Radiol Phys Technol (2014) 7:95–101 DOI 10.1007/s12194-013-0236-z
Heating and safety of a new MR-compatible guidewire prototype versus a standard nitinol guidewire Malgorzata Wolska-Krawczyk • Martin A. Rube Erwin Immel • Andreas Melzer • Arno Buecker
Received: 29 June 2013 / Revised: 2 October 2013 / Accepted: 3 October 2013 / Published online: 8 November 2013 Ó Japanese Society of Radiological Technology and Japan Society of Medical Physics 2013
Abstract Our purpose in this study was to examine heating of nitinol and polyetheretherketone (PEEK) guidewires during near-real-time MR imaging in an artificial vascular model an ‘‘aorta phantom’’. The first 100 cm of the nitinoland PEEK-based guidewires both 145 9 0.08 cm were immersed in a saline-filled aorta phantom. The probes of a fiber-optic thermometer were positioned at the tips of both wires. Balanced steady-state free precession (bSSFP) [TE 1.6 ms; TR 3.5 ms; flip angle (FA) 60°; field of view (FOV) 40 cm; matrix 256 9 256; specific absorption rate (SAR); 1.15 Watt (W)/kg] and spoiled gradient-echo (SPGR) (TE 1.8 ms; TR 60 ms; FA 60°; FOV 40 cm; matrix 256 9 256; SAR 1.15 W/kg) pulse sequences were acquired in a 1.5-T MR scanner with use of an 8-channel array coil. Temperatures were recorded while the phantom was placed centrally in the bore of a MR scanner and in an off-center position (x = 24 cm, y = -5 cm, z = -10/10 cm). The temperature of the nitinol guidewire increased by 0.3 °C (center) and 1.1 °C (off-center position) with use of the bSSFP and by 9.6 and 13 °C (off-center position) with use of the SPGR sequence. Only minor temperature changes up to a maximum of 0.4 °C were observed with the MR-compatible PEEK guidewire when any position or sequence was applied. The
M. Wolska-Krawczyk (&) Clinic of Diagnostic and Interventional Neuroradiology, Saarland Medical Center, Homburg, Germany e-mail: [email protected]
PEEK guidewire showed substantially lower heating as compared to the nitinol guidewire in near-real-time imaging sequences in a phantom. Keywords Radiofrequency heating MR safety Specific absorption rate Abbreviation ASTM American society for testing and materials bSSFP Balanced steady-state free precession CT Computed tomography FA Flip angle FIESTA Fast imaging employing steady-state acquisition FLASH Fast low angle shot FOV Field of view f/s Frames per second GW Guidewire iMRI Interventional magentic resonance imaging MRI Magnetic resonance imaging PEBAX Polyether block amide PEEK Polyetheretherketone RF Radiofrequency SAR Specific absorption rate SPGR Spoiled gradient-echo SPIO Superparamagnetic iron oxide T Tesla TE Time to echo TR Repetition time
M. Wolska-Krawczyk A. Buecker Clinic of Diagnostic and Interventional Radiology, Saarland Medical Center, Homburg, Germany
M. A. Rube E. Immel A. Melzer Institute of Medical Science and Technology, Dundee University, Dundee, UK
Interventional magnetic resonance imaging (iMRI) is not a well-recognized field in medicine, but has been introduced as a major subject heading in PubMed in 2007 . Clinical
progress has been made for non-vascular interventions, but the advent of vascular interventions is still hindered mainly by the lack of an MR-safe guidewire (GW) and other interventional devices and implants. Radiofrequency (RF)induced heating of commercially available metallic wires inserted in the MR field makes these devices (i.e., catheters and GWs) unsafe due to possible damage to surrounding tissues . Stainless steel-braided catheters can be replaced by plastic catheters endowed with radiopaque markers, but eliminating the heating of metallic GWs is still a major challenge. According to Maxwell’s theory, the temperature in a MR scanner can arise from eddy currents induced by RF pulses and from induction loops built, for example, from electrocardiographic electrodes . In these two mechanisms, there is no storage of electrical energy within the device [3, 4]. The energy, however, accumulates when excessive heating is evoked around and in conductors (i.e., GWs) by resonating waves along them . If resonance of the GW occurs, the reflected waves are spreading along the longitudinal axis of such a wire, which acts like a dipole antenna, forming standing RF waves . The occurrence of resonance, and hence the heating of the GW exposed to RF pulses, depends on its position in the magnetic field, its length, and, if examined in a solution, on its immersed length . For example, in the 1.5-Tesla (T) MR scanner, an excessive heating hazard occurs when conductors longer than 15–18 cm are used . According to theoretical models, it can be assumed that GWs shorter than a quarter of the wavelength should be MR-safe . If the minimum length for resonance is exceeded, the occurrence is unpredictable due to environmental factors like the patient’s body shape or the position within the MR scanner. Experiments performed in a swine model repeatedly demonstrated sparks at the GW end [6, 7]. Thermal effects of the applied RF power in the MR field also depend on the specific absorption rate (SAR) . The SAR represents the RF energy absorbed per unit mass of the tissue  and is a pivotal parameter for high RF–MR applications. The local SAR limits, which indicate the heating of specific points within the MRI field, can only be estimated. Especially the local maxima of the SAR, the so-called ‘‘hot spots’’, are hazardous for the tissues [10, 11]. Several studies have addressed this issue of RF heating on simulation models and on healthy volunteers in 1.5- and 3-T MR fields. In these studies, the authors made use of temperature mapping or performed the post-processing of the measured amplitudes and phase of the MR field [9, 12]. The limitations of these methods include the possible underestimation of the local SAR by neglect of the non-measurable magnetic field components. For this reason, novel approaches for the design of MR-compatible and MR-safe metal-free GWs are desirable . According to the ICNIRP (International commission on non-ionizing radiation protection)
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statement on medical magnetic resonance procedures: protection of patients , no adverse health effects are expected when the body temperature in an MR field does not increase by 1 °C, except for infants, pregnant women, and patients with cardiovascular impairment. Moreover, the allowable body temperature is 38 °C for the head, 39 °C for the trunk and 40 °C for the extremities. Adair et al.  demonstrated that, in patients exposed to a whole-body SAR of 4 W/kg, the body temperature will raise by 0.6°. However, the increase is dependent on environmental conditions, which are difficult to predict. On this account, the SAR levels in the ICNIRP statement are valid at environmental temperatures below 24 °C for average scan times of 6 min. Depending on the body region, the whole-body SAR should not exceed 2 W/kg in normal operating mode (suitable for all patients) and 4 W/kg in controlled operating mode (under medical supervision), and the local SAR averaged over 10 g of tissue should not exceed 10 W/kg for the head and trunk and 20 W/kg for the extremities in normal and controlled operating modes. The materials used in the cores of commercially available GWs are mostly nitinol and stainless steel. Nitz et al.  demonstrated the potential hazards of nitinol-based GWs in the MR environment. The authors performed several experiments with a standard 0.08-cm nitinol GW (Terumo, Tokyo, Japan) of different lengths that was placed in different positions (in and off-center) of the scanner and suspended in the air as well as immersed in a saline solution filled phantom simulating patient tissue conductivity. The temperature alteration during real-time gradient-echo sequences was measured. The results demonstrated that the voltage spreads toward both ends of a conducting wire. The farther away from the magnet’s isocenter the wire is placed, the greater the temperature measured at its distal tip. On the other hand, the temperature increased in this setting as a function of the flip angle, and this increase did not exceed 1° for flip angles below 30°. Thus, non-conductive and non-ferromagnetic materials such as polymers [polyetheretherketone (PEEK), Aramid (KEVLAR), or glass fibers embedded in epoxy resin] were introduced in the first prototypes of MR-compatible GWs and are considered to be metal-free alternatives . Our aim in the present study was to compare the heating of a standard commercially available nitinol-based GW (Terumo, Tokyo, Japan) with a novel PEEK-based GW (EPflex, Dettingen, Germany) on an aorta phantom at 1.5 T during two ‘‘nearreal-time’’ imaging sequences.
2 Materials and methods 2.1 MR-compatible guidewire The novel PEEK GW of 0.08-cm diameter and 145-cm length was applied (Fig. 1) in all experiments. The GW
Nitinol vs. PEEK guidewire in 1.5-T MR scanner
Fig. 1 MR-compatible PEEK guidewire (EPflex, Dettingen, Germany)
consists of a high-strength para-aramid synthetic fiber (KevlarÒ) core, surrounded by a bending-resistant highperformance polymer (PEEK) and a hydrophilic coating. The tip section is conically tapered to create an atraumatic tip. Six markers consisting of superparamagnetic iron oxide (SPIO) markers are then applied. The first SPIO marker is positioned at the tip, and the other markers are positioned along the tip section of the wire. The tip section is then coated with a copolymer [polyether block amide (PEBAX)] to generate the atraumatic tip section. In a static magnetic field B0, the SPIO markers will be magnetized up to saturation magnetization, while each particle distorts the local magnetic field. This leads to a magnetic dipole–dipole interaction between the markers in such a manner that they are aligned following the magnetic field lines of B0 . Therefore, SPIO markers create susceptibility artifacts in MR images and allow passive visualization of the GW during MRI. Due to the small size of the SPIO particles and the very limited area coated with these markers, resonance effects cannot occur, thus excluding the possibility of RF heating . Besides a hydrophilic coating, which provides gliding properties, the GW also has torque stability and bending resistance. 2.2 Nitinol-based guide wire The nitinol-based GW (Terumo, Tokyo, Japan) was used as a standard GW. The same diameter and length of the nitinolbased MR-compatible GW were selected (0.08, 145 cm). The GW has a nitinol alloy core, which restores the shape during a procedure. The core is covered by tungsten powder in a polyurethane jacket, which results in radiopacity under X-ray fluoroscopy. The hydrophilic coating allows easy gliding of the GW inside the vessel and provides elasticity and torque. 2.3 Acceptance testing The compatibility of the MR GW for a field strength of 1.5 T was assessed in a clinical 1.5 T MR system (Signa
HDxt, GE Medical Systems, Waukesha, WI, USA) and in accordance with the following standard test methods proposed by the American Society for Testing and Materials (ASTM): the standard test method for measurement of magnetically induced displacement force on medical devices in the MR environment (ASTM International. F2052-06: standard test method for measurement of magnetically induced displacement force on medical devices in the magnetic resonance environment 2006) and for measurement of magnetically induced torque on medical devices in the MR environment (ASTM International. F2213–06: standard test method for measurement of magnetically induced torque on medical devices in the magnetic resonance environment 2006). The whole GW and also the tip section, separated from the rest of the GW, were tested for evaluation of the influence of the SPIO markers on MR compatibility. 2.4 Phantom experiments An aorta-shaped elastic vascular phantom of natural size was used during the experiments (Fig. 2). The model consisted of a polymer tube 2.5 cm in inner diameter resembling the abdominal aorta and rubber tubes of 0.8-cm inner diameter representing the iliac vessels. The wall thickness of all tubes was 2.3 mm. All vessels were connected with 0.8-cm plastic tubes, allowing fluid circulation between them. The phantom was completely filled with 0.9 % saline solution to mimic the human body conductivity. The first 100 cm of 145-cm long nitinol-based GW was inserted in the phantom. The temperature probe of a fiber-optic thermometer (Fotemp 4, OPTOcon AG, Dresden, Germany) was attached to the distal tip of the GW by a thin thread. Only one probe was attached to the GW examined because it was difficult to insert such a device via the 6F sheath into the phantom. One of the probes (called reference probe) was also attached to the MR table as a reference probe for the temperature in the MR room. The accuracy of the
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The first sequence was run continuously for exactly 60 s, the second for a total of 12 min (720 s). The SAR values were calculated by the MR scanner.
3 Results 3.1 Acceptance testing
Fig. 2 Aorta-shaped phantom with nitinol-based guidewire. Black arrow points to guidewire
measuring device was ±0.2 °C. The aorta bifurcation of the phantom was exactly in the middle of a multichannel coil, which was wrapped around the phantom. The GW tip was placed 3 cm above the aorta bifurcation in all experiments. The phantom was then placed in a 1.5-T MR system (Signa HDxt, software release: 15.0M4A0947, GE Medical Systems, Waukesha, WI, USA), with the artificial aorta in the isocenter and subsequently in an off-center position (x = 24 cm, y = -5 cm, z = -10/10 cm), as close as possible to the MR scanner bore walls, but still within the field of view (FOV) of the 1.5-T MR scanner (48 9 48 9 48 cm). The z-position describes the tip section of the GW that is inserted into the phantom. The temperature measurements were recorded with 1-s temporal resolution during two gradient-echo sequences, with selection of only one slide. The GW temperature was recorded between the MR scans, so that the next scanning began after the GW had cooled down to an initial temperature of approximately 20 °C: 20.3–20.8 °C. The experiment was then repeated under the same conditions with an MR-compatible GW. 2.5 MR imaging The experiments were conducted in a specially designed MR surgical suite (IMSaT, Dundee, UK) equipped with a 1.5-MR scanner (Signa HDxt, software release: 15.0M4A0947, GE Medical Systems, Waukesha, WI, USA). An eight-channel phase-array coil was applied in every sequence (8Ch Body FullFOV, GE Medical Systems, Waukesha, WI, USA). Two gradient-echo sequences were run continuously, namely, bSSFP (FIESTA: fast imaging employing steady-state acquisition: TE 1.6 ms; TR 3.5 ms; FA 60°; FOV 40 cm; matrix 256 9 256; SAR 1.15 W/kg) and spoiled gradient-echo (SPGR: TE 1.8 ms; TR 60 ms; FA 60°; FOV 40 cm; matrix 256 9 256; SAR 1.15 W/kg).
No magnetically induced displacement angle was measured according to ASTM standards (ASTM International. F2052-06: standard test method for measurement of magnetically induced displacement force on medical devices in the magnetic resonance environment 2006) at the position of the strongest magnetic field gradient for the whole GW and also for the separated tip section containing the SPIO markers. Therefore, we conclude that, the GW with the SPIO markers is not affected by any magnetically induced displacement force at a field strength of 1.5 T. No magnetically induced deflection angle of the equilibrium position was measured according to ASTM standards (ASTM International. F2213–06: standard test method for measurement of magnetically induced torque on medical devices in the magnetic resonance environment 2006); we used a torsional pendulum method, while the whole GW or the GW tip section was positioned in the isocenter of the magnet. Therefore, we conclude that, the GW with the SPIO markers is not affected by any magnetically induced displacement torque at a field strength of 1.5 T. 3.2 Balanced steady-state free-precession (FIESTA) sequence acquisition for 60 s The results are presented in Figs. 3 and 4. In the first scenario (Fig. 3), when the nitinol-based GW was placed in the isocenter of a 1.5-T MR scanner during the 60-s bSSFP sequence, only minor temperature changes of both GWs were observed. The temperature of the nitinol-based GW increased from an initial 20.3 °C to a maximum of 20.6 °C. Temperature changes within the measurement error of the fiber-optic thermometer (±0.2 °C) of 0.1 °C were observed for the PEEK-based GW. The reference probe attached to the MR scanner table did not change the initial temperature (20.3 °C). The second position in the MR scanner, when the aorta phantom was placed close to the MR scanner walls, was performed to simulate a worst-case scenario for possible heating (Fig. 4). The temperature of a nitinol-based GW in this off-center position rose from an initial 20.3 to 29.9 °C. The PEEK-based GW demonstrated no relevant
Nitinol vs. PEEK guidewire in 1.5-T MR scanner
35 Nitinol guidewire
Nitinol guidewire PEEK guidewire Reference probe
20 0 0
Fig. 3 Temperature measurements of PEEK and nitinol guidewires during FIESTA sequences run for 60 s in IN-CENTER position
Fig. 5 Temperature measurements of PEEK and nitinol guidewires during SPGR sequences run for 720 s in IN-CENTER position
Nitinol guidewire PEEK guidewire Reference probe
30 Nitinol guidewire PEEK guidewire 25
Fig. 4 Temperature measurements of PEEK and nitinol guidewires during FIESTA sequences run for 60 s in OFF-CENTER position
Fig. 6 Temperature measurements of PEEK and nitinol guidewires during SPGR sequences run for 720 s in OFF-CENTER position
temperature elevation. The reference probe showed a minimal increase in temperature of 0.1 °C, which was within the measurement error interval of the fiber-optic thermometer (±0.2 °C).
0.1 °C. The reference probe showed an increase from 21.2 to 21.3 °C. Both temperature changes of the PEEKbased GW and the reference probe were within the measurement error of the fiber-optic thermometer (±0.2 °C). The estimated SAR for both sequences equaled 1.15 W/kg. The peak local SAR for every sequence equaled 2.30 W/kg.
3.3 Spoiled gradient-echo (SPGR) sequence acquisition for 720 s The second sequence, i.e., spoiled gradient-echo, was run continuously for 720 s (i.e., 12 min) to imitate the duration of imaging during an endovascular intervention. The results of these experiments are presented in Figs. 5 and 6. As demonstrated in Fig. 5, during 720 s the temperature at the nitinol-based GW tip increased from 20.3 to 21.4 °C when the phantom was placed centrally in the magnet bore. The PEEK-based GW demonstrated a temperature elevation of maximally 0.4 °C. The reference probe showed a temperature increase from 20.4 °C up to 20.8 °C. The off-center position demonstrated temperature changes for the nitinol GW from 21.8 to 34.8 °C (Fig. 5). The same experiment repeated with the PEEK-based MRcompatible GW demonstrated a temperature elevation of
4 Discussion The availability of MR-compatible and safe GWs is a prerequisite for MR-guided vascular interventions. Vascular procedures performed without a GW, especially in small-caliber vessels, may result in an increased risk of complications such as dissection or perforation of the vessel. There are a few exceptions where interventions can be performed without a GW, and these, mostly in pediatric patients, have already been examined successfully . The conductive properties of commercially available GWs exclude their safe use in iMRI. Excessive RF heating
may cause injury of the vessel wall, bleeding, and even patient death due to arrhythmia in a worst-case scenario. The technique of combining polymers and Aramid appears to yield a non-metallic GW with properties adequate to replace metallic wires. The tested novel GW with a PEEK core is non-conductive and at the same time visible in MR sequences. It has been demonstrated previously [20, 21] as a proof of concept in an animal model that such a GW— from a mechanical point of view—can be used for vascular MR-guided interventions. In these animal studies, however, the temperature of the wires was not recorded. Before the GW is applied in patients, potential heating has to be examined. In our study, the two applied sequences were designed to simulate a real-time imaging scenario during an MR-guided intervention. The nitinol GW showed a maximum temperature of 13 °C during the 12 min of spoiled gradient-echo (SPGR) sequence and 9.6 °C during the 60-s bSSFP (FIESTA) sequence. Both temperature elevations were observed while the phantom was placed eccentrically in the magnet bore. The amount of applied RF energy, and thus, the SAR value, is influenced by various parameters such as the pulse sequence, the transmit coil, and the size, shape, and electromagnetic and thermodynamic properties of the object within the FOV . Most frequently, a body coil is used for transmission in MRI, and in this case the SAR is a function of the radial distance to the center of magnetic field . Hence, the offcenter position close to the magnet bore, is the worst-case scenario for RF-induced heating in the MRI environment [16, 23, 24]. During most interventions, if not all—such a position will not be achieved. However, we deem any kind of major temperature elevation (i.e., more than 1 °C) to be hazardous. Consequently, we have examined the worstcase to be sure that there will be no harmful coupling of the RF energy with the GW. The nitinol GW did not show any substantial temperature elevation when the phantom was placed in the magnet isocenter. The minor temperature fluctuations of 0.6 °C during the bSSFP (FIESTA) sequence and of 1.1 °C during the spoiled gradient-echo (SPGR) sequence, should not be harmful to tissue and do not exceed the limits proposed by the ASTM (F2182-09) of more than 0.6 °C [0.5 °C for the normal mode (suitable for all patients), and 1 °C in the first level controlled mode (under medical supervision)]. The temperature, however, could increase when pulse sequences with a generally high SAR level are used, i.e., fast spin-echo sequences that use many 180° RF pulses for image acquisition. These two sequences (bSSFP, SPGR) were used because they were applied for MR guidance of interventions at the time. In the experiments in which we used the PEEK-based GW, we demonstrated that this GW does not show any
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substantial elevation of the temperature. A maximum temperature rise of 0.4 °C was observed during the 12-min SPGR sequence, while the wire and phantom were placed in a central position of the magnet bore. The reference probe showed the same temperature elevation, proving that the measured temperature for the PEEK-based GW was in the range of measurement variation. It was demonstrated that the PEEK GW showed no temperature increase beyond the measurement error of ±0.2 °C. One limitation of our study is the single temperature measuring point at the distal tip of both GWs. However, it was demonstrated before that heating occurs mainly at the GW tip . A further drawback is the lack of a humanshaped phantom, because it is known that the object measured in the MR scanner influences the magnetic fields (B0 and B1). A human-shaped phantom (i.e., an aqueous gel-filled phantom of head and torso [25, 26]), however, would not allow the extreme off-center position of our aorta-shaped phantom, which we used as a worst-case scenario. Other limitations of our study are the lack of cooling of the blood flow in the vessel model, and the difference in conductivity as compared to the human body. Our study might, however, mimic the frequent presence of an occluded vessel, which is often the case in elderly patients. In conclusion, our investigation supports the lack of conductivity of the PEEK GW in a 1.5 T field. No relevant temperature elevation was measured during near-real-time gradient-echo sequences including SPGR and FIESTA. Acknowledgment The research leading to these results has received funding from the European Community’s Seventh Framework Programme (FP7/2007-2013) under Grant Agreement no 238802 (IIIOS project). Conflict of interest of interest.
The authors declare that they have no conflict
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