publications/13407/spect/ct-atlas-of-quality-control-and-image-artefacts (accessed 2020 05 07)
Misalignments between SPECT and CT
Although SPECT and CT studies are performed on the same occasion, the studies are sequential, and therefore mis-registrations can occur due to breathing, patient movements, scanner bed bending, or differences in acquisition protocols between SPECT and CT[148].