Radiopharmaceuticals are radioactive compounds used for both diagnostic and therapeutic purposes. Radiopharmaceuticals generally mimic naturally occurring chemicals with specific biological functions. Tissues with unique properties, such as a high metabolic rate, accumulate the radionuclide at far higher rates than tissues without these properties. This feature allows for imaging and treatment based on biological activity.
Uses of Radiopharmaceuticals
Radiopharmaceuticals are classified as being for either therapeutic or diagnostic use.
Therapeutic radiopharmaceuticals are intended to deliver a sufficiently large dose of the radiopharmaceutical to the target tissue, while minimizing the dose to non-target tissues. This requirement limits the choice of available nuclides to those that decay via short range emissions, such as beta particles and alpha particles.
Key Point: All currently approved therapeutic radiopharmaceuticals are beta or alpha emitters whose short ranged radiation reduces dose to non-target tissue and the general public.
Diagnostic radiopharmaceuticals are used to determine the presence and position of biological processes or features. A good diagnostic radiopharmaceutical should have the following properties:
- Low radiation dose
- Radiation should be of a minimal dose while producing photons with a high enough energy to readily escape the patient’s body and be detected.
- High radiation detection efficiency
- Radiation escaping the patient should be of low enough energy that it is readily detectable.
- High specificity
- The radiopharmaceutical should readily accumulate in the tissue of interest and should not accumulate elsewhere.
- The radiopharmaceutical should be chemically non-toxic and safe for patient administration.
- Cost effective
- The production of radiopharmaceuticals should be cost effective in clinically useful quantities.
Key point: Most nuclear imaging equipment, except for PET scanners, is optimized to collect photons of about 140 keV. This is because 140 keV has been found to be a good trade-off between minimizing patient dose and a high radiation detection efficiency. Lower energy photons would increase the radiation dose to the patient due to attenuation within the body, while higher energy photons are more difficult to detect.
Radiopharmaceutical Mechanisms of Localization
The utility of a radiopharmaceutical is driven by its ability to selectively accumulate within desired tissues. Therefore, significant attention is paid to the biochemical mechanisms which fulfill this role. Below are some common methods of localization.
Active transport involves the use of biological mechanisms that selectively uptake the radiopharmaceutical into a cell against a concentration gradient. That is, active transport requires cellular components which transport the drug from areas of low concentration to areas of high concentration.
I-123 is commonly used to assess the structure and function of the thyroid gland. Iodine is actively transported into the thyroid gland, where a process termed the organification of iodine allows for the incorporation of iodine into thyroid hormones.
F-18 FDG is a glucose analog that concentrates in highly active cells using glucose for energy. F-18 FDG is useful for imaging active areas of cancer growth. The FDG is actively transported across the cell membrane, along with glucose, to fuel the cell’s elevated activity levels.
Key Point: Active transport, the mechanism by which F-18 FDG functions, can be thought of as reverse osmosis by which concentrations are increased on one side of a barrier (cell wall) rather than equalized.
Compartment localization generally involves placing a radiopharmaceutical into a specific anatomic compartment and imaging for signs of leakage.
Xe-133 gas which is inhaled into the lungs.
Tc-99m labeled red blood cells which are introduced into the circulatory system.
Receptor binding radiopharmaceuticals adhere to specific binding sites located on cells of interest. Receptor binding holds promise in targeting specific abnormal cell types based on their receptors.
In-11-octreotide is used for the localization of neuroendocrine tumors because of its ability to bind to somatostatin receptor sites.
Physiochemical absorption involves the incorporation of radiopharmaceuticals into the structure of a target cell.
Ra-223, in the form of Ra-223 dichloride, mimics the properties of calcium allowing it to be incorporated into the mineral structure of bone. Ra-223 is an alpha emitter and is used for the treatment of prostate cancer which has invaded boney anatomy.
Radionuclides used in nuclear medicine have a short half-life which allows for faster treatments, faster imaging, and lower public exposure. As a result of the need for such short half-lives, most radiopharmaceuticals in use today must use artificially generated radionuclides. The vast majority of these nuclides are produced in cyclotrons, nuclear reactors, and radionuclide generators.
Nuclear Transformation Notation
Nuclear transformations are denoted T(P, E)R where T is the target material, P is the accelerated particle type, E is the number and type of emitted particles during transformation, and R is the resulting radionuclide.
For example, the notation for Gallium-67 production is 68Zn (p, 2n) 67Ga which indicates that a proton (p) was accelerated into 68Zn which produced 67Ga by the emission of two neutrons (2n).
Often, the desired radionuclide is produced after decay of a bombardment produced radionuclide.
For example, Iodine-123 is produced via the (p, 5n) transformation to Xe-123 which then undergoes electron capture (EC) to yield I-123.
Nuclear fission resultant of neutron absorption is indicated as in the below example. The creation of Mo-99 (a parent of Tc-99m) via thermal neutron bombardment and subsequent decay of U-235. Note that in addition to Mo-99 and Sn-134, three neutrons and additional gamma photons are produced with a net energy release of about 200 MeV.
A cyclotron is a charged particle accelerator consisting of an evacuated cylinder divided into two sections referred to as "Dees" (because they look like a capital D).The evacuated cylinder is placed between two large magnets producing a constant magnetic field. Charged particles are injected into the center of the cylinder and an oscillating electric field is applied to the Dees. The electric field causes the charged particles to accelerate and the constant magnetic field causes that acceleration to have a circular trajectory. The energy and velocity of the particles increase as they pass from one Dee to the other. As the particles accelerate, the frequency of the electric field oscillation must increase until a desired energy is reached.
Radionuclides are produced by colliding these high energy-charged particles into target nuclei. The accelerated particle type, its energy, and the target material all impact what nuclides will be produced. Cyclotron-produced nuclei will be created in an excited state and will decay either by particle (protons and neutrons) or electromagnetic (photon) emissions. Most cyclotron-produced radionuclides are neutron deficient and decay via either electron capture (EC) or positron emission (β+).
Key Point: Most Positron Emission Tomography (PET) radionuclides are produced in cyclotrons.
Nuclear reactors are commonly used to produce radionuclides via nuclear fission or nuclear activation.
Nuclear fission is the splitting of a heavy atom into multiple smaller nuclei. Nuclear fission is induced in a nuclear reactor by the absorption of thermal (low energy) neutrons. Medical radionuclides may be produced as a decay product of these nuclear fusions.
Neutrons produced in a nuclear reactor may also be used to bombard stable nuclei in a process known as nuclear activation. When the stable nucleus absorbs a neutron, it produces an isotope with a moderate half-life.
The most common neutron activation methods are the (n, γ) reaction, in which neutron capture is immediately followed by emission of a gamma photon, and the (n, p) or (n, α) reactions in which a proton or alpha particle is emitted immediately after neutron absorption.
Radionuclide generators produce a short-lived medical radionuclide by decay of a longer-lived radionuclide. The most common example of a clinically useful radionuclide generator is the Molybdenum-99/Technitium-99m generator often referred to as Moly Cows.
Moly Cows (Tc-99m Generators)
In a Moly Generator, a sample of Mo-99 (in the form of ammonium molybdenate) is loaded into a porous column containing aluminum oxide resin. The ammonium molybdenite becomes attached to the surface of the resin, greatly increasing its surface area. When the Mo-99 decays to Tc-99m, the Tc-99m is less tightly bound to the resin than its parent, Mo-99. This allows the Tc-99m to be washed from the resin and collected while leaving the Mo-99 behind. This washing process is referred to as “milking” and gives rise to the name Moly Cow (i.e. “milking the cow”).
Assessing the dose of radiopharmaceuticals to internal organs for un-encapsulated radiopharmaceutical sources is a challenging task because the radiopharmaceuticals move dynamically through the body over time. Two methodologies to assess internal dose have been put forth by the Society of Nuclear Medicine; the Medical Internal Radiation Dosimetry (MIRD) formalism and the Radiation Dose Assessment Resource (RADAR) formalism. Additionally, the International Commission on Radiological Protection (ICRP) report 53 and report 80 contain tabulated absorbed dose values for many radiopharmaceuticals.
Medical Internal Radiation Dosimetry (MIRD) Method
The MIRD method uses a simple model of the human body and considers organs to be source organs, which contain the radiopharmaceutical, and target organs, for which the absorbed dose is calculated. An organ can be both a source and a target.
MIRD - Step 1. Compute accumulated activity
Accumulated activity (As) depends on both administered activity (A) and on the fraction of that activity that is taken into the source organ (F). Accumulated activity at a given time is determined as:
- As is the accumulated source activity at time (t)
- A0 is the administered activity
- Fs is the fraction of pharmaceutical which is the fraction of radiation accumulated in the organ
- λe is the effective decay constant
Note that effective half-life and decay constant are defined as below where physical half-life is the half-life of the radionuclide and biological half-life is the time required for half of the radionuclide to be expelled from the body.
Step 2. Determine the S-factor
The S-factor is the mean dose per unit of activity and has units of Gy/Bq×s. Although the S-factor can be calculated, it is most often found in tables as a function of the radionuclide, source organ, and target organ. The MIRD phantom used in these computations are based on a greatly simplified model of a 70kg adult male.
Key Point: Tabulated S-factors are determined by Monte Carlo simulation for an assumed 70kg mean man phantom.
Step 3. Compute dose to the target organ
Dose to the target organ is determined by:
Step 4. Compute effective dose to the whole body
Effective dose may then be computed as:
- Wt is the tissue weighting factor.
Scintillation (Gamma) Cameras
Scintillation cameras, sometimes referred to as gamma cameras, are scintillation-type detectors which convert radiation to light and, ultimately, to an electric signal which may be read out digitally. The scintillator is coupled to a set of photomultiplier tubes (PMTs) which convert light into electrical signals. These electric signals are then digitally processed to produce an image.
Anger Scintillation Camera Design
The most common scintillation camera design was created by Hal Anger in the 1950s. The Anger camera features a single scintillation crystal coupled to multiple photomultiplier tubes which are able to encode the location of the radiation interaction with the crystal. Other designs, such as the multi-crystal design, are less commonly used.
The scintillation crystal emits visible or ultraviolet light after interacting with ionizing radiation. High detection efficiency requires a scintillator with the following properties:
- A high conversion factor (>10%) meaning that a large fraction of radiation energy deposited in the scintillator is converted to light.
- A short decay time which limits the delay between irradiation and scintillation.
- Material should be transparent to its own scintillation wavelength to prevent self-attenuation.
- The scintillation wavelength should be readily detectable by the readout device (PMT, photodiode, etc).
- A large attenuation coefficient for the radiation being measured.
NaI(Tl) Scintillation Crystals
Sodium iodine crystals doped with thallium are the most common scintillation crystals used in nuclear medicine.
Advantages of NaI(Tl) Scintillators
- High conversion efficiency (13%)
- High attenuation coefficient for photons in the range of most radiopharmaceuticals (70-365keV)
- Prompt photon emission with a decay constant of 250ns
- High light yield (38 photons/keV is absorbed)
Disadvantages of NaI(Tl) Scintillators
- The attenuation coefficient is too low to efficiently collect the 511 keV annihilation photons produced during PET imaging studies
- The crystal is hydroscopic meaning that it must be enclosed to prevent atmospheric water absorption
- Crystals are fragile and may shatter if struck or subjected to extreme temperature changes
- Susceptible to accumulated radiation damage including that from ultraviolet light (fluorescent lamps and sunlight)
Photomultiplier tubes (PMTs) are used to convert scintillation light into electrical signals. The basic PMT design consists of an evacuated tube containing a electrodes referred to as dynodes, and an anode. The photocathode emits a small number of electrons (~1 electron/5eV) when struck by radiation. A voltage difference of approximately 100 V is used to accelerate these electrons into each dynode stage meaning that a total voltage of over 1,000 V is needed across the PMT. Each electron incident on a dynode causes the release of about 5 additional electrons, meaning that the net effect for a 10 dynode PMT is an amplification factor of 510. These additional electrons are then collected by the anode.
Gamma camera collimators are devices used to restrict the angle of incidence of radiation upon a portion of the detector. Collimators must attenuate photons in the kV range, so they are typically constructed of materials with a high atomic number, such as lead. All collimators consist of an aperture, through which radiation enters, and lead walls referred to as septa.
Image Formation Process
1. Radiopharmaceuticals are administered to the patient and differentially accumulate based on biological activity.
2. Gamma photons are generated by radioactive decay.
Radioactive decay may involve direct emission of a photon, as is the case for Tc-99m, or the photons may be generated by annihilation, as is the case for F-18 used in PET imaging. In the case of therapeutic radiopharmaceuticals, the decay is often via β- with photons emitted as the daughter nuclide returns to ground state.
3. Gamma photons exit the patient and are incident upon the collimator.
The septa of the collimator blocks photons that are not parallel to the hole of the collimator. This focuses the resulting image but also reduces the number of photons reaching the scintillator requiring a higher administered activity.
4. The remaining photons interact with the scintillation crystal, producing scintillation photons.
Conversion from keV energies to visible wavelengths allows for collection in the photomultiplier tube. Each keV of gamma photon energy creates approximately 38 photons (assuming it is a NaI scintillator), which are projected in radially from the site of interaction.
Four types of events may be detected in the scintillation crystal:
- Valid events - generated when an unscattered photon oriented parallel to the collimator septa produces scintillation.
- Septal penetration events - occur when a photon that is not oriented parallel to the collimator septa penetrates a septal wall and produces scintillation.
- Object scatter events - occur when photons are scattered within the body before passing the collimator and producing scintillation.
- Detector scatter events - occur when a photon passes the collimator, but is multiply scattered within the scintillation crystal. Detector scatter appears as multiple coincident low energy events.
5. Scintillation photons are converted to electronic signals by a photomultiplier tube.
The PMT converts scintillation photons into electronic signal and amplifies the signal by a factor of about 510. This step allows for the digital signal processing that follows.
6. An energy discrimination circuit rejects the scattered signals.
Energy discriminators, also referred to as pulse height analyzers, are electronic circuits used to reject scatter counts and coincidence interactions by rejecting counts that fall outside of a narrow energy window. Multiple counts are rejected because they produce too great a signal, while scattered counts are rejected because they create too small a signal. The energy discriminator reduces noise in the final image.
7. A position logic circuit determines the origin location of each signal on an x, y plane. Each localized signal is known as a count.
Many scintillation photons are created at each interaction in the scintillator. These photons spread radially from their source and produce signal in multiple PMTs. The precise location of a scintillation can be determined by interpolating the relative signal intensities and positions of the PMTs. The position of the scintillation between two PMTs can be determined as:
Clinical Gamma Cameras
Most clinical gamma cameras are either single or dual headed designs. In both cases, the gamma camera is mounted to a gantry which allows the gamma camera’s imaging angle to be adjusted.
In frame mode, counts are digitized into the appropriate image matrix bin immediately after detection. These counts are accumulated in the image matrix over a given amount of time, building up an image directly.
In list mode, individual x,y coordinates of each count are stored along with a time marker. This allows retrospective framing of the data after acquisition.
Frame rate is the length of time over which counts are accumulated in an image. Framing rate is impacted by the expected counting rate and the purpose of the study. High count rates and dynamic studies (such as cardiac exams) allow for faster frame rates. Cardiac studies may be as short as 0.2 to 0.5 seconds per frame.
Appearance: Bowing in or out of straight lines and/or regular repeating distortion lines in a flood field.
Cause: Imperfect mapping of scintillation location within the positional logic system.
Resolution: Perform linearity calibration, correcting for imperfect positional mapping.
Nuclear Tomographic Imaging
Both Single Photon Emission Computed Tomography (SPECT) and Positron Emission Tomography (PET) involve the creation of 3-dimensional images via computed tomographic reconstruction. The tomographic reconstruction of SPECT and PET images are very similar to CT reconstruction and typically use either filtered back projection or iterative reconstruction. Nuclear tomographic imaging differs from CT imaging in three key areas:
- Nuclear tomographic imaging is used to produce a map of biological activity identified by radiopharmaceutical markers. In contrast, CT imaging produces a map of the linear attenuation coefficient and the density of the body.
- Unlike CT imaging, which uses an external X-ray source, nuclear tomographic imaging relies on radionuclides within the body to generate its information carriers (gamma photons).
- In nuclear imaging, the photon source is within the body. As a result, photon attenuation depends strongly on the source’s path length to the detector. This requires an attenuation correction to prevent sources deep within the patient from appearing darker than superficial sources.
|Radiation Detectors||Planar Anger Camera||Full 360 degree scintillators|
|Scintillator||NaI(Tl)||LSO, GSO, LYSO|
|Most Common Nuclide||Tc-99m||F-18 (FDG)|
|Attenuation Correction||Patient thickness|
|Photon Vector Determination||Collimator||Coincident photon detection|
|Spatial Resolution||~10mm FWHM|
Dependent on orbit and collimator choice
Single Photon Emission Computed Tomography (SPECT)
A modern SPECT system consists of 1 to 3 gamma cameras with the dual image design being the most common. The gamma cameras are mounted on a complex gantry system allowing them to image from multiple angles around the patient.
SPECT Imaging Parameters
- Image array
- 64x64 or 128x128
- Parallel collimators are the most common
- Fan beam collimators (a cross between a converging collimator and a pinhole collimator) are used to provide magnification in brain imaging scans
- Data collection modes
- Step-and-shoot collection – data is collected at discrete angles (typically 64-128 positions)
- Continuous collection – data is collected continuously during imager rotation
- Circular – the simplest orbit with the lowest collection efficiency and signal-to-noise ratio
- Elliptical/Contoured – the imager path closely follows the patient’s body. This orbit requires a complex gantry, but provides the highest collection efficiency and signal-to-noise ratio
- Degrees of rotation
- 360° - most common
- 180° - commonly used in cardiac imaging to improve temporal resolution
- Typical imaging dose: 1-5 mSv
Photon attenuation must be determined to avoid a cupping artifact, in which deep seated objects appear darker than superficial objects. Three techniques are commonly used to correct for attenuation:
Thickness Based Estimates
Attenuation may simply be estimated and corrected for based on a standard model and a measure of the patient’s “thickness”.
Attenuation scans are generated using the SPECT/PET scanner’s imager, but using an external photon source. Attenuation scans are able to generate maps of the linear attenuation coefficient called attenuation maps which may be used to correct more accurately for attenuation.
CT imaging is performed in addition to nuclear tomographic imaging, as it provides both anatomical information and a direct map of attenuation.
Common SPECT Nuclides
The most commonly used nuclear imaging radionuclide. Uses include:
- Renal imaging and function
- Cardiac function
- Breast cancer imaging
- Cerebral perfusion
- Hepatic function
Gamma energy: 140.5 keV
Half-life: 6.007 hours
Production: Tc-99m's short half-life makes remote production and transportation impractical. Instead its parent nuclide, Mo-99, is produced and transported to hospitals as Tc-99m generators. These generators, colloquially known as Moly Cows, produce Tc-99m via beta decay with a half-life of 66 hours. Mo-99 is itself produced in nuclear reactors.
Positron Emission Tomography (PET)
Like SPECT and CT, positron emission tomography (PET) generates 3-dimensional images via tomographic reconstruction of photon projection data. PET imaging is unique in that the photons collected are generated by positron-electron annihilation processes which produce two 511 keV photons which project away from the annihilation site at approximately opposite angles. This unique photon arrangement (two photons traveling in opposite directions) has the unique ability of localizing the site of annihilation to a single point along a path by analyzing the location and timing of photon detection events. PET imaging has experienced great success owing largely to its ability to image metabolic activity via Fluorine -18 fluorodeoxyglucose (FGD).
PET systems are comprised of several rings of gantry-mounted detectors which surround the patient during imaging. The use of multiple detector rings allows for the acquisition of many (over 100) transverse slices simultaneously. The design also improves system sensitivity by being able to detect coincident photon events which are oriented slightly off-axis (nearly, but not quite orthogonal to the bore axis).
Common NaI(Tl) scintillators used in gamma cameras and SPECT imaging have a low detection efficiency for 511 keV photons generated in PET imaging. As a result, several other scintillators are used; including Lu2SiO2 (LSO), Gd2SiO4O (GSO), and LuxY2-xSiO4O (LYSO). These crystals have higher densities requiring smaller crystals and less self-attenuation of scintillation light. This means that these specialty crystals have higher collection efficiency and superior spatial resolution for PET applications.
PET scanners can identify multiple detection events occurring at the same time in different areas. These coincident events are assumed to result from a single positron annihilation emitting two photons in opposite directions. Therefore, the source can be assumed to lie on the line connecting the coincident detections. This means that PET scanners do no need collimators to determine the path of incident photons, which greatly increases their collection efficiency.
While coincidence detection removes the need for a physical collimator, it does introduce two new forms of noise:
- Scatter coincidence occurs when one of the annihilation photons is scattered prior to detection
- Random coincidence occurs when two photons arising from separate annihilations are detected simultaneously
Time of Flight
Annihilations photons travel at the speed of light from their point of origin to the detector. The time required for this journey, called the time of flight, is predominately dependent upon the distance traveled. This allows advanced PET scanners to determine the approximate point of origin along the photon’s path by computing the time between coincident events.
Key point: If time-of-flight calculations could perfectly determine the point of annihilation along the photons path, tomographic reconstruction would not be necessary to produce 3D images. Uncertainty in scintillation decay time, along with the speed of light, limits position determination to a few cm and necessitates tomographic reconstruction.
All modern PET scanners are coupled with a CT scanner in a PET/CT system. In a PET/CT system, the patient passes through both the PET and CT bores in a single scan. This CT image allows for precise correction of photon attenuation within the patient.
Limitations on Spatial Resolution
PET scanner spatial resolution is fundamentally limited to a few millimeters for several reasons:
- Positrons travel a small, but non-zero, distance between emission and annihilation. For F-18, this distance is approximately 2.4 mm. This means that annihilations take place outside of tissues which absorb PET radiopharmaceutical markers.
- The 511 keV annihilation photons do not travel in exactly opposing directions. This introduces uncertainty in the exact path that the annihilation photons traveled along.
- Time of flight calculations are dependent upon the time required for the scintillator to emit scintillation photons after an interaction. Although PET scintillation crystals are very prompt (quick to emit photons), scintillation is a decay phenomenon which has some fundamental uncertainties. This dependence limits the determination of the point of origin to within a few centimeters for LSO and LYSO detectors.
Common PET Nuclides
Uses: F-18 is used as the marker in fluorine-18 fluorodeoxyglucose (FDG). FDG is a glucose analogue which is absorbed in tissues with high rates of metabolic activity. This makes FDG ideal forimaging many forms of cancer, in differentiating viable tissue from scar tissue in the myocardium, and several other applications.
Half-life: 109.8 minutes
Emission Energy: Positron - 0.633MeV (approximate range of 2.4mm in tissue)
Production: Cyclotron bombardment of O-18 via 18O (p, n) 18F reaction
Typical dose: 7mSv for adult scan (370MBq administered)
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