Patent Description:
It is known that medical practitioners have found it useful to use surgical instruments to assist in the performance of surgical procedures. A surgical instrument is designed to be applied to a surgical site on the patient. The practitioner is able to position the surgical instrument at the site on the patient at which the instrument is to perform a medical or surgical procedure. Endoscopic surgical procedures are routinely performed in order to accomplish various surgical tasks. In an endoscopic surgical procedure, small incisions, called portals, are made in the patient. An endoscope, which is a device that allows medical personnel to view the surgical site, is inserted in one of the portals. Surgical instruments used to perform specific surgical tasks are inserted into other portals. The surgeon views the surgical site through the endoscope to determine how to manipulate the surgical instruments in order to accomplish the surgical procedure. An advantage of performing endoscopic surgery is that, since the portions of the body that are cut open are minimized, the portions of the body that need to heal after surgery are likewise reduced. Moreover, during an endoscopic surgical procedure, only relatively small portions of the patient's internal organs and tissue are exposed to the open environment. This minimal opening of the patient's body lessens the extent to which a patient's organs and tissue are open to infection.

Many tube devices have been developed for use in surgical procedures. They are valuable because they facilitate reduced incision size, improved access and visibility, while enhancing surgical outcome and quicker recovery. Some are surgical cutting tools that use a burr or shaver to remove bone or tissue. The surgical cutting tools include a driveshaft that rotates at a high rate of speed, such as <NUM>,<NUM> revolutions per minute. The driveshaft is enclosed by an outer tube that supports the driveshaft and protects the driveshaft from contacting tissue within the body. The close fit between the driveshaft and the outer tube generates a large amount of friction and heat. A surgical instrument that overcomes these challenges is desired.

Surgical instruments are known from <CIT>, <CIT> and <CIT>.

The invention provides a surgical cutting tool according to claim <NUM>. The dependent claims refer to developments of the invention. A first aspect of the present disclosure is directed to a surgical cutting tool is configured for use with a handpiece having a motor. The surgical cutting tool includes an outer tube extending longitudinally between proximal and distal ends. The outer tube includes an inner surface defining a lumen therethrough. The surgical cutting tool further includes a flexible driveshaft including an inner tube rotatably disposed in the lumen and extending in a longitudinal direction past the distal end to a cutting end outside of the lumen. The driveshaft further includes at least two torsion sections spaced from one another longitudinally along the inner tube. The at least two torsion sections each have at least one channel extending through the inner tube. The driveshaft further includes a bearing section disposed longitudinally between the torsion sections. The channels are configured to collapse under rotational load in response to transmission of torque as the driveshaft rotates such that the torsion sections are spaced apart from the inner surface of the outer tube to a greater extent than the bearing section for reducing friction between the inner tube and the outer tube.

In certain implementations, the inner tube includes a pair of channel walls individually corresponding to and defining each of the at least one channel. The pair of channel walls may oppose and face toward one another to define a width of the channel therebetween. The width may be at least <NUM> microns. The pair of channel walls may extend substantially parallel to one another. The pair of channel walls may be configured to draw toward one another and collapse the channel when under rotational load in response to transmission of torque with rotation of the driveshaft. The outer diameter of the torsion sections is configured to reduce at least five percent when the channels collapse. The outer diameters of the torsion and bearing sections may be substantially equal when the driveshaft is stationary.

In certain implementations, the at least one channel includes a plurality of channels in each torsion section and evenly spaced from one another. The plurality of channels may have a uniform size and shape. The plurality of channels may extend in a helical configuration longitudinally along and about the inner tube. For example, the plurality of channels may extend about the inner tube less than <NUM> degrees or less than <NUM> degrees. The plurality of channels in the helical configuration may extend at an angle at or between <NUM> and <NUM> degrees from an axis orthogonal to the longitudinal direction of the inner tube.

In certain implementations, the outer tube and the lumen are disposed in a non-linear configuration between the ends, and the driveshaft may be correspondingly disposed in the non-linear configuration within the lumen. A length of the torsion sections may have a ratio of at least <NUM>:<NUM> relative to a length of the bearing section.

A second aspect of the present disclosure is directed to a flexible driveshaft is configured for use with a surgical cutting tool. The driveshaft is configured to be rotatably disposed in a lumen of an outer tube. The driveshaft includes an inner tube extending in a longitudinal direction to a cutting end, and a torsion section having an outer diameter and a plurality of channels extending through the inner tube. The plurality of channels extends in a helical configuration longitudinally along and about the inner tube in a first rotational direction. The driveshaft is configured to rotate in a second rotational direction, opposite the first rotational direction. The channels are configured to collapse under rotational load in response to transmission of torque as the driveshaft rotates in the second rotational direction such that the outer diameter of the torsion section is reduced.

In certain implementations, the flexible driveshaft according to the second aspect of the present disclosure may be included on the surgical cutting tool of the first aspect, the third aspect, and/or the fourth aspect, and optionally, any of their corresponding implementations.

A third aspect of the present disclosure is directed to a surgical cutting tool is configured for use with a handpiece having a motor. The surgical cutting tool includes an outer tube extending longitudinally between proximal and distal ends. The outer tube includes an inner surface defining a lumen therethrough. The surgical cutting tool further includes a flexible driveshaft including an inner tube rotatably disposed in the lumen and extending in a longitudinal direction past the distal end to a cutting end outside of the lumen. The driveshaft includes pairs of channel walls individually corresponding to and defining channels. The pairs of channel walls may oppose and face toward one another to define a width of the channels therebetween. The width may be at least <NUM> microns. The channels are configured to collapse under rotational load in response to transmission of torque.

A fourth aspect of the present disclosure is directed to a surgical cutting tool is configured for use with a handpiece having a motor. The surgical cutting tool includes an outer tube extending longitudinally between proximal and distal ends. The outer tube includes an inner surface defining a lumen therethrough. The surgical cutting tool further includes a flexible driveshaft including an inner tube rotatably disposed in the lumen and extending in a longitudinal direction past the distal end to a cutting end outside of the lumen, wherein a portion of the flexible driveshaft is configured to collapse under rotational load in response to transmission of torque.

A method of operating a surgical cutting tool for use with a handpiece having a motor is further disclosed herein. The surgical cutting tool includes an outer tube extending longitudinally between proximal and distal ends and including an inner surface defining a lumen therethrough. The surgical cutting tool further includes a flexible driveshaft including an inner tube rotatably disposed in the lumen and extending in a longitudinal direction past the distal end to a cutting end outside of the lumen. The driveshaft further includes at least two torsion sections spaced from one another longitudinally along the inner tube and each a having at least one channel extending through the inner tube. The driveshaft further includes a bearing section disposed longitudinally between the torsion sections. The method includes the steps of rotating the driveshaft to transmit torque, applying a rotational load to the driveshaft, collapsing the channels, and spacing the torsion sections from the inner surface of the outer tube to a greater extent than the bearing section for reducing friction between the inner tube and the outer tube.

Referring to <FIG>, a surgical cutting tool is shown at <NUM> for use with a handpiece <NUM> having a motor. The surgical cutting tool <NUM> as shown in the Figures and described below is used in a medical procedure for a patient (not shown). The surgical cutting tool <NUM> may comprise an attachment <NUM>, such as burr (shown in the <FIG> and <FIG>) or a shaver. The motor of the handpiece <NUM> drives the surgical cutting tool <NUM> at a high rate of speed. Exemplary uses of the attachment <NUM> include abrading and resecting bone and tissue, such as during endoscopic sinus surgery. However, the surgical cutting tool <NUM> may also be adapted for other medical procedures, including but not limited to, spinal, neuro, and endoscopic applications. It should be appreciated that the surgical cutting tool <NUM> may be operated by a user (not shown) such as a surgeon.

As shown in <FIG>, the surgical cutting tool <NUM> includes an outer tube <NUM> extending longitudinally between proximal and distal ends <NUM>, <NUM> and including an inner surface <NUM> defining a lumen <NUM> therethrough. The surgical cutting tool <NUM> further includes a flexible driveshaft <NUM> including an inner tube <NUM> rotatably disposed in the lumen <NUM>. The driveshaft <NUM> extends in a longitudinal direction X past the distal end <NUM> to a cutting end <NUM> outside of the lumen <NUM>. The aforementioned attachment <NUM> may be coupled to the cutting end <NUM> of the driveshaft <NUM>. Although the attachment <NUM> is shown in the Figures as a burr, any suitable attachment may be coupled to the cutting end <NUM> of the driveshaft <NUM>. The driveshaft <NUM> further includes a motor end <NUM> opposite the cutting end <NUM>. The driveshaft <NUM> is coupled to the motor of the handpiece <NUM> at the motor end <NUM>, with the motor configured to selectively rotate the driveshaft <NUM>. The motor transfers torque to the driveshaft <NUM>, which rotates the attachment <NUM> disposed at the cutting end <NUM> of the driveshaft <NUM>. The motor may rotate the driveshaft <NUM> and the attachment <NUM> at speeds greater than <NUM>,<NUM> revolutions per minute. The high-speed rotation and torque transfer from the motor to the attachment <NUM> allows the surgical cutting tool <NUM> to accurately and efficiently abrade or resect the bone and tissue during the surgical procedure as described above. Although the attachment <NUM> is configured to be rotated by the driveshaft <NUM> in the Figures, the attachment <NUM> may be configured for non-rotatable motion (such as reciprocation) in other examples not shown herein.

The outer tube <NUM> and the inner tube <NUM> may have a corresponding slip fit which allows the inner tube <NUM> to rotate axially within the lumen <NUM> while laterally retaining the inner tube <NUM> within the lumen <NUM>. As such, the outer tube <NUM> supports the inner tube <NUM> and reduces wobbling of the inner tube <NUM> during rotation. Furthermore, the outer tube <NUM> protects the inner tube <NUM> from invasion of material (e.g., tissue of the body) that could inhibit the operation of the inner tube <NUM>, such as binding the outer and inner tubes <NUM>, <NUM> and slowing the rotation of the inner tube <NUM>. The outer tube <NUM> also protects the patient from tissue degradation due to contact with the inner tube <NUM> (i.e., friction and heat from the rotating inner tube <NUM>).

The surgical cutting tool <NUM> may further comprise a nose tube <NUM> surrounding the outer tube <NUM> and extending longitudinally between the attachment <NUM> and the handpiece <NUM>. The nose tube <NUM> may be tapered as shown in the Figures to ease the insertion of the surgical cutting tool <NUM> into the body. The outer tube <NUM> may be fixed to the nose tube <NUM>, with the nose tube <NUM> further supporting the outer tube <NUM>. The nose tube <NUM> may further cover and protect the rotatable inner tube <NUM> from the body of the patient and vice-versa. Although the outer tube <NUM> and the nose tube <NUM> are shown as separate components, the outer tube <NUM> and the nose tube <NUM> may be an integral, unitary component in other examples not shown herein. As such, in other examples the outer tube <NUM> is the outer most component that interfaces the body of the patient.

As shown in <FIG>, the outer tube <NUM> and the lumen <NUM> may be disposed in a linear configuration between the ends, and the driveshaft <NUM> may be correspondingly disposed in the linear configuration within the lumen <NUM>. On the other hand, the outer tube <NUM> and the lumen <NUM> may be disposed in a non-linear configuration between the ends, and the driveshaft <NUM> may be correspondingly disposed in the non-linear configuration within the lumen <NUM>, as shown in <FIG>. Moreover, the outer tube <NUM> and the driveshaft <NUM> may vary between linear and non-linear configurations at different portions of the surgical cutting tool <NUM>. For example, the outer tube <NUM> and lumen <NUM> may be disposed in the linear configuration at the proximal end <NUM> of the outer tube <NUM> and in the non-linear configuration at the distal end <NUM> of the outer tube <NUM>, as shown in <FIG>. Transnasal applications of the surgical cutting tool <NUM> will often employ such a bend at the distal end <NUM>. On the other hand, spinal applications of the surgical cutting tool <NUM> will often employ a bend at the proximal end <NUM>.

As shown in <FIG>, the driveshaft <NUM> may further comprise at least two torsion sections <NUM> spaced from one another longitudinally along the inner tube <NUM> and each having at least one channel <NUM> extending through the inner tube <NUM>, and a bearing section <NUM> disposed longitudinally between the torsion sections <NUM>. In the example shown in the Figures, the at least one channel <NUM> includes a plurality of channels <NUM> evenly spaced from one another. More specifically, each torsion section <NUM> has a plurality of channels <NUM> extending through the inner tube <NUM>. In the example shown in the Figures, each torsion section <NUM> includes eight channels <NUM>. The number of channels <NUM> may vary depending on the diameter of the inner tube <NUM>. For exemplary purposes, the number of channels <NUM> may be within the range of <NUM> to <NUM>, within the range of <NUM> to <NUM>, or within the range of <NUM> to <NUM>. However, the torsion sections <NUM> may have any number of channels <NUM>, and the number of channels <NUM> may be based on imparting the desired flexural and bearing properties to be described in detail.

Moreover, the at least two torsion sections <NUM> and the bearing section <NUM> may be further defined as a plurality of torsion sections <NUM> and a plurality of bearing sections <NUM>, with the torsion sections <NUM> and bearing sections <NUM> individually staggered along driveshaft <NUM> (i.e., a bearing section <NUM>, followed by a torsion section <NUM>, followed by a bearing section <NUM>, followed by a torsion section <NUM>, etc.). As shown in <FIG>, each of the torsion and bearing sections <NUM> may have a length L1, L2 along the longitudinal direction X of the inner tube <NUM>, with the length L1 of the torsion sections <NUM> having a ratio of at least <NUM>:<NUM> relative to the length L2 of the bearing section <NUM>. For exemplary purposes, the length L1 of the torsion sections <NUM> may have a ratio relative to the length L2 of the bearing section <NUM> ranging from <NUM>:<NUM> to <NUM>:<NUM>. In one non-limiting example, the torsion sections <NUM> have a length L1 of <NUM> and the bearing sections <NUM> have a length L2 of <NUM>. In general, the greater the ratio between the torsion and bearing sections <NUM>, <NUM>, the greater the flexibility between of the driveshaft <NUM>. The greater the flexibility, the more the driveshaft <NUM> is able to bend to accommodate non-linear configurations of the outer tube <NUM>. However, the torsion and bearing sections <NUM> may have any suitable length L1, L2 and ratio. Furthermore, although the torsion sections <NUM> are shown in the Figures to have a uniform length L1 and the bearing sections <NUM> are shown in the Figures to have a uniform length L2, each of the lengths L1, L2 may vary. Therefore, the ratio described above may apply to one torsion section <NUM> and one bearing section <NUM> that are adjacent to one another. Furthermore, because the lengths L1, L2 may vary, the ratios between the adjacent torsion and bearing sections <NUM>, <NUM> may vary as well.

The bearing and torsion sections <NUM>, <NUM> may be integrally formed of a single material. More specifically, the driveshaft <NUM> may be formed of the single material. In one example, the single material is stainless steel. However, any material capable of transmitting torque may be utilized.

The inner tube <NUM> may have an interior surface defining an inner diameter and an exterior surface defining an outer diameter. For exemplary purposes, the outer diameter may range from <NUM> to <NUM>. However, the outer diameter may be any suitable measurement depending on the application of the driveshaft <NUM>. Moreover, the interior and exterior surfaces may define a cross-sectional thickness T of the inner tube <NUM>, therebetween, as shown in <FIG>. In one non-limiting example, the inner tube <NUM> is comprised of a <NUM> gauge hypotube. However, any suitable tubing may be used.

Turning to <FIG>, because the driveshaft <NUM> is typically comprised of stainless steel, the attachment <NUM> may fixed to the cutting end <NUM> of the driveshaft <NUM> by welding. Moreover, when the driveshaft <NUM> is comprised of the hypotube, the ends of the hypotube (i.e., the cutting and motor ends <NUM>, <NUM>) provide sockets that facilitate an improved welding interface. Known devices may utilize a cable the driveshaft with the attachment welded to strands of the cable. In such an arrangement, the strands are prone to fraying at ends, and welding involves attaching each of the strands of the cable to the attachment. The welding process may undesirably melt the individual strands. Utilizing the inner tube <NUM> possibly of unitary construction and having the aforementioned socket may require welding only two components together with greater contact area between the two for the improved welding interface.

In one example, the channels <NUM> are formed by laser cutting the inner tube <NUM>. However, the channels <NUM> may be formed using other suitable machining/forming processes, including electric discharge machining and 3D printing.

As shown in <FIG>, the channels <NUM> are configured to collapse under rotational load in response to transmission of torque as the driveshaft <NUM> rotates such that the torsion sections <NUM> are spaced apart from the inner surface <NUM> of the outer tube <NUM> to a greater extent than the bearing section <NUM> for reducing friction between the inner tube <NUM> and the outer tube <NUM>. As shown in the schematic view of <FIG>, the inner tube <NUM> may comprise a pair of channel walls <NUM> individually corresponding to and defining each of the at least one channel <NUM>. More specifically, each of the channels <NUM> includes a pair of channel walls <NUM>. The pair of channel walls <NUM> may oppose and face toward one another to define a width W of the channel <NUM> therebetween. Said differently, the pair of channel walls <NUM> are spaced from one another and thus define the channel <NUM>.

The pair of channel walls <NUM> may extend substantially parallel to one another. Moreover, the plurality of channels <NUM> may have a uniform size and shape. Said differently, the channels <NUM> are shown in the Figures to have a uniform spacing between the pair of channel walls <NUM> along the entirety of the channel walls <NUM> (i.e., the width W of the channel <NUM> may be consistent along the length of the channel <NUM>). Moreover, the width W of each of the channels <NUM> shown is shown in the Figures to be substantially equally. As shown in the schematic view of <FIG>, the channels <NUM> in the helical configuration extend linearly along the inner tube <NUM>. However, the channels <NUM> may be curved or any other suitable shape.

However, in other examples not shown herein the width W of each of the channels <NUM> may vary along the length of the channel <NUM>, as well as vary between different channels <NUM>. The width W between the pair of channel walls <NUM> may be at or between <NUM>-<NUM> microns. The greater the width W, the more that the torsion sections <NUM> are able to space from the outer tube <NUM> under rotational load. However, the width W between the pair of channel walls <NUM> may be any suitable value that permits the collapse of the channel <NUM> when under rotational load.

As shown in <FIG>, portions of the inner tube <NUM> between adjacent channels <NUM> may have a second width W2 (i.e., the distance between the adjacent channels <NUM>). The second width W2 of the inner tube <NUM> may be greater than the aforementioned cross-sectional thickness T of the inner tube <NUM> to facilitate the collapse of the channels <NUM> in the torsions sections <NUM> while ensuring localized stresses on the torsion sections <NUM> are managed. As one non-limiting example, the second width W2 and the cross-sectional thickness T have a ratio of at least <NUM>:<NUM>. However, any suitable sizing of the second width W2 and the cross-sectional thickness T may be used. Moreover, the second width W2 between adjacent channels <NUM> and the cross-sectional thickness T of the inner tube may vary around the inner tube <NUM>.

As shown in <FIG>, the pair of channel walls <NUM> are configured to draw toward one another and collapse the channel <NUM> when under rotational load in response to transmission of torque as the driveshaft <NUM> rotates. More specifically, the localized reduction of material in the torsion section <NUM> at the channels <NUM> locally weakens the torsion section <NUM> and facilitates deflection of the driveshaft <NUM> at the torsion section <NUM>. The rotational load occurs when the rotating driveshaft <NUM> experiences an opposing torque. Most commonly, the opposing torque is exerted by the material that the attachment <NUM> engages to remove (i.e., bone and tissue). The rotating attachment <NUM> engages the stationary material, which exerts a torque on the attachment <NUM> which opposes the rotation of the attachment <NUM>. The opposing torque is transmitted to the driveshaft <NUM>. The channels <NUM> collapse causing the torsion section <NUM> to deflect under the rotational load. The collapse of the channel <NUM> causes the opposing pair of channel walls <NUM> to draw toward one another. The channel <NUM> may collapse until the pair of channel walls <NUM> contact. The contact between the pair of channel walls <NUM> may restore the torsional rigidity to the torsion sections <NUM> (i.e., the lack of the channel <NUM> does not promote further deflection in the torsion sections <NUM> and therefore strengthens the torsional rigidity of the torsion sections <NUM>). The torsional rigidity of the torsion region may become substantially equal to the bearing sections <NUM> of the driveshaft <NUM>. However, the pair of channel walls <NUM> need not contact to continue to transmit rotational motion to the attachment <NUM>. If the rotational load exerted through the contact of the attachment <NUM> with the material (i.e., the tissue or bone) is of a small enough value, the torsion sections <NUM> may have sufficient rigidity even with the channel walls <NUM> spaced from one another to continue to rotate the attachment <NUM>.

The collapse of the channels <NUM> under rotational load causes the torsion sections <NUM> to be spaced apart from the inner surface <NUM> of the outer tube <NUM> to a greater extent than the bearing sections <NUM> for reducing friction between the inner tube <NUM> and the outer tube <NUM>. More specifically, each of the torsion and bearing sections <NUM> may have an outer diameter D1, D2, as shown in <FIG>. The outer diameters D1, D2 of the torsion and bearing sections <NUM> may be substantially equal when the driveshaft <NUM> is stationary. However, the outer diameter D1 of the torsion sections <NUM> may be configured to reduce at least five percent when the channels <NUM> collapse. More specifically, the outer diameter D1 of the torsion sections <NUM> may be configured to reduce at or between five and ten percent when the channels <NUM> collapse. The reduction in diameter causes the torsion sections <NUM> to be spaced from the outer tube <NUM> when under rotational load. Therefore while the entire driveshaft <NUM> engages the outer tube <NUM> when the driveshaft <NUM> is stationary, only portions of the driveshaft <NUM> engage the outer tube <NUM> when the driveshaft <NUM> is under rotational load (namely, the bearing sections <NUM>). As such, the surface area of the driveshaft <NUM> in contact with the outer tube <NUM> reduces under rotational load, which reduces friction and heat between the driveshaft <NUM> and outer tube <NUM> during operation of the surgical cutting tool <NUM>.

To facilitate the reduction in the outer diameter D1 of the torsion sections <NUM> when the channels <NUM> collapse, the channels <NUM> may extend in a helical configuration longitudinally along and about the inner tube <NUM>. More specifically, the channels <NUM> may extend in the helical configuration longitudinally along and about the inner tube <NUM> in a first rotational direction R1. The first rotational direction R1 of the channels <NUM> refers to the direction around the driveshaft <NUM> that the channels <NUM> progress from the motor end <NUM> toward the cutting end <NUM>. The first rotational direction R1 is illustrated in <FIG>. In <FIG>, the motor end <NUM> is disposed on the left side of the page, with the channels <NUM> twisting about the longitudinal direction X in a clockwise direction as the channels <NUM> progress away from the motor end <NUM>. The driveshaft <NUM> is configured to rotate in a second rotational direction (shown at R2 in <FIG>), opposite the first rotational direction R1. The channels <NUM> are configured to collapse under rotational load in response to transmission of torque as the driveshaft <NUM> rotates in the second rotational direction R2 such that the outer diameter D1 of the torsion section <NUM> is reduced. More specifically, when the driveshaft <NUM> rotates in the first rotational direction R1, the rotational load (i.e., the opposing torque) that is exerted on the driveshaft <NUM> acts in the first rotational direction R1. As such, the opposing channels walls <NUM> are pushed toward one another causing the channels <NUM> to collapse. The channels <NUM> collapse causing the torsion sections <NUM> to deflect under the rotational load.

The collapse of the channels <NUM> facilitates the reduction in the outer diameter D1 of the torsion sections <NUM> because the circumference of the driveshaft <NUM> in the torsion sections <NUM> reduces when the channels <NUM> collapse. More specifically, as is commonly known in geometry, the circumference of a circle is equal to the diameter of the circle multiplied by pi (i.e., C = πd). Therefore, a reduction in the circumference of a circle would proportionally reduce the diameter of the circle. In the example shown in the Figures, the channels <NUM> occupy and define a portion of the circumference of the driveshaft <NUM> when no rotational load is exerted on the driveshaft <NUM>. When the channels <NUM> collapse under the rotational load, the circumference of the driveshaft <NUM> reduces, which correspondingly reduces the outer diameter D1 of the torsion sections <NUM>.

As shown in <FIG>, the channels <NUM> in the helical configuration may extend at an angle A at or between <NUM> and <NUM> degrees from an axis O orthogonal to the longitudinal direction X of the inner tube <NUM>. The angle A of the channels <NUM> in the helical configuration may dictate how the driveshaft <NUM> reacts when the channels <NUM> collapse. More specifically, the angle A of the channel <NUM> may affect the flexibility of the torsion sections <NUM> and the reduction of the outer diameter D1 of the torsion sections <NUM> when under rotational load. For example, increasing the angle A between the channel <NUM> and the axis O orthogonal to the longitudinal direction X may result in torsion sections <NUM> becoming more flexible but may decrease the amount that the outer diameter D1 constricts when the channels <NUM> collapse. Conversely, decreasing the angle A between the channel <NUM> and the axis O orthogonal to the longitudinal direction X may result in torsion sections <NUM> becoming less flexible but may increase the amount that the outer diameter D1 constricts when the channels <NUM> collapse. As such, the angle A of the channels <NUM> may be configured to provide the desired flexibility and outer diameter D1 reduction of the torsion sections <NUM> depending on the application of the surgical cutting tool <NUM>. For example, configurations in which the outer tube <NUM> and the lumen <NUM> are non-linear as described above may require a large angle A between the channels <NUM> and the axis O (such as <NUM> degrees) to emphasis the flexibility of the driveshaft <NUM> within the non-linear lumen <NUM>. On the other hand, configurations in which the outer tube <NUM> and the lumen <NUM> are linear as described above may require a small angle A between the channels <NUM> and the axis O (such as <NUM> degrees) to emphasis the friction reduction between of the driveshaft <NUM> and the outer tube <NUM>. One having skill in the art will appreciate that the channels <NUM> may be positioned at any suitable angle A relative to the axis O, including outside of the bounds of the exemplary range given above.

The channels <NUM> may extend about the inner tube <NUM> less than <NUM> degrees, as shown in <FIG>. More specifically, the channels may extend about the inner tube <NUM> less than <NUM> degrees. Even more specifically, the channels <NUM> may extend about the inner tube <NUM> less than <NUM> degrees. The angle that the channels <NUM> extend about the inner tube <NUM> may be dictated by the length of the channels <NUM> and the angle A of the channels <NUM> relative to the axis O (i.e., the helical angle described above). More specifically, the greater the angle A of the channels <NUM> relative to the axis O, the longer the channels <NUM> must be to extend the same radius about the inner tube <NUM>. As the channels <NUM> increase in length, more material must be removed from the inner tube <NUM> to form the channels <NUM>, which improves the flexibility of the inner tube <NUM> but also weakens the inner tube <NUM>. Therefore, while exemplary angles have been given above, the radius that the channels <NUM> extend about the inner tube <NUM> may be any suitable value depending on the application of the surgical cutting tool <NUM>.

An exemplary method, not forming part of the invention, of operating the surgical cutting tool <NUM> is described below and shown in <FIG>. The method includes the steps of rotating the driveshaft <NUM> to transmit torque and applying the rotational load to the driveshaft <NUM>. As described above, the rotational load occurs when the rotating driveshaft <NUM> experiences an opposing torque. Most commonly, the opposing torque is exerted by the material that the attachment <NUM> engages to remove (i.e., bone and tissue). The rotating attachment <NUM> engages the stationary material, which exerts a torque on the attachment <NUM> which opposes the rotation of the attachment <NUM>. The opposing torque is transmitted to the driveshaft <NUM>.

The method further includes the step of collapsing the channels <NUM>. As described above, the channels <NUM> collapse causing the torsion section <NUM> to deflect under the rotational load. The collapse of the channels <NUM> causes the opposing pair of channel walls <NUM> to draw toward one another. The channels <NUM> may collapse until the pair of channel walls <NUM> contact. The contact between the pair of channel walls <NUM> may restore the torsional rigidity to the torsion sections <NUM> (i.e., the lack of the channel <NUM> does not promote further deflection in the torsion sections <NUM> and therefore strengthens the torsional rigidity of the torsion sections <NUM>). The torsional rigidity of the torsion sections <NUM> may become substantially equal to the bearing sections <NUM> of the driveshaft <NUM>.

The method further includes spacing the torsion sections <NUM> from the inner surface <NUM> of the outer tube <NUM> to a greater extent than the bearing section <NUM> for reducing friction between the inner tube <NUM> and the outer tube <NUM>, which reduces friction between the inner tube <NUM> and the outer tube <NUM>. More specifically, each of the torsion and bearing sections <NUM> may have the outer diameter D1, D2 described above. The outer diameters D1, D2 of the torsion and bearing sections <NUM> may be substantially equal when the driveshaft <NUM> is stationary. However, the outer diameter D1 of the torsion sections <NUM> may be configured to reduce at least five percent when the channels <NUM> collapse. The reduction in diameter causes the torsion sections <NUM> to be spaced from the outer tube <NUM> when under rotational load. Therefore while the entire driveshaft <NUM> engages the outer tube <NUM> when the driveshaft <NUM> is stationary, only portions of the driveshaft <NUM> engage the outer tube <NUM> when the driveshaft <NUM> is under rotational load (namely, the bearing sections <NUM>). As such, the surface area of the driveshaft <NUM> in contact with the outer tube <NUM> reduces under rotational load, which reduces friction and heat between the driveshaft <NUM> and outer tube <NUM> during operation of the surgical cutting tool <NUM>.

The step of applying the rotational load to the driveshaft <NUM> may be further defined as applying a first rotational load to the driveshaft <NUM>, as shown in <FIG>. The method may further comprise the steps of applying a second rotational load to the driveshaft <NUM> that is greater than the first rotational load, further collapsing the channels <NUM>, and further spacing the torsion sections <NUM> from the inner surface <NUM> of the outer tube <NUM> to a greater extent than the bearing section <NUM> and the torsion sections <NUM> under the first rotational load for further reducing friction between the inner tube <NUM> and the outer tube <NUM>, as shown in <FIG>. Said differently, the spacing of the torsion sections <NUM> from the inner surface <NUM> of the outer tube <NUM> is proportional to amount of rotational load exerted on the driveshaft <NUM>. The greater the rotational load (e.g., the second rotational load compared to the first rotational load), the greater the spacing of the torsion sections <NUM> from the inner surface <NUM> of the outer tube <NUM>. In other words, as will become apparent below, the driveshaft <NUM> works more efficiently (i.e., reduced friction and heat) under higher rotational speeds and rotational loads.

More specifically, the collapse of the channels <NUM> begins at a center <NUM> of each of the channels <NUM> (and respectively the center <NUM> of the torsion sections <NUM>) and moves outwardly toward a pair of ends <NUM> of each of the channels <NUM> (schematically shown in <FIG>). As such, greater rotational load results in a greater portion of each of the channels <NUM> collapsing from their respective centers <NUM> toward the pair of ends <NUM>. Therefore, as more of the channels <NUM> collapse, more of the outer surface of the inner tube <NUM> in the torsion sections <NUM> becomes spaced from the outer tube <NUM>, which further reduces friction between the driveshaft <NUM> and the outer tube <NUM>. Accordingly, the efficiency of the driveshaft <NUM> increases with the increase of rotational load due to the proportional decrease in friction.

Claim 1:
A surgical cutting tool (<NUM>) for use with a handpiece (<NUM>) having a motor, the surgical cutting tool (<NUM>) comprising:
an outer tube (<NUM>) extending longitudinally between proximal and distal ends (<NUM>, <NUM>) and comprising an inner surface (<NUM>) defining a lumen (<NUM>) therethrough;
a flexible driveshaft (<NUM>) comprising an inner tube (<NUM>) rotatably disposed in said lumen (<NUM>) and extending in a longitudinal direction (X) past said distal end (<NUM>) to a cutting end (<NUM>) outside of said lumen (<NUM>), said driveshaft (<NUM>) further comprising:
at least two torsion sections (<NUM>) spaced from one another longitudinally along said inner tube (<NUM>) and each a having at least one channel (<NUM>) extending through said inner tube (<NUM>); and
a bearing section (<NUM>) disposed longitudinally between said torsion sections (<NUM>),
wherein said channels (<NUM>) are configured to collapse under rotational load in response to transmission of torque as said driveshaft (<NUM>) rotates such that said torsion sections (<NUM>) are spaced apart from said inner surface (<NUM>) of said outer tube (<NUM>) to a greater extent than said bearing section (<NUM>) for reducing friction between said inner tube (<NUM>) and said outer tube (<NUM>).