15th European Molecular Imaging Meeting
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MRI, MRS & Hyperpolarization Technologies

Session chair: Eric Ahrens (La Jolla, USA); Ruiqing Ni (Zurich, Switzerland)
Shortcut: PS 11
Date: Wednesday, 26 August, 2020, 3:45 p.m. - 5:15 p.m.
Session type: Parallel Session


Abstract/Video opens by clicking at the talk title.

3:45 p.m. PS 11-01

Introductory Lecture

Kevin Brindle1

1 University of Cambridge, Cambridge, United Kingdom

4:00 p.m. PS 11-02

Development of non-invasive porcine ex- and in-vivo sodium imaging acquisition and analysis methods using 23Na-MRI

James T. Grist1, 3, Esben Hansen2, Nikolaj Boegh2, Rasmus Tougaard2, Ferdia A. Gallagher3, Christoffer Laustsen2

1 University of Birmingham, Institute of Cancer and Genomic Sciences, Birmingham, United Kingdom
2 Aarhus University, Clinical Medicine, Aarhus, Denmark
3 University of Cambridge, Radiology, Cambridge, United Kingdom


Renal sodium handling is highly dependent on renal oxygenation and metabolism(1). In particular, renal sodium transport is mediated by ATP dependent transport proteins and injuries such as acute kidney injury and chronic renal disease impair the function of these transporters(2).
Sodium MRI is a promising tool for studying a range of clinical areas including the renal system(3).
Here we provide the first in vivo and ex vivo porcine renal sodium imaging data at 3T, demonstrating a change in renal sodium in response to diuretics, paving the way for further clinical studies.


Six pigs were imaged supine in a 3T MRI scanner (GE Healthcare, WI) and sodium imaging (3D cones, Time resolution = 5minutes, Echo time = 0.5ms, Repetition time = 50ms, Voxel volume = 8mm3) performed using a Helmhotlz loop coil pair (Pulseteq, Cobham) before and after a frusemide bolus (0.5 mgkg-1). Sodium concentration maps were calculated from phantoms placed in the imaging FOV (4% agar, 32 and 80mML-1)(4).

Porcine kidneys were excised and placed into a phantom containing NaCl (150mML-1). Sodium T1 mapping was performed (Inversion times = 20-42ms). Regions of interest were used to segment the medulla, and cortex and cortico-medullary sodium gradient(5). Statistical analysis was performed to assess for alterations in renal sodium before and after furosemide.


Sodium imaging was successful performed in vivo (before and after furosemide - Figure 1A and B, respectively) and demonstrated a significant difference between the medulla and cortex sodium distribution (Figure 1C –92±4vs51±5mML-1, respectively, p<0.05) and between the medulla sodium pre and 30 minutes post furosemide (92±4vs.69±3mML-1, respectively, p<0.05). Furthermore, significant differences in the renal sodium gradient before and 20 minutes after the introduction of furosemide was observed (Figure 1E -2.76±0.15vs2.03±0.23mML-1mm-1, respectively, p<0.05). Renal T1 mapping revealed a very isotropic relaxation distribution across the ex vivo renal system (26±2ms) with T1 weighted and T1 map images shown in figure 2A and B, respectively. Here we have demonstrated the ability of sodium imaging at 3T to detect alterations in renal sodium handling, as well as ex vivo methods to further probe renal microstructure, allowing for further studies of renal pathology.


This study has combined both ex vivo and in vivo sodium imaging to analyse the healthy porcine kidney. The results showed highly reproducible sodium concentration and relaxation values. A particular strength of this study is the use of a clinical MRI scanner, allowing for the translation of methods to human studies in both health and disease.


This study was funded by the Lundbeck Foundation, Medical Research Council, and Aarhus University.

[1] Deng A, Miracle CM, Suarez JM, et al. Oxygen consumption in the kidney: Effects of nitric oxide synthase isoforms and angiotensin II. Kidney Int. 2005;68:723–730.
[2] Zhou HY, Chen TW, Zhang XM. Functional Magnetic Resonance Imaging in Acute Kidney Injury: Present Status. Biomed Res. Int. 2016
[3] Madelin G. Sodium Magnetic Resonance Imaging: Biomedical Applications. 2012.
[4] Christensen JD, Barrère BJ, Boada FE, Vevea JM, Thulborn KR. Quantitative tissue sodium concentration mapping of normal rat brain. Magn. Reson. Med. 1996;36:83–89.
[5] Piskunowicz M, Hofmann L, Zuercher E, et al. A new technique with high reproducibility to estimate renal oxygenation using BOLD-MRI in chronic kidney disease. Magn. Reson. Imaging 2015;33:253–261.
Figure 1 - Porcine renal sodium imaging
A: Example fused baseline proton-sodium imaging from a different subject demonstrating higher sodium concentration in the medulla compared to the cortex. B: Fused proton-sodium imaging 30 minutes post furosemide from the sample subject as B. C: Sodium concentration derived by manual segmentation results before (black) and 30 minutes after (red) furosemide introduction, revealing a significant decrease in medullary sodium concentration. *= Significant difference, p<.05. D: Dynamic changes in the cortico-medullary sodium gradient *= Significant difference vs baseline, p < 0.05.
Figure 2 - ex vivo porcine kidney measurements
A: Example T1-weighted image (inversion time = 30ms) of ex vivo kidneys
B: Example T1 map of ex vivo kidneys
Keywords: Sodium, Metabolism, Renal, Imaging, MRI
4:12 p.m. PS 11-03

Minimally invasive implantable NMR microcoils for in vivo metabolic profiling of microliter volumes

Justine Deborne1, Noël Pinaud1, Luisa Ciobanu2, Alan Wong3, Yannick Crémillieux1

1 Université de Bordeaux, Institut des Sciences Moléculaires, Bordeaux, France
2 CEA Saclay, Neurospin, Gif-sur-Yvette, France
3 CEA Saclay, Nimbe, Gif-sur-Yvette, France


The use of implanted NMR microcoils still remains a relatively unexploited research area, without emerging or significant biomedical applications.  The limitations inherent to implanted NMR coils derive obviously from the relatively weak detection sensitivity of NMR, hindering great challenges for nano-volume analyses. In addition, the necessity to preserve tissue during microprobe implantation imposes severe constraints on the geometry and structure of the NMR microcoil.  In this study, we present in vitro and in vivo results obtained with innovative and minimally invasive microcoils.


An example of implantable NMR microprobe is shown in Figure 1. This filar-type architecture is based on the use of twisted copper microwires (diameter of 150 μm). The twisted wires are inserted inside a polyamide tubing (outer diameter of 380 μm). A biocompatible glue is used to seal the polyamide tube, while tuning and matching capacitors are connected to the two sides of the wire. For in vivo experiments, cannulae were stereotaxically positionned the day before the insertion of the NMR microcoils in the brain of male wistar rats. Experiments were performed at 7 T and 17.2 T. NMR spectra were acquired using a PRESS sequence. 3D MRI acquisitions were performed using a ZTE (zero echo time) sequence.


The quality factor of the loaded coils was ranging between 100 and 120. The full width at half maximum of water peak were measured to 6 Hz. In vivo results are illustrated in Figure 2 with a PRESS NMR spectrum obtained in the rat brain with a twisted microcoil (a volume coil was used for selective excitation). In this particular example of water-suppressed acquisition (240 averages, 10-minutes acquisition) at 7 Tesla, main peaks of brain metabolites (NAA, glu, gln, pCr, Cr, etc) can be easily identified and quantified. The ZTE MRI image (right side of Figure 1) shows the sensitive detection zone of the microprobe (volume evaluated to 500 nL) extending to about 200 μm away from the wire.


The MRS/MRI results obtained in vitro and in vivo illustrate the relevance of the microcoil design with respect to spectral resolution, detection sensitivity, spatial selectivity and limited invasiveness. Foreseen applications include the investigation of metabolism in microliter volumes in physiological conditions and in diseases with metabolic dysfunctions (tumoral environement, neurodegenerative pathologies, etc).


The study received financial support from the « Laboratory of Excellence » TRAIL ANR-10-LABX-57 (research program Insight) and from France Life Imaging (FLI).

Figure 1
Left: implantable twisted microcoil with 150-micrometer diameter copper wire and a total length of 3 mm. Right: MRI image obtained with the microcoil, illustrating the sensitive detection zone of the NMR probe.
Figure 2
NMR spectrum acquired at 7 Tesla in the rat brain using a PRESS sequence with water suppression. Total acquisition time: 10 minutes.
Keywords: MRS and MRI, metabolic profiling, implanted NMR microcoil
4:24 p.m. PS 11-04

Simulation of Focused RF Heating using a High Channel Count MRI RF Coils and a Maximum SAR Algorithm

Koray Ertan1, 2, Joshua D. Bever1, Mihir Pendse1, Paolo Decuzzi2, Brian Rutt1

1 Stanford University, Department of Radiology, Palo Alto, United States of America
2 Italian Institute of Technology, Genoa, Italy


Ultra-high field MRI scanners commonly use parallel transmit RF coil arrays to mitigate transmit magnetic field inhomogeneity while limiting absorbed power (SAR). Parallel transmit coils can also be used to intentionally generate targeted hyperthermia [1-2] by maximizing SAR in the target region while constraining SAR in background tissues. A 70 channel loop coil array was simulated at multiple frequencies including 7T and 10.5T MRI Larmor frequencies and a maxSAR algorithm [3] used to compute optimal channel weights for controlled focal SAR and thermal hot spot formation.


A 70-channel loop array consisting of 5 rows and 14 columns was designed to fit human head. EM design, modeling and thermal simulations were performed using Sim4Life and Ella body model [4] at several frequencies (298, 447, 600, 900, 1200, 1500 MHz). The maxSAR algorithm [3] was used to maximize the peak target 10g-SAR while peak background 10g-SAR was limited to 20W/kg. SAR values in the transition region surrounding the target region were not included in the optimization. Local optimization was solved in less than 4min thanks to GPU-accelerated computation of SAR matrices [3]. Optimized SAR maps were used as heat source terms in a Pennes’ bioheat model that included perfusion and metabolic heat generation terms to obtain temperature and cumulative equivalent minutes at 43oC (CEM43) maps.


Figure 1 shows 10g-SAR, temperature and CEM43 maps resulting from maxSAR optimization for an example target region near the center of the brain. For this specific target region, 900 MHz has advantages over the other simulated frequencies, generating the highest maximum target SAR (~100W/kg), temperatures (~45oC) and CEM43 (~16min) values. Results are reported for 10 minutes of RF application because this was long enough to reach steady state temperature in target tissues. High levels of SAR, temperature and CEM43 focusing are achieved in the target region with focal spot diameters of approximately 1.6 cm in all directions, although performance varies significantly across different target locations. Figure 2 compares the results for all simulated frequencies at all target locations. Overall, our results show that target temperatures of 41-45 oC and target CEM43s of 0.2-24.7 min can be achieved within highly localized volumes (~4cc) with only 10 minutes of RF exposure.


We demonstrate that very high target SAR values (~100W/kg) are possible with background SAR no higher than 20W/kg.  Thermal and CEM43 results suggest that this RF focusing performance should be enough for clinical hyperthermia applications, especially for targeted blood brain barrier permeability modulation [5]. The efficiency of the maxSAR algorithm allows brute force optimization over a wide parameter space to find the optimal hardware design.

[1] Winter, Lukas, et al. "Design and evaluation of a hybrid radiofrequency applicator for magnetic resonance imaging and RF induced hyperthermia: electromagnetic field simulations up to 14.0 Tesla and proof-of-concept at 7.0 Tesla." PloS one 8.4 (2013): e61661.
[2] Guérin, Bastien, et al. "Computation of ultimate SAR amplification factors for radiofrequency hyperthermia in non-uniform body models: impact of frequency and tumour location." International Journal of Hyperthermia 34.1 (2018): 87-100.
[3] Pendse, M., and B. Rutt. "An algorithm for maximum-SAR targeted RF hyperthermia." Proc Intl Soc Mag Reson Med. 2015. Abstract No: 3224.
[4] Gosselin, Marie-Christine, et al. "Development of a new generation of high-resolution anatomical models for medical device evaluation: the Virtual Population 3.0." Physics in Medicine & Biology 59.18 (2014): 5287.
[5] Yarmolenko, Pavel S., et al. "Thresholds for thermal damage to normal tissues: an update." International Journal of Hyperthermia 27.4 (2011): 320-343.
Figure 1
Maximum intensity projections of optimized (left) 10g-SAR distribution as well as resulting (middle) temperature and (right) CEM43 maps for 298, 447, 600, 900, 1200, 1500 MHz. Temperature and CEM43 results are provided for 10 minutes of RF application. Each row shows the results for target volume located near the center of the brain. Throughout the study, the target region was defined by a sphere of 2 cm radius while a surrounding sphere of 4 cm radius was used to bound the transition region.
Figure 2
Evaluation of maxSAR optimization performance for 298, 447, 600, 900, 1200 and 1500 MHz. Maximum and mean 10g-SAR values in target and background tissues for (left) 10g-SAR, (middle) temperature and (right) final CEM43 values after 10 minutes of RF application are plotted. Each row shows the results for different target locations. First and second set of four target locations are in the central and peripheral parts of the brain respectively. (Vertical axes of CEM43 results are plotted in logarithmic scale due to very high difference between CEM43 in target and background tissue.)
Keywords: ultra-high field MRI, parallel transmit RF coil arrays, focused RF heating, targeted drug delivery, targeted hyperthermia
4:36 p.m. PS 11-05

Cerebral metabolism of hyperpolarized [2H7, U-13C6] glucose in the healthy mouse under different anesthetic conditions

Emmanuelle Flatt1, Bernard Lanz1, Andrea Capozzi1, Magnus Karlsson2, Mathilde Hauge Lerche2, Rolf Grütter1, Mor Mishkovsky1

1 Ecole Polytechnique Fédérale de Lausanne (EPFL), LIFMET, Lausanne, Switzerland
2 Danmarks Tekniske Universitet (DTU), Lyngby, Denmark


Glucose is the primary fuel for the brain and its metabolism is linked with cerebral function1. Isoflurane anesthesia is commonly employed in preclinical MRS but affects cerebral hemodynamics2 and functional connectivity3. The combination of isoflurane and medetomidine is routinely used in rodent fMRI and show similar functional connectivity as in awake animals3. Given that glucose metabolism is tightly linked to neuronal activity1, the aim of this study was to compare the cerebral metabolism of hyperpolarized (HP) [2H7, U-13C6] glucose under these two anesthetic conditions.


Brain metabolism of HP glucose was monitored in two groups of 12h-fasted male C57BL6/J mice in a 9.4T MRI system. In group 1 (ISO), mice were kept under 1.3-1.6% isoflurane for the entire duration of the experiment (N=8). In group 2 (MED+ISO), anesthesia was switched to a combination of medetomidine and 0.25-0.5% isoflurane3 one hour before injection of glucose (N=6). Single voxel 1H MRS were carried out in the hippocampus of each mouse 10 minutes before the glucose bolus. Blood analysis were performed before the mouse entered the scanner and ˜2 minutes after the end of the 13C MRS acquisition. 540μL of 44±10mM HP [2H7,13C6] glucose was injected through the femoral vein and the signal was acquired every 1s for 70s. Lactate to glucose ratio (LGR) was calculated from the summed spectra.


In both anesthetic conditions, the injection of HP [2H7,13C6]glucose resulted in de novo synthesis of [1-13C]lactate and the timecourses were well reproducible(Fig.1). A larger amount of lactate was produced in the case when the functional connectivity was similar to the awake animals3(+128% in LGR in MED+ISO, Fig.2A), in line with studies suggesting that astrocytic lactate production can be stimulated during neuronal activation4,5. The correlation between isoflurane concentration and LGR suggests that little variations in isoflurane concentration directly influence the dynamic of lactate synthesis (Fig.2B).
The Cr/PCr ratio, related to ATP and ADP balance, was significantly different between the groups, indicating a difference in the energetic state. Also, endogenous lactate concentration was higher in the ISO group(Fig.2C), implying that [1-13C]lactate production from HP glucose is not a reflection of the steady-state pool-size but rather related to the dynamic of the glycolytic flux.


Changing the anesthesia from isoflurane to a combination of isoflurane and medetomidine is known to influence brain activity3 and had high impact on cerebral uptake and metabolic flux of HP [2H7,U-13C6]glucose in the mouse. When using the combination, the [1-13C]lactate signal and lactate to glucose ratio were more than doubled. The higher SNR reported in this work is an important step toward real-time imaging of cerebral glycolysis.


Supported by CIBM of the UNIL, UNIGE, HUG, CHUV, EPFL, the Leenaards and Jeantet Foundations.

[1] Sokoloff L., Neurochem Res., 1999
[2] Todd, M. et al., J Neurosurg Anesthesiol., 1996
[3] Paasonen, J. et al., NeuroImage., 2018
[4] Hertz, L. et al., Neurochem Res., 1988
[5] Magistretti, P.J. et al., Dev. Neurosci., 1993
Figure 1

A,B - Characteristic spectra measured in a mouse brain under isoflurane and under medetomidine + isoflurane anesthesia following infusion of HP [2H7,13C6] glucose. The glycolytic intermediate 3PG (179.8ppm) can be identified in the summed spectra in addition to the lactate at 183.5ppm. The broad peak at 175 ppm designated by (*) is an impurity in the [2H7,13C6] glucose powder.

C,D - The corresponding timecourses (mean ± SD) of [1-13C] lactate and glucose are shown for each group.

Figure 2

A - Comparison of LGR between isoflurane (ISO) and  medetomidine + isoflurane (MED+ISO) groups.

B – Correlation between LGR and isoflurane concentration(%).

C - Brain Lac and Cr/PCr comparison (mean±SD) between the ISO group and the MED+ISO group.

D - 1H-MRS spectra in the hippocampus of an ISO mouse (green) and a MED+ISO mouse (red). Part of the spectra corresponding to Lac and Cr&PCr are shown in blue and dark, respectively.

Statistical analysis was performed using one-way ANOVA in GraphPad Prism software.

Keywords: cerebral metabolism, dDNP, glucose, 13C MRS, anesthesia
4:48 p.m. PS 11-06

Real Time Nuclear Magnetic Resonance Detection of Fumarase Activity using Parahydrogen-Hyperpolarized [1-13C]fumarate

James Eills2, Eleonora Cavallari1, Carla Carrera1, 4, Dmitry Budker2, 3, Silvio Aime1, Francesca Reineri1

1 University of Torino, Dept. of Molecular Biotechnology and Health Sciences, Torino, Italy
2 Johannes Gutenberg University, Helmholtz Institute, Mainz, Germany
3 University of California, Department of Physics, Berkeley, United States of America
4 National Research Council of Italy, Institute of Biostructures and Bioimaging, Torino, Italy


Nuclear magnetic resonance experiments are often limited by the intrinsic low sensitivity of the technique. The sensitivity can be improved by ‘hyperpolarizing’ the nuclear spins of a compound under study [1,2]. A promising candidate metabolite is [1,4-13C2]fumarate, which is used in carbon-13 MRI experiments as a sensitive marker of cell necrosis [3], and is currently being assessed for application in clinical trials [2].
In this work, we generate [1-13C]fumarate via PHIP at a concentration of 45 mM in aqueous solution, and demonstrate a 13C polarization level of 25%.


We trans-hydrogenate [1-13C]acetylene dicarboxylate (50 mM in D2O) with para-enriched hydrogen (86 % enriched) using a commercially available Ru catalyst in water to produce hyperpolarized [1-13C]fumarate.
In our experiment, the proton singlet order was transformed into hyperpolarized magnetization on the 1‑13C spin by subjecting the sample to a 'costant-adiabaticity' magnetic field cycle (MFC). The chemical reaction and MFC profile are shown in Fig. 1.
To study PHIP-polarized fumarate as a probe of cellular necrosis, the reaction mixture containing HP [1‑13C]fumarate was added to EL-4 cell suspensions.Three experiments were performed: one with media containing 107 lysed EL-4 tumour cells, one with 107 intact (healthy) EL-4 tumour cells, and one with no cells in the media.


The observed signal enhancement factor was calculated to be 20,000 at 14.1 T, corresponding to a polarization level of 24%. The ratio of peak integrals between fumarate and succinate in the thermal spectrum indicates 89% fumarate yield (45 mM), and 11% succinate yield (5 mM), and there is no detectable unreacted starting material.
After perfusion of the [1‑13C]fumarate containing solution through the cells suspension, the 13C NMR signal was acquired every 2 s using 15° flip-angle pulses. From the control experiment without cells in the media the fumarate T1  under our experimental conditions was measured to be 28 ± 2 s.
In the experiment with intact cells, fumarate transport across the cell membrane is slow relative to , so no malate signal is detected. In the experiment with lysed cells, the fumarate has access to the enzyme and reacts with water to form malate with a reaction constant k = 1.62 ± 0.01 x 10-3 s‑1.


In conclusion, we have demonstrated that PHIP allows for the production of highly polarized [1‑13C]fumarate (24%) at a concentration of 45 mM in aqueous solution, using relatively inexpensive and easily operated PHIP equipment. Following the perfusion of lysed cells metabolic transformation into malate was observed. This work paves the way to accelerate studies with HP fumarate, to better undertand what role it can play in diagnostics.


This project has received funding from the European Union’s Horizon 2020 research and innovation program under the Marie Skłodowska-Curie Grant Agreement No. 766402, and Compagnia di San Paolo (Athenaeum Research 2016, n. CSTO164550). The authors would like to thank Bogdan Rodin for calculating the constant adiabaticity field-cycling profile.

[1] Wang, Z. J.; Ohliger, M. A.; Larson, P. E. Z.; Gordon, J. W.; Bok, R. A.; Slater, J.; Villanueva-Meyer, J. E.; Hess, C. P.; Kurhanewicz, J.; Vigneron, D. B., 2019, 'Hyperpolarized 13C MRI: State of the Art and Future Directions', Radiology, 291, 273-284.
[2] Kurhanewicz, J., Vigneron, D. B., Ardenkjaer-Larsen, J. H.; et al., 2019, 'Hyperpolarized 13C MRI: Path to Clinical Translation in Oncology', Neoplasia, 21, 1-16
[3] Gallagher, F. A., Kettunen, M. I., Hu D.-E., et al., 2009, 'Production of hyperpolarized [1,4-13C2]malate from [1,4-13C2]fumarate is a marker of cell necrosis and treatment response in tumors', Proc. Natl. Acad. Sci., 106, 19801-19806
[4] Ripka, B., Eills, J., Kourilova, H., Leutzsch, M., Levitt, M. H., Münnemann, K., 2018, 'Hyperpolarized fumarate via parahydrogen', Chem. Commun., 54, 12246-12249
[5] Eills, J.; Cavallari, E.; Carrera, C.; Budker, D.; Aime, S.; Reineri, F., 2019, 'Real Time Nuclear Magnetic Resonance Detection of Fumarase Activity using Parahydrogen-Hyperpolarized [1-13C]fumarate', JACS (to press)
Figure 1
Reaction scheme showing the chemical addition of para-enriched hydrogen to an unsaturated [1-13C]acetylene dicarboxylate precursor, to yield [1-13C]fumarate. The proton singlet order is then transformed into 13C magnetization by applying a 'constant-adiabaticity' magnetic field cycle, optimized for the [1-13C]fumarate J-couplings. The field-cycling profile is shown in the inset.
Figure 2
Top: The enzymatic conversion of [1-13C]fumarate into both [1-13C]malate and [4-13C]malate. Bottom: Flux of 13C label between fumarate and malate in a suspension of lysed EL4 tumour cells. The inset spectrum shows a single acquisition from the dataset . Fits to the data are shown by dashed blue (fumarate) and magenta (malate) lines Image taken with
permission from Ref [5].
Keywords: Fumarate, Parahydrogen, Hyperpolarization, Metabolism, 13C NMR
5:00 p.m. PS 11-07

A multi spin echo pulse sequence with optimized excitation pulses and a 3D cone readout for hyperpolarized 13C imaging

Vencel Somai1, 2, Alan Wright1, Maria Fala1, Friederike Hesse1, Kevin Brindle1, 3

1 University of Cambridge, Cancer Research UK Cambridge Institute, Cambridge, United Kingdom
2 University of Cambridge, Department of Radiology, Cambridge, United Kingdom
3 University of Cambridge, Department of Biochemistry, Cambridge, United Kingdom


Imaging tumour metabolism in vivo using hyperpolarized [1-13C]pyruvate is a promising technique for detecting disease, disease progression and assessing treatment response1,2. However, the transient nature of the hyperpolarization and its depletion following excitation limits the available time for imaging3. We describe here a single shot multi spin echo sequence, which improves on previously reported sequences, with a shorter readout time, isotropic point spread function (PSF) and field of view (FOV) and better signal-to-noise ratio (SNR).


The sequence uses numerically optimized excitation pulses and hyperbolic secant adiabatic refocusing pulses, all applied in the absence of slice selection gradients. The excitation pulses were designed to give constant phase in the pass-band immediately after the pulse in order to increase signal in the presence of large B0 and B1 field inhomogeneities and to minimize the achievable echo time. The gradient readout uses a 3D cone trajectory distributed among 7 spin echoes (Figure 1). Experiments were performed at 7T (Agilent, Palo Alto, CA). A 42 mm diameter birdcage volume coil was used for 1H transmit and receive and a similar volume coil for 13C transmit. A 20 mm diameter surface coil was used for 13C receive (Rapid Biomedical GMBH, Rimpar, Germany).


The maximal gradient amplitude and slew-rate were set to 4 G/cm and 20 G/cm/ms respectively to demonstrate the feasibility of clinical translation. This gave a minimal repetition time of 165 ms. The pulse sequence had an isotropic FOV of 32 mm and nominal resolution of 2 mm and gave an isotropic PSF of 2.8 mm when reconstructed at the 0.125 mm in-slice resolution of the anatomical reference 1H image. The lack of slice selection, which was possible because of the isotropic FOV and the localized sensitivity profile of the receiver coil, enabled the design of a pulse optimised for frequency selectivity. Echo formation at the end of the pulse enabled very short echo times. The optimized spectral profile showed a high degree of spectral selectivity. The sequence was demonstrated with dynamic imaging of hyperpolarized [1-13C]pyruvate and [1-13C]lactate in a murine tumour model, and the data obtained were in good agreement with previous findings in this tumour model(Figure 2).


The pulse sequence was capable of dynamic imaging of hyperpolarized 13C labelled metabolites with relatively high spatial and temporal resolution. The segmented k-space readout fully exploited the long T2 relaxation time by sampling at multiple spin echoes to maximize the SNR. With a 32 cm FOV and 2 cm resolution, typical of a clinical scanner, the readout takes only 29.133 ms, and should preserve the high degree of robustness to imperfections.

[1] Gram A, Hansson G, Hansson L, et al. Increase in signal-to-noise ratio of. 2003:1-6. papers2://publication/uuid/B5F38F51-C552-4E4E-8B3C-6AC6B9109E56.
[2] Brindle KM. Imaging Metabolism with Hyperpolarized 13C-Labeled Cell Substrates. J Am Chem Soc. 2015;137(20):6418-6427. doi:10.1021/jacs.5b03300
[3] Day SE, Kettunen MI, Gallagher FA, et al. Detecting tumor response to treatment using hyperpolarized 13C magnetic resonance imaging and spectroscopy. Nat Med. 2007;13(11):1382-1387. doi:10.1038/nm1650
[4] Wang J, Hesketh RL, Wright AJ, Brindle KM. Hyperpolarized 13 C spectroscopic imaging using single-shot 3D sequences with unpaired adiabatic refocusing pulses. NMR Biomed. 2018;31(11):1-12. doi:10.1002/nbm.4004
[5] Wang J, Wright AJ, Hu DE, Hesketh R, Brindle KM. Single shot three-dimensional pulse sequence for hyperpolarized 13C MRI. Magn Reson Med. 2017;77(2):740-752. doi:10.1002/mrm.26168
Figure 1
The pulse sequence starts with an optimized excitation pulse, without a slice selection gradient, and contains 7 unpaired adiabatic refocusing pulses to generate 7 spin echoes and an optional 8th adiabatic pulse at the end of the sequence to flip back the spins parallel to the +z-direction in case of very long injection. At each echo a pair of two identical  cones were acquired in decreasing order with respect to cone angle. The fine structure of the first cone is not visible due to the plot linewidth. The whole sequence takes 165 ms.
Figure 2
Hyperpolarized [1-13C]pyruvate (A) and [1-13C]lactate (B) images from a tumor-bearing mouse overlaid on the corresponding 1H images (2 mm slice thickness). The 13C images were interpolated to the 128x128 in-plane matrix size of the 1H image and summed in time over the first 20 s. The relatively small signal leakage from neighbouring slices in slices 6 and 12 suggests that the simulated sampling PSF profile is well preserved.
Keywords: hyperpolarized, imaging, cone trajectory, tumour
5:12 p.m. PS 11-08

INDUSTRY TALK (Bruker) |Hyperpolarized 129Xe MRI of Pulmonary Gas-Exchange in Rodents

Rohan Virgincar1

1 Bruker BioSpin, Ettlingen, Germany


This lecture will give insight into the fundamentals and practical considerations for high-resolution hyperpolarized 129Xe MRI in rodents, based on his 7 years of research in this field at Duke University, USA. Important factors such as hyperpolarization, xenon gas-delivery using a hyperpolarized-gas-compatible ventilator, imaging considerations for ventilation and dissolved-phase imaging using the 3D ultrashort echo time method, image analysis techniques, and present examples of applications in clinically relevant rodent models of lung disease will be discussed.

Keywords: hyperpolarization