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The goal of this work is to demonstrate that a CCD-based system can be used as a unified device which allows visible, β, X and γ rays imaging. A system composed of a CCD coupled with lens mounted on a black light-tight box and a high resolution intensifying screen for the radiations conversion were used. In order to investigate the detection of different type of radiations in vitro and in vivo experiments were performed. The comparison of the results obtained with our prototype and those obtained with dedicated commercial devices showed a good agreement.
© 2015 Optical Society of America
1. Introduction
Different kinds of imaging techniques are currently available for clinical and preclinical applications.
The commonly used technologies for the acquisition of projection radiographs are the computed radiography (CR) and the digital radiography (DR). CR employs photostimulable phosphor (PSP) plates and a separate image readout process. The PSP exposed to X rays stores the absorbed energy then released when the plate is stimulated by a laser beam of a raster scanner. The photostimulated luminescence is collected and then detected by a photomultiplier tube (PMT). This signal is digitalized by an analog-to-digital converter to form the image. Details on the physics and on the characteristics of CR can be found on the review article by Rowlands [1]. DR flat panel systems integrate an X ray-sensitive layer and an electronic readable system based on thin film transistor (TFT) arrays formed of amorphous silicon (a-Si). Detectors using a scintillator layer and a light-sensitive TFT photodiode are called indirect-conversion TFT detectors. Those using an X ray-sensitive photoconductor layer, such as amorphous selenium (a-Se), and a TFT charge collector are called direct-conversion TFT detectors. An overview and a comparison between the two technologies can be found in [2] and [3]. Digital systems that couple a scintillating phosphor layer to multiple charged coupled device (CCD) cameras by lenses or fibres optics were also developed [4].
A widespread technique developed over the last decades for molecular imaging (MI) of small animals for preclinical studies is bioluminescence imaging (BLI) [5]. This non-invasive method is based on the detection of visible light produced in the oxidation of a substrate (luciferin) catalysed by luciferase enzymes using a CCD camera. In recent years, another technique suitable for preclinical studies on small animals is Cerenkov luminescence imaging (CLI) which combines optical and nuclear imaging together [6–9]. Cerenkov radiation is a well-known phenomenon induced by charged particles that travel in a medium with a velocity v greater than the speed of light c in the medium. A CCD camera is used to detect the Cerenkov radiation produced by β emitting radionuclides allowing imaging their spatial distribution in vitro and in vivo [10–15]. In order to improve optical imaging of Cerenkov radiation quantum nanoparticles can be used as Strokes shifters [16,17]. Recently a proof of the possibility to perform human Cerenkov imaging was given by Spinelli et al. [18]. Another clinical application of Cerenkov luminescence is the endoscopy. Kothapalli et al. [19] and Liu et al. [20] demonstrated the feasibility of the intraoperative tumour imaging using optical fibres coupled to a CCD camera placed in a dark chamber to simulate anatomic cavity. First results of the Cerenkov luminescence endoscopy on patient was obtained by Hu et al. [21].
A CCD-based system can also imaging α and γ emitting radionuclides exploiting the radioluminescence imaging (RLI) with and without the use of scintillating materials [22–25].
Multimodal systems that combine optical and X rays imaging, such as the IVIS Lumina XRMS (Perkin-Elmer) and the Xtreme (Bruker), are commercially available. However, those are closed systems that do not allow customising the material used for radiation conversion based on the experimental needs.
In the present work, we proposed a unified CCD-based approach suitable for the visible light, β particles, X and γ rays imaging. As a proof of principle, we performed in vitro experiments exploiting the CLI and the scintillation imaging with the use of a mammographic intensifying screen and in vivo experiments for the luciferin imaging. We showed that a high resolution intensifying screen, even if optimized for mammographic X rays, is suitable for the conversion of γ and β radiations into optical photons. These experiment are performed with a system that is up to ten times cheaper that the commercial ones. In order to validate our approach each experiment was performed also with the standard and the state of the art dedicated techniques.
The paper is organized as follows: in section 2, we present our system and we describe the experiments performed to validate the proposed approach for β particles, γ and X rays and bioluminescence imaging. The experimental results are illustrated on section 3. Discussion and conclusions then follow.
2. Materials and methods
2.1 Optical system
The optical system used in this work is composed of a cooled (−80°C) electron multiplied charge coupled device (EMCCD) (Andor iXon Ultra) coupled with an f/0.95, 17mm C-mount lens (Schneider Optics). The EMCCD has a back-illuminated 512 x 512 sensor with pixel size equal to 16 µm, resulting in a detector with total dimensions equal to 8.2 x 8.2 mm2. The smallest achievable field of view (FOV) is about 6.1 x 6.1 cm2 resulting in an image pixel size of 120 µm. To exclude ambient light, the system is mounted on a black light-tight enclosure, shown in Fig. 1, in which an adjustable stage is placed for sample positioning. In order to investigate the absence of any external light contamination, five minutes images with a phantom and no light sources were acquired. In this case, the CCD camera detected no light signals. Dark measurements were acquired before any image acquisition in order to perform background subtraction. For each acquisition, the lens aperture diaphragm was set full open in order to maximize the light collection. The quantum efficiency (QE) of the EMCCD is >80% in the visible range (400-800 nm), as shown in Fig. 2.
Fig. 1 Pictures of the imaging system: the CCD camera coupled with the lens is mounted on a black light-tight enclosure. An adjustable stage for the positioning of the sample is placed inside the enclosure.
Fig. 2 Quantum efficiency of the EMCCD camera iXon Ultra as a function of the photons wavelength (adapted from Andor specification sheet). The QE is > 80% in the 450-750nm range.
As described by Robbins et al [26], the noise of the acquired image is dependent on the readout noise of the CCD camera, , that is the minimum electronic noise floor of the device, following the expression below:
where is the noise on the optically generated signal, is the noise on the dark signal and M is the EM gain. is a noise factor that accounts for the noise introduced by the gain and tends to two when the gain is large. Since the decreases reducing the readout rate, not always the use of an EM gain is convenient to reduce the image noise.Considering that acquiring one image at a time at a fast frame rate is not needed, in our experiments we chose not to use any EM gain as also suggested by the vendor. With the conventional amplifier and a 0.08 MHz rate, the performance sheet supplied by the vendor indicates a read noise equal to 3.11 photons/pixel. We chose anyway to use an EMCCD since is a flexible and versatile detector.
As will be discussed later in section 2.3 and 2.4, the optical system was used with and without a high resolution intensifying screen, depending on the experiment performed. This intensifying screen (MAMORAY Detail R screen, AGFA) is composed of green-emitting rare earth phosphors (Gd2O2S:Tb) and it is optimized and commonly used for mammography. The interesting aspect of using such a screen rather than a scintillator slab is the possibility to cover a wider area with a net reduction of the costs [27].
The analysis of the images acquired with our system were performed with the use of the IDL (interactive data language) 8.2 software (Exelis VIS).
2.2 Detection of Cerenkov radiation
As well known, Cerenkov radiation in the visible range has a threshold for production that depends on the refractive index n of the medium in which the particle travels, expressed as
In the acrylic plastic, which has a mean refractive index of 1.49, electrons or positrons threshold energy for the Cerenkov production is about 178 keV. The β- particles emitted by 137Cs have an average energy of 188.4 keV and a maximum energy of 1175.6 keV and thus enough to produce Cerenkov light in the acrylic.
As a proof of principle experiment of β- particles detection with the optical system, the Cerenkov luminescence was exploited. A source of 137Cs sealed in a 2 mm thick acrylic slab was placed on the adjustable stage in the light-tight box and, after acquiring a photographic image of the sample, 300 seconds acquisition was performed. As verified in Cerenkov luminescence experiments previously performed by our group [25], 300 seconds is a reasonable exposure time to have a good light collection. The used 137Cs source had originally an activity of 1 µCi, as we can observe in the image below (Fig. 4), but at the moment of the experiment its activity was 0.7 µCi.
In order to validate our experiment, the same source was imaged with the IVIS SpectrumCT (Perkin-Elmer) used in bioluminescence mode. This system is equipped with a cooled (−90°C), back-thinned, back-illuminated 2048 x 2048 CCD camera with pixel size equal to 13.5 µm, resulting in a detector with total dimensions equal to 2.7 x 2.7 cm2. The measurement was performed setting the field A (40 mm) and a 2 x 2 binning, resulting in an image pixel size of 0.041 mm. The exposure time was set to 300 seconds. Furthermore, in order to investigate the source dimensions, a CT image was obtained with the IVIS SpectrumCT.
2.3 Radiopharmaceuticals QA
The thin-layer chromatography is the standard technique described in the European Pharmacopoeia [28] to perform radiopharmaceutical QA, both for diagnostic and therapeutic radiopharmaceuticals. A small amount (usually few kBq) of radiopharmaceutical is placed at the extremity of a chromatographic strip representing the stationary phase. The strip is then soaked in a solvent (that depends on the radiopharmaceutical) that travels along the strip by capillary action. The radiopharmaceutical that is soluble in the solvent will migrate with it from the original point of deposition. Otherwise, the not soluble impurities will remain at the origin.
Before the administration to a patient, the radiopharmaceutical routine recommends the evaluation of the radiochemical purity (RCP), defined as the percent of the total radioactivity presents in the desired chemical form in a radiopharmaceutical and evaluated as follows:
where R and I are the amount of radiopharmaceutical and impurity along the strip, respectively. The most accurate method to measure these two values is to obtain the profile of the activity distribution along the strip performing a linear scan of the chromatographic strip with a conventional radiation detector (mainly NaI(TI) scintillation counter) [29].2.3.1 Detection of β+ radiation with intensifying screen
In order to detect β+ radiation with the optical system the intensifying screen was used for the conversion into optical photons. As source of this radiation, a 68Ga-labelled radiopharmaceutical was taken into account. More precisely a chromatographic strip where 7 kBq of 68Ga-DOTANOC was previously deposited was imaged. In this case, even if the β+ particles could be imaged with CLI, we decided to use the intensifying screen in order to reduce the acquisition time.
The strip was placed inside the light-tight box and a photographic image of the sample is first acquired. The sample is then covered with the intensifying screen. Considering the strip dimensions (few cm wide and 10-12 cm long), for this measurement, a 15x15 cm2 FOV resulting in an image pixel size of about 290 µm was used. The acquisition was performed with an exposure time of 10 seconds.
In order to validate our approach, the same strip was imaged with the Cyclone Plus Storage Phosphor system (Perkin-Elmer), originally developed for autoradiography and routinely used to perform radiopharmaceutical analysis on chromatographic strips [30, 31]. In the system the chromatographic strip is placed for few minutes on a conventional PSP that records the two dimensional spatial distribution of the radioisotope activity. After removing the strip from the plate, it can be read with a laser beam.
2.3.2 Detection of γ rays with intensifying screen
In order to detect γ rays with the optical system, the same procedure described in section 2.3.1 was adopted for the simultaneous acquisition of two chromatographic strips where 5 kBq of 99mTc-labelled radiopharmaceutical (99mTc-HMDP) was previously deposited. Also in this case, the strips were imaged with the Cyclone Plus Storage Phosphor system.
2.4 Detection of X rays
In order to investigate the detection of X rays, measurements with X rays beams produced by a standard diagnostic tube at 41 kV were performed. After testing different values of charge, it was observed that 20 mAs is a trade-off between a good quality of the image and the presence of hot spots caused by X rays. The use of removal noise tools is required in case of an excessive presence of hot spots; however, this technique can cause image degradation. Furthermore, this charge value is sufficiently low to spare the CCD camera from radiation damages.
The focus to detector distance was set at 1.5 m. In order to convert the X rays into visible light photons then detected by the CCD camera, the intensifying screen was placed between the sample and the detector. Since our purpose is to analyse the maximum resolution achievable by the system, we used the as small as possible FOV.
With this experimental set-up, schematically represented in Fig. 3, we acquired images of a lead grid phantom dedicated for line pair resolution evaluation (with a range from 0.6 to 10 lp/mm) and of a few millimetre thick lead plate used as an edge-test device for the evaluation of the modulation transfer function (MTF) of the system. The lead plate was tilted of few degrees with respect to the axes of the pixel matrix of the detector allowing the use of the slant edge technique for MTF evaluation, as will be discussed later in section 3.3.
In order to validate the performances of the optical system, the same phantoms were imaged using a 35 x 43 cm PSP plate read by the CR-35X (AFGA), with the same exposure parameters and acquisition time. Also in this case the focus to detector distance was set at 1.5 m.
2.5 Bioluminescence imaging
In order to investigate the bioluminescence detection in vivo, a luciferase-tagged myeloma model mouse was injected in the tail vein with D-luciferin (150 mg/kg body weight). The maximum time to peak in not known a priori as shown, for example, in [32], since it depends on the experimental conditions. In order to obtain the maximum signal, we previously performed time study for the specific mouse model and we found that the maximum is reached 10 minutes after D-luciferin injection and the signal is stable approximately for 15 minutes.
Ten minutes after the injection, the mouse, anesthetized with 10 minutes of 2-3% isoflurane and 1 l/min oxygen, was placed in the optical system. A photographic image was first taken and then a 60 seconds emissive image was acquired.
The same mouse was imaged with the IVIS SpectrumCT described in section 2.2, which is the state of the art for the preclinical in vivo optical imaging.
The experiment was conducted following the principles of the NIH Guide for the Use and Care of Laboratory Animals, and the European Community Council (86/609/EEC) directives.
3. Results
3.1 Detection of Cerenkov radiation
In Fig. 4 the fusion of the black and white photographic and of the corresponding emissive image of a 137Cs source obtained with the optical system (a) and with the IVIS SpectrumCT (b) are shown. As one can see, in both cases the signal is localized just inside the circle where the source is positioned. The normalized signal profiles of the two images are reported in Fig. 5. From the analysis of the profiles, we measured a full width half maximum (FWHM) equal to 0.94 mm in both cases. The source diameter, measured from the CT image, was equal to 0.9 mm.
Fig. 4 Fusion of the black and white photographic and of the corresponding emissive image of a 137Cs source obtained with our optical system (a) and with the IVIS SpectrumCT (b).
Fig. 5 Normalized profile of the signal of the 137Cs source detected by our optical system (blue solid line) and by the IVIS SpectrumCT (red dashed line). The measured FWHM is 0.94 mm with both the acquisition systems.
3.2 Radiopharmaceuticals QA
3.2.1 Detection of β+ radiation with intensifying screen
Figure 6(a) shows the fusion of the black and white photographic and of the corresponding emissive image of the chromatographic strip where few kBq of 68Ga-DOTANOC was deposited. In order to enhance the SNR, we observed that the 4x4 binning is the optimal value to set during the acquisition. The normalized profile evaluated along the yellow line drawn on the emissive image is shown in Fig. 6(b). Analysing this profile, that is evaluating the peaks of the radiopharmaceutical and of the impurity, we measured a RCP equal to 97.9%. Figure 7 shows the emissive image of the same chromatographic strip recorded by the Cyclone Plus Storage Phosphor system (a) and the corresponding normalized profile (b). Evaluating the peaks of this profile, we measured a RCP equal to 96.8%.
Fig. 6 The fusion of the black and white photographic image and of the corresponding emissive image of the strip where 7 kBq of 68Ga-DOTANOC was deposited obtained with our optical system (a). The normalized profile of the activity shown on b) is evaluated along the yellow line drawn on the fused image.
Fig. 7 On the left the activity along the chromatographic where 7 kBq of 68Ga-DOTANOC was deposited read using the Cyclone system is shown. The activity profile, normalized and background subtracted, shown on the right is evaluated along the yellow line drawn on the emissive image.
3.2.2 Detection of γ rays with intensifying screen
Figure 8(a) shows the fusion of the emissive and of the black and white photographic image of the two chromatographic strips where few kBq of 99mTc-HMDP were deposited. The normalized profile evaluated along the drawn yellow line is shown in Fig. 8(b). The emissive image of the same chromatographic strips recorded with the Cyclone Plus Storage Phosphor system and the corresponding normalized profile is shown in Figs. 9(a) and 9(b) respectively.
Fig. 8 The fusion of the black and white photographic and the emissive image of the strips where 5 kBq of 99mTc-HMDP was deposited obtained with our optical system (a). The normalized profile of the activity shown on b) is evaluated along the yellow line drawn on the fused image.
Fig. 9 The activity along the chromatographic strips where 5 kBq of 99mTc-HMDP was deposited read using the Cyclone Plus Storage Phosphor (a). The normalized activity profile shown on (b) is evaluated along the yellow line drawn on the emissive image.
3.3 Detection of X rays
Figure 10 shows the X rays image obtained with the optical system of the lead grid phantom (a) and a zoom of the observable finest grid (b). The X rays image of the same phantom obtained with the CR-35X system is shown in Fig. 10(c). The analysis of the images of the dedicated phantom allows characterizing the resolution performances of the systems. The normalized intensity profiles, evaluated along the yellow lines drawn Figs. in 10b and 10c, are compared in Fig. 11. The blue curve, referred to the optical system, highlights that the last well distinct peaks and valleys are those of the fourth grid from the left, corresponding to a maximum resolution of 3.7 lp/mm. The maximum resolution of the CR-35X system results slightly above the 3.4 lp/mm.
Fig. 10 X rays image of lead grid phantom obtained with our optical system (a) and a zoom of the observable finest grid are shown (b). X ray image of the same phantom obtained with the CR-35X is also shown (c). The intensity profiles were evaluated along the yellow lines drawn on the image.
Fig. 11 Normalized intensity profiles evaluated along the yellow lines drawn in Figs. 10(b) and 10(c) on the lead grid phantom X rays image obtained with the optical system (blue solid line) and with the CR-35X (red solid line).
In addition, as for standard radiographic systems, the MTF was evaluated using the slant edge target technique [33] derived from the technique of Fujita et al. [34]. The MTF is evaluated as the Fourier transform of the line spread function (LSF) obtained as the derivative of the edge spread function (ESF). Figure 12 shows the image of the lead plate used as an edge-test device obtained with the optical system (a) and with the CR-35X system (b). The regions of interest (ROIs) considered for the evaluation of the edge profiles are also shown. In Fig. 13, the plots of the normalized MTFs obtained for the two systems are compared.
Fig. 12 X rays images of a few millimeters lead thick plate tilted of few degrees with respect to the axes of the detector pixel matrix used as an edge-test device obtained with the optical system (a) and with the CR-35X system (b). The MTFs of the systems are evaluated starting from the ESF measured in the drawn ROIs.
Fig. 13 Normalized MTF of the optical system (blue solid line) and of the CR-35X (red dashed line) obtained with the slant edge technique.
3.4 In vivo small animal bioluminescence
Figure 14 shows the fusion of the black and white photographic image and the corresponding bioluminescence signal obtained with the optical system (a) and with the IVIS SpectrumCT (b). The emissive images are reported in normalized values. A quantitative comparison between the two systems is not possible since our system is not calibrated in light radiance (p/s/cm2/sr) as the IVIS SpectrumCT output. Nevertheless, as one can see the signal has the same localization in the two images.
Fig. 14 In vivo BLI image acquired respectively with the optical system (a) and with the IVIS SpectrumCT used in bioluminescence mode (b). The emissive images were fused with the black and white photographic image of the mouse taken with the two systems.
4. Discussion
The experimental results presented in the previous section demonstrated with respect to the state of the art techniques the feasibility of a unified CCD-based approach for the imaging of different type of radiations, ranging from visible to X, γ and β radiation.
The results obtained with the 137Cs showed the detection of Cerenkov radiation. The detected signal, shown in Fig. 4, is due to Cerenkov light induced by the emitted β- particles travelling in the acrylic thickness over the source. As one can see from Fig. 5 the FWHM measured from the intensity profile of the source signal detected with our optical system and with the IVIS SpectrumCT were equal. We observe that the FWHM value (0.94 mm) is mainly affected by the physical dimensions of the source itself that cannot be considered as a point source. Actually, a CT image showed that the source has a diameter equal to 0.9 mm.
The imaging of the chromatographic strip where few kBq of 68Ga-DOTANOC were deposited showed the detection of β+ radiation. We observed that an acquisition time of 10 seconds is sufficient to detect the signal needed to perform a quantitative analysis of the strip, providing a very quick method to evaluate the RCP required by the routine radiopharmaceutical QA. Comparing the RCP values obtained with our approach and with the Cyclone Plus Storage Phosphor system, commonly used for such analysis, we observed a 1.1% difference, denoting a good accuracy of the optical method. It is worth noting that, also in this case, rather than using an intensifying screen, it would be possible to image the β+ emitting radiopharmaceuticals exploiting the CLI, using a transparent Cerenkov radiator with high refractive index. In this experiment the use of the intensifying screen was preferred to the CLI in order to enhance the signal and to reduce the acquisition time, mandatory when short half-life radioisotopes have to be studied.
The imaging of the chromatographic strips where few kBq of 99mTc-HMDP were deposited showed the detection of γ rays. The comparison of the profiles of the signals detected with our system and with the Cyclone Plus Storage Phosphor system (Figs. 8(b) and 9(b)) showed a good agreement between the two, confirming the efficiency of the optical approach. Furthermore, this experiment showed the possibility to perform the analysis of different samples simultaneously, without the need to increase the measurement time. In this case, the fusion of the photographic image with the corresponding emissive image allows spatially localizing the different signals. This tool is also useful to properly identify the strip.
It is interesting to note that despite the intensifying screen used in this work is optimized for low energy X rays used in mammography, the obtained results showed that it is also suitable for the conversion of different kind of radiations, from X and γ rays to β particles. Nevertheless, it will be interesting to further investigate the use of different screens.
The data obtained with X ray beam showed the possibility to acquire X ray images. The achievable resolution with the optical system is 3.7 lp/mm and it is comparable with the resolution achievable with the conventional CR system as the CR-35X. The MTFs of the systems was evaluated using the slant edge technique. The analysis of the MTF confirms the resolution observed from the lead grid phantom. We believe that our system could be used in the preclinical to perform radiographies of small animals (mainly mice).
Lastly, the in vivo result showed the detection of bioluminescent signal produced in small animal. Even though our system is not calibrated in light radiance as the IVIS SpectrumCT output, a qualitative comparison between the signals acquired by the two systems showed a good agreement in the signal localization. It is important to note that for in vivo experiment a gas anaesthesia integrated system could make it possible to increase the exposure time and thus to enhance the detected signal.
Once installed such a system, it could be possible to use the bioluminescence imaging in combination with the X rays modality of the optical system, allowing the localization of the emissive signal on the internal anatomy of the mouse.
5. Conclusions
In this paper, we demonstrated that a CCD-based system can be used as a unified device that allows luciferin, Cerenkov luminescence, β, X and γ rays imaging. In vitro and in vivo experiments were performed in order to validate our prototype with the state of the art imaging systems and methods.
Acknowledgments
AES and CRG are funded by the Italian ministry of health grant no. GR-2010-2309585. The authors would also like to acknowledge Dr. Giancarlo Gorgoni for the useful discussions and Dr. Michela Frenquelli for providing the mouse model.
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