Comparison of 3He, 129Xe, inert fluorinated gas and propane.
Hyperpolarized gas MRI of the mouse lung is of great interest due to the urgent need for novel biomarkers for the assessment of respiratory-disease progression and development of new therapies. Small animal 3He lung MRI requires high-spatial-resolution imaging (<500 μm) to obtain acceptable images for visualization of all branches of lung microstructure from the mouse trachea to lung parenchyma. The use of conventional fast-gradient-recalled echo (FGRE) pulse sequences for high-spatial-resolution mouse lung imaging is challenging due to the signal loss caused by significant diffusion-weighting by the imaging gradients, particularly in larger airways where 3He diffusion is maximized. In this chapter, a modified FGRE approach called X-Centric is described for acquiring 3He mouse lung MRI. X-Centric is a center-out technique, allowing high-spatial-resolution, and high signal-to-noise ratio density-weighted imaging, as it is a short-TE method minimizing diffusion decay. Here, we also take advantage of a high-performance insertable-gradient-set interfaced with a clinical MRI system and a custom-built constant-volume ventilator to get the maximum benefits of X-Centric. High-spatial-resolution 3He X-Centric imaging was performed in a phantom and mouse lungs, yielding a nominal resolution of 39 μm and 78 μm respectively. We also demonstrate the feasibility of 129Xe/19F X-Centric MRI in a phantom and in rat lungs.
- 19F mouse
- alpha-1 antitrypsin deficiency
Inhaled hyperpolarized gas lung MRI  was proven to be useful for the observation and treatment planning of several pulmonary diseases including chronic obstructive pulmonary disease (COPD) [2, 3, 4, 5, 6], asthma, [7, 8, 9, 10, 11] and lung cancer . With the combined economic burden of COPD and asthma in Canada, Ontario being $5.7 billion (2011) , and lung cancer being the leading cause of cancer deaths worldwide,  there has been a growing interest in developing new lung imaging techniques such as hyperpolarized 3He and 129Xe MRI as well as hyperpolarized propane (1H) MRI [15, 16] and thermally polarized inert fluorinated gas (19F) MRI [17, 18, 19] to better understand various lung diseases.
Our main focus will be on 3He gas MRI of small animal models of lung disease [20, 21], as it is a promising tool to quantify airway and ventilation abnormalities associated with bronchoconstriction, airway narrowing, and subsequent bronchodilation [22, 23]. Small animal 3He lung MRI normally assumes high spatial resolution imaging (<500 μm ) in order to obtain acceptable images for the quantitative analysis and visualization of rodent airways and lung parenchyma. However, high-resolution hyperpolarized gas MR imaging has a number of signal-to-noise ratio (SNR) limitations due to the non-renewable nature of the non-equilibrium magnetization, requiring a minimized number of small flip angle radio-frequency (RF) pulses and quite short image acquisition times during a breath-hold which should be as short as a second in duration for the case of mouse lung imaging. Coupled with the issues of high-resolution imaging are the effects of fast diffusion of 3He (∼2 cm2/s) through strong spatial encoding gradients, which can completely destroy the MRI signal in the airways and significantly attenuate the signal in the lung parenchyma [24, 25, 26]. There is a need for an imaging method which incorporates specific hardware and novel rapid image acquisition approaches that minimize the signal loss due to all of these mitigating factors and permit fast and high spatial resolution 3He MR imaging of small animal airways and lung microstructure . Such imaging modalities can be used for observation of ventilation heterogeneity, such as in an ovalbumin asthmatic mouse model [22, 27], as well as for pediatric lung applications  often requiring much smaller field-of-view (FOV).
The projection-reconstruction (PR) approach is the well-known  acquisition method that minimizes 3He diffusion-induced signal decay, as it acquires the k-space center immediately following the excitation pulse and before the imaging gradient starts which avoids significant signal loss occurs . Both PR and spiral/cones acquisitions belong to free induction decay based or apparent transverse relaxation time (
An alternative to PR (which may suffer from poor edge resolution and sampling/reconstruction artifacts ) for high spatial resolution single-breath 2D imaging of small animal lung employs two separate excitations with inverted read-out gradients to acquire both halves of Cartesian k-space in a center-out fashion to minimize both diffusion-weighting [38, 39] and
In this chapter we present an X-Centric approach developed for a single breath-hold high spatial resolution 3He MRI of mouse lungs. The method takes advantage of a high performance insertable gradient set interfaced with a clinical MRI system and precise custom built constant-volume ventilator. The X-Centric approach is compared to partial-echo FGRE  on the basis of SNR efficiency over a range of spatial resolutions in a phantom . The robustness of the X-Centric technique for 3He in-vivo lung imaging of mice is demonstrated . The feasibility of the X-Centric approach to image dissolved hyperpolarized 129Xe in a phantom and with 19F gas in a rat lung  is also demonstrated in this chapter.
2. Theoretical background
2.1 3He MRI
Table 1 summarizes the physical properties of the 3He isotope along with other gases used for inhaled lung MRI  such as 129Xe [1, 41], inert fluorinated gases possessing 19F spins [17, 18, 19] and propane  which is proton-based (1H). It is well-known, that helium gas has the highest self-diffusion coefficient (
|Parameter||C3H8 (1H)||SF6 (19F)||3He||129Xe|
|Nuclear spin, I||1/2||1/2||1/2||1/2|
|Gyromagnetic ratio, (MHz rad/T)||267.5||251.6||203.7||74.5|
|Natural abundance (%)||100||95||1.4 × 10−4||26|
|Apparent transverse relaxation time, ||>1.0||0.5a/2.0b||1.5a/3.2b||2.8a/2.9c|
|Longitudinal relaxation time, ||5.0||30||90d||25d|
|Self-diffusion coef., ||10−3||0.033||2.05||0.061|
|Diffusion coef. in air, ||0.098||0.094||0.86||0.14|
|Chemical shift range, Δδ (ppm)||20||150||0.8||220|
|Ostwald solubility × 10−4, L||61||54||850||1700|
|Cost, per liter||$20||$20||$800||$20/200e|
2.2 Relaxation times
Generally, both longitudinal and apparent transverse relaxation  times (
This suggests that a small number of anoxic pre-breaths or wash-out breaths prior the actual 3He MRI measurements can be helpful for minimizing the oxygen concentration in lung prior to acquisition and administration of hyperpolarized gas. Table 1 suggests that the
2.3 Signal to noise ratio
An imaging sequence should be optimized to simultaneously ensure maximum SNR and minimal blurring in order to successfully perform high-resolution hyperpolarized 3He MR measurements. Understanding of the mechanisms that affect SNR and contribute to image blurring is the basis for the developing an imaging technique appropriate for hyperpolarized gas imaging. Generally, an expression for the SNR of a FGRE sequence can be expressed as :
where ΔxΔyΔz is the voxel size, NxNyNz is the number of k-space samples in each direction, BWread readout bandwidth and is the pulse-sequence-dependent function that determines the signal amplitudes at readout for the center of k-space :
As it can be seen, the signal amplitudes at readout depend on the density (ρ) and polarization (P) of 3He, flip angle (α), TE, repetition time (TR) and finally, diffusion attenuation  introduced by slice select and/or frequency encoding gradients. In order to simultaneously minimize
2.4 Flip angle
Let us consider each factor in Eq. (4) individually in order to optimize this function. As the hyperpolarized magnetization is not renewable due to the its non-equilibrium nature, it needs to be spent (flipped by RF pulses) very efficiently. A single 90o RF pulse can effectively waste all of the available magnetization. We start with the two approaches of setting flip angles in hyperpolarized gas MRI. The first method is the Constant Flip Angle (CFA) approach  or small flip angle approach (1o < flip angle <10o). Using a CFA approach, the optimal flip angle (αopt), which provides the highest signal for sequential phase encode ordering, is expressed as :
Unfortunately, the use of the optimal flip angle given by Eq. (5) leads to significant signal decay due to the constant-value RF pulses applied during k-space acquisition. As an example, for a 2D case where Ny = 128 and αopt = 7.1o the signal decays by ∼60% across the k-space). An undesirable consequence of such decay is reduced image resolution as result of the RF pulse “history” during k-space acquisition. Point-spread-function (PSF) simulations for a 2D case with center-out phase encode ordering (Figure 1) confirmed that the CFA approach leads to image blurring (Full Width at Half Maximum (FWHM) = 1.5 pixels) . This signal decay-induced image blurring can be eliminated by using a Variable Flip Angle (VFA) approach (1o < flip angle <90o) . VFA mitigates signal loss during image acquisition (i.e., blurring) by ensuring a constant signal for each phase-encode line by starting with a low flip angle and incrementing the flip angle of each RF pulse, line after line . The flip angle for each ith phase-encode line of a VFA sequence can be calculated from the following equation :
For a 2D case (Ny = 128), α1 = 5.1o and α128 = 90o, producing no CFA-like signal decay, and consequently, no image blurring due to RF pulse “history” . Simulations have shown that VFA only shows blurring of 1.2 pixels due to discrete sampling  and because of these benefits this method is preferable for high spatial resolution 3He imaging.
Eq. (4) suggests that signal decay due to diffusion is the main reason the MR signal degradation (assuming minimal
where γ is gyromagnetic ratio,
where δ is the diffusion gradient duration, Δ is a distance between two gradient pulses and
As 3He gas has the highest self-diffusion coefficient of the usable hyperpolarized gases and it is a physical constant, the diffusion weighting
The X-Centric approach requires two discrete acquisitions for each k-space line making the method twice as long in acquisition time compared to a conventional partial-echo FGRE approach for a chosen spatial resolution and matrix size [17, 27, 39]. Though, this is not a significant cost to achieve high-resolution imaging of the lungs of small animals, the scan time can be further reduced by using a partial-echo method in the phase-encoding direction [17, 27, 39]. Note, that the use of a partial-echo method in the phase-encoding direction makes the X-Centric scan time only 10% longer than the partial-echo FGRE sequence for a similar spatial resolution and matrix size .
2.7 Diffusion attenuation
According to Eq. (2), for the case of the high-resolution 2D 3He phantom imaging, the SNR should generally depend on the diffusion attention, image resolution, BWread and the first flip angle in the VFA scheme :
as Δx, Δy can be expressed through FOVs and Nx, and Ny Eq. (11) can be rewritten as
If one keeps FOV, BWread, 3He polarization and volume constant across all phantom measurements, then the SNR comparison between the partial-echo and X-Centric FGRE can be done by normalizing the experimental SNR by the first flip angle in the VFA scheme (α1) and matrix size (). Thus, for a SNR normalized by the respective first applied flip angle and matrix size, one can write the following expression :
where . This final SNRnor equation can be used for SNR comparisons of the phantom images obtained for different resolutions with the partial-echo and X-Centric FGRE methods; also this equation is similar to Eq. (10).
2.8 Phantom imaging
A plastic 10 mL syringe filled with hyperpolarized 3He was used to validate the X-Centric sequence for high spatial resolution imaging . Figure 5 shows a 2D coronal view whole projection phantom images obtained for partial-echo (top panel) and X-Centric (bottom panel) FGRE for five different resolutions. Images start from 64 × 64 matrix size and end with 512 × 512 from left to the right . The images indicate that at low nominal resolution (625 μm, 64 × 64 matrix size) there is no qualitative or visual difference in the SNR between the images obtained for the partial-echo and X-Centric FGRE. However, there is virtually no signal in phantom images obtained with partial-echo FGRE when nominal resolution is greater than 256 × 256 (78 μm nominal resolution). As expected, the signal intensity of the images obtained with X-Centric gradually decays with increasing nominal resolution and decreasing flip angle (because the first pulse for VFA is a function of resolution, Eq. (7)) but even for 512 × 512 matrix size (39 μm nominal resolution) it does not go below the noise floor as is seen in the case of the partial-echo FGRE. The SNRnor dependence on resolution (bottom axis) and/or b-value (top axis, partial-echo case only) for the phantom images obtained with the partial-echo (solid squares and solid line) and X-Centric FGRE (solid circles and solid line) is plotted in Figure 6. Clearly, the SNRnor obtained for the X-Centric did not depend on image resolution, i.e., diffusion attenuation was minimal even for very high image resolution (39 μm). The experimental phantom results were consistent with the predicted SNR (SNRtheor) calculated based on Eq. (10) for the partial-echo (open squares and dashed line) and X-Centric FGRE (open circles and dashed line). Eqs. (10) and (13) indicate that for the case of unrestricted 3He diffusion (
|3He MRI||129Xe MRI||19F MRI|
|Flip angle||3.2o (VFA)||6.4o (CFA)||70 o (EA)|
|FOV (mm2)||20 × 20||250 × 250||60 × 60|
|Matrix size||256 × 160||128 × 80||64 × 40|
|Number of averages||1||1||60|
|Pixel size (mm3)||0.078 × 0.078||1.9 × 1.9 × 300||0.94 × 0.84|
|Scan time (s)||1.1||>8||>10|
|MRI scanner||GE 3 T MR750||Philip. 3 T Achieva||Philip. 3 T Achieva|
3. High spatial resolution imaging of mouse lung
3.1 Animal preparation
The animal protocol was approved by the Animal Use Subcommittee of the University Council on Animal Care at the University of Western Ontario, London, Ontario, Canada and the Animal Care Committee at Merck Frosst Canada, Kirkland, Quebec, Canada. Each mouse (BALB/c) was pretreated with Midazolam (1 mg/kg) intraperitoneally (i.p.) 5 min before anesthesia with ketamine (95 mg/kg, i.p.) and xylazine (6.4 mg/kg, i.p.). Following anesthesia, the tail vein was cannulated (26 GA Abbott catheter) to maintain anesthesia with ketamine (30 mg/kg) and xylazine (2 mg/kg) every 40 min. To allow artificial mechanical ventilation, the trachea was cannulated by tracheotomy using an 18 GA Teflon catheter. A line for administration of pancuronium (0.8 mg/kg, i.p.) was established using a 26 GA cannula to allow for injections after the mouse was being ventilated inside the scanner. Other physiological instrumentation included ECG and a rectal temperature probe (SA Instruments Inc., Stony Brook, NY).
It was a very challenging task to design a ventilator to ventilate mice inside an MRI because of the restrictions due to the high magnetic field environment and the requirement to accurately deliver very small tidal volumes of gas to the lungs of mice at a high ventilation rate. Under spontaneous breathing, the mouse maintains its oxygenation with a tidal volume less than 0.1 ml [54, 55]. Because of the dead space volume within the tubing necessary for connecting mouse airways to the ventilator, the mouse was normally ventilated with tidal volume of 0.2 ml (8 ml/kg). In this range, the amount of the ventilating gas (or fresh gas) getting compressed in the gas line could significantly reduce the portion of fresh gas entering the lungs for gas exchange. Without a proper correction for this effect, adequate ventilation would not be achieved at this ventilating setting. In many cases, the ventilation was set with a much larger tidal volume (>10 ml/kg) at a slower ventilation rate (60–90 bpm). Since lung volume under spontaneous breathing is only 0.3–0.4 ml , the larger tidal volume approach could overextend the airways. As a result, many mice were observed to expire shortly after the ventilation, impeding the completion of studies.
To resolve these problems, we modified a flexiVent ventilator (Figure 7, Scireq, Montreal, QC, Canada) that incorporated real-time monitoring of volume and pressure to allow corrections of the gas compression effect. To minimize the gas line volume and ensure the effectiveness of ventilation, a non-metal swing valve was designed and placed close to the animal (inside the scanner) to regulate the gas flow. Different than the pneumatic valves normally used in the magnetic environment, the swing valve was connected to two driving solenoid valves (which reside far away from the high magnetic field) via two multiple-section light-weight carbon fiber rods. This mechanical design enabled fast open/close activations. The activation of a solenoid valve pulls the swing valve to close the ports on one side and open the ports on the other side. Two solenoid valves were activated in sequence to generate a swing motion that opens and closes the ports on the respective sides alternately. This swing motion was synchronized with the plunging motion of the piston that drives the gas into the lung to form a ventilation cycle. Because of the fast response of the swing valve, the modified flexiVent ventilator is capable of performing the forced oscillation method to access pulmonary mechanics. To minimize oxygen-induced depolarization (
3.3 3He polarization and delivery
3He gas was polarized using an optical pumping spin-exchange system (Helispin™, GE Healthcare, Durham, NC). The 3He gas was typically polarized for at least 8 hours to achieve a polarization over 35%. Prior to transferring the hyperpolarized 3He gas to a 150 mL Tedlar bag (Jensen Inert Products, Coral Springs, FL), the bag was carefully washed with N2 gas three times to eliminate any oxygen contamination. Following the gas filling, the 3He bag was placed inside a custom-made plexiglass chamber and connected to a 3He intake line of the ventilator. The chamber was then placed within the homogenous B0 field inside the scanner to conserve 3He polarization (
To validate the X-Centric imaging method, 2 mL at 1 atm of the hyperpolarized 3He gas was drawn into a 10 mL syringe for phantom testing .
3.4 MRI hardware and sequence parameters
Hyperpolarized 3He MR imaging was performed on a GE clinical scanner (3 T, Excite 12.0) which was converted using a home-made insert gradient set (Figure 8, maximum gradient at 50 G/cm, 17 cm in diameter, slew rate = 2100 mT/m/s)  to allow high spatial resolution imaging for mice. This high performance insert gradient set was used for phantom and mouse imaging . A quadrature birdcage RF coil for mice (Figure 9, 3 cm in diameter and 6 cm in length, Morris Instruments, Ottawa, ON) was used for 3He imaging (97.3 MHz) . The coil was driven using a 8 kW AMT 3 T90 RF power amplifier (GE Healthcare, Waukesha, WI) using a manufacturer-supplied T/R switch and preamplifier tuned to 97.3 MHz. Typical rectangular 90° RF pulse lengths were 100 μs . A VFA RF pulse trajectory was employed to reduce blurring due to RF de-polarization [47, 52]. Center-out sampling in the phase-encoding direction was used . The VFA pulses were calibrated as described in .
3.5 Image reconstruction
A custom-made IDL6.4 (ITT Visual Information Solutions, Boulder, CO) routine was used for off-line Fast Fourier transformation (FFT) of the k-space data. For reconstruction of the X-Centric data, a combination of both halves of k-space was applied prior to the inverse Fourier transformation as follows. To form a full line of k-space data in the read-out direction, the two half-echo data sets were combined . Each half-echo data set included two additional points (BWread = 62.5 kHz dwell time was 8 μs) prior to
Acquisition of a few extra
3.6 In-vivo imaging
X-Centric sequence parameters for the highest image resolution (78 μm) are shown in Table 2.
The larger SNR in the major airway (X-Centric mouse lung images) compared to the lung parenchyma is likely due to the higher 3He density in trachea and reasonably short TE used in X-Centric sequence (0.7 ms ).
It is useful to calculate the SNR efficiency (ϒ) of the X-Centric and partial-echo FGRE acquisitions was using the following equation :
One of the limitations of the X-Centric method is the need to use the FFT for image reconstruction from two half-echo k-space samples rather than a single k-space. Instead, one can use a single half-echo k-space and Partial Fourier Reconstruction (PFR) . To test this idea we used the Projection Onto Convex Sets (POCS) method  to reconstruct a phantom image for 130 × 256 (2 + 128 × 256) matrix size (Figure 5b). Figure 12 shows phantom images obtained with PFR (left column) and FFT (right column) for three different matrix sizes. The following SNR values were obtained for the presented images: SNR = 39 (160 × 256, PFR); SNR = 54 (160 × 256, zero-filling in x-direction and then FFT); SNR = 45 (130 × 256, PFR); and SNR = 58 (256 × 160, zero-filling in y-direction and then FFT). The results confirm that the FFT with zero-filling approach gives better SNR than the PFR method for 160 × 256 matrix (39 vs. 54). We have calculated the efficiency of half-echo acquisition (130 × 256, SNR = 45) reconstructed with PRF and X-Centric (256 × 160, 62.5% under-sampling in y-direction, SNR = 58) reconstructed with FFT in order to compare imaging methods. Thus, the calculated efficiency ratio is 0.87 (Eq. (14), with
3.8 Future role of X-Centric
The focus of this chapter is a presentation of a fast and SNR efficient imaging method for high spatial resolution imaging of mouse lungs that include both the airways and lung parenchyma. Such a technique can be potentially used for mapping the morphological changes and ventilation heterogeneity associated with rodent models of asthma or ovalbumin-challenged model (OVA). Figure 13 shows an example of ventilation heterogeneity maps obtained for a sham mouse (right column) and OVA treated mouse (left column). The figure compares increased (a and c) and decreased (b and d) ventilation heterogeneity indexes from a sham mouse (left column) and an OVA mouse (right column) under similar increases in airways resistance (104% vs. 93%, respectively) and elastance (21% vs. 27%, respectively). The white and red lines in the figure outline the lung contour before and after the Methacholine inhalation (MCh) challenge, respectively. Airway resistance and tissue elastance were measured immediately following each scan. Lung images were then analyzed using local standard deviation (SD) in sliding (9 × 9 pixels2) ROIs. The ventilation heterogeneity index (
The post-MCh lung function data were normalized by the pre-MCh values to give the percent change. The sham mouse and an OVA mouse data are shown only as an example of the potential X-Centric pre-clinical application. However, data suggested that ventilation heterogeneity following MCh challenge may be commensurate with traditional lung function testing to access the airway hyper-responsiveness in the OVA mouse model. The observed heterogeneity in ventilation distribution could potentially provide a novel endpoint to study disease modification in asthma, as well as for better diagnosis and classification of asthma in the clinic.
The X-Centric method described here should also be beneficial for clinical hyperpolarized lung MRI and other applications where short-TE techniques are needed. Table 2 helps to estimate the acquisition time for BWread = 62.5 kHz and FOV = 200 × 200 mm2 (pediatric study ) for two matrix sizes frequently used in human studies : 128 × 128 × 14 slices and 256 × 256 × 14 slices (3D case) yielding resolutions of 3.1 × 3.1 × 15 mm3 and 1.6 × 1.6 × 15 mm3 at FOV = 400 × 400 × 15mm3. With the assumption of partial sampling in the phase-encoding direction and taking into account that the maximum strength of the clinical gradient system is typically 4–5 G/cm with a maximum slew rate of 200 T/m/s , one can calculate the acquisition to be around 5 s for a 128 × 128 matrix and 16.7 s for a 256 × 256 matrix. As the 256 × 256 matrix at a FOV of 20 × 20 cm2 corresponds to 781 μm nominal resolution, this should lead to 70% signal decay in airways due to diffusion-weighting according to  for the case of full-echo FGRE. The X-Centric approach could eliminate 90% of the signal loss due to incidental diffusion-weighting in the major airways. Parallel imaging approaches  or compressed sensing approaches [65, 66] can also speed up data acquisition even more and are a straight-forward extension of the X-Centric approach described here.
Until now, X-Centric has not been used in any pediatric studies; nevertheless, hyperpolarized gas MRI is gradually becoming an important research tool, allowing for longitudinal observation of lung diseases in children such as Cystic Fibrosis (CF). Figure 14 shows 3He MRI lung images obtained from a 17-month old patient with CF. Note, the regions of hypo-intensity are difficult to see at this resolution, but using X-Centric, one should be able to gain diagnostic value and better enhance the contrast in these regions. We believe that X-Centric can be very useful for not just CF but the longitudinal observation of other lung diseases like Bronchopulmonary Dysplasia (BPD). Abnormal prematurity at birth (<28 weeks gestation) frequently requires neonatal intensive care and mechanical ventilation due to respiratory distress syndrome due to limited or lack of lung surfactant and incomplete lung development. Ventilator-associated lung injury, along with the extreme prematurity of the lungs themselves, causes abnormal and irregular lung growth (BPD) . The relatively short signal life-time caused by the need to scan newborns inside their neonatal incubators (causing
4. Use of X-centric with other MR-visible nuclei
4.1 Gas phase 129Xe MRI
3He is a rare and expensive isotope (Table 1), so a worldwide clinical translation of the 3He MRI method is quite questionable, especially when there is a relatively inexpensive alternative—129Xe MRI. In contrast to the 3He isotope, the 129Xe isotope is much more abundant (∼26% natural abundance) and cost efficient (Table 1). In addition, recent improvements in the xenon polarization process have enabled sufficient increase in the level of 129Xe polarization (>30%) and production volume of the polarized gas (∼2 L per hour) [68, 69]. Presently, 129Xe lung MRI is translating towards a clinical tool, and it has recently been used as a clinical tool (along with 3He MRI) in the United Kingdom . The North-American xenon consortium  expects 129Xe MRI to be the Food and Drug Administration approved in the first quarter of 2020.
Presently, the main directions of 3He MRI are pediatric and neonatal studies (using relatively small doses of 3He) mainly due to the cost of the 3He isotope and the high quality of the 3He lung images. In turn, 129Xe MRI is much more suited for adult lung imaging; however, the low SNR of 129Xe images (compared to 3He) impede the development of many novel acquisition schemes that are highly sensitive to SNR , such as isotropic voxel static-ventilation imaging and accelerated multiple
4.2 Dissolved phase 129Xe MRI
Unlike its counterpart 3He, 129Xe is highly soluble in a variety of solvents and biological materials, and exhibits a large range of chemical shifts within these distinct chemical environments (Ostwald solubility coefficients of 0.17 ). Of particular interest is dissolved phase 129Xe residing in and exchanging with lung tissue (tissue barrier) and red blood cells (RBC) [76, 77, 78]. The exchange can be measured using MRI and it is also possible distinctly resolve and quantify 129Xe within the tissue barrier and RBC. Exploiting these signals for diagnostic imaging purposes can be difficult since for example, only about 2% of inhaled xenon dissolves into the tissue barrier and RBC [76, 79], and the corresponding
The feasibility of the X-centric method was initially demonstrated using 3He for high-resolution rat lung imaging  indicating its ability to overcome diffusion induced signal attenuation within the trachea and substantially reducing TE and
Figure 17b shows dissolved phase hyperpolarized xenon obtained from within a water-filled resolution phantom (Figure 17a) on a clinical 3 T MRI scanner (maximum gradient strength was 5 G/cm, sequence parameters are shown in Table 2). The 129Xe image shows that X-Centric can be used to image dissolved phase xenon efficiently (within a breath-hold) and with reasonable signal-to-noise ratio (SNR = 12). After reconstruction, the smallest geometric features of the resolution phantom in the X-Centric image are discernable (yellow arrows). It should be noted, that the available SNR achievable when imaging dissolved phase 129Xe principally depends on concentration, polarization and relaxation times and even physiological parameters such as blood flow and blood oxygenation. The quality of the polarizer, solubility of the solvents used and chemical environments all play a role. Nevertheless, techniques employing significant reductions in echo-times can deliver higher SNR than is typically achievable for spins within environments where their transverse relaxation times are significantly shortened. Though, the phantom dissolved-phase spins had estimated
4.3 Static-ventilation 19F MRI
The limited use of the hyperpolarized 3He/129Xe MRI imaging modality can be partially explained by its need for utilizing expensive isotopes and polarizers, the latter requiring specially-trained personnel. Inert inhaled fluorinated gas MRI can be a promising and less expensive (Table 1) alternative to hyperpolarized 3He/129Xe MRI. 19F MRI does not require the use of rare isotopes (100% naturally abundant, Table 1) or expensive polarizers. 19F gases (such as perfluoropropane (C3F8) and sulfur hexafluoride (SF6) [17, 18]) can be imaged using high field (≥1.5 T) clinical MRI scanners; however, a multi-nuclear amplifier and receiver along with RF coils tuned to the 19F frequency are still required [18, 19]. Fluorine-19 has a large gyromagnetic ratio (∼95% of 1H, Table 1), both SF6 and C3F8 have several 19F atoms per molecule, and have relatively short
4.4 Dynamic 19F MRI
19F gas MRI also has a high potential for successful dynamic lung imaging [88, 90] due to the fact that fluorinated gases can be mixed with O2 which helps to restore the initial magnetization faster (decreasing
where 0 < β ≤ 1, and
The remaining obstacle preventing wide-spread high-resolution 19F MRI with or without X-Centric is the generally low SNR of 19F lung images. In this particular case, the further improvement of the image quality strongly depends on a hardware component, specifically from RF coil systems. It has been recently demonstrated that the use of a multi-channel phased-receive array can significantly improve the image quality of the 19F MRI lung images . This is a very promising result suggesting that the combination of the X-Centric method with advanced array-type RF coil system will permit future high-resolution 19F MRI human lung imaging in both static-ventilation and dynamic studies.
In summary, hyperpolarized 3He X-Centric MR imaging with high spatial resolution was proven to be a robust technique in phantom and mouse lung measurements yielding a nominal resolution of 39 μm and 78 μm respectively. In particular, mouse major airways with less restricted diffusion of 3He (
Another beneficial feature of the X-Centric sequence is the significantly reduced echo-time, so the method can be used for imaging of nuclei with short signal life-times such as hyperpolarized 129Xe dissolved within the lung tissue barrier and/or Red Blood Cells, and inert fluorinated gas (SF6) inside the lungs. In both cases the
The authors would like to thank the following individuals for assistance with MRI experiments and data analysis: Grace E. Parraga, Giles E. Santyr, Mitchell S. Albert, Marcus J. Couch, Tao Li and Iain K. Ball. A special thanks to Ben T. Chen for providing support for the hyperpolarized 3He gas, animals, and ventilator.
Conflict of interest
The authors do not have any conflict of interest.
We thank Michael Völker for providing the MatLab code (Projection onto Convex Sets (POCS), 2013) for image reconstruction. We thank Abascal et al. for providing the MatLab code (Signal Decay Into the Reconstruction (SIDER), 2017) for image reconstruction.
acceleration factor bandwidth laboratory-bred strain of the house mouse bronchopulmonary dysplasia perfluoropropane cystic fibrosis constant flip angle chronic obstructive pulmonary disease self-diffusion coefficient fluorine-19 fast Fourier transformation fast gradient recalled echo field of view full width at half maximum general electric hydrogen-1 helium-3 methacholine inhalation challenge magnetic resonance imaging oxygen ovalbumin-challenged model partial Fourier reconstruction projection onto convex sets red blood cells region of interest radio frequency standard deviation post-methacholine challenge SD pre-methacholine challenge SD hexafluoride signal to noise ratio normalized signal to noise ratio theoretical signal to noise ratio longitudinal relaxation time apparent transverse relaxation time total scan (imaging) time echo time transmit and receive repetition time variable flip angle ventilation heterogeneity index xenon-129
laboratory-bred strain of the house mouse
constant flip angle
chronic obstructive pulmonary disease
fast Fourier transformation
fast gradient recalled echo
field of view
full width at half maximum
methacholine inhalation challenge
magnetic resonance imaging
partial Fourier reconstruction
projection onto convex sets
red blood cells
region of interest
post-methacholine challenge SD
pre-methacholine challenge SD
signal to noise ratio
normalized signal to noise ratio
theoretical signal to noise ratio
longitudinal relaxation time
apparent transverse relaxation time
total scan (imaging) time
transmit and receive
variable flip angle
ventilation heterogeneity index