The demonstration of visible light emission during the radioactive decay of positron-emitting radionuclides represents a new method of molecular imaging, which we have termed Cerenkov Luminescence Imaging (CLI). and demonstrate clearly the potential for in vivo CLI with widely available optical instrumentation as a molecular imaging modality.
demonstrate that (1) increased activity increases luminescence by increasing the number of events and (2) that higher energy positron-emitters produce more luminescence. We also confirmed that other 18F tracers—[18F]NaF and 3’-deoxy-3’-[18F]fluorothymidine (FLT)—emit light, showing that the observation is linked with the positron-emitter rather than the chemical nature of the tracer. Further, the role of the positron-emitter is supported by the observed luminescence signal decay as a function of time.
Another explanation considered for the source of the signal was bremsstrahlung, which is radiation emitted during the deceleration of high-energy, charged particles due to interactions with nuclei in the surrounding medium. There is a low probability of significant bremsstrahlung radiation production in low atomic number (Z) materials (such as water and tissue), and the bremsstrahlung would not change as a function of the refractive index of the material as we have observed in . Furthermore, the spectral distribution of the observed signal is inconsistent with bremsstrahlung radiation. That said, there are examples of bremsstrahlung imaging used with high energy beta emitters such as 32
P and 90
Y and detection resulting high energy photons. A major distinction is that bremsstrahlung imaging is not performed using CCD detection but rather using gamma detection equipment, such as in single photon emission computed tomography (SPECT) imaging (Clarke 1992
). It is worth noting that these high-energy β−
emitting radionuclides should also generate Cerenkov radiation, which would be detectable using the methods demonstrated herein; additional studies are ongoing to characterize the relative role of each process.
The high energy 18
F or 13
N positrons are governed by relativistic physics and eject from the nucleus near the speed of light, c
. The speed of light in a fluid or tissue medium, vm
, is less than in a vacuum (i.e., vm
) and for some fraction of the positron energy spectrum, the positron velocity is initially greater than the speed of light in the medium. The reduction in the speed of light in a medium is due to the interaction of light with the material and is proportional to the refractive index, η
, of the medium:
Water has an η
of 1.33, thus the speed of light in water is vm
. Biological tissue has an η
of ~1.36–1.45 (Mobley 2003
), so the speed of light in tissue is vm
. When the speed of a particle nears the speed of light, the relativistic kinetic energy, E
, is calculated by:
is the particle mass and v
is the velocity. Solutions to Equation 3
reveal that a proportion of the positrons are traveling faster than the speed of light in the medium. For example, for the 0.635 MeV maximum energy of the 18
F positron, the velocity, v
, is 0.895 c
. For a 1.19 MeV positron from 13
N, the velocity is 0.954 c
. During the short time that the charged positron is traveling faster than the speed of light in tissue, there is a burst of Cerenkov radiation. The threshold energy for production of Cerenkov radiation by a positron in tissue (v=vm
in Eq. 3
) is 0.26 MeV (Jelley 1955
, Elrick 1968
, Ross 1968). This is approximately equal to the mean energy of the distribution for 18
F positrons. During transit, the positron will lose energy, eventually falling below the threshold, and will no longer satisfy the Cerenkov criteria. The amount of light (i.e., number of photons) produced by a charged particle with energy exceeding the Cerenkov threshold was first calculated by Frank and Tamm (1937), who shared the Nobel Prize with Cerenkov in 1958. Later, Ross (1969)
noted analytical solutions are complicated by range and energy degradation factors but published a table of some numerically integrated solutions to the Frank-Tamm equations. Interpolating from the table, we estimated that a 0.635 MeV positron would produce approximately 20 photons in the 250–600 nm wavelength range. This is likely a maximum estimate as Collins and Reiling (1939) measured 40 photons from a 1.9 MeV electron whereas an estimate based on the Ross table is 89. Estimates of Cerenkov light emitted from the 18
F used in these studies are approximate and require additional modeling; however, an initial estimate based on the Ross (1969)
tables and the energy spectrum of the decay is 3 photons per decay (or photons per second per Bq) of 18
F activity in water. Experiments and more detailed calculations are being performed to confirm this estimate and to model the effects of photon absorption in the medium and the efficiencies of the CCD camera.
For optical imaging studies, it is vital to consider the role of tissue absorption and scattering if in vivo imaging applications are to be feasible. As expected for Cerenkov radiation, the majority of the light is produced in the blue portion of the spectrum and follows an inverse relationship with the square of the wavelength. An important consideration is that Cerenkov light extends into the green and red region of the EM spectrum as shown in . Whereas blue light has a short range in tissue, limiting in vivo applications, green and red light have deeper tissue penetration, increasing the range of applications for imaging studies. But, like other optical imaging methods, the absorption and scattering of the light within the tissue results in relatively low levels of light output and a blurring effect. We have considered that future study is warranted to determine if for deep structures, such as tissues in vivo, the Cerenkov radiation may stimulate local autofluorescence that is then subsequently detected. While this phenomenon may increase depth of light penetration, it may also complicate the localization of signal. In addition, we are performing experiments to demonstrate that the Cerenkov radiation can be used with UV-excitable fluorophores, such as lanthanides, and the Stokes shift results in an enhanced release of light in a tissue-penetrating band.
The first demonstration of CLI as a viable in vivo
imaging method is presented in . Although the FDG broadly distributes through the animal, the image is thresholded to show the enhanced signal—hence accumulation of FDG—in the xenograft, a CWR22-RV1 (human prostate tumor) xenograft injected subcutaneously. We have previously shown the use of the CWR22-RV1 model for in vivo
FDG PET imaging in conjunction with bortezomib treatment (Zhang 2005
). The CWR22-RV1 xenograft is passaged by tumor fragment from animals bearing tumors and the cells have not been engineered for bioluminescence by luciferase, thus only the FDG and Cerenkov luminescence could be responsible for the emitted light. Although the optical image shown in is compelling evidence of luminescence imaging with Cerenkov radiation, it is interesting to compare this method with PET scanning, since they are both imaging the same tracer. Of course, the method and process of detection and image generation for PET and CLI are different. The light produced as a result of Cerenkov radiation would precede the annihilation event and production of paired photons. Further, minimal tissue absorption and scattering will occur on the 511 keV photons, whereas this is likely to be a major factor in the CLI method. Despite these caveats, both imaging methods rely on the radiotracer in tissue, and one would expect significant agreement between data acquired from each scan.
Note the intensely hot bladders in the PET image and the high degree of luminescence in this region on the optical image. This data supports that the optical measurement, at least for relatively superficial tissues, is a good surrogate for the radiotracer accumulation in the tissue. Experiments are currently being conducted to compare the relative level of sensitivity and performance of PET and optical imaging with PET tracers in drug treatment studies. Modern preclinical PET scanners offer up to ~10% sensitivity, which we expect will greatly exceed the sensitivity of an optical CLI experiment for in vivo applications due to tissue absorption and scattering. Thus as a direct comparison to PET scanning, CLI may best be used as a high throughput proof-of-concept method. For example, in a typical CLI scan, up to 5 animals can be imaged simultaneously in a 3–5 minutes; conversely, PET scans are often 1–2 animals in 10–20 minutes. Conservatively, CLI is 4× faster—but could be 30× depending on experimental design—than small animal PET scanning.