Magnetic resonance imaging (MRI) has great potential as an imaging modality for guiding minimally invasive interventions because of its superior soft tissue contrast and the possibility of arbitrary slice positioning while avoiding ionizing radiation and nephrotoxic iodine contrast agents. The major constraints are: limited patient access, the insufficient assortment of compatible instruments and the difficult device visualization compared to X-ray based techniques. For the latter, resonant MRI markers, fabricated by using the wire-winding technique, have been developed. This fabrication technique serves as a functional model but has no clinical use. Thus, the aim of this study is to illustrate a four-phase design process of resonant markers involving microsystems technologies. The planning phase comprises the definition of requirements and the simulation of electromagnetic performance of the MRI markers. The following technologies were considered for the realization phase: aerosol-deposition process, hot embossing technology and thin film technology. The subsequent evaluation phase involves several test methods regarding electrical and mechanical characterization as well as MRI visibility aspects. The degree of fulfillment of the predefined requirements is determined within the analysis phase. Furthermore, an exemplary evaluation of four realized MRI markers was conducted, focusing on the performance within the MRI environment.
Magnetic resonance imaging (MRI) proves to be an excellent modality for guiding minimal-invasive interventions because of its unique capabilities, such as superior soft tissue contrast, the possibility of arbitrary slice positioning, functional information and the absence of ionizing radiation and iodine contrast agents. Additionally, recent development in MR hardware imaging techniques allows for obtaining MR images in near real-time (above 1 fps) with satisfactory resolution. New wide-bore MR scanners and open MR scanners provide improved access to the patient. Within the open MR environment, interventions have been performed, for example, endoscopic surgery , drug injections, spine therapy  and the placement of afterloading catheters for brachytherapy  using MR-compatible puncture needles and catheters. Thereby, non-magnetic metallic instruments, made of, for example, titanium or nitinol, cause hypointense artifacts . In contrast, polymer catheters can be almost completely invisible in the MR image. Several techniques have been proposed to highlight the instrument tip, thus enabling the precise positioning of the instrument. Some of these techniques allow MR tip tracking for automatic repositioning of scan planes in near real-time during interventions. The methods of monitoring and visualizing the instrument tip in MRI can be broadly categorized into active and passive approaches [8, 16, 22, 36, 42]. Active tracking provides superior visualization and allows for automatic tracking of absolute coordinates using small receive coils at the catheter tip with coaxial cable connection to the MR scanner. Another active method for device tracking is based on local magnetic fields generated by direct current (DC) fed conducting structures . A subtraction of MR images with and without the DC yields an image from which the device location can be extracted. The main drawback of active methods is the need for long conductors, which can result in induced radio frequency (RF) heating.
Passive approaches are more cost effective and easier to implement, but are less robust than active methods. A common passive approach is to embed a layer or coating containing paramagnetic particles onto the device. The homogenous static magnetic B0 field is locally distorted because of the magnetism of the particles. The corresponding signal loss results in a hypointense artifact within the MR image . However, passive device tracking remains challenging in particular for in vivo applications, e.g., where other unwanted inhomogeneities, such as air-tissue boundaries, are present. Furthermore, the artifact of such passive markers can obstruct the relevant anatomy  in particular during gradient echo-based imaging sequences and are difficult to detect in anatomical areas with low MR signal.
An attractive third method, which allows a direct depiction of the marker without relying on additional hardware and post processing modification , is based on resonant circuits tuned to the Larmor frequency of the protons that experience an external magnetic field . These circuits have already been used for interventional devices, such as catheters  and implants . Several studies [11, 19] have demonstrated an accurate automatic instrument tracking comparable to active methods. This enables a more autonomous MR image acquisition and navigation during interventional procedures. Most MRI markers examined in the literature were handcrafted using copper wire and soft soldering of non-biocompatible surface mounted devices (SMD) capacitors [12, 52, 58]. So far, no paper has been published that summarizes biocompatible alternatives and the whole manufacturing process of the resonant markers. Therefore, this paper is focused on the strategy of designing miniaturized and integrated resonant circuits using microsystems technologies. A selection of suitable fabrication technologies, such as aerosol-deposition process, hot embossing technology and thin film technology is explained in detail. Based on this guideline, the reader will be able to adapt the design strategy to a certain clinical application. Following this, the results of the evaluation of four types of MRI markers manufactured in different ways is described. Thereby, the tested MRI markers were designed with regards to a specific clinical application and, therefore, differ in their coil and capacitor designs.
The tuned MRI marker should resonate at the Larmor frequency, which is defined by
where represents the gyromagnetic ratio for nuclei of hydrogen atoms. MRI is based on the excitation of nuclear spins leading to deflection from their equilibrium positions to the flip angle α due to the incident circularly polarized B1 field at the Larmor frequency.
During their return to the energy-efficient state (relaxation), an echo signal can be measured. The transmit coil couples with the resonant circuit and an RF current IIND is induced in the micro coil of the resonator. Considering an equivalent parallel circuit (Figure 1), the micro coil can be represented as a parallel connection consisting of an inductance LP and an ohmic resistance RP . The following condition is applied for the magnitude of the current flowing through the equivalent inductance:
where Q is the quality factor of the resonant circuit. This current generates a linear polarized magnetic field locally, leading to an increased excitation of the spins in the adjacent surroundings of the micro coil. This substantially enhances the incident B1 field locally and, thus, the flip angle in close proximity to the marker. As a result, a hyperintense signal is obtained locally. The signal amplification is most prominent in gradient echo sequences with a low flip angle, however, it is present in most typical MRI pulse sequences [10, 13, 40]. The signal amplification highly depends on the quality factor of the resonant circuit, the geometry and orientation of the coil with regards to the B1 field and the applied flip angle α.
The design process for resonant circuits can be divided into four stages: planning and specification, realization of samples and prototypes, evaluation and analysis phase. Within the planning phase, requirements of the MRI markers for the respective MRI-guided intervention are defined. Based on this, the resonator design can be engineered and verified by simulations. The realization phase involves the fabrication of the MRI markers. Several tests are conducted throughout the evaluation phase in order to characterize the properties of the fabricated MRI markers. These results serve as basis for the subsequent analysis phase in which the degree of fulfillment of the requirements is determined.
The requirements for MRI markers are mainly influenced by the MR scanner configuration being used, the type of intervention and the instrument that needs to be equipped by the circuit. An overview of the influences on the MRI marker design is shown in Table 1. The local signal amplification effect depends mainly on the chosen MR sequence, the electric parameters of the MRI marker and the relative orientation of coil within the B1 field. Conversely, the instrument properties, i.e., the geometrical dimensions, the surface material and the stiffness, have a major impact on the selection of suitable fabrication technologies.
|Influencing factor||Effect (on)||Compensation (by)|
|B0 magnetic field strength||fL||–||Adjustment of coil design and/or capacitor design for tuning to the desired fL|
|B1 orientation||Magnetic flux (B1) through the coil surface||–||Adjustment of coil design for maximizing IIND|
|Orientation of the instrument||Magnetic flux (B1) through the coil surface||–||Adjustment of coil design for maximizing IIND|
|Conductivity and permittivity of the surrounding tissue||fres of the MRI marker||–||Adjustment of coil design and/or capacitor design for tuning|
|Blood flow||Mixture of locally excited and non-excited spins ||–||Increasing quality factor for compensating this flow effect|
|Geometrical dimensions||Acceptable increase of the instrument diameter in absolute terms||–||Selection of a fabrication technology with corresponding layer thickness|
|–||Embedding of the MRI marker into the instrument|
|Surface material||Adhesion||–||Pretreatment of the instrument surface [e.g., chemical/mechanical/physical (deposition of extra layer)]|
|–||Fabrication of carrier element with MRI marker and fixation on the instrument|
|Stiffness||fres of the MRI marker when bending flexible instruments||–||Total length of the MRI marker smaller than the bending radius of the instrument|
|–||Selection of a coil design that is less affected by the bending|
|Acceptable change of the instrument stiffness in absolute terms||–||Selection of a fabrication technology with corresponding layer thickness and material flexibility|
|–||Meander structure instead of straight-shaped structures for flexible instruments|
|Other instruments in the direct surrounding (e.g., guide wires)||fres of the MRI marker||–||Adjustment of coil design and/or capacitor design for tuning|
The resonant MRI markers are generally designed with a specific MR scanner in mind, as orientation of the static B0 field and the Larmor frequency are essential influencing factors. Nowadays, the most widely spread MR systems used for diagnostics are closed bore 1.5T and 3T scanners with longitudinal B0 field orientation, such as MAGNETOM Espree 1.5T and MAGNETOM Skyra 3T (Siemens Healthcare, Erlangen, Germany), 450 WB 1.5T and WB 750 3T (GE Healthcare, Milwaukee, IL, USA), Achieva 1.5T and Achieva 3.0T TX (Philips Healthcare, Best, The Netherlands). However, for interventional applications, alternative configurations such as open MR scanner (e.g., 1T Panorama HFO, Philips Healthcare, Best, The Netherlands) should also be factored in. The starting point of the coil design is the common orientation of the instrument during the intervention with respect to the static B0 field. Figure 2 gives an overview of standard coil designs, such as saddle coils (Figure 2A), spiral coils (Figure 2B), single loop coils (Figure 2C) and cylindrical coils (Figure 2D). The fabrication and examination of single loop, spiral, as well as cylindrical coils have already been presented by our group in [14, 17, 54] or [35, 60], respectively. Also, other groups demonstrated the suitability of such coil designs for MRI markers, such as single loop coils [12, 52], spiral coils  and cylindrical coils [11, 13, 40, 41, 58]. Figure 3 illustrates the most suitable coil designs for the two most relevant MR scanner configurations and typical instrument orientations. In the case of an intravascular intervention inside a closed bore MR scanner, the instrument is oriented parallel to the B0 field. Therefore, a saddle, spiral or single loop coil should be taken into consideration in order to provide the most efficient coupling. If the instrument is aligned perpendicular to the B0 field typically, for instance during needle insertion within an open scanner, a cylindrical coil is the means of choice. Finally, if there is no predictable orientation of the instrument during the intervention and a wide range of device angles have to be taken into account, a combination of two or more coil designs is recommended .
Most interventional instruments, such as needles, catheters and guide wires, have a cylindrical shape. Therefore, cylindrical or rolled-up plate [18, 57] capacitors may be used. Furthermore, the capacitance per area can be increased by realizing ultrathin multilayer capacitors . This reduces the size of the resonant circuit and significantly improves the form factor of the added circuit on the sensitive tip of interventional devices. In contrast, an SMD capacitor is bulky and can be fixed to one side of the device only by challenging soft soldering. This construction is mechanically delicate. Furthermore, SMD capacitors are not biocompatible. Besides using a single capacitor, it can be advantageous to integrate several smaller capacitors with detachable connections for tuning to the Larmor frequency .
In order to identify the required parameters of the components of the resonant circuit, simulations of the inductance, resistance and capacitance have to be carried out. The micro coil can be characterized numerically with the partial element equivalent circuits (PEEC) method [21, 48]. By applying the discretized PEEC model (Figure 4A), the inductance and resistance of the conductor structure can be estimated by solving a dense system of discretized equations iteratively using, for example, a generalized minimal residual (GMRES) algorithm . Furthermore, open source software packages (e.g., FastHenry2, FastFieldSolvers S.R.L., Italy) allow a segmentation of each element into several filaments. This segmentation enables taking the skin effect into account leading to reliable results for high frequencies in the MHz-range (Figure 4B).
For completion of the resonant circuit, a capacitor with the appropriate capacitance is required. Purpose-built capacitors require additional simulations for determining the capacity, considering different parameter combinations. The simulation is necessary in particular if an inhomogeneous material of the dielectric layer or complex geometries have to be taken into account. One way of determining the capacity is to discretize the electrodes into several elements. Then, the partial capacitances are calculated by computing the charge distribution and the voltage potential at each element. This can be realized with the open source software package FastCap2 (FastFieldSolvers). Figure 5A shows the model of a rolled-up plate capacitor with a total length of 5 mm and an inner diameter of 2 mm. The relative permittivity εr was set to 2.5. The gap d between the two electrodes was varied from 1 μm to 20 μm. The resulting capacity was computed using FastCap2 (Figure 5B). As to be expected, the capacity increased inversely proportional to the gap d.
Electromagnetic field simulations (e.g., HFSS, CST, Semcad, openEMS) of the entire MRI markers enable a precise estimation of their electrical properties as well as their impact within the MR environment. Simulations facilitate a realistic testing environment for the determination of the resonant frequency fL and quality factor Q. Thereby, the entire MRI marker can be excited by an additional test coil. The computation of the frequency-depended scattering parameters on the primary side allows for an analysis of the electrical properties of the MRI marker .
Additional simulations of the entire MRI markers within their aimed MR environment focusing on the influence on the B1 field distribution have to be carried out in order to assess the risk associated with RF-induced heating. Such simulations can be performed using phantom setups or human torso models. Figure 6A shows the model of a cylindrical coil tuned to the Larmor frequency of a 1.0 T MR scanner using a lumped capacity . The signal enhancing effect of the resonant MRI marker is shown in Figure 6B. A 90° shift between the exciting field B1 and the linearly polarized field of the marker was observed.
Based on simulation results, the resonant circuits can be fabricated. For the wire-winding as the simplest technique, an insulated wire is bent to the desired coil shape and attached to the instrument. Afterwards the coil is soft soldered to a non-magnetic SMD capacitor. This method serves as reference for the following methods: aerosol-deposition process, hot embossing technology and thin film technology. Table 2 summarizes various strengths and weaknesses of the presented technologies.
|Wire-winding technique||–||Simple manufacturing of 3D structures||–||Low reproducibility|
|–||Easy tuning (shapeability of coil)||–||High expenditure of time (manual work)|
|–||Low material costs||–||Requires soft soldering of additional capacitor|
|–||Low technological requirements||–||Enlargement of the instruments diameter|
|–||High mechanical stability under dynamic fatigue loading|
|Aerosol-deposition process||–||Simple manufacturing of 3D structures||–||High material costs|
|–||Adaption to diverse instrument shapes||–||Adhesion of the deposited layer is strongly depending on the surface character|
|–||Independence on instrument material||–||Low mechanical stability under dynamic fatigue loading|
|Hot embossing technology||–||High reproducibility||–||Dependence on instrument material (only thermoplastics)|
|–||Realization of embedded structures||–||Challenging manufacturing of 3D structures|
|–||Low material costs|
|–||High mechanical stability under dynamic fatigue loading|
|Thin film technique||–||High reproducibility||–||High material costs|
|–||Precise structures with high accuracy||–||Expensive equipment|
|–||Highly flexible substrates||–||Only fabrication of 2D structures onto carrier foils|
|–||Low mechanical stability under dynamic fatigue loading|
The aerosol-deposition process enables the fabrication of a resonant circuit directly on the instrument in three dimensions . For an improved manufacturing process, aerosol-like catalytic nano ink containing palladium seeds and adhesive is deposited on the surface. This serves as a starting layer for the subsequent electroless plating processes. The first deposition defines the geometry of one electrode of the cylindrical capacitor and the return path. After the metallization process a dielectric layer, e.g., a polyolefin tube or a lacquer, is attached to the instrument. In the last step, the geometries of the cylindrical coil and the second electrode of the capacitor are preset by another deposition process followed by a metallization process.
The approach of hot embossing enables the embedded fabrication of conducting paths into thermoplastic instruments . In case the instrument contains one or more lumina, a customized embossing core has to be inserted in order to avoid a permanent deformation of the instrument. A metallic foil with a typical thickness of 12 μm to 35 μm is attached to the surface of the instrument. Afterwards, a two-piece negative embossing stamp is positioned under as well as above the instrument. This stamp can be fabricated by milling a metallic tool, e.g., made of aluminum, or by structuring a silicon wafer . With adjusted pressure and temperature the embossing stamp presses the metallic foil into the instrument while simultaneously shearing the foil with its sharp edges. Subsequently to the embossing process, the remaining non-embossed foil is removed. The integration of a plate capacitor can be realized by embossing a pre-structured foil containing a dielectric interlayer.
The manufacturing process of the thin film technique  is divided in two steps; the fabrication of the carrier foil and the metallization of the conducting structures. The layer build-up is conducted onto a silicon wafer with a sacrificial layer that enables a removal of the completed resonant circuit. A photoresist polyimide layer is then applied by a spin coating process. The geometry of the carrier foil is defined by exposing UV light to the photoresist-coated polyimide while a pre-formed mask is affixed. Etching and cleaning processes remove the undesired areas of the polyimide layer. As a basis for the metallization, a plating of chrome and gold is applied to the carrier foil. Following stages are coating with photoresist, a masked UV exposure and an etching process. The electrically conductive pathways are then realized by galvanization of gold. Finally, a cleaning process removes the remaining photoresist and the plating.
Firstly, it needs to be ensured that the predefined geometrical dimensions are met. Further typical test scenarios of the realized MRI markers consider the investigation of electrical and mechanical properties and the estimation of the behavior in the MR environment, as well as biocompatibility and sterilization capabilities. These explanations are focused on technical preliminary investigations. In a further step, animal trials should follow as they enable testing of such instruments under more realistic conditions. These tests are aimed at the proof of functionality and safety. For a subsequent evaluation with patients, the equipped instrument must be conformal to the local regulatory authority (e.g., CE, FDA) and the clinical trials have to be approved by an Ethics Committee.
As the MRI markers are to be inserted into patients, they have to be encapsulated to avoid contact with the surrounding body fluids and tissue in compliance with ISO norm 13485 . This requires an indirect measurement of electrical parameters using an inductive coupling method. The method presented in  deals with this aspect. A resonant circuit is aligned centrically to a coupling coil connected to a vector network analyzer. If the quality factor is significantly higher than 1, the resonant frequency can be identified at the maximum value of the real part of the input impedance. By measuring the full bandwidth at half maximum (FWHM) of the curve, the quality factor amounts to the ratio of the resonant frequency and the bandwidth.
Test methods for the mechanical characterization of MRI markers can be inferred from the qualification process of flexible printed circuits. This allows a qualitative and quantitative assessment of the mechanical stability, the adhesive strength as well as ductility of the components and assemblies. The quality of the fabricated assemblies can be evaluated e.g., by peel test [2, 59], pull-off test , cross-cut test [4, 28, 33], bending test [29, 32], or scotch-tape-test [3, 27]. Furthermore, the influence of the MRI marker in regards to the mechanical properties of the instrument (e.g., axial stiffness, bending stiffness, torsional stiffness) has to be investigated. Suitable test methods have been presented in [7, 62].
In general, the main objective is to guarantee the safety for the patient. As electrically conducting structures might cause potentially RF-induced heating , the fabricated resonant MRI marker has to be tested accordingly. Thereby, the maximum temperature rise within a body-like phantom has to be measured according to ASTM standard F2182 . Additionally, the resonant MRI markers and the used materials must not generate significant susceptibility artifacts within the MR image . Artifacts reduce the diagnostic quality of the images and also reduce or cancel the signal amplification of the resonant marker. Thus, material tests within the MR environment according to the ASTM standard F2119  should be conducted at an early stage of the design process. Finally, the proof of functionality has to be provided. This includes tests of the marker’s visibility within the MR image for several coil orientations with respect to the B1 field using a homogenous phantom . In order to guarantee the marker’s functionality during a specific MR-guided interventional procedure, the test has to be conducted while applying clinically used pulse sequences. A quantitative analysis of the MRI markers quality can be achieved by recording a B1 field map. Alecci et al.  presented a method to record a B1 field map using a birdcage coil in transmit/receive mode using a spin echo pulse sequence and varying the transmitter power. By use of this method, the signal enhancing effect in the proximity of the MRI marker could be estimated.
Consequently, tests with inhomogeneous phantoms have to be conducted, and the automatic detection and tracking of the markers via image processing algorithms has to be verified [11, 19]. In case of cardiovascular instruments, tests within a vascular phantom, providing controllable blood flow and fluids with similar electromagnetic properties and relaxation times as blood, should be applied.
The testing of materials for biocompatibility is a multistep process. Hence, physical, chemical and biological studies have to be included. The process starts with the physical-chemical material characterization, followed by in vitro experiments. If it is impossible to use biocompatible materials only, a biocompatible coating of the circuit has to be applied. In this case, the biocompatibility tests can be employed either on the components, the assemblies or the final product. If there are no specific regulations for the physical and chemical tests related to medical products, the biological test is defined in ISO 10993 . In any case it has to be investigated if exclusively non-toxic, non-carcinogenic as well as haemo-compatible materials have been used at the interface between technical and biological systems. Furthermore, the final instrument must not cause mechanical stress of the tissue. Figure 7 shows a differentiation of the materials used for resonant MRI markers.
For the successive clinical use, the fabricated device including one or several resonant MRI markers must enable sterilization. Potential issues have to be considered for common sterilization methods. Typically, chemical sterilization methods applied to catheters and electronics utilize low temperature formaldehyde or ethylene oxide (EtO). For that reason, a chemical resistance is a necessity for the employed materials and electronics. Furthermore, plasma sterilization methods such as low-pressure inductively coupled plasma (ICP) or electron beam sterilization are applied. The occurring UV exposition, ion bombardment and the generation of free radicals are the cause of the sterilizing effect. These radiation types must not influence the functionality of the fabricated markers. We do not consider steam sterilization (or hot air sterilization) as sufficient heat resistance of the MRI marker would be very difficult to achieve due to the polymer metal compound.
The purpose of the analysis phase is to determine the degree of fulfillment of functional and non-functional requirements. If the predetermined specifications cannot be achieved, potential improvements, such as the selection of materials, parameter adaption for the fabrication technology and modification of the geometry, have to be identified. With regards to medical electronics packaging the highest level of reliability has to be guaranteed for a medical product. The fabrication process of resonant MRI markers implies different levels of system integration, such as electrical functional elements (e.g., conducting paths), physical elements (e.g., micro coil) and the whole assemblies group. Each level has its own sources of errors. For that reason, failures, driving forces and their mechanisms have to be detected for each level with the help of investigations and tests or a failure mode and effect analysis (FMEA). Following this analysis phase, the design cycle starts again with the planning phase.
The following section gives a brief description of the evaluation of four MRI markers manufactured with different technologies and with regards to a specific clinical application. Thereby, the most important aspect is the technical feasibility of the applied microsystems technology in combination with the functionality of the MRI marker. Therefore, our examination focuses on the performance within the MR environment and initially excludes aspects as mechanical characterization, biocompatibility and sterilization. The examined fabrication technologies were wire-winding technique as reference, aerosol-deposition process, hot embossing technology as well as thin film technique. All necessary parameters are summarized in Table 2.
As there is a relatively high number of studies focusing on wire-winded MRI markers [12, 24, 40, 41], this technique serves as reference to microsystems technologies. Our investigation targeted the equipment of a 5F intravascular catheter for use inside a 1.5T closed bore MR system. Therefore, the requirements were defined as stated in Table 3. The most appropriate coil designs for this setup are single loop, spiral and saddle coils. In consequence, the polymer catheter has been equipped with wire-winded resonant MRI markers consisting of a saddle coil and SMD capacitors (Figure 8A). The catheter was tested inside a water phantom using a balanced steady state free precession (bSSFP) sequence (sequence 1, Table 3). All resonant MRI markers generated a significant local signal enhancement (Figure 8B, top). In order to estimate the value for clinical applications, tests have been conducted on a Thiel cadaver model  (Figure 8B, bottom). Therefore, a 5F catheter containing one resonant MRI marker was tested while imaging with a fast gradient echo (FGRE) sequence (sequence 2, Table 3). As a result of the high quality factor, the tested resonant MRI marker generated a hyperintense signal enhancement.
|Wire-winding technique||Aerosol-deposition process||Hot embossing technology||Thin film technique |
|MR scanner unit||Closed bore system: 1.5T Signa HDx (GE, Waukesha, WI, USA)||Closed bore system: 3.0 T Achieva scanner (Philips Healthcare, Best, The Netherlands)||Closed bore system: MAGNETOM Skyra 3T (Siemens Healthcare, Erlangen, Germany)||Closed bore system: 1.5 T Achieva scanner (Philips Healthcare)|
|Targeted intervention||Intravascular||Placement of afterloading catheters for a subsequent brachytherapy||Intravascular||n/a|
|Instrument orientation||Parallel to||Perpendicular to||Parallel to||n/a|
|Instrument||5F polymer catheter||6F polymer catheter||6F polymer catheter||5F catheter|
|Coil||Shaping a copper wire (∅ 0.2 mm) to a saddle coil (l=11 mm) with two windings||Cylindrical coil with 16 windings (l=5 mm)||Single loop coil (l=23 mm)||Spiral coil
|Capacitor||Attaching MR-compatible SMD capacitors by soft soldering||Cylindrical capacitor with polyolefin tube as dielectric layer||Attaching MR-compatible SMD capacitors by soft soldering||Polyimide layer with one main capacitor and several detachable capacitors|
|Functional testing in MR unit|
|MR sequence||Seq. 1: bSSFP (TE/TR/α=3 ms/1.6 ms/25°; slice 8 mm; square FOV 200 mm; matrix 256×256)||Seq. 1: FFE (TE/TR/α=2 ms/6 ms/5°; slice 10 mm; square FOV 190 mm; matrix 432×432)||FLASH (TE/TR/α=2.5 ms/350 ms/1° and 8°; slice 3 mm; square FOV 220 mm; matrix 320×320)||GRE (α=20°)
(further details of the MR sequence n/A)
|Seq. 2: (FGRE) sequence (TE/TR/α=1.3 ms/3.8 ms/30°; slice 7 mm; square FOV 480 mm; matrix 256×128)||Seq. 2: FFE with background canceling using multi-cycle projection dephasers  (TE/TR/α=2…3.6 ms/4…7.4 ms/8°; slice 10 mm; square FOV 190 mm)|
|Test environment||Water phantom (seq. 1), Thiel cadaver (seq. 2)||Water phantom||Water phantom||Phantoms with Ringer’s solution and deionized water|
The objective of our investigation was to equip a 6F afterloading catheter with MRI markers using the aerosol-deposition process. Afterloading catheters are used for internal radiotherapy and serve as applicators for the insertion of radioactive seeds. The requirements for the design process were formulated focusing on the treatment of liver tumors within a wide-bore MR system. As the main orientation of the instrument during the placement of the catheter is perpendicular to the B0 field, a cylindrical coil design was realized. The functionality of MRI markers consisting of such a micro coil equipped with an SMD capacitor has been shown previously by our group . In order to transfer the realized MRI marker to a clinical application, the relatively large SMD capacitor has been substituted by a cylindrical capacitor fabricated by the same technology (Figure 9A). The main problems of the fabrication were insufficient adhesion (Figure 9B, top) and short circuits between adjacent conductors (Figure 9B, bottom). The functionality of the accurately manufactured MRI markers was proven within a 3.0 T Achieva scanner (Philips Healthcare, Best, The Netherlands) involving a fast field echo (FFE) sequence (sequence 1, Table 3). A significant signal enhancing effect of the MRI markers could be recorded (Figure 9C). These markers were also tested by applying a FFE sequence with background canceling using multi-cycle projection dephasers as described in  (sequence 2, Table 3). The functionality of one resonant MRI marker was tested comprising three different orientations with respect to the B0 field: parallel, angled by 45° and perpendicular. Furthermore, pixel sizes range from 1.0×1.0 mm2, and 1.5×1.5 mm2 up to 3.0 ×3.0 mm2. Figure 10 shows the result of this investigation. It can be stated, that the local signal enhancement in the surrounding of the resonant MRI marker can be seen clearly in all cases.
The fabrication of MRI markers onto planar substrates using the hot embossing technology has been presented by our group in . Our latest development was the adaption to the three-dimensional surface of a 6F polymer catheter for intravascular procedures. So far a single loop coil has been fabricated (Figure 11A) and tuned using a MR-compatible SMD capacitor. The main problems of the fabrication were conductor breaks caused by insufficient adhesion (Figure 11B) and unsatisfactory detachment of the metallic foil. MR visibility tests of several hot embossed MRI markers were performed inside a MAGNETOM Skyra 3T (Siemens Healthcare, Erlangen, Germany) using a fast low angle shot (FLASH) sequence (Table 3). We fabricated several MRI markers, which vary in their trace width (200 μm, 300 μm and 400 μm). The tested MRI markers generated a significantly amplified signal in their vicinity (Figure 11C). The observed signal losses occurred because of the susceptibility of the SMD capacitors.
The MR visibility of the MRI marker realized with the thin film technique was investigated by another research group and published in . The MRI marker was composed of a spiral coil with a dimension of 1.3×1.3 mm2, a main capacitor and several detachable capacitors for frequency tuning (Figure 12A, top). The foil was attached to a 5F catheter and sealed with a parylene layer (Figure 12B, bottom). The test was conducted within a 1.5 T Achieva scanner (Philips Healthcare, Best, The Netherlands) by using a gradient echo sequence with a flip angle of 20° (further details of the MR sequence were not mentioned). The fabricated MRI marker generates a clearly visible local signal enhancement (Figure 12B).
A design process particularly for the fabrication of MRI markers used for instrument visualization during MR-guided minimally invasive interventions has been presented. This process consists of four phases: planning and specification, realization of samples and prototypes, evaluation and analysis phase. Requirements for the MRI marker design and influencing factors on their usage during the intervention were described within the planning phase. The realization phase gives a brief introduction to the following fabrication technologies: wire-winding techniques, the aerosol-deposition process, the hot embossing technology as well as the thin film technique. The fabricated MRI markers have to fulfill several functional and non-functional requirements, such as technical functionality, biocompatibility and sterilization capability. Therefore, appropriate testing methods have been compiled from relevant literature and norms. Deduced from the test results, the degree of fulfillment of the predefined requirements is determined within the analysis phase. In case, the requirements are fulfilled insufficiently only, the design process restarts with the planning phase aimed at optimizing the current design.
As a further result, we evaluated four types of MRI markers realized by using the fabrication techniques mentioned above. The focus was to prove their functionality within the MR environment, i.e., the good visibility of the markers within the images. In fact, our MRI markers and also the thin film setups by Ellersiek et al.  showed a strong B1 enhancing effect. However, the current comparison suffers from a small numbers of samples and also from different MR systems and sequences. The limitations of this study can be overcome only by testing all fabricated MRI markers in vitro using identical test conditions. Furthermore, a measurement of a B1 field map would result in a quantitative and more comparative analysis of the fabricated MRI markers. Additional tests regarding the functional reliability, patient safety, biocompatibility and sterilization capability are necessary to bring the MRI markers to market readiness. This also includes animal studies and clinical trials.
Besides the mentioned fabrication technologies, other previously unapplied technologies such as micropenning should be taken into consideration. With this micro-capillary technology, flowable materials including electrically conductive, resistive and dielectric ink can be deposited onto an irregular 3D surface. This may open up a further possibility of fabricating resonant circuits directly onto the instrument. Further fabrication techniques might be low temperature cofired ceramics (LTCC) , flexible printed circuits (FPC)  and polymer thick film (PTF) technology [34, 43].
Nevertheless, it can be stated, that a fabrication directly onto the instrument is preferred. Using this manufacturing method a significant enlargement of the instrument’s diameter should be avoided. The main challenge is to guarantee a sufficient adhesion to the instrument, especially when mechanical stress is to be expected.
Overall, our paper gives an overview of a variety of different aspects that need to be considered during the design process of MRI markers. Thus, it enables a more structured development of resonant MRI markers, which will further advance the research in this field.
The research leading to these results has received funding from the Federal Ministry of Education and Research (BMBF) in context of the ‘INKA’ project (‘03IP71’) and in context of the Forschungscampus STIMULATE under grant number ‘03FO16102A’ as well as from the European Community’s Seventh Framework Programme (FP7/2007-2013) under grant agreement number 238802 (IIIOS project). The contribution of Andreas Brose to the fabrication of aerosol deposed MRI markers is gratefully acknowledged. IMSaT, Dundee, UK and the Intervention Centre, Oslo, Norway are gratefully acknowledged for provision of MRI expertise and testing of the MRI markers.
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