MEDICAL DC CURRENT GENERATOR AND BIPOLAR MEDICAL IMPLANT FRAGMENTATION DEVICE EQUIPPED THEREWITH

Abstract

The present invention is directed to an endoscopic implant cutting and/or fragmenting apparatus of the bipolar type, operating on direct current, comprising an endoscope instrument having at least two opposing electrodes at its distal instrument head forming a cutting gap inbetween for receiving an electrically conductive implant or implant section to generate punctiform physical contact with the implant, and a DC-impulse generator having or connected to a control device adapted to generate a direct current in a pulsed way being controlled by the control device such that in a first phase of physical contact, the current pulse is adjusted preferably by controlling the current value at the electrodes to induce electric energy into the implant material being sufficient to melt the implant material exclusively in the area of the contact portion and in a second phase of physical noncontact, the current pulse is adjusted preferably by controlling the voltage value at the electrodes to generate an electric arc between at least one electrode and the melted implant material being sufficient to cut the melted implant material.

Claims

1. A medical DC-impulse generator of a medical/surgical bipolar fragmentation device, operating on direct electric current, and adapted to fragment medical implants that are made of an electrically conductive material, comprising an electric current source and an electric current control device with the DC-impulse generator being adapted to be connected to a medical endoscopic instrument having at least two electrodes at its distal end portion, which are adapted to physically contact the medical implant, and to apply a direct electric current of predetermined or adjustable strength in a pulsed or timed way to the electrodes, such that the electric current flows from one electrode to the other electrode via the implant material and thereby applies electrical energy onto the implant material for cutting said material in a cutting process, wherein said electric current control device of the medical DC-impulse generator has a first control portion or an implemented first control process being adapted to apply in a first phase (t.sub.0−t.sub.1) an electric current after physical contact between the electrodes of the medical endoscopic instrument and the medical implant is established for melting the implant material and a second control portion or an implemented second control process being adapted to maintain in a second phase (t.sub.1−t.sub.2) the current flow through the implant material after a loss of the physical contact between at least one of the electrodes and the implant material maximal for a predetermined maintaining duration, wherein the predetermined maintaining duration is pre-adjusted or pre-adjustable to a value safely avoiding re-solidification of the implant material before the separation/dissection of the implant segments is completed.

2. The medical DC-impulse generator according to claim 1, wherein in the first phase (t.sub.0−t.sub.1) of direct contact between the electrodes and the implant, the current value of the electric current flowing between the electrodes through the material of the implant is controlled by the electric current control device to generate an electrical energy density at the contact portion between the electrodes and the implant adapted to melt the implant material exclusively at the contact portion, said electric current control device being adapted to detect a physical contact loss between the electrodes and the implant preferably by measuring an abrupt decrease of a current value by a current meter and/or an abrupt increase of a voltage value by a volt meter, and in the second phase (t.sub.1−t.sub.2) of a physical contact loss between the electrodes and the implant, the voltage value induced by the electric current is controlled by the electric current control device to allow an electric arc to extend between at least one of the electrodes and the implant, wherein the voltage value is preferably limited to a predetermined maximum value below a biocompatible voltage value.

3. The medical DC-impulse generator according to claim 1, wherein the electric current control device contains a safety device, which is electrically connected to the current source and adapted to limit the maximal duration of the current pulse in the first phase to a maximum first pulse width.

4. The medical DC-impulse generator according to claim 1, wherein the electric current control device contains a lagging element, which is electrically connected to the current source and adapted to maintain and limit the maximal duration of the current pulse in the second phase to a maximum second pulse width.

5. The medical DC-impulse generator according to claim 4, wherein the second phase of the two consecutive phases has a duration or pulse width between 50 μs and 600 μs.

6. The medical DC-impulse generator according to claim 1, wherein the sum of the duration of the two consecutive phases is shorter than 120 ms, preferably shorter than 100 ms and most preferably equal to or shorter than 80 ms.

7. The medical DC-impulse generator according to claim 1, wherein the electric current source is adapted to limit the voltage value at the electrodes in the second phase to a maximum value of approximately 4 V.

8. The medical DC-impulse generator according to claim 2, wherein the electric energy density introduced into the implant by the electric arc is adapted to dissect the metal of the implant.

9. The medical DC-impulse generator according to claim 4, wherein the lagging element is activated by either a voltage rise at the electrodes to a predetermined value or by a voltage reduction at a shunt which is electrically connected in series with wires connecting the electric current source with the electrodes.

10. The medical DC-impulse generator according to claim 2, wherein the current value is controlled to be >100 amperes, preferably between 140-155 amperes.

11. The medical DC-impulse generator according to claim 2, wherein the voltage value is controlled to be below a limit of approximately 4 volts at the electrodes for a low voltage, preferably between 2-4 volts at the electrodes.

12. The medical DC-impulse generator according to claim 1, wherein the electrodes are made from a heat-resistant material, in particular silver, whose resistance-adjusted specific melting energy is larger than 35.Math.10.sup.15 J m.sup.−4Ω.sup.−1 preferably larger than 100.Math.10.sup.15 J m.sup.−4Ω.sup.−1.

13. The medical DC-impulse generator according to claim 1, wherein the instrument is an endoscopic instrument being adapted to be inserted into a working channel of a standard endoscope.

14. The medical DC-impulse generator according to claim 1, wherein the instrument further comprises a distal instrument head on which at least two mutually opposing instrument branches are arranged which define between them a cutting gap for receiving an electrically conductive implant or implant section, wherein the mutually facing sides of the at least two instrument-longitudinal branches each form the electrode or are each equipped with the at least one electrode, and the electrodes are shaped to form a punctiform contact area with the metallic implant material at their mutually facing electrode sides.

15. A bipolar fragmentation device operating on direct electric current and being adapted to fragment implants that are made of an electrically conductive material such as metal, comprising an endoscopic instrument having at least two electrodes at its distal end and a DC-impulse generator, the DC-impulse generator comprising an electric current source and an electric current control device with the DC-impulse generator being adapted to be connected to the bipolar fragmentation device, the at least two electrodes being adapted to physically contact the medical implant, and to apply a direct electric current of predetermined or adjustable strength in a pulsed or timed way to the at least two electrodes, such that the electric current flows from one of the at least two electrodes to another of the at least two electrodes via the implant material and thereby applying electrical energy onto the implant material for cutting said material in a cutting process, wherein said electric current control device of the medical DC-impulse generator has a first control portion or an implemented first control process being adapted to apply in a first phase (t0−t1) an electric current after physical contact between the at least two electrodes and the medical implant is established for melting the implant material and a second control portion or an implemented second control process being adapted to maintain in a second phase (t1−t2) the current flow through the implant material after a loss of the physical contact between at least one of the at least two electrodes and the implant material maximal for a predetermined maintaining duration, wherein the predetermined maintaining duration is pre-adjusted or pre-adjustable to a value safely avoiding re-solidification of the implant material before the separation/dissection of the implant segments is completed.

Description

[0072] The invention will now be explained in more detail below with reference to preferred embodiment examples with reference to the accompanying drawings.

[0073] FIG. 1 shows an illustration of a typical DC pulse of a medical implant cutting/fragmentation device according to the present invention, consisting of a heating phase (t.sub.0−t.sub.1) and a cutting phase (t.sub.1−t.sub.2),

[0074] FIG. 2 shows the DC-impulse generator and accessories and a magnification of the tip of the medical implant fragmentation device according to the invention,

[0075] FIG. 3 shows a table with specifications of different materials,

[0076] FIG. 4A shows a schematic depiction of the material specifications along the current path and specifically the relevant segments of the DC cutter;

[0077] FIG. 4B shows a schematic depiction of an electrical equivalent circuit;

[0078] FIG. 4C shows the resistance adjusted specific melting energy along the current path;

[0079] FIG. 4D shows the current density along the current path

[0080] FIG. 5 shows an example for a re-solidified implant segment,

[0081] FIG. 6 shows the conceptual construction of two instrument branches at the distal instrument head of the device in physical contact engagement (short circuit engagement) with an electrically conductive implant,

[0082] FIG. 7A shows the establishment of contact between an implant segment and the medical implant fragmentation device according to the invention;

[0083] FIG. 7B shows a cross-section of an implant segment to be cut and marking of current path between electrodes.

[0084] FIG. 8 shows the current control device of the medical implant fragmentation device according to the invention.

[0085] As shown in FIG. 1, according to the present invention at t.sub.0, a DC impulse is released when two distal electrodes of a surgical instrument of the medical implant fragmentation device, represented by two semi-spherical elements 1, are in physical (punctiform) contact with an implant (stent) 2 being sandwiched by the two electrodes 1, wherein melting of the implant material in the contact area between the two opposing electrodes 1 and the implant 2 starts (first/melting phase). In case of an interruption of direct (physical) contact between at least one of the electrode 1s and the implant/stent material (at t.sub.1), for example because of a movement of the already melted implant material, electric arcs between the electrodes 1 and the implant/stent material arise and maintain an electrical connection between the electrodes 1 and the implant/stent (second/cutting phase). These electric arcs lead to an abrupt increase of the voltage U.sub.Gr between the electrodes 1, and at the same time to an abrupt decrease of the current I.sub.S. Both the current I.sub.S and the voltage U.sub.Gr become noisy signals during the second/cutting phase due to the instability of the current conduction through electric arcs (the electric arcs are unpredictably “dancing” within the space between the electrodes 1 and the material of the implant 2, much like the electric arc generated in argon plasma coagulation).

[0086] The figures are only illustrative drawings; diagram axes are not to scale. Nevertheless, it shall be clear from FIG. 1 that the subsequent second/cutting phase is much shorter than implied by the illustrative diagram (second/cutting phase is about 3 orders of magnitude shorter than the first/melting or heating phase). At t.sub.2 at the latest (or even earlier), the electric arc finally collapses since the material of the implant 2 has been separated and the distance between the two opposing electrodes 1 becomes too large to maintain the electric arc.

[0087] It shall be noted here, that the current value and the voltage value during the above-mentioned first phase is held (adjusted) at a level being sufficient to melt the material of the implant 2 exclusively in a physical contact area between the two opposing electrodes 1 because of a short circuit therebetween. Preferably, the current source is controlled or selected such that a maximum current value of about 150 A and a voltage value of 36 to 48V is generated, which leads to a voltage value at the electrodes of about 2V. At the time point at which the physical contact gets lost, the current value drops and the voltage value rises automatically. However, according to the present invention, during the above-mentioned second phase directly following the first phase current value is (automatically) held (adjusted) at a level just being sufficient to generate/keep the electric arc between the electrodes 1 and the material of the implant 2. This can be done by a current control or by the respective selection of the suitable current source. Preferably, the voltage value at the electrodes rises up to about 4V (because of the limits of the selected current source or its control) wherein the current value adjusts itself respectively because of the generated electric arcs.

[0088] Furthermore, the maximal duration of the second phase is selected to be about 500 μs in this preferred embodiment which maximal duration is pre-selected based on tests with different implants/implant materials substantially guaranteeing the cutting success but also the protection of the patient, It shall be clear here, that the duration of the second phase could be shorter than the preferred maximal 500 μs because of an earlier cut of the implant.

[0089] The medical implant fragmentation device in FIG. 2 represents an example for fragmentation of stents, such as tissue clips in the digestive tract. In order to be applicable with conventional flexible endoscopes, gastroscopes and colonoscopes, a cutting instrument (instrument shaft adapted to be introduced into the working channel of a well-known endoscope and having a distal instrument tip being equipped with the two opposing branches 2 as shown in FIG. 1) preferably has to have a length of min. 220 cm and a diameter of max. 2.6 mm. The DC pulse generated by a DC-impulse generator as shown in FIG. 2 must be conducted through the instrument shaft to the tip and back again, which leaves a total conductive length of min. 440 cm (2×min. 220 cm) for the DC pulse. Considering a required isolation and at least two additional sensing wires (as shown in FIG. 8), the (copper) cross-section available for conducting the DC pulse is not more than 2.5 mm.sup.2 per single wire (i.e. 2.5 mm.sup.2 over 440 cm).

[0090] The medical implant fragmentation device is a bipolar device with at least the two electrodes 1 provided in/at the two instrument branches at the distal end of the instrument shaft shown in FIG. 2. In case of three electrodes 1 (for improved gripping), two of the three electrodes have the same polarity). The instrument shaft contains at least four, preferably six conductive wires 10, 11, 12 as shown in FIG. 8: [0091] two wires 11 for high current (large cross-section) connected to each one of two poles, and [0092] at least two low diameter wires 10, 12 for voltage and/or current sensing, respectively, connected to one of the two electrodes 1 (poles) and/or a shunt 3.

[0093] The sensing wires 10 are required since the voltage decrease along the high current wires (about approx. 25 V along the entire high current path). The voltage, which is measured at the instrument tip, varies from approximately 2 V during the melting phase of the stent (first/melting phase (t.sub.0−t.sub.1)) to approximately 4 V. Approximately 4 V are measured when the second phase (second/cutting phase, (t.sub.1−t.sub.2)) starts. The current source in the DC-impulse generator then (automatically) tries to maintain the current, when the material starts melting and the electrodes lose contact to the melting metal. In trying to maintain the current, the current source shows the tendency to increase the voltage. To avoid a voltage which is too high and therefore dangerous for the patient, the voltage at the electrodes 1 is limited to approx. 4 V during the second phase.

[0094] The DC-impulse generator is designed to send an electrical direct current through the bipolar, surgical instrument. This DC pulse flows through the clip segment/stent/implant, wherein the two opposing electrodes 1 at the distal tip of the instrument are establishing physical contact with the implant 1 (see also FIGS. 6 and 7), resulting in localized heating and melting of the implant material. The DC-impulse generator delivers a direct current pulse of preferably optionally between 100 A-150 A.

[0095] The instrument branches are basically spaced or spaceable such that a cutting gap inbetween has a gap width which permits/ensures an introduction of an implant or implant section 2 (OTSC (Clip) or stent wire) into the gap when coming into physical contact with the two mutually opposing longitudinal branches/electrodes.

[0096] Now, if a metallic implant or an implant section 2 is introduced into the gap, generally the implant material already at the distal end portion of the instrument branches comes into contact with the respective electrodes 1 and short-circuits them, whereby, because of the applied electric current, the implant material between the electrodes 1 is heated and melted.

[0097] The contact resistance between the electrodes 1 and the implant 2 should be as high as possible in order to securely melt the implant material (exclusively) in the contact region (i.e. from the outside to the inward), but additionally to leave implant regions further away to be unheated as much as possible.

[0098] The energy input into the implant material should take place such that heat dissipation into the surrounding patient tissue remains as small as possible, even in the absence of additional protection measures.

[0099] Melting of one volume element of implant material by DC current can be calculated as follows:

[0100] The thermal energy E.sub.th that is required to melt one volume element of a material depends on the starting temperature T.sub.0, the specific melting temperature T.sub.S, the specific heat capacity c and the specific weight p:

E.sub.th=c.Math.ρ.Math.V.Math.(T.sub.S−T.sub.0)

[0101] Given that T.sub.0=38° C. and one volume element V is 1 mm.sup.3, the values for the required melting energy per mm.sup.3 calculate as given in FIG. 3.

[0102] The disposition of a material to convert electrical energy into thermal energy is proportional to the ohmic resistance of the material. In order to consider this for the case of the present invention, where the same current I.sub.S runs through different materials, the specific melting energy is adjusted by the ohmic resistance (by multiplication with the reciprocal value of the specific ohmic resistance). This calculates a value, which allows for comparing different materials regarding their individual willingness to melt at a given current, and therefore regarding their suitability for being used in the current path of the present invention.

[0103] A relevant value for material selection (considering only electrical properties) is the resistance adjusted specific melting energy e. The aim of this selection is to maximize the difference between the material to be melted (the implant material) and the material used in the DC cutter instrument, in order to achieve very selective melting of the implant material, while the material used in the DC cutter instrument does not come even close to its melting temperature. The resistance adjusted specific melting energy is constant for a material and calculates as follows:

[00003]$e=c.Math.\rho .Math.V.Math.\left(T_s-T_0\right).Math.\frac{1}{r}$

with the following variables:
e resistance adjusted specific melting temperature
c specific heat capacity
ρ specific weight
V volume element (here: V=1 mm.sup.3)
T.sub.S melting temperature
T.sub.0 temperature before energy intake (here: T.sub.0=38° C.)
r specific ohmic resistance

[0104] An overview over relevant specifications of materials is given in FIG. 3.

[0105] Although copper has the highest resistance adjusted specific melting energy e, silver has been chosen as electrode material due to its biocompatibility. Consequently, the conductive wires inside the instrument shaft are made of copper, as they do not come in contact with the patient's tissue.

[0106] Material specifications along the current path of the DC cutter are given in FIG. 4. The upper schematic constructive illustration shows the relevant segments of the DC cutter. Below, an electrical equivalent circuit is depicted. The third diagram shows the “resistance adjusted specific melting energy” along the current path. A fourth illustrative diagram shows the current density along the current path.

[0107] With the material selection in FIG. 4, the relation of the resistance adjusted specific melting energy e of the implant material (Nickel-Titanium) and the adjacent electrode material (silver) is 4.86 vs. 143. This means that the tendency of the implant material to melt at the given current is 29.4 times higher than the tendency of the electrode material to melt. This provides a large safety margin and ensures that the electrode material does not heat up close to melting temperature, while the implant material reaches melting temperature much faster. Additionally, the wire material has an even higher resistance adjusted specific melting energy e of 210, thus the tendency of the wires to melt is even lower than the one of the electrodes.

[0108] FIG. 5 shows an example where an implant segment has re-solidified after unsuccessful heating to melting temperature. During the heating phase (first phase as shown in FIG. 1), the material of the medical implant 2 in the heating zone reaches melting temperature; this temperature must be maintained for a short time in order to allow the implant segment to separate before the melted region (about the contact region between the electrodes and the implant) solidifies again. However, during the first phase separation of the segment cannot be safely achieved for the implant as shown in FIG. 5. In order to avoid re-solidification of the melted material (after losing physical contact between the electrodes 1 and the implant 2 especially because of movements of the already melted implant material), and to ensure effective separation of the implant segment, the cutting process according to the invention comprises the above mentioned second/cutting phase after the heating/melting (first) phase: The cutting phase which supports the cutting with an electric arc, has an abrasive effect on the melted material. Since medical implants generally have elastic properties (pre-tension), these properties normally support the fragmentation of the material when melting. In case the current pulse (during physical contact) is not sufficient (until the physical contact interrupts), the implant material cools down after melting, and the supporting elastic characteristic of the implant 2 is lost or at least reduced. No/low pre-tension is left in the implant material and it is very hard to fragment this re-solidified implant 2 again. This can happen if a medical cutting instrument according to EP 2967712 A1 is used. The additional abrasive electric arc, which is produced in the second (cutting) phase in the present invention, however, overcomes this drawback.

[0109] In FIG. 6, the punctiform physical contact between the implant material and the electrodes 1 is shown. Additionally, it can be seen, that a leakage current I.sub.L occurs if both poles/electrodes 1 of the DC cutting instrument of the bipolar medical implant fragmentation device are in direct contact with adjacent tissue.

[0110] The DC pulse current I.sub.S (specified to max. 155 A) is fed through the implant and generates a voltage drop U.sub.Gr between the electrodes 1. This voltage drop leads to the leakage current I.sub.L being fed through the tissue (if the electrodes 1 are in physical contact with tissue) and is a reason for limiting the voltage in the second/cutting phase. The leakage current I.sub.L spreads over the adjacent tissue, while the current density J decreases with increasing distance to the electrode pair. The current density J depends on electrode spacing and distance to the electrode pair as well as on the total leakage current I.sub.L. In the consideration of the DC pulse current for cutting, I.sub.L is not considered, since I.sub.L is lower than I.sub.S by several orders of magnitude.

[0111] In FIG. 7, the left drawing shows how a physical contact between an implant and DC cutting instrument is established. The right drawing shows the cross-section of an implant segment 2 to be cut and marking of current path I.sub.S between electrodes 1.

[0112] In FIG. 8, the control or working mechanism of the DC-impulse generator is shown in a block diagram and described in detail in the following.

[0113] The implant 2 is situated between the electrodes 1. The electrodes 1 are in electric contact (physical contact) with the implant 2. The electrodes 1 are further connected to the current source 4 via the two high current wires 11. The current from the current source 4 flows through the wires 11 to the electrodes 1 and therefore also through the (metallic) implant 2 to heat it up and melt the implant material. The shunt 3 is connected in series with one of the wires 11, which means that the current flowing through the implant 2 to melt it also flows through the shunt 3.

[0114] First ends of the two sensing wires 12 are connected to a first and a second end of the shunt 3, second ends of the sensing wires 12 lead to an unit 9 for current control. This current control 9 also measures the voltage drop over the shunt 3 when current is flowing through the wires 11. This voltage is proportional to the current flowing through the wires 11 and the implant 2. A first end of the sensing wires 10 is coupled to a distal end of the wires 11, respectively, preferably to the electrodes 1. A second end of the sensing wires 10 is connected to the voltage control unit 8, respectively.

[0115] If an activation switch 5 (e.g. a foot switch) is activated, the current source in the DC-impulse generator is activated. At the same time a safety device 6 is activated. The safety device 6 defines the maximum time the current source 4 is activated. The maximum time up to which the current source 4 remains active is defined as 120 ms, but preferably the current source 4, and therefore the DC-impulse, is deactivated after 80 ms. The reason being that the long wires 11 cause a large voltage drop over the wires 11. In connection with the current flowing through the wires 11, the wire material has to bear a power load of over 4 kW. It is clear that this would heat up the wires 11 and, in connection, also the complete instrument.

[0116] When the current flows through the wires 11, the current control unit 9 and the voltage control unit 8 detect a voltage and a current curve. The voltage curve is measured parallel to the electrodes 1, the current curve is measured via a voltage drop over the shunt 3.

[0117] When the current source 4 is activated via the activation switch 5, the metallic material of the implant 2 is in physical contact with the electrodes 1. The implant material (metal) is heated up and starts to melt. If the material melts, it loses contact to the electrodes 1, since the material “pinches” and an arc is formed between the electrodes 1 and the melted material (metal) of the implant 2. In other words, the melting material loses its form and does not physically contact the electrodes anymore.

[0118] In this moment, when the electrodes 1 lose contact to the material/metal of implant 2, an abrupt change in current and voltage occurs. Since this process takes place in fractions of a second, this is preferably measured by two instances: course of the current and course of the voltage. The current control unit 9 as well as the voltage control unit 8 can both emit a signal due to a “positive criterion” after current/voltage measurement. This means, that the current control unit 9 emits the signal according to a dl/dt measurement. If the electrodes 1 loses physical contact to the material/metal of the implant 2, the current strength in the wires 11 drops.

[0119] More or less simultaneously, the voltage, which is measured preferably at the electrodes 1 (the distal end of wires 11), rises from a first value to a second value. The voltage between the electrodes 1 during the first/melting phase (physical contact between electrodes and implant material/metal) is in a range from about 1.5 to 2.5 V, dependent on the current and the resistance between the electrodes 1 and the implant 2. The “positive criterion” for detecting a change in voltage is dependent on dU/dt (rising edge). Once the electrical arc starts at the beginning of the second phase (melted material loses physical contact to electrodes 1), the voltage between electrodes 1 rises and is preferably limited in a range from 3.5 to 4.5 V by the control device (or more concretely voltage control unit 8) of the DC current generator. Preferably, the voltage is (automatically) limited to a maximum value of about 4 V (biocompatible value) especially in the second phase. This is done in the current source 4 (DC current generator) and is necessary, since an ideal current source tries to maintain its load-independent direct current by rising its output voltage as high as necessary. This would most probably be dangerous for a human being.

[0120] Once the voltage control unit 8 detects the rise of the voltage, it emits a control signal, like the current control unit 9 does when detecting the drop of current at the wires 11. This starts the second/cutting phase.

[0121] The control signals of the voltage control unit 8 and the current control unit 9 are fed to a lagging element 7. This lagging element 7 is connected to the current source 4.

[0122] When the lagging element 7 is activated, the so-called “cutting-phase” (t.sub.1−t.sub.2) which has been previously mentioned in the description of the two consecutive phases, is maintained by the lagging element 7. The lagging element 7 can be programmed with different “time constants”. The time constant defines how long the lagging element 7 allows the current source 4 to maintain the second/cutting phase. The maximum time constant in the lagging element 7 is about 600 μs, which is not exceeded. Preferably, the lagging element 7 maintains the second/cutting phase for 50 μs to 500 μs.

[0123] The choice of the timings is caused by the compromise between the “faradic effect” and cutting efficiency. Generally it can be said, the higher the lagging time is, the higher is the cutting efficiency, but also the risk caused by the faradic effect. The lower the lagging time, the lower the risk of the faradic effect, but at the same time the cutting efficiency decreases.

[0124] The duration of the time constants is a good compromise between the faradic effect and the electrolytic effect.

[0125] If, for whatever reason, the lagging element 7 is defective and is about to maintain the cutting phase for too long, the safety device 6 assures that the complete process time (melting phase 1+cutting phase 2) does not exceed 120 ms. Preferably, the current source 4, or rather the DC-impulse has a maximum length (start to stop) of not more than 80 ms.

[0126] To summarize the above explanation, the present invention is generally directed to a medical endoscopic implant cutting and/or fragmenting device of the bipolar type, operating on direct current, comprising an endoscope instrument being adapted to be inserted into the working channel of an endoscope and having at least two opposing electrodes at its distal instrument head forming a cutting gap inbetween for receiving an electrically conductive implant or implant section to generate punctiform physical contact with the implant, and a DC-impulse generator having or connected to a control device adapted to generate a direct current in a pulsed way being controlled by the control device such that in a first phase of physical contact, the current pulse is adjusted preferably by controlling the current value at the electrodes to induce electric energy into the implant material being sufficient to melt the implant material exclusively in the area of the contact portion and in a second phase of physical noncontact, the current pulse is continued wherein the duration of continuing the current pulse is adjusted not to exceed a limit value of maximal about 600 μs after the loss of physical contact between the electrodes and the implant material is detected.

LIST WITH REFERENCE SIGNS

[0127] 1 electrode [0128] 2 implant (metallic) [0129] 3 shunt [0130] 4 current source [0131] 5 activation switch (e.g. foot switch) [0132] 6 safety device [0133] 7 lagging element [0134] 8 voltage control unit [0135] 9 current control unit [0136] 10 sensor wires (voltage control) [0137] 11 current wires (high current to electrodes) [0138] 12 sensor wires (current control)