Leads with high surface resistance

Implantable medical leads having resistance characteristics adapted to dissipate radio frequency (RF) electromagnetic energy during medical procedures such as magnetic resonance imaging (MRI) are disclosed. An illustrative medical device includes a lead having an inner electrical conductor operatively coupled to an electrode and a pulse generator, and one or more outer resistive shields that radially surround the inner conductor and dissipate RF energy into the surrounding body tissue along the length of the lead.

TECHNICAL FIELD

The present invention relates to implantable medical devices. More specifically, the present invention relates to implantable medical leads having impedance characteristics adapted to dissipate radio frequency (RF) electromagnetic energy during medical procedures such as magnetic resonance imaging (MRI).

BACKGROUND

Magnetic resonance imaging (MRI) is a non-invasive imaging procedure that utilizes nuclear magnetic resonance techniques to render images within a patient's body. Typically, MRI systems employ the use of a magnetic coil having a static magnetic field strength of between about 0.2 to 3 Teslas. During the procedure, the body tissue is briefly exposed to RF pulses of electromagnetic energy in a plane perpendicular to the magnetic field. The resultant electromagnetic energy from these pulses can be used to image the body tissue by measuring the relaxation properties of the excited atomic nuclei in the tissue.

During imaging, the electromagnetic radiation produced by the MRI system may be picked up by implantable device leads used in implantable medical devices such as pacemakers or cardiac defibrillators. This energy may be transferred through the lead to the electrode in contact with the tissue, which may lead to elevated temperatures at the point of contact. The degree of tissue heating is typically related to factors such as the length of the lead, the conductivity or impedance of the lead, and the shape and surface area of the lead electrodes. In some cases, exposure to electromagnetic energy may also induce an undesired voltage on the lead.

SUMMARY

The present invention relates to implantable medical leads having impedance characteristics adapted to dissipate RF electromagnetic energy during medical procedures such as magnetic resonance imaging (MRI). An illustrative implantable medical device (IMD) configured for use in a magnetic resonance imaging environment includes a lead having an inner electrical conductor operatively coupled to an electrode, and at least one resistive shield that radially surrounds the inner, electrical conductor along all or a portion of the length of the lead. The inner electrical conductor can comprise a material having a relatively low resistance to facilitate energy transmission of IMD electrical signals through the conductor to the lead electrode. The inner conductor may have a relatively low impedance at the IMD such that it does not attenuate electrical energy transmitted by the IMD (e.g., electrical pulses transmitted by a pulse generator).

The outer resistive shield has a resistance that is relatively large in comparison to the resistance of the inner conductor, which dissipates RF electromagnetic energy received on the lead during an MRI scan along the length of the lead. In some embodiments, the outer resistive shield includes a layer or coating of resistive material radially disposed about at least a portion of the inner conductor. In other embodiments, the outer resistive shield includes a helically-shaped coil radially disposed about at least a portion of the inner conductor. The resistive shield can comprise a single resistive shield that extends continuously along the length of the lead, or can comprise multiple resistive shields each spaced apart from each other along the length of the lead via a gap, which serves to electrically isolate the resistive shields from each other. In use, the resistive shields minimize the energy pickup by the inner portion of the lead, and the high impedance of the shields at the frequency of the MRI RF energy minimizes the transfer of any energy picked up by the lead to the lead electrode.

A medical device in accordance with another illustrative embodiment includes a lead having an electrical conductor wire operatively coupled to an electrode. The conductor wire can vary in resistivity either continuously or at one or more locations across the width of the lead such that an outer portion of the conductor has a greater resistivity than the resistivity at a center portion of the conductor. In some embodiments, for example, the conductor wire includes an inner conductor core surrounded radially by one or more outer resistive shields configured to dissipate RF electromagnetic energy along the length of the lead. In other embodiments, the resistivity of the conductor wire varies continuously across its width between the center portion and the outer portion.

DETAILED DESCRIPTION

FIG. 1is a schematic view of an illustrative medical device12having a lead implanted within the body of a patient. In the illustrative embodiment depicted, the medical device12comprises a pulse generator implanted within the body. The medical device12can be coupled to a lead14inserted into the patient's heart16. The heart16includes a right atrium18, a right ventricle20, a left atrium22, and a left ventricle24. The pulse generator12can be implanted subcutaneously within the body, typically at a location such as in the patient's chest or abdomen, although other implantation locations are possible.

A proximal portion26of the lead14can be coupled to or formed integrally with the pulse generator12. A distal portion28of the lead14, in turn, can be implanted at a desired location within the heart16such as in the right ventricle20, as shown. Although the illustrative embodiment depicts only a single lead14inserted into the patient's heart16, in other embodiments, however, multiple leads can be utilized so as to electrically stimulate other areas of the heart.16. In some embodiments, for example, the distal portion of a second lead (not shown) may be implanted in the right atrium18. In addition, or in lieu, another lead may be implanted at the left side of the heart16(e.g., in the coronary veins) to stimulate the left side of the heart16. Other types of leads such as epicardial leads may also be utilized in addition to, or in lieu of, the lead14depicted inFIG. 1.

During operation, the lead14can be configured to convey electrical signals between the heart16and the pulse generator12. For example, in those embodiments where the pulse generator12is a pacemaker, the lead14can be utilized to deliver electrical therapeutic stimulus for pacing the heart16. In those embodiments where the pulse generator12is an implantable cardiac defibrillator, the lead14can be utilized to deliver electric shocks to the heart16in response to an event such as a heart attack. In some embodiments, the pulse generator12includes both pacing and defibrillation capabilities.

FIG. 2is a schematic view showing a simplified equivalence circuit30for the lead14representing the RF energy picked up on the lead14resulting from RF electromagnetic energy produced by an MRI scanner. As shown inFIG. 2, Vi34in the circuit30represents an equivalent source of energy picked up by the lead14from the MRI scanner.

During magnetic resonance imaging, the length of the lead14functions similar to an antenna, receiving the RF energy that is transmitted into the body from the MRI scanner. Voltage34inFIG. 2may represent, for example, the resultant voltage received by the lead14from the RF energy. The RF energy picked-up by the lead14may result, for example, from the rotating RF magnetic field produced by an MRI scanner, which generates an electric field in the plane perpendicular to the rotating magnetic field vector in conductive tissues. The tangential components of these electric fields along the length of the lead14couple to the lead14. The voltage Vi34is thus equal to the integration of the tangential electric field (i.e., the line integral of the electric field) along the length of the lead14.

The Zl parameter32in the circuit30represents the equivalent impedance exhibited by the lead14at the RF frequency of the MRI scanner. The impedance value Zl32may represent, for example, the inductance or the equivalent impedance resulting from the parallel inductance and the coil turn by turn capacitance exhibited by the lead14at an RF frequency of 64 MHz for a 1.5 Tesla MRI scanner, or at an RF frequency of 128 MHz for a 3 Tesla MRI scanner. The impedance Zl of the lead14is a complex quantity having a real part (i.e., resistance) and an imaginary part (i.e., reactance).

Zb38in the circuit30may represent the impedance of the body tissue at the point of lead contact. Zc36, in turn, may represent the capacitive coupling of the lead14to surrounding tissue along the length of the lead14, which may provide a path for the high frequency current (energy) to leak into the surrounding tissue at the RF frequency of the MRI scanner. Minimizing the absorbed energy (represented by source Vi34) reduces the energy that is transferred to the body tissue at the point of lead contact with the tissue.

The circuit represented inFIG. 2and the associated equation described below are for the purpose of illustrating the concept of lead heating in an MRI environment. At frequencies where the wavelength of induced voltage (or current) is close to the size of the circuit, a simple lumped sum system such as that illustrated inFIG. 2may not accurately model the behavior of the lead14in the MRI environment. Consequently, in those circumstances, a distributed model should be used along with Maxwell's equation for a proper mathematical description of the circuit. In some cases, the approximating distributed model can be created using field solvers or simulators.

As can be further seen inFIG. 2, the lead14has some amount of leakage36into the surrounding tissue at the RF frequency of the MRI scanner. As further indicated by38, there is also an impedance at the point of contact of the lead electrode to the surrounding body tissue within the heart16. The resulting voltage Vb delivered to the body tissue may be related by the following formula:
Vb=ViZbe/(Zbe+Zl), where Zbe=Zb in parallel with Zc.
The temperature at the tip of the lead14where contact is typically made to the surrounding tissue is related in part to the power dissipated at38(i.e., at “Zb”), which, in turn, is related to the square of Vb. To minimize temperature rises resulting from the power dissipated at38, it is thus desirable to minimize Vi (34) and Zc (38) while also maximizing the impedance of the lead Zl (32). In some embodiments, the impedance Zl (32) of the lead14can be increased at the RF frequency of the MRI scanner, which aids in reducing the power dissipated into the surrounding body tissue at the point of contact38.

In some embodiments, the impedance of the lead14can be increased by adding inductance to the lead14and/or by a suitable construction technique. For example, the inductance of the lead14can be increased by increasing the diameter of the conductor coil and/or by decreasing the pitch of the conductor coil. Decreasing the coil pitch may result in increasing capacitance between successive turns of the coil (i.e., coil turn by turn capacitance). The parallel combination of inductance (from the helical shape of the coil) and the turn by turn capacitance constitutes a resonance circuit. For a helically coiled lead construction, if the resonance frequency of the lead is above the RF frequency of the MRI, then the helical coil acts as an inductor. For an inductor, increasing the cross section of the coil area and/or reducing the coil pitch increases the inductance and, as a result, increases the impedance of the lead14.

Similar to an antenna, the energy pickup from a lead is related to its resonance length with respect to the wavelength of the frequency of interest. For example, for a dipole antenna, the antenna is considered tuned, or at resonance, when the antenna length is half the wavelength or an integer multiple of the wavelength. At resonance lengths, the energy pickup of the antenna is maximized. In a similar manner, and in some embodiments, the lead14can be detuned so as to prevent resonance within the lead14, and thus minimize the voltage Vi. For the illustrative embodiment shown inFIG. 1, for example, the lead14functions as an antenna having a resonance frequency at length L=integer×λ/2. In some embodiments, the length of the lead14and/or the construction parameters of the lead14affecting the wavelength can be chosen so as to avoid resonance within the lead14.

In some embodiments, in addition to detuning the length of the lead14with respect to the wavelength of the MRI induced RF energy, shielding can also be added to the lead14to further reduce the amount of electromagnetic energy picked-up from the lead14. For example, the energy picked up from the shielding can be coupled to the patient's body along the length of the lead14, preventing the energy from coupling to the lead tip. The transfer of intercepted energy by the shielding along the length of the shielding/lead can also be inhibited by dissipating the energy as resistive loss, using resistive material for the shielding construction.

FIG. 3is a longitudinal cross-sectional view showing a portion of the lead14ofFIG. 1in greater detail.FIG. 4, in turn, is a transverse cross-sectional view showing the lead14across line4-4inFIG. 3. As shown inFIGS. 3-4, the lead14includes an inner conductor core40and an outer resistive shield42that radially surrounds the inner conductor40along at least a portion of the length of the lead14. The inner conductor core40may function as an electrical conduit for supplying energy from the pulse generator12to one or more electrodes (not shown) located on the distal portion28of the lead14. In those embodiments where the lead14is a bradycardia lead, for example, the inner conductor core40may serve as an electrical conduit for supplying therapeutic energy to one or more electrodes used for pacing the patient's heart16and/or for sensing electrical activity occurring within the heart16. Alternatively, in those embodiments where the lead14is a tachycardia lead, the inner conductor core40may serve as an electrical conduit for supplying shocking energy to one or more electrodes coils located on the distal portion28of the lead14. Although a single inner conductor core40is shown in the embodiment ofFIGS. 3-4, in other embodiments multiple inner conductors may be provided for transmitting electrical energy to multiple pacing/sense electrodes located on the lead14.

In the illustrative embodiment ofFIGS. 3-4, the lead14has a substantially coaxial configuration with the inner conductor core40extending co-linearly with the outer resistive shield42along all or a portion of the length of the lead14. In other embodiments, the inner conductor core40can comprise a helically-shaped conductor coil (or multiple co-radially wound helically-shaped conductors) extending through the interior of the outer resistive shield42along all or a portion of the length of the lead14.

The outer resistive shield42can comprise a layer or coating of resistive material that radially surrounds the inner conductor core40. In one embodiment, for example, the outer resistive shield42comprises a resistive jacket that is formed integrally with the inner conductor40(e.g., as a single conductor) that radially surrounds the inner conductor core40. In some embodiments, the lead14may further include a layer or coating of insulative material about the resistive shield42to electrically isolate the conductor core40from the surrounding body tissue and blood. In further embodiments discussed herein in which there is a layer or coating of insulation between the conductor and the resistive shield, the layer or coating of insulative material disposed about the resistive shield42may be omitted, provided the length of the conductor is sufficiently small.

The outer resistive shield42may have a relatively high resistance in comparison to the inner conductor core40in order to facilitate dissipation of RF electromagnetic energy received along the length the lead14, which can cause heating of body tissue in contact with the lead14. In some embodiments, for example, the ratio of the resistance of the outer resistive shield42to the resistance of the inner conductor core40may be in the range of between about 2 to 10. An example resistance of the outer resistive shield42may be approximately 1 kΩ for a 50 cm long length of lead, although other values are possible.

In some embodiments, the difference in the resistance between the outer resistive shield42and the inner conductor core40can depend at least in part on the type of material(s) used, the dielectric constant of those materials, as well as other factors. In certain embodiments, for example, the inner conductor core40comprises a relative low resistance material configured to facilitate low-energy transmission of electrical signals along the core40whereas the outer resistive shield42comprises a relatively high resistance material configured to dissipate RF electromagnetic energy received on the lead14along the length of the lead14during magnetic resonance imaging. High resistance materials suitable for use as an outer resistive shield42can include, for example, metals, conductive polymers, and/or composite materials. In one exemplary embodiment, the inner conductor core40is a silver-filled MP35N wire containing approximately 28% to 30% silver whereas the outer resistive shield42comprises a layer or coating (e.g., a tubular jacket) of a different, more resistive material. An example conductive polymer suitable for use as the outer resistive shield42is polyphenylenevinylene or polyfluorene.

Although a higher conductivity shield generally provides better shielding of the inner conductor, and also permits thinner shielding to be used due to the lower skin depth, the higher conductivity of such shielding may also transfer the RF energy more easily along the length of the lead. Thus, even though the RF energy does not couple from the shield to the inner conductor, the shield itself may transfer the RF energy along the length of the lead towards the electrode at the lead tip. At or near the electrode, this high concentration of energy either directly heats the surrounding tissue (e.g., by capacitively coupling to the tissue), or couples the energy back to the lead where the inner conductor is exposed and contacts the tissue.

FIGS. 5 and 6are longitudinal and transverse cross-sectional views, respectively, showing the energy flow path of RF electromagnetic energy received on the lead14ofFIGS. 3-4in the presence of an MRI field. As shown inFIGS. 5 and 6, RF energy transmitted into the patient's body during magnetic resonance imaging is received along the length of the lead14, which acts as an antenna. Due to the “skin effect” property in conductive wires, in which alternating currents are limited to conduction at or near the surface of a conductor, the resistance increases towards the center of the lead. The increase of the resistance of the shield toward the center may also be enhanced by providing insulating material about the shielding in some embodiments. As energy travels along the length of the lead14, as indicated generally by reference arrows46, the magnitude of the energy is reduced (due to dissipation), and the dissipated energy is converted to heat along the length of the lead14. This can be seen generally by the reduction in the vector length of the arrows46from left to right inFIG. 5. As can be further seen inFIG. 6, the energy46takes the lowest impedance path of the resistive shield42instead of coupling to the inner conductor40due to the skin effect. Because the RF energy is dissipated along the length of the lead14and not at the electrode/body tissue interface, the heat absorbed by the lead from the RF energy is attenuated along the length of the lead14.

FIG. 7is a transverse cross-sectional view showing a lead48in accordance with another illustrative embodiment having a low-thermal impedance insulator radially disposed about the resistive shield. The lead48is similar to the lead14described with respect toFIGS. 3-4, including an inner conductor core50that can be used as an electrical conduit for supplying energy to one or more electrodes on the lead48, and an outer resistive shield52having a relatively high resistance compared to the resistance of the inner conductive core50for dissipating RF electromagnetic energy along the length of the lead48.

In the embodiment ofFIG. 7, the lead48further includes a low thermal impedance insulator54adapted to transfer heat generated along the length of the lead48at the lead/tissue interface. In some embodiments, the insulator54comprises a layer or coating of an insulative material radially disposed about the outer resistive shield52. In certain embodiments, for example, the insulator54is a thin layer or coating of silicone or polyurethane, although other configurations are possible. In other embodiments, the insulator54is a thin layer or coating of metal radially disposed about the outer resistive shield52. In some embodiments, the thickness of the metal insulator54is within the range of between about 10 μm to about 10,000 μm, although other configurations are possible.

In some embodiments, the material used for the insulator54has a relatively high dielectric constant in the range of between about 6 to 100. In some embodiments, the relatively high dielectric constant for the insulator material can be achieved by adding carbon particles, boron nitride particles, aluminum oxide particles, or the like to the insulator material. The insulator54can have a gradient of the dielectric constant that is constant along the length of the lead48, constant across the width (or transverse cross section) of the lead48, or a combination of both. During magnetic resonance imaging, the properties of the insulator54, including its thickness, dielectric constant, or a combination of both, can provide a means for transferring electromagnetic energy and/or the heat generated by the resistive shield52to the surrounding body tissue along the length of the lead48.

FIG. 8is a transverse cross-sectional view showing a lead56in accordance with another illustrative embodiment having a gradual change in resistivity across its width. In the embodiment ofFIG. 8, the lead56includes a single conductor58having a variable resistivity across its width. As indicated generally by reference arrow60, for example, the resistivity of the conductor58can vary across its width such that the resistivity is greatest at the outer surface62of the lead56where the lead56contacts the surrounding body tissue. In some embodiments, the resistivity of the lead56continuously increases from a center portion64of the conductor58to the outer surface62. In other embodiments, the resistivity may increase at one or more finite locations across the width of the lead56.

During an MRI scan, the increased resistivity towards the outer surface62of the lead56serves to dissipate the RF energy received from the MRI device at or near the outer surface62along the length of the lead56, thus minimizing the amount of energy transmitted into the interior of the lead56. This attenuation of the RF energy at or near the outer surface62prevents alternating currents from being transmitted through the interior conductor58to the electrodes located at the lead tip.

FIG. 9is a transverse cross-sectional view showing a lead66in accordance with another illustrative embodiment having multiple layers of conductors each with a different resistivity. As shown inFIG. 9, the lead66includes an inner conductor68surrounding radially by a number of outer, relatively high resistance conductors70,72,74. Each of the outer conductors70,72,74can have a different resistivity such that the resistance of the lead66varies across its width. In some embodiments, and as indicated generally by reference arrow76, the resistivity of each of the outer conductor layers70,72,74may successively increase across the width of the lead66such that the resistivity is greatest at or near the outer surface78of the lead66. In certain embodiments, for example, a first outer conductor layer70may have a first resistivity, a second outer conductor layer72may have a second resistivity greater than the first resistivity, and a third outer conductor layer74may have a third resistivity greater than the first and second resistivities. The number and arrangement of the conductor layers may differ, however. For example, while three outer conductor layers70,72,74are depicted inFIG. 9, in other embodiments a greater or lesser number of outer conductor layers each having a successively larger resistivity towards the outer surface78may be provided to dissipate RF energy received on the lead66during an MRI scan.

AlthoughFIGS. 8 and 9illustrate embodiments in which the lead resistivity varies either continuously or at one or more finite locations across the width of the lead56,66, other embodiments in which the impedance varies along all or a portion of the length of the lead56,66are also possible. In some embodiments, for example, the impedance of the lead56,66increases along the length of the lead56,66such that the distal portion of the lead has a greater impedance than at the proximal portion of the lead. The change in impedance along the length of the lead can be achieved by the selection of materials having a particular characteristic (e.g., a high dielectric, resistivity, etc.), by the construction of the lead (e.g., inductance), by the dimensions of the lead (e.g., surface area of the lead), as well as other factors. Creating impedance discontinuities along the length of the lead by changing the impedance of the lead along the lead length has an effect on the energy pickup of the lead during an MRI scan. In some embodiments, these discontinuities can be distributed along the length of the lead so as to prevent a standing wave from being generated along the length of the lead, thus minimizing tissue heating at the lead electrode.

FIG. 10is longitudinal cross-sectional view showing a lead80in accordance with another illustrative embodiment having an insulator between the resistive shield and the inner conductor.FIG. 11, in turn, is a transverse cross-sectional view showing the lead80across line11-11inFIG. 10. The lead80is similar to the lead14described with respect toFIGS. 3-4, including an inner conductor core82that can be used as an electrical conduit for supplying energy to one or more electrodes on the lead80, and an outer resistive shield84having a relatively high resistance compared to the resistance of the inner conductor core82for dissipating RF energy along the length of the lead80.

In the embodiment ofFIGS. 10-11, the lead80further includes a layer of insulation86disposed between the outer resistive shield84and the inner conductor core82. The layer of insulation86is configured to electrically isolate the inner conductor core82from RF energy received on the outer resistive shield84. An example of a layer of insulation86suitable for electrically isolating the inner conductor core80is a thin layer less than or equal to about 10 mils. If another insulation layer or coating is placed about the resistive shield84, then the layer of insulation86employed may be thinner, in some embodiments less than or equal to about 1 mil thickness. In certain embodiments, the outer diameter of the lead80, including the inner coil conductor82, the resistive shield84, and the insulation86is about 50 to 100 mils.

In various embodiments, the resistive shielded wire can be wound to make a coiled conductor, which adds further impedance to the lead by increasing the inductance. In the embodiment ofFIG. 3, for example, the conductor40and resistive shield42can be helically coiled, similar to the illustrative lead118discussed further herein, for example, with respect toFIG. 18, thus adding an inductance to the conductor40. In some embodiments, the resistive shielded wire used for coil construction may be relatively thin (e.g., less than 10 mils). In some embodiments, the coiled conductor can be inserted into insulation tubing or sheathing (e.g., 10 mil insulative tubing) to prevent the coil conductor from contacting the tissue and blood at the implantation site. If each of the wires forming the conductor has a layer of insulation, then the insulation tubing or sheathing placed about the coil conductor is generally very thin (e.g., not more than about 1 mil).

FIG. 12is a longitudinal cross-sectional view showing a lead88in accordance with another illustrative embodiment having a number of resistive shields spaced apart and electrically isolated from each other along the length of the lead. As shown inFIG. 12, the lead88is similar to that shown inFIG. 10, including an inner conductor core90that can be used as an electrical conduit for supplying energy to one or more electrodes on the lead88, and a layer of insulation92radially surrounding the inner conductor core90.

In the embodiment ofFIG. 12, the lead88further includes a number of outer resistive shields94a,94b,94cradially disposed about the layer of insulation92and the inner conductor core90. Each of the outer resistive shields94a,94b,94ccan comprise a layer or coating of material having a resistance that is relatively large in comparison to the resistance of the inner conductor core90. In some embodiments, for example, the ratio of the resistance of one of the outer resistive shields94a,94b,94cto the inner conductor core90can be in the range of between about 2 to 10.

The outer resistive shields94a,94b,94ceach extend along a portion of the length of the lead88, and are separated from each other via a number of small gaps G, as shown. The gap G between each longitudinally adjacent shield94a,94b,9ccan be sufficient such that each shield94a,94b,94cis electrically isolated from the other shields94a,94b,94c. In some embodiments, for example, the outer resistive shields94a,94b,94ccan be separated from each other by a gap G of approximately 4 mm to 5 mm. In other embodiments, the gap G separating each of the outer resistive shields94a,94b,94cmay be greater or lesser depending on the electrical characteristics of the shields94a,94b,94c(e.g., the material and thickness of the shields94a,94b,94c), the amount of RF energy received on the lead88, as well as other factors.

The length L of each of the resistive shields94a,94b,94ccan be selected to detune sections of the shields94a,94b,94cand prevent resonance based on the frequency of the RF energy provided by the MRI device. In some embodiments, for example, each of the outer resistive shields94a,94b,94chas a length L that is less than or equal to ¼ of the wavelength of the RF energy received on the lead88, thus detuning each of the shields94a,94b,94c. In use, the picked up energy can be evenly distributed along the length of the lead (and dissipated evenly in the resistive material) instead of concentrating near the ends of each shield94a,94b,94c, which, in turn, can capacitively coupled to an adjacent shield94,94b,94cand travel to the lead tip. The gaps G also serve to prevent the picked up energy from traveling to the end of the lead and dissipating at the location where the lead electrode contacts the surrounding body tissue, which can cause a temperature rise in the body tissue at this location.

FIG. 13is another longitudinal cross-sectional view of the lead88ofFIG. 12, showing the energy flow path of the RF electromagnetic energy received by the lead88in the presence of an MRI field. As shown inFIG. 13, RF energy transmitted into the patient's body during an MRI scan is received on each of the outer restive shields94a,94b,94c. Due to the relatively high resistance of the shields94a,94b,94c, and as indicated generally by reference arrows98, the RF energy is dissipated at or near the surface of the lead88along only the length L of each shield94a,94b,94c. The small gap G between each of the shields94a,94b,94cprevents the RF energy induced on one of the shields (e.g., shield94b) from being transmitted to an adjacent shield (e.g., shield94c). As a result, the flow of RF energy on one section of the lead88is interrupted and prevented from being transmitted along the entire length of the lead88to the lead tip. The interrupted RF energy is thus reflected back due to the mismatch in the impedance and is eventually dissipated within the resistive shield94a,94b,94c.

FIG. 14is a longitudinal cross-sectional view showing a lead100in accordance with another illustrative embodiment having a helically-shaped inner conductor coil surrounded radially by a resistive shield. As shown inFIG. 14, the lead100is similar to the lead80described with respect toFIGS. 10-11, with like elements labeled in like fashion in the drawings. In the embodiment ofFIG. 14, however, the inner conductor102is a helically-shaped conductor coil that extends through the interior of the lead100, and which adds inductance to the lead100at MRI RF frequencies, thus increasing the impedance and inhibiting transmission of the RF energy along the length of the lead100. In contrast to the outer resistive shield84, the conductor coil102is fabricated from an electrically conductive, low resistance material configured to facilitate low-energy transmission of therapeutic energy to the electrodes on the lead100. In some embodiments, for example, the inner conductor coil102is fabricated from a silver filled MP35N wire containing approximately 28% to 30% silver whereas the outer resistive member84comprises different, more resistive material.

In the embodiment ofFIG. 14, the outer resistive shield84may extend continuously and uninterrupted along all or a portion of the length of the lead100. In another illustrative lead104depicted inFIG. 15, the lead104includes a number of outer resistive shields106a,106b,106ceach spaced apart and electrically isolated from each other along the length of the lead104via a number of small gaps G, as shown. The length L of each of the outer resistive shields106a,106b,106ccan be selected to detune the lead104and prevent resonance based on the frequency of the RF energy provided by the MRI device. In some embodiments, for example, each of the outer resistive shields106a,106b,106chas a length L that is less than or equal to ¼ of the wavelength of the RF energy received on the lead104.

FIG. 16is a longitudinal cross-sectional view showing a lead106in accordance with another illustrative embodiment having a helically-shaped inner conductor coil surrounded radially by a helically-shaped resistive coil. The lead106is similar to the lead100described with respect toFIG. 14, including a helically-shaped inner conductor coil108that can be used as an electrical conduit for supplying energy to one or more electrodes on the lead106, a layer of insulation110radially surrounding the inner conductor coil108, and an outer resistive shield112having a relatively high resistance compared to the resistance of the inner conductor coil108for dissipating absorbed RF energy along the length of the lead106. In some embodiments, the inner conductor coil108is a single-filar wire coil. In other embodiments, the inner conductor coil108is a multi-filar wire coil.

In the embodiment ofFIG. 16, the outer resistive shield112comprises a helically-shaped resistive coil that radially surrounds the inner conductor coil108and the layer of insulation110. The resistive coil112has a relatively high resistance in comparison to the resistance of the inner conductor coil108to facilitate dissipation of RF energy received on the lead106into the surrounding body tissue along the length of the lead106. In some embodiments, for example, the ratio of the resistance of the resistive coil112to the resistance of the inner conductor coil108may be in the range of between about 2 to 10. In one embodiment, the inner conductor coil102is fabricated from a silver filled MP35N wire containing approximately 28% to 30% silver whereas the outer resistive member84comprises different, more resistive material such as a pure MP35N.

In the embodiment ofFIG. 16, the resistive coil112may extend continuously and uninterrupted along all or a portion of the length of the lead106. In another illustrative lead114depicted inFIG. 17, the lead114includes a number of resistive coils116a,116b,116ceach spaced apart and electrically isolated from each other along the length of the lead114via a number of small gaps G, as shown. The length L of each of the resistive coils116a,116b,116ccan be selected to detune the lead114and prevent resonance based on the frequency of the RF energy transmitted by the MRI device. In some embodiments, for example, each of the resistive coils116a,116b,116chas a length L that is less than or equal to ¼ of the wavelength of the RF energy received by the lead114.

FIG. 18is a longitudinal cross-sectional view showing a lead118in accordance with another illustrative embodiment. As shown inFIG. 18, the lead118includes a helically-shaped conductor coil120that extends along at least a portion of the length of the lead118. The conductor coil120is configured as an electrical conduit for supplying energy to the one or more electrodes on the lead118. In certain embodiments, the conductor coil120is encased within an outer member121which serves to radially constrain the conductor coil120along the length of the lead118. In some embodiments, for example, the outer member121comprises a jacket of silicone or polyurethane disposed radially about the conductor coil120.

FIG. 19is a transverse cross-sectional view showing the configuration of the conductor coil120ofFIG. 18in greater detail. As further shown inFIG. 19, the conductor coil120includes an inner conductive core122and an outer resistive shield124radially disposed about the inner core122. The resistive shield124can be formed integrally with the inner conductive core122(e.g., by a co-extrusion process), and comprises a material having a resistance that is greater than the resistance of the conductor core122. In some embodiments, for example, the ratio of the resistance of the resistive shield124to the resistance to the conductor core122can be in the range of between about 2 to 10.

In the embodiment ofFIGS. 18-19, the conductor coil120further includes a layer or coating of insulation126disposed between the inner conductor core122and the outer resistive shield124. In some embodiments, a second layer or coating of insulation128may also be provided over the outer, resistive shield124to further insulate the shield124.

If certain embodiments, and as further shown inFIG. 20, the first layer or coating of insulation126can be eliminated between the inner conductor core122and the outer resistive shield124. If, for example, the skin depth of the resistive shield124is sufficiently large (e.g., greater than 4), than the first layer or coating of insulation126may be eliminated since the alternating currents produced by the RF energy are unable to penetrate sufficiently through the depth of the shield124due to the decline in current density at the interface between the shield124and the inner conductor core122.

In those embodiments in which the conductor is coiled, the geometry of the coil conductor, including the outer diameter and pitch of the coil conductor, can be configured so as to increase the inductance and hence the impedance of the lead in order to inhibit the transfer of energy along the length of the lead. In some embodiments, for example, the inductance of the coil conductor can be increased by increasing the number of coil turns (e.g., by decreasing the pitch of the coil conductor), by increasing the outer diameter of the coil conductor, or by a combination of both. Since the impedance of a coil conductor is based in part on its inductance, increasing the inductance of the coil conductor by increasing the number of coil turns and/or increasing the outer diameter of the conductor results in an increase in the overall impedance of the conductor. Since the overall impedance of the lead at RF frequencies in MRI applications (e.g., 64 MHz) is partly a function of the inductance of the lead, this increase in the coil conductor inductance results in a decrease in the transfer of absorbed RF energy by the lead along its length towards the lead electrode.

In some embodiments, the coil conductor can comprise a helically-shaped wire coil conductor having a width of about 0.005 inches and a coil diameter (i.e., outer diameter) in the range of between about 0.016 inches to about 0.066 inches. In certain embodiments, for example, the coil conductor can have a coil diameter of at least 0.036 inches, 0.050 inches, 0.060 inches, or 0.066 inches. Other coil diameter configurations, however, are possible.

The pitch of the coil conductor can also be configured so as to increase the inductance and hence the impedance of the lead. In some embodiments the pitch of conductor can be in the range of between about 0.005 inches to 0.160 inches. Other coil pitch configurations, however, are possible. In general, as the coil pitch increases, the heating at the lead electrode also increases. For a 0.035 inch outer diameter coil, and in some embodiments, the pitch of the conductor should be no greater than about 0.008 inches, and more specifically, about 0.005 inches. For larger outer diameter coils, however, the minimum pitch can be larger, in some embodiments up to and including about 0.025 inches.

A strong interdependence exists between the coil pitch and the coil diameter of the coil conductor as the coil pitch increases and the coil diameter decreases.FIG. 21is a chart showing the interdependence of coil pitch and coil diameter on the amount of absorbed RF energy transferred by a helically-shaped wire coil conductor to a distal lead tip.FIG. 21may represent, for example, the peak specific absorption rate (SAR) as a function of coil pitch and coil diameter for a 0.036 inch coil diameter wire conductor and a 0.066 inch coil diameter wire conductor, each conductor having a length of about 50 cm and comprising an MP35N material. As the coil pitch of each of the conductors increases, the peak SAR, representing the amount of RF energy absorbed by the conductor, increases as a quadratic function. As shown inFIG. 21, a coil pitch greater than about 0.020 inches (20 mils) for a 0.036 inch wire coil, and a coil pitch greater than about 0.050 inches (50 mils) for a 0.066 inch wire coil, results in a significant rise in peak SAR that can increase the temperature of the body tissue at the lead electrode. As further shown inFIG. 21, at equivalent coil pitch values, the peak SAR for the larger coil diameter conductor is generally smaller than the peak SAR for the smaller coil diameter conductor.

Other design parameters of the lead can also be selected so as to reduce lead heating by the lead. In some embodiments, for example, the insulation provided about the coil conductor can be selected so as to reduce lead heating. The insulation thickness changes how much energy is coupled into or out of the surrounding body tissue along the length of the lead. In some cases, a relatively thin insulation, or insulations with higher dielectric constants, can minimize the temperature rise of a lead at the point of contact with the body tissue. An example of a relatively thin insulation for a coiled wire conductor has a wall thickness of less than about 0.015 inches, although other insulation thicknesses are possible.