Correction of the post-irradiation T1 relaxation effect for chemical exchange-sensitive MRI: A phantom study

Purpose: In many pulse sequences of chemical exchange-sensitive MRI including multi-slice chemical exchange saturation transfer (CEST) or chemical exchange sensitive spin-lock (CESL), there is a finite time delay between the irradiation preparation and the imaging acquisition, during which the T₁-relaxation reduces the chemical exchange contrast and affects the accuracy for volumetric imaging. We propose a simple post-acquisition method to correct this contamination.

Methods: A simple formula was derived to evaluate the cross-slice T₁-relaxation contamination in multi-slice echo-planar imaging (EPI) after the irradiation preparation. CEST and CESL experiments were performed on phantoms to examine the accuracy of this approach.

Results: Theoretical derivation showed that the cross-slice T₁-relaxation contamination in multi-slice EPI imaging can be corrected by the signals of each slice at a parameter that suppresses the signal, e.g., at the water frequency for CEST, or with very long spin-lock pulse for CESL. This formula was confirmed by the results of phantom experiments, for both long and short irradiation durations with and without a steady-state, respectively. To minimize the effect of B₀ inhomogeneity in the CEST experiment, a more accurate measurement of the signal at water frequency can be achieved with a higher pulse power and shorter duration.

Conclusion: We proposed and validated a simple approach to correct the cross-slice T₁-relaxation effect, which can be applied to volumetric CEST and CESL studies acquired by multi-slice EPI, or other imaging modalities with similar T₁-relaxation contamination.

Introduction

Chemical exchange saturation transfer (CEST) MRI is an emerging molecular imaging technique that can measure certain amounts of biomolecules with labile protons, including glucose, glycogen, amino acids, creatine, and phosphocreatine, as well as exchange rate-related environmental parameters such as pH, temperature, and the concentration of exchange catalysts [1–6]. Due to its versatility, CEST has shown great potential in the study of many diseases such as tumor, stroke, muscle pathology, cartilage and spinal cord injuries, and neurodegenerations [5–25]. In CEST, the labile proton of the biomolecule of interest is usually saturated by an irradiation module during which the bulk water signal is reduced due to the chemical exchange process, and the water magnetization is measured by an imaging module. To avoid the T₁-relaxation effect and fully capture the water magnetization, the post-irradiation delay, i.e., the time delay between the irradiation module and the k-space center of the imaging module, should be minimized.

CEST measures the Z-magnetization of the bulk water and is suitable for studies where the direct water saturation effect is small, e.g., the nutation frequency of the saturation pulse is much smaller than the chemical shift difference between the labile proton and water. In contrast, CESL is another chemical exchange-sensitive MRI technique that images the whole magnetization rather than the Z-component [26–29], and therefore, can be applied to studies where the direct water saturation is large, or even on the resonance of the water frequency. In on-resonance CESL [9, 30], the water magnetization after a spin-lock pulse is in the X-Y plane. For multi-slice imaging, it is often beneficial to flip the magnetization back to the Z-axis because the T₁-relaxation of water along the Z-axis is much slower than the T₂-relaxation in the X-Y plane.

A common imaging readout method for CEST and CESL MRI is EPI where the post-irradiation delay is negligible for single slice imaging. However, in multi-slice EPI, T₁-relaxation may contaminate the chemical exchange-sensitive signals for the slices acquired after the first slice, especially for the latter slices if the post-irradiation delay time is not negligible as compared to T₁. To minimize this cross-slice T₁-relaxation effect, post-processing using a T₁ map to compensate for the relaxation effect on CEST contrast has been reported [31], which needs a separate acquisition of the T₁ map and would thus lengthen the scan time. A similar T₁-relaxation effect is present in the multi-slice fast spin-echo sequence, and a combination of primary saturation and secondary saturation pulses has been proposed to compensate for this effect [16, 32], however, the choice of primary and secondary saturation may depend on pulse power and tissue properties, therefore careful calibration may be necessary. For multi-slice spin-lock studies, the T₁ contamination in the T_1ρ-weighted image may be eliminated by an RF cycling technique where the longitudinal magnetization is inverted immediately after alternate spin-lock preparation [33].

In this work, we propose a simple post-acquisition correction method to compensate for the relaxation effect which can be used in multi-slice CEST or CESL MRI, and validate the method by phantom experiments. Part of the results has been reported in ISMRM 2020 [34].

Material and methods

Theory

Assuming the delay between the end of irradiation and the beginning of the acquisition of the ith slice is t_delay (i) in a multi-slice CEST experiment, the saturated signal of the ith slice acquired at an RF offset of Ω can be expressed as

$S_{sat}\left(\Omega ,i\right)=S_0+\left\{S_{sat}^{corr}\left(\Omega ,i\right)-S_0\right\}\cdot e^{-t_{delay}\left(i\right)/T_1}(1)$

where S₀ is the fully relaxed signal, and $S_{sat}^{corr}$ is the saturated signal after the T₁ relaxation effect is corrected, or the ideal case when the T₁ relaxation effect is minimal, i.e., t_delay = 0, which is approximately the case for the first slice. At an offset of Ω = 0 (i.e., the water frequency), the water magnetization is fully saturated so that the signal in the absence of the T₁ relaxation effect should be zero, i.e., $S_{sat}^{corr}\left(\Omega =0,i\right)=0$. Thus

$S_{sat}\left(\Omega =0,i\right)=S_0+\left\{S_{sat}^{corr}\left(\Omega =0,i\right)-S_0\right\}\cdot e^{-\left[\frac{t_{delay}\left(i\right)}{T_1}\right]}=S_0\left(1-e^{-\left[\frac{t_{delay}\left(i\right)}{T_1}\right]}\right)(2)$

This signal is only related to T₁ and t_delay(i), and can fully account for the T₁ recovery during the post-irradiation delay. Combine Eqs 1, 2 to eliminate the exponential term, we have

$S_{sat}^{corr}\left(\Omega ,i\right)/S_0=\frac{S_{sat}\left(\Omega ,i\right)/S_0-S_{sat}\left(\Omega =0,i\right)/S_0}{1-S_{sat}\left(\Omega =0,i\right)/S_0}(3)$

Note the correction factor used here is

$\frac{S_{sat}\left(\Omega =0,i\right)}{S_0}=1-e^{-\left[\frac{t_{delay}\left(i\right)}{T_1}\right]}\approx \frac{t_{delay}\left(i\right)}{T_1}(4)$

when t_delay (i) << T₁. The asymmetric magnetization transfer ratio (MTR_asym) has been widely used to measure the CEST contrast, which is defined as

$MT{R}_{asym}\left(\Omega \right)=\frac{S_{sat}\left(-\Omega \right)-S_{sat}\left(\Omega \right)}{S_0}(5)$

From Eq. 3, the T₁ relaxation-corrected MTR_asym can be derived as

$MT{R}_{asym}^{corr}\left(\Omega ,i\right)=\left\{S_{sat}^{corr}\left(-\Omega ,i\right)-S_{sat}^{corr}\left(\Omega ,i\right)\right\}/S_0=\frac{MT{R}_{asym}\left(\Omega ,i\right)}{1-S_{sat}\left(\Omega =0,i\right)/S_0}(6)$

Eqs 3, 4 show that for both the Z-spectrum and the MTR_asym spectrum, the T₁-relaxation effect of multi-slice imaging can be easily corrected using the signals acquired from each slice at Ω = 0. Another index to describe the CEST contrast is the difference between the CEST signal at a specific frequency and a baseline with minimal CE signal. The latter can be obtained by model fitting, or by a specifically designed saturation scheme, e.g., a low duty cycle saturation pulse train in the ASEF approach [40]:

$∆S_{sat}\left(\Omega ,i\right)/S_0=\left\{S_{sat}^{baseline}\left(\Omega ,i\right)-S_{sat}\left(\Omega ,i\right)\right\}/S_0(7)$

Similarly, we can get

$∆S_{sat}^{corr}\left(\Omega ,i\right)/S_0=\frac{∆S_{sat}\left(\Omega ,i\right)/S_0}{1-S_{sat}\left(\Omega =0,i\right)/S_0}(8)$

The same correction approach can be applied to a CESL experiment with a finite spin-lock time (TSL), similar to the derivation above. A very long spin-lock time, e. g, five times the T_1ρ value, can be used to suppress the CESL signal, $S_{CESL}^{corr}\left(TSL=5T_{1\rho },i\right)=0$. Thus, we have

$S_{CESL}^{corr}\left(TSL,i\right)/S_0=\frac{S_{CESL}\left(TSL,i\right)/S_0-S_{CESL}\left(TSL=5T_{1\rho },i\right)/S_0}{1-S_{CESL}\left(TSL=5T_{1\rho },i\right)/S_0}(9)$

By analogy with the CEST case, the chemical exchange sensitive contrast in a CESL experiment can also be obtained as:

$∆S_{CESL}\left(TSL,i\right)/S_0=\left\{S_{CESL}^{baseline}\left(TSL,i\right)-S_{CESL}\left(TSL,i\right)\right\}/S_0(10)$

where the baseline indicates the signal with minimal CE effect. This can be obtained by a spin-lock pulse with a nutation frequency much higher than the chemical exchange rate of interest, such as >4,000 Hz as reported in recent studies [35, 36]. We can also get

$∆S_{CESL}^{corr}\left(TSL,i\right)/S_0=\frac{∆S_{CESL}\left(TSL,i\right)/S_0}{1-S_{CESL}\left(TSL=5T_{1\rho },i\right)/S_0}(11)$

For both CEST and CESL, the correction term is a signal that is fully suppressed at the first slice. Thus, the $S_{sat}\left(\Omega =0,i\right)$ term can also be used for the correction of CESL data to replace the $S_{CESL}\left(TSL=5T_{1\rho },i\right)$ term in Eqs 8, 10, if all the other imaging parameters are kept the same, such as the repetition time (TR) and echo time (TE). Note that the formula above is independent of the irradiation duration or TR, thus, the correction approach is fully applicable to cases of both CEST and CESL experiments where TR is not sufficiently long for water magnetization to fully relax before the next saturation, as well as CEST experiments where saturation pulse duration is not long enough to achieve a steady state.

MRI experiments

Phantom experiments were performed on a 9.4 T Bruker system at room temperature. A 4.0-cm inner diameter quadrature volume coil provided RF transmission and detection. After the volume of interest was shimmed, a B₀ map was obtained by the WASSR method, and a T₁ map was acquired using an inversion-recovery pulse sequence.

Three phantom experiments were performed:

1) 40 mM of creatine was dissolved in 1× phosphate-buffered saline (PBS) and titrated to pH = 7.4, with and without the addition of 0.15 mM MnCl₂ to reduce the water T₁ and T₂ relaxation times. CEST Z-spectra was acquired by a 2 μT pulse with RF offset ranging from 7 ppm to −7 ppm. The duration of saturation pulse is T_sat = 8 s, and the repetition time between two consecutive images of the same slice is TR = 13 s. The post-saturation delay t_acq was set to 50 ms or 100 ms for each slice.

2) 20% of bovine serum albumin (BSA) was dissolved in PBS, titrated to pH = 7.0, and heated to 75°C for 15 min. This heating process can only denature part of the BSA so that the resulting phantom will be a mixture of mobile and immobile proteins, which is similar to the tissue. CEST Z-spectra was acquired by a 1.5 μT pulse, with RF offset ranging from 7 ppm to −7 ppm. To confirm that our correction method applies to the transient state, a short T_sat of 2 s and TR = 4 s was used. In addition, the spectra near the water resonance were measured with 0.8 μT and 5 s pulse, and a 3 μT and 1 s pulse to compare the susceptibility of the correction factor to B₀ shift. The post-saturation delay t_acq was set to 100 ms for each slice.

3) For the 20% BSA phantom, CESL data were acquired using an on-resonance spin-lock sequence, where the spin-lock time (TSL) was varied from 0 to 100 ms. In addition, an image with TSL = 300 ms was used for correction following Eq. 8. Data were acquired for two repetition times of TR = 4 and 8 s, and t_acq of 126 ms for each slice.

Immediately after the irradiation preparation, a single-shot spin-echo EPI was used to acquire five slices. Other imaging parameters are 64 × 64 matrix size, 4 × 4 cm² field of view, 0.6-mm or 1.0 mm slice thickness with a 0.2 mm gap, echo time = 25 ms. Quantitative data were analyzed based on the region of interest (ROI), where only pixels with a B₀ shift of less than 10 Hz were included.

Results

Figure 1A shows the Z-spectra of creatine in PBS with t_acq = 50 ms. The cross-slice T₁-relaxation effect leads to a small difference in the Z-spectra of the five slices. Due to the relative long T₁ of 2.88 s, the largest delay after saturation is t_delay = 200 ms for the fifth slice, which is 7% of the T₁ value. Indeed, the residue signal detected at the water resonance with slice number and reaches a ∼7% for the fifth slice (Inset), in good agreement with Eq. 4. For a longer t_acq of 100 ms, the T₁-relaxation effect is more significant (Figure 1B), and the residue signal detected at the water resonance reaches a ∼14% for the fifth slice (Inset), doubling the value for t_acq = 50 ms. After correction with Eq. 3, the Z-spectra for all slices overlap very well for both t_acq = 50 and 100 ms (Figures 1C,D). The addition of 0.15 mM MnCl₂ shortens the T₁ of the creatine phantom to 0.65 s so that the cross-slice T₁-relaxation effect is amplified. With t_acq = 100 ms, there are large distinctions among the Z-spectra (Figure 2A) and the MTR_asym (Figure 2B) spectra measured from five slices. The peak of MTR_asym at 1.9 ppm for the last slice is reduced by 48% as compared to the first slice (#5 versus #1). Using the data at Ω = 0, the Z-spectra and the MTR_asym spectra (Figures 2C,D) from the five slices can be corrected well with Eqs 3, 5, respectively.

FIGURE 1

FIGURE 1. Multi-slice CEST Z-spectra of creatine in PBS. The raw Z-spectra show small difference of a few percent among the five slices with t_acq = 50 ms (A), and the difference becomes larger with a longer t_acq of 100 ms (B). The T₁-relaxation effects can be clearly seen near the water frequency in the Insets. (C,D) These differences can be satisfactorily corrected by Eq. 3, leading to overlapping Z-spectra for the five slices.

FIGURE 2

FIGURE 2. Multi-slice CEST MRI of creatine in PBS with MnCl₂. The T₁-relxation effect is very significant with t_acq = 100 ms in both the Z-spectra (A) and the MTR_asym spectra (B), which can be satisfactorily corrected by Eqs 3, 6, respectively, leading to overlapping spectra for the five slices (C,D).

Figure 3A shows the cross-slice T₁-relaxation effect in a BSA phantom heated to 75°C. This phantom is closer to in vivo tissue due to the presence of both mobile and immobile protein, as shown by the CEST and NOE signal dips at 2.8 ppm, −3.5 ppm, and the magnetization transfer (MT) effect at larger offsets (i.e., 6.5 ppm), and the separation between slices can be clearly seen for the whole −7 to 7 ppm range (Figure 3A). With a T₁ value of 1.9 s and a t_acq of 100 ms, $\frac{t_{delay}\left(\#5\right)}{T_1}=21\%$, the difference between the first and fifth slices is ∼22% at Ω = 0 which matches well with Eq. 4, and reduce to 3.2% at 7 ppm. Again, the T₁-relaxation effect can be corrected using Eq. 3 for the multi-slice spectra (Figure 3B). For CEST study, our approach relies on an accurate measure of S_sat (Ω = 0). When significant B₀-inhomogeneity is present, this data point can be obtained by interpolation and fitting of Z-spectrum, including the use of high order polynomials, or by a separate WASSR measurement. Alternatively, because the bandwidth of a saturation pulse is correlated with the B₁ power, a high-power pulse can saturate the water signal for a wider B₀ range. Figure 3C indicated that for a saturation pulse of 0.8 μT and 5 s, there was significant variation in S_sat (Ω = 0) of the heated BSA sample if the B₀ shift is large (e.g., >50 Hz). In contrast, the S_sat (Ω = 0) is insensitive to the B₀ shift when a pulse of 3 μT and 1 s were used (Figure 3D).

FIGURE 3

FIGURE 3. (A) Multi-slice Z-spectra of partially-denatured BSA shows the cross-slice T1-relaxation affects multiple signal contributions, including CEST, magnetization transfer (MT), NOE, and direct water saturation. (B) The cross-slice T1-relaxation can be satisfactorily corrected. The saturated signal near the water frequency is strongly dependent on the B₀ shift with a lower saturation power of 0.8 μT (C), and becomes insensitive to the B₀ shift with a higher saturation power of 3 μT (D).

In CESL experiments, the TSL-dependent signal of the BSA phantom is shown for TR = 8.5 s (Figure 4A). Similar to the CEST results, there is a clear T₁-relaxation effect for later slices. A mono-exponential fitting of the TSL-dependent data yields T_1ρ value of 48.7 ms and 68.8 ms for the first and fifth slice, respectively. For a shorter TR = 4 s, the T₁-relaxation effect is slightly larger than that of TR = 8.5 s for later slices (Figure 4B), as shown by the signals at TSL = 300 ms. After a correction with Eq. 8, the TSL-dependent data overlaps well for both TR values (Figures 4C,D), and the fitted T_1ρ value, averaged for the 5-slices, are 47.40 ± 0.75 ms and 47.38 ± 0.73 ms for TR = 8.5 s and 4 s, respectively, indicating good consistency.

FIGURE 4

FIGURE 4. CESL signal as a function of TSL for (A) TR = 8.5 s and (B) TR = 4 s before and (C,D) after the correction of the cross-slice T₁-relaxation effect.

Discussions

In CEST or CESL MRI, chemical exchange contrast can be affected by the T₁-relaxation between the irradiation preparation and the imaging module. The contamination is dependent on the ratio of t_delay and T₁, and thus, may be exacerbated at lower magnetic field strengths where T₁ is shorter. In this work, we reported a simple post-acquisition correction method with good accuracy. Multi-slice EPI was used as an example because the T₁-relaxation is minimal in the first slice, and thus, its first slice data can be used to determine the accuracy of the correction approach. Our approach only corrects the T₁-relaxation occurring between the irradiation pulse and the image acquisition, rather than the T₁-relaxation, if any, during the irradiation pulse. The CEST contrast competes with T₁ relaxation because the saturation transfer reduces the Z-magnetization of the water, whereas the T₁ relaxation brings the magnetization back to equilibrium. Therefore, the CEST effect is strongly affected by the water T₁, which is often corrected by a normalization of the water T₁ in CEST studies.

Besides EPI, imaging modules that have been used for CEST or CESL MRI include rapid acquisition with refocusing echoes (RARE), steady-state free precision (FISP), etc. In some of these imaging methods, there is a non-negligible delay between the end of saturation and the beginning of the central k-space even for a single slice imaging. For example, In FISP-CEST [37], the T₁-relaxation effect at the water resonance can be 10%–20% (as determined from the Z-spectrum signal at the water resonance frequency) for centric-encoding and even higher for linear-encoding, depending on the matrix size.

3D-CEST has recently been developed where the short saturation pulse and the imaging acquisition module are interleaved, so that the saturation reaches a steady state after a certain number of repetitions. Compare to multi-slice MRI, 3D acquisitions often require dedicated hardware for parallel imaging to speed up the acquisition. One drawback of this pulsed saturation approach is that the saturation frequency bandwidth is broader for shorter RF irradiation pulses which can affect the specificity of the CEST signal. Moreover, this approach has limitations because the optimal sensitivity for some CEST applications, such as those with a fast exchange, occurs with a relatively high power pulse at the transient state before reaching the steady-state.

This method is highly dependent on the accuracy of the correction term, i.e., the signal at 0 frequency offset for CEST or with long TSL for CESL. Thus, a high SNR for this correction image is critical. In addition, while correction methods that minimize the T₁-relaxation effect using specially designed irradiation schemes may be able to maintain CE contrast during acquisition [16, 32], post-acquisition correction such as the current method cannot artificially restore lost chemical exchange effects as the correction signal only contains T₁-information. Thus, our approach should only be applied to compensate for T₁-relaxation in cases where this relaxation has not fully diminished the chemical exchange contrast to noise ratio.

Although we only demonstrated our correction for CEST and CESL, this approach can be extended to other variants of chemical exchange sensitive MRI such as chemical exchange rotation transfer [38], on-resonance variable delay multiple pulse [39], or average saturation efficiency filter [40]. More generally, this method can also be applied to other imaging modalities where the imaging is acquired after a magnetization preparation and there is a delay between the preparation and the imaging module, such as multi-slice T₁-weighted imaging or diffusion-weighted imaging.

Conclusion

We reported a post-acquisition method to compensate for the T₁-relaxation effect in the CEST signal with multi-slice EPI imaging. This method is simple and easy to perform and can also be applied to other imaging modalities where there is a delay between the saturation and the imaging of the center k-space line.

Data availability statement

The raw data supporting the conclusion of this article will be made available by the authors, without undue reservation.

Author contributions

TJ derived the formula, designed the experiment, performed part of the experiments, and write the manuscript. JC performed part of the experiments and manuscript writing.

Funding

This work is supported by NIH grant NS100703.

Conflict of interest

The authors declare that the research was conducted in the absence of any commercial or financial relationships that could be construed as a potential conflict of interest.

Publisher’s note

All claims expressed in this article are solely those of the authors and do not necessarily represent those of their affiliated organizations, or those of the publisher, the editors and the reviewers. Any product that may be evaluated in this article, or claim that may be made by its manufacturer, is not guaranteed or endorsed by the publisher.

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