If you are familiar with MRI (or NMR in general), then probably also with the relaxation time constants T1 and T2. These tissue-specific (or substance-specific) constants describe how fast the nuclear magnetization returns to its equilibrium value, M0, after excitation by a pulsed radio-frequency (rf) field. Shortly summarized, T1 describes the exponential recovery of the longitudinal magnetization ML (i. e., of the magnetization parallel to the external static magnetic field B0), and T2 describes the exponential decay of the transverse magnetization MT (that is precessing in the plane orthogonal to B0).
Typically (as shown in the first figure), T2 values of tissue are considerably lower than T1 values, i. e., the transverse magnetization decays quicker than the longitudinal relaxation needs for recovery. For most tissues in vivo, T1 varies between about 300 ms and 3 s, while T2 varies between about 10 ms and 200 ms. Longer T2 relaxation times (up to about 3 s as well) are found for liquids.
So one may ask if there are good physical reasons for T2 values being shorter than or – at most – equal to T1. As a physicist, I’d start with checking some extreme cases, e. g., assuming that T2 is much longer than T1, i. e. T2 ≫ T1. Then the longitudinal magnetization can fully recover while at the same time some transverse magnetization would be preserved. As a result, the magnitude of the total magnetization (ML2 + MT2)0.5 would become greater than the equilibrium value M0 – which is physically impossible.
To analyze these properties of T1 and T2 in more detail, longitudinal and transverse relaxation can also be plotted together in a diagram showing the transverse magnetization on the horizontal axis and the longitudinal magnetization on the vertical axis. These diagrams show the evolution of the magnetization (for experts: in a rotating frame of reference) after a 90° rf pulse (all trajectories start at the lower right corner of the diagram). The left-hand side of the following figure shows three cases T2 = T1/2 (blue), T2 = T1 (green), and T2 = 2 T1 [sic!] (cyan); and all three curves show a “benign” behavior in that they lie in the shaded area below the black circle segment ML2 + MT2 = M02. This means that the magnitude of the total magnetization vector is always smaller than M0.
However, the right-hand side of this figure shows what’s happening if T2 becomes greater than 2 T1 – in this case, T2 = 3 T1 (red curve): Now the magnitude of the total magnetization vector increases above the physical limit of M0, i. e., the red line crosses the black border of physically benign behavior!
In fact, it can be shown (see appendix if interested) that the maximum T2 value, for which the red curve stays always below the black line, is exactly T2 = 2 T1. And as so often, almost everything that is physically possible is also realized in nature (although the case T1 < T2 < 2 T1 is really extremely rare), as described by Malcolm H. Levitt in his highly recommendable NMR text book “Spin dynamics” (2nd ed., section 11.9.2, note 13):
The case where T2 > T1 is encountered when the spin relaxation is caused by fluctuating microscopic fields that are predominantly transverse rather than longitudinal. One mechanism which gives rise to fields of this form involves the antisymmetric component of the chemical shift tensor (not to be confused with the CSA). […] Molecular systems in which this mechanims is dominant are exceedingly rare (see F. A. L. Anet, D. J. O’Leary, C. G. Wade and R. D. Johnson, Chem. Phys. Lett., 171, 401 (1990)).
So, the answer to the title question is: No, T2 can in fact be greater than T1 in very special circumstances, but it can never be greater than 2 T1.
The maximum T2 value, for which the red curve stays always below the black line, is exactly T2 = 2 T1. This can be seen by analyzing the inequality
First, we divide by M02 and set T2 = αT1 as well as β = exp(–t/T1), yielding
(β1/α)2 + (1 – β)2 = β2/α + 1 – 2β + β2 ≤ 1
which is (after subtraction of 1 and division by β)
β2/α – 1 – 2 + β ≤ 0 or β2/α – 1 ≤ 2 – β.
β is by definition (for positive t) between 0 and 1, so the right-hand side of the last inequality is a linear function descending from 2 to 1 (i. e. always ≤ 2). Its left-hand side has very different shapes depending on α: it is increasing from 0 to 1 for 0 < α ≤ 2 (since then the exponent 2/α – 1 ≥ 0); but it is going to infinity for β → 0 if α > 2 (since then the exponent 2/α – 1 < 0). So, the last inequality will not hold in the latter case for sufficiently small values of β, which means that non-physical behavior occurs if α > 2 or, using the definition from above, if T2 > 2 T1.
You may have heard that gadolinium-based MRI contrast agents can enhance or increase the signal of tissue. This is generally a good description of what’s going on. Here, however, I would like to argue why this is – strictly speaking – not true: contrast agents cannot really increase the signal available for MRI.
Some basic facts first: gadolinium-based contrast agents are very frequently used in clinical MRI to improve the image contrast as illustrated in the following example.
An important fact about (conventional) MRI contrast agents is that it’s never the contrast agent itself that is visible in MR images. Instead, the contrast agent changes the behavior of the atomic nuclei in its neighborhood – in clinical MRI, these are the nuclei of hydrogen, i. e. the protons. As a consequence, these protons now appear brighter in T1-weighted MRI than protons which are not influenced by the contrast agent.
So, apparently the proton signal is increased by gadolinium? Yes, apparently … Actually, there is always a maximum signal that is available for MRI and that depends on three major factors:
the number of available protons, which is related to the proton density ρ of the tissue: the more protons (per voxel), the higher the signal;
the magnetic field strength B0: the higher the field strength, the higher is also the (thermal) nuclear magnetization and, hence, the measured signal;
the receiver coil: the more efficient the receive system (the radio-frequency coil), the higher the signal.
But the presence of a contrast agent does not increase this maximum signal.
Instead, we are cheating: First, we artificially decrease the MRI signal by choosing short repetition times (TR). And only afterwards, a certain part of this suppressed signal is recovered due to the influence of the contrast agent! This is illustrated in the following diagram:
Obviously, the maximum signal, Smax, with contrast agent is exactly the same as the maximum signal without contrast agent – but we can obtain this maximum signal considerably faster (i. e, at shorter TRs). That’s why gadolinium can be described to increase the speed, but not the MRI signal. In agreement with this observation, no additional gadolinium-induced signal enhancement can be found in proton-density-weighted MR images (with very long TRs). But, hypothetically, if contrast agent could be distributed homogeneously in the tissue (which in reality is not possible), then PD-weighted MRI could be accelerated by using shorter TRs without changing the contrast.
In MRI, we are frequently interested in data from a single two-dimensional slice of the imaged subject or object – or actually in data from several such slices. These slices are then displayed as conventional 2D MR images. A procedure called slice selection is used to restrict our data acquisition to each single slice. To understand slice selection, one has to know that MRI is based on the resonant excitation of spins in a static magnetic field B0. Spins have a characteristic, so-called Larmor frequency ω = γB depending on the magnetic field B (the constant of proportionality γ is the gyromagnetic ratio). By applying a radio-frequency (rf) field with exactly this Larmor frequency, spins can be excited (i. e., they can be made to generate a measurable signal).
The basic idea of slice selection is to excite only spins in a single slice by (first) superposing a linear magnetic field gradient gz e. g. in z direction, resulting in the spatial field distribution (shown in the first figure):
B(x,y,z) = B0 + gzz.
(These linear magnetic fields or gradient fields are one of the basic ingredients of MR imaging. Each MR imager comes with built-in coils to apply gradient fields in all possible spatial orientations.)
Then, by choosing an rf frequency ωslc = 2πfslc, we can select a plane in space where
ωslc = γB(x,y,z) = γB0 + γgzzslc,
and only spins in this plane centered around zslc = (ωslc /γ – B0)/gz are excited (see first figure).
More advanced MRI techniques can excite several spatially separated slices at once by applying rf fields with more than one frequency – the difficult part is then to separate data acquired from these slices for reconstruction.
The nice idea realized in the paper by Koray Ertan et al. is to excite multiple slices not by applying rf fields with a mixture of several frequencies, but instead by modifying the gradient field to a spatially non-linear magnetic field. Consequently the mapping between frequencies ω and spatial positions (z) is no longer one-to-one, but several positions can correspond to a single frequency:
So, depending on the shape of the magnetic field variation two or more slices can be excited using a single excitation frequency. This has the advantage that we can use short and simple standard rf pulses for slice excitation. The obvious disadvantage is that it requires additional gradient hardware providing the non-linear magnetic fields, which is currently not available at existing MRI systems.
However, this may change in future, since there are currently some promising approaches for non-linear encoding fields. In particular, I’m thinking about an impressive MRM paper (doi: 10.1002/mrm.26700) by Sebastian Littin and colleagues published also this year, in which an 84-channel matrix gradient coil is presented, which is capable of providing very flexible linear or non-linear field configurations.