According to the MR method in general, the body of the patient to be examined is arranged in a strong, uniform magnetic field whose direction at the same time defines an axis (normally the z-axis) of the co-ordinate system on which the measurement is based.  The magnetic field produces different energy levels for the individual nuclear spins in dependence on the magnetic field strength which can be excited (spin resonance) by application of an electromagnetic alternating field (RF field) of defined frequency (so-called Larmor frequency, or MR frequency).  From a macroscopic point of view, the distribution of the individual nuclear spins produces an overall magnetization which can be deflected out of the state of equilibrium by application of an electromagnetic pulse of appropriate frequency (RF pulse) while the magnetic field of the RF pulse extends perpendicular to the z-axis, so that the magnetization performs a precession about the z-axis.  This motion of the magnetization describes a surface of a cone whose angle of aperture is referred to as flip angle.  The magnitude of the flip angle is dependent on the strength and the duration of the applied electromagnetic pulse.  In the case of a so-called 90° pulse, the spins are deflected from the z axis to the transverse plane (flip angle 90°).  The RF pulse is radiated toward the body of the patient via a RF coil arrangement of the MR device.  The RF coil arrangement typically surrounds the examination volume in which the body of the patient is placed. 
After termination of the RF pulse, the magnetization relaxes back to the original state of equilibrium, in which the magnetization in the z direction is built up again with a first time constant T1 (spin lattice or longitudinal relaxation time), and the magnetization in the direction perpendicular to the z direction relaxes with a second time constant T2 (spin-spin or transverse relaxation time).  The variation of the magnetization can be detected by means of receiving RF coils which are arranged and oriented within the examination volume of the MR device in such a manner that the variation of the magnetization is measured in the direction perpendicular to the z-axis.  The decay of the transverse magnetization is accompanied, after application of, for example, a 90° pulse, by a transition of the nuclear spins (induced by local magnetic field inhomogeneities) from an ordered state with the same phase to a state in which all phase angles are uniformly distributed (dephasing).  The dephasing can be compensated by means of a refocusing pulse (for example a 180° pulse).  This produces an echo signal (spin echo) in the receiving coils. 
In order to realize spatial resolution in the body, linear magnetic field gradients extending along the three main axes are superposed on the uniform magnetic field, leading to a linear spatial dependency of the spin resonance frequency.  The signal picked up in the receiving coils then contains components of different frequencies which can be associated with different locations in the body.  The signal data obtained via the receiving coils corresponds to the spatial frequency domain and is called k-space data.  The k-space data usually includes multiple lines acquired with different phase encoding.  Each line is digitized by collecting a number of samples.  A set of k-space data is converted to a MR image by means of Fourier transformation or other suitable algorithms. 
The lack of harmful effects on the examined patient and the operator make MR imaging well-suited for “interventional radiology”, wherein the acquired and reconstructed MR images are used to guide invasive procedures.  The general goal of imaged guidance is to apply imaged-based information to the processes of diagnosis and therapy.  Known MR imaging-guided therapy systems use pre-operatively acquired MR images to create anatomic models, which provide localization, targeting, and visualization of the 3D anatomy.  These models support pre-operative planning to define and optimize access strategies and to simulate planned interventions.  These models connect image coordinates with the actual position defined by an instrument's location in the surgical field.  Thus, they enable a surgeon to navigate and execute procedures with full knowledge of the surrounding anatomy. 
In a plurality of practical applications, shifts and deformations of soft tissues occur during surgery because of mechanical factors, physiological motion, swelling, or hemorrhage.  These changes may displace organs or their tissue components to such a degree that pre-operatively acquired MR imaging-based 3D models cannot be registered with the patient's actual anatomy.  In this situation the ultimate solution for accurate MR imaging-guided surgery is real-time intra-operative MR imaging or at least frequent updating of the volumetric MR images during interventional procedures.  This results in methods that can continuously detect changes of the position of various tissue components and locate the targets of the interventions and their environments in order to define trajectories to the lesion to be treated.  Hence, the justification of intra-operative MR imaging is the change in anatomy during surgeries or the change of tissue integrity during therapy.  The goal is to allow MR imaging-guided therapy to make full use of the anatomic and functional information accessible by current MR imaging methods.  By providing the physician with current MR image information, safety and efficiency of surgical or interventional procedures is significantly improved. 
A problem is that it is difficult in a surgical setting to optimally place the RF coils required for MR signal acquisition around the respective body portion in such a fashion that (i) a good signal-to-noise ratio (SNR) is obtained and (ii) a good access to the interventional field is assured for the physician.