Synopsis

We propose a novel, efficient diffusion method, called isotropic diffusion weighted MRI (IDWI), for measuring orientationally-averaged properties of tissue water diffusion, free from modulations due to anisotropy. Using efficient diffusion gradient sampling schemes, IDWI rapidly and accurately quantifies the mean apparent diffusion coefficient (mADC) over a wide range of b-values, along with other important rotation-invariant intrinsic microstructural parameters, such as the mean t-kurtosis. The ability to efficiently and effectively remove modulations due to anisotropy in images with high-b values may improve existing diffusion MRI techniques and spur the development and clinical translation of new methods with improved biological specificity.

Purpose

Methods

We acquired high-quality preclinical and clinical diffusion MRI datasets in fixed ferret brain (250x250x250µm³, TE/TR=36/700ms, bmax=13500mm²/s) and in vivo human brain (2.5x2.5x5mm³, TE/TR=93/7000ms, b_max=6000mm²/s), respectively. For both ex vivo and in vivo experiments we obtained two DWI datasets:

1. One dataset was acquired using dense angular sampling at multiple b-values to allow measurement of mADC-weighted diffusion weighted images (DWIs). The dataset was analyzed with generalized diffusion tensor imaging (GDTI)² to explicitly measure the HOTs up to order n=6, their generalized Traces, $$$TrD^{(n)}$$$, and the mean t-kurtosis, ($$$\overline{W}$$$). The generalized trace $$$TrD^{(n)}$$$ was computed form the elements of the n-th order diffusion tensor $$$D_{n_{x}n_{y}n_{z}}^{(n)}$$$ $$TrD^{(n)}=\sum_{n_{x}+n_{y}+n_{z}=n}^{}K_{n_{x}n_{y}n_{z}}\mu_{\frac{n_{x}}{2}\frac{n_{y}}{2}\frac{n_{z}}{2}}D_{n_{x}n_{y}n_{z}}^{(n)}$$, where $$$\mu_{n_{x}n_{y}n_{z}}=\frac{n!}{n_{x}!n_{y}!n_{z}!}$$$, and $$$K_{n_{x}n_{y}n_{z}}=1$$$ when $$$n_{x}$$$, $$$n_{y}$$$, and $$$n_{z}$$$ are all even, and 0 otherwise.

2. A second dataset was acquired using efficient sampling Schemes 1, 2, and 3 (Fig.1) with a maximum of 13 orientations, at each of the 3 or 5 different b-values over the same range as in the GDTI experiments. From this IDWI data we generated mADC-weighted DWIs at each b-value using linear combinations of the log signal attenuations averaged over the signals acquired with orientations from Schemes 1, 2, and 3, denoted by M₃(b), M₄(b), and M₆(b). respectively (Fig.2). From the mADC-weighted DWIs we computed $$$TrD^{(n)}$$$ and $$$\overline{W}$$$ for comparison with GDTI-derived values.

Results

The high accuracy and rotation invariance of ID-MRI over a large range of b-values is validated in Fig.3. In fixed ferret and live human brain tissues orientationally-averaged (i.e., mADC-weighted) DWIs derived from densely sampled GDTI data and from IDWI data acquired with different rotations of the sampling schemes in Fig.1, are in excellent agreement. Moreover, despite requiring significantly fewer DWIs, rotation-invariant microstructural parameters such as and derived with IDWI showed similar values to those computed by explicitly measuring the HOT components with GDTI (Fig.4). Fig.4 also illustrates the benefit of including a larger number of b-values to stabilize the measurement of from orientation-averaged DWIs, in agreement with^5,6. Compared to the fixed-brain experiment, in vivo IDWI showed slightly larger errors, especially at tissue boundaries, likely due to subject motion during the duration of the scan (Fig.3 and Fig.4).

The clinical potential of IDWI is best illustrated with measurements obtained at very high b-values (Fig.5). The in vivo (mADC-weighted) IDWI signal at b = 8500s/mm² reveals a tissue contrast that resembles the fractional anisotropy (FA)⁷ but may be less modulated by architectural features of the tissue (Fig.5 yellow arrows). At the same time, the mADC measured at high b-value (Fig.5D) shows improved contrast around the putamen and globus pallidus compared to the T2-weighted image (Fig.5 red arrows).

Discussion

Our results indicate that in both micro-imaging and clinical experiments, signal modulations from bulk anisotropy of water diffusion can be accurately eliminated using a 6^th-order tensor model. Despite the structural and architectural complexities of brain tissues, orientationally-averaged diffusion signals can be obtained from as few as 13 measurements even at large diffusion sensitization, while for smaller b-values fewer measurements are needed.

The efficient IDWI sampling schemes can reduce total scan duration (and subject/physiological motion) enabling a rapid and accurate assessment of orientationally-averaged water mobilities in clinical exams. Further acceleration can be achieved with single-shot isotropic diffusion techniques^8-10.

IDWI extends the clinical assessment of eloquent rotation-invariant parameters, such as the mADC, which has provided a sensitive, robust, reliable and quantitative clinical imaging biomarker for non-invasive detection and characterization of hypoxic ischemic brain injury¹¹, cancers, and other pathologies¹². Moreover, IDWI may enable new whole-brain clinical applications that require dense sampling of the orientation-averaged diffusion signal decays as a function of b-value^13,14.

Conclusions

Acknowledgements

References

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