|
|
| Fig. 1: Siemens SPECT scanner that has two gamma cameras. (Source: Wikimedia Commons ) |
Radionuclide imaging techniques like single-photon emission computed tomography (SPECT) and positron emission computed tomography utilize radioisotope decay to visualize specific internal processes and particular tissue types by localizing the site of each decay event using tomographic reconstruction. SPECT is a radionuclide imaging method that measures &gamkma;-ray emissions from radioisotopes within a patient. In contrast, PET measures the annihilation photons generated when a positron emitted by the decay of a radioisotope interacts with an atomic electron in an annihilation event. [1]
PET has superior spatial resolution because two collinear photons are generated with each annihilation event whereas in SPECT, only a single γ-ray photon is generated for each nuclear de-excitation event. When localizing the site of each positron decay event in PET, electronic collimation is applied where a tomographic line of response is drawn between the detector element pairs. Because only a single γ-ray photon is generated for each nuclear de-excitation event, the photons have to be physically collimated. However there is a large disparity in cost: SPECT gamma cameras tends to cost far less than PET scanners. [2] Furthermore, SPECT is capable of multi-radioisotope resolution because the energy of the γ-ray photon generated in the de-excitation event depends on the radioisotope used. In PET, the annihilation photons generated by the interactions between positrons and electrons is always 511 keV so it cannot distinguish signal from multiple radiotracer species. [3] Consequently, SPECT is still a widely used imaging modality even with all of PET's advantages.
A radiotracer containing a γ-ray emitting radioisotope such as Tc-99, I-123, and In-111 is injected into a patient's bloodstream. A nuclear de-excitation event occurs where the short-lived excited nucleus decays into a more stable form and releases a γ-ray photon. These individual photons traverse the patient's tissues and interact with the electrons and nuclei of atoms in the tissue through Compton scattering or photoelectric absorption. Compton scattering is the inelastic scattering of the photon where it is deflected from its original direction and transfers some of its energy to a scattered electron. Photoelectric absorption occurs when an atom absorbs the γ-ray photon, and the energy results in the emission of a photoelectron from one of the inner atomic shells. [1] Both effects lead to attenuation of signal and require corrections in order to maintain image resolution.
Once the γ-ray photon exits the patient's body, it interacts with a collimator. In order to determine the location of the original de-excitation event, the photon's direction of incidence must be at specific fixed angles. Collimators consist of a base of high density and high atomic number elements like lead or tungsten with a set arrangement of perforations for photons to enter the detector. Photons that interact with the collimator at the wrong angle are absorbed by the collimator and do not reach the detector. Photons that succeed in traversing the collimator are those that have the appropriate angle of incidence for optimum image resolution. [1]
The detector consists of either a scintillation crystal coupled with photomultiplier tubes (PMT) or semiconductors. A scintillation crystal is a compound that emits visible light when it absorbs a high-energy photon like an X-ray or γ-ray. This visible light then interacts with a PMT, which generates an amplified electronic signal from the absorbed visible light through the photoelectric effect. One drawback to this detection system is that some of the visible light is lost at the interface of the scintillation crystal and the PMT, limiting the energy resolution of the detection system. In contrast, in a semiconductor detector a γ-ray is absorbed, and the resulting energy frees electrons within the charge-free depletion zone of the semiconductor. This generates an induced charge at the terminals, which can then produce electronic pulses proportional to the energy of the incident γ-ray. Without a conversion stage, the fidelity of the resulting electronic signal is superior in the semiconductor system than in the scintillation crystal and PMT apparatus. [1]
Although perfecting the detector or collimator systems can further enhance SPECT, significant improvements lie in the development of better radionuclides. Currently, there are three predominant radionuclides used in clinical SPECT imaging: Tc-99m, I-123, and In-111.
|
|||||||||||||||
| Table 1: Decay properties of the most common radionuclides used in SPECT and a proposed new clinical radionuclide, Tb-155. [4] |
One clinical application of SPECT is calculating the pre-therapeutic dosage of antibodies and peptides that utilize therapeutic radiometals such as Y-90 and Lu-177 in a particular patient. A radioisotope whose properties closely match the therapeutic radiometal is substituted into the antibody or peptide, and SPECT is used to measure the subsequent tissue distribution. Currently, In-111 is the major radionuclide used in these dosage and tissue distribution studies because its half-life and emission profile, outlined in Table 1, are most conducive for longer term studies of a drug's tissue distribution. However, the main issue with In-111 usage is that its coordination chemistry is slightly different than the radiolanthanides typically coordinated in these antibodies and peptides, suggesting that the tissue distribution and chemical properties measured for these compounds using an In-111 radiotracer are not completely accurate. One advancement in this field is that Muller et al. have developed a methodology that utilizes Tb-155 rather than In-111. [4] Tb-155 is the only radiolanthanide that emits γ-rays appropriate for SPECT without the emission of positrons and stronger, more harmful γ-rays. As a congener for the radiometals like Y-90 and Lu-177 typically used in therapies, Tb-155 more accurately reflects the tissue distribution and pharmacokinetics of these radiometal antibodies and peptides.
Despite the lower spatial resolution afforded by SPECT as compared to PET, SPECT remains a viable and widespread imaging modality because of its lower cost and larger suite of radioisotopes with potential implementation. While optimizing the SPECT device itself can improve its imaging resolution, further development of radioisotopes and ligands capable of more effectively targeting biological markers and producing stronger signal with a larger half-life are also possible avenues towards improvement of SPECT.
© Jeremy Uang. The author grants permission to copy, distribute and display this work in unaltered form, with attribution to the author, for noncommercial purposes only. All other rights, including commercial rights, are reserved to the author.
[1] S. S. Gambhir and S. S. Yaghoubi, Molecular Imaging with Reporter Genes (Cambridge University Press, 2010), p. 120.
[2] E. E. van der Wall, "Cost Analysis Favours SPECT Over PET and CTA For Evaluation of Coronary Artery Disease: the SPARC Study," Neth. Heart J. 22, 257 (2014).
[3] B. Xu et al., "Utilizing the Multiradionuclide Resolving Power of SPECT and Dual Radiolabeled Single Molecules to Assess Treatment Response of Tumors," Mol. Imaging Biol. 17, 671 (2015).
[4] C. Muller et al., "Future Prospects for SPECT Imaging Using the Radiolanthanide Terbium-155 - Production and Preclinical Evaluation in Tumor-Bearing Mice," Nucl. Med. Biol. 41, e58 (2014).
