Figures
Abstract
Myocardial tissue characterization using T2* relaxation mapping techniques is an emerging application of (pre)clinical cardiovascular magnetic resonance imaging. The increase in microscopic susceptibility at higher magnetic field strengths renders myocardial T2* mapping at ultrahigh magnetic fields conceptually appealing. This work demonstrates the feasibility of myocardial T2* imaging at 7.0 T and examines the applicability of temporally-resolved and high spatial resolution myocardial T2* mapping. In phantom experiments single cardiac phase and dynamic (CINE) gradient echo imaging techniques provided similar T2* maps. In vivo studies showed that the peak-to-peak B0 difference following volume selective shimming was reduced to approximately 80 Hz for the four chamber view and mid-ventricular short axis view of the heart and to 65 Hz for the left ventricle. No severe susceptibility artifacts were detected in the septum and in the lateral wall for T2* weighting ranging from TE = 2.04 ms to TE = 10.2 ms. For TE >7 ms, a susceptibility weighting induced signal void was observed within the anterior and inferior myocardial segments. The longest T2* values were found for anterior (T2* = 14.0 ms), anteroseptal (T2* = 17.2 ms) and inferoseptal (T2* = 16.5 ms) myocardial segments. Shorter T2* values were observed for inferior (T2* = 10.6 ms) and inferolateral (T2* = 11.4 ms) segments. A significant difference (p = 0.002) in T2* values was observed between end-diastole and end-systole with T2* changes of up to approximately 27% over the cardiac cycle which were pronounced in the septum. To conclude, these results underscore the challenges of myocardial T2* mapping at 7.0 T but demonstrate that these issues can be offset by using tailored shimming techniques and dedicated acquisition schemes.
Citation: Hezel F, Thalhammer C, Waiczies S, Schulz-Menger J, Niendorf T (2012) High Spatial Resolution and Temporally Resolved T2* Mapping of Normal Human Myocardium at 7.0 Tesla: An Ultrahigh Field Magnetic Resonance Feasibility Study. PLoS ONE 7(12): e52324. https://doi.org/10.1371/journal.pone.0052324
Editor: Wolfgang Rudolf Bauer, University Hospital of Würzburg, Germany
Received: July 23, 2012; Accepted: November 16, 2012; Published: December 14, 2012
Copyright: © 2012 Hezel et al. This is an open-access article distributed under the terms of the Creative Commons Attribution License, which permits unrestricted use, distribution, and reproduction in any medium, provided the original author and source are credited.
Funding: This study was founded through institutional funding provided by the Max Delbrück Centrum for Molecular Medicine. The funders had no role in study design, data collection and analysis, decision to publish, or preparation of the manuscript.
Competing interests: TN is founder of MRI.Tools GmbH. This does not alter the authors' adherence to all the PLOS ONE policies on sharing data and materials. The authors would also like to point out that there are no patents, products in development or marketed products to declare.
Introduction
Emerging cardiovascular magnetic resonance (CMR) imaging applications include T2* relaxation sensitized techniques, which are increasingly used in basic research and (pre)clinical imaging. Methodological developments in T2* sensitized imaging [1]–[4] and simulations of myocardial vasculature [5], [6] have been indispensable. Applications include investigation of the microstructure of the isolated rat heart [7], detection of myocardial ischemia [8]–[14], probing of vasodilator or dipyridamole-induced changes in myocardial perfusion [15]–[18], visualization of scarred myocardium [19], imaging of capillary recruitment [20] and assessment of tissue oxygenation related to endothelium-dependent blood flow changes [21]. T2* mapping has also been shown to be of substantial clinical value for the ascertainment of myocardial iron levels [22]–[29].
The most widely used methods for T2* mapping are echo planar imaging (EPI) and gradient echo based techniques. Unlike conventional CINE gradient echo imaging, the relatively strong T2*-weighting required to make gradient echo sequences sensitive to changes in magnetic susceptibility asks for a long evolution time (TE) between RF excitation and data acquisition. Consequently, gradient echo based myocardial T2* mapping is commonly restricted to a single slice and single cardiac phase that can be accommodated in a single breath-hold at 1.5 T and 3.0 T [30]–[32].
The linear relationship between magnetic field strength and microscopic susceptibility [33]–[35] renders it conceptually appealing to pursue myocardial T2* mapping at ultrahigh magnetic field strengths. Realizing the opportunities and challenges of T2* mapping, this pilot study demonstrates the feasibility of ultrahigh field susceptibility-weighted myocardial imaging and examines its applicability for temporally-resolved and high spatial resolution myocardial T2* mapping at 7.0 T. To meet this goal, the applicability of 2D spoiled gradient-echo multi-echo based techniques for T2* mapping at 7.0 T is closely investigated in phantom experiments. The feasibility of gradient-echo multi-echo based techniques for fast CINE T2* mapping of the human heart is demonstrated at 7.0 T. We also present the suitability of this technique for high spatial resolution myocardial T2* mapping by using thin slices (slice thickness = 2.5 mm) and in-plane spatial resolution of (1.1×1.1) mm2. Our initial volunteer studies serve as a mandatory precursor to a broader clinical study. The merits and limitations of T2* mapping using 2D spoiled gradient-echo multi-echo imaging at 7.0 T are discussed and implications for cardiac MR at 7.0 T are considered.
Methods
MR-Hardware
Imaging was conducted using a 7.0 T whole body MR scanner (Magnetom, Siemens Healthcare, Erlangen, Germany) equipped with a gradient system (Avanto, Siemens Healthcare, Erlangen, Germany) capable of supporting a slew rate of 200 mT/m/ms and a maximum gradient strength of 40 mT/m. A 16 channel transmit/receive coil array was used for excitation and signal reception. The coil was designed for cardiac imaging and comprises an anterior and posterior former, each laid out on a two-dimensional 2 by 4 grid of loop elements. For further details about the coil please see [36], [37]. An MR stethoscope (EasyACT, MRI.TOOLS GmbH, Berlin, Germany) was used for cardiac triggering [38], [39].
T2* Mapping Techniques
Myocardial T2* mapping is commonly conducted with cardiac triggered, segmented multi-echo spoiled gradient echo (ME) techniques that use breath-held acquisitions for respiratory motion compensation [8], [9], [15], [18], [40]. In this work, various ME configurations have been used for (i) time resolved CINE and for (ii) single cardiac phase acquisitions (Figure 1).
A). Conventional multi-echo (ME) gradient echo for single cardiac phase myocardial T2* mapping. Multiple echoes are acquired after excitation to obtain a set of T2* weighted images. The competing constraints of inter echo time and spatial resolution inherent to the ME approach are addressed by the B) interleaved multi-shot multi-echo (MS) gradient echo technique. In MS a set of excitations is employed together with echo interleaving echoes to acquire a set of T2* weighted images. C) The multi-breath-hold multi-echo (MB CINE) gradient echo technique allows myocardial CINE T2* mapping by interleaving the echoes over several breath-holds. For benchmarking D) multi-echo CINE (ME CINE) gradient echo and E) multi-shot multi-echo CINE (MS CINE) were applied for T2* mapping in phantom studies. To guide the eye vertical dashed lines refer to k-space lines. Vertical solid lines refer to cardiac phases. A unipolar readout using gradient flyback was applied for all strategies.
For single cardiac phase imaging, the acquisition period is commonly placed into end-diastole, which limits the viable window of data acquisition to 100 ms to 200 ms. Data acquisition is segmented over a series of cardiac cycles with each segment acquiring a set of echoes during the quiescent interval (Figure 1a). The number of segments and echoes per segment are dictated by the longest TE used for T2* weighting as outlined in Figure 1a. To avoid T2* errors due to signal modulations induced by fat-water phase shift, it is essential to choose echo times where fat and water are in-phase [4]. At 1.5 T and 3.0 T TE increments equivalent to the fat-water shift are 4.4 ms and 2.2 ms, respectively. At 7.0 T this inter echo time is 1.02 ms, which is beneficial for rapid multi-echo acquisitions. However, if a larger data matrix size is needed for high spatial resolution T2* mapping, the readout/acquisition window can easily exceed 1.02 ms even when short dwell times are used. Consequently, it is elusive to accomplish inter-echo time increments of 1.02 ms for high spatial resolution T2* mapping at 7.0 T using sequential multi echo gradient echo imaging.
CINE T2* mapping might be feasible at 7.0 T assuming a T2* reduction at 7.0 T versus 1.5 T and 3.0 T and considering the relationship between proper T2* weighting and range of echo times to be covered. To this end, a maximum TE = 10 ms would be compatible with the needs of CINE imaging but would also provide sufficient coverage of the T2* decay at 7.0 T.
For all these reasons, two imaging strategies were employed at 7.0 T:
- Interleaved multi-shot multi-echo (MS) gradient echo technique for single cardiac phase myocardial T2* mapping (Figure 1b). This approach addresses the competing constraints of inter echo time and spatial resolution of the ME approach by adding more excitations and by interleaving the echoes.
- Multi-breath-hold multi-echo (MB CINE) gradient echo technique for CINE myocardial T2* mapping (Figure 1c). This approach runs the trait that all k-space lines required to form the final image for a given echo time are acquired in a single breathhold.
Phantom Studies
For the evaluation of MS and MB CINE T2* mapping strategies, phantom experiments were conducted using a long T2* and medium T2* phantom. For long T2*, a cylindrical water phantom (diameter = 15 cm) containing an agarose copper sulfate solution (4 g CuS04 +2 g NaCl +2 g agarose dissolved in 1.0 l H2O) was used. A glass capillary (inner diameter = 0.5 mm) filled with air and a tube (inner diameter = 5 mm) filled with water were placed inside the phantom to create strong susceptibility gradients of limited spatial extension within the uniform phantom. For the medium T2* phantom, a cylindrical water phantom (diameter = 8 cm) containing agarose (5 mg agarose dissolved in 250 ml H2O) was used. T2* was reduced by ultrasmall superparamagnetic iron oxide particles (500 µl Molday ION (10 mg Fe/ml), BioPal, Worchester, USA), which afforded a T2* of approximately 20 ms.
In the phantom experiments, MS (Figure 1b) and MB CINE (Figure 1c) T2* mapping strategies were benchmarked against other T2* mapping techniques, which are already established at 1.5 T and 3.0 T but are unsuitable for myocardial T2* mapping at 7.0 T due to echo time and acquisition time constraints. These reference methods include:
- conventional multi-echo (ME) gradient echo for single cardiac phase T2* mapping (Figure 1a).
- multi-echo CINE (ME CINE) gradient echo (Figure 1d).
- multi-shot multi-echo CINE (MS CINE) gradient echo (Figure 1e).
For phantom T2* mapping, an image matrix of 320×240, a field of view of (360×270) mm2, an in-plane resolution of (1.1×1.1) mm2, and a slice thickness ranging from 2.5 mm to 8 mm were used. A unipolar readout using gradient flyback was applied together with echo times ranging from 2.04 ms to 10.20 ms. This approach results in nine equidistant echoes with an inter echo time of 1.02 ms with the exception of ME and ME CINE due to gradient switching induced peripheral nerve stimulation constraints. For ME and ME CINE 6 echoes with an inter echo time of 3.06 ms and a TEmin = 2.04 ms were used. For the MS, MS CINE and MB CINE techniques three excitations together with 3 echoes were used to ensure an inter-echo time of 1.02 ms. With the first excitation echo 1, 4 and 7 were acquired. The second excitation covered echo 2, 5 and 8 while echo 3, 6, and 9 were recorded after the third excitation. A simulated heart rate of 60 bpm was used for prospective triggering of the phantom experiments.
Ethics Statement
For the in vivo feasibility study, 8 healthy subjects (mean age: 27±3 years, 5 females, mean BMI: 24 kg/m2, mean heart rate: 78 bpm) without any known history of cardiac disease were included after due approval by the local ethical committee (registration number DE/CA73/5550/09, Landesamt für Arbeitsschutz, Gesundheitsschutz und technische Sicherheit, Berlin, Germany). Informed written consent was obtained from each volunteer prior to the study.
Volunteer Studies
For each volunteer, slice positioning was carried out following international consensus by the same technician to omit inter-operator variability. Myocardial T2* mapping was conducted using the MS and the MB CINE imaging strategies for all subjects. For this purpose, a mid-ventricular short axis view and a four chamber view were used. Imaging parameters were set to: acquisition data matrix = 256×224, FOV = (288×252) mm2, in-plane resolution = (1.1×1.1) mm2, slice thickness = 4 mm if not otherwise stated. A nominal flip angle of α = 20° has been applied; and electro magnetic field (EMF) simulations using the human voxel model “Duke” were conducted for transmission field shaping to enhance B1+ uniformity across the heart [36], [37], [41]. Echo times ranging from 2.04 ms to 10.20 ms were applied. Three excitations together with 3 echoes per excitation were used to ensure an inter-echo time of 1.02 ms. With the first excitation, echo 1, 4 and 7, were acquired. The second excitation covered echo 2, 5 and 8. Echo 3, 6, and 9 were recorded after the third excitation. Moderate acceleration (R = 2 for MS and R = 3 for MB CINE) in conjunction with GRAPPA reconstruction [42] was applied to reduce the breath-hold time. For single cardiac phase acquisitions, the MS protocol was prospectively triggered to place data acquisition at end-diastole or end-systole. For prospectively triggered CINE T2* mapping, 25 cardiac phases were acquired for a heart rate of 60 bpm. Phase images of the first two echoes (TE1 = 2.04 ms, TE2 = 3.06 ms) of MS were used to determine B0 field maps offline.
Prior to T2* mapping, volume selective B0 shimming was conducted to reduce static magnetic field inhomogeneities [43], [44]. In doing so, the susceptibility weighting will be dictated by microscopic B0 susceptibility gradients, rather than by macroscopic B0 field inhomogeneities. For this purpose, a 2D multi-slice, cardiac gated, breath-hold double echo (DE) gradient echo sequence (TE1 = 3.06 ms, TE2 = 5.10 ms) was used for B0 field mapping. Cardiac gating and breath-holding were applied to reduce and possibly eliminate phase contributions induced by cardiac and respiratory motion. The shim volume was adjusted to cover the left and right ventricle in the four chamber view and in the short axis view of the heart. Data acquisition for B0 shimming was adjusted to diastole. This choice is based on previous reports which demonstrated that field maps showed a negligible temporal variation across the cardiac cycle [43]. For the shim volumes, linear and second order room temperature shims were calculated to reduce the frequency shift across the phantom or across the heart with the goal to render B0 uniform. The overall protocol time, including localizer, slice angulation, volume selective B0 mapping routine, 2D FLASH CINE imaging ( 4 chamber view and short axis view) and T2* mapping using MS and MB CINE was approximately 30 minutes. Table 1 outlines the breath hold times used for MS and MB CINE T2* mapping in healthy volunteers.
Post-Processing and Image Analysis
Image datasets were transferred to a MATLAB (The Mathworks, Natick, USA) workstation and processed offline. For all datasets, T2* was estimated based on a linear equation set obtained from the logarithm of Equation 1, where S(0) was estimated through S(TEmin). The T2* and S0 estimation was used as initialization values to fit the data points to a mono exponential decay (Equation 1) based on the MATLAB region trust algorithm.(1)
For T2* assessment of the phantom data, a ROI covering the entire central axial view of the phantom was used. Average and standard deviation of T2* were determined.
For examination of the in vivo data, an affine registration of the MB CINE datasets was incorporated into the post-processing procedure to compensate for misalignments due to the use of multiple breath-hold periods. The affine registration is landmark based. It shifts and shears the datasets derived from multiple breath-held acquisitions. Landmarks were set manually. Mid-ventricular short axis T2* maps were segmented according to the standardized myocardial segmentation and nomenclature for tomographic imaging of the heart [45]. For each segment of the mid-ventricular slice (segment 7–12 according to [45]), T2* values were calculated during end-diastole and end-systole for the single cardiac phase approach. For assessment of temporal changes in T2* throughout the cardiac cycle ROIs encompassing segment 7 to 12 were defined and analysed for all cardiac phases derived from MB CINE. For this purpose, the position and shape of the ROI was carefully adjusted throughout the cardiac cycle to account for myocardial contraction and relaxation. Also, this approach was used to include only compact myocardium into the analysis so that blood or trabecular tissue contributions can be eliminated. For careful delineation of the myocardial borders 2D CINE FLASH (flip angle = 32°, acquisition data matrix = 256×224, FOV = (288×252) mm2, in-plane resolution = (1.1×1.1) mm2, slice thickness = 4 mm, TE = 2.8 ms, TR = 4.2 ms) was used. Mean values and standard deviation of T2* were calculated for all ROIs. Statistical analysis was performed to test for data distribution and group differences using R project for statistical computing (OpenSource: www.r-project.org). A p-value below 0.05 was considered as statistically significant.
Results
Phantom Studies
T2* maps derived with all imaging strategies for the phantom experiments are surveyed in Figure 2 and Table 2 and show that all imaging strategies provide similar T2* maps. For the long T2* phantom, T2* varied from 13.6 ms to 37.4 ms across the entire central axial slice when using a slice thickness of 8 mm. The non-uniformity in T2* was reduced for 6 mm slices with T2* ranging from 19.7 ms to 33.9 ms across the entire central axial slice. T2* non-uniformity across the central axial slice was further reduced upon further reducing the slice thickness to 4 mm and 2.5 mm, which showed T2* of (29.1±1.5) ms and (29.9±1.7) ms, respectively. T2* values derived from a ROI (d = 2 cm) placed in the iso-center of the central axial slice of the phantom varied between 26.1 ms (ME) and 28.3 ms (MB CINE) for a slice thickness of 8 mm. In comparison, T2* values ranging from 28.7 ms (MS) to 30.7 (MB CINE) were observed for the same small ROI when using a slice thickness of 2.5 mm. T2* values derived from CINE imaging remained constant (std <1 ms) throughout the cycle given by the gating paradigm.
T2* maps obtained for all imaging strategies using a long T2* (A) and a medium T2*phantom (B). Slice thicknesses ranging from 8 mm to 2.5 mm (top to bottom) were applied. T2* analysis revealed similar results for all T2* mapping strategies. For a slice thickness of 8 mm T2* varied substantially across both phantoms. The uniformity in T2* was improved for a slice thickness of 6 mm and even further enhanced for a slice thickness of 4 mm or 2.5 mm.
For the medium T2* phantom, mean T2* varied between 18.3 ms and 19.3 ms for a ROI covering the entire central axial slice (slice thickness = 8 mm). For a slice thickness of 6 mm, mean T2* values were ranging from 19.0 ms to 19.9 ms. T2* mapping using a slice thickness of 4 mm yielded mean T2* values ranging from 19.3 ms to 19.8 ms. For a 2.5 mm slice thickness the range of mean T2* values encompassed 19.5 ms to 20.1 ms. For this slice thickness the standard deviation of T2* across the central slice of the phantom was approximately 2 ms for MS, MB and MS CINE due to SNR constraints. T2* values derived from CINE imaging remained constant (std <1 ms) throughout the cycle given by the cardiac gating paradigm.
In all phantom experiments, the acquisition time of the MS approach was doubled versus the ME approach, as summarized in Table 1. In MS, only 5 views per segment were recorded while ME used 10 views per segment. This approach has been deliberately chosen already at this stage to ensure that the acquisition windows do not exceed the cardiac rest period in the in vivo studies. For ME CINE, only two views per segment were used to accomplish an acquisition window of 38 ms which increased the total scan duration to 121 s in the phantom studies. This scan duration was doubled for MS CINE since only one view per segment could be used for this approach.
Volunteer Studies
For the in vivo studies, MS and MB CINE were applied using a slice thickness of 4 mm to balance the competing constraints between SNR and B0 background gradients. A reduction in slice thickness helps to reduce intra-voxel dephasing due to B0 gradients along the slice direction. This slice thickness is afforded by the SNR advantage inherent to 7.0 T and is smaller than that commonly used for T2* mapping at 1.5 T and 3.0 T. Localized shimming was performed to reduce static magnetic field inhomogeneities to make sure that the susceptibility weighting is not dominated by macroscopic B0 field inhomogeneities but rather by microscopic B0 susceptibility gradients. Figure 3 and Figure 4 depict B0 maps together with B0 profiles across the heart and frequency histograms of the heart obtained prior to and after volume selective shimming for a four chamber (Figure 3) and short axis view (Figure 4). The B0 field maps following global shimming showed a mean peak-to-peak field difference of approximately 400 Hz across the heart for a four chamber view (Figure 3B) and approximately 300 Hz for a mid-ventricular short axis view (Figure 4B). After volume selective shimming, a mean peak-to-peak B0 difference of approximately 80 Hz was found across the entire heart for a four chamber view (Figure 3B) and a mid-ventricular short axis view (Figure 4B) of the heart. For the left ventricle a B0 peak-to-peak difference of approximately 65 Hz was observed after volume selective shimming. For both short axis view and 4 chamber view, a maximum in-plane field gradient of approximately 20 Hz/mm (through-plane approximately 80 Hz/voxel for a 4 mm slice thickness) was observed at the epicardial fat/lung interface of the inferior and inferolateral segment as indicated by the B0 maps and frequency profiles shown in Figure 3B,C and Figure 4B,C. This local B0 gradient translates into a phase loss of approximately 80% at the maximum echo time of TE = 10 ms. However, the through-plane field gradient at the epicardium/lung interface is much more pronounced versus the through-plane field gradient obtained for the left and right ventricle, as demonstrated in Figure 3B/C. For myocardial anterior, anterolateral and inferoseptal segments a mean in-plane B0 gradient of approximately 3 Hz/mm was obtained which translates into an through-plane B0 dispersion of approximately 12 Hz/voxel for a 4 mm slice thickness. This B0 gradient implies that macroscopic intravoxel dephasing effects are of minor effect for the TE range used. Inspite of B1+ shaping using EMF simulations the signal intensity change across the myocardium of a mid-ventricular short axis slice was found to be approximately 40%.
A) Four chamber view of the heart illustrating the positioning of the volume (marked in red) used for global (left) and volume selective (right) shimming. B) B0 field variation derived from global and volume selective shimming. For this subject the global shim provided a peak-to-peak field variation of about 400 Hz across the entire heart. After volume selective shimming peak-to-peak B0 variation across the heart was reduced to approximately 80 Hz. The direction of the maximal B0 gradient is illustrated by the dashed black line in B) and the corresponding profile of B0 field distribution is plotted in C). To guide the eye the epicardial borders are marked in B) and C) by two triangles. The histogram of the field distribution over the left ventricle is shown in D). The full width at half maximum is approximately 200 Hz for the globally shimmed B0 field map and was reduced to about 80 Hz after volume selective shimming.
A) Mid-ventricular short axis view of the heart illustrating the positioning of the volume (marked in red) used for volume selective shimming. B) B0 field maps. C) B0 profile along the direction of the strongest B0 gradient which is highlighted by the dashed black line in B). To guide the eye the epicardial borders are marked in B) and C) by two triangles. D) Frequency histogram across the left ventricle. After volume selective shimming a strong susceptibility gradient at the inferior region of the heart could be reduced. The full width at half maximum is approximately 300 Hz for the globally shimmed field map and was reduced to about 80 Hz after volume selective shimming.
We next employed MS and MB CINE to determine the limits of susceptibility artifacts. Figure 5 shows end-diastolic short axis views derived from MS and MB CINE acquisitions using echo times ranging from 2.04 ms to 10.20 ms. With the flip angle of 20°, RF power deposition was well in the SAR limits given by the IEC guidelines [46], which were confirmed by rigorous electromagnetic field simulations [36], [37]. The low flip angle was deliberately chosen to preserve myocardial signal by reducing T1-saturation effects. Although this approach results in a low contrast between the blood pool and the surrounding myocardium (Figure 5), it is beneficial for myocardial T2* quantification. Unlike previous reports at 1.5 T and 3.0 T double inversion recovery prepared blood suppression was not used since this approach does not meet the requirements of CINE T2*mapping. It should be noted that no differences in mean T2* values were found for black blood versus white blood acquisitions at 1.5 T [47]. No severe susceptibility artifacts were detected in the septum (segment 8–9) and in the lateral wall (segment 11–12) for TEs ranging between 2.04 ms to 10.20 ms. For anterior (segment 7) and inferior (segment 10) myocardial areas, which encompass major cardiac veins, a signal void related to the susceptibility weighting was observed for TE >7 ms (Figure 5).
Echo times ranging from 2.04 ms to 10.20 ms were used for MS and MB CINE acquisitions. A low nominal flip angle of 20° was used to preserve myocardial signal. Image quality observed for MS and MB CINE acquisitions is comparable. No severe susceptibility artifacts were detected in the septum and in the lateral wall for TEs ranging between 2.04 ms to 10.20 ms. For anterior and inferior myocardial areas encompassing major cardiac veins susceptibility weighting related signal void was observed for TE >7 ms as highlighted by white arrows.
We next explored myocardial T2* using single cardiac phase (MS) and dynamic CINE (MB CINE) acquisitions. Figure 6 shows end-systolic and end-diastolic short axis and four chamber view T2* maps. Regions of T2* reduction were observed in areas adjacent to the large cardiac vein and the posterior vein. A paired t-test showed no significant differences in T2* values observed for MS and MB CINE strategies. (p = 0.8 for diastole, p = 0.7 for systole). The differences between the MS the MB CINE imaging strategy including all mid-ventricular segments were found to be T2*diff = (0.2±3.7) ms at end-diastole and T2*diff = (−0.2±3.4) ms at end-systole. MS and MB CINE showed a mean T2* of approximately 14 ms for mid-ventricular myocardium. Regional T2* variation was observed for mid-ventricular myocardium. The longest T2* values were found for segment 8 (MS: T2* = 17.2 ms, MB CINE: T2* = 17.3 ms), segment 9 (MS: T2* = 16.5, MB CINE: T2* = 16.3) and segment 7 (MS: T2* = 14.0 ms, MB CINE: T2* = 16.8 ms). For segment 10 and segment 11 lower T2* values were observed (segment 10: MS: T2* = 10.6 ms, MB CINE: T2* = 12.0 ms, segment 11: MS: T2* = 11.4 ms, MB CINE: T2* = 11.4 ms). A synopsis of T2* values averaged over all subjects for MS and MB CINE acquisitions at end-diastole is given in Table 3.
Four chamber (top) and short axis view T2* colour maps obtained from MS and MB CINE superimposed to anatomical 2D CINE FLASH gray scale images. For MB CINE a systolic and diastolic phase was chosen to match the cardiac phase with the end-systolic and end-diastolic phase derived from MS. T2* maps deduced from MS and MB CINE showed no significant differences between both methods in the segmental analysis of T2* values. When comparing systolic and diastolic T2* maps significant differences were found with p = 0.002 for MS and p = 0.01 for MB CINE.
A closer examination revealed a significant difference between myocardial T2* obtained at end-diastole and those derived from end-systolic acquisitions. Indeed, an increase in T2* can be clearly observed, particularly at the septum, during systolic phases (Figure 6, Figure 7). Figure 7 shows exemplary CINE T2* maps for all phases of the cardiac cycle and illustrates the changes of T2* over the cardiac cycle. A paired t-test comparing end-diastolic and end-systolic phase T2* values for all segments revealed p = 0.002 for single cardiac phase MS acquisitions and p = 0.01 for MB CINE acquisitions. When averaging the T2* time course from all subjects an increase in T2* of 27%±6% in all cardiac segments could be observed over the cardiac cycle (Figure 8). The shortest T2* values were noted for a cardiac phase placed in systole. For all cardiac segments, the longest T2* values were found after the onset of diastole. The largest T2* increase over the cardiac cycle was found for segment 7 (ΔT2* = +3.7 ms) and segment 10 (ΔT2* = +3.4 ms).
Short axis view T2* colour maps derived from MB CINE acquisitions across the cardiac cycle overlaid to conventional 2D CINE FLASH images. T2* values are increasing from diastole to systole, especially for endocardial layers. Macroscopic susceptibility induced T2* reduction effects were present at the epicardium at inferior regions.
Synopsis of the evolution of mean T2* averaged over all subjects for standard mid-ventricular segments of the heart. T2* derived from each cardiac segment are plotted versus the cardiac cycle. T2* changes over the cardiac cycle. Averaging T2* over all mid-ventricular myocardial segments revealed that T2* increases approximately 27% between systole and diastole. Myocardial T2* was derived from MB CINE acquisitions. Prospective triggering was used which resulted in a gap at end-diastole of approximately 100 ms depending on the heart rate. For this reason the cardiac cycle is normalized for all subjects without including this gap.
To demonstrate the baseline SNR advantage of 7.0 T and to further reduce intravoxel dephasing along the slice direction, the slice thickness was reduced to 2.5 mm while maintaining the in-plane spatial resolution of (1.1×1.1) mm2 in the MB CINE protocol. Figure 9 shows short axis and four chamber view T2* maps, which demonstrate the high spatial resolution. This approach further manifests the observed changes in T2* during systole and diastole and also indicates the high sensitivity of susceptibility mapping by visualizing differences in T2* between endocardial and epicardial layers of the myocardium.
For MB CINE slice thickness was reduced to 2.5 mm while maintaining the in-plane spatial resolution of (1.1×1.1) mm2. Compared to the results obtained with MB CINE using a slice thickness of 4 mm, changes in T2* from epicardial to endocardial septal myocardial layers are more pronounced, in particular during systole.
Discussion
This work shows the feasibility of high spatially and temporally resolved myocardial T2* mapping at 7.0 T. For this purpose T2* weighted, gradient echo based imaging techniques using single cardiac phase (MS) and CINE (MB CINE) acquisition regimes were benchmarked against T2* mapping techniques commonly used in current clinical practice at 1.5 T and 3.0 T. These two imaging techniques were first examined in detail in phantom experiments.
It might be considered as a remaining limitation that our results might be affected by residual macroscopic B0 gradients. However, our B0 mapping results suggest that a reasonable B0 uniformity across the heart and the left ventricle can be achieved at 7.0 T which is embodied by a mean through-plane gradient of 3 Hz/mm across the left ventricle. In this regard it should be also noted that our measurements of the B0 field distribution after volume selective shimming of our uniform phantom provided a through slice peak-to-peak B0 variation of 80 Hz along a distance of 4 cm which translates into 2 Hz/mm. This B0 field gradient is similar to what has been observed for the left and right ventricle which showed a mean of 3 Hz/mm.
The frequency shift across the heart reported here compares well with previous 3.0 T studies which reported a peak-to-peak off-resonance variation of (262±58) Hz over the left ventricle and the right ventricle (basal short axis view) for a global shim [48]. This B0 inhomogeneity was improved to (176±30) Hz and (121±31) Hz with the use of localized linear and second-order shimming [48]. The use of an enhanced locally optimized shim algorithm, which is tailored to the geometry of the heart, afforded a reduction of the peak-to-peak frequency variation over the heart from 235 Hz to 86 Hz at 3.0 T [49]. Another pioneering study showed a peak-to-peak off-resonance of (71±14) Hz for short axis views acquired at 1.5 T [40] using global shimming.
Previous reports on brain imaging/spectroscopy suggest that third-order and even higher order shims help to further enhance B0 uniformity across the target anatomy [50], [51]. For this purpose extra shim drum inserts are retrofitted to the scanner. Notwithstanding its utility the current implementation limits the available space inside of the MR scanner bore which would be prohibitive for cardiac or body MRI at 7.0 T. Other attempts to integrate third order shim coils into high performance 7.0 T whole body gradient coil designs were found to show pronounced gradient non-linearity for spherical volumes with a diameter larger than (20–25) cm; a behavior which does not meet the requirements of cardiac or body MR. Obviously, another approach to further reduce the residual impact of through-plane gradients and intra-voxel dephasing B0 gradients is the use of even thinner slices and the reduction in voxel size. To meet this goal we pushed the envelope by using a slice thickness as thin as 2.5 mm together with an in-plane resolution of (1.1×1.1) mm2. This slice thickness and in-plane resolution is afforded by the SNR advantage inherent to 7.0 T. The corresponding voxel size is by a factor of five smaller than commonly used for T2* mapping at 1.5 T and 3.0 T. However, it should be noted that the move to even thinner slices and smaller voxel sizes – ideally one might opt to use an infinitesimal small voxel – would disturb the balance dictated by the competing constraints of SNR and background gradients effects.
This study sheds further light to the current literature since it demonstrates the applicability of MS and MB CINE for T2* mapping of normal myocardium at 7.0 T. While we recognize a limitation due to the limited number of healthy subjects studied, we believe this feasibility study to be an essential precursor to a larger 7.0 T study involving healthy and patient cohorts. Such a study would aid to establish the lower limits for normal myocardial T2* values versus the clinically established normal values for T2* of healthy myocardium at 1.5 T and 3.0 T. To this end, T2* mapping at 7.0 T may be useful to extend the capabilities and the dynamic range of the sensitivity of the established approach used for quantification of myocardial iron content. With this in mind, we anticipate to extend our efforts towards clinical studies at 7.0 T including thalassemia major patients, whose T2* relaxation times will be benchmarked against the normal values of healthy subjects.
Our results show that T2* obtained for human myocardial muscle tissue at 7.0 T ranges from 9 ms to 18 ms. This is in line with T2* = (15.8±0.2) ms recently observed for hind limb skeletal muscle in rats at 7 Tesla [52] Admittedly, the absolute spatial resolution demonstrated for T2* mapping of the human heart at 7.0 T is still by an order of magnitude below that previously reported for ex vivo MR microscopy based T2*mapping of the isolated rat heart [7], which demonstrated that T2*mapping provides an insight into the complex architecture of the heart musculature. However, the effective anatomical spatial resolution – voxel size per anatomy – is getting close to what has been demonstrated for animal models. This improvement might be beneficial to gain a better insight into the myocardial microstructure in vivo with the ultimate goal to visualize myocardial fibers or to examine helical angulation of myocardial fibers using T2* mapping, since the susceptibility effects depend on the tilt angle between blood filled capillaries and the external magnetic field [53]. Myocardial fibre tracking using T2* mapping holds the promise to be less sensitive to bulk motion than diffusion-weighted MR of the myocardium [54], [55]. Our results also suggest that the increased susceptibility contrast available at 7.0 T could be exploited to quantitatively study iron accumulations in organs other than the heart with high sensitivity and temporal and spatial resolution superior to what can be achieved at 1.5 T and 3.0 T.
For normal myocardium a T2* value of approximately 37 ms was found at 1.5 T [23]. At 3.0 T a T2* of approximately 27 ms was observed for normal myocardium [56]. These measurements are usually limited to the septum, which shows the lowest spatial variation in T2* [57]. It is elusive to study temporal changes in T2* at 1.5 and 3.0 T due to scan time constraints which are prohibitive for CINE T2* mapping. Of course, single cardiac phase T2* mapping can be applied to diastole and systole as reported previously [58]. This 1.5 T study with thalassemia patients demonstrated mean T2* values of (26.4±14.2) ms for early systole and (25.2±13.1) ms for late diastole, which were found to be not significantly different (P = 0.27). However, the limited T2* sensitivity together with the temporal resolution used in the study presents a challenge for tracking temporal changes in T2*. Please also note, that these data exhibit a rather large standard deviation of approximately ±13.0 ms, which presents another challenge for assessment of temporal T2* changes.
A careful literature research revealed that no 1.5 T and 3.0 T T2* mapping study has been reported yet which uses a high spatial resolution accomplished here. We would also like to point out that our study is the first study which affords CINE T2* mapping due to inter echo time shortening at 7.0 T. To this end it is interesting to note that the myocardial BOLD effect has been investigated using SSFP imaging, which is sensitive to changes in the relaxation times T2 and T1. In this regard it has been shown recently that the signal intensity derived from SSFP imaging of the myocardium varies across the cardiac cycle [59]. This study showed a systole-to-diastole T2 ratio of approximately 1.1 for normal myocardium.
The ability to probe for changes in tissue oxygenation using T2* sensitized imaging/mapping offers the potential to address some of the spatial and temporal resolution constraints of conventional first pass perfusion imaging and holds the promise to obviate the need for exogenous contrast agents. Since microscopic susceptibility increases with field strength, thus making the BOLD effect due to (patho)physiology of interest more pronounced, T2* mapping at 7.0 T might be beneficial to address some of the BOLD sensitivity constraints reported for the assessment of regional myocardial oxygenation changes in the presence of coronary artery stenosis [60] or for the characterization of vasodilator-induced changes of myocardial oxygenation at 1.5 T and at 3.0 T [17].
Conclusion
Our results underscore the challenges of myocardial T2* mapping at 7.0 T due to the propensity to macroscopic susceptibility artefacts and T2* shortening, but demonstrate that these issues can be offset by using tailored shimming techniques together with dedicated acquisition schemes.
Acknowledgments
We wish to acknowledge Peter Kellman (Laboratory of Cardiac Energetics, NHLBI, Bethesda, MD, USA) for his valuable support and Antje Els and Julia Skodowski for technical assistance.
Author Contributions
Conceived and designed the experiments: FH CT JSM TN. Performed the experiments: FH CT TN. Analyzed the data: FH JSM. Contributed reagents/materials/analysis tools: FH SW. Wrote the paper: FH CT SW JSM TN.
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