Performance evaluation of the CT component of the ...

Nuclear Instruments and Methods in Physics Research A 805 (2016) 135–144

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Nuclear Instruments and Methods in Physics Research A journal homepage: www.elsevier.com/locate/nima

Performance evaluation of the CT component of the IRIS PET/CT preclinical tomograph Daniele Panetta a,n, Nicola Belcari b, Maria Tripodi a, Silvia Burchielli c, Piero A. Salvadori a, Alberto Del Guerra b a

CNR Institute of Clinical Physiology (IFC-CNR), v. G. Moruzzi 1, I-56124 Pisa, Italy Department of Physics “E. Fermi”, University of Pisa, L.go B. Pontecorvo 3, I-56127 Pisa, Italy c Fondazione CNR/Toscana “G. Monasterio” – FTGM, v. G. Moruzzi 1, I-56124 Pisa, Italy b

art ic l e i nf o

a b s t r a c t

Available online 29 August 2015

In this paper, we evaluate the physical performance of the CT component of the IRIS scanner, a novel combined PET/CT scanner for preclinical imaging. The performance assessment is based on phantom measurement for the determination of image quality parameters (spatial resolution, linearity, geometric accuracy, contrast to noise ratio) and reproducibility in dynamic (4D) imaging. The CTDI100 has been measured free in air with a pencil ionization chamber, and the animal dose was calculated using Monte Carlo derived conversion factors taken from the literature. The spatial resolution at the highest quality protocol was 6.9 lp/mm at 10% of the MTF, using the smallest reconstruction voxel size of 58.8 μm. The accuracy of the reconstruction voxel size was within 0.1%. The linearity of the CT numbers as a function of the concentration of iodine was very good, with R2 40.996 for all the tube voltages. The animal dose depended strongly on the scanning protocol, ranging from 158 mGy for the highest quality protocol (2 min, 80 kV) to about 12 mGy for the fastest protocol (7.3 s, 80 kV). In 4D dynamic modality, the maximum scanning rate reached was 3.1 frames per minute, using a short-scan protocol with 7.3 s of scan time per frame at the isotropic voxel size of 235 μm. The reproducibility of the system was high throughout the 10 frames acquired in dynamic modality, with a standard deviation of the CT values of all frames o 8 HU and an average spatial reproducibility within 30% of the voxel size across all the field of view. Example images obtained during animal experiments are also shown. & 2015 Elsevier B.V. All rights reserved.

Keywords: Micro-CT Preclinical imaging Contrast agents Modulation transfer function Dynamic imaging

1. Introduction Animal models of human diseases play a pivotal role in understanding the underlying mechanism of several pathologies, as well as assessing the effectiveness of new drugs and therapeutic approaches [1,2]. A wide range of imaging instrumentation dedicated for the study of small rodents, mainly mice and rats, have been built and validated in the last two decades by adapting the performance of equivalent instrumentation for clinical use to the higher spatial resolution and sensitivity needed for such small subjects [3–8]. In particular, several micro-computerized tomography (microCT) scanners have reached the preclinical imaging market so far, featuring spatial resolutions roughly 10 times better and temporal resolutions at least 10 times worse than the clinical counterparts [9,10]. Unlike micro-CT scanners dedicated for e.g. material analysis, non destructive testing (NDT) and very high resolution ex-vivo imaging of biological samples, featuring resolutions of the order of n

Corresponding author.

http://dx.doi.org/10.1016/j.nima.2015.08.044 0168-9002/& 2015 Elsevier B.V. All rights reserved.

1–10 μm [11–13] or even sub-micron [14–16], the main design target of most recent micro-CT for in vivo imaging of small animals is the reduction of the scanning time and radiation dose while preserving a spatial resolution in the order of magnitude of 0.1 mm. Keeping the animal dose at low levels (o100 mGy) is very important to enable longitudinal studies avoiding radiation-related injuries, in agreement with the 3 R's law (reduction, refinement and replacement) [17] and with the most recent recommendations and directives on the use of animals in experimental research [18–20]. The last few years have seen a growing interest in the development and validation of new contrast agents (CA) for preclinical micro-CT [21] and, at the same time, in the implementation of dynamic microCT instrumentation and scanning protocols enabling functional imaging via kinetic analysis of CA's time attenuation curves (TAC) [22,23]. Nevertheless, few commercial micro-CT scanners to date allow dynamic imaging with sufficient temporal resolution. In this paper, we aim to evaluate the image quality performance and dynamic imaging capability of a new commercial micro-CT scanner (IRIS) integrated in a multimodal PET/CT scanner for preclinical imaging of small rodents. Attention was paid in the determination of

D. Panetta et al. / Nuclear Instruments and Methods in Physics Research A 805 (2016) 135–144

the mean animal dose for the available scanning protocols implemented in the scanner; moreover, the emphasis was given to the relationship between reconstructed CT numbers and iodine concentration at different X-ray tube voltages (from 35 to 80 kV) and the reproducibility of reconstructed images in dynamic modality in terms of CT numbers and spatial reproducibility. The paper is structured as follows: in Section 2, we first describe the micro-CT scanner and list the main technical specifications; then we describe the phantoms employed for the image quality performance assessment and the instrumentation used for dosimetry, as well as the methodology of such measurements. In Section 3, the results of these physical measurements are presented. Section 4 is devoted to a qualitative description of the scanner potentiality in experiments with small animals. Finally, all the presented results are discussed in Section 5.

2. Materials and methods 2.1. Description of the IRIS micro-CT scanner The IRIS scanner is a novel multimodal preclinical tomograph for high resolution PET/CT imaging of small animals. Both the PET ring and the CT components are placed on the same rotating gantry having a total rotation range in the interval of [
190°; þ190°]. The longitudinal offset between the centers of the PET and CT fields of view (FOV) is 120 mm. The full description of the PET section is beyond the purpose of this article and can be found elsewhere [24]. The CT section is equipped with a microfocus X-ray source (35–80 kV, 80 W) with fixed tungsten anode and a nominal maximum focal spot size of 50 μm, and with a flat-panel CMOSbased X-ray detector coupled to a 150 μm thick CsI(Tl) scintillator. The detector is subdivided in 1536  1944 square pixels (transaxial x longitudinal) with side length of 75 μm, for a total active surface of 115  145 mm2; working at full resolution (i.e., without rebinning) the minimum exposure time per frame in sequence mode is 39 ms, which can be shortened to 15 ms and 12 ms using 2  2 and 4  4 hardware rebinning, respectively. The nominal source-toaxis distance (SAD) and source-to-detector distance (SDD) are 206 mm and 262 mm, respectively, resulting in a transaxial FOV diameter of 90 mm. Due to the cone-beam geometry, the longitudinal FOV length for single rotation scans depends on the distance r ¼(x2 þy2)1/2 from the axis of rotation (AOR) ( ¼z axis) and ranges from 90 mm at r ¼45 mm (only available without cropping) to 110 mm along the AOR. The more relevant hardware specifications of the IRIS micro-CT are listed in Table 1. Images are reconstructed using a Feldkamp-type filtered backprojection algorithm [25] on a dedicated workstation equipped with two Intel Xeon CPU, 2 GHz clock frequency, 8 cores per CPU and 64 GB of RAM, working under a Linux x64 operating system, where data are automatically transferred and processed at Table 1 Summary of the hardware specification of the IRIS CT scanner. Parameter

Value

Source anode material Source voltage Max. anode current Detector matrix size Detector pixel size Source-axis distance Axis-detector distance Transaxial field of view (max) Axial field of view (r¼ 0 mm/r¼ 45 mm) Maximum scan length (multi-bed acquisition) Gantry rotation range

W 35–80 kV 1 mA 1536  1944 75 μm 206 mm 56 mm 90 mm 110 mm/90 mm 250 mm 380°

the end of each study. The options for data preprocessing (cropping and selection of voxel size) and reconstruction (apodization window and pre-corrections for ring artifacts and beam hardening) are all user-selectable via a dedicated graphical user interface. All reconstructed images are calibrated in Hounsfield units (HU), converted in a DICOM3 compliant image format and exported to a local DICOM server. Typical reconstruction times range from 20 s for a low resolution scan to 2–10 minutes for the most used medium to high resolution protocols (see next section); the maximum reconstruction time observed for the workstation in use was 45 minutes for a final volume of 1536x1536  972 voxels from 2000 uncropped projections, without rebinning. 2.2. Scanning protocols The assessment of the imaging performance of the IRIS microCT scanner has been carried out using a set of scanning protocols covering the whole range of scan times, photon energy and spatial resolutions available for this tomograph. Even though this was the “standard” (i.e., pre-loaded) set of imaging protocols, new protocols can be added by the user by editing the relevant parameters for acquisition, preprocessing and reconstruction. The set of standard protocols is reported in Table 2. The available tube voltage settings in the standard set of protocols are 35, 50, 65 and 80 kV, with a number of projections for Table 2 Summary of the standard set of acquisition protocols implemented in the IRIS CT scanner. For each combination of tube voltage and voxel size, the corresponding number of projections, acquisition time and time–current product are reported. Voxel size (nominal) (μm)

Voxel size (DICOM) (μm)

No. of projections

Rotation typea

60

58.812

2000

FS

80

78.416

1280

FS

120

117.624

800

FS

160

156.833

576

FS

240

235.249

400

FS

240

235.249

228

SS

a

Tube voltage (kV) Total scan time (s)/Time– current product (mA s) 35 560/ 560 358/ 358 224/ 224 147/ 147 112/ 112 –

50 280/ 280 224/ 224 140/ 140 98/ 98 40/ 40 –

65 140/ 140 89.6/ 89.6 51.2/ 51.2 35.1/ 35.1 25.6/ 25.6 –

80 112/ 94 64/ 56.3 40/ 35 19.6/ 19.6 16/16 7.3/ 7.3

FS: full-scan; SS: short-scan.

CTDI100,air

2.0

Fit of CTDI100,air

CTDI100,air (mGy/mAs)

136

1.5

1.0 Equation

y = a + b*x^c

Adj. R-Squa

0.5

1 Value

-- a

-0.3544

Standard Err 0.00729

-- b

0.0012 6.08905E-5

-- c

1.7082

0.01032

0.0 30

40

50

60

70

80

Tube voltage (kV) Fig. 1. CTDI100 values normalized for the time–current product, measured free in air at the center of the FOV.

D. Panetta et al. / Nuclear Instruments and Methods in Physics Research A 805 (2016) 135–144

full-scan mode (2π) ranging from 400 to 2000 depending on the reconstruction voxel size. Five nominal voxel sizes are available: 60, 80, 120, 160 and 240 μm. The actual voxels sizes reported in the DICOM header are slightly smaller (see Table 2), but we will use these values in the following text for the sake of brevity, as they are the nominal values reported in the scanner user interface. All the protocols are created by combining together all the available tube voltages and voxel sizes. Twenty out the 21 standard protocols were full-scan (FS), whereas the only short-scan (SS) protocol implemented is recommended for low-resolution, dynamic (4D) imaging. The rotation range is [
180°;þ180°] for all FS protocols and [
110°;þ110°] for the SS protocol, excluding the extra angles for gantry acceleration and deceleration. The total scan time varies from 16 s to 560 s for the FS protocols, and is 7.3 s for the SS protocol. In the following sections, we will use the following notation to denote protocols: P[voxel size]-[kV]-[scan mode]. For example, P120-65-FS will denote the protocol with nominal voxel size of 120 μm, acquired at 65 kV in full-scan mode. 2.3. Phantoms To evaluate the linearity of the reconstructed CT numbers with respect to the concentration of a iodinated contrast agent (CA), a hollow PMMA phantom with inner diameter of 40 mm (typical rat size) was modified by adding inside its water filled cavity 5 Eppendorf 3810X micro test tubes containing mixtures of water and CA with the following concentrations: 1.96, 3.85, 10.33, 18.39, 34.78 mg I/mL. The accuracy of the reconstructed voxel size was measured by mean of an in-house built phantom, consisting in three small zeolite beads (1 mm of diameter) placed at the three vertices of a planar L-shaped piece of a resin matrix board, with a hole pitch of 2.54 mm. The distances between each pair of beads were 35.56, 40.64 and 54 mm. The spatial resolution was measured with a gold wire with a cross sectional diameter of 20 μm, suitably tensioned on a PMMA framework that allowed to put it in various positions of the scanning FOV. 2.4. Dosimetry The dosimetric measurements have been performed for all available voltage settings using a DCT10-RS pencil ion chamber connected to a Barracuda quality assurance (QA) multimeter (RTI Electronics AB, Mölndal, Sweden). The computed tomography dose index (CTDI) over a length of 100 mm is defined as:

CTDI100 =

1 NT

50 mm

∫−50 mm D (z) dz

(1)

where NT is the nominal beam aperture along the z axis and D(z) is the measured dose profile. Due to the relatively large diameter of the ion chamber (1 cm) with respect to the typical size of a small rodent, in-phantom measurements of the CTDI were considered meaningless especially at the edge of the FOV; for this reason, we have not performed them [26]. Only the CTDI measured free in air at the center of the FOV (CTDI100,air), normalized for unit time– current product, was considered in this study. As the longitudinal X-ray beam width of the IRIS scanner is longer than the active length of the pencil chamber, we have used NT ¼100 mm thus obtaining a CTDI100,air that is a good approximation of the dose D (0) free in air at the isocenter. These measurements allowed us to calculate the mean dose to animals of standard diameters (20 mm for mice, 40 mm for rats) for all protocols, by using Monte Carlo conversion factors from dose in air to mean animal dose [27].

137

2.5. Geometric accuracy One of the key features of micro-CT is its ability to perform quantitative measurements of distances, surface areas and volumes. Hence, it is very important to assess the accuracy of the reconstructed voxel size. To this purpose, the L-shaped bead phantom was placed approximately at the axial center of the scanner FOV with random orientation of the lines joining the three pairs of beads. A scan with the highest quality protocol was performed (P60-80-FS) and reconstructed with a standard filter. The resulting DICOM series was then imported in ImageJ [28] and the distance between each pair of beads was computed after measuring the center of mass (COM) of each bead using the following formula:

Dij, meas (xCOM, i − xCOM, j )2 + (yCOM, i − yCOM, j )2 + (zCOM, i − zCOM, j )2 ,

=

i, j = 1, 2, 3, i ≠ j,

(2)

where the subscripts “meas” denotes a measured distance (i.e., based on reconstructed images). Then, the geometric accuracy Ageom was computed as:

Ageom = 1 −

1 3

∑ i≠j

Dij, meas , Dij, theor

(3)

with “theor” denoting the theoretical distance between the beads i and j. 2.6. Spatial resolution The spatial resolution of the micro-CT was measured on the transaxial plane using the gold wire phantom, at three different distances from the AOR: 1 mm (D0), 25 mm (D1) and 35 mm (D2). The measurements were made with protocols employing the maximum rating of the X-ray tube for all voxel sizes (P60-80-FS; P80-80-FS; P120-80-FS; P160-80-FS; P240-80-FS). For the P60-80FS protocol, the reconstruction was done using the three available apodization windows (standard, smooth and edge), whereas only the standard filter was used for all other protocols. The in-plane oversampled line spread function (LSF) was obtained from reconstructed images as described by Kwan et al [29]. After that, the radial and tangential LSFs where fitted to a 1D Gaussian function and the mean full-width at half maximum (FWHM) was determined. The Gaussian shape of the LSF was exploited to compute the spatial frequency at 10% of the modulation transfer function (MTF), denoted by f10, and the corresponding spatial resolution (i.e., low contrast detectability) denoted by R10:

f10 =

R10 =

2 ln 10 2 ln 2⋅ln 10 0.804 = ≈ 2πσ FWHM π⋅FWHM 1 ≈ 0.622⋅FWHM 2f10

(4)

(5)

2.7. Linearity In many practical applications, in vivo CT and micro-CT imaging studies make use of iodinated CA's administered via a tail vein. For this reason, the knowledge of the relationship between CA concentration and reconstructed CT numbers at various scanning conditions might be very important for setting up a successful experiment. The linearity phantom was scanned with all the described protocols, and tomographic images were reconstructed

138

D. Panetta et al. / Nuclear Instruments and Methods in Physics Research A 805 (2016) 135–144

with the standard filter. A cylindrical ROI was placed in each Eppendorf tube and in the water background. For each ROI and protocol, the mean and standard deviation (SD) of the CT numbers Pxx-35-FS - Mouse Pxx-35-FS - Rat Pxx-50-FS - Mouse Pxx-50-FS - Rat Pxx-65-FS - Mouse Pxx-65-FS - Rat Pxx-80-xS - Mouse Pxx-80-xS - Rat

160

Mean dose to animal (mGy)

140 120

have been computed. A linear fit of the obtained ROI values was performed to determine the slope and linear correlation coefficient (R2) for all the available tube voltage settings. The consistency of the reconstructed CT numbers for different voxel sizes at a fixed voltage and CA concentration was also evaluated. 2.8. Contrast to noise ratio An important figure of merit, useful to quantitatively discriminate object features that are “visible” in the image from those that cannot be distinguished from the background, is the contrastto-noise ratio (CNR):

100 80 60 40

CNR =

20 0 60

80

120

160

240

240 (SS)

Nominal voxel size (μm) Fig. 2. Calculated animal dose values for the standard set of acquisition protocols. These values were calculated using the measured CTDI100,air values, considering 20 mm and 40 mm for the diameter of mice and rats, respectively.

#CTobj − #CTbkg 1 (σ 2 obj

+ σ bkg )

(6)

To evaluate the low contrast performance of IRIS scanner for various doses and spatial resolutions, we have measured the CNR for the CA insert having the second lowest concentration (3.85 mg I/ml) of the linearity phantom. A CNR 43 was used as a threshold for visibility of this insert.

2.9. Dynamic imaging capability A useful feature of the IRIS micro-CT is its ability to acquire a sequence of tomographic images (dynamic mode) for application in perfusion CT studies. The dynamic mode was tested using the fastest protocol available (P240-80-SS), with scan duration of 7.3 s. We have assessed the dynamic capability by acquiring 10 consecutive frames, measuring the interscan delay, the geometric reproducibility and the CT number reproducibility. The interscan delay was measured using the timestamp recorded at the beginning of each scan, with a time resolution of 1 s; the delay was calculated by subtracting the time difference between two consecutive scans and the scan duration. The geometric reproducibility was evaluated by measuring the spatial shift of a single bead of the geometric accuracy phantom among all scans along x, y and z. For the CT number reproducibility, a ROI placed at the center of one insert of the linearity phantom was used to evaluate the constancy of the reconstructed values among all the time frame, so as to identify any drift of the system due to e.g., temperature increase of the X-ray detector, fluctuations of the X-ray tube output or image lag.

Fig. 3. The geometric accuracy phantom.

300 Standard filter Smooth filter Edge filter

1.0

250 200

0.6

R10 (μm)

Normalized LSF

0.8

D0 D1 D2

0.4 0.2

150 100

0.0

50

-0.2

0 -0.4

-0.3

-0.2

-0.1

0.0

0.1

x-xCOM (mm)

0.2

0.3

0.4

60

80

120

160

240

Nominal voxel size (μm)

Fig. 4. Left: normalized oversampled line spread function (LSF) for the high quality protocol P60-80-FS, reconstructed with the three available filters. Right: low contrast resolution, as defined in Eq. (5), measured for all the scanning protocols using the standard reconstruction filter, for the gold wires placed at distance D0¼ 1 mm, D1 ¼25 mm and D2 ¼35 mm from the AOR.

D. Panetta et al. / Nuclear Instruments and Methods in Physics Research A 805 (2016) 135–144

58.85 μm, 58.81 μm and 58.61 μm along the x, y and z direction, respectively. The geometric accuracy calculated as explained in Section 2.5 was Ageom ¼9.96  10
4, meaning that the voxel size is accurate within 0.1% of that reported in the DICOM header.

3. Results 3.1. Dosimetry The results of the CTDI100,air along the AOR normalized for unit time–current product are shown in Fig. 1. The dose values varied from a minimum of 0.196 mGy/mA s at 35 kV, to a maximum of 1.902 mGy/mA s at 80 kV. The measured values for all the voltages were fitted with a non linear function of type CTDI100,air(kV)¼ aþ b  kVc using a Levenberg–Marquardt algorithm, as reported in Fig. 1. Combining the normalized CTDI with the reported time– current products (Table 2) we have calculated the CTDI100,air for all protocols, ranging from 165 mGy (P60-80-FS) to 13 mGy (P240-80SS). The value of CTDI100,air for each protocol was used as the best approximation of the dose in air at the CFOV, Dair(0), which in turn was used to calculate the average dose to the animal Dmouse, and Drat . The results are shown in Fig. 2. The maximum dose of 158 mGy was obtained with the P60-80-FS protocol for the mouse (20 mm diameter), whereas for the fastest protocol used in dynamic modality (P240-80-SS) the average dose was as low as 12 mGy. Apart for the 35 kV setting, all the implemented protocols gave similar doses among the different kV settings, for a fixed voxel size. 3.2. Geometric accuracy The geometric accuracy phantom is shown in Fig. 3. For the high-quality protocol (P60-80-FS), the voxel size reported in the DICOM header was 58.812 μm while the measured voxel size was Table 3 Spatial resolution for the high quality (P60-80-FS) protocol, for different reconstruction filters and positions in the FOV. Distance from center (mm)

Filter type FWHM (μm) / Resolution at 10% MTF, f10 (lp/mm) / Low contrast resolution, R10 (μm)

1 25 35

139

Standard

Smooth

Edge

116.9/6.9 /72.7 141.8/5.7/88.2 166.8/4.8/103.7

125.1/6.5/77.8 145.8/5.5/90.7 170.6/4.7/106.1

104.4/7.7/65.0 114.1/7.1/71.0 128.2/6.3/79.7

In Fig. 4 (left), we have reported the oversampled LSF for the highest quality protocol, P60-80-FS, measured at D0 ¼ 1 mm from the AOR using the gold wire phantom. A comparison of the three available reconstruction kernel is shown. The edge-enhancing behavior of the edge filter is clearly visible, even though no evident differences can be appreciated in terms of LSF width. The measurements of the FWHM resolution after fitting of the LSF with a 1D Gauss function gave 117 μm, 125 μm and 104 μm for the standard, smooth and edge filter, respectively. These results correspond to a spatial resolution of 6.9 line pairs per mm (lp/mm) (R10 ¼73 μm) at 10% MTF using the standard reconstruction filter. The numerical results of the spatial resolution for all the three distances from the AOR and the three filters for the P60-80-FS protocol are reported in Table 3. For all other voxel sizes, the obtained values of R10 are shown in Fig. 4 (right). These results refer to a tube voltage of 80 kV and were obtained with the standard reconstruction kernel. For all protocols, there is an increase 20–40% of the LSF width moving away from the AOR. This can be attributed to the finite integration time of each projection and the continuous rotation of the gantry, and is expected even though the angular sampling appears to be sufficient for all protocols. 3.4. Linearity A cross-section of the linearity phantom passing from the inserts at various concentration of iodinated CA is shown in Fig. 5 (left). The standard deviation of the average ROI values was always less that 8 HU among all the available voxels sizes, for a given CA concentration. The reconstructed CT numbers showed a strong linear correlation with respect to the CA concentration for all voltages, as shown in Fig. 5 (right). The Pearson's correlation coefficient was R2 40.996 for all voltages. This good linearity with the C5 insert placed at the center of the phantom, differently from all other inserts that are more close to the periphery of the

80kv 65kv 50kv 35kv Linear Fit of 80kv Linear Fit of 65kv Linear Fit of 50kv Linear Fit of 35kv

1800 1600 1400

CT Number (HU)

3.3. Spatial resolution

1200 1000 800 600 400 200 0 -200 0

5

10

15

20

25

30

35

Iodine concentration (mg/mL) Fig. 5. Left: transaxial slice of the linearity phantom, acquired with the P120-80-FS protocol (voxel size: 117.6 μm) (display window:
500 HU; 1500 HU-min; max). Right: plot of the CT numbers measured for all iodine concentrations and for all tube voltages.

140

D. Panetta et al. / Nuclear Instruments and Methods in Physics Research A 805 (2016) 135–144

phantom, was also an indirect evidence of the good image uniformity at all kilovoltages. The slopes of the linear fits of Fig. 5 were used to measure the sensitivity of the system to iodine concentration. The results are graphically reported in Fig. 6 for all voltages, and show no big differences in sensitivity for the 50, 65 and 80 kV settings (47, 49 and 46 HU/mgI/mL, respectively). On the other hand, a

considerable decrease of sensitivity was observed for the 35 kV settings (21 HU/mgI/mL). This fall of sensitivity is expected due to the high filtration of the X-ray tube (2 mm Al) that narrowed the X-ray spectrum just below the iodine k-edge (33 keV). For our purposes, we felt that the 35 kV setting is less useful than other settings, also because of the long scan times required (see Table 2). Nevertheless, the ability of the scanner to work at 35 kV can be useful for dual energy imaging protocols and will be the subject of further investigations.

55

3.5. Contrast to noise ratio Iodine sensitivity (HU/mg/mL)

50

The CNR values of the C2 insert (3.85 mGI/mL) vs. water background in the linearity phantom are displayed in Fig. 7 (left), for all the protocols. For the same protocols and phantom measurements, the corresponding noise levels in water, calculated as the standard deviation of the CT numbers in the C0 ROI of the linearity phantom, have been reported on Fig. 7 (right). It can be seen that the visibility threshold (CNR 43) is reached for the protocols with voxel size ≧120 μm and for tube voltage ≧50 kV. At 120 μm, the CNR at 65 kV and 80 kV was barely below the threshold (2.87 and 2.84, respectively). Indeed, it must be noted that the selected concentration is generally lower that that achieved in common studies in vivo and was chosen just to emphasize the trade-off between dose, noise, scan time and spatial resolution. For instance, we have observed that the protocol P160-65-FS seems attractive for CA-enhanced longitudinal studies, providing a good trade-off between dose (31 mGy), scan time (35 s), and spatial resolution (3 lp/mm at 10% MTF), with a CNR44.5 at a concentration below 4 mGI/mL.

45 40 35 30

hν < iodine k-edge for almost the whole spectrum

25 20 30

40

50

60

70

80

Tube voltage (kV)

35 kV

80 kV

3.6. Dynamic imaging capability

Fig. 6. Top: iodine sensitivity as a function of tube voltage, calculated as the slope of the least-square fits shown in Fig. 5. Bottom: comparison of two transaxial slices of the linearity phantom, acquired with the protocols P120-35-FS (left) and P12080-FS (right). The drop of the attenuation coefficient of iodine below the iodine k-edge is evident at 35 kV. For both images the display window (min; max) is (
500 HU;1500 HU).

The dynamic modality was tested placing both the linearity phantom and the geometric accuracy phantom in the scanner FOV, and then acquiring 10 consecutive frames with the P240-80-SS protocol (scan time ¼ 7.3 s). With this protocol, the average7standard deviation of the CNR of the C2 insert vs. water was 3.49 70.39 and the noise on the water ROI was 52 76 HU (n ¼10). The reproducibility of the reconstructed CT numbers was measured on the linearity phantom in water (C0) and for the C2 and C5 inserts, and is shown in Fig. 8. No appreciable drifts of the

Water+Iodine, 3.85 mg I /mL, rat phantom

1

40 s; 17 mGy 26 s; 23 mGy 16 s; 24 mGy

240 220

Pxx-35-FS Pxx-50-FS Pxx-65-FS Pxx-80-FS

200 180 160

Noise (HU)

147 s; 14 mGy

224 s; 21 mGy

2

224 s; 96 mGy 90 s; 80 mGy 64 s; 83 mGy

3

358 s; 34 mGy

4 560 s; 53 mGy 280 s; 121 mGy 140 s; 125 mGy 112 s; 139 mGy

CNR

5

140 s; 60 mGy 51 s; 46 mGy 40 s; 52 mGy

6

260

112 s; 11 mGy

7

Pxx-35-FS Pxx-50-FS Pxx-65-FS Pxx-80-FS

98 s; 42 mGy 35 s; 31 mGy 20 s; 29 mGy

8

140 120 100 80 60 40 20 0

0 60

80

120

Voxel size (μm)

160

240

60

80

120

160

240

Voxel size (μm)

Fig. 7. Contrast to noise ratio (left) and noise in water (right) for all the full-scan protocols. The total scan times and mean dose to rat are also reported for each protocol (only for the graph on the left).

D. Panetta et al. / Nuclear Instruments and Methods in Physics Research A 805 (2016) 135–144

reconstructed values can be seen throughout the 10 frames. More specifically the standard deviation of the mean ROI values measured for all frames was 5 HU in water, 3 HU for the C2 insert and 6 HU for the C5 insert. The average time interval between frames was 19.1 70.3 s, corresponding to a frame rate of 3.147 0.05 frames per minute (fpm) and an interscan delay of 11.770.3 s. The ability to perform dynamic scans with a relatively short interscan delay even without slip ring and with the PET detector mounted on the same gantry is also possible by alternating the rotation direction of the gantry (anticlockwise for even frames, clockwise for odd frames). For this reason, we have sought to assess the geometric reproducibility in dynamic mode by measuring the shift of a off-center detail of the geometric accuracy phantom placed at point (x,y,z) ¼(
3 mm, 31 mm, 27 mm) from the CFOV. Fig. 9 shows the reconstruction of one bead in three different time frames, demonstrating that no evident spatial shifts are present in the dynamic reconstruction. By measuring the center of mass (COM) coordinates of the bead in all frames and by calculating the shift of the bead COM with respect to the first frame (frame 0), we have found out that the root mean square (RMS) value of the bead shift was 70 μm (¼ 29% of the voxel size) while the maximum shift in the reconstructed image was 130 μm

4. Animal experiments 4.1. Ethical statement All images presented in this Section have been obtained during animal experiments that were conducted in accordance with the Italian Law D.L. 26/2014, implementation of the 2010/63/EU directive for animal protection in scientific experiments. All animals were handled by qualified veterinary personnel following internal standards of IFC-CNR, after approval of the experimental protocol by the Italian Ministry of Health.

1750

0.30

1500

0.25

1250

Dx Dy Dz Dr

0.20

1000

Shift (mm)

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(¼54% of the voxel size). The results are shown in Fig. 10 for all frames. This type of spatial shift can be attributed to the relatively high rotation speed of the gantry (30 deg/s) and consequently to the steep acceleration and deceleration ramps used for the alternate rotation, hence they are expected to decrease when higher resolution protocols (with slower rotation speed and smoother ramps) are used. Anyway, we consider subvoxel uncertainties of such magnitude negligible in practical applications.

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Frame midpoint (s) Fig. 10. Inter-scan spatial shift for the bead placed at (x,y,z) ¼ (1 mm, 31 mm, 27 mm) from the CFOV. The dashed line represents the voxel size.

Fig. 9. Left: transaxial slice of the phantom used for geometric reproducibility, acquired in dynamic modality with the protocol P240-80-FS (7.3 s scan time per frame). The cross mark indicate one bead placed at 31 mm from the CFOV, used to calculate the geometric reproducibility of the scanner in dynamic mode as explained in Section 2.9. Right: close-up of the bead image for three different time frames in transaxial and coronal view, as opened in AMIDE without additional spatial coregistration between frames; the cross mark is centered in the same (x,y,z) position for all time frames.

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Fig. 11. Example of images obtained in animal studies. (A) In vivo image of a mouse acquired with a modified version of the P80-50-FS protocol (scan time: 117 s, dose: 80 mGy, pixel size: 78.8 μm, slice thickness 236 μm, window min/max:
300 HU/400 HU). (B) In vivo image of a mouse knee (P60-80-FS). (C,D) Images of mice injected with iodinated contrast agent, acquired with the protocols P120-80-FS (C, in vivo) and P60-80-FS (D: post mortem). (E) Maximum intensity projection (MIP) of a whole body multi-bed scan of a rat (512 g) scanned over three bed positions with the protocol P160-80-FS (20 s of scan time per bed position). The overall scan length is 250 mm.

4.2. In vivo imaging experiments Examples of application of the IRIS CT scanner in preclinical imaging experiments with various scanning protocols, conducted at IFC-CNR, are shown in Fig. 11. We have omitted images obtained in PET/CT experiments because the emphasis of this paper is in CTonly studies. Fig. 11A shows a transaxial thoracic image of a mouse (37 g), acquired in vivo at 50 kV with a modified low-dose version of the P80-50-FS protocol, with total scan time of 117 s and a calculated mean dose of 80 mGy. The image shows the good discrimination of muscle, skin, lungs, white adipose tissues (WAT) and interscapular brown adipose tissue (BAT) allowing reliable quantification of lean mass and fat. Almost no ring artifacts are visible. In Fig. 11B, a volume rendering (VR) of a mouse knee obtained in vivo with the high-quality P60-80-FS protocol is shown.

Fig. 11C–E refer to contrast enhanced studies, obtained with continuous infusion of Iomeron 400 (Bracco, Italy) through a tail vein. Fig. 11C was obtained in vivo on a CB57/BL mouse after injection of 0.2 mL of CA over 2 minutes, 300 mg I/mL, using the P120-80-FS protocol (40 s scan time) starting 90 s after the beginning of the CA infusion. The main vessels and perfused tissues are well visible and distinguished from bones. In Fig. 11D, an ex-vivo image of a different mouse is shown, acquired with the P60-80-FS protocol. The multi-bed capability was also used for rat imaging: in Fig. 11E, a Wistar rat (512 g) was acquired with the P160-80-FS protocol over 3 bed positions, reaching the maximum allowed scanning length of the scanner (250 mm). The image was obtained after 5 minute of continuous infusion of CA, 200 mGI/mL, at the infusion rate of 24 mL/h. The maximum intensity projection (MIP) visualization emphasizes the renal excretion of the CA and the abdominal vasculature.

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5. Discussion and conclusions The performance evaluation of a novel commercial micro-CT scanner has been reported. Even though the scanner in use was installed as the integrated CT component of a multimodal PET/CT scanner, this study has validated it in terms of its capability for standalone CT studies, with regards to image quality, dynamic imaging capability and usefulness in practical preclinical experiments with mice and rats. Indeed, the overall performances of the IRIS CT were comparable, and in some case outperformed, those found in the literature for standalone preclinical CT scanners. The obtained spatial resolution at 10% of the MTF at the FOV ranged from 73 μm ( ¼6.9 lp/mm) at 60 μm voxel size and 80 kV to 234 μm ( ¼2.2 lp/mm) for the fastest modality (7.3 s); the maximum values of CTDI100,air were 165 mGy and 170 mGy for the highest quality protocol at 80 kV and 50 kV, respectively. This ranges of spatial resolution and doses appear completely adequate for in vivo imaging, with or without the use of contrast agents, as also evident from the animal images described in Section 4. Furthermore, these values compare well with the spatial resolution of other preclinical CT scanners: for instance, a spatial resolution of 6 lp/mm is reported for the Tomoscope scanner (CT Imaging, Erlangen, Germany) [30], with a CTDI100,air of 290 mGy for the high-quality protocol at 40 kV [26]. Even higher doses have been reported for the Skyscan 1178 scanner (Bruker, Belgium), for which the CTDI100 was 490 mGy inside a PMMA phantom of 27 mm of diameter (50 kV, 0.615 μA) for a 121 s long protocol at the voxel size of 83 μm. [31]. At the time of writing this article, we were unable to perform a direct measurement of the animal dose either in vivo or in phantom. In fact, the measurement of the CTDI using conventional pencil ion chambers designed for clinical CT scanners prevented us to do realistic dose measurements with animal-sized phantoms. In fact, the diameter of the pencil chamber (1 cm) is not negligible with respect to the typical size of a small animal, which makes impracticable the measurement of the CTDI at the phantom periphery needed for the calculation of e.g. the weighted CTDI (CTDIw), as already observed by Hupfer et al. Unpackaged thermoluminescent dosimeters (TLD) were also not available. Nevertheless, the CTDI100 measured free-in-air was a good metric that enabled us to calculated realistic values of the average animal dose through the application of precalculated conversion factors obtained under very similar conditions (i.e., geometry, filtration, X-ray spectrum) via Monte Carlo simulations [27]. The calculated average dose in mice reported for the different protocols implemented in our scanner range from 158 mGy (high quality, P60-80FS, 112 s scan time) to 12 mGy (fastest modality, P240-80-SS, 7.3 s scan time). These doses are well below the lethal dose of about 6– 9 Gy for mice and rats [32,33], and also considerably lower than doses that could have an impact on the results of a longitudinal experiments, in the order of about 1 Gy [34–35]; they are comparable to those reported in vivo by other investigators with different scanners [36] and appear suitable for repeated, longitudinal scans in accordance with the 3 R's principle. The wide FOV in both transaxial (90 mm) and axial (90– 110 mm) direction allows to image even heavy rats weighting more than 0.5 kg (see Fig. 11E) with a maximum scan length of 25 cm in just three bed positions. On the other hand, the performance of this scanner seems to be suboptimal for some studies requiring much better spatial resolution, such as e.g. quantitative measurement of cancellous bone microarchitecture in mice [37,38]. An attractive feature of the IRIS CT scanner was the possibility to perform dynamic scans (4D), useful for functional CT studies (e.g., perfusion imaging). The maximum frame rate was 3.14 fpm (i.e., 1 time point every 19 s) using the short-scan protocol with scan time of 7.3 s at 80 kV and 240 μm of voxel size. The

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main limiting factors of the dynamic frame rate are the mechanical design (no slip ring, with the PET detector mounted on the same rotating gantry of the CT scanner) and the data transfer speed between consecutive scans. The phantom measurements evidenced a good stability of both CT numbers and spatial reproducibility throughout a dynamic scan of 3 min (10 frames). Further studies on animals will be performed to assess the usefulness of this modality for perfusion imaging. In conclusion, the IRIS CT scanner provides spatial resolution at 10% of the MTF in the range of 73–234 μm, with mean animal dose in the range of 13–158 mGy. The geometric accuracy is within 0.1%; good linearity with respect to all X-ray tube voltage settings (35, 50, 65, 80 kV) was observed (R2 40.996). The minimum time interval between consecutive scans in dynamic modality is 19 s, corresponding to 3.1 fpm: this feature seems promising for perfusion imaging, in both standalone CT or in conjunction with PET.

Acknowledgments The authors would like to thank Dr. Claudio A. Traino, Director of the Health Physics Unit of the University Hospital of Santa Chiara in Pisa for his assistance during the CTDI measurements.

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