JP6600632B2 - Application of electric field therapy to body parts - Google Patents

Description

This application claims the benefit of priority over US Provisional Patent Application No. 61 / 921,304, filed Dec. 27, 2013. This patent application is related to International Application No. PCT / US2012 / 55726 filed on September 17, 2012. It now claims the benefit of priority over the international application PCT / US2012 / 31923 filed on April 2, 2012, and is a continuation-in-part of that. This application is related to International Application PCT / US2012 / 55726 filed on September 17, 2012. This time, in contrast to US Provisional Patent Application No. 61 / 535,986, filed on September 17, 2011, and US Provisional Patent Application No. 61 / 584,399, filed January 9, 2012. Claim priority interests. This application is also related to US patent application Ser. No. 14 / 215,214, filed Mar. 17, 2014. This time, priority is given to US provisional application 61 / 803,775 filed March 20, 2013 and US provisional application 61 / 803,776 filed March 20, 2013. Insist on the interests of. The disclosure of each of the aforementioned patent applications is incorporated herein by reference in its entirety for all purposes.

FIELD OF DISCLOSURE The present disclosure relates to methods and systems for treating various tissue disorders. In particular, the present disclosure is directed, inter alia, to the treatment of tissue using strong electric field applications.

  For some time now, medical professionals have applied electric fields to living tissue in a variety of ways. The present disclosure improves the state of the art.

  Objects and advantages of the present disclosure are set forth in and will be apparent from the following description. Additional advantages of the disclosed embodiments may be realized and attained by the method and system particularly pointed out in the written description herein and also from the appended drawings.

  Of the known techniques for applying electric fields to tissue, the majority of these various types of electric field applications in the medical field can be classified as 1) thermal and 2) non-thermal.

  Applicants have recognized that most thermal electric field applications significantly increase the temperature of the target tissue, usually causing the tissue to die through the process of necrosis, ultimately resulting in ablation and cauterization. Applicants have observed that such processes affect not only the cells in the tissue but also the extracellular tissue matrix that contains various biological fibers to a very significant extent. They are also poorly distinguished between different cells and extracellular material. Many proteins in affected tissues denature through increasing temperature.

  The electric field that results in a temperature rise in the target tissue can be applied in a number of different ways, and various parameters may be different. Some have frequencies in the radio frequency and even microwave range (1 MHz to 300 MHz). Applications where an electric field in this frequency range is applied and its intensity and duration are sufficient to provide tissue heating for more than a few seconds above 40 degrees Celsius above normal body temperature is usually This is called RF or microwave ablation. RF and microwave thermotherapy are applications that increase the target tissue temperature by about 5-10 degrees for a few minutes. Powerful tissue heating and ablation have also been achieved through conducting relatively low frequency currents through the tissue by means of an electric field. In some cases (such as when the Bovie sword is used in surgery), an ionized gas called plasma can be generated between the target tissue and the electrode supplying the current. In thermal therapy, the penetration of the electric field and the heat generated thereby into the tissue is highly dependent on the details of the kinetics of the thermal diffusion and tissue transformation processes. As a result, the depth of the thermal effect is highly dependent not only on time but also on power density.

  Although it is perhaps more recently that non-thermal field effects have become more widely used in the medical field, non-thermal effects have been developed by using Volta and the electric field to generate neural tissue excitement and muscle contraction. Known since the time of Galvani. Clearly, today non-thermal applications of electric fields include nerve, muscle and brain stimulation, wound and bone healing stimulation, pain management, genetic modification and anti-cancer therapy. An important feature of non-thermal application of electric fields in the medical field is that different cell and tissue types can have different effects.

  For example, when the electric field is directed along nerve fibers, the magnitude is about 0.5-5 V / cm, and the pulse duration is several microseconds, the effect of exciting action potentials in nerve and muscle fibers is On the other hand, other cells and tissues are known to have no immediate effect. For stimulation of nerve and muscle tissue, in most cases, an exposed implantable conductive electrode (such as used with a Medtronic pacemaker device), conduction of current through the skin (percutaneous neural electrical stimulation or TENS) and transcranial magnetic stimulation ("TMS", eg, as described in US Pat. No. 4,672,951, which is incorporated herein by reference in its entirety for all purposes). Magnetic field pulses are used to induce an electric field in neural tissue as in some cases.

  More recently, an alternating electric field of about 1 V / cm with a frequency range of about 100-200 KHz and hours of exposure (about 24 hours) has been shown to prevent cell growth ( Eilon D. Kirson et al., PNAS, 104, 24, 10152-10157). This is used, for example, as a basis for the treatment of glioblastoma (a type of brain cancer) developed by Novocure. There, an electrode is placed on the skin of the head and a voltage within a few volts is applied at a frequency between 50 and 300 KHz to generate an electric field of about 1 V / cm within the target range in the brain. To do.

  Electroporation effects that are thought to generate pores in the cell plasma membrane that last from a few microseconds to a few milliseconds and result in about 200-1000 mV across the cell membrane (usually about 5 nanometers thick) It is known to bring Electroporation has been used for several decades in cell culture to transfer various molecules, including DNA, into cells. This is often aimed at transfection and subsequent genetic modification. It is known that the electroporation effect is complexly dependent on the magnitude of the electric field, the pulse repetition frequency and the total number of pulses. When the frequency, duration or number of pulses becomes sufficiently large, cells often die through cell self-destruction mechanisms.

Electroporation therapy leading to cell death is known as the irreversible electroporation effect (“IRE”). More recently, electroporation has been used in the medical field not only to modify cells in vitro, but also to treat tissues in human and animal bodies. Examples of such applications include treatment with a device called the Nanoknife ^™ device by Angiodynamics. There, an exposed solid conductive electrode is inserted into a target tissue, such as a tumor or an enlarged prostate, and a voltage pulse between the electrodes generates a conducting current that provides an electric field between the electrodes. In normal circumstances, the electroporation electrodes are separated by a gap of 1-5 cm and a voltage of 100-2000 volts is applied. Electroporation therapy has also been proposed and used to assist in the intake of chemotherapeutic agents in medical applications.

  Recent advances in electroporation include shortening of voltage pulse length as well as pulse rise and fall times. The faster the voltage application process, the greater the penetration of the electric field into the cells. This is because the current through the cell membrane does not have time to shield the applied electric field. However, applying a strong electric field with an uninsulated electrode as described above can cause local tissue necrosis and electrolysis (via heat), which can lead to excessive damage and scar formation near the electrode. So-called non-thermal irreversible electroporation (IRE or NTIRE) is possible with bare electrodes, but the field strength and pulse repetition frequency are severely limited [Davalos, 2012]. The practical limit is about 100 pulses with a repetition frequency of 1 Hz, an electric field strength of 1000 to 2000 V / cm, a pulse width of 50 to 100 microseconds. These treatment parameters can result in very long treatment times and / or limited treatment ranges.

Another example of non-thermal electric field therapy is applying radio frequency or microwave frequency electromagnetic field pulses to cells and tissues in the body. The intensity of the electric field in such treatment is similar to that employed to heat the tissue. However, the heating effect can be avoided by making the pulse shorter than a few milliseconds and limiting the pulse repetition frequency. Such applications of non-thermal electric field treatment are applied primarily in wound care and have been suggested to help wounds heal more quickly. As an example, such treatment using a Diapulse ^™ device has been reported. There, an electromagnetic field with a frequency range of 25-30 MHz is pulsed at a frequency of 10-800 pulses per second, and each pulse typically lasts less than 100 microseconds (eg, “Wound Care: A Collaborative Practical Manual for Health Professionals”). (See Carrie Sussman and Barbara Bates-Jensen, Kluwer Health, pages 580-583). It appears that the faster the pulse frequency, or the longer the pulse, leads to tissue heating. The electromagnetic field is applied to the tissue by coupling a pulsed electromagnetic field generator to a waveguide (which may be terminated by an antenna). Although contact between the tissue and the antenna is not required, it may be desirable to place the antenna in close proximity to the target tissue.

  In some cases, an electric field can be generated in the tissue without having to pass a conduction current from the electrode into the tissue. Typically, this is the case for electromagnetic therapy based on RF and microwaves. In some cases, an electric field in the tissue is induced using a pulsed magnetic field. In such cases, the magnitude of the electric field in the tissue is usually relatively weak and treatment often requires long-term (several hours) electric field exposure.

  In other cases, particularly when a relatively low frequency, long (greater than 1 second) pulse length, or large electric field is required, the current through the exposed conductive electrode is directed through the tissue to another electrode. When the required electric field and the required conduction current are relatively small, the electrochemical process occurs at a relatively low intensity in the vicinity of the exposed conductive electrode. A more powerful current that results in a stronger electric field (as often required for electroporation treatment with microsecond pulses) may result in a significantly stronger electrochemical process and heat generation near the electrode. Absent. This can cause significant undesirable tissue damage near the electrodes. This can be particularly problematic if one wants to treat tissue very locally without any substantial heating, or if one wants to control the heating effect separately. Some examples of medical applications where such effects are highly undesirable include treatment of brain tissue and in blood vessels, ventricles, intestinal cavities, esophagus, colorectal passages, bronchi, urinary passages, nasal passages and skin Or treatment of a thin tissue layer near the tissue wall, such as that found in the surroundings. A pulse of electric field lasting several microseconds can also excite nerves and muscles. This may be avoided by employing a nanosecond long pulse (especially when the electric field direction is perpendicular to the nerve fiber direction). However, there is still a problem with the current flowing between the electrodes in the tissue.

  The purpose of this disclosure is to avoid conduction currents or bubbles from being generated near the electrodes (which may occur due to electrochemical processes at the electrodes driven by conduction currents). It is to provide a system and method for applying a strong electric field in tissue without using a conducting current carrying electrode.

  In some embodiments, the present disclosure provides a system for treating a region of biological tissue. This includes an electrical signal generator configured to generate a time-varying electric field having a plurality of pulses. Each of the plurality of pulses lasts for a rise time and fall time of about 1-100 microseconds and a duration of about 2-200 microseconds. The system may further include at least one electrode disposed proximate to the target anatomical region and configured to apply a time-varying electric field to the target anatomical region. The electric field can have a peak magnitude between about 100 and 100,000 V / cm. The peak value of the electric field in the tissue in the target anatomical region can be controlled by the rate of voltage variation and by the peak voltage imposed by the electrical signal generator. The system may further include a current limiting barrier layer configured to be disposed between the at least one electrode and the target anatomical region.

  The present disclosure also provides an embodiment of a system for treating a region of biological tissue. This includes an electrical signal generator configured to generate a time-varying electric field having a plurality of pulses. Each of the plurality of pulses has a rise and fall time that lasts between about 1-1000 nanoseconds and a duration between about 2-2000 nanoseconds. The system may further include at least one electrode disposed proximate to the target anatomical region and configured to apply a time-varying electric field to the target anatomical region. The electric field can have a peak magnitude between about 100 and 100,000 V / cm. The peak value of the electric field in the tissue in the target anatomical region can be controlled by the rate of voltage variation and by the peak voltage imposed by the electrical signal generator. The system may further include a current limiting barrier layer configured to be disposed between the at least one electrode and the target anatomical region.

  Current limiting barriers include, among other things, (i) an insulating material having a thickness of between about 1 and 2000 micrometers, and (ii) an electric field across the gas filling gap exceeds a gas ionization threshold in the gas filling gap. A gas-filled gap that begins to conduct current therethrough, and (iii) an ionically conductive liquid or polymer. The at least one electrode may be configured to apply an electric field to the target anatomical region. This is oriented mainly in the direction perpendicular to or in contact with the electrodes in the target anatomical region. At least one electrode may be formed proximate to the surface of the inflatable member. This can be selectively inflated to bring at least one electrode proximate to the target anatomical region. The at least one electrode can be patterned. The patterning can be defined by, inter alia, areas with alternating exposed conductors and insulating materials, or areas with alternating insulating materials of different thicknesses. The at least one electrode may be flexible and / or elastic.

  In some embodiments, the at least one electrode can include a plurality of elongated conductors surrounding the central axis. This can be deployed radially outward from the central axis. The plurality of elongated conductors can include an insulating layer substantially along its entire length. The thickness of the insulating layer can vary linearly, non-linearly or in steps along the length of the plurality of elongated conductors. The thickness of the insulating layer may decrease along portions of the plurality of elongated conductors that are configured to be placed proximate to the target anatomical region. The plurality of elongate conductors may be partially insulated along portions of the plurality of elongate conductors configured to be disposed proximate to the target anatomical region.

  The present disclosure further provides a method that includes applying an electric field to a target anatomical region using any of the systems disclosed herein. The target anatomical region is between about one-tenth and one-hour, about five-second and about two-tenth, about thirty-second and about thirty-minute, about three-minute and about seven-minute. , Or any one second unit of time between 1 and 10,000 seconds. The target anatomical region can be on or in an animal or person.

  It should be understood that both the foregoing general description and the following detailed description are exemplary and are intended to provide further description of the embodiments disclosed herein. is there. The accompanying drawings are incorporated in and constitute a part of this specification, and are included to illustrate and provide a further understanding of the methods and systems of the present disclosure. Together with the description, the drawings serve to explain the principles of the disclosed embodiments.

1 is a schematic diagram illustrating an exemplary geometrically simplified model for calculating electric field generation in tissue. FIG. Schematic showing a typical frequency response of the relative permittivity and conductivity of skin tissue to an electric field. Schematic of typical voltage pulse shape. 1 is a schematic diagram of an exemplary patch electrode. FIG. 1 is a schematic diagram of an exemplary apparatus made in accordance with the present disclosure. FIG. 3 is a schematic diagram of a further exemplary apparatus made in accordance with the present disclosure. FIG. 3 is a schematic diagram of a further exemplary apparatus made in accordance with the present disclosure. FIG. 3 is a schematic diagram of a further exemplary apparatus made in accordance with the present disclosure. Further embodiments are illustrated according to the present disclosure. 6 illustrates further embodiments of electrode probes in accordance with the present disclosure. The long insulated electrode which has a concave treatment edge is illustrated. 10B illustrates a cross section of the embodiment of FIG. 10A. Furthermore, 10A illustrates a modification of the electrode of FIG. 10A. Here, the concave treatment end is on the side of the probe. 6 illustrates a further embodiment of an electrode head. 1 is a perspective view of a first embodiment of a magnetically assisted steering catheter system. FIG. FIG. 12B is a cross-sectional detail of the distal end of the system of FIG. 12A, showing a guidewire inserted through the system. FIG. 12B is a cross-section along line 3-3 of the system of FIG. 12B. 6 illustrates a further embodiment of an apparatus for applying an electric field to tissue. FIG. 6 is a front cross-sectional view of the tip of a further apparatus according to the present disclosure. Fig. 4 illustrates a further treatment device according to the present disclosure. Further exemplary treatment devices are further illustrated in accordance with the present disclosure. Further therapeutic devices are illustrated further according to the present disclosure. 6 illustrates a further embodiment of the present disclosure. Further examples of the present disclosure are further illustrated.

  Reference will now be made in detail to the presently preferred embodiments of the disclosure, examples of which are illustrated in the accompanying drawings. The methods and corresponding steps of the disclosed embodiments are described along with a detailed description of an exemplary system.

  A significant challenge in the field of medical use is the proximity of the electrodes while controlling the magnitude of the electric field pulse, rise and fall times, duration and time waveform, and the depth at which the electric field penetrates into the tissue. It is related to avoiding the thermal and electrochemical effects associated with the conduction current. Here, some of the ways in which a simplified model can be used to control these electric field parameters are described.

  In this simplified model illustrated in FIG. 1, a layer of tissue is placed between two conductive electrodes. By placing one or more separate layers of dielectric insulating material, resistive, semiconducting or other material between the tissue and the electrode, 1) the conduction of current through the tissue and associated with such conduction Heat or chemical processes may be prevented or limited, and 2) an irregularly shaped tissue surface and the surface of the electric field control device may be contacted appropriately. In addition to solid insulating materials, semiconducting materials that can control field penetration via thermal heating or photoexcitation may be used. Liquid, creamy or gel type insulating, conductive or semiconductive materials may be used that can be adapted to irregular tissue surfaces. Such liquids, creams or gels can be applied to the target treatment surface via suitable application means (manual application, spraying, infusion, irrigation, etc.). Furthermore, a gas or ionized gaseous material (plasma) may be used as a separation layer between the tissue and the conductive solid electrode. These materials can also be used in any combination, such as two, three, four.

The effect of the electrode shape on the electric field parameters is not considered in this simplified model. Rather, this will be described later. Instead, the electrodes, separating material, and tissue assume an interface that is planar and extend indefinitely within the boundary plane. This effectively allows the electric field perpendicular to the material interface to change only at the material interface. Let the combined thickness of the dielectric insulators be λ _d and the tissue thickness be λ _t .

The material in this model is described by parameters of relative permittivity k and conductivity σ. The relative dielectric constant is the ratio of the dielectric constant ε of the material to the dielectric constant ε ₀ ≈10-11 farads / m in vacuum. Conductivity is a proportionality factor between the effective conduction current density and the electric field involved in energy dissipation to heat.

In the following discussion, for the sake of brevity, it is assumed that the conductivity of the separation layer disposed between the conductive electrode and the tissue is zero. The frequency independent dielectric constant k _d of this dielectric material may vary in size from about 2 to 10 (or any 0.1 unit therebetween) depending on the actual material. In general, the conductivity of the separation layer contributes to energy dissipation to heat and a slower response to the applied voltage. Tissue electrical conductivity can vary significantly depending on cell density, cell type and water content. However, for most tissues, it remains below 1 S / m at frequencies below 1 GHz. The relative dielectric constant of the tissue may also vary significantly. In many cases, it is extremely large (10 ³ to 10 ⁷ ) at frequencies below the so-called α dispersion frequency (usually about 10 ⁴ to 10 ⁶ Hz), and about the β dispersion frequency (usually about 10 ⁷ to 10 ⁸ Hz). It decreases to 10 ² (relative dielectric constant of water), and then decreases to less than 10 ² at a frequency exceeding the γ dispersion frequency (about 10 ¹⁰ Hz).

As an example, if the skin is considered the target tissue, the relative permittivity k _t and conductivity σ _t are collected in FIG. 2 (collected from multiple sources and published in Phys. Med. Biol. 41 (1996) 2231-2249. (Data examined by C. Gabriel, S. Gabriel and E. Corthout), given as a function of frequency.

（Ｒは、パルスのなかでの振動周期の数）のような関数との積として表現できるからである。この数は、パルスのなかで１未満の振動周期が生じる場合、１未満であり得る。したがって、電圧パルスの簡略化モデルは、

であってもよい。 A typical time-dependent waveform of the voltage applied to the electrodes may have a shape as shown in FIG. This can be interpreted as a pulse. Because, normally, a vibration function such as a sinusoidal function sin (2πf ₀ t) (f ₀ is a vibration frequency) and a rectangular pulse

This is because it can be expressed as a product with a function such as (R is the number of vibration periods in a pulse). This number can be less than 1 if an oscillation period less than 1 occurs in the pulse. Therefore, a simplified model of the voltage pulse is

It may be.

と書くことができる。ただし、

In mathematical models of linear systems, it is often convenient to use a Fourier transformed variable called a phasor instead of a time-varying function. The Fourier transform of the above pulse is

Can be written. However,

The Fourier transform of a typical pulse is that in the phasor display, the center frequency of the pulse is the oscillation frequency f ₀ , and the bandwidth of the frequency associated with the pulse ranges from about Δf = 2π / τ = 2πf ₀ / R. Explain that. This justifies the use of a phasor with a frequency of about 1 / τ for pulses of approximate duration τ.

ただし、ｊ＝√（−１）、ｆは電界フェーザ周波数、ｋ_ｔ０＝１０^４は、α分散周波数ｆ_α＝１０^６Ｈｚ未満の周波数における組織の比誘電率、ｆ_β＝１０^８はβ分散周波数である。上記関係の漸近ボード線図を、図２における実際のデータの上に黒い実線で重ね合させて、この関係が、文献に見られるデータを、いかによく近似しているかを説明する。ｊ２πｆを乗じる代わりに時間微分を使用し、フェーザの代わりに時間依存量それ自体を使用することにより、フェーザの関係を実際の時間依存量の間の関係に変換してもよい。 In the various cases considered below, tissue conductivity is assumed to be constant at all frequencies of interest and have a value of about 1 S / m. On the other hand, the relative dielectric constant of the tissue takes a value consistent with the following relationship between the phasors of the electric field substitution “D tilde _t ” and the electric field “E tilde _t ”.

Where j = √ (−1), f is the electric field phasor frequency, k _t0 = 10 ⁴ is the relative dielectric constant of the tissue at a frequency less than α dispersion frequency f _α = 10 ⁶ Hz, and f _β = 10 ⁸ is β dispersion Is the frequency. The asymptotic Bode diagram of the above relationship is superimposed on the actual data in FIG. 2 with a solid black line to explain how well this relationship approximates the data found in the literature. Instead of multiplying j2πf, the phasor relationship may be transformed into a relationship between actual time-dependent quantities by using time derivatives instead of phasors and using time-dependent quantities themselves instead of phasors.

From the boundary conditions between the tissue and the dielectric insulator, the following relationship between the dielectric insulator and the phasor of the electric field in the tissue is obtained.

が得られる。ここで、

とする。 When “E tilde _d ” is found from (5), k _t (f) is found from (3), and is substituted into (4),

Is obtained. here,

And

Using the numbers typical for wet skin as described above and assuming an intermediate value k _d = 5 for the relative dielectric constant of the dielectric insulator,

と書ける。 From equation (6) it is clear that when the voltage V changes very slowly in time and the frequency f is close to zero, there is no electric field in the tissue. This is a basic characteristic of a system where conduction current cannot flow from the electrode into the tissue. If a conducting current can flow from the electrode through the tissue, the electric field can be maintained in the tissue for an extended period of time. However, equation (6) explains that as the frequency increases, an electric field can be applied into the tissue and maintained in it for some time. In this case, the characteristic time τ _d can be physically interpreted as the time it takes for the electric field energy to be transferred from the tissue into the dielectric insulator. The characteristic frequency f _d is the reciprocal of this time. The characteristic time τ _t can be viewed as the maximum time in the tissue where the electric field is dissipated to heat. _ft is the reciprocal of this electric field energy dissipation time. Using this characteristic frequency, equation (6) becomes

Can be written.

  Here, it is useful to consider several different modes of operation. This can be used to apply electric fields of varying magnitude and duration to the tissue using insulated electrodes.

Mode 1 : electric pulse of microsecond length, ratio of tissue to dielectric thickness is small to medium (λ _t / λ _d = 10 to 50)
Using voltage pulses with durations on the order of 1-10 microseconds, the frequency of operation is approximately f = 0.1-1 MHz. In this case,

となる。 The above equation explains that in the microsecond pulse mode of operation, the electric field in the tissue is proportional to the rate of change of the voltage applied between the electrodes, rather than the voltage itself. The tissue thickness is about 1 cm, the dielectric insulator thickness is about λ _d = 200 micrometers (λ _t / λ _d = 50), the magnitude of the voltage is about 10 kV, and the pulse length ranges from 1 to 10 microseconds Assuming

It becomes.

  This electric field can be further increased to about 45 V / cm by reducing the dielectric insulator thickness to about 100 micrometers and increasing the peak voltage to about 20 kV. Interestingly, according to equation (11), the electric field in the tissue is virtually independent of the dielectric constant of the tissue in this mode. This can be advantageous in some situations where the tissue to be electroporated has a very inhomogeneous dielectric constant but still wants to apply the same electric field to every corner.

Mode 2 : Nanosecond long electric field pulse, ratio of tissue to dielectric thickness is small to medium (λ _t / λ _d = 10 to 50)
Using voltage pulses with durations on the order of 10 nanoseconds, the frequency of operation is approximately f = 0.1 GHz. In this case,

となる。無次元量

は、重要なシステムパラメータと見ることができる。これは、ナノ秒パルス状電圧モードにおいて組織のなかで適用された電圧を電界に変換する有効性を制御する。この量は、通常、誘電絶縁体の誘電率だけでなく、組織最大誘電率、様々な臨界組織分散周波数の値、組織伝導率にも依存して変化する。よって、方程式（１５）が説明するところによれば、ナノ秒モードにおいて、組織のなかでの電界は、組織の電気的特性のすべてに敏感であり、組織伝導率のみに敏感なマイクロ秒動作モードとは対照的である。また、これが説明するところによれば、適切な電圧パルスパラメータを選択することにより、組織選択的電界浸透を増減することが可能である。 Substituting the above into (6), the second term on the right side of (6) is dominant, and as a result,

It becomes. Dimensionless quantity

Can be viewed as an important system parameter. This controls the effectiveness of converting the voltage applied in the tissue to an electric field in nanosecond pulsed voltage mode. This amount usually varies depending not only on the dielectric constant of the dielectric insulator, but also on the maximum tissue dielectric constant, various critical tissue dispersion frequency values, and tissue conductivity. Thus, equation (15) explains that in the nanosecond mode, the electric field in the tissue is sensitive to all of the electrical properties of the tissue and is sensitive only to tissue conductivity. In contrast to This also explains that tissue selective field penetration can be increased or decreased by selecting appropriate voltage pulse parameters.

となる。 According to equation (15) above, the electric field in the tissue in this mode of operation is almost independent of the rate of voltage increase (the increase and decrease in voltage ensure a 10 nanosecond pulse). As long as it is fast enough). Instead, the electric field actually increases with voltage. Physically, this is due to the fact that the conduction current through the tissue is negligible in this mode of operation. When the tissue thickness is about 1 cm, the dielectric insulator thickness is about λ _d = 200 micrometers (λ _t / λ _d = 50), and the magnitude of the voltage is about 10 kV,

It becomes.

Thus, using pulses of 10 nanoseconds or less can achieve peak electric field magnitudes on the order of electroporation fields, and the duration of the electric field pulses is about 10 ³ to 10 ⁶ times shorter than typical electroporation pulses. Such electric field pulses may not actually electroporate cells in the target tissue, but can be used to seriously affect cell behavior. Assuming that the rate of change of the electric field in the tissue corresponds to the magnitude of the electric field in (14) and that the rate of increase is about 5 nanoseconds (half the pulse length), the rate of change of the electric field is 5 × 10 5. ¹¹ V / (cm · s).

から推定できる。ただし、ｆ_ｐはパルス適用周波数である。ナノ秒モードにおいて、例えば、ピーク電界が約１×１０^３（Ｖ／ｃｍ）になるとき、時間平均出力堆積は、約１０^−１（Ｗ／ｃｍ^３）である一方、マイクロ秒モードにおいて、ピーク電界が約５０（Ｖ／ｃｍ）であり、パルスが同じ割合で適用されると、０．２５×１０^−１（Ｗ／ｃｍ^３）である。出力堆積率は、どちらも、同じオーダーの大きさである。このような出力堆積による組織温度上昇を推定するため、伝熱の問題を簡略化し、代謝活性による発熱と組織の血液灌流による除熱とを無視し得る。この仮定により、熱は、電界を施す領域から出るほうにのみ伝導し得る。簡略化のため、更に、約ａ＝１ｃｍ半径の球状組織体積の中心で電界が熱を発生している状況を考える。定常状態において、電界加熱された球状体積の周りの温度分布は、

となる。ただし、ξ≒０．５Ｗｍ^−１Ｋ^−１＝５０Ｗｃｍ^−１Ｋ^−１である。この定常状態分布における最大温度は、電界加熱された球状体積の表面で生じ、

となる。 When non-thermal treatment is desired, it is important to understand how the electric field contributes to tissue heating and what are the limits of pulse frequency, length and field strength. Time average power density deposition by electric field in the tissue is

Can be estimated from However, _{f p} is the pulse applied frequency. In nanosecond mode, for example, when the peak electric field is about 1 × 10 ³ (V / cm), the time average power deposition is about 10 ⁻¹ (W / cm ³ ), while in microsecond mode, the peak When the electric field is about 50 (V / cm) and the pulses are applied at the same rate, it is 0.25 × 10 ⁻¹ (W / cm ³ ). Both output deposition rates are of the same order of magnitude. In order to estimate the tissue temperature rise due to such power deposition, the problem of heat transfer can be simplified, and heat generation due to metabolic activity and heat removal due to blood perfusion of tissue can be ignored. With this assumption, heat can only conduct away from the area where the electric field is applied. For the sake of simplicity, further consider the situation where the electric field generates heat at the center of a spherical tissue volume with a radius of about a = 1 cm. In steady state, the temperature distribution around the spherical volume heated by electric field is

It becomes. However, ξ≈0.5 Wm ⁻¹ K ⁻¹ = 50 Wcm ⁻¹ K ⁻¹ . The maximum temperature in this steady state distribution occurs at the surface of the spherical volume heated by the electric field,

It becomes.

Obviously, if the time-averaged power deposition due to the electric field is on the order of P _E = 10 ⁻¹ (W / cm ³ ), the maximum temperature is only a few Kelvin higher than normal body temperature. This increase in temperature is further reduced by blood perfusion. Thus, nanosecond and microsecond modes of treatment can be considered non-thermal when the tissue volume is limited to a few cubic centimeters.

  The model simplifies the calculation of the electric field in the tissue by assuming that the electrodes are planar and spread infinitely. In more realistic situations, limit the depth of field penetration into the tissue or its spread along the tissue interface by selecting the appropriate size of electrodes and spacing the electrodes appropriately The shape of the electrode can be selected to do so.

  One advantage of generating an electric field in the tissue while limiting the conduction current between the electrodes is to use only a single electrode and avoid placing a second electrode on or in the tissue. It is to become. This is particularly advantageous when the electric field is applied near the skin surface or near the surface of various other tissues including lung, intestinal tissue, esophageal tissue, peritoneal tissue and other similar tissues. . An example of an insulating patch electrode is one in which a substantial electric field can be generated using a single electrode without perforating the surface in the tissue.

  For the purpose of illustrating the operation of the insulating patch electrode, consider one possible implementation of this embodiment. Here, a thin conductor, which can be made from a metallic material, is surrounded by a thin layer of dielectric insulator and placed on the surface of the skin. For the purpose of deriving an analytical approximation model, it is assumed that the thin conductor is in the form of a thin circular disk and the insulator surrounding it has the shape of a spheroid spheroid. This is illustrated in FIG. Here, the cross-section of this patch above the surface of the tissue has a thickness much greater than the size of the patch electrode (which is assumed to be infinite for the purpose of deriving an approximate analytical model) Assume.

となる。ただし、ａは円盤の半径であり、変数ηは各表面を記述する。円盤の表面は、η＝０に対応し、ηの値が増えると、それに対応して、扁球状回転楕円体表面が大きくなる。（２１）から示し得るところによれば、

かつ、ρ＝０であるｚ軸において

As can be shown, constant potential surfaces (equipotential surfaces) form a group of oblate spheroids (the dielectric insulator is also an oblate spheroid whose surface is the same group of equipotential surfaces). If it belongs to). If the origin is located at the center of the conductive disk electrode and ρ and z represent the distance from the axis perpendicular to the disk and the distance from the deduction point along that axis, respectively (see FIG. 4), A group of oblate spheroids that are equipotential is

It becomes. Where a is the radius of the disk and the variable η describes each surface. The surface of the disk corresponds to η = 0, and as the value of η increases, the oblate spheroid surface increases correspondingly. According to what can be shown from (21)

And in the z-axis where ρ = 0

の形式で表現し得る。ただし、「φチルダ」は時変電位のフェーザを示し、係数Ａ及びＢは境界条件を満足するよう選択される。この境界条件には、ａ）円盤電極上の電位を指定する値、ｂ）組織と誘電絶縁体との間の界面を横切る電位が連続していること、ｃ）組織と誘電絶縁体との間の界面に対して垂直な電界変位が、表面電荷密度の量により、不連続であること、及び、ｄ）電極円盤及び誘電絶縁体から遠く離れると電位が零になること、が含まれる。上記静電問題に対する最終解を見つけるため、解の体積を２つに分割して、誘電体内及び組織内とすると便利である。（２４）を用いると、誘電体内の電位のフェーザは、

と書ける。 As known (see, for example, WR Smythe “Static and Dynamic Electricity”, 3rd edition, McGraw-Hill, New York, 1968), the potential at the disc electrode is

Can be expressed in the form However, “φ tilde” indicates a phasor with a time-varying potential, and the coefficients A and B are selected so as to satisfy the boundary condition. The boundary conditions include: a) a value specifying the potential on the disc electrode, b) a potential across the interface between the tissue and the dielectric insulator, and c) between the tissue and the dielectric insulator. The electric field displacement perpendicular to the interface is discontinuous depending on the amount of surface charge density, and d) the potential becomes zero away from the electrode disk and the dielectric insulator. In order to find the final solution to the electrostatic problem, it is convenient to divide the volume of the solution into two, in the dielectric and in the tissue. Using (24), the phasor of the potential in the dielectric is

Can be written.

となり、この軸上における電界のフェーザは、（２６）から微分により、

となる。 Using (23), on the disk axis (z-axis)

The phasor of the electric field on this axis is derived from (26) by

It becomes.

と書ける。 Within the tissue, the potential phasor is

Can be written.

となり、この軸上における電界フェーザは、

となる。 Using (23), on the disk axis (z-axis)

The electric field phasor on this axis is

It becomes.

となる。 The electric field type form in the tissue automatically satisfies the boundary condition (potential becomes zero) far away from the electrode. Therefore, the remaining three coefficients are obtained from the boundary conditions. That _is, A d, _{B d} and _{A t.} When the dielectric thickness on the axis is λ _d , the boundary condition for potential continuity is

It becomes.

と書ける。 The boundary condition for the electric field displacement at the interface between the tissue and the dielectric is the same as condition (4),

Can be written.

となる。 Using the above,

It becomes.

である。 Finally, the boundary condition for the phasor potential “V tilde” specified on the disk conductor is:

It is.

となり、パッチの軸上における組織のなかの電界フェーザに対する式は、

である。 When all these boundary conditions are combined,

And the equation for the electric field phasor in the tissue on the axis of the patch is

It is.

であり、ｚ＝０．５ａにおいて、

である。 Equation (36) reveals that the electric field in the tissue is approximately constant if the axial distance from the electrode is a fraction of the electrode radius. For example, at distance z = 0.33a, the quantity

And when z = 0.5a,

It is.

これが示すところによれば、電界は、軸方向距離の二乗に反比例して素早く減衰する。軸方向距離が電極円盤半径の約３倍に等しいと、電界は、円盤から近接したところで見られる値の約１０％まで減衰する。これが示唆するに、その領域のなかの組織が電極径の数倍のオーダー厚さである場合、パッチ電極又は他の任意の絶縁表面電極は、電極径の約半分未満の距離においてほぼ均一な電界を組織のなかに提供する。 Thus, even though the axial distance is about half the electrode radius, the electric field in the tissue changes slowly. However, if z >> a,

This indicates that the electric field decays quickly in inverse proportion to the square of the axial distance. When the axial distance is equal to about three times the electrode disk radius, the electric field decays to about 10% of the value seen close to the disk. This suggests that if the tissue in that region is on the order of several times the electrode diameter, the patch electrode or any other insulating surface electrode will have a substantially uniform electric field at a distance of less than about half the electrode diameter. In the organization.

となり、（３９）は、

となる。 In order to discuss the magnitude of the electric field that can be obtained using a patch, it is worthwhile again to consider different electric field application modes. For all modes, the case of a thin dielectric λ _d / a << 1 is particularly useful. In this case, when z / a <0.5,

(39) becomes

It becomes.

となる。ただし、パラメータｆ_ｔ及びｆ_ｄは方程式（７）で定められ、ｆ_α及びｆ_βは組織モデルにおける低い方及び高い方の分散周波数であり、量ｇは皮膚モデルについて方程式（１６）で定められる。一例として、具体的なパラメータａ＝０．５ｃｍ、λ_ｄ＝５００μｍ＝５×１０^−２ｃｍを考えると、

となる。 Here, consider an operation mode that employs nanoseconds (on the order of 10 nanoseconds). In this mode, using the tissue model in (3),

It becomes. Where the parameters f _t and f _d are defined by equation (7), f _α and f _β are the lower and higher dispersion frequencies in the tissue model, and the quantity g is defined by equation (16) for the skin model. . As an example, considering specific parameters a = 0.5 cm and λ _d = 500 μm = 5 × 10 ⁻² cm,

It becomes.

  If the peak voltage is 10 kV, this will result in an electric field magnitude of about 1 kV / cm to a depth of about 0.25 cm under the patch in the tissue.

  This electrode can be applied to the exterior or interior of the body. This can be positioned and held in place by various positioning and holding means. Positioning means include manual, catheter and / or guidewire positioning and deploying devices, or endoscopes or bronchoscopes or other similar positioning and deployment means. Holding means include those by hand pressure, adhesives, rubber or Velcro strings, or other mechanical attachment means. The electrodes are made from a rigid, flexible or compatible material (such as a thin flexible sheet or mask filled with a conductive gel) for intimate contact with the target treatment area of the body. Optionally, a gel such as a hydrogel may be applied between the electrode and the target area to ensure that the applied electric field is properly distributed and minimizes the possibility of generating a surface plasma that can cause heat generation. .

  Further and other configurations of “patch-like” electrodes are described below. When the electric field described above is applied to tissue using the electrodes described below, target cells (such as tumors, carcinomas, lesions, cysts or other unwanted growth) are killed by cell self-destruction or necrosis and are either surrounding or below. Tissues can be kept intact. Furthermore, for example, in order to interfere with melanogenesis, an electric field can be applied to the skin for cosmetic treatment. Locally kills over-pigmented tissue or interferes with melanin status, ie, healing returns the skin to normal pigmentation, among other things, sunbeam (brown “liver” plaques), freckles Can treat hyperpigmented conditions such as, spots, melasma, pigmented nevus (nevus or black bruises), bears under the eyes, post-inflammation hyperpigmentation and black epidermoma. A specific goal of electrode design and treatment parameters is to minimize the generation of heat that can diffuse into the surrounding tissue and cause damage beyond the desired therapeutic range. Protection of surrounding tissue is particularly desirable in sensitive areas of the body such as blood vessels, ventricles and arteries, esophagus, colorectal passages, stomach, urethra, gynecological passages and tissues, nasal passages and bronchi.

  All electrode shapes described herein have a separation layer (eg, an insulating layer). This drastically limits the amount of conducting currents (which can lead to tissue necrosis) that lead to undesirable electrochemical reactions, creation of reactive species, electrolysis and heating. Such an electrode design can generate a much higher strength electric field. This allows the targeted cells to be killed faster and over a wider area without harming the walls of the delicate tubes and cavities.

  Various needle-based electrodes or multi-needle arrays have also been proposed in the prior art. These configurations are designed to be inserted into a target tissue (such as a tumor) to be treated. Such needle-based techniques are not well suited to killing target cells near a delicate, thin-walled inner tube or body cavity. This is because the risk of perforating these structures is high. If the target range is multiple or large, the treatment time can be delayed. This is because needle-based systems must be carefully repositioned and reinserted into each target. Finally, such an electrode can generate a very non-uniform and concentrated electric field. This results in uneven treatment with minimal or local “sparks” and, in the worst case, damage.

The following criteria help to select the optimal electrode shape.
1. The distance between the electrode and the target treatment area is clearly defined and repeatable.
2. Adaptable to changes in size, curvature and local distribution of target treatment area (curvature of tube or chamber, diameter and length, etc.).
3. Maximize the range of treatment while meeting the goals of (1) and (2).
4). Easy to manufacture / assemble.
5. Minimize the voltage required to generate a sufficient electric field and / or plasma for therapeutic purposes. Lowering the operating voltage can reduce the cost and size of the power supply and associated electronics.
6). Minimize the occurrence of electric field inhomogeneities. This can result in the generation of a plasma spark (a “flow” of high temperature plasma such as lightning) where electrical energy is concentrated within one or more highly localized areas.
7). Electrodes can be scanned when the entire target area cannot be treated while still.

  One embodiment of an electrode that meets these requirements is an insulated conductor needle, rod or wire. It has a pointed or rounded effector with a slightly thinner dielectric coating than along the needle / rod / wire length. In the case of an insulated needle, the needle is inserted into the center of the tumor or other target tissue to be treated. Prior to activation, the generator settings are adjusted to match the magnitude of the applied electric field with the size of the target tissue. In the case of a rod, it is inserted into the desired body structure (cavity or tube) and a round end effector is placed in contact with the wall of the cavity or tube.

  When the high voltage waveform generator is activated, the electric field is concentrated in a circular area around the contact area. U.S. Pat. Nos. 8,583,260, 7,850,685, 8,394,091, 8,419,681, 6,585,718, 5,480,382 and Steering and deflection mechanisms as well as electrode configurations as shown in US Pat. No. 5,318,525 can be used with such application catheters. Further embodiments of interest are US patent application Ser. No. 10 / 426,923 filed Apr. 29, 2003 and US Patent Application No. 10 / 754,445 filed Jan. 9, 2004. Is disclosed. All patents and publications referred to herein are hereby incorporated by reference in their entirety for all purposes. With these mechanisms, the treatment electrode can be inserted into and brought into contact with the body toward the target area. In the above-mentioned patent, the bare electrode is connected to the radio frequency generator, whereas the electrode in this embodiment is insulated and connected to the high voltage waveform generator as described above. It should further be noted that the wiring in the application device must be sufficiently insulated to prevent dielectric breakdown along the application device. The amount of insulation required depends on the voltage applied. Such a system still relies on the operator to ensure that the effector maintains contact with the body vessel or cavity during treatment.

  Alternatively, the catheter is pre-shaped and adapted to the target physiology, and the catheter is treated (such as the heart wall as shown in US Pat. Nos. 8,535,303 and 8,486,062). The treatment electrode may automatically be placed in close contact with the target area. The above mentioned patents are incorporated herein for all purposes.

  In another embodiment shown in FIG. 5, the above-described conductive rod may include a plurality of insulated electrical structures (wires or wire tracking) extending axially along the effector (rounded or flat at the ends). Have In particular, FIG. 5 shows a catheter having a steerable electrode tip (1). When the catheter (2) is advanced to the desired position, the tip (1) is extended by holding the catheter hub (3) and moving the stylet hub (4) distally to extend the tip (1). The tip can then rotate as desired to the target range. The stylet is connected to a waveform generator (not shown) as described herein.

  The wire tracking pairs are alternately electrically connected to the high voltage and ground from the pulse generator. As described above, the rod is inserted into the desired body structure (cavity or tube) and the effector is placed in contact with the target area. U.S. Pat. Nos. 8,583,260, 7,850,685, 8,394,091, 8,419,681, 6,585,718, 5,480,382 and Steering and deflection mechanisms as well as electrode configurations as shown in US Pat. No. 5,318,525 can be used with such application catheters. With these mechanisms, the treatment electrode can be inserted into and brought into contact with the body toward the target area. The above mentioned patents are incorporated herein for all purposes. In the above-mentioned patent, the bare electrode is connected to the radio frequency generator, whereas the electrode in this embodiment is insulated and connected to the high voltage waveform generator as described above. It should further be noted that the wiring in the application device must be sufficiently insulated to prevent dielectric breakdown along the application device. The amount of insulation required depends on the voltage applied. Alternatively, the catheter may be preformed to match the target physiology and the treatment electrode may be automatically placed in close contact with the target area when the catheter is inserted into the treatment site. Actuation of the pulse generator can apply an electric field to the portion of the target area that is in contact with the wire tracking.

  In another embodiment, the helical electrode is screwed into a target tissue site (such as a tube or cavity) to be treated. Spiral electrodes, such as those shown in FIGS. 1-6 of US Pat. No. 5,431,649 and accompanying specifications, are configured to provide a larger surface area for treatment. Preferably, as described above, it is coated with a separating material and connected to a high voltage waveform generator. The above mentioned patents are incorporated herein for all purposes.

  In another embodiment, shown in FIG. 6 herein, a balloon made from a separating material is placed near the target area by hand or via a delivery catheter and / or guidewire. The balloon is then inflated with a conductive material electrically connected to the pulse generator. When the high voltage waveform generator is activated, an electric field is applied to the portion of the target area that touches the side of the balloon or is likely to touch the balloon. The conductive material can be an electric plasma, a liquid, a particle, or a composite thereof.

  As illustrated in FIG. 6, the balloon (1) is introduced to a desired position by the guide wire (2). The catheter (3) provides a conduit for introducing the conductive fluid (4) into the balloon prior to actuating the waveform generator (not shown). This balloon-based treatment can be particularly useful as an alternative to post-operative therapy after tumor removal. After tumor removal, radiation or chemotherapy is often used to ensure that any residual tumor within that range is killed. Placing a balloon inflated with a conductive material near its surface and treating it with electric fields may avoid such radiation or chemotherapy associated with dangers and side effects. If desired, balloons can be implanted and filled after extraction and operated wirelessly as needed.

  In order to make intimate contact between the electrode and the irregularly shaped local distribution of the tissue surface (such as the surface that can be seen after the excision surgery described above), this local distribution can be adapted between the electrode and the tissue. It is useful to provide a separating material. Examples of such isolation materials may include liquids, gels and gases that act as dielectric insulators or electrical resistance materials. This can include, but is not limited to, oils, conductive fluids, fluids containing conductive materials or particles, gels and plasmas. These materials can also be used in combination with a solid separation layer.

  In another embodiment, shown in FIG. 7 herein, the balloon is coated with a conductive layer and then with a separating layer. The balloon is connected to a pulse generator, inflated and activated as described above. It should further be noted that the wiring in the application device must be sufficiently insulated to prevent dielectric breakdown along the application device. The amount of insulation required depends on the voltage applied.

  In another embodiment, a balloon having an insulated electrical structure independent of each other is shown in FIG. 8, FIGS. 1A and 1B of US Patent Application No. 2013/0184702, and FIG. 11 of US Pat. No. 850,685. In proximity to the target area, via a delivery catheter and / or guidewire. The aforementioned patents and applications are incorporated herein for all purposes. One variation of a balloon having an insulated electrical structure includes a plurality of insulated wires that run at defined intervals along the outside of the balloon. The depth at which the electric field penetrates can be set based on the interval between the wirings. Other structural configurations can also be designed by those skilled in the art. The wires are connected to a high voltage waveform generator, and alternately the wires are connected to a high voltage output and ground. The balloon is then inflated with gas, and the insulated electrical contact contacts the target and is electrically connected to the pulse generator. As in the previous embodiment, when the pulse generator is activated, an electric field is applied to the target area.

  In particular, FIG. 8 shows a balloon treatment catheter. The guide wire (1) provides internal support when inserting the catheter (2) into place. The balloon (3) has an insulated wire (4). This surrounds the balloon and is joined at the connector (5) to a wire that is connected back to the generator (not shown). . The catheter is retracted from the initial position (6) to deploy the balloon.

  Alternatively, round or semi-round wires, multiple wires, insulated conductive sheets, or insulated sheets with a wire tracking pattern are each bent into a wound shape and into the body via an application catheter Insert to target range. When the target area is reached, the rolled wire is deployed from the application catheter as shown in FIGS. 12 and 14 of US Pat. No. 7,850,685. Such wires are made to match the diameter of the target tube or cavity, or to be slightly larger, so that the wrap is compressed and preloaded against the target area. When the pulse generator is activated, an electric field is applied to the target area via the wound wire.

  Similarly, one or more wrapping wires can be provided and inserted into the body to the target area. As shown in FIG. 6 of US Pat. No. 7,850,685, one wire or set of wires is connected to the high voltage output and the other is connected to ground. In all cases, the electrode described in US Pat. No. 7,850,685 is modified and coated with a desired separation layer and connected to a suitable high voltage waveform generator as described herein above. . It should also be noted that the wiring in the application device must be sufficiently insulated to prevent dielectric breakdown along the application device (not in the target range). The amount of insulation required depends on the voltage applied.

  The electrode configuration described above can also be delivered into the body via an endoscopic, arthroscopic or laparoscopic means, for example as shown in FIGS. 1-6 of US 2013/0261389. The above-mentioned patent applications are incorporated herein for all purposes. It should be noted that the electrodes in the above mentioned application are modified to include a separation layer. Optionally, one or more of these wires can be used (both monopolar and bipolar operation is possible) and connected to the desired high voltage waveform generator.

  All application devices may provide additional capabilities. This can assist in identifying target treatment areas, targeting and diagnosis (such as imaging and electrical detection). Such an application device may provide the ability to apply liquids, creams and / or gels. This becomes a separation layer and assists in adaptively contacting the local distribution of the target tissue that is bent or otherwise irregular. Such application means include injectors, needles, hoses, nozzles and the like. This may be driven via a syringe, pump, pressurized gas cartridge or other fluid motion guidance system.

  Control of the treatment procedure was originally filed on April 2, 2012 as International Application No. PCT / US2012 / 031923 and filed on September 17, 2012 as International Application No. PCT / US2012 / 055726. This is accomplished through a variety of methods as described in US patent application Ser. No. 13 / 943,012, filed on Jan. 16. Each of the aforementioned applications is incorporated herein by reference for all purposes.

  Remotely conductive electrodes can be employed because they generate relatively strong electric field pulses without the need to transfer current from the electrodes into the tissue. It is not connected to a voltage or current pulse source by a current carrying conductor to generate a sufficiently strong pulse of the electric field in the target tissue. Some examples of using remote electrodes are shown in FIG. 9A-C show that when an external electric field is applied by shaping and placing a remote electrode on or in tissue and connecting a generator of pulsed voltage to other control electrodes, 6 illustrates that it can be focused or strengthened around a tissue region that has been made. In some cases, as illustrated in FIG. 9, the remote electrode may be connected to the control electrode by a brief generation of plasma that transfers charge to the remote electrode. The remote electrode may be used without insulation in some cases. This is because it is not directly electrically connected to an external generator capable of supplying a conduction current.

  10A-10C illustrate further embodiments of electrode probes according to the present disclosure. FIG. 10A illustrates a long insulated electrode having a concave treatment end. FIG. 10B illustrates a cross section of the embodiment of FIG. 10A. Further, FIG. 10C illustrates a variation of the electrode of FIG. 10A. Here, the concave treatment edge is on the side of the probe. As illustrated, each probe has a proximal end attached to a high voltage connector, a distal end having a treatment end, and an outer circumferential generally cylindrical surface. This is covered with a layer of electrically insulating material in the form of a dielectric coating. A concave treatment end is provided at the distal end of the probe of FIG. 10A. It has a dielectric coating. This may be thinner (or otherwise less resistive) than the dielectric coating covering the probe. If desired, the concave probe may be placed on the side of the probe as shown in FIG. 10C. If desired, this probe may be used to generate a local plasma discharge in the concave surface and apply an electric field to the tissue. Alternatively, it may be configured to generate substantially no plasma and simply apply an electric field to the tissue. The recesses in the electrode can be patterned to create regions of varying electrical resistance across the dielectric (such as changing the thickness of the dielectric to form a grid pattern ridge).

  11A-11J illustrate various examples of embodiments of insulator / electrode arrangements for applying electric field pulses into tissue. The brighter color in this figure corresponds to an insulating material such as silicone, PEEK, PVDF. The black color in this figure corresponds to a solid conductive material such as copper. Where indicated, a fluid or gel material ("G") may be provided. This can be a conductive material that can flow or otherwise change shape three-dimensionally.

  FIG. 11A illustrates an insulated multiconductor electrode. Due to air breakdown in the form of electronic avalanches and ionization, the gap between the insulated surface of the electrode and the surface of the tissue may be electrically shorted, as shown in FIGS. This air breakdown typically occurs during a sub-interval that is shorter than the voltage pulse. Whether the ionized gas fills the gap between the surface of the insulating electrode and the tissue may depend on the rise and fall times of the voltage. In some cases, typically when the rise time is longer than a few nanoseconds, the ionized gas tends to occur more likely closer to the ridge. Different conductors may be electrically connected to the same or different potentials. When connected to the same potential source, the electric field is mainly perpendicular to the electrode surface over the entire electrode surface, as shown in FIG. 11G. When the electrodes are connected to different potential sources (high potential and ground, respectively), the electric field is mainly perpendicular to the electrode surface where the conductor spreads, as illustrated in FIG. It is mainly parallel to the electrode surface.

  FIG. 11B illustrates an insulated electrode having a continuous conductor indicated by a dark arcuate line. When a high voltage pulse is connected to the conductor, the electric field is generated mainly perpendicular to the electrode surface over almost the entire electrode surface.

  The pattern of ionized gas formation under the electrode may also depend on the relative position of the conductor with respect to the bumps or protrusions on the insulator surface. For example, as shown in FIG. 11H, the plasma breakdown region may be confined near the surface of the ridge or protrusion when the conductor is placed over the ridge or protrusion. FIG. 11D illustrates that all conductors may not be insulated to avoid conducting currents passing through the tissue. FIG. 11I illustrates how this type of electrode may be used when applying a pulse of electric field into the tissue.

  FIG. 11C shows that the conductive material surrounded by the insulator need not necessarily be a solid conductor. Solid conductive materials such as aluminum or copper metal are highly conductive, but may suffer from undesirable mechanical failures (eg, cracks) when the electrode is pulled. If this is a problem, it may be convenient to replace the solid conductor with a conductive gel, fluid or clay. One advantage of using a fluid is the possibility of filling different flow paths in the insulator, thereby providing the possibility of having a reconfigurable structure of conductive material in the insulated electrode. . The possibility arises of using a material with a relatively low conductivity to form the conductive structure in the insulated electrode. The application discussed here is because the electric field needs to be strong during the pulse duration, but the output may be relatively low. In some cases, there is no current through the electrode conductor or the current is very low, thereby avoiding concerns about electrode heating and conductor voltage drop typical of RF ablation applications. Using a conductive gel electrode, an example that helps to adapt the electrode to the tissue is shown in FIG. 11J. It should be noted that the electric field distribution in materials with relatively low conductivity can change more significantly as the amount of material in the stretchable electrode increases. In such a case, a highly conductive material in the form of a wire or coil or other shape, such as a metal or other method, is inserted into a material with low conductivity (gel, fluid, etc.) and the entire stretchable electrode is Some variations in the electric field may be minimized.

  FIG. 12A is a perspective view of an embodiment of a magnetically assisted steering catheter system. This is configured to apply a time-varying electric field to the tissue to be treated located in the anatomical target area. FIG. 12B is a cross-sectional detail of the distal end of the system of FIG. 12A showing a guidewire inserted through the system. FIG. 12C is a cross-section along line 3-3 of the system of FIG. 12B. Referring to FIG. 12A, a magnetically assisted steering catheter system 10 is shown. System 10 typically includes a catheter 20. This is through the patient's vasculature into a three-dimensional anatomical cavity (such as the heart) or through a naturally or surgically created opening into the gastrointestinal tract, thoracic cavity, abdominal cavity, etc. Configured to be introduced, where an electric field can be applied to the tissue and used to survey the tissue if desired. As illustrated, the system may further include an electrophysiology survey processor 14 that is used to electrophysiologically measure tissue with the catheter 20. The system further includes an electrical energy source as described herein. This applies a time-varying voltage waveform to the catheter 20 to impose an electric field on the desired target location. A guide wire 30 is also provided to facilitate introduction of the catheter 20 through the patient's tortuous vasculature.

  Still referring to FIGS. 12B and 12C, the guidewire 30 is configured to be slidably disposed within the catheter 20. Guidewire 30 includes an elongate flexible body 31 having a proximal end 32 and a distal end 34. The guide wire body 31 is made of a flexible and elastic material, for example, a superelastic alloy such as Nitinol, or alternatively, a material such as titanium, tantalum, or stainless steel. The outer diameter of the guide wire body 31 is about 0.020 inch (0.51 mm). The length of the guide wire body 31 is sufficient to pass through the entire length of the catheter 20. In the exemplary embodiment, the length of the guidewire body 31 is about 180 to 220 cm.

  If desired, the guidewire 30 may further comprise a magnetic attraction element 36 disposed at the distal end 34 of the guidewire 30. The magnetic attraction element 36 may be secured to the distal end 34, such as by welding, brazing, gluing or other suitable adhesive. The magnetic attraction element 36 may take the form of an element that moves in response to a magnetic field. For example, the magnetic attraction element 36 may include a permanent magnet material (such as neodymium-iron-boron). Alternatively, it may include a steel material (such as cold rolled steel or iron-cobalt alloy). The magnetic attraction element 36 may also take the form of an electromagnet connected to a wire (not shown) that passes through a lumen (not shown) that extends across the guidewire 30 in a conventional manner.

  Catheter 20 typically includes an elongate flexible catheter body 21 having a proximal end 22 and a distal end 24 and a guidewire lumen 35 extending the entire length of the catheter body 21. Guidewire lumen 35 is sized to receive guidewire 30. In the illustrated embodiment, the catheter body 21 comprises a unibody design. That is, it is formed by a single extrusion according to the method used by those skilled in the art. Alternatively, a two-part catheter body (not shown), such as a proximal member (not shown) and a distal member (not shown), can be bonded at the interface (not shown) using overlapping thermal bonding, or a sleeve The end may be bonded to the end by what is called “butt bonding” throughout.

  In the illustrated embodiment, the catheter body 21 has a diameter of about 5 to 9 French and a length of about 80 to 115 cm. The catheter body 21 is made of a biocompatible thermoplastic material (such as a Pebax (registered trademark) material (polyether block amide) and a stainless steel braided composite). This has excellent torque transmission characteristics. In some embodiments, an elongate guide coil (not shown) may be provided in the catheter body 21. The distal end 24 of the catheter body 21 preferably includes a radiopaque compound (such as barium) so that the catheter 20 can be viewed using fluoroscopy or ultrasound imaging or the like. Alternatively, a radiopaque marker (not shown) can be placed along the distal end 24 of the catheter body 21.

  The catheter body 21 has an elastic structure. This facilitates the functionality of the magnetically assisted steering catheter system 10. For this purpose, the catheter 20 includes a spring member 29 disposed within the catheter body 21 and passing through its entire length. The shape of the catheter body 21 is achieved by using a spring member 29. The torque transferability of the catheter body 21 is important for predictable and controlled movement of the distal end 24, and to improve it, the spring member 29 in the illustrated embodiment is formed from a unibody structure. The torsional force that is fixed and applied to the proximal end 22 is transmitted to the distal end 24 of the catheter body 21 without significant loss.

  The catheter 20 further includes an electromagnet 26 carried at the distal end 24 of the catheter body 21, one or more actuating elements, and in particular an electrode 25 surrounded by an insulating sleeve or layer “I”, if desired. A surveying element 27 carried on the distal end 24 of the catheter body 21 and a manipulation assembly 40 mounted on the proximal end 22 of the catheter body 21 may be included. The electromagnet 26 may take the form of a wound solenoid or an induction electromagnet. This is secured to the distal end 24 using any suitable means (such as welding, brazing, gluing or other suitable adhesive) depending on the material from which the electromagnet 26 and catheter body 21 are made. The catheter 20 further includes two conductors 23a and 23b that extend from the electromagnet 26 toward the proximal end 22 of the catheter body 21 throughout the catheter body 21 and for coupling to an electrical energy source.

  Current is supplied to the electromagnet 26 via the conductor 23a and returned by the return wire 23b to induce a first magnetic field. The polarity passing through the electromagnet 26 is reversed to induce the second magnetic field in the opposite direction from the first magnetic field. The electromagnet 26 is configured to generate a magnetic field sufficient to distort (shown as a phantom in FIGS. 12A-12C) or otherwise manipulate the magnetic attraction element and / or tip 36 at the distal end 34 of the guidewire 30. Has been. The electromagnet 26 may be configured to generate a higher magnetic field due to the heavier guidewire and / or the guidewire with the distal end 34 farther away from the electromagnet 26.

  The strength of the magnetic field generated by the electromagnet 26 depends on many factors (such as the shape of the electromagnet 26, the material from which the electromagnet 26 is made, and / or the amount of output supplied to the electromagnet 26). In the illustrated embodiment, the power applied to the electromagnet 26 is fixed at a set level and, when activated, distorts the guidewire tip 36 within a predetermined distance from the electromagnet 26 (as a phantom in FIGS. 12A-12C). Apply sufficient power to (shown). Alternatively, the output applied to the electromagnet 26 may be adjustable. In this case, in use, the power increases until the desired magnetic field strength is achieved, thereby sufficient to distort the guidewire tip 36 away from the electromagnet 26.

  In another embodiment, the electromagnet 26 may include multiple portions, each portion being electrically isolated from the other portions. Each part may be electrically connected to an energy source (not shown) and may operate individually or as a whole to generate a magnetic field. If a weak magnetic field is desired, only one part is activated. If a relatively strong magnetic field is desired, one or more parts are added to operate. Further details in one embodiment of the electromagnet 26 are disclosed in US Pat. No. 6,961,620. This is expressly incorporated herein by reference.

  In an alternative embodiment, the distal end 24 of the catheter body 21 can carry a magnetic attraction element 36 instead of an electromagnet 26. The distal end 34 of the guidewire body 31 can carry an electromagnet 26 instead of a magnetic attraction element 36.

  In the illustrated embodiment, the insulated electrode 25 takes the form of a linear electrode assembly. This includes insulated annular electrodes 25a, 25b mounted at a distal end 24 surrounded by an insulating layer I. In particular, because the electrode 25 is split, selective monopolar and bipolar functions are provided to the catheter 20. That is, one or both of the insulated annular electrodes 25a and 25b are configured as one pole in a monopolar arrangement, and the energy released by one or both of the electrodes 25a and 25b is attached to the patient's skin from the outside. It may be returned through an electrode (not shown). Alternatively, the annular electrodes 25a and 25b may be configured as two poles in a bipolar arrangement so that energy released by one of the annular electrodes 25a and 25b returns to the other electrode. The combined length of the electrodes 25a, 25b is preferably about 6 mm to about 10 mm in length. In one embodiment, each electrode 25a, 25b has a length of about 4 mm and a spacing of 0.5 mm to 3.0 mm.

  In the illustrated embodiment, the electrodes 25a, 25b may include a solid ring of conductive material such as platinum. Alternatively, a conductive material such as platinum-iridium or gold coated on the device using conventional coating techniques or ion beam assisted deposition (IBAD) processes may be provided. A nickel or titanium primer may be applied to improve adhesion. Any combination of electrodes can also be in the form of a spiral ribbon. Alternatively, it can be formed of a conductive ink compound pad printed on a non-conductive tubular body. A preferred conductive ink compound is a silver-based flexible adhesive conductive ink (polyurethane binder). However, other metal-based adhesive conductive inks such as platinum-based, gold-based and copper-based are also considered. Such ink is more flexible than epoxy-based ink. Insulation material disposed over the annular electrodes 25a, 25b preferably facilitates the application of the electric field to the target anatomical region without substantially heating the surrounding tissue via Joule heating. Any suitable insulating material (silicone, PEEK, PVDF, etc.) and can be any suitable thickness.

  The electrodes 25a, 25b may be electrically coupled to individual wires (not shown) disposed within the catheter 20. Here, the wire (not shown) is directly connected to the connector (not shown) received in the port on the operation assembly 40 or indirectly to the connector via the PC board (not shown) in the operation assembly 40. Are electrically coupled. A connector (not shown) plugs into the voltage waveform generator 16 (shown in FIG. 12A).

  The operating assembly 40 includes a handle 41, a steering mechanism 42, an electrical connector 52, and a guide wire doorway 50 associated with the handle 41. The handle 41 is preferably made of a durable rigid material (such as medical plastic) and is ergonomically shaped so that the physician can more easily manipulate the catheter 20. The steering mechanism 42 includes a steering switch 44. This is because the first position (off) (prevents electrical energy from the energy source from exciting the electromagnet) to the second position (on) (electrical energy from the energy source (not shown) can excite the electromagnet 26. To operate). As a result, the excited electromagnet 26 pulls the magnetic attraction element 36, thereby distorting the distal end 34 of the guidewire body 31 from a straight shape, resulting in a simple curve (ie, a curve in a single plane). ) (Shown as a phantom in FIGS. 12A-12C). The steering switch 44 is configured to operate and return to the off position, so that the guide wire body 31 can be bent by elasticity and return to a straight shape.

  The steering mechanism 42 is described as having a curved shape by bending the proximal end 34 of the guide wire main body 31 to only one side. However, the steering mechanism 42 is, for example, a third position (exciting the electromagnet 26 to the opposite polarity ), The distal end 34 of the guidewire 30 can be modified into two oppositely bent shapes along a single plane. As a result, the excited electromagnet 26 may push the magnetic attraction element 36, thereby distorting the distal end 34 of the guidewire body 31 opposite to the simple curve described above. In this case, the off position may be between two on positions so that the switch 44 must pass through the off position when changing the polarity of the electromagnet 26.

  In the illustrated embodiment, the guidewire doorway 50 may include a luer lock connector (not shown) through which the guidewire 30 is inserted. Optionally, the guidewire entry / exit 50 may have multiple uses as well as facilitate insertion of the guidewire 30. For example, the guidewire doorway 50 may provide access for inserting other flexible elongated instruments or devices during a surgical procedure. It may also provide a connection to a fluid source such as saline or contrast (not shown). The conduit 52 facilitates electrical connection between the processing device 14, the waveform generator 16, and the energy source (not shown) and the actuating elements of the catheter 20 by providing a path for a conductor or wire (not shown). A wire (not shown) may be coupled to the connector (not shown) or the PC board (not shown).

  13A-D illustrate a further embodiment of a catheter according to the present disclosure having a tip. Here, the carrier assembly includes a plurality of carrier arms. Catheter 50 is constructed of a biocompatible material suitable for advancing percutaneously through the patient's vasculature and navigating through the patient's heart. Various tubular body members and shafts are constructed from injection molding materials such as Pebax, silicone, polyurethane, polymers, elastomers, flexible plastics and combinations thereof. Catheter 50 includes a tip 94. It is made from a material that is atraumatic to the tissue. Catheter 50 further includes an outer shaft 76. This is preferably between 8-9 Fr in diameter and is constructed to provide sufficient stability and torque throughout the procedure. Catheter 50 includes a carrier assembly 85. This includes a plurality of electrode elements 92, each surrounded by an insulating layer attached to the distal carrier arm 88.

  Insulated electrode 92 and other components of carrier assembly 85 are configured to bend to fit the pulmonary vein ostium and other applicable tissues. The outer shaft 76 can be advanced forward to change the shape of the carrier assembly 85 and bring one or more insulated electrodes 92 into contact with the tissue. The outer shaft 76 slidably receives the inner shaft 78. This is fixedly attached to the proximal carrier arm 86. The distal carrier arm 88 is fixedly attached to the tip 94 and the distal end of the control shaft 84. Proximal carrier arm 86 is pivotally attached to distal carrier arm 88 and advances and retracts control shaft 84 relative to inner tube 78 to reduce and expand the diameter of carrier assembly 85, respectively (4). Expand to a diameter of ~ 5mm). Inner shaft 78 also provides cylindrical strength to allow an operator to advance inner shaft 78 and ensure that carrier assembly 85 is in contact with tissue, such as to fit a non-circular pulmonary venous portal. It can be so. Inner shaft 78 is preferably attached to a pull wire (not shown) near the distal end, which is operably connected to a controller at the proximal end of device 50. This allows the operator to controllably distort the distal portion of the device 50.

  The distal end of the outer shaft 76 includes a shaft end 82. This is configured to expand radially when the carrier assembly 85 is retracted. The proximal end of outer shaft 76 preferably includes a handle (not shown). This includes one or more controllers (such as knobs or levers) for advancing and retracting the inner shaft 78 and the control shaft 84. The proximal end of device 50 includes one or more connectors. This connects to the waveform generator and applies the applied voltage to the insulating electrode 92. In an alternative embodiment, one or more proximal control arms 86 are attached to the second control axis and the symmetry of the shape of the carrier assembly 85 is adjusted to accommodate an asymmetric pulmonary vein portal. In another alternative embodiment, device 50 is configured to be inserted over a pre-positioned guidewire.

  Referring now to FIGS. 13A and 13B, the distal portion of device 50 is illustrated. FIG. 13A depicts the fully expanded carrier assembly 85 with the inner shaft 78 fully advanced. FIG. 13B depicts the inner shaft 78 retracted halfway. As a result, the proximal end 86 is captured by the shaft end 82 and compressed in the radial direction. This expands as shown, which causes the carrier assembly 85 to smoothly transition into the inner lumen of the outer shaft 76.

  Referring now to FIG. 13C, the catheter of the present disclosure is illustrated, which comprises two carrier assemblies arranged in series along a single axis. Each carrier assembly is a tip configuration. Catheter 40 includes an elongate tube or outer shaft 36. This is preferably constructed from Pebax material, has a diameter of about 6-8 Fr, and slidably receives the first control shaft 48. The first control shaft 48 is attached to the first carrier assembly 45 at a distal portion. This comprises a plurality of carrier arms and an insulation coating electrode. It is configured to apply a time varying electric field via an applied time varying voltage waveform. The proximal end (not shown) of the first control shaft 48 is attached to the controller at the proximal end of the catheter 40. This is configured so that the operator can accurately advance and retract the first control shaft 48. The first control shaft 48 includes a ring 52 at the distal end. This attaches one end of each distal carrier arm segment 44 fixedly to the first control shaft 48. Each distal carrier arm segment 44 is pivotally attached to one end of the proximal carrier arm segment 42 at its opposite end. The opposite end of each proximal arm segment 42 is fixedly attached to the distal end of outer shaft 36 via ring 38. The distal carrier arm segment 44 and the proximal arm segment 42 are constructed from a flexible material (such as Nitinol) and can be elastically biased straight or in a tip configuration. As the first control shaft 48 advances and retracts, the diameter of the carrier assembly 45 changes. When the first control shaft 48 is completely advanced, it is in a state of being completely miniaturized (minimum diameter) in the radial direction, and when the first control shaft 48 is fully retracted, it is in a maximum diameter state.

  An insulated electrode 46 is fixedly attached to the distal arm segment 44. This is configured to apply a time-varying electric field to the tissue. If desired, the electrode 46 may include a fin 54. This is configured to be present in the bloodstream during electric field application. Electrode 46 is configured to apply monopolar, bipolar or a combination of monopolar and bipolar energy. Electrode 46 may include an integrated temperature sensor (such as a thermocouple welded to the inner portion of electrode 46) if desired. Electrode 46 and any integral temperature or other sensor is attached to a wire (not shown). This travels close to the proximal portion of the catheter for attachment to an energy application unit, surveying unit and / or other electronic device for transmitting or receiving signals and / or outputs.

  The first control shaft 48 slidably receives the second control shaft 57. The second control shaft 57 is attached to the second carrier assembly 55 at the distal portion. This includes a plurality of carrier arms and electrodes configured to survey electrical activity. A proximal end (not shown) of the second control shaft 57 is attached to the controller at the proximal end of the catheter 40. This is configured so that the operator can accurately move the second control shaft 57 forward and backward. The second control shaft 57 includes a tip 62 at the distal end. This fixes and attaches one end of each distal carrier arm segment 56 to the second control shaft 57. The tip 62 is preferably constructed from a soft or flexible material, such as a soft plastic or elastomer selected to be atraumatic to the tissue, preferably like a Pebax material supplemented with barium sulfate. , Radiopaque. The tip 62 is constructed to help navigate into and stabilize within the pulmonary veins. The tip 62 includes a guidewire lumen 63. This connects the inner lumen of the second control shaft 57 and the flow path. This lumen travels to the proximal portion of the catheter 40 and exits therefrom. This allows the catheter 40 to be inserted percutaneously over the guidewire into the patient's vasculature.

  Each distal carrier arm segment 56 is pivotally attached to one end of the proximal carrier arm segment 54 at its opposite end. The opposite end of each proximal arm segment 54 is fixedly attached to the distal end of the first control shaft 48 via a ring 52. The distal carrier arm segment 56 and the proximal arm segment 54 are constructed from a flexible material (such as Nitinol) and can be elastically biased straight or in a tip configuration. As the second control shaft 57 advances and retracts, the diameter of the carrier assembly 55 changes. When the second control shaft 57 is fully advanced, it is in a state of being completely miniaturized (minimum diameter) in the radial direction, and when the second control shaft 57 is fully retracted, it is in a maximum diameter state.

  Fixedly attached to the distal arm segment 44 is an insulated electrode 58. It is configured to apply an electric field and / or survey electrical activity present in tissue. The insulated electrode 58 can be constructed from a conductive material such as platinum or a combination of platinum and iridium and coated with a dielectric material. Electrode 46 may preferably include an integral temperature sensor (such as a thermocouple welded to the inner portion of electrode 58). Electrode 58 and any integral temperature or other sensor are attached to a wire (not shown). This travels close to the proximal portion of the catheter 40 for attachment to a surveying unit, energy application unit and / or other electronic device for transmitting or receiving signals and / or outputs.

  The catheter 40 of FIG. 13C includes a handle (not shown) at the proximal end. This is preferably of the type described with reference to FIG. 13D and includes a plurality of controllers. Thereby, the operator moves the first control shaft 48 forward and backward, moves the second control shaft 57 forward and backward, operates the voltage waveform application to one or more of the electrodes 46 or 58, and applies the energy application unit or surveying. The user interface of the unit (both not shown) can be operated or another function can be performed. The handle includes an exit. Through this, a guide wire (such as a guide wire placed in the patient's pulmonary vein) can be removed. The carrier assembly 55 is sized to engage the lumen wall of the pulmonary vein. The carrier assembly 45 is sized to be able to engage a pulmonary vein ostium including a non-circular hole and is sufficiently flexible thereto. The outer shaft 36 is constructed of a material sufficient to apply a force that matches the carrier assembly 55 and / or the carrier assembly 45, and the handle may be manipulated as such. Both the first control shaft 48 and the second control shaft 57 are configured to transmit sufficient torque so that the operator can accurately rotate and position the carrier assembly 45 and the carrier assembly 55, respectively.

  With further reference to FIG. 13D, a catheter 40 is illustrated, which includes the dual carrier assembly of the catheter of FIG. 13C. Catheter 40 includes a tubular body member or outer shaft 36. This includes a proximal carrier assembly 45 and a distal carrier assembly 55 at the distal end, as described in detail with reference to FIG. 13C. The proximal end of the outer shaft 36 is attached to a handle 66. This includes a plurality of controllers. That is, the slide 67, the slide 68, and the button 69. The slide 67 is operably attached to the first control shaft 48. The slide 68 is operably attached to the second control shaft 57. As the slides 67 and 68 are moved, the shapes of the first carrier assembly 45 and the second carrier assembly 55 change as described in detail with reference to FIG. 13C. Many types of mechanical mechanisms can be incorporated into the handle 66 to operably advance one or more control shafts (linear knobs, rotary knobs or rotary levers such as knobs connected to the cam assembly) And other mechanisms used to move the shaft back and forth). When the first carrier assembly 45 is placed in contact with the pulmonary vein portal and the catheter 40 is electrically connected to an energy application unit (not shown), the button 69 is used to initiate voltage waveform application.

  The handle 66 includes two pigtails. One ends with a luer 74 and the other ends with an electrical connector 72. The luer 74 is in fluid communication with a guidewire lumen 63 exiting from the tip 62 to allow the catheter 40 to advance over the wire into the patient's vasculature. The electrical connector 72 includes a plurality of connection points for a plurality of wires. This travels through the outer shaft 36 and is connected to an insulated electrode and one or more sensors, such as temperature sensors included in the first carrier assembly 45 and the second carrier assembly 55. The electrical connector 72 is configured to electrically connect to one or more of the following. That is, an energy application unit, a surveying unit, an electronic device that receives and / or transmits signals and / or outputs, such as signals received from the temperature of catheter 40 or other physiological sensors, and combinations thereof.

  Referring now to FIG. 13E, the catheter of the present disclosure is illustrated. Here, the carrier assembly includes an inflatable balloon having a plurality of insulated electrodes attached or embedded on its outer surface. Catheter 100 includes a carrier assembly 175. This includes a balloon 174 and an insulated electrode 172 that applies an electric field to the tissue. Catheter 100 further includes an elongate tubular body, i.e., outer shaft 166. This includes an inflation lumen 176 from the proximal portion to the interior cavity of balloon 174. Balloon 174 is sealed and fixedly attached to the distal portion of outer shaft 166. As fluid such as air or saline passes through the inflation lumen of the outer shaft 166, the balloon 174 is inflated and remains expanded as long as fluid pressure is maintained. Balloon 174 may be a stretchable or non-stretchable balloon and may be shown as a disc or donut shape, while having a specific contour that mimics the pulmonary venous ostium and tissue extending therefrom.

  A centering post 168 extends coaxially from the distal end of the outer shaft 166. This may include a protrusion that traverses from the proximal end to the distal end of balloon 174 and is configured to engage the pulmonary vein lumen. In a preferred embodiment, a guidewire lumen is included from the distal end to the proximal portion of the catheter 100, such as a guidewire in which the catheter is pre-placed into the pulmonary vein portal or other applicable hole, etc. Allows advancement over the guidewire.

  Referring now to FIG. 14, a further embodiment of a catheter (illustrating only the distal catheter region) is presented. This has a shaft 10, which in turn includes an outer covering. This may be nylon, urethane or other plastic. The outer covering 34 surrounds the reinforcing material 36. This usually involves a stainless steel braiding or winding in turn. The reinforcing material 36 is disposed around the base layer 38. This preferably comprises a tube of polyimide or other relatively rigid or high durometer material. The shaft stiffness and torque transfer characteristics can be varied by changing the type of material used for the outer covering 34, the stiffener 36 and the base layer 38. In addition, by using differently shaped reinforcements 38. For example, the reinforcement 36 can be a braid or a coil. The number of single fibers, the shape of the single fibers, the winding or weaving pattern, the number of windings, etc. can be varied individually or in combination to provide the desired stiffness. Preferably, the polyimide tube of the base layer 38 has a thickness in the range of 0.002 inches to 0.005 inches.

  Outer covering 34, reinforcement 36, and base layer 38 define a central lumen 40 that extends the length of shaft 10. Disposed within the central lumen 40 is a core wire 42, a pull wire 44, an electrode wire 46, and a thermocouple wire 48 if desired. The tip 16 includes tubing 50 of low durometer flexible plastic (such as Pebax® material, silicone rubber, or other elastic material). Preferably, tip 16 has a durometer in the range of 30A to 60D. Tubing 50 typically has at least four lumens extending in length parallel to the longitudinal axis. That is, a central lumen 54 and at least three radially offset lumens 56 (four are illustrated). The core wire 42, along with the electrode wire 46 and the thermocouple wire 48, extends through the central lumen 54. The pull wire 44 extends from the central lumen of the shaft 10 to a lumen 56 that is offset in the radial direction of the distal end portion 16.

  The spring tube 58 is also disposed within the central lumen 54 of the tip 16. The spring tube 58 fits snugly against the wall of the inner lumen 54 and has a hollow center. Core wire 42, electrode wire 46 and thermocouple wire 48 extend therethrough. The spring tube 58 typically comprises a polyimide tube. This provides the tip 16 with lateral and torsional stiffness as well as torsional resistance. The spring tube 58 can also be a braided or wound structure, or a composite of multiple layers. The tip 16 is fixed to the shaft 10 by butt joint 18, preferably by thermal welding. The central lumen 54 of the tip segment 16 is smaller in diameter than the central lumen 40 of the shaft 10 and the spring tube 58 extends into the central lumen 40 of the shaft 10 a distance, typically about 0.5 inches. Yes. Such an extension limits the twist at or near the butt joint 18 when the tip 16 is distorted. The proximal end 60 of the spring tube 58 extends into the central lumen 40, thereby improving the stiffness at the transition between the tip 16 and the remainder of the shaft 10.

  Core wire 42, electrode wire 46 and thermocouple wire 48 extend from central lumen 40 of shaft 10 through the center of spring tube 58 and into central lumen 54 of tip 16. At the distal end of the tip 16, the spring tube 58 exits the central lumen 54 and enters the opening 64 in the tip electrode 20. The output wire 66 (one of the electrode wires 46) is coupled to the insulated tip electrode 20 and applies a time-varying electric field. The insulation may be complete or incomplete as desired and can be applied in any desired pattern (grating, spiral, radial, etc.).

  The thermocouple wire 48 terminates in a thermocouple 68 disposed in the opening 64. Preferably, the opening 64 is filled with a high temperature adhesive so that the thermocouple 68 is maintained in place. The electrode band wire 70 exits from the central lumen 54 in the spring tube 58 and couples to the strip electrode 22. The core wire 42 extends through the central lumen 54 and into the opening 64 of the tip electrode 20.

  An electrically and thermally insulated fuser plate 72 is bonded to the distal end 74 of the tubing 50, and the tip electrode 20 is bonded to the distal side of the fuser plate 72. The fixing plate 72 has a central passage 78 corresponding to the central lumen 54 of the tip 16 and four radially offset openings 76 through which the pull wire 44 passes. Pull wire 44 terminates at anchor 80. This typically comprises a steel ball formed or welded to the end of the pull wire 44. Anchor 80 is larger in diameter than opening 76, thereby providing a strong and pivotable connection between pull wire 44 and the distal end of tip 16. The fuser plate may be formed from any polymer or ceramic material having the required mechanical strength and electrical and thermal insulation properties. Preferably, polyether ether ketone is used. It is available from ICI Americas (Wilmington, Del.) Under the trade name Victrex.

  FIGS. 15A-15B illustrate using a further embodiment of a catheter to treat diseased tissue 48 in the lower esophagus 42. One embodiment of the system 10 may be attached to a flexible endoscope 12 (also referred to as endoscope 12). The flexible endoscope 12 includes an endoscope handle 34 and a flexible shaft 32. The endoscope system 10 typically includes one or more treatment / diagnostic probes 20, a plurality of conductors 18, a handpiece 16 having a switch 62, and an electrical waveform generator 14. The treatment / diagnostic probe 20 is located at the distal end of the flexible shaft 32. The conductor 18 is attached to the flexible shaft 32 using a plurality of clips 30. The treatment / diagnostic probe 20 includes an elongated member. This is attached to an electrical energy application device comprising one or more insulated electrotherapy electrodes 28. In one embodiment, treatment / diagnostic probe 20 extends through a hole in flexible shaft 32, such as working path 36. In one embodiment, the treatment / diagnostic probe 20 may comprise an elongated diagnostic probe 26. It is attached or joined to an electrode 28 that extends through the working path 36. In another embodiment, the flexible shaft 32 includes two working paths 36, a first diagnostic probe 26 joined to a first electrode 28 extending through the distal end of the first working path 36, and a second working path. A second diagnostic probe 26 joined to a second electrode 28 extending through the distal end of the path 36 may be provided. In one embodiment, the diagnostic probe comprises one or more diagnostic probes 26. It is attached or joined to electrode 28 in any suitable manner. For example, the diagnostic probe 26 may be joined or attached to the electrode 28 by welding, soldering, brazing, or other known techniques. Many different energy sources may be used for welding, soldering, or brazing, for example, gas flames, electric arcs, lasers, electron beams, friction, and ultrasound. Accordingly, in one embodiment, treatment / diagnostic probe 20 may be used to obtain a biopsy specimen of diseased tissue using diagnostic probe 26 in diagnostic mode, and in one embodiment, electrode 28 in therapeutic mode. The treatment / diagnostic probe 20 may be used by treating the affected tissue with an electric field applied by. In other embodiments, the therapeutic / diagnostic probe 20 may be used in a combination of therapeutic and diagnostic modes. The therapy / diagnostic probe 20 may pass through one or more working paths 36 located within the flexible shaft 32. A treatment / diagnostic probe 20 is applied endoscopically to the tissue treatment area and placed on top of the affected tissue to be electrically treated. When the treatment / diagnostic probe 20 is properly positioned by the operator, the switch 62 on the handpiece 16 is manually operated to electrically connect or disconnect the electrode 28 from the electrical waveform generator 14. As an alternative, the switch 62 may be mounted on a foot switch (not shown), for example.

  The operator positions the treatment / diagnostic probe 20 using endoscopic visualization such that the affected tissue 48 to be treated is within the field of view of the flexible endoscope 12. When the operator positions the treatment / diagnostic probe 20 and positions the insulated electrotherapy electrode 28 above the diseased tissue 48, the operator energizes the insulated electrode 28 with the electrical waveform generator 14 to provide a tissue treatment region. A strong electric field may be applied to the affected tissue 48. Electrode 28 may be energized with a microsecond or nanosecond electrical waveform, as described elsewhere herein. In this manner, the affected tissue 48 in the tissue treatment area may be electroporated and lead to cell death.

  The device 700 illustrated here in FIG. 16 provides a minimally invasive microsurgical instrument. Typically, the electrode 700 is a catheter-type device having an insulated activity 712 and a grounded 711 electrical conducting wire embedded therein. The electric conductive wires 711 and 712 have an annular shape, and are arranged in the length direction along the length of the tip 710 of the electrode 700. It may be entirely or partially covered by a dielectric material. The insulated conductive wires 711, 712 are separated from each other longitudinally along the length of the electrode 700 by a selected distance and separated by insulation. Electrode 700 is compatible with existing electroporation electronics and includes a universal connector 750 that connects proximal end 718 of electrode 700. This operably communicates with an electrical pulse generator configured to apply a waveform, as described elsewhere herein. To operably communicate with the electrical pulse generator, a portion of the electrically conductive wires 711, 712 is placed in the outer body 770 of the electrode 700, or the electrically conductive wire is placed from the electrically conductive wires 711, 712 and closer to the tip 710. Run along the length of the electrode 700 up to the distal end 718. The electrically conductive wire can comprise any type of conductive metal (such as platinum / iridium). Different materials have different radiopacity and can be selected according to this property to achieve specific results. For example, silver is much more radiopaque than titanium, and therefore some embodiments of the electrodes of the present invention may use silver as a marker while having a titanium conductive surface / wire.

To ensure biocompatibility, electrode embodiments may be covered with an insulating polyurethane coating 770 to confine the electrically conductive wire leading to the electrical pulse generator.
In embodiments, the electrically conductive wire need not be entirely conductive. For example, the conductive surface may comprise one or more portions that are insulated (particularly near the base). The exposed portion or surface of the electrically conductive wires 711, 712 at the tip 710 can be of any thickness and width. Similarly, the amount of separation distance between the electrically conductive wires at the tip 710 can be any amount. The electrode may comprise any number of conductive wires. In embodiments, the electrodes can be configured such that the separation distance between the electrically conductive wires is adjustable and / or the amount of exposed conductive surface is adjustable.

  The embodiment of FIG. 16 is constructed as a thin device. This facilitates direct navigation through the cardiovascular system to the treatment site. If desired, the electrode may comprise a J-shaped tip (or similar shape), as is common in angioplasty and catheter-based interventions. This can guide the electrode through the vascular system and reach the target site. Such an asymmetric tip may also be used as a starting point for determining the direction of rotation in this particular embodiment. Typical dimensions of the electrode device (such as about 0.5 mm in diameter) can be reliably placed in the coronary artery. This is because it is smaller than that already used in catheterization. This diameter or width is therefore on the order of 5 mm to 1 cm. Preferably, the diameter or width is from about 0.5 mm to about 5 mm (such as about 1 mm, 2 mm, 3 mm, or 4 mm). When used in humans, the device is typically on the order of 40 cm or less in length (such as about 30 cm, 25 cm, or 15 cm), and when used in veterinary medicine, this length is the size of the animal to be treated. Depending on it, it can be much larger. For the treatment of human brain tumors, this length can be on the order of 40 cm.

  Referring now to FIG. 17A, the catheter 100 includes an elongate member 102, typically a tubular polymer extrudate, and includes a deployable arcuate electrode delivery structure 104. FIG. This structure can be collapsed by retracting into the tube 102 or can be deployed by advancing away from the tube, as shown in FIG. 17A. A partially or fully insulated electrode array 106 is formed on the outer surface of the arcuate support 104 so that it can engage the outer periphery of the mucosal inner wall M of the lower esophagus E, as shown in FIG. After applying a time-varying electric field through the electrode array 106, the arcuate deployment structure 104 may be rotated so that the electrode array 106 engages a new area of the mucosal wall of the esophagus E. The new area is treated as indicated by TR2 in FIG. 17C. Subsequent rotation of the arcuate deployment structure 106 may treat the entire outer wall of the esophagus. Catheter 100 can be axially repositioned to treat consecutive axial sections of the esophagus as well. Referring now to FIG. 17D, the treatment catheter 120 is similar to the catheter 100 and a pair of deployable arcuate structures 122 and 124 may be provided. A partially / fully insulated array 126 of electrodes may be formed on the outer and / or inner surface of the deployable arcuate structure. Using a pair of deployment structures 122 and 124 reduces the total treatment time required when compared to the catheter 100 of FIG. 17A.

  As shown in FIG. 17E, an inflatable balloon 130 may be used to sequentially treat the mucosal wall of the esophagus as well. Although the majority of the balloon 130 is formed from an elastomer, a non-inflatable dimensionally stable portion 132 is formed as part of the balloon and spaced apart in a fixed and dimensionally stable manner. It carries a partially or fully insulated electrode 134. Therefore, the elastomeric balloon may be used by being completely inflated in the esophagus to treat the outer periphery of the esophagus. After the first treatment is completed, the balloon may be rotated by a desired amount and used to treat the next outer periphery. Such sequential treatment is continued until the entire circumference of the esophagus is treated.

  According to a further embodiment, as illustrated in FIG. 18, the treatment device 10 includes an elongate shaft 22. This can be flexible and is configured to be inserted into the body in any of a variety of ways selected by a healthcare provider. The shaft 22 may be placed (i) endoscopically, eg, through the esophagus 14, (ii) surgically, or (iii) by other means. When using an endoscope (not shown), the shaft 22 can be inserted into the lumen of the endoscope. Alternatively, the shaft 22 can be placed outside the endoscope. Alternatively, an endoscope may be used to visualize the path that the axis 22 should follow during placement. In addition, the shaft 22 can be inserted into the esophagus 14 after removing the endoscope.

  An energy applicator 24 may be provided and placed at the distal end 26 of the shaft 22 to provide the appropriate energy for the desired electric field application. In various embodiments, the energy applicator 24 is coupled to an energy source as described elsewhere herein. This drives the energy applicator 24 at a level appropriate to provide the desired electroporation of the tissue. In one embodiment, shaft 22 includes a cable. This includes a plurality of conductors surrounded by an electrically insulating layer, with an energy applicator 24 disposed at the distal end 26. A positioning and inflation device may be coupled to the shaft. The positioning and inflation device may be configured and sized to contact and expand the wall of the body cavity in which it is placed, including but not limited to the esophagus. The positioning and expansion device may be at different locations on the energy application device 24. This includes, but is not limited to, proximal and / or distal ends and sides.

  The energy applicator 24 may be supported at a controlled distance from the wall of the tissue site or in direct contact therewith. This can be accomplished by coupling the energy applicator 24 to the expandable member 28. Suitable expandable members 28 include balloons, extensible balloons, tapered balloons, ridges, multiple struts, expandable members with rolled and unwound states, one or more springs, foams, bladders, restraints Including, but not limited to, a backing material that expands to an expanded configuration when not done. The expandable member 28 can be utilized to adjust and control the amount of energy transferred to the tissue at the tissue site. This can occur with the esophageal wall. The expandable member 28 can be joined to a portion of the shaft 22 at a point away from the distal end 26. In any of the embodiments of FIG. 18, the conductive elements can be partially or fully insulated to apply an electric field. In another embodiment, the expandable member 28 is utilized to apply the electric field energy itself. The expandable member can be made from a variety of different materials. This includes, but is not limited to, conductive elastomers such as mixtures of polymers, elastomers and conductive particles. Non-stretchable bladders have a fully expanded form shape and size that stretches in a manner suitable for contacting tissue.

  In another embodiment, the conductive member may be formed from an electrically insulating elastomer having a conductive material (such as copper) deposited on the surface. The electrode pattern can then be etched into the material. And then, the conductive member can be attached to the outer surface of the balloon. Thereafter, a further partial or total coating of insulating material may be deposited on top of the electrically conductive material. In one embodiment, the conductive member can be a balloon 28 and expandable to a shape that matches the size of the expanded (uncollapsed) inner lumen (such as the human lower esophageal canal) of a tissue site or structure Have Further, the conductive member can include a plurality of electrode range segments 30. This is partially or totally covered by an insulating material.

  Further exemplary electrodes according to embodiments of the present disclosure are shown in FIGS. Some of the components of the electrode device include guide wires, tips, outer bodies, connectors, balloons, and partially / fully insulated conductive wires. The configuration depicted in this embodiment is 2 mm in diameter and is therefore designed for larger coronary vessels (such as the right or left aorta). This has proximal to intermediate lumen diameters of about 3.6 and 4.3 mm, respectively. Dodge, J .; T.A. , Jr. , B. G. Brown, E.I. L. Bolson and H.C. T.A. Dodge “Lumen diameter of normal human coronary arts, Inflation of age, sex, atomic variation, and left ventricular hypertropy ord. Such a tube may have a typical restenosis lumen diameter of about 2.5 mm. Radke, P.M. W. A. Kaiser, C.I. Frost and U.S. Sigwart “Outcome after treatment of coronary in-stent restenosis- Results from a systematic review using meta-analysis techniques, 27, 66, Europe 3 Hur.

  Even smaller designs can be built with modern microfabrication techniques. This may target smaller tubes (such as distal portions and branches of the coronary arteries). Furthermore, the apparatus of the present disclosure may be enlarged for other applications. Accordingly, the diameter of the outer body of the electrode of the present invention can range, for example, from about 0.1 mm to about 5 cm. Preferably, electrode embodiments of the present invention have an outer diameter from about 0.5 mm to 5 mm (such as from about 1 mm to about 3 mm, such as from about 1.5 mm to about 2 mm), or from about 0.75 mm to 3 cm. Up to.

  In an embodiment, the guide wire is the thinnest physical component and is used to guide the surgeon into the appropriate tube. Although not typical, it can be hollow. This may be used with soluble contrast agents used for angiography and fluoroscopy as well as typical endovascular therapy. Schwartz, R.A. S. J. et al. G. Murphy, W.M. D. Edwards, A.D. R. Camrud, R.A. E. Vliestra and D.C. R. See Holmes, “Restenosis after ballon angioplasty, Apractical proliferative model in porcine coronary artists” Circulation, 1990, 82 (6), 2190-200. The guide wire is usually smaller in diameter than the diameter of the outer body of the balloon.

  The guide wire or catheter forms a support for which all other components of the device are placed. An inflatable balloon is placed over the portion of the catheter between the tip and proximal end of the device. The balloon is secured to the catheter at the distal and proximal ends of the device. An inflation mechanism is also provided that provides fluid in the region between the balloon and the catheter. The balloon can be inflated with any inert fluid (such as saline, contrast fluid, air, or even a low electrical conductivity aqueous sucrose solution) during use. In embodiments where an electrically conductive wire is placed inside the balloon (between the balloon and the catheter) and there is a highly conductive fluid in the lumen (inside the balloon), the current flow is more than through the wire. It may preferentially pass through the fluid and thereby obtain a more diffused electric field. In a preferred embodiment, a low conductivity buffer is preferred as a fluid for inflating the balloon and treating the target tissue more accurately.

  A fully / partially insulated conductive wire or electrode can apply a time-varying voltage waveform to apply electric field target tissue from an electrical pulse generator. During use of the device, the proximal end of the wire / electrode is in operative communication with the pulse generator. The wire can be connected directly to the pulse generator. Alternatively, in a preferred embodiment, the electrode comprises a universal connector (or other connection structure) to secure the electrically conductive wire / electrode to the electrical pulse generator in an operatively connected manner. If the electrode has a greater number of electrically conductive wires than the number of outputs available on the pulse generator that it wants to use together, the electrode and / or pulse generator can be modified or modified to operably cooperate with each other. obtain. In embodiments, if there are more conductive wires (eg, 8) on the electrode than there are outputs (eg, 6 outputs) on the pulse generator, the system of the present invention may comprise an electrode-generator interface. This works with the generator to switch which wire is activated for a given set of pulses in the overall flow. More specifically, the interface may comprise a switch that switches between wires 4 and 8 for an 8-electrode system. This is because these wires are unlikely to be energized simultaneously 180 degrees apart. Another example of the operation of the therapeutic interface is basically using two pairs of pulses at the same time, the system connects to all eight wires at the output while only receiving positive and negative inputs from the generator Can have. Here, the pair 1-2 to 2-3 etc. are automatically switched, and the positive and negative outputs (ports 1 and 2) of the generator change the pulse set 1 at 2000V (wires 1-2) (for example) Then, the voltage is changed to 1500 V (wires 2 to 3) second.

  In this embodiment, the electrically conductive wire is long and is placed along the length of the balloon. The wires are placed around the circumference of the electrode at a selected distance from each other. 19A-19B, there are eight electrically conductive wires / electrodes spaced circumferentially around the balloon. The distal and proximal ends of the electrically conductive wire / electrode are protected or covered by a relatively thick outer insulating body, while the middle portion of the wire / electrode to be placed against the target anatomical region is It contains a relatively thin insulating layer or is not partially insulated and is directly exposed to the atmosphere or tube wall in use.

  The tip of the balloon preferably includes a conical distal end so that it can be easily advanced through the lumen. The conical distal end is preferably formed from, or in operative communication with, a portion of the outer body of the device. The outer body of the device is an insulative container that protects the components of the device and controls the exposure of the electrically conductive wires. Another part of the outer body of the device provides a proximal container disposed at the proximal end of the electrode. In embodiments, there is a gap or separation distance between the containers or portions distal and proximal of the outer body. This gap or unprotected area of the electrode exposes the electrically conductive wire to the atmosphere.

  A connector is disposed on the proximal side of the balloon. The connector is hollow and can move relative to the guide wire and has an internal chamber. This extends all the way back to the proximal end of the balloon catheter and can carry fluid (such as saline or air or a low electrical conductivity sucrose solution).

  The balloon is attached to the connector. Running through the body / connector assembly is an array of eight conductive 35 gauge wires (0.15 mm). These wires are insulated through the entire electrode device until they reach a highly insulating balloon and are then exposed. In embodiments, a conductive wire may be attached to the proximal portion of the connector, but the distal end is free and confined within the tip container. Thereby, a wire can be freely expanded with expansion | swelling of a balloon. In other embodiments, the wire may be attached to the surface of the balloon and / or the distal end of the wire is attached to the distal portion of the tip or distal outer body or electrode connector.

Accordingly, in some exemplary embodiments, the present disclosure provides:
1. A method of generating a time-varying electric field in a living tissue of an animal or a human body. The time variation of the electric field takes the form of a pulse. This is a rise and fall time lasting between about 1 and 100 microseconds, lasting between about 2 and 200 microseconds (or any 1 microsecond unit in between), with a peak size of 100-100. , 000 V / cm (or any 1.0 V / cm unit in between). A pulsed or oscillating voltage of approximately the same duration is applied between the conductive electrodes. At least one electrode is placed near the tissue and conduction current flowing from the electrode into the tissue is substantially reduced or prevented by the presence of a dielectric insulator surrounding the electrode. The peak value of the electric field in the tissue in the target anatomical region is controlled by the rate of voltage variation and the peak voltage imposed by the electrical signal generator. A current limiting barrier layer is configured to be disposed between the at least one electrode and the target anatomical region.

  2. A method of generating a time-varying electric field in a living tissue of an animal or a human body. The time variation of the electric field takes the form of a pulse. This is a rise and fall time lasting between about 1 and 100 nanoseconds, lasting between about 2 and 200 nanoseconds (or any 1 nanosecond unit in between), with a peak size of 100-100. , 000 V / cm (or any 1.0 V / cm unit therebetween). A pulsed or oscillating voltage of approximately the same duration is applied between the conductive electrodes. At least one electrode is placed near the tissue and conduction current flowing from the electrode into the tissue is substantially reduced or prevented by the presence of a dielectric insulator surrounding the electrode. The peak value of the electric field in the tissue in the target anatomical region is controlled by the rate of voltage variation and the peak voltage imposed by the electrical signal generator. A current limiting barrier layer is configured to be disposed between the at least one electrode and the target anatomical region.

  3. A device that generates a time-varying electric field in a living tissue of an animal or a human body. The time variation of the electric field takes the form of a pulse. This is a rise and fall time lasting between about 1 and 100 microseconds (or any 1 microsecond unit therebetween), lasting between about 2 and 200 microseconds, with a peak size of 100-100. , 000 V / cm (or any 1.0 V / cm unit in between). A pulsed or oscillating voltage of approximately the same duration is applied between the conductive electrodes. At least one electrode is placed near the tissue and conduction current flowing from the electrode into the tissue is substantially reduced or prevented by the presence of a dielectric insulator surrounding the electrode. The peak value of the electric field in the tissue in the target anatomical region is controlled by the rate of voltage variation and the peak voltage imposed by the electrical signal generator. A current limiting barrier layer is configured to be disposed between the at least one electrode and the target anatomical region.

  4). A device that generates a time-varying electric field in a living tissue of an animal or a human body. The time variation of the electric field takes the form of a pulse. This is a rise and fall time lasting between about 1-1000 nanoseconds (or any 1 nanosecond unit therebetween), lasting between about 2-2000 nanoseconds, with a peak size of 100-100 , 000 V / cm (or any 1.0 V / cm unit therebetween). A pulsed or oscillating voltage of approximately the same duration is applied between the conductive electrodes. At least one electrode is placed near the tissue and conduction current flowing from the electrode into the tissue is substantially reduced or prevented by the presence of a dielectric insulator surrounding the electrode. The peak value of the electric field in the tissue in the target anatomical region is controlled by the rate of voltage variation and the peak voltage imposed by the electrical signal generator. A current limiting barrier layer is configured to be disposed between the at least one electrode and the target anatomical region.

  All statements herein (in addition to specific examples) that enumerate the principles, aspects, and embodiments of the present disclosure are intended to include structural and functional equivalents thereof. Moreover, such equivalents are intended to include currently known equivalents. Of course, it includes both equivalents that will be developed in the future, that is, any elements that are developed to achieve the same function regardless of structure.

  Descriptions herein of circuits and method steps represent conceptual embodiments of exemplary circuits and software that embody the principles of the disclosed embodiments. Accordingly, the functionality of the various elements shown and described herein may be provided through dedicated hardware. Of course, it may be hardware capable of executing software in connection with appropriate software according to this specification.

  In this disclosure, any element expressed as a means of realizing a particular function is intended to encompass all ways of realizing that function. This may include, for example, a) a combination of circuit elements that implement the function and associated hardware, or b) software in any form shown herein (and thus including firmware, microcode, etc.) It includes a combination with a circuit that executes the software, suitable for realizing the function. Applicant therefore regards any means that can provide those functionalities as equivalent to those shown herein.

  Similarly, it is understood that the systems and process flows described herein represent various processes. This is substantially represented in a computer readable medium and may be performed as such by a computer or processing device (whether or not such computer or processing device is explicitly indicated). Further, various processes may be understood not only as representing processes and / or other functions, but instead as blocks of program code that perform such processes or functions.

  Among other things, the methods, systems, computer programs and portable devices of the present disclosure, as described above and illustrated in the drawings, provide improved methods, systems and machine readable programs for performing therapeutic procedures. It will be apparent to those skilled in the art that various modifications and variations can be made in the devices, methods, software programs and portable devices of the present disclosure without departing from the spirit or scope of the disclosure. Accordingly, this disclosure is intended to include modifications and variations within the scope of subject disclosure and equivalents. All documents referred to herein are expressly incorporated herein by reference in their entirety for all purposes.

Claims (22)

A system for treating a region of biological tissue,
with an electric signal generator configured to generate a substation field when having a) a plurality of pulses, said plurality of pulses, respectively, the rise and fall time lasts between 1 to 100 microsecond, 2 200 persisted for the duration of between microseconds,
b) is arranged close to the target anatomical region comprising at least one electrode configured to apply the time-substation field to the subject anatomical region, said electric field, the magnitude of the peak 1 00 It is between ~100,000V / cm, peak value of the electric field among the tissue in the subject anatomical region, voltage regulation, and, by the peak voltage imposed by the electrical signal generator Controlled ,
c) a current limiting barrier layer configured to be disposed between said at least one electrode the subject anatomical region,
system. A system for treating a region of biological tissue,
with an electric signal generator configured to generate a substation field when having a) a plurality of pulses, said plurality of pulses, respectively, the rise and fall time lasts between 1 to 1000 nanoseconds, 2 2000 and lasts for a duration of between nanoseconds,
b) is arranged close to the target anatomical region comprising at least one electrode configured to apply the time-substation field to the subject anatomical region, said electric field, the magnitude of the peak 1 00 It is between ~100,000V / cm, peak value of the electric field among the tissue in the subject anatomical region, voltage regulation, and, by the peak voltage imposed by the electrical signal generator Controlled ,
c) a current limiting barrier layer configured to be disposed between said at least one electrode the subject anatomical region,
system. The system of claim 1 or 2,
The current limiting barrier layer is configured to prevent conduction current through the biological tissue;
system. The system according to any one of claims 1 to 3 ,
The current limiting barrier layer comprises an insulating material having a thickness between 1 and 2000 micrometers;
system. The system according to any one of claims 1 to 3 ,
The current limiting barrier layer includes a gas fill gap and begins to conduct current therethrough when the electric field across the gas fill gap exceeds a gas ionization threshold in the gas fill gap;
system. The system according to any one of claims 1 to 3 ,
The current limiting barrier layer comprises an ion conductive liquid or polymer,
system. The system according to any one of claims 1 to 3 ,
The at least one electrode is configured to apply an electric field to the target anatomical region, which is oriented in a direction that is primarily perpendicular to the electrode within the target anatomical region;
system. The system according to any one of claims 1 to 3 ,
The at least one electrode is configured to apply an electric field to the target anatomical region, which is oriented in a direction that is primarily in contact with the electrode within the target anatomical region;
system. The system according to any one of claims 1 to 3 ,
The at least one electrode is formed proximate to a surface of an inflatable member that can be selectively expanded to bring the at least one electrode into proximity with the target anatomical region;
system. The system according to any one of claims 1 to 3 ,
The at least one electrode is patterned, the patterning being defined by regions of alternating exposed conductors and insulating materials, or regions of alternating insulating materials of different thicknesses;
system. The system according to any one of claims 1 to 3 ,
The at least one electrode is flexible;
system. The system according to any one of claims 1 to 3 ,
The at least one electrode is elastic;
system. The system according to any one of claims 1 to 3 ,
The at least one electrode includes a plurality of elongated conductors surrounding a central axis, which can be deployed radially outward from the central axis.
system. The system of claim 13 , wherein
The plurality of elongated conductors includes an insulating layer substantially along their entire length,
system. 15. The system of claim 14 , wherein
The thickness of the insulating layer varies along the length of the plurality of long conductors.
system. The system of claim 15 , wherein
A thickness of the insulating layer is reduced along a portion of the plurality of elongated conductors configured to be disposed proximate to the target anatomical region;
system. The system of claim 13 , wherein
The plurality of elongated conductors are partially insulated along portions of the plurality of elongated conductors configured to be disposed proximate to the target anatomical region,
system. 16. A method comprising applying an electric field to the target anatomical region on or in a non-human animal using the system according to any of claims 1-15 . The method of claim 18 , wherein
The target anatomical region, time between one second and one hour of ten minutes, exposed to the electric field,
Method. The method of claim 18 , wherein
The target anatomical region, time between five seconds and secondary sufficiently exposed to the electric field,
Method. The method of claim 18 , wherein
The target anatomical region, time between thirty seconds and ten minutes, exposed to the electric field,
Method. The method of claim 18 , wherein
The subject anatomical region is exposed to the electric field for a time between three and seven minutes;
Method.

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