# The ultra high sensitivity blood counter: a compact, MRI-compatible, radioactivity counter for pharmacokinetic studies in µl volumes

L. Convert, Otman Sarrhini, Maxime Paillé, Nicolas Salem, Paul Charette, Roger Lecomte

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L. Convert, Otman Sarrhini, Maxime Paillé, Nicolas Salem, Paul Charette, et al. The ultra high sensitivity blood counter: a compact, MRI-compatible, radioactivity counter for pharmacokinetic studies in µl volumes. Biomedical Physics & Engineering Express, 2022, 8 (3), pp.035022. 10.1088/2057-1976/ac4c29. hal-03701266

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# The ultra high sensitivity blood counter: a compact, MRI-compatible, radioactivity counter for pharmacokinetic studies in µl volumes

To cite this article: Laurence Convert *et al* 2022 *Biomed. Phys. Eng. Express* **8** 035022

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PAPER

# The ultra high sensitivity blood counter: a compact, MRI-compatible, radioactivity counter for pharmacokinetic studies in µl volumes

Laurence Convert<fnref n="1" />,<fnref n="2" />,<fnref n="3" /> Ⓓ, Otman Sarrhini<fnref n="1" /> Ⓓ, Maxime Paillé<fnref n="1" />, Nicolas Salem<fnref n="4" />, Paul G Charette<fnref n="2" />,<fnref n="3" /> Ⓓ and Roger Lecomte<fnref n="1" /> Ⓓ

E-mail: Laurence mil Convert@USherbrooke.ca

Keywords: β-detector, blood sampling, MRI-compatible, pharmacokinetic modeling, small-animal PET, small animal SPECT

Abstract

Quantification of physiological parameters in preclinical pharmacokinetic studies based on nuclear imaging requires the monitoring of arterial radioactivity over time, known as the arterial input function (AIF). Continuous derivation of the AIF in rodent models is very challenging because of the limited blood volume available for sampling. To address this challenge, an Ultra High Sensitivity Blood Counter (UHS-BC) was developed. The device detects beta particles in real-time using silicon photodiodes, custom low-noise electronics, and 3D-printed plastic cartridges to hold standard catheters. Two prototypes were built and characterized in two facilities. Sensitivities up to 39% for ¹⁸F and 58% for ¹¹C-based positron emission tomography (PET) tracers were demonstrated. ⁹⁹mTc and ¹²⁵I based Single Photon Emission Computed Tomography (SPECT) tracers were detected with greater than 3% and 10% sensitivity, respectively, opening new applications in nuclear imaging and fundamental biology research. Measured energy spectra show all relevant peaks down to a minimum detectable energy of 20 keV. The UHS-BC was shown to be highly reliable, robust towards parasitic background radiation and electromagnetic interference in the PET or MRI environment. The UHS-BC provides reproducible results under various experimental conditions and was demonstrated to be stable over days of continuous operation. Animal experiments showed that the UHS-BC performs accurate AIF measurements using low detection volumes suitable for small animal models in PET, SPECT and PET/MRI investigations. This tool will help to reduce the time and number of animals required for pharmacokinetic studies, thus increasing the throughput of new drug development.

**1. Introduction**

Small animal nuclear imaging using Positron Emission Tomography (PET) and Single Photon Emission Computed Tomography (SPECT) is widely used to support the development of new drugs, diagnostic procedures, and therapies [1]. Traditionally used as stand-alone modalities for high sensitivity functional imaging, PET and SPECT are increasingly associated with high spatial resolution morphological imaging techniques like Computed Tomography (CT) and Magnetic Resonance Imaging (MRI) [1-3].

Nuclear imaging relies on injecting a radiolabeled
molecule, the radiotracer, in the living subject vascular

system and tracking it inside the body using dedicated
PET or SPECT scanners. Quantification of the radio-
tracer uptake is critical in many preclinical studies for
kinetic modeling of the fate of the radiolabeled mole-
cule [4]. Kinetic modeling requires the determination
of the radiotracer concentration in plasma as a func-
tion of time after injection, the so-called arterial input
function (AIF). Currently, manual sampling remains
the gold standard to measure the AIF, but it is very
challenging in small animal models due to the fast
kinetics and limited blood volume [5].

Various approaches have been developed over the
last decades to obtain the AIF in small animal models
[6]: image-derived input function (IDIF) [7, 8],

<footnote n="1">¹ Sherbrooke Molecular Imaging Centre of CRCHUS and Department of Nuclear Medicine and Radiobiology, Université de Sherbrooke, Sherbrooke, QC, Canada</footnote>

<footnote n="2">² Laboratoire Nanotechnologies Nanosystèmes (LN2) - CNRS IRL-3463, Université de Sherbrooke, Sherbrooke, QC, Canada</footnote>

<footnote n="3">³ Institut Interdisciplinaire d'Innovation Technologique (3IT), Université de Sherbrooke, Sherbrooke, QC, Canada</footnote>

<footnote n="4">⁴ Biogen, Cambridge, MA, United States of America</footnote>
<figure><img src="image_4.png" /><figcaption>**Figure 1.** Schematic view of the Ultra High Sensitivity Blood Counter (UHS-BC). The detection unit is shown semi-transparent and partially cut-out to display the major components: detection module, heat sinks, fans, and plastic enclosure).</figcaption></figure>

machine learning derived AIF [9], population mean [10], intravascular β probe [11, 12] and external blood radioactivity counters connected to an artery via a catheter [13–19]. The latter approach can provide a continuous input function independently of the imaged organ. However, current devices are cumbersome, use large blood volumes or have limited sensitivity, and are generally not suitable to perform measurements with multimodality scanners (PET/SPECT, PET/MRI, and SPECT/MRI).

In previous work, a compact catheter-based β blood counter [20] was developed for continuous measurement of the AIF in µl blood volumes. However, the detection efficiency was limited to a few percent for low β-energy PET isotopes, and the counter could not detect SPECT tracers. An integrated microfluidic chip-based β counter [21, 22] was also developed, which enabled the sampled blood to flow directly onto the detector surface to increase β particle collection efficiency. Though the system yielded a 2- to 10-fold improvement in sensitivity, depending on experimental conditions, it was fragile, built from custom components, which made it hardly reproducible with reliable characteristics, and therefore unsuitable for intensive use or industrial production.

In the present work, the catheter-based β blood counter concept was upgraded with a larger solid angle for detection of β particles and with low noise differential electronics allowing a lower detection threshold to reach the higher sensitivity of a chip-based system. Two prototypes (ProtoA and ProtoB) of this new Ultra High Sensitivity Blood Counter (UHS-BC) were built and characterized in two separate facilities. This article starts with a description of the UHS-BC. Then, the system characteristics and its validation are presented, as well as data corrections required for quantitative AIF measurement. Finally, typical animal studies are presented in combination with PET and MRI imaging.

## 2. Methods

### 2.1. System overview

The UHS-BC comprises three separate hardware units, a detection unit, a control unit, and a pump to draw blood through a catheter, all controlled through a USB link by software running under Windows (figure 1).

The detection module contains two large-area silicon p-i-n photodiodes (S3204–09, unsealed, 18 mm × 18 mm active area, Hamamatsu Corp.) for direct β detection. Unlike photomultiplier tubes, photodiodes are not affected by magnetic fields so that the detection module can be made fully MRI-compatible. The photodiodes are placed face-to-face, 2.4 mm apart, on each side of the catheter to maximize β particle collection efficiency (figure 2(a)). The resulting detection area is augmented compared to previous prototypes, increasing the detection solid angle and thus sensitivity. The photodiodes are shielded by 25 µm thick aluminum foils connected to the case ground to minimize electromagnetic interference (EMI), as well as for light and dust protection.

A custom-made low noise electronic board amplifies, shapes, and digitizes the electrical signals from the photodetectors for transmission to the control unit. Each detector is connected to a commercial charge-sensitive preamplifier (CR110, Cremat Inc.), respectively providing positive and negative pulses. Differential amplification is then used to remove common mode noise on both channels before further amplification and shaping. This essential step removes most of the EMI noise picked up from the fluidic channel by both detectors, allowing the detection threshold to be significantly reduced and the overall sensitivity to be improved. An aluminum and non-magnetic tungsten alloy case (HD 17, Mi-Tech Tungsten Metals,
<figure><img src="image_5.png" /><figcaption>**Figure 2.** (a) Detection module including a plastic cartridge for consistent positioning of the catheter between the two photodetectors, engineered tungsten slabs for radiation shielding, and an aluminum case for EMI shielding (top case, top shield, and top detector are partially cut-out and transparent to display the detection area). (b) Schematic view of the plastic cartridge for PE50 catheter with a secure lock of the catheter.</figcaption></figure>

Indianapolis, IN, USA) provides additional EMI and radiation shielding for the entire module (figure 2(a)).

A 3D printed plastic cartridge holds the catheter (Intramedic PE50 or PE10, Becton Dickinson) (figure 2(b)), exposing a 2.5 cm piece of tubing between the photodetectors (detection volume = 6.7 µl for PE50 and 1.6 µl for PE10). Finally, a 3D-printed plastic enclosure houses heat sinks and fans above the detection module to stabilize the temperature and avoid unwanted electrical contact with the EMI shielding case (figure 1). Standard brushless fans are compact and well-suited to use with PET or SPECT scanners. The operation close to an MRI scanner requires piezoelectric fans or a constant airflow over the heat sink.

The control unit comprises a low noise power supply (HAA15–0.8-AG, Bel Power Solutions, Santa Clara, CA, USA), an isolated control board for the pump, and a circuit board for digital signal acquisition and communication with the computer. As this unit does not have to be close to the scanner, its MRI compatibility is less stringent. The peristaltic miniature pump (P625/900.133, Instech Laboratories Inc., Plymouth Meeting, USA) is MRI-compatible and can accommodate 8–400 µlmin⁻¹ withdrawing speed depending on the tubing. The custom-made C# control software sets and monitors experiments with a user-friendly interface. It also provides data correction (See Data corrections section). For more flexibility, two acquisition modes are available in the software. The Normal Mode (NM) automatically sets the detection threshold to avoid counts due to electronic noise (<1 count s⁻¹ (CPS)). NM is intended for use with typical tracers. The Sensitive Mode (SM) has a slightly lower detection threshold allowing some background counts (~30 CPS) that can be subtracted from the data. SM is suitable for tracers with low energy emissions.

## 2.2. System calibration and validation

The UHS-BC (ProtoA) was calibrated and validated with energy spectra derived from the analog output. Energy calibration was performed by fitting the position of known energy peaks for three sealed sources: ¹³³Ba (γ 81 keV, x-ray 31 keV), ²⁴¹Am (γ 60 keV), ¹⁰⁹Cd (x-ray 22 keV, γ 88 keV). Only half of the detection module (half of the casing and only one detector) was used to be able to position the sealed source above the detector. All parts were placed inside an EMI and light-tight shielding box. The shaping amplifier output was connected to a multichannel analyzer (Trump 2K-W3, ORTEC, Oak Ridge, TN, USA), and energy spectra were recorded for 5 min. Using the resulting calibration factor, the noise level was calculated in keV from the threshold level giving less than 1 CPS.

For quality control, the detection unit was fully assembled (both detectors and all parts of the casing), and energy spectra were recorded for the most common PET (¹⁸F) and SPECT (⁹⁹mTc) radiotracers, as well as ¹²⁵I, injected inside a PE10 cartridge. A thin wall (20 µm) polyimide catheter, Poly120, (120–0.050, microLumen, FL, USA) was also tested with ⁹⁹mTc.

## 2.3. System characteristics

The performance of the UHS-BC was characterized according to the following operational parameters: detection efficiency, η (cps Bq⁻¹ or %) and detection limit, Amin (Bq µl⁻¹) for the most common isotopes, noise level due to parasitic radioactivity and EMI in the environment, and response stability over time. The measured responses for the different isotopes were individually fitted with high-order polynomials to model sensitivity and instrumentation-related static non-linearities (ex: dead time).

Polynomial model coefficients were determined from decay measurements made with both prototypes using initial activity concentrations of 35 MBq ml⁻¹ of
$^{18}$F and $^{11}$C (13 half-lives) and 80 MBq ml$^{-1}$ of $^{99m}$Tc (12 half-lives). Measurements were performed in triplicate for each catheter type and acquisition mode. To facilitate quantitative comparisons of UHS-BC system sensitivity for common PET and SPECT tracers, detection efficiency values were derived from the initial linear portions in the system responses with PE10 and PE50 catheters. This sensitivity is expressed as the slope of the curve $n$ (Bq ml$^{-1}$) = $f$ ($m$ (CPS ml$^{-1}$)) from $n$ = 0.04 to 8 MBq ml$^{-1}$ for PET tracers and $n$ = 1.5 to 10 MBq ml$^{-1}$ for $^{99m}$Tc.

Minimum detectable activity, $A_{min}$, was then derived from Currie equation [23], where $A_{min}$ can be interpreted as the minimum activity ensuring a false-positive readout less than 5% of the time. Sensitivity to $^{125}$I was computed from a 10 min acquisition of an 11 MBq ml$^{-1}$ source (single measurement for each acquisition mode and catheter type: PE50, PE10 & Poly120). The sensitivity was also assessed with Poly120 catheter for $^{99m}$Tc (single measurement, 17 MBq ml$^{-1}$, 5 min acquisition, SM & NM).

Background count rate due to parasitic radioactivity in the environment was assessed using a 43 MBq source of $^{18}$F and a 70 MBq source of $^{99m}$Tc placed 10 to 40 cm away from the detection unit in each direction (1 min acquisition, NM and SM). Noise due to EMI was evaluated by withdrawing (100 µl min$^{-1}$) saline solution from a vial placed inside a pre-clinical PET/CT scanner (5 min acquisition, NM & SM). Finally, response stability was assessed by comparing three $^{11}$C (10 to 20 h acquisition, SM, ProtoB) and six $^{99m}$Tc decay measurements (72 h acquisition, SM & NM, ProtoB).

## 2.4. Data corrections

Typical data corrections required to the AIF in pharmacokinetic models include radionuclide decay time, propagation delay, dead time, and dispersion. Decay time is corrected using the known half-lives of each isotope [24]. Propagation delay is corrected using the pump speed and the known fixed animal-to-detector distance. Dead time is corrected for by the polynomial model of the system response. For evaluation purposes, dead-time was assessed using paralyzable and non-paralyzable models [23] fitted with Origin 2018 on six radioactive decay measurements of $^{11}$C in different conditions (ProtoB). Decay measurements were performed on PE10 and PE50 catheters, with 70 to 850 MBq ml$^{-1}$ starting activity and for 15 to 59 half-lives. Finally, dispersion of the radioactivity in the catheter between the animal and the detector is corrected by fitting the AIF as presented in [22].

## 2.5. Animal handling

Animal experiments were performed following the recommendations of the Canadian Council on Animal Care. They were approved by the Université de Sherbrooke in-house Ethics Committee for Animal

Experiments under Protocol 199–18R. Animals had free access to food and water before the experiments. Animals were anesthetized with isoflurane (2% + 1.5 l min$^{-1}$ O$_2$) and cannulated in the caudal vein for injection (prefilled with heparinized saline, 0.9%, 50 U ml$^{-1}$). For blood withdrawal, animals were cannulated in the caudal or carotid artery. The animal temperature was regulated, and the heartbeat and breathing were monitored to ensure that physiological conditions were maintained stable during the scan.

## 2.6. Animal experiments

A first experiment was performed to measure the $^{18}$F-FDG influx constant in a mouse (ProtoB, SM). The anesthetized mouse (BALB/c, male, 20.8 g, PE10 withdrawing cannula in carotid) was placed in a LabPET8 scanner [25] with the heart centered in the field of view. The UHS-BC was placed on a table in front of the scanner (detector to animal distance, DAD = 60 cm). Both PET and counter acquisition were started 30 s before injection of 13.6 MBq of $^{18}$F-FDG (0.2 ml @ 0.4 ml min$^{-1}$, no flush, 20-cm long PE10 catheter). The pump speed was set to 50 µl min$^{-1}$ up to 2.5 min, then 10 µl min$^{-1}$ for 17.5 min (total 300 µl including blood volume drawn pre-injection).

A second experiment was performed to measure the myocardial blood flow of a rat (ProtoB, SM). A male Wistar rat (403 g, PE50 withdrawing cannula in caudal artery) was imaged with the LabPET8 scanner with the heart in the center of the field of view (DAD = 60 cm). Both PET and counter acquisition were started 30 s before injection of 37.2 MBq of $^{11}$C-Acetate (0.3 ml @ 0.9 ml min$^{-1}$, no flush, 20-cm long PE50 catheter). The pump speed was set to 200 µl min$^{-1}$ for 2.5 min, then 50 µl min$^{-1}$ for 8 min (total 900 µl including pre-injection volume).

A third experiment was performed to assess the influence of an MRI scanner on the UHS-BC (ProtoA, NM). An input function was measured with the UHS-BC while acquiring an MRI image of the animal. An anesthetized female Sprague & Dawley rat (268 g, PE50 withdrawing cannula in carotid) was imaged in the dorsal recumbency in a small animal 7T MRI system (Unity INOVA, Varian Inc., CA, USA) equipped with 210/105 mm Magnex gradient coils and a 63 mm diameter volume coil. The detection module of the UHS-BC was placed on a table in front of the MRI system with compressed air cooling. The main unit and computer were placed on another table around 1 m from the MRI system. The DAD was 105 cm, which is longer than the typical length for experiments on PET scanners because of the long bore size of the MRI system. Blood was initially withdrawn through the PE50 catheter at 50 µl min$^{-1}$ for background measurements, the first 2 min before the MRI acquisition and 2 min while acquiring the image. After the first 4 min of background acquisition, the animal was injected
<figure><img src="image_6.png" /><figcaption>**Figure 3.** System validation. Energy spectra recorded with the UHS-BC for ¹⁸F (4 MBq ml⁻¹, 10 min acquisition), ⁹⁹mTc (17 MBq ml⁻¹, 30 min acquisition) and ¹²⁵I (8 MBq ml⁻¹, 15 min acquisition) using a cartridge with PE10 or Poly120 catheter.</figcaption></figure>

with 18.6 MBq of ¹⁸F-FDG (0.4 ml @ 0.6 ml min⁻¹, 102-cm long PE10 catheter, caudal cannula), and the withdrawing speed was changed to 250 µl min⁻¹ for 5 min, then lowered to 50 µl min⁻¹ for 10 min (total blood withdrawn = 1.95 ml).

A last experiment was performed to assess the influence of the UHS-BC (ProtoA, NM) on the MRI images. An anesthetized male Sprague & Dawley rat (394 g, PE50 withdrawing cannula in carotid) was imaged in the same MRI scanner using the same coil as in the previous experiment. Images of the rat brain were recorded without the UHS-BC (Without), with the UHS-BC turned off (Off) and with the UHS-BC and its pumping unit turned on (On). During this last imaging sequence, blood was withdrawn at 50 µl min⁻¹ for 15 min (total blood withdrawn = 0.75 ml). No radioactivity was injected.

## 2.7. Image and data analysis

PET images were reconstructed using a maximum-likelihood expectation-maximization algorithm with 20 iterations and using the following dynamic sequence: 1 × 30 s (before radiotracer injection), 18 × 5 s, 4 × 30 s, 4 × 60 s, and 5 × 150 s (32 frames) for the ¹⁸F-FDG experiments and 1 × 30 s, 18 × 5 s, 5 × 30 s and 6 × 60 s (30 frames) for the ¹¹C-Acetate experiment. Cross-calibration of PET images was performed using a cylindrical phantom loaded with a known activity concentration to convert image counts/s/voxel values to MBq ml⁻¹ values.

Myocardial tissue time-activity curves (mTAC) and IDIF were obtained by manually drawing regions-of-interest (ROIs) on the myocardium and left ventricle cavity, respectively, on the last frame for ¹⁸F-FDG and at ~1.5 min post-injection for the ¹¹C-Acetate experiment, using the PMOD software (PMOD Technologies Ltd, Zurich, Switzerland, version 3.9). Mouse IDIF was further corrected for partial volume effect and spillover of the activity from the myocardium using the hybrid

method of Tantawy and Peterson [26]. Briefly, to construct the 'hybrid' IDIF (HIDIF), the initial part of the IDIF, containing the peak concentration of ¹⁸F-FDG in the left ventricle blood pool (0–1.6 min), was corrected for partial volume (RC = 0.7) and then joined (linear interpolation, 1.6–2.5 min) to the liver TAC immediately after its peak (>2.5 min). For the rat experiment, the ROI was larger, and a value of 0.9 was used for RC to correct the IDIF for partial volume effect. RC coefficients were derived from dedicated imaging experiments with the scanner on phantoms with rods of different sizes [25]. The ¹⁸F-FDG influx constant $K_i$ was computed using the Patlak graphical method [27], and MBF was computed using a 3-compartment kinetic model [28].

For the first MRI image, a dynamic acquisition consisting of 10 complete T1-weighted images of the rat brain was performed using the following parameters: TR/TE: 180/3.44 ms, FOV: 3.5 cm × 3.5 cm, matrix: 256 × 256, flip angle: 30°, NA: 4, 10 slices of 2 mm, 46 s per slice. For the MRI images in the last experiment, three dynamic acquisitions consisting of 5 complete T1-weighted images of the rat brain were performed using the following parameters: TR/TE: 180/3.38 ms, FOV: 4.2 cm × 4.2 cm, matrix: 256 × 256, flip angle: 30°, NA: 4, 15 slices of 1 mm, 3 min acquisition. Each image set was corrected for motion (first frame as reference) and registered with the second image set (Off) using SPM8. Signal to noise ratio was approximated on each image set (Without, Off, and On) by taking the ratio of the mean signal in a tissue region of interest to the mean signal in a background region.

# 3. Results

## 3.1. System calibration and validation

Energy spectra of sealed sources showed well-defined peaks with calibration of 18.6 ± 0.3 mV keV⁻¹ and a noise level of 12 keV. Energy spectra of liquid sources (figure 3) showed expected shapes for each isotope: a
continuous spectrum for $\beta^+$ emitted by $^{18}$F and well-resolved X and $\gamma$ rays energy peaks for $^{99m}$Tc and $^{125}$I. The 18–22 keV x-ray peak of $^{99m}$Tc is visible but partially overshadowed by the electronic noise. Indeed, the noise level of the whole system with two detectors was higher than in the previous experiment with only one detector. The resulting noise level is ~20 keV, yet still masking less than 4% of the full $^{18}$F spectrum. The detection threshold in NM was set to 20 keV, just above the electronic noise level. In SM, the threshold is slightly lowered to 17 keV to enhance sensitivity (see table 1) but allows some electronic noise background that can be subtracted by the software.

The peak at the end of the $^{18}$F spectrum in figure 3 is due to saturation of signals exceeding the dynamic range, but it did not prevent events from being recorded. Unlike typical beta spectra from $^{18}$F, there is no peaking around one-third to one-half the maximum beta emission energy of 634 keV. This behavior is attributed to Compton interactions of the 511 keV photons in the Si photodiodes and Compton backscattered events from the 511 keV photons interacting in the shielding surrounding the detectors. Compton and backscattered $\gamma$ were visible in $^{99m}$Tc spectra, but $^{99m}$Tc conversion electrons (emitted at 119–140 keV) more easily cross the thin Poly120 catheter walls. The ~60 keV peak in the $^{99m}$Tc spectra results from the characteristic tungsten x-rays excited by primary gamma rays undergoing a photoelectric interaction in the detection module shielding. The $^{125}$I spectrum shows two energy peaks. The first one is attributed to X and $\gamma$ rays at 27 to 35 keV (mainly ~27 keV). The second one at ~55 keV results mainly from an accidental sum of 27 keV x-rays.

## 3.2. System characteristics

Detection efficiency, $\eta$, for the two most common PET tracers ($^{18}$F, $^{11}$C) ranges from 28 to 58% (table 1). This represents a factor $R_1 \approx 2$ to 4.5 times better than a previous catheter-based counter [20], thanks to the use of unsealed diodes (no epoxy protection layer) and the optimization of the electronics and mechanical designs. The sensitivity of the UHS-BC is comparable ($R_2 \approx 0.6–1.2$) to a previous chip-based microfluidic system [22], which had only one detector but a very thin channel wall (10 $\mu$m).

Detection efficiency for the most common SPECT isotope ($^{99m}$Tc) was low because most of its low energy conversion electrons are absorbed by the catheter wall, the x-ray has a low emission probability, and the $\gamma$-ray has a small interaction probability in the silicon detector. Detection efficiency of $^{99m}$Tc was significantly improved with the thin wall Poly120 catheter compared to PE50 and PE10 catheters due to conversion electrons reaching the detectors. The resulting sensitivities are 4.4% in SM and 3.3% in NM, an increase of

30%. The detection efficiency for $^{125}$I was around 10%.

Minimum detectable activity ($A_{min}$) ranged from 2 to 47 Bq $\mu l^{-1}$ for PET tracers and from 39 to 531 Bq $\mu l^{-1}$ for the $^{99m}$Tc SPECT tracer under the various experimental conditions. These ranges of activities are suitable for typical pharmacokinetic studies in nuclear imaging [14]. As for $^{125}$I, a minimum detectable activity down to 10 Bq $\mu l^{-1}$ can be measured. Despite the higher detection efficiency, the minimum detection limit is higher due to the background noise count rate allowed in the SM mode.

The UHS-BC did not record any event with the $^{99m}$Tc external source in NM except when it was placed 10 cm in front of the plastic cartridge slot (figure 4). In SM, the mean background counts above the electronic noise were slightly higher, not exceeding ~1.5 CPS MBq$^{-1}$ for a source further away than 10 cm from the counter. As expected with its higher $\gamma$ energy, the background counts were higher with the $^{18}$F source, reaching ~25 CPS MBq$^{-1}$ in the worst case of a source at 10 cm in front of the cartridge slot (figure 4).

In typical animal studies, the detection unit is usually placed on a table in front of the scanner while the animal is on the bed inside the detection ring and the injection pump on another table. The resulting distance between the external radioactive sources (animal and injection syringe) and the unit is usually over 40 cm, leading to a negligible background count rate even for PET tracers. The radiation shielding is then excellent for SPECT tracers and acceptable for PET tracers. As a good practice, it is recommended to avoid placing the animal or the injection syringe in front of or behind the detection unit and to always record the background counts before injection. The acceptable count rate before injection depends on the experiment parameters (isotope and amount of radioactivity). Still, it is usually close to the electronic noise count rate (no counts in NM and a few tens of CPS in SM).

The UHS-BC did not record any significant counts due to EMI pickup with saline solution flowing from inside a PET scanner to the counter. Finally, the UHS-BC showed high stability over time: a 2 CPS variation in mean electronic noise was recorded over 14 h without activity. The last 30 min of three $^{11}$C acquisitions showed a tiny variation of the mean electronic noise level (19 ± 0.2 CPS, *ProtoB*). Owing to the temperature control of the detection unit, the signal baseline and response to events detection were stable across all experiments, including the decay of $^{99m}$Tc sources continuously recorded for up to 72 h.

The UHS-BC dead time was found to be paralyzable because the measured versus expected count rate curve decreased at very high-count rates (> 500 kCPS). The dead time was estimated at $\tau = 3.6 \pm 0.3$ $\mu$s with linearity better than 1% up to 18 kCPS and 10% up to 37 kCPS. No correction would normally be required for typical pharmacokinetic studies (typical
<table><caption>**Table 1.** Detection efficiency, $\eta$ (cps Bq$^{-1}$ expressed in %), and detection limit, $A_{min}$ (Bq $\mu l^{-1}$), for the UHS-BC compared to previous catheter-based (relative efficiency factor $R_1$) [20] and microfluidic chip-based ($R_2$) counters [22]. Mean value of triplicate measurements for PET tracers and $^{99m}$Tc and single measurement for $^{125}$I, both prototypes. Triplicate measurements are indicated with their mean deviation.</caption><thead><tr><th rowspan="2">Isotope</th><th rowspan="2">E (keV)<sup>a</sup></th><th rowspan="2">β/CE Emission probability (%)<sup>a</sup></th><th rowspan="2">Mode</th><th colspan="4">PE50 (6.7 μl)</th><th colspan="4">PE10 (1.6 μl)</th></tr><tr><th>η (%)</th><th>A<sub>min</sub> (Bq μl<sup>-1</sup>)</th><th>R<sub>1</sub></th><th>R<sub>2</sub></th><th>η (%)</th><th>A<sub>min</sub> (Bq μl<sup>-1</sup>)</th><th>R<sub>1</sub></th><th>R<sub>2</sub></th></tr></thead><tbody><tr><td rowspan="2"><sup>18</sup>F</td><td rowspan="2">β0–634</td><td rowspan="2">96.7</td><td>SM</td><td>31.6 ± 0.4</td><td>14</td><td>4.5</td><td>0.8</td><td>39.4 ± 0.1</td><td>47</td><td>4.1</td><td>1.0</td></tr><tr><td>NM</td><td>27.9 ± 0.7</td><td>4</td><td>3.9</td><td>0.7</td><td>35.1 ± 0.8</td><td>12</td><td>3.7</td><td>0.9</td></tr><tr><td rowspan="2"><sup>11</sup>C</td><td rowspan="2">β0–961</td><td rowspan="2">99.8</td><td>SM</td><td>52.9 ± 1.4</td><td>8</td><td>2.5</td><td>1.1</td><td>57.7 ± 1.5</td><td>32</td><td>2.4</td><td>1.2</td></tr><tr><td>NM</td><td>43.8 ± 1.2</td><td>2</td><td>2.1</td><td>0.9</td><td>48.1 ± 0.2</td><td>9</td><td>2.0</td><td>1.0</td></tr><tr><td rowspan="2"><sup>99m</sup>Tc</td><td rowspan="2">EC 115–140 × 18–22 γ 141</td><td rowspan="2">227.7<sup>b 88.5</sup></td><td>SM</td><td>3.3 ± 0.2</td><td>130</td><td>NA</td><td>0.8</td><td>3.5 ± 0.1</td><td>531</td><td>NA</td><td>0.8</td></tr><tr><td>NM</td><td>2.5 ± 0.3</td><td>39</td><td>NA</td><td>0.6</td><td>2.6 ± 0.2</td><td>164</td><td>NA</td><td>0.6</td></tr><tr><td rowspan="2"><sup>125</sup>I</td><td rowspan="2">X 27–32 γ 35</td><td rowspan="2">137.9<sup>b 6.6</sup></td><td>SM</td><td>10.2</td><td>42</td><td>NA</td><td>NA</td><td>10.9</td><td>168</td><td>NA</td><td>NA</td></tr><tr><td>NM</td><td>9.7</td><td>10</td><td>NA</td><td>NA</td><td>10.0</td><td>42</td><td>NA</td><td>NA</td></tr></tbody></table>

<footnote><sup>a</sup> From Laboratoire National Henri Becquerel, n.d. [24].</footnote>

<footnote><sup>b</sup> Sum of K x-rays.</footnote>
<figure><img src="image_7.png" /></figure>

<figure><img src="image_8.png" /></figure>

20 kBq µl⁻¹ blood concentration in mice gives ~18 kCPS for ¹¹C).

### 3.2.1. Animal experiments

The data obtained from the mouse experiment are illustrated in figure 5, together with the ROI used to compute mTAC and IDIF from the left ventricle. AIF<sub>BC</sub> is the original AIF recorded by the UHS-BC corrected for decay and propagation delay. AIF<sub>U</sub> is AIF<sub>BC</sub> further corrected for dispersion. The hybrid IDIF (HIDIF), built from the initial (0–1.6 min) IDIF from the left ventricular blood pool corrected for partial volume concatenated to the liver TAC

(>2.3 min), shows a good correlation with AIF<sub>U</sub>. The resulting <sup>18</sup>F-FDG influx constant K<sub>i</sub> is 0.20 ml min⁻¹ g⁻¹ using AIF<sub>U</sub> and 0.22 ml min⁻¹ g⁻¹ using HIDIF, which is within the statistical range of reported values of 0.14 ± 0.06 ml min⁻¹ g⁻¹ [29] for the analysis from the AIF<sub>U</sub>.

The data obtained from the rat experiment are displayed in figure 6, together with the ROI used to compute mTAC and IDIF. AIF<sub>U</sub> and IDIF show a good correlation in the peak region and the last portion of the curve. Spillover from myocardial uptake likely explains the differences between the two curves from minutes 1 to 6. Oscillations of the IDIF in the 1.2 to
<figure><img src="image_9.png" /><figcaption>**Figure 6.** (a) AIF~BC~, AIF~U~, IDIF, and mTAC obtained after bolus injection of $^{11}$C-acetate (starting at 30 s, 37.2 MBq in 0.3 ml @ 0.9 ml min$^{-1}$, no flush) in a 403 g Wistar rat. AIF~BC~ was obtained at 200 µl min$^{-1}$ for 2.5 min, then 50 µl min$^{-1}$ (DAD = 60 cm, PE50). (b) ROI used for mTAC and IDIF.</figcaption></figure>

<figure><img src="image_10.png" /><figcaption>**Figure 7.** (a) AIF~BC~ obtained during an MRI scan, after bolus injection of $^{18}$F-FDG (at 4.0 min, 18.6 MBq in 0.4 ml @ 0.6 ml min$^{-1}$, no flush) in a 268 g Sprague & Dawley rat (DAD = 105 cm, PE50). AIF~BC~ is corrected for radioactive decay but not for propagation delay. (b) AIF~BC~ further corrected for the change in withdrawing speed at 9 min and then for dispersion (AIF~U~). (c) MRI image of the rat head acquired during blood withdrawal with the UHS-BC.</figcaption></figure>

2.2 min timeframe can be partly attributed to statistical error due to the short image time frames (5 s). Still, they may also be caused by physiological fluctuations of the blood radioactivity resulting from the second pass of the bolus injection. Such oscillations are often observed in small animal experiments. The extracted MBF was 3.6 ml g⁻¹ min⁻¹ using AIF_U and 3.7 ml g⁻¹ min⁻¹ using IDIF, which is comparable to other experiments performed with a similar protocol

in which reported MBF = 3.5 ± 1.0 ml g⁻¹ min⁻¹ [30].

The AIF~BC~ obtained from the animal injected with radioactivity while it was imaged in the MRI system is presented in figure 7(a). The background acquisition (first 4 min) shows that the MRI system has no influence on the signal, and the AIF~BC~ shows the standard peak due to bolus injection, demonstrating the compatibility of UHS-BC with MRI systems. The plateau
<figure><img src="image_11.png" /><figcaption>**Figure 8.** MRI images (first image, slice 7) of the rat brain taken (a) without the UHS-BC, (b) with the UHS-BC turned off next the MRI scanner and (c) with the UHS-BC turned on next the MRI scanner. ROIs used to compute the signal-to-noise ratio are indicated on each set with the calculated values.</figcaption></figure>

between 9 and 14.5 min is an artifact due to the change in withdrawing speed and the very long catheter. At 9 min after acquisition start, the pump speed was lowered from 250 µl min⁻¹ to 50 µl min⁻¹, changing the propagation delay within the 105 cm long catheter from 1.1 to 5.6 min. The blood monitored during the following 5.6 min was actually withdrawn from the animal at 250 µl min⁻¹, translating into a 1.1 min delay. Data in this 5.6 min time window were then resampled in a 1.1 min time window. Such artifact is usually negligible with smaller DAD. The entire curve was then corrected for dispersion to obtain the AIFU shown in figure 7(b).

The MRI images recorded during the experiment presented typical artifacts like wrap-around (due to the field of view smaller than the rat) or motion artifacts (due to blood flow inside arteries), but no spikes or visible metallic interferences.

The MRI images recorded during the last experiment are presented in figure 8. The three image sets were visually analyzed, and no spike or metallic interference was visible. The signal-to-noise ratio was not significantly different in each data set (9.40 ± 0.12 for set *Without*, 9.44 ± 0.14 for set *Off*, and 9.68 ± 0.13 for set *On*). The UHS-BC had no visible influence on the recorded MRI images. As is well known, MRI images are susceptible to the presence of metallic objects inside the antenna but not outside. Metallic parts outside the antenna will distort the field lines, but this can be easily corrected during pre-acquisition adjustments. The hazards with metallic accessories outside the antenna lie more in the risk of objects being thrown than in the degradation of image quality. As the detection unit and the pump are made of non-magnetic materials, they can be placed in front of the MRI system, which reduces the distance between the animal and the detection unit. The control unit and the computer have to be further away. The specific distance is very dependent on the MRI system used. In the present study, the Unity INOVA is actively shielded, so the magnetic field is below the 5-gauss safety limit one meter away from the MRI scanner.

## 4. Discussion

The present work describes an optimized catheter-based β blood counter for PET and SPECT pharmacokinetic studies with improved stability, reliability, and sensitivity. The UHS-BC has a better sensitivity than any small animal blood counters currently found in the literature: 39% compared to 1%–33% for ¹⁸F (table 2). A higher sensitivity helps reduce the injected dose [15] and/or the blood volume required for detection. Whereas a higher detection efficiency can be reached in SM mode by lowering the noise threshold, this does not translate into a lower detection limit than the NM mode. The higher noise floor of the SM mode significantly influences the calculation of this limit. There is most likely an optimal point of the noise threshold to obtain the best detection limit for each radioisotope. This optimization was not part of the present work.

The UHS-BC also has a smaller detection volume than most small animal blood counters (table 2), whereas still being compatible with large animals and humans. This low volume has several advantages: lower animal blood loss, higher temporal resolution, lower dispersion, sharper AIF peak, and, as a result, more accurate pharmacokinetic modeling. Furthermore, in addition to be compatible with MRI systems, the UHS-BC can detect SPECT tracers such as ⁹⁹mTc and ¹²⁵I, which only one other non-MRI compatible blood counter has been shown to do [13], hence opening new opportunities for pharmacokinetic studies in microSPECT imaging [38, 39] and dual-modality imaging.

Whereas there is little gain in performance relative to previous microfluidic-based prototypes [21, 22] in terms of sensitivity and detection volume, the UHS-BC uses standard catheters, is MRI-compatible, and can be built from off-the-shelf components and inexpensive 3D-printed cartridges as catheter holders, making the concept more suitable for intensive use or as a pre-commercial prototype.

Several detection principles are found in the literature for small animal blood counters (table 2). In the present work, direct β particle detection with p-i-n
<table><caption>**Table 2.** Characteristics of the UHS-BC versus other blood counters in literature (non-exhaustive).</caption><thead><tr><th>Reference</th><th>Detection principle</th><th>Catheter wall thickness (μm)<sup>b</sup></th><th>η for <sup>18</sup>F (%)</th><th>Detection volume (μl)<sup>b</sup></th><th>η for <sup>99m</sup>Tc (%)</th><th>MRI comp.</th></tr></thead><tbody><tr><td>This work</td><td>β with p-i-n photodiodes</td><td>165</td><td>39.4</td><td>1.6</td><td>3.4</td><td>Yes</td></tr><tr><td>Microfluidic β counter [22]</td><td>β with p-i-n photodiodes</td><td>10</td><td>39</td><td>0.66</td><td>4</td><td>—</td></tr><tr><td>Catheter-based β blood counter [20]</td><td>β with p-i-n photodiodes</td><td>165</td><td>7</td><td>1.85</td><td>—</td><td>—</td></tr><tr><td>ART [16]</td><td>β with p-i-n photodiodes</td><td>25</td><td>33</td><td>23.6</td><td>—</td><td>—</td></tr><tr><td>β microprobe [31]</td><td>β with Bicron/PMT</td><td>0</td><td>3–9</td><td>120 (blood + tissue)</td><td>—</td><td>Yes<sup>a</sup></td></tr><tr><td>Positron detector [19]</td><td>β with GSO/PMT</td><td>165</td><td>8.6</td><td>1.2</td><td>—</td><td>No<sup>c</sup></td></tr><tr><td>Multi-wire propor. Counter [18]</td><td>β with argon methane gas</td><td>250</td><td>4</td><td>15.3 (rat)</td><td>—</td><td>No<sup>c</sup></td></tr><tr><td>Swisstrace [32]</td><td>γ coinc. with glass fiber/PMT</td><td>Non- dependent</td><td>3</td><td>6</td><td>—</td><td>Yes</td></tr><tr><td>MR-Compatible BS [15]</td><td>γ coincidence with LSO/APD</td><td>Non- dependent</td><td>23</td><td>6</td><td>—</td><td>Yes</td></tr><tr><td>ABSS [33]</td><td>γ coincidence with LSO/APD</td><td>Non- dependent</td><td>18.5</td><td>13.2 (human)</td><td>—</td><td>Yes</td></tr><tr><td>Online blood sampler [17]</td><td>γ coincidence with LYSO/PMT</td><td>Non- dependent</td><td>7</td><td>33.4 (rat)</td><td>—</td><td>No<sup>c</sup></td></tr><tr><td>GSO detector assembly [13]</td><td>γ coinc/single with GSO/PMT</td><td>Non- dependent</td><td>6.2</td><td>5.3 (rat)</td><td>70</td><td>No<sup>c</sup></td></tr><tr><td>CZT-based blood counter [34]</td><td>γ coinc/single CZT detector</td><td>Non- dependent</td><td>12.5</td><td>56.9 (human)</td><td>—</td><td>No<sup>c</sup></td></tr><tr><td>PBS-101 [35]</td><td>γ single with BGO/PMT</td><td>Non- dependent</td><td>20/70</td><td>47.1 (human)</td><td>—</td><td>No<sup>c</sup></td></tr><tr><td>Pico-count flow-through d. [36]</td><td>γ coincidence with BGO/PMT</td><td>Non- dependent</td><td>6.9</td><td>110 (human)</td><td>—</td><td>No<sup>c</sup></td></tr></tbody></table>

<footnote>a [37].</footnote>

<footnote>b The detection volume and catheter wall thickness are given for mice (mainly PE10 catheter) unless otherwise noted.</footnote>

<footnote>c Data not included in the reference but intrinsically PMT are not compatible with MRI systems unless they are connected to the scintillator via an optical fiber as in [31, 32].</footnote>
photodiodes has been used instead of indirect β or 511 keV annihilation photon detection. Direct β detection allows for more compact radiation shielding due to low sensitivity to γ rays, more compact detectors (no scintillator), and more straightforward electronic (no coincidence detection). Direct β detection also avoids the efficiency losses due to conversion to scintillation photons and light collection in scintillation detectors [15, 40].

The major drawbacks of β detection are the energy loss in intermediate materials before reaching the detector (e.g., blood, catheter wall, detector protective layer or foil) and the need for calibration with each radioisotope. However, unlike previous designs, the UHS-BC may not require daily detection threshold adjustment and frequent recalibrations. Nevertheless, a calibration should initially be done on-site for each isotope to account for potential environment-specific variables. For example, we measured a 12% difference in the UHS-BC FDG calibration factor for ¹⁸F between the two facilities where the same device (ProtoA) was used.

Polyimide catheters, having a thinner wall thickness than PE10 and PE50, have already been used to increase the detection efficiency [16] and were tested in the present work (energy spectra, Poly120). Such catheters are more rigid and less suitable than PE tubing for animal cannulation and installation in plastic cartridges. However, due to their increased detection efficiency, they can be a useful alternative for ⁹⁹mTc and ¹²⁵I detection.

The UHS-BC is expected to detect any isotope emitting CE, β⁺ or β⁻ particles with $Eβ_{\text{mean}} > 100$ keV and any X- or γ-rays with energy between 20 and 200 keV. Higher energy photons will have a minimal interaction probability in the thin Si detector. The device could also be used in other applications than PET and SPECT pharmacokinetic studies, such as metabolic radiotherapy [41] or basic research in physiology/pharmacology using ¹²⁵I labeled molecules [42–44]. Other potential applications are capillary electrophoresis and High-Performance Liquid Chromatography (HPLC) of radiolabeled compounds where a radiation counter can be coupled to those instruments for quality assessment in the development and production process of radiopharmaceuticals for preclinical and clinical imaging.

The optimized counter uses 3D printed plastic cartridges to hold the catheter. These cartridges have the advantage of being cheap and easy to adapt to any catheter with the desired detection volume. However, the tension applied to the catheter manually placed inside the cartridge may result in slight differences in detection length, leading to minor differences in detection volume and detection efficiency. Even if the overall mean deviation is low (between 0.1% and 1.5% for values in table 2), the UHS-BC would benefit from a disposable plastic cartridge with a microfluidic channel. Indeed, three efficiency measurements with the

same cartridge and the same catheter just emptied, rinsed, and filled again gave a very low mean deviation (0.4%, PE10, ¹¹C, SM, ProtoB), whereas triplicate measurements with different cartridges and catheters gave a much higher mean deviation (1.5%, PE10, ¹¹C, SM, ProtoA).

In addition to a more precise detection volume, a microfluidic cartridge could benefit from thinner walls (below 50 µm) and a more advantageous shape factor (flat rectangular rather than round) [22]. The overall detection efficiency could then be increased by up to a factor of 4, especially for isotopes emitting low-energy particles.

As with any continuous blood counters for small animals, the withdrawing speed with the UHS-BC is a compromise between (1) low speed to limit the total blood volume withdrawn while maintaining physiological conditions and (2) high speed to limit the distortion of the AIF inside the catheter due to dispersion. Even if dispersion can be corrected, it is desirable to keep the distortion as low as possible. A common way to solve this problem is to withdraw blood at high speed in the first few minutes corresponding to the peak of the AIF and then to lower the speed for the remaining of the experiment where the AIF is more stable. However, as demonstrated in the rat experiment with MRI, this change in speed introduces an artifact in the AIF, especially for very long catheters. The correction applied in this work is simple and could benefit from further improvement. Since shorter catheters were used in the PET experiments, the delayed measurement artifact is much less important, and no correction was applied using this simplistic approach.

Experiments with mice using a continuous blood counter are particularly challenging regarding the total blood volume withdrawn. In the present work, the experiment with FDG on a mouse (figure 5) was limited to 20 min, and the experimental parameters were chosen accordingly to achieve the best possible temporal resolution for demonstration purposes. For the longer acquisition times typical of pharmacokinetic studies with ¹⁸F-labeled tracers, it is always possible to lower the withdrawing speed or to perform an arteriovenous shunt [32]. However, it is worth noting that withdrawing speed below 8 µl min⁻¹ should be avoided to prevent blood coagulation in the catheter. In addition to the fact that the pump speed is no more linear with the voltage setpoint, experiments with the embedded pump or syringe pumps showed that blood clotting often occurs in the catheter at such low speed.

The choice of the cannulation site also implies a compromise. Cannulation in the caudal artery is non-lethal but hard to perform on mice or small rats, and the blood flow is sometimes limited. Cannulation in the femoral or carotid artery is terminal since the animal cannot be wakened at the end of the experiment. Still, it is easier to perform in small rodents, and the blood flows more easily, especially for the carotid.
## 5. Conclusion

This paper presents the UHS-BC, an optimized blood radioactivity counter based on β-particle detection through a catheter to measure the AIF in pharmacokinetic studies with small animal models in PET and SPECT. The compact system occupies a small footprint next to the scanner, is robust towards noise from parasitic radioactivity and electromagnetic interference, making it compatible with MRI systems for multimodality PET/MR and SPECT/MR kinetic imaging studies. The UHS-BC achieves the highest reported detection efficiency in a μl range detection volume, reaching 39% and 58% for ¹⁸F and ¹¹C PET radioisotopes and over 3% and 10% for ⁹⁹mTc and ¹²⁵I SPECT radioisotopes with standard PE10 catheters. The system is mostly built from off-the-shelf parts and uses affordable 3D printed plastic catheter cartridges for highly reproducible results. It was demonstrated to be stable over days of continuous acquisition under various experimental conditions. The UHS-BC will contribute to increase the throughput of pharmacokinetic studies in drug discovery by reducing the time and number of animals required and by avoiding operator radiation exposure. Various scenarios are being considered to make the device available to the scientific community.

## Acknowledgments

The authors would like to thank Jean-Nicolas Melançon-Edmond (Institut Interdisciplinaire d'Innovation Technologique, Université de Sherbrooke) for the initial electrical/mechanical design and burn-in, Jean-François Beaudoin (Sherbrooke Molecular Imaging Centre, CIMS) for the design of the plastic cartridge and the help in the prototype optimization, Mélanie Archambault (CIMS) for the support with animal experiments as well as Luc Tremblay (CIMS) for the MRI scans.

## Data availability statement

The data that support the findings of this study are available upon reasonable request from the authors.

## Funding

This research was funded in part by Biogen MA and the Natural Science and Engineering Research Council of Canada. The Research Centre of CHUS (CRCHUS) hosting the Sherbrooke Molecular Imaging Centre is a FRQS-funded research organization.

## ORCID iDs

Laurence Convert et al. https://orcid.org/0000-0003-2370-4062

Otman Sarrhini et al. https://orcid.org/0000-0001-6231-2630

Paul G Charette et al. https://orcid.org/0000-0003-3604-469X

Roger Lecomte et al. https://orcid.org/0000-0002-8541-0783

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