Simulation and optimization of the Cherenkov TOF whole-body PET scanner

Accepted Manuscript Simulation and Optimization of the Cherenkov TOF whole-body PET scanner Marharyta Alokhina, Clotilde Canot, Oleg Bezshyyko, Igor Kadenko, Gérard Tauzin, Dominique Yvon, Viatcheslav Sharyy

PII: DOI: Reference:

S0168-9002(18)30044-5 https://doi.org/10.1016/j.nima.2018.01.027 NIMA 60453

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Nuclear Inst. and Methods in Physics Research, A

Received date : 29 September 2017 Revised date : 10 January 2018 Accepted date : 10 January 2018 Please cite this article as: M. Alokhina, C. Canot, O. Bezshyyko, I. Kadenko, G. Tauzin, D. Yvon, V. Sharyy, Simulation and Optimization of the Cherenkov TOF whole-body PET scanner, Nuclear Inst. and Methods in Physics Research, A (2018), https://doi.org/10.1016/j.nima.2018.01.027 This is a PDF file of an unedited manuscript that has been accepted for publication. As a service to our customers we are providing this early version of the manuscript. The manuscript will undergo copyediting, typesetting, and review of the resulting proof before it is published in its final form. Please note that during the production process errors may be discovered which could affect the content, and all legal disclaimers that apply to the journal pertain.

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Simulation and Optimization of the Cherenkov TOF Whole-body PET Scanner Marharyta Alokhinaa,b,∗, Clotilde Canota , Oleg Bezshyykob , Igor Kadenkob , G´erard Tauzina , Dominique Yvona , Viatcheslav Sharyya a IRFU, b Taras

DSM, CEA, 91191, Gif-sur-Yvette, France Shevchenko National University of Kyiv, Kyiv, Ukraine

Abstract The present work describes the GATE/Geant4 simulation of the TOF wholebody scanner constructed with PbF2 crystals and motivated by the on-going developments of the efficient and fast Cherenkov detector at the IRFU, CEA. We consider different geometrical configurations of the elementary detectors, different options of the optical interface and optimize the scanner construction using TOF-modified noise equivalent count rate calculation. Keywords: PET scanner, Time-of-flight, Cherenkov radiation, Lead fluoride

1. Introduction Positron emission tomography (PET) is a powerful nuclear imaging technique widely used nowadays in oncology, cardiology and neuropsychiatry. The PET technology consists in injecting the patient with a radioactive tracer. The 5

decay of the tracer emits a positron that annihilates with an electron. As a result of the annihilation, two 511 keV gamma rays are emitted back-to-back and registered by the dedicated detectors. The line-of-response (LOR) connects the two points where the photons are detected and allows one to reconstruct the tracer distribution when sufficient number of LORs are accumulated. When

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a whole-body scan is performed, the typical dose required to obtain an image ∗ Corresponding

author Email address: [email protected] (Marharyta Alokhina)

Preprint submitted to NIM A PROCEEDINGS

January 2, 2018

with a good quality ranges from 3 to 5 MBq/kg, which corresponds to a dose of about 150–400 MBq for an adult. This limits the use of the whole-body PET to cases with a positive risk-benefit ratio. The main objective in the development of PET scanners is to reduce the radiation dose received by the patient while 15

keeping the same image quality, or, alternatively, to improve the image quality without increasing the received dose. It is known since the 80s that image quality is improved by adding the time-of-flight (TOF) capability to the PET scanner technology [1, 2]. More precisely, the scanner should have the capacity to measure the difference in the arrival time of two photons with good precision.

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The TOF capability provides information on the localization of the annihilation vertex within the LOR. Only during the last decade the first commercial PET scanner was equipped with a TOF capability. It was reported that the best commercially available scanners reach the coincidence resolving time (CRT) of 325 ps (FWHM) [3]. Commercially available cameras as well as most of the

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laboratory PET devices use scintillator crystals to detect the gamma quanta. Scintillation is a rather slow process and for the fastest scintillators the decay time of the fast component of the signal is of the order of 1 ns. An alternative approach consists in detecting Cherenkov photons [4–6]. The 511 keV gamma is converted to an energetic electron through the photoionization or Compton

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effects. If the material has a sufficiently large refractive index and consequently small speed of light, the recoil electron is relativistic in the media. In such a condition, the recoil electron produces photons. These photons are detected by a micro-channel-plate photomultiplier tube (MCP-PMT) attached to the crystal. The Cherenkov process is extremely fast and photons are radiated at the

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timescale of several picoseconds. This allows to achieve very fast detection with the resolution in time limited mainly by two effects: dispersion of the photon pathlengths and time resolution of the photodetector device. One of the best candidate as a Cherenkov radiator is crystalline lead fluoride, PbF2 . It produces no scintillation light, but only Cherenkov radiation [7]. It is very dense

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(7.8 g/cm3 ) and has one of the highest photoelectric fraction, 46 %. Due to these properties, it is possible to create efficient gamma detectors with a very 2

thin crystal of the order of 10 mm thick and hence minimize the length and dispersion of the photon trajectories. The ability to detect 511 keV photons has been demonstrated in [8]. Since this crystal radiates only Cherenkov light, the 45

overall number of photons is small and the total detection efficiency is limited to 10 % or smaller [8]. This low detection efficiency is a major limiting factor for making very fast TOF-PET devices.

2. Simulation We have developed a simulation of the whole-body TOF-PET scanner using 50

the GATE [9] and Geant4 [10] simulation softwares. For this, we have simulated and propagated every photon generated in the crystals. This method differs from the conventional usage of GATE, with which a parameterized approach is used to simulate detector response consecutive to energy deposition. In case of the Cherenkov radiation, the small number of photons allows to track each

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individual photon in the crystal and make a precise simulation of the detector response. Geometrical Parameters. The foreseen PET tomograph, Fig. 1 consists of 9216 PbF2 crystals with dimensions of 6.5 x 6.5 x 10 mm3 , 3 rings of detectors for an axial field-of-view (FOV) of 180 mm and the transaxial FOV of 91 cm. The

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PET detection unit is a block of 64 (8 x 8 x 1) individual PbF2 crystals optically coupled to a single photomultiplier tube, Fig. 2. As a benchmark PMT we chose to use the PlanaconT M by Photonis an active surface of 53 mm x 53 mm and 64 anodes. Our choice is motivated by the excellent time resolution of this device ∼ 80 ps (FWHM) and its good quantum efficiency ∼ 25 % for a wavelength

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400 nm. Each anode corresponds to a single crystal. To improve the detection efficiency we have considered to use a PMT with a sapphire window (thickness is 1.3 mm, density 3.98 g/cm3 , refractive index 1.78 at 400 nm. This configuration is motivated by developments that are currently on-going at IRFU-CEA [11]. The conventional method to couple a PMT to a crystal consists in using

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an optical media. We have estimated the performance of this configuration 3

Figure 1: Layout of the 3-rings whole-body TOF-PET scanner with lead shields described in the GATE simulation.

and compared it to an ambitious idea of improving the detection efficiency by attaching the sapphire window directly onto the crystal by using a molecular bonding process (that has been patented by the CEA). We have also simulated the use of shields on both sides of the scanner to protect our scanner from 75

photons out of FOV. It is 3 cm thick and has a ring shape with inner diameter 70 cm and external diameter 108 cm. The coincidence time window is 4.0 ns. Crystal Coating. We have simulated three types of crystal coating: diffuse white, black and polished. The diffuse white coating is an optical surface that reflects photons uniformly and independently of their incidence angle with a

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probability of 95 %. The black surface absorbs 100 % of the incident light. The polished coating obeys the Fresnel’s laws for transmission and reflection with a specular reflection probability of 95 %.

Figure 2: Sceme of the PET detection unit, which consists of a block of 64 (8 x 8 x 1) individual PbF2 crystals molecularly bonded with a single photomultiplier tube.

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Dead Time. We did not simulate the dead-time effect in the current simulation. It is expected that the dead time will be caused mainly by the ADC electronics, 85

since the Cherenkov process is extremely fast and the MCP signal is of the order of 5 ns. The dead time depends on the concrete realization of the electronic chain. As an estimation, we can consider a readout module SAMPIC used currently in the laboratory tests [11]. This module can read out with a speed up to 10 kHz per channel, corresponding to an activity of about 60 kBq/cm3 ,

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so we did not expect a significant dead time effects on NECR up to this value.

3. Results Coincidence resolving time calculation. The measurement of the time difference between the arrival of the two photons improves the signal to noise ratio. The resolution in time difference (known as coincidence resolving time, or CRT) de95

pends on the physical properties of the radiator (e.g, the yield and the emission time of the photons), the radiator geometry, the detection efficiency, the time resolution (time jitter) introduced by the photosensors and the readout electronics. Taking into account the quantum efficiency of the photocathode, [12], and transit time spread of the PMT, 80 ps FWHM, we have estimated the CRT for

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different configurations. Distributions of the time differences for two thicknesses of 10 mm and 20 mm with white coating of the crystal molecularly bonded to the PMT are shown on Fig. 3. The crystals with black coating have better CRT, below 40 ps FWHM, but the detection efficiency is very low. The crystals with polished coating have the worst CRT, therefore, diffuse white coating was

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chosen as an optimal. NECR estimation. Coincidence events in PET fall into 3 categories: true, scattered and random. A true coincidence is an event for which both photons from an annihilation event are detected, none of them undergoing any form of interaction prior to detection. We consider only one pair of photons with the highest

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total energy within the coincidence time window. A random coincidence is an event when two photons arising from different annihilations are detected within 5

Coincidences

6000 5000

FWHM = 180 ps FWTM = 740 ps

4000 3000 2000 1000 0−1 −0.8 −0.6 −0.4 −0.2

0

0.2

0.4

0.6

0.8

1

Coincidences

∆t, ns 3500 3000

FWHM = 380 ps FWTM = 1500 ps

2500 2000 1500 1000 500 0−1 −0.8 −0.6 −0.4 −0.2

0

0.2

0.4

0.6

0.8

1

∆t, ns

Figure 3: Distributions of the time differences for two thicknesses of 10 mm (up) and 20 mm (down) with white coating of the crystal.

the coincidence time window. The distribution of random coincidences is fairly uniform across the FOV, and will cause isotope concentrations to be overestimated if not corrected for. Random coincidences also add statistical noise to the 115

data. A scattered coincidence is an event for which at least one of the detected photons undergoes a Compton scattering in the patient’s body or in the phantom. Since the direction of the photon is changed by the Compton scattering process, it is likely that the resulting coincidence event will be assigned to the wrong LOR. Scattered coincidences add a background to the true coincidence

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distribution that changes slowly with position, decreasing contrast and resulting in an underestimation of the isotope concentrations. They also add statistical noise to the signal. Both the random and the scattered coincidences add blurring and therefore degrade image quality. Noise Equivalent Count Rates (NECR) is an ”effective” count rate, which allows to compare theoretical signal-to-noise

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ratios that can be achieved in images from different scanners. It is defined as (see Eq. 1): N ECR =

T2 T + S + 2R

(1)

where T, S and R are the number of the true, scatter and random coincidences respectively. We have calculated the NECR curves as a function of the activity

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concentration in the FOV for the different scanner configurations following the 130

NEMA Standard NU2-2007 [13]. To do this we have simulated a test cylindrical phantom composed of polyethylene with an outside diameter of 200 mm, and an overall length of 700 mm. A 6.4 mm hole parallel to the central axis of the cylinder at a radial distance of 45 mm is filled with radioactivity. We have simulated a source activity range from 10 to 55 kBq/cm3 . Results of

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the simulation are shown in Fig. 4 and compared to the performance of GE’s

NECR, kcps

Discovery-690 scanner [14]. 300

10mm 10mm & TOF 20mm 20mm & TOF Discovery-690

250 200 150 100 50 0

10

15

20

25

30

35

40

45

50

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Activity, kBq/cm3

Figure 4: NECR of the simulated Cherenkov PET scanner and comparison with Discovery690 [14]. Two configurations are concidered with crystal thickness of 10 mm and 20 mm, and with and without TOF correction (see text).

Applying TOF technique for NECR. The standard NECR calculation does not take into account the TOF information, and correspondingly, the Cherenkov detector are expected to have lower NECR compare to conventional scanners 140

due to the lower detection efficiency. We have also calculated the modified NECR (see Eq. 2), which takes into account the TOF information N ECRT OF =

D N ECR ∆x

(2)

where D is the radial dimension of the subject to be imaged, and ∆x is the spatial uncertainty associated with the CRT of the scanner. For example, in the case of 10 mm thick for the phantom diameter D = 20 cm, a CRT ∼ 145

180 ps FWHM corresponding to a spatial uncertainty of ∆x ∼ 2.7 cm FWHM, N ECRT OF will be ∼ 7.4 × N ECR. 7

At 1 cm radius Transverse

2.8 mm

Axial

3.0 mm At 10 cm radius

Transverse radial

3.0 mm

Transverse tangential

3.0 mm

Axial resolution

3.0 mm

Table 1: Spatial resolution

Spatial resolution. Spatial resolution characterizes the width of the point spread function resulting from the reconstruction of a compact radioactive point source. According to the NEMA prescriptions [13], six capillary sources shall be used 150

to measure the spatial resolution (see Fig. 5).

Figure 5: Positions of the radioactive sources for resolution measurement. In our case the FOV of the scanner is 18 cm, 1/4 FOV = 4.5 cm.

Image reconstruction. For image reconstruction we have used the CASToR platform (Customizable and Advanced Software for Tomographic Reconstruction) [15]. The current version of CASToR includes 2D iterative MLEM algorithms. We have modeled a spatially uniform point spread function using a 3D 155

Gaussian of FWHM 2.5 mm transaxially and axially. The reconstructed images have 512 x 512 x 256 voxels 0.5 x 0.5 x 0.5 mm3 . Spatial resolution obtained for a configuration with molecular bonding are shown in Tab. 1. As we can see spatial resolutions do not vary much from the center to the periphery of the

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scanner and are comparable with spatial resolution of conventional scanners.

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4. Conclusion In this work we are investigating the feasibility of developping a Cherenkov whole-body TOF-PET scanner using lead fluoride crystals. This crystal has a high photoelectric fraction and a high crystal density, and hence, provide excellent stopping power for the 511 keV gamma rays. The number of optical

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photons at the photocathode, and thus, the detection efficiency, depends on the crystal size and its optical coating. We have estimated some performance characteristics of such a scanner based on Monte Carlo simulations with the GATE/Geant4 simulation software. An optimal configuration of the detector can be achieved by using of diffuse white coating of the crystal surfaces with

170

which a coincidence resolving time of about 180 ps FWHM or better can be achieved with a reasonable detection efficiency for a 10 mm thick crystal. The spatial resolution of this detector of ∼ 3 mm FWHM is limited by the size of

the crystals 6.5 x 8.5 mm2 in a similar way as for conventional PET scanners. The CRT for 10 mm (1-interaction length) thick crystals is about twice better 175

than for 20 mm thick crystals (2-interaction lengths). Therefore, for the further studies we will focus on the configuration with 10 mm thick crystals. NECR values are comparable with conventional scintillator based PET scanners when TOF is taking into account and could even become better if one improves either detection efficiency (e.g. the quantum efficiency of the photocathode) or

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transition time spread of photomultiplier tube.

5. Acknowledgments We acknowledge the financial support by the LabEx P2IO R&D project of the region Ile-de-France, the ”IDI 2015” project funded by the IDEX ParisSaclay, ANR-11-IDEX-0003-02, the PhD financial support by the French em185

bassy in Ukraine, DRI University Paris-Saclay and the Programme Transversal

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Technologies pour la Sant´e, of CEA. This work is conducted within the scope of the IDEATE International Associated Laboratory.

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