Rapid T1 quantification based on 3D phase sensitive inversion recovery
© Warntjes et al; licensee BioMed Central Ltd. 2010
Received: 26 October 2009
Accepted: 17 August 2010
Published: 17 August 2010
In Contrast Enhanced Magnetic Resonance Imaging fibrotic myocardium can be distinguished from healthy tissue using the difference in the longitudinal T 1 relaxation after administration of Gadolinium, the so-called Late Gd Enhancement. The purpose of this work was to measure the myocardial absolute T 1 post-Gd from a single breath-hold 3D Phase Sensitivity Inversion Recovery sequence (PSIR). Equations were derived to take the acquisition and saturation effects on the magnetization into account.
The accuracy of the method was investigated on phantoms and using simulations. The method was applied to a group of patients with suspected myocardial infarction where the absolute difference in relaxation of healthy and fibrotic myocardium was measured at about 15 minutes post-contrast. The evolution of the absolute R 1 relaxation rate (1/T 1) over time after contrast injection was followed for one patient and compared to T 1 mapping using Look-Locker. Based on the T 1 maps synthetic LGE images were reconstructed and compared to the conventional LGE images.
The fitting algorithm is robust against variation in acquisition flip angle, the inversion delay time and cardiac arrhythmia. The observed relaxation rate of the myocardium is 1.2 s-1, increasing to 6 - 7 s-1 after contrast injection and decreasing to 2 - 2.5 s-1 for healthy myocardium and to 3.5 - 4 s-1 for fibrotic myocardium. Synthesized images based on the T 1 maps correspond very well to actual LGE images.
The method provides a robust quantification of post-Gd T 1 relaxation for a complete cardiac volume within a single breath-hold.
Contrast Enhanced Magnetic Resonance Imaging (CEMRI) is the preferred modality for the detection and characterization of myocardial viability [1–6]. At 10-30 minutes after the administration of a T 1 contrast medium fibrotic or otherwise damaged myocardium exhibits hyper-enhancement in comparison with healthy tissue, owing to differences in wash-out kinetics of the contrast agent. Typically, a Phase Sensitive Inversion Recovery (PSIR) sequence is applied for high image contrast between healthy and fibrotic myocardium. In such an acquisition an inversion pulse is applied followed by two acquisitions, one at a short inversion delay time T inv and a second during the same heart phase at the subsequent heart beat. The total kernel time of the acquisition spans two cardiac RR intervals. The latter acquisition is used to correct the background phase of the former such that a real image is reconstructed instead of a modulus. The advantage of this procedure is that there is no contrast degradation due to signal rectification of the original modulus image .
The contrast in the PSIR image is governed by the longitudinal T 1 relaxation of the various tissues. Signal intensity differences in the images indicate differences in T 1 but, since the image is arbitrarily scaled, no absolute numbers can be retrieved. Changes in the absolute relaxation rate R 1 (= 1/T 1) provides a measure for absolute local contrast media concentration [8, 9]. Monitoring R 1 over time pictures the actual contrast medium dynamics without the potential offset intensity bias or changes caused by heart rate variability, as seen with conventional dynamic T 1-weigted imaging. Quantification may even improve the stability of segmentation of healthy myocardium and scar tissue. Especially in follow-up studies on the volume of the myocardium and scar, a reliable segmentation is required, independent of scanner settings [10–12]. Finally, T 1 maps are independent of RF coil sensitivity. This may be important for imaging using (phased array) coils with a strong spatial sensitivity gradient and without proper intensity scaling (such as CLEAR, Constant Level Appearing). It may also reduce the need for a fat suppression technique to remove the high-intensity fat signal that may disturb the image reading. A variety of T 1 mapping methods exists (see e.g. Refs. [13–17]). A number of these methods rely on strategies with continuous acquisition, which leads to movement artifacts for the heart. Others require several breath-holds to cover the complete cardiac volume. In this work a method is described to retrieve the absolute T 1 relaxation based on a 3D PSIR acquisition, which can cover the complete cardiac volume within a single breath hold.
A perfect inversion pulse is assumed, a small deviation of the complete inversion has only a negligible effect on the calculations. Typically, a 10% reduction of the inversion angle (162 degrees rather than 180) results in an overestimation of 2-3% for a T 1 in the range 200-800 ms. For a single shot PSIR acquisition, with sufficient time between subsequent measurements for the magnetization to fully recover, M A equals -M 0. For a multi-shot acquisition, either a multi-slice or a 3D sequence, where an inversion pulse is applied every 2RR, in clinical practice a steady-state is reached where M A = -M F and (-M A) < M 0.
where exp(TC) equals exp(-(T RR - T acq)/T 1). Finally M F is calculated using the new M 0 in Eqs. 5 and 6. The second iteration starts with the estimated M 0 from Eq. 8 and a new M A = -M F.
Using Eq. 9 the T 1* relaxation can be retrieved from a LL sequence. Subsequently the actual T 1 can be calculated using Eq. 1. From these equations it can directly be seen that the intensity zero-crossing of a LL sequence, in general, does not coincide with the intensity zero-crossing of a PSIR sequence. The continuous acquisition of a LL leads to a T 1* relaxation which is shorter than T 1 and the saturation and acquisition effects lead to a different steady-state magnetization. Therefore care should be taken in using the LL to find the intensity zero-crossing for another sequence, although this is commonly done.
As an additional confirmation of the accuracy of the T 1 quantification the approach of Synthetic MRI [21–24] is applied: The T 1 maps are used as input to simulate a 3D spoiled gradient echo sequence (Inversion Recovery Turbo Field Echo or IR-TFE). Based on the T 1 values of the quantification scan the expected image intensity of an IR-TFE can be calculated for any chosen T inv using Eqs. 1-6 (using M D = M E and a kernel time of a single RR). The synthesized images are compared with the actual ones with the same T inv. The IR-TFE is interesting since this sequence does not have the second acquisition, like the PSIR method, to restore phase and hence potentially has up to 41% more SNR within the same scan time compared to a PSIR sequence of equal geometry. A prerequisite, however, is that the optimal inversion delay is applied since the IR-TFE scans do suffer from signal rectification in case T inv is chosen too short . To ensure an optimal T inv the synthetic images can be set first such that the healthy myocardium appears black. Subsequently this value for T inv can be used as input for the actual IR-TFE scan.
All experiments were performed on a 1.5T Achieva scanner (Philips Healthcare, Best, The Netherlands). Phantoms were made that matched the relevant cardiac relaxation rates as good as possible. Water was used with a 2.5% Agerose solution (Sigma-Aldrich, St. Lious, USA) and different concentrations of the Gadolinium contrast agent (Magnevist, 0.5 mmol/mL, Bayer Healthcare, Germany, diluted to 0.06 - 0.3 mmol/L) resulting in T 1 = 228, 298, 411, 539, 638 and 754 ms. The T 2 relaxation times of all phantoms was in the range 42-59 ms. The 3D PSIR protocol for the T 1 quantification was a segmented 3D Turbo Field Echo Planar Imaging (TFEPI) sequence with an EPI factor 3 and a TFE factor 23. The echo time (TE) was 4.2 ms and the repetition time (TR) 9.4 ms leading to an acquisition phase of 215 ms per heart beat. The matrix size was 228×138 (reconstructed 320×320) over a Field of View (FOV) of 350 × 320 mm. The slices had a thickness of 5 mm (overcontiguous, i.e. the slices overlap). Using a Sense factor 2 a volume of 12-18 slices can be acquired within 24 seconds, depending on the heart rate (2 heart beats per slice). For the phantoms physiology simulation was used for artificial heart triggering. A heart rate of 60 beats per minutes was set, resulting in 12 slices in 24 seconds.
The absolute T 1 relaxation time of the phantoms was validated using a standard inversion recovery sequence with 9 separate measurements at an inversion delay time of 50, 100, 150, 200, 250, 300, 500, 1000 and 2900 ms. The TE was 29 ms (EPI factor 13), TR = 3000 ms and the flip angle 90°.
For the in-vivo measurements the T inv was by default set to 300 ms and the flip angle to 18 degrees. The acquisition was performed during diastole. The number of slices was adjusted in the range 12-18 slices to fit within a 24 seconds breath-hold. The quantification method was added to routine clinical examinations of patients that were followed up after primary PCI for ST-elevation myocardial infarction. The study was approved by the regional ethics committee and complied with the declaration of Helsinki. All patients gave written informed consent. They were given 0.2 mmol/kg (max. 15 mmol) Gadolinium contrast agent (Magnevist 0.5 mmol/ml, Bayer Healthcare, Germany). On one patient, the T 1 quantification protocol was performed every 2 minutes. The quantification scan was interleaved with a LL sequence, a single slice inversion recovery scan with a flip angle of 15 degrees and a TR of 25 ms. The LL resulted in 29 heart phases within a breath-hold of 17 seconds. The matrix size was 228 × 201, reconstructed to 320×320.
On all other patients the quantification method was applied once, about 15 minutes after contrast injection. The Synthetic MRI approach was used to establish the optimal inversion delay. Directly after the 3D PSIR acquisition the images were sent to the PACS system (IDS5, Sectra Imtec, Sweden) where the optimal inversion delay time was retrieved from the synthetic images using a dedicated cardiac package (SyMRI Cardiac Studio, SyntheticMR AB, Sweden). The value for the optimal T inv was used as input for the IR-TFE sequence to ensure black myocardium for this protocol. To compensate for the time delay between the 3D PSIR and the IR-TFE, in general about 1 minute, 20 ms was added to the suggested T inv.
The IR-TFE sequence was a segmented 3D spoiled gradient echo sequence with TE = 1.3 ms, TR = 4.4 ms and TFE factor 43, leading to an acquisition phase time of 188 ms, also acquired during diastole. In total 17 slices were acquired with a thickness of 5 mm (overcontiguous) and Sense factor 2. The matrix size was 256×172 (reconstructed to 320×320) over a FOV of 350 mm resulting in a scan time of 17 heart beats.
In order to compare the imaging methods two regions of interest were placed in all images, one in healthy and one in fibrotic myocardium. The Contrast to Noise Ratio (CNR) was defined as the difference in signal intensity of the two ROI's divided by the standard deviation in the ROI over the healthy (black) myocardium. Both methods turn out to have equal CNR (R2 = 0.90). A significant difference is, however, that the IR-TFE acquires more slices (22 - 17 compared to 18 - 10 of the PSIR) and a higher resolution (acquisition voxel size 1.5 × 1.6 mm compared to 1.5 × 2.3 mm of the PSIR) in the same scan time.
The validation of the absolute T 1 relaxation time in phantoms, as depicted in Figs.2 and 3, shows that our method is consistent over a large range of T 1 values, flip angles and inversion delay times. As can be seen in Fig. 2 the optimal flip angle of the method is around 15-20 degrees where the signal to noise ratio is highest and the deviation from the expected value is low. The range of inversion delays that can be chosen is large. The fitting algorithm consistently removes the saturation and acquisition effects. The observation that short T 1 values in the order of 200 ms are consistently measured even with an inversion delay of up to 600 ms implies that the assumption of a perfect inversion pulse is valid. An imperfect inversion would lead to an overestimation of T 1. Possibly this issue has a larger influence at higher field strengths.
Typical T 1 values for an LGE measurement are in the range 200 - 500 ms. Based on Fig 3 the inversion delay of the method can therefore be chosen anywhere in the range 300 - 600 ms. An inversion delay of 300 ms was selected for the in-vivo experiments mainly because it is close to the commonly used inversion delay. With these settings the proposed method correctly estimates T 1 values even up to 800 ms although the kernel time is only 2 s. Although it was not the focus of this study it is likely that the method would work for pre-GD myocardial T 1 values as well. In that case a more natural choice for the inversion delay would be higher, e.g. 600 ms.
In clinical practice other parameters are important, such as the variability of the heart-rate during breath-hold, which might decrease the accuracy of the T 1 map. The Monte-Carlo simulation displayed in Fig. 4 shows that small changes in heart rate have less influence on the typical T 1 values than the noise level of the measurement. Furthermore, the resulting error shifts all tissue of interest in a similar fashion and leaves the differences in T 1 between the healthy and fibrotic myocardium virtually unaffected.
Care has to be taken in the interpretation of the absolute T 1 relaxation time. A tissue voxel comprises of many T 1 components rather than the mono-exponential decay that is assumed in the method. Furthermore, at the selected echo time the short-lived T 1 components might be underestimated and the frequency difference between water and fat might lead to a spatial shift of the intensity in the images. Our 2-point method is designed to rapidly estimate the predominant component of T 1 relaxation and for the given scanner parameters this is achieved.
A potential disadvantage of the method is the long breath-hold time (24 s). This results in 12-18 slices of 5 mm depending on the heart rate. Should this be too long the acquisition time may be decreased by reducing the number of slices. The reduced heart volume coverage may be compensated by increasing the slice thickness.
An example of the application of the method is given in Figs. 5, 6, 7 where a patient was monitored every 2 minutes post-Gd. A clear evolution of R 1 is observed over time in the heart and the liver. The change of the relaxation rate ΔR 1 is taken here, rather than T 1, since ΔR 1 is proportional to the absolute amount of contrast medium present in the tissue and should therefore also represent the severity of fibrosis on a microscopic level. As known from practice the ΔR 1 rapidly decreases in the first 5-10 minutes to remain relatively flat in the following 10-30 minutes. The late enhancement contrast already appears after a few minutes. For our patient group the amount of contrast media was about 2.5 times higher in the fibrotic areas compared to the healthy areas (Fig. 8) at about 15 minutes post-Gd. Note that fat is unaffected by the contrast medium and hence can serve as a reference signal intensity for relative measurements of signal intensity during a contrast bolus for perfusion.
The approach of synthetic MRI is used for a direct comparison of the T 1 quantification maps and conventional imaging resulting in very similar images as shown in Fig. 9. Pathology shows up similarly and the method has a good CNR. Interestingly this approach also means that the quantification method may provide both the T 1 maps and the relevant clinical images in one single scan. The ability to synthetically vary T inv after the actual acquisition may optimize the image quality separately for both ventricles . Moreover the method may serve as a test scan to optimize the T inv for subsequent IR-TFE sequences.
We present a method to quantify cardiac T 1 relaxation in a large volume within a single breath-hold, based on a 3D Phase Sensitive Inversion Recovery sequence. The fitting algorithm takes acquisition and saturation effects into account and is robust against variation of scan parameters and heart rate. The method is independent of RF coil sensitivity issues such that high-SNR phased array coils can be employed. The absolute relaxation rate R 1 is monitored over time and over a group of patients. The T 1 maps can be used for 3D segmentation and synthesis of conventional LGE images with a free choice of inversion delay.
Parts of this work were funded by the University Hospital Research Funds, the Medical Research Council of Southeast Sweden and the Swedish Heart-Lung foundation. The synthetic MRI software has been provided by SyntheticMR AB, Sweden, http://www.syntheticmr.se
- Kim RJ, Fieno DS, Parrish TB, Parrish TB, Harris K, Chen EL, Simonetti O, Bundy J, Finn JP, Klocke FJ, Judd RM: Relationship of MRI delayed contrast enhancement to irreversible injury, infarct age and contractile function. Circulation. 1999, 100: 1992-2002.View ArticlePubMedGoogle Scholar
- Wendland MF, Saeed M, Lund G, Higgins CB: Contrast-enhanced MRI for quantification of myocardial viability. J Magn Reson Imaging. 1999, 10: 694-702. 10.1002/(SICI)1522-2586(199911)10:5<694::AID-JMRI12>3.0.CO;2-J.View ArticlePubMedGoogle Scholar
- Wagner A, Mahrholdt H, Sechtem U, Kim RJ, Judd RM: MR imaging of myocardial perfusion and viability. Magn Reson Imaging Clin N Am. 2003, 11: 49-66. 10.1016/S1064-9689(02)00048-X.View ArticlePubMedGoogle Scholar
- Hunold P, Schlosser T, Vogt FM, Eggebrecht H, Schmermund A, Bruder O, Schüler WO, Barkhausen J: Myocardial late enhancement in contrast-enhanced cardiac MRI: distinction between infarction scar and non-infarction-related disease. Am J Roentgenol. 2005, 184: 1420-1426.View ArticleGoogle Scholar
- Isbell DC, Kramer CM: Magnetic resonance for the assessment of myocardial viability. Curr Opin Cardiol. 2006, 21: 469-72. 10.1097/01.hco.0000240584.68720.ae.View ArticlePubMedGoogle Scholar
- Tatli S, Zou KH, Fruitman M, Reynolds HG, Foo T, Kwong R, Yucel EK: Three-dimensional magnetic resonance imaging technique for myocardial delayed hyperenhancement: A comparison with the two-dimensional technique. J. Magn Reson Imaging. 2004, 20: 378-382. 10.1002/jmri.20124.View ArticlePubMedGoogle Scholar
- Kellman P, Arai AE, McVeigh ER, Aletras AH: Phase-sensitive inversion recovery for detecting myocardial infarction using gadolinium-delayed hyperenhancement. Magn Reson Med. 2002, 47: 372-282. 10.1002/mrm.10051.View ArticlePubMedPubMed CentralGoogle Scholar
- Kiss P, Suranyi P, Simor T, Saab-Ismail NH, Elgavish A, Hejjel L, Elgavish GA: In Vivo R1-Enhancement Mapping of Canine Myocardium Using CeMRI With Gd(ABE-DTTA) in an Acute Ischemia-Reperfusion Model. J. Magn. Reson. Imag. 2006, 24: 571-579. 10.1002/jmri.20661.View ArticleGoogle Scholar
- Judd RM, Reeder SB, May-Newman K: Effects of water exchange on the measurement of myocardial perfusion using paramagnetic contrast agents. Magn Reson Med. 1999, 41: 334-342. 10.1002/(SICI)1522-2594(199902)41:2<334::AID-MRM18>3.0.CO;2-Y.View ArticlePubMedGoogle Scholar
- Choi CJ, Haji-Momenian S, Dimaria JM, Epstein FH, Bove CM, Rogers WJ, Kramer CM: Infarct involution and improved function during healing of acute myocardial infarction: the role of microvascular obstruction. J Cardiovasc Magn Reson. 2004, 6: 917-925. 10.1081/JCMR-200036206.View ArticlePubMedGoogle Scholar
- Petersen SE, Voigtlander T, Kreitner KF, Horstick G, Ziegler S, Wittlinger T, Abegunewardene N, Schmitt M, Schreiber WG, Kalden P, Mohrs OK, Lippold R, Thelen M, Meyer J: Late improvement of regional wall motion after the subacute phase of myocardial infarction treated by acute PTCA in a 6-month follow-up. J Cardiovasc Magn Reson. 2003, 5: 487-495. 10.1081/JCMR-120022264.View ArticlePubMedGoogle Scholar
- Petersen SE, Mohrs OK, Horstick G, Oberholzer K, Abegunewardene N, Ruetzel K, Selvanayagam JB, Robson MD, Neubauer S, Thelen M, Meyer J, Kreitner KF: Influence of contrast agent dose and image acquisition timing on the quantitative determination of nonviable myocardial tissue using delayed contrast-enhanced magnetic resonance imaging. J Cardiovasc Magn Reson. 2004, 6: 541-548. 10.1081/JCMR-120030581.View ArticlePubMedGoogle Scholar
- Wansapura J, Gottliebson W, Crotty E, Fleck R: Cyclic variation of T1 in the myocardium at 3 T. Magn Reson Imaging. 2006, 24: 889-893. 10.1016/j.mri.2006.04.016.View ArticlePubMedGoogle Scholar
- Freeman AJ, Gowland PA, Mansfield P: Optimization of the ultrafast Look-Locker echo-planar imaging T1 mapping sequence. Magn Reson Imaging. 1998, 16: 765-772. 10.1016/S0730-725X(98)00011-3.View ArticlePubMedGoogle Scholar
- Bokacheva L, Huang AJ, Chen Q, Oesingmann N, Storey P, Rusinek H, Lee VS: Single breath-hold T1 measurement using low flip angle TrueFISP. Magn Reson Med. 2006, 55: 1186-1190. 10.1002/mrm.20845.View ArticlePubMedGoogle Scholar
- Cernicanu A, Axel L: Theory-based signal calibration with single-point T1 measurements for first-pass quantitative perfusion MRI studies. Acad Radiol. 2006, 13: 686-693. 10.1016/j.acra.2006.02.040.View ArticlePubMedGoogle Scholar
- Messroghli DR, Greiser A, Fröhlich M, Dietz R, Schulz-Menger J: Optimization and validation of a fully-integrated pulse sequence for modified look-locker inversion-recovery (MOLLI) T1 mapping of the heart. J Magn Reson Imaging. 2007, 26: 1081-1086. 10.1002/jmri.21119.View ArticlePubMedGoogle Scholar
- Deichmann R: Fast high-resolution T1 mapping of the human brain. Magn Reson Med. 2005, 54: 20-27. 10.1002/mrm.20552.View ArticlePubMedGoogle Scholar
- Warntjes JBM, Dahlqvist O, Lundberg P: A Novel Method for Rapid, Simultaneous T1, T2* and Proton Density Quantification. Magn Reson Med. 2007, 57: 528-37. 10.1002/mrm.21165.View ArticlePubMedGoogle Scholar
- Look DC, Locker DR: Time saving in measurement of NMR and EPR relaxation times. Rev Sci Instrum. 1970, 41: 250-251. 10.1063/1.1684482.View ArticleGoogle Scholar
- Riederer SJ, Suddarth SA, Bobman SA, Lee JN, Wang HZ, MacFall JR: Automated MR image synthesis: feasibility studies. Radiology. 1984, 153: 203-206.View ArticlePubMedGoogle Scholar
- Bobman SA, Riederer SJ, Lee JN, Suddarth SA, Wang HZ, Drayer BP, MacFall JR: Cerebral magnetic resonance image synthesis. Am J Neuro Rad. 1985, 6: 265-269.Google Scholar
- Redpath TW, Smith FW, Hutchison JM: Magnetic resonance image synthesis from an interleaved saturation recovery/inversion recovery sequence. Br J Radiol. 1988, 61: 619-24. 10.1259/0007-1285-61-727-619.View ArticlePubMedGoogle Scholar
- Zhu XP, Hutchinson CE, Hawnaur JM, Cootes TF, Taylor CJ, Isherwood I: Magnetic resonance image synthesis using a flexible model. Br J Radiol. 1994, 67: 976-982. 10.1259/0007-1285-67-802-976.View ArticlePubMedGoogle Scholar
- Gupta A, Lee VS, Chung YC, Babb JS, Simonetti OP: Myocardial infarction: optimization of inversion times at delayed contrast-enhanced MR imaging. Radiology. 2004, 233: 921-926. 10.1148/radiol.2333032004.View ArticlePubMedGoogle Scholar
- Desai MY, Gupta S, Bomma C, Tandri H, Foo TK, Lima JA, Bluemke DA: The apparent inversion time for optimal delayed enhancement magnetic resonance imaging differs between the right and left ventricles. J Cardiovasc Magn Reson. 2005, 7: 475-479. 10.1081/JCMR-200053534.View ArticlePubMedGoogle Scholar
- The pre-publication history for this paper can be accessed here:http://www.biomedcentral.com/1471-2342/10/19/prepub