Fast, quantitative, murine cardiac 19F MRI/MRS of PFCE-labeled progenitor stem cells and macrophages at 9.4T

Purpose To a) achieve cardiac 19F-Magnetic Resonance Imaging (MRI) of perfluoro-crown-ether (PFCE) labeled cardiac progenitor stem cells (CPCs) and bone-derived bone marrow macrophages, b) determine label concentration and cellular load limits, and c) achieve spectroscopic and image-based quantification. Methods Theoretical simulations and experimental comparisons of spoiled-gradient echo (SPGR), rapid acquisition with relaxation enhancement (RARE), and steady state at free precession (SSFP) pulse sequences, and phantom validations, were conducted using 19F MRI/Magnetic Resonance Spectroscopy (MRS) at 9.4 T. Successful cell labeling was confirmed using flow cytometry and confocal microscopy. For CPC and macrophage concentration quantification, in vitro and post-mortem cardiac validations were pursued with the use of the transfection agent FuGENE. Feasibility of fast imaging is demonstrated in murine cardiac acquisitions in vivo, and in post-mortem murine skeletal and cardiac applications. Results SPGR/SSFP proved favorable imaging sequences yielding good signal-to-noise ratio values. Confocal microscopy confirmed heterogeneity of cellular label uptake in CPCs. 19F MRI indicated lack of additional benefits upon label concentrations above 7.5–10 mg/ml/million cells. The minimum detectable CPC load was ~500k (~10k/voxel) in two-dimensional (2D) acquisitions (3–5 min) using the butterfly coil. Additionally, absolute 19F based concentration and intensity estimates (trifluoroacetic-acid solutions, macrophages, and labeled CPCs in vitro and post-CPC injections in the post-mortem state) scaled linearly with fluorine concentrations. Fast, quantitative cardiac 19F-MRI was demonstrated with SPGR/SSFP and MRS acquisitions spanning 3–5 min, using a butterfly coil. Conclusion The developed methodologies achieved in vivo cardiac 19F of exogenously injected labeled CPCs for the first time, accelerating imaging to a total acquisition of a few minutes, providing evidence for their potential for possible translational work.

Introduction Implantation of stem cells (SCs) has provided a methodological pathway that promises cardiac tissue regeneration and structural and functional improvements following injury. The basic approach of SC therapy involves the direct transplantation of cells, followed by their migration, differentiation, and proliferation, ultimately attaining homing and engraftment. However, while the feasibility of SC technologies has been proven, efficacy is still in question [1].
Within the realm of SC therapies, non-invasive imaging and tracking of labeled SCs, and their functional impact, has taken a prominent role in recent years. The visualization of the implanted SCs to define optimal therapy strategies (dose, timing, delivery) using pre-labeled cells with fluorescent probes [2], transduced expression of fluorescent proteins [3], or iron oxide particles (MPIOs) [4], and their assessment for temporal label persistence, has become a subject of intense research. Over the past decade, nanoparticles (NPs) containing perfluorocrown-ethers (PFCE) have led to direct tracking and quantification of exogenously labeled cell populations [5,6,7,8] with 19 F magnetic resonance imaging (MRI).
Despite the implementation of 19 F MRI early on in the development of MRI, exploitation efforts had languished until recent years [5,6,9,10]. The resurgence of interest in 19 F imaging arose further to initiatives in molecular imaging, and capitalized on the exogenously injected fluorine's 100% abundance, and the high relative sensitivity and contrast with respect to the 1 H nucleus. The lack of endogenous fluorine provides fluorinated labels an added advantage as tracking agents. Consequently, the technique has found applicability in cellular labeling and tracking applications in vivo [5,11], with potential for translational value [12].
Furthermore, prior applications were confined to either direct injections of neural SCs [13], immune cells [6,7,9], hematopoietic SCs [14], or on direct intravascular administrations of NP emulsions [15,16,17,18,19,11]. Correspondingly, there have been no prior reported 19 F MRI preclinical applications in normal or infarcted hearts using exogenously administered, labeled progenitor SCs, while direct applications of other types of SCs in the rodent heart have been limited [14].
Prior efforts attempted to optimize fluorine acquisitions in terms of speed, evoked MR signal, and cellular detectability [11], by focusing on spectroscopy [17,18,19] or on dedicated pulse sequences [20][21][22][23][24][25][26], and by selecting imaging parameters that elicited maximum signal responses, despite the prohibitively long acquisition times. To our knowledge, there is no prior 19 F MRI study on the use of labeled cardiac progenitor cells (CPCs) (previously used to show efficacy of regeneration and cardiac functional improvements [27]). Certainly, lacking are also detailed relaxometric studies in these cells post-labeling.
We present a comprehensive methodology that applies 19 F MRI aiming to achieve: a) fast imaging of PFCE-labeled CPCs within clinically relevant times (of the order of a few minutes) 090532/Z/09/Z), and British Heart Foundation (BHF) grant FS/11/50/29038 (JS). The funders had no role in study design, data collection and analysis, decision to publish, or preparation of the manuscript.
Competing interests: I would also like to respectfully indicate that I am currently affiliated with Chi Biomedical Ltd. (since my departure from U. Oxford and UK on 7/8/17). However, the presented work in this manuscript was completed during the period of July 2015-June 2017 at the U. Oxford during my employment at U. Oxford as a Marie-Sklodowska Curie fellow. Correspondingly, there are no issues relevant to funding or competing interests. Chi Biomedical has been in a financial dormant status for a number of years and despite its viable legal status, it has been financially inactive. Correspondingly, there was no salary contributed by Chi Biomedical directors/staff, and for this submitted work there are no declarations pertaining to employment, consultancy, patents, products in development, or marketed products. The current commercial affiliation of the lead author [CC] commercial affiliation does not alter the adherence of all the authors/coauthors of this work to PLOS ONE policies on sharing data and materials.
in the in vivo mouse, b) determination of detection limits of label cellular load with clinically applicable surface and volume coils, and c) spectroscopy and image-based quantification validated in phantoms, CPCs, labeled bone-marrow-derived murine macrophages, and in the post-mortem mouse. The stated objectives were investigated based on theoretical and simulation comparisons of pulse sequence performances, in vitro relaxation value characterization of PFCE-labeled CPCs, experimental concentration validations, and post-mortem and in vivo applicability of the imaging approach in the cardiac and skeletal muscles of the C57BL/6 mouse.

Animal ethics
All procedures were in accordance with the Home Office (UK) guidelines under The Animals (Scientific Procedures) Act, 1986 (Permit Number: PIL30/3322), the European Animal Research Directive, and with local institutional guidelines. All surgery and live animal imaging was performed under isoflurane (ISO) anesthesia, and all efforts were made to minimize suffering. Animals were euthanized using cervical dislocation.

Cardiac progenitor stem cells and bone marrow-derived macrophages
Nanoparticle synthesis. Particles were synthesized in accordance to previously published methodologies [7]. All particles were then extensively washed with distilled water and lyophilized for 2-3 d. For the cell labeling, nanoparticles were synthesized with the addition of fluorescent dye (Atto647, ATTO-TEC, GmbH, Germany) to the organic phase. Prior 19 F Magnetic Resonance Spectroscopy (MRS) characterization of PFCE NP labels confirmed the presence of a single spectral peak at -91.8 ppm (with respect to CFCl 3 ) [11].
Cell isolation. Cardiac progenitor cells (CPCs, comprising either cardiosphere-derived (CDC) or collagenase-trypsin (CT)) were isolated from adult, C57BL/6, green fluorescent protein (GFP) positive or GFP negative, mouse atria, using standard protocols [28], and were maintained in Iscove's Modified Dulbecco's Medium (IMDM) media (Thermo Fisher Scientific, UK). CDCs have been previously used to show efficacy of regeneration and cardiac functional improvements [27], while CTs have been recently shown to express similar cardiac phenotypic characteristics to CDCs [29]. Additionally, bone marrow-derived macrophages were cultured from bone marrow harvested from C57BL/6 mouse hindlimbs. Bone marrow cells were washed with phosphate buffer solution (PBS, Sigma-Aldrich, UK), passed through a cell strainer to produce a single cell suspension, and differentiated for a week in petri-dishes in Dulbecco's modified eagle's medium (DMEM) that contained L-cell conditioned media. At that point, adherent bone-marrow derived macrophages were harvested. They were subsequently washed with PBS and were re-suspended in pellets in improved minimal essential medium (OPTIMEM).
Cellular culture and labeling. Cells were plated in IMDM solutions and incubated with PLGA-PFCE-Atto647-containing NPs for approximately 24 h before isolation and pelleting. Addition of a fluorescent dye (Atto647) allowed independent flow cytometry and confocal microscopy validation studies. Cell pellet suspensions (CPCs or macrophages) were maintained in media (IMDM or OPTIMEM) and were subsequently used for MRI, flow cytometry, or confocal microscopy, after fixation in 2% methanol-free paraformaldehyde solution (Thermo Scientific Pierce, UK) mixed with PBS (1:7 v/v).
FuGENE labeling. The commercially available DNA transfection agent FuGENE (Promega, Madison, WI, USA) was used to label the CPCs (both CT and CDC cells in separate cultures) with the NPs (using 25 μl of FuGENE in~10 6 cells). FuGENE was pre-mixed and incubated with the NPs before cell transfection for~20 min. Cells were then labeled overnight [30].
Confirmation of cellular label uptake and viability. Successful labeling was confirmed with a CyAn ADP flow cytometer (Beckman Coulter, USA) using control and labeled cell samples. Cellular viability was determined with the Trypan Blue exclusion assay, directly after labeling, and at the completion of MRI studies (wherever applicable), using a cell counter.
High-content epifluorescence imaging. Live cells were plated in 6-well plates and stained with Calcein (CellTrace™ Calcein Red-Orange, ThermoFisher Scientific, UK) and Hoechst (ThermoFisher Scientific, UK) for cytoplasmic and nuclear high-content imaging (Operetta, Perkin-Elmer, UK). Fluorescence was assessed based on Atto647. The Operetta's Harmony software was used for image analyses. Imaging was based on a randomized field analysis methodology that covered each of the studied wells.
MRI/MRS. All experiments were conducted on a 9.4 T Agilent scanner equipped with a DirectDrive console and a 1000 mT/m actively shielded gradient set (internal diameter = 60 mm) (Agilent Technologies, USA). For comparative tests (pulse sequences, radiofrequency (RF) coils), the same acquisition parameters and total acquisition times were used.
Aqueous and cellular phantoms. Twenty-four cylindrical (15-50 ml) phantoms (1-100 mM), containing trifluoro-acetic acid (TFA), PFCE NPs, and labeled CPCs or macrophages, were used to test RF coil responses at 9.4 T (without, and with the use of adiabatic excitations), determine detection limits (TFA, NP solutions mixed in water and IMDM, labeled cells), and for image-and spectroscopy-based quantification.
RF coils. Coils comprising an eight-rung, low-pass, quadrature birdcage (diameter = 34 mm), an 40×20 mm 2 butterfly constructed on a 28 mm diameter former, and a 5 (diameter) ×8 (length) mm 2 solenoid prototypes, were constructed. All coils were tuned at 375.88 MHz (fluorine resonance), and were matched to 50 O. The broad frequency response of the coils (3dB range spanning a few tens of kHz) permitted imaging on both the 1 H and 19 F nuclei.
The butterfly coil was used for murine studies owing to its increased B 1 detection sensitivity. The increased B 1 homogeneity of the birdcage and solenoid coils allowed the a) assessment of the minimum detectable number of 19 F atoms (and the fluorine content of the PLGA-NPs), and b) the estimation of relaxation times.
Adiabatic excitation. Phantom studies (six Eppendorf vials containing aqueous TFA solutions at 5-10 mM) were conducted without (Gaussian RF pulse excitation) and with the use of hyperbolic adiabatic full-passage (HS-AFP) RF pulses [31] using the butterfly coils to determine ultimate concentration detection limits and RF B 1 penetration. Since HS-AFP pulses are non-ideal in achieving proper slice selection, three-dimensional (3D) adiabatic acquisitions were performed, whereby excitation of a thick tissue slab was accompanied by imaging of only a relatively smaller tissue area (relevant to the injected labeled CPCs).
Adiabatic pulse parameters were chosen appropriately to achieve adiabaticity, yet maintaining power to comparable levels as those used for Gaussian (non-adiabatic) excitations [31]  Simulations. SPGR, rapid acquisition with relaxation enhancement (RARE), and fid/ echo steady state free precession (fid-SSFP, echo-SSFP) sequences were simulated in accordance with steady-state, closed-form, signal and signal-to-noise ratio (SNR) equation formulations, as described in the S1 Appendix. Parametric SNR maps were generated in MATLAB (Version 2010b, Mathworks, Natick, MA, USA) using typical 19 F relaxation times and imaging parameters for TFA solutions and CPCs, as these were determined in this study. Estimated SNR values for SPGR and SSFP sequences were normalized to the maximum signal over the entire parametric space. SNR normalization in the case of RARE adhered to the recent analysis presented by Mastropietro et al. [25].
Pulse sequence comparison. Extensive prior in vivo preclinical work has favored RARE imaging [25], and more recently bSSFP [24]. The choice of the optimal pulse sequence for 19  SNR maps were generated by dividing the reconstructed images by the noise SD, estimated from background regions using standard methodologies. Given the relatively large SNR values (>10) of reconstructed images for cell and phantom studies, no bias corrections for the magnitude reconstruction were applied.
For the relaxation measurements of CTs, cells were suspended in Eppendorf tubes and were maintained in IMDM media in ice-cold baths at 4˚C and under normoxic conditions (normal oxygen tension). The tubes were then allowed to reach room temperature before measurements were conducted. The partial pressures of oxygen (pO 2 ) of the tested solutions were not monitored during experiments.
In comparison to the T 1 values, the T 2 values of labeled CT cells were not quantified owing to the low 19 F signal elicited from these labeled cells, often leading to exceedingly long acquisition times using nonlocalized spectroscopy that may lead to compromised oxygenation status and viability.

Detection threshold in labeled cells in vitro.
Multiple CT cells (1, 0.75, 0.5, 0.25 million cells) were seeded, and labeled with FuGENE with NPs at a concentration of 10 mg/ml/million cells. The cells were subsequently suspended, maintained in IMDM media, and placed in multiple Eppendorf tubes for imaging using SPGR with the butterfly coil (  MRS and image-based concentration quantification. Implementation was achieved in phantom TFA solutions, in labeled CPCs, and in the post-mortem mouse following intracardial administration of CPCs using a setup that could be used in vivo. In vitro quantification: Direct MRS and SPGR-image-based quantification was implemented using the birdcage coil and the multivial TFA phantom described above, based on region-of-interest (ROI) estimation. A reference calibration curve was generated with four TFA solutions (25- To facilitate extension of the methodology to the post-mortem mouse, a 3D SPGR acquisition protocol was successfully tested in vitro (flip angle = 200˚) using the butterfly coil and reference (800k FuGENE-labeled CPCs) and test (~1.3 million CPCs before their injectionplease see (c) below) Eppendorf vials. The two vials were positioned on the surface of the butterfly coil at comparable positions to those used in the post-mortem case. T 1 and T 2 correction was achieved based on fully relaxed MRS or the steady-state, closed form, signal equations of SPGR and echo-SSFP, in accordance to Eqs 8 and 16 (S1 Appendix). CPC cell numbers and 19 F content from samples with unknown content were extrapolated based on known cell densities of reference standards (validated using Trypan Blue), against a 25 mM TFA phantom.
Post-mortem mouse: For these tests, the butterfly coil was used. The reference vial was positioned on the anterior thorax. The cell pellet position was approximately at the same level (along the inferior-superior direction of the animal, inclined~-45˚) as the injection site location. Estimation of the injected cell number was based on the ratio of the total 19 F MRI signal of the injected cells (upper mid-ventricular myocardium) and the reference CPC signal, obtained using the 3D SPGR adiabatic excitation (optimized in vitro-see (b) above) using a 220˚flip angle.
Post-mortem and in vivo animal models-Skeletal and cardiac muscle applications. To assess the feasibility and reproducibility of the imaging protocol, labeled cells (~1.5-2. In vivo cardiac muscle applications. Labeled cells (~1.5 million suspended in~50 μl of IMDM media) were injected in two C57BL/6J mice (male weight range = 20-30 g) that underwent thoracotomy, followed by recovery. Induction was achieved using 4% ISO, and the mice were maintained with 1.5-2.0% ISO, mixed in 100% oxygen.
Histology. Post-mortem histological evaluation was performed in mouse hearts on the same day following cellular implantation to confirm CPC localization. In brief, the excised hearts were dehydrated and fixed (in a 4% methanol-free formaldehyde solution), processed, embedded in paraffin, and stored. Serial transverse paraffin sections were subsequently cut, processed, and imaged on a Leica bright-field optical microscope, Image processing. 19 F images were processed in MATLAB (Mathworks, Natick, MA, USA) or ImageJ (NIH, Bethesda, MD, USA), and MRS in CSX (Johns Hopkins, USA) and IDL (Harris Geospatial, USA). The overlay of the 19 F and 1 H images was achieved using up-interpolation of 19 F MRI using bicubic splines followed by merging (at an opacity of 50%) in Ima-geJ. Flow cytometric data processing and exporting was achieved using FlowJo (FlowJo LLC, Version 10, Ashland, OR, USA).
The field uniformity of the birdcage coil was assessed following high-order shimming (B 0 homogeneity linewidths of~30-70 Hz), based on the signal variability (coefficient of variation [CV] = SD/mean) from multiple one-dimensional (1D) profiles spanning the central regions of multiple 2D images (along both image dimensions), acquired using high-concentration TFA phantoms. SNR was estimated as the mean intensity from selected ROIs divided by the background standard deviation.
Statistical analyses. All results are reported as mean±standard deviation (SD). Two-tailed Student's t-tests, were also used (XLSTAT, Addinsoft, New York) to determine whether labeling led to significant changes in transverse relaxation times.

Results
Highest field uniformity was achieved by the birdcage coil that yielded a coefficient of B 1 variation in the central 25×32×32 mm 3 region of the coil of 2.5% (Fig 1). Additionally, the butterfly yielded the highest SNR, while increased field penetration was noted when adiabatic (HS-AFP) excitation was used. Successful implementation of the adiabatic pulses is also demonstrated with 1 H and 19 F MRI of the multivial sensitivity phantom using the butterfly coil, indicating the ability to image 19 F concentrations of aqueous TFA solutions down to approximately 1 mM, at increased depths of penetration compared to nonadiabatic imaging ( Fig  1C-1E).
Normalized parametric SNR maps of the flip angle and ETL versus the normalized TR/T 1 values for SPGR, RARE, and SSFP imaging, are shown in Fig 2, depicting the zones from which acquisition parameters were selected for fast imaging.
To allow a direct, experimental comparison of SPGR, RARE, and SSFP sequences, solution phantoms (75-100 mM TFA) were used to perform a direct comparison using the homogeneous birdcage coil (Fig 3). Maximum SNR values (100 mM TFA) are elicited by the fid-SSFP sequence (139±10) that outperformed both the SPGR (106±7) and RARE (48±4) sequences. Based on TFA phantom imaging, SSFP achieved the highest mean SNR (1.3-and 2.9-fold higher than SPGR and RARE) (Fig 3).  (Table 1) measured primarily owing to the low 19 F signal of labeled cells, and the prohibitively long measurement times using CPMG (that would impose cellular viability risks). Labeling led to significant increases in longitudinal relaxation values in CPCs, compared to NPs in media solutions (p<0.00047 (labeled), p<0.001 (FuGENE-labeled), α = 5%).  Based on phantom imaging and MRS, the minimum detectable NP dose was 2.5 mg/ml (0.8mM) (birdcage) (Fig 4). As before, superior SNR performance is demonstrated by echo-SSFP (compared to SPGR) in 19 F imaging of the NP solutions, whereas lack of substantially improved signals is also documented for NP loading beyond 7.5 mg/ml for the studied CPCs ( Fig 4E).  (Fig 5A-5F), in justification of the lower sensitivity of detection of the labeled population by confocal microscopy (and correspondingly by MRI). The estimated percentage of viable labeled cells using confocal microscopy was 10% using simple labeling, and significantly increased upon use of FuGENE to 80% or higher [30]. A significant labeling heterogeneity was also noted for CPCs (Fig 5G-5I). Additionally, MRS/MRI allowed imaging and quantification of CPC NP concentrations (fully relaxed MRS for a label loading concentration of 7.5-10 mg/ml/million cells) (Fig 5J-5L) of the order of 0.3-0.5 mM (solenoid).

. T 2 values in labeled cells could not be
Image-based quantification (SPGR images, error of 3%, 10-100 mM) (Fig 6A-6C) confirmed linear signal-concentration dependence in TFA solutions (Actual concentration [mM] = 1.11×Estimated concentration [mM]-5.5, R 2 = 0.997), with detection limits for the butterfly/birdcage equal to~0.5 and 10 mM in imaging acquisitions that spanned~3 min). Fig  6F-6H confirmed that the cellular load detection threshold for the butterfly coil was approximately 0.5 million labeled CPCs (without FuGENE) ( Table 2) using the imaging protocol. More importantly, the ability to conduct in vitro, image-based cell quantification was confirmed using TFA phantoms (Fig 6A and 6B), labeled CPC ( Fig 6E) and macrophage cells (Actual concentration [mM] = 0.17×Estimated concentration [mM]+0.03, R 2 = 0.99), using fast imaging and MRS. The effort was extended successfully in the post-mortem mouse using the butterfly coil.
The applicability of the presented methodologies and optimization strategies for in vivo imaging are justified by the fast, post-mortem imaging of injected labeled CPCs in cardiac muscle (in vivo and post-mortem) and femoral areas (post-mortem) of the C57BL/6 mouse  19 F T 1 and T 2 relaxation values in phantoms and in labeled cells (immune, neural stem, and cardiac progenitor) from this and prior published studies. Significantly increased T 1 values (p<0.00047 (labeled), p<0.001 (FuGENE-labeled), α = 5%) were measured for the NP-labeled cells compared to NPs in solution. [37] (Fig 7). Reported findings were confirmed with histology, whereby cellular injection localization was identified using bright field histological imaging (Fig 7). From the quantification viewpoint, the in vitro measurement using 19 F MRI (based on the labeled CPC reference standard) estimated the number of labeled cells as 1.26 million based on the labeled CPC reference standard (compared to the actual number of 1.3 million estimated based on Trypan Blue). Only 1 million cells were quantified from the post-mortem images.  The discrepancy is attributed to injection losses (resuspension fluid loss before injection, and other injection losses).

Compound T 1 (ms) T 2 (ms) Field strength (T) Comment Reference
From the post-mortem and in vivo imaging viewpoints, the total 19 F signal of injected CPCs in mice (post-mortem vs. in vivo) yielded a CV of 24%, whereas the maximum SNR values were comparable (10±1.8, range = 8-11.8). The reported variability is a result of the expected discrepancies in the injected number of cells and injection losses in the four studied cases.
While Faber and Schmid [26] have recently reported theoretical and experimental comparisons of 19 F pulse imaging strategies, the approach adopted herein is mathematically rigorous and analytic, and includes reference to the explicit mathematical formulations, allowing theoretical comparison of SSFP, SPGR, and RARE sequences. Additionally, the 2D/3D comparison is also formulated mathematically, thereby justifying the efficiency and expected SNR improvements in 3D (as tabulated in Faber and Schmid [26]). Additionally, there are important practical implications relevant to the translatability and implementation of the bSSFP sequence for cardiac 19 F MRI that have not been addressed previously as they pertain to artifacts and SNR performance for exogenously administered SCs in the in vivo murine beating heart, while the explicit mathematical formulation for the signal of the SSFP sequence (fid vs. echo) is lacking (useful and relevant for quantification). Furthermore, there have been no prior direct, experimental comparisons of image SNR under controlled phantom conditions for SPGR, SSFP, and RARE. Prior publications on the performance of SPGR imaging in 19 F MRI, including the recent work by Faber and Schmid [26], have shown merit only for Ernst angle imaging. Nevertheless, in this work, the inefficient aspects of (slow) Ernst angle and the beneficial aspects of fast SC imaging are emphasized using SPGR.
Technical development of cardiac 19 F MRI using exogenously administered SCs has been lacking, while optimizations of MR imaging acquisitions in association with SCs have been limited [26].
More importantly, imaging times for adopted methodologies in most applications in vivo were excessively long (often > 60 min) [36] due to the low cellular label concentration and the choice of the imaging sequences, often becoming prohibitive for translational work. Interestingly, recent implementations of compressed sensing in 19 F MRI [41,42] have led to reduced imaging times, albeit limited by the low SNR of multinuclear in vivo studies.
We have presented the first in vivo cardiac 19 F MRI data from labeled CPCs injected in the murine myocardium. We have demonstrated feasibility and reproducibility of fast (of the order of a few minutes), in vivo cardiac 19 F-MRI of exogenously administered CPCs in the murine heart using SPGR/SSFP sequences.
The imaging protocol can be easily and readily adopted for any other labeling agent in preclinical work, and has potential for use in translational work. Additionally, the presented theoretical and experimental schemes are label-independent and can be readily applied to any other fluorinated compounds. using the solenoid coil showing excellent 19 F signal localization. (L) 19 F magnitude spectrum in labeled CTs using the solenoid coil (line broadening = 30 Hz, zero reference frequency set to the NP-labeled CT cell resonance).
https://doi.org/10.1371/journal.pone.0190558.g005 In contrast to imaging of nuclei other-than-protons (e.g., 23 Na MRI), where the intrinsically abundant sodium nucleus exhibits fast longitudinal relaxation (and where the optimal SNR is elicited at the Ernst angle favoring TR%T 1 ), 19 F MRI depends on exogenously administered NPs, emulsions, or labeled cells, with relatively long T 1 values and small T 2 /T 1 ratios. Despite the expected maximal SNR increases for SPGR sequences at long TRs, it is recommended that fast acquisitions are used in conjunction with averaging and with an appropriate choice of the flip angle, based on theoretical evaluations. This study has also assessed the cellular load detection thresholds (0.5 mM for the butterfly versus 10 mM for the birdcage coil) for 19 F NP labels, based on fluorinated phantom solution comparisons and fast acquisitions of the order of a few minutes at voxel resolutions of~2 μl or less.
Despite recent efforts to chemically modify the T 1 /T 2 characteristics of labeled NPs [43] to speed up acquisitions, such approaches require complex chemical syntheses. The PLGA-PFC NPs used in this study have elicited T 1 and T 2 values that are in agreement with prior reports [36], and a T 1 -effect is demonstrated post-cellular loading. The inability to quantify T 2 upon cellular loading is primarily attributed to the low labeling efficiency (and hence the low 19 F signal), and to the intracellular endosomal/vesicular packaging of the NP label leading to a relatively short transverse relaxation times (of the order of a few ms) [39,36]. Correspondingly, prolonged spectroscopic acquisition times will invariably impose additional issues in terms of the cellular oxygenation status (oxygen tension, leading to hypoxia [41]), and altered viability, compromising T 2 estimates. Given the limited in vivo/post-mortem visibility of labeled cells (without FuGENE), and the reported dependence of the T 2 values of PFCE-NPs on i) temperature [44,45], ii) label concentration [46], and iii) cell type, and the iv) extremely low 19 F signal of labeled cells, increased complexity and limited usefulness and consistency, are anticipated from the measurement of these values.
Our reported relaxation values in NP solutions (T 1 and T 2 ) are smaller in value than those recently reported by Colotti et al. [46] at 24˚C at 3T. This difference, however, is expected in  view of the B 0 -field relaxation dependence trends reported by de Vries et al. for 19 F emulsions [44]. Our successful labeling protocol was confirmed using flow cytometric and confocal microscopy validations. Based on validation studies presented herein, no additional benefits are expected by increasing the label concentration beyond 7.5-10 mg/ml per million cells, given the constancy of elicited signal responses in NP solutions using both the SPGR and SSFP acquisitions.
The minimum cellular detectable load (no FuGENE) was determined to be approximately 500k cardiac stem cells (or equivalently~10k cells per voxel) in fast acquisitions (~3-5 min) using the butterfly coil. This finding can be justified by the inefficient process of cellular label uptake in these cardiac stem cells in association with their much smaller cellular size (~30 μm 3 isotropic) compared to dendritic cells or macrophages. Evidently, FuGENE significantly decreases this detection limit, thereby achieving 19 F cardiac MRI [30].
Regarding quantification, we have demonstrated direct spectroscopy, and direct imagebased quantification of absolute 19 F concentration in TFA solutions, in labeled CPCs and macrophage cells, and injected cells post-mortem, with responses that scale linearly with increased fluorinated label or fluorine concentrations ex vivo. The ability to conduct both 19 F MRI/MRS provides an easy/efficient methodological pathway to study focal myocardial disease.
However, unlike prior quantification studies [6,47,48], a major limitation of the quantitative capacity of in vivo 19 F MRI is the fact that it cannot be applied to CPCs given their low and heterogeneous label uptake at insufficient levels for their MRI detection. To overcome such a limitation, we have validated an in vitro quantification scheme (in combination with adiabatic excitation) that can be adopted to in vivo applications, where a secondary reference phantom containing a known number of labeled CPCs can be used.
The present study is associated with various limitations, including the focus on two particular CPC types that consist of a heterogeneous cell mixture, and their low label uptake that ultimately hinders T 2 measurements. Given the stringent spatial requirements of the high-field bore system, temperature and oxygen tension effects cannot be easily monitored during relaxation measurements, or following cellular injection. Furthermore, our study has focused on PFCE with a single resonant peak, compared to other fluorinated compounds (e.g., PFOB) that exhibit complex, multi-peak responses.
Another major limitation is the lack of quantitative accuracy in vivo primarily owing to the stringent time limitations for additional data acquisitions for achieving adiabaticity and B 1 and motion corrections. Even still, an additional confounding factor that may limit quantification is the diminished viability of injected cells as a result of the hypoxic environment in which they are injected [1]. Correspondingly, evoked 19 F signal hyper-enhancements that may be attributed to viable cells, or released NPs from lysed cells (immediately following injection, or in tracking studies), ought to be interpreted with care. Such present the primary limitations in translating this work to a true experimental setting in humans.
Furthermore, the spatio-temporal effects of motion on 19 F cardiac MRI may need to be addressed in more detail. However, preliminary tests with the implemented protocols indicate lack of spatial discrepancies between ungated and gated 1 H MRI scans (voxel size = 0.2 μl) in normal mice. Correspondingly, possible motion effects between ungated and gated 19 F MRI MRS from the upper thorax showing the two isoflurane (ISO) and the labelled CT cell peaks. All 1 H images were acquired when the coil was tuned/matched at the 19 F resonance. (G) Indicative optical bright field histological image from the mouse heart in (D) above. The dotted square box indicates the area where cells were localized within the left ventricular myocardium. https://doi.org/10.1371/journal.pone.0190558.g007 acquisitions are expected to be minimal, considering the large voxel (voxel volume>2 μl), and the spatio-temporal averaging over the acquisition time intervals. Although gated cardiac 19 F scans are possible, they are prohibitively long (exceeding at least 30 min, thereby imposing beat-to-beat variability issues) and yield inadequate image SNR responses that would disallow NP detectability and localization.
Supporting information S1 Appendix. Theoretical background: MRI signal-to-noise ratio maximization: Theoretical and experimental considerations.