High-sensitivity in vivo contrast for ultra-low field magnetic resonance imaging using superparamagnetic iron oxide nanoparticles

Superparamagnetic nanoparticles will boost image contrast on portable MRI scanners operating at low magnetic fields.


Supplementary Note 1: Magnetization Calculations Based on Susceptibility Effects
To determine the magnetization (M ) of contrast agents (CAs) at 6.5 mT, bSSFP MRI images with banding artifacts caused by CA magnetization were acquired. The measured data were fit with simulated images calculated from equations for magnetization of CA vials and the bSSFP MRI signal equation. Examples of acquired and simulated images are shown in Figure 3.
To simulate image artifacts we used the model presented in reference 27, which is an analytical calculation of the magnetic field produced by an infinitely long vial of CA placed transverse to a static, x-directed, magnetic field (B 0 ). Defining B 0 in the x-direction is unconventional in MRI but appropriate in these calculations as it means the z-direction is parallel to the axis of the CA vial and conventional cylindrical coordinates may easily be used. For the water region of the phantom shown in Fig. 3b, the change in magnetic field (∆B x ) induced by magnetization of the CA is given by: where D 3 is a constant given by equation (2) for gadolinium based contrast agents and equation (3) for iron-oxide based contrast agents.
In these equations, χ 1 is the magnetic susceptibility of the region containing contrast agent (0 < r < R 1 ), χ 2 is the susceptibility of the glass vial (R 1 < r < R 2 ) and χ 3 the susceptibility of water (r > R 2 ). The variables r and ϕ are cylindrical coordinates defined relative to the center of the phantom and the B 0 field direction. R 1 and R 2 are the inner and outer radii of the CA vial.
H 0 is the applied magnetic field strength and M s is the CA magnetization. µ 0 is the permeability of free space.
From Equation 1, the shift in Larmor resonance frequency can be calculated (∆f L = γ H ∆B x ).
This frequency shift was used in conjunction with the repetition time (T R ) and tip angle (α) to produce an analytical bSSFP image using the bSSFP signal equation [20], assuming T 1 /T 2 = 1 which is true at 6.5 mT [17]:

Supplementary Note 2: Relaxivity
The longitudinal relaxivity (r 1 ) and the transverse relaxivity (r 2 ) of contrast agents were determined by fitting T 1 (inversion recovery) and T 2 (Hahn echo) measurements, respectively, to the concentration dependent relaxivity equation: Where T 1,2 is the 1 H relaxation time of the solution with CA, T 0 1,2 is the 1 H relaxation time in the absence of contrast agent, r 1,2 is the relaxivity coefficient, and [CA] is the concentration of contrast agent. 7 Table S1. SPION relaxivity at clinical field strengths. This table shows the relaxivity values of carboxylated, highly susceptible (HS) SPIONs measured in preclinical 3 T and 7 T MR Solutions MRI scanners. T 1 and T 2 values were obtained by fitting phantom imaging data acquired with a fast spin echo (FSE) sequence to standard magnetization recovery (T 1 ) and magnetization decay (T 2 ) models for various  due to the relatively low r 2 relaxivity of Gd-DTPA at 6.5 mT. We estimate that contrast would arise above 3.6 mM Gd-DTPA concentration (when T 2 becomes smaller than T E ) based on relaxivity data in Table I) but Gd-DTPA only reaches an in vivo blood concentration of 1.2 mM (assuming all contrast agent is initially distributed in the ∼25 mL blood volume of the rat [28]).      Figure S8. Rat body coil. The custom-built imaging coil was designed to accommodate a rat body with a high filling factor and resonated to the 1 H frequency of 276 kHz using an external capacitor board (series-match, parallel-tune). The nose cone at top was used for animal alignment and to deliver isoflurane anesthesia during imaging. This probe was built following the design process described in detail in Ref. 30.
Photo Credit: David Waddington, The University of Sydney.