TRANSLUMENALLY IMPLANTABLE HEART VALVE WITH FORMED IN PLACE SUPPORT
A cardiovascular prosthetic valve, the valve comprising an inflatable cuff comprising at least one inflatable channel that forms, at least in part, an inflatable structure, and a valve coupled to the inflatable cuff, the valve configured to permit flow in a first axial direction and to inhibit flow in a second axial direction opposite to the first axial direction, the valve comprising a plurality of tissue supports that extend generally in the axial direction and that are flexible and/or movable throughout a range in a radial direction.
This application is a continuation of U.S. patent application Ser. No. 15/061,743, filed Mar. 4, 2016, which is a continuation of U.S. patent application Ser. No. 12/976,671, filed Dec. 22, 2010, now issued U.S. Pat. No. 9,308,360, which is a continuation of U.S. patent application Ser. No. 12/197,172, filed Aug. 22, 2008, which claims priority to U.S. Provisional Application No. 60/957,691, filed Aug. 23, 2007, all of these applications are hereby incorporated by reference herein in their entirety.
BACKGROUND OF THE INVENTION Field of the InventionThe present invention relates to medical methods and devices, and, in particular, to methods and devices for percutaneously implanting a stentless valve having a formed in place support structure.
Description of the Related ArtAccording to recent estimates, more than 79,000 patients are diagnosed with aortic and mitral valve disease in U.S. hospitals each year. More than 49,000 mitral valve or aortic valve replacement procedures are performed annually in the U.S., along with a significant number of heart valve repair procedures.
The circulatory system is a closed loop bed of arterial and venous vessels supplying oxygen and nutrients to the body extremities through capillary beds. The driver of the system is the heart providing correct pressures to the circulatory system and regulating flow volumes as the body demands. Deoxygenated blood enters heart first through the right atrium and is allowed to the right ventricle through the tricuspid valve. Once in the right ventricle, the heart delivers this blood through the pulmonary valve and to the lungs for a gaseous exchange of oxygen. The circulatory pressures carry this blood back to the heart via the pulmonary veins and into the left atrium. Filling of the left atrium occurs as the mitral valve opens allowing blood to be drawn into the left ventricle for expulsion through the aortic valve and on to the body extremities. When the heart fails to continuously produce normal flow and pressures, a disease commonly referred to as heart failure occurs.
Heart failure simply defined is the inability for the heart to produce output sufficient to demand. Mechanical complications of heart failure include free-wall rupture, septal-rupture, papillary rupture or dysfunction aortic insufficiency and tamponade. Mitral, aortic or pulmonary valve disorders lead to a host of other conditions and complications exacerbating heart failure further. Other disorders include coronary disease, hypertension, and a diverse group of muscle diseases referred to as cardiomyopothies. Because of this syndrome establishes a number of cycles, heart failure begets more heart failure.
Heart failure as defined by the New York Heart Association in a functional classification.
-
- I. Patients with cardiac disease but without resulting limitations of physical activity. Ordinary physical activity does not cause undue fatigue, palpitation, dyspnea, or anginal pain.
- II. Patient with cardiac disease resulting in slight limitation of physical activity. These patients are comfortable at rest. Ordinary physical activity results in fatigue, palpitation, dyspnea, or anginal pain.
- III. Patients with cardiac disease resulting in marked limitation of physical activity. These patients are comfortable at rest. Less than ordinary physical activity causes fatigue palpitation, dyspnea, or anginal pain.
- IV. Patients with cardiac disease resulting in inability to carry on any physical activity without discomfort. Symptoms of cardiac insuffiency or of the anginal syndrome may be present even at rest. If any physical activity is undertaken, discomfort is increased.
There are many styles of mechanical valves that utilize both polymer and metallic materials. These include single leaflet, double leaflet, ball and cage style, slit-type and emulated polymer tricuspid valves. Though many forms of valves exist, the function of the valve is to control flow through a conduit or chamber. Each style will be best suited to the application or location in the body it was designed for.
Bioprosthetic heart valves comprise valve leaflets formed of flexible biological material. Bioprosthetic valves or components from human donors are referred to as homografts and xenografts are from non-human animal donors. These valves as a group are known as tissue valves. This tissue may include donor valve leaflets or other biological materials such as bovine pericardium. The leaflets are sewn into place and to each other to create a new valve structure. This structure may be attached to a second structure such as a stent or cage or other prosthesis for implantation to the body conduit.
Implantation of valves into the body has been accomplished by a surgical procedure and has been attempted via percutaneous method such as a catheterization or delivery mechanism utilizing the vasculature pathways. Surgical implantation of valves to replace or repair existing valves structures include the four major heart valves (tricuspid, pulmonary, mitral, aortic) and some venous valves in the lower extremities for the treatment of chronic venous insufficiency. Implantation includes the sewing of a new valve to the existing tissue structure for securement. Access to these sites generally include a thoracotomy or a sternotomy for the patient and include a great deal of recovery time. An open-heart procedure can include placing the patient on heart bypass to continue blood flow to vital organs such as the brain during the surgery. The bypass pump will continue to oxygenate and pump blood to the body's extremities while the heart is stopped and the valve is replaced. The valve may replace in whole or repair defects in the patient's current native valve. The device may be implanted in a conduit or other structure such as the heart proper or supporting tissue surrounding the heart. Attachments methods may include suturing, hooks or barbs, interference mechanical methods or an adhesion median between the implant and tissue.
Although valve repair and replacement can successfully treat many patients with valvular insufficiency, techniques currently in use are attended by significant morbidity and mortality. Most valve repair and replacement procedures require a thoracotomy, usually in the form of a median sternotomy, to gain access into the patient's thoracic cavity. A saw or other cutting instrument is used to cut the sternum longitudinally, allowing the two opposing halves of the anterior or ventral portion of the rib cage to be spread apart. A large opening into the thoracic cavity is thus created, through which the surgical team may directly visualize and operate upon the heart and other thoracic contents. Alternatively, a thoracotomy may be performed on a lateral side of the chest, wherein a large incision is made generally parallel to the ribs, and the ribs are spread apart and/or removed in the region of the incision to create a large enough opening to facilitate the surgery.
Surgical intervention within the heart generally requires isolation of the heart and coronary blood vessels from the remainder of the arterial system, and arrest of cardiac function. Usually, the heart is isolated from the arterial system by introducing an external aortic cross-clamp through a sternotomy and applying it to the aorta to occlude the aortic lumen between the brachiocephalic artery and the coronary ostia. Cardioplegic fluid is then injected into the coronary arteries, either directly into the coronary ostia or through a puncture in the ascending aorta, to arrest cardiac function. The patient is placed on extracorporeal cardiopulmonary bypass to maintain peripheral circulation of oxygenated blood.
Since surgical techniques are highly invasive and in the instance of a heart valve, the patient must be put on bypass during the operation, the need for a less invasive method of heart valve replacement has long been recognized. At least as early as 1972, the basic concept of suturing a tissue aortic valve to an expandable cylindrical “fixation sleeve” or stent was disclosed. See U.S. Pat. No. 3,657,744 to Ersek. Other early efforts were disclosed in U.S. Pat. No. 3,671,979 to Moulopoulos and U.S. Pat. No. 4,056,854 to Boretos, relating to prosthetic valves carried by an expandable valve support delivered via catheter for remote placement. More recent iterations of the same basic concept were disclosed, for example, in patents such as U.S. Pat. Nos. 5,411,552, 5,957,949, 6,168,614, and 6,582,462 to Anderson, et al., which relate generally to tissue valves carried by expandable metallic stent support structures which are crimped to a delivery balloon for later expansion at the implantation site.
In each of the foregoing systems, the tissue or artificial valve is first attached to a preassembled, complete support structure (some form of a stent) and then translumenally advanced along with the support structure to an implantation site. The support structure is then forceably enlarged or allowed to self expand without any change in its rigidity or composition, thereby securing the valve at the site.
Despite the many years of effort, and enormous investment of entrepreneurial talent and money, no stent based heart valve system has yet received regulatory approval, and a variety of difficulties remain. For example, stent based systems have a fixed rigidity even in the collapsed configuration, and have inherent difficulties relating to partial deployment, temporary deployment, removal and navigation.
Thus, a need remains for improvements over the basic concept of a stent based prosthetic valve. As disclosed herein a variety of significant advantages may be achieved by eliminating the stent and advancing the valve to the site without a support structure. Only later, the support structure is created in situ such as by inflating one or more inflatable chambers to impart rigidity to an otherwise highly flexible and functionless subcomponent.
SUMMARY OF THE INVENTIONIn accordance with one aspect of present invention, there is provided an inflatable or formed in place support for a translumenally implantable heart valve, in which a plurality of tissue supports are flexible and/or movable throughout a range in a radial direction. As used herein, a radial direction is a direction which is transverse to the longitudinal axis of the flow path through the valve.
Further features and advantages of the present invention will become apparent from the detailed description of preferred embodiments which follows, when considered together with the attached drawings and claims.
One cause of heart failure is failure or malfunction of one or more of the valves of the heart 10. For example, the aortic valve 34 can malfunction for several reasons. For example, the aortic valve 34 may be abnormal from birth (e.g., bicuspid, calcification, congenital aortic valve disease), or it could become diseased with age (e.g., acquired aortic valve disease). In such situations, it can be desirable to replace the abnormal or diseased valve 34.
Inflatable Prosthetic Aortic Valve Implant
With continued reference to
In the description below, the present invention will be described primarily in the context of replacing or repairing an abnormal or diseased aortic valve 34. However, various features and aspects of methods and structures disclosed herein are applicable to replacing or repairing the mitral 30, pulmonary 22 and/or tricuspid 20 valves of the heart 10 as those of skill in the art will appreciate in light of the disclosure herein. In addition, those of skill in the art will also recognize that various features and aspects of the methods and structures disclosed herein can be used in other parts of the body that include valves or can benefit from the addition of a valve, such as, for example, the esophagus, stomach, ureter and/or vesice, biliary ducts, the lymphatic system and in the intestines.
In addition, various components of the implant and its delivery system will be described with reference to coordinate system comprising “distal” and “proximal” directions. In this application, distal and proximal directions refer to the deployment system 300, which is used to deliver the implant 100 and advanced through the aorta 36 in a direction opposite to the normal direction of blood through the aorta 36. Thus, in general, distal means closer to the heart while proximal means further from the heart with respect to the circulatory system.
With reference now to
In the illustrated embodiment, the cuff 102 comprises a thin flexible tubular material 106 such as a flexible fabric or thin membrane with little dimensional integrity. As will be explained in more detail below, the cuff 102 can be changed preferably, in situ, to a support structure to which other components (e.g., the valve 104) of the implant 100 can be secured and where tissue ingrowth can occur. Uninflated, the cuff 102 is preferably incapable of providing support. In one embodiment, the cuff 102 comprises Dacron, PTFE, ePTFE, TFE or polyester fabric 106 as seen in conventional devices such as surgical stented or stent less valves and annuloplasty rings. The fabric 106 thickness may range from about 0.002 inches to about 0.020 inches of an inch depending upon material selection and weave. Weave density may also be adjusted from a very tight weave to prevent blood from penetrating through the fabric 106 to a looser weave to allow tissue to grow and surround the fabric 106 completely. Additional compositions and configurations of the cuff 102 will be described in more detail below.
With continued reference to
The illustrated inflatable structure 107 also includes inflatable struts 114, which in the illustrated embodiment are formed from an annular zig-zag pattern having three proximal bends 116 and three distal bends 118. As best seen in
As mentioned above, the inflatable rings 108 and struts 114 form the inflatable structure 107, which, in turn, defines the inflation channels 120. The inflation channels 120 receive inflation media 122 to generally inflate the inflatable structure 107. When inflated, the inflatable rings and struts 108, 114 provide can provide structural support to the inflatable implant 100 and/or help to secure the implant 100 within the heart 10. Uninflated, the implant 100 is a generally thin, flexible shapeless assembly that is preferably uncapable of support and is advantageously able to take a small, reduced profile form in which it can be percutaneously inserted into the body. As will be explained in more detail below, in modified embodiments, the inflatable structure 107 may comprise any of a variety of configurations of inflation channels 120 that can be formed from other inflatable members in addition to or in the alternative to the inflatable rings 108 and struts 114 shown in
With particular reference to
With reference to
In prior art surgically implanted valves, the valve generally includes a rigid inner support structure that is formed from polycarbonate, silicone or titanium wrapped in silicone and Dacron. These surgical valves vary in diameter for different patients due to the respective implantation site and orifice size. Generally the largest diameter implantable is the best choice for the patient. These diameters range from about 16 mm to 30 mm.
As mentioned above, the implant 100 allows the physician to deliver a valve via catheterization in a lower profile and a safer manner than currently available. When the implant 100 is delivered to the site via a delivery catheter 300, the implant 100 is a thin, generally shapeless assembly in need of structure and definition. At the implantation site, the inflation media 122 (e.g., a fluid or gas) may be added via a catheter lumen to the inflation channels 120 providing structure and definition to the implant 100. The inflation media 122 therefore comprises part of the support structure for implant 100 after it is inflated. The inflation media 122 that is inserted into the inflation channels 120 can be pressurized and/or can solidify in situ to provide structure to the implant 100. Additional details and embodiments of the implant 100, can be found in U.S. Pat. No. 5,554,185 to Block, the disclosure of which is expressly incorporated in its entirety herein by reference.
With reference to
Any number of additional inflatable rings or struts may be between the proximal and distal end 126, 128. The distal end 126 of the implant 100 is preferably positioned within the left ventrical 34 and can utilize the aortic root for axial stabilization as it may have a larger diameter than the aortic lumen. This may lessen the need for hooks, barbs or an interference fit to the vessel wall. Since the implant 100 may be placed without the aid of a dilatation balloon for radial expansion, the aortic valve 34 and vessel may not have any duration of obstruction and would provide the patient with more comfort and the physician more time to properly place the device accurately. Since the implant 100 is not utilizing a support member with a single placement option as a plastically deformable or shaped memory metal stent does, the implant 100 may be movable and or removable if desired. This could be performed multiple times until the implant 100 is permanently disconnected from the delivery catheter 300 as will be explained in more detail below. In addition, the implant 100 can include features, which allow the implant 100 to be tested for proper function, sealing and sizing, before the catheter 300 is disconnected. When the disconnection occurs, a seal at the device may be required to maintain the fluid within the inflation channels 120. Devices for providing such a seal will be described in more detail below.
With reference to
In yet another embodiment of the implant 100, the implant 100 is configured such that it does affect the mitral valve 22. In such an embodiment, the distal end 128 of the implant 100 has a protrusion or feature that pushes on the annulus of the mitral valve 22 from the aortic root or aortic valve annulus. In this way, mitral regurgitation is treated by pushing the anterior leaflet closer 22a to the posterior leaflet 22b and improving the coaptation of the valve. This feature can be a separate device from the implant 100 and/or it may be actuated by a secondary mechanism, or it may simply be a function of the shape of the implant 100.
In yet another modified embodiment the implant 100 (see
With reference back to
Various shapes of the body 102 may be manufactured to best fit anatomical variations from person to person. As described above, these may include a simple cylinder, a hyperboloid, a device with a larger diameter in its mid portion and a smaller diameter at one or both ends, a funnel type configuration or other conforming shape to native anatomies. The shape of the implant 100 is preferably contoured to engage a feature of the native anatomy in such a way as to prevent the migration of the device in a proximal or distal direction. In one embodiment the feature that the device engages is the aortic root or aortic bulb 34 (see e.g.,
Different diameters of valves will be required to replace native valves of various sizes. For different locations in the anatomy, different lengths of valves or anchoring devices will also be required. For example a valve designed to replace the native aortic valve needs to have a relatively short length because of the location of the coronary artery ostium (left and right arteries). A valve designed to replace or supplement a pulmonary valve could have significantly greater length because the anatomy of the pulmonary artery allows for additional length.
Other variations of inflatable valve shapes may include an implant 100 in which entire or substantially the entire cuff 102 forms an cylindrical pocket that is filled with fluid creating a cylinder shape with commissural supports defined by sinusoidal patterns cut from a cylindrical portion of the body 102. In such an embodiment, there may be a desire to seam or join the body 102 together at points or areas to provide passageways for fluid to flow or be restricted. This may also allow for wall definition of the body 102 defining a thickness of the cylinder. It may be desired to maintain a thin body wall allowing the largest area where blood or other fluids may pass through the valve. The wall thickness of the inflated implant 100 may vary from 0.010 to 0.100 of an inch depending upon construction, pressures and materials. There also may be a desire to vary the thickness of the cuff wall from distal to proximal or radially. This would allow for other materials such as fixed pericardial tissue or polymer valve materials to be joined to the wall where support is greatest, or allow the maximum effective orifice area in the area of the implant 100 its self. The implant 100 may be sealed fluid tight by glue, sewing, heat or other energy source sufficient to bond or fuse the body material together. There can be secondary materials added to the cuff for stiffness, support or definition. These may include metallic elements, polymer segments, composite materials.
The stiffening members 140 can be metallic wire, ribbon or tube. They may vary in thickness from 0.005 to 0.050 inches and taper or vary in thickness, width or diameter. As mentioned embodiment, the members 140 can be used to support the valve commissars 136, and/or define the height of the cuff or be attachment points for the deployment catheter. These members 140 may be sewn to or woven into the cuff material 106 through conventional techniques as described above and may be shaped with hoops to accept thread or wires. The members 140 may also be formed from a hypotube, allowing deployment control wires or a deployment control system as will be described below to pass through the stiffening wires or to attach to them. Other lengths of stiffening wires are also possible, in some instances a shorter wire may be preferred, either to allow a smaller profile, better conform to a calcified valve annulus, or to ensure positive engagement of an anchor. Short sections of stiffening wires may also be positioned in directions other than the axial direction. Positioning wires off axis may allow the valve to move more naturally relative to the native tissue, or prevent anchors from rotating and disengaging. The stiffening members 140 may be substantially straight pieces of wire.
In the embodiments described herein, the inflation channels 120 may be configured such that they are of round (see
With reference now to
In other embodiments configured for maintaining patent flow through the coronary arteries 152, the cuff 102 has an open mesh structure that allows patent flow in any orientation. The mesh structure is preferably sufficiently configured that not more than one or two of its threads or wires would cross an ostium at any position. It is also possible to access the coronary arteries with an angioplasty balloon and deform the mesh structure away from the ostium, provided that the mesh is manufactured from a plastically deformable material, such as stainless steel, or any of the biocompatable materials with similarly appropriate mechanical properties.
In order to visualize the position and orientation of the implant 100, portions of the body 102 would ideally be radio-opaque. Markers made from platinum gold or tantalum or other appropriate materials may be used. These may be used to identify critical areas of the valve that must be positioned appropriately, for example the valve commissures may need to be positioned appropriately relative to the coronary arteries for an aortic valve. Additionally during the procedure it may be advantageous to catheterize the coronary arteries using radio-opaque tipped guide catheters so that the ostia can be visualized. Special catheters could be developed with increased radio-opacity or larger than standard perfusion holes. The catheters could also have a reduced diameter in their proximal section allowing them to be introduced with the valve deployment catheter.
As mentioned above, during delivery, the body 102 is limp and flexible providing a compact shape to fit inside a delivery sheath. The body 102 is therefore preferably made form a thin, flexible material that is biocompatible and may aid in tissue growth at the interface with the native tissue. A few examples of material may be Dacron, ePTFE, PTFE, TFE, woven material such as stainless steel, platinum, MP35N, polyester or other implantable metal or polymer. As mentioned above with reference to
The cuff 102 would ideally have a diameter of between 15 and 30 mm and a length of between 6 to 70 mm. The wall thickness would have an ideal range from 0.01 mm to 2.00 mm. As described above, the cuff 102 may gain longitudinal support in situ from members formed by fluid channels or formed by polymer or solid structural elements providing axial separation. The inner diameter of the cuff 102 may have a fixed dimension providing a constant size for valve attachment and a predictable valve open and closure function. Portions of the outer surface of the cuff 102 may optionally be compliant and allow the implant 100 to achieve interference fit with the native anatomy.
Many embodiments of inflatable structure 107 shapes have been described above. In addition, as described above, the implant 100 can have various overall shapes (e.g., an hourglass shape to hold the device in position around the valve annulus, or the device may have a different shape to hold the device in position in another portion of the native anatomy, such as the aortic root). Regardless of the overall shape of the device, the inflatable channels 120 can be located near the proximal and distal ends 126, 128 of the implant 100, preferably forming a configuration that approximates a ring or toroid. These channels 120 may be connected by intermediate channels designed to serve any combination of three functions: (i) provide support to the tissue excluded by the implant 100, (ii) provide axial and radial strength and stiffness to the 100, and/or (iii) to provide support for the valve 104. The specific design characteristics or orientation of the inflatable structure 107 can be optimized to better serve each function. For example if an inflatable channel 120 were designed to add axial strength to the relevant section of the device, the channels 120 would ideally be oriented in a substantially axial direction. If an inflatable channel 120 were designed primarily to add radial strength to the relevant section of the device the channel would ideally be oriented generally circumferentially. In order to prevent tissue from extending between the inflatable channels the channels 120 should be spaced sufficiently close together to provide sufficient scaffolding.
Additionally depending on the manufacturing process used certain configurations may be preferred. For example a single spiraling balloon (see e.g.,
In other embodiments, the implant 100 is manufactured from multiple layers that are selectively fused together, then the inflation channels 120 are defined by the unfused or unjoined areas between fused areas 152. In this case any of a variety configurations of inflation channels 120 can be used. For example, as shown in
The cuff 102 and inflation channels 120 of the implant 100 can be manufactured in a variety of ways. In one embodiment the cuff 102 is manufactured from a fabric, similar to those fabrics typically used in endovascular grafts or for the cuffs of surgically implanted prosthetic heart valves. The fabric is preferably woven into a tubular shape for some portions of the cuff 102. The fabric may also be woven into sheets. The yarn used to manufacture the fabric is preferably a twisted yarn, but monofilament or braided yarns may also be used. The useful range of yarn diameters is from approximately 0.0005 of an inch in diameter to approximately 0.005 of an inch in diameter. Depending on how tight the weave is made. Preferably, the fabric is woven with between about 50 and about 500 yarns per inch. In one embodiment, a fabric tube is woven with a 18 mm diameter with 200 yarns per inch or picks per inch. Each yarn is made of 20 filaments of a PET material. The final thickness of this woven fabric tube is 0.005 inches for the single wall of the tube. Depending on the desired profile of the implant 100 and the desired permeability of the fabric to blood or other fluids different weaves may be used. Any biocompatible material may be used to make the yarn, some embodiments include nylon and PET. Other materials or other combinations of materials are possible, including Teflon, floropolymers, polyimide, metals such as stainless steel, titanium, Nitinol, other shape memory alloys, alloys comprised primarily of a combinations of cobalt, chromium, nickel, and molybdenum. Fibers may be added to the yarn to increases strength or radiopacity, or to deliver a pharmaceutical agent. The fabric tube may also be manufactured by a braiding process.
The cut edges of the fabric are melted or covered with an adhesive material, or sutured over, in order to prevent the fabric from unraveling. Preferably the edges are melted during the cutting process, this can be accomplished using a hot-knife. The blade of the tool is heated and used to cut the material. By controlling temperature and feed rate as well as the geometry of the blade, the geometry of the cut edge is defined. In one embodiment the hot knife blade is 0.060 inches thick sharpened to a dull edge with a radius of approximately 0.010 inches. The blade is heated to approximately 400 degrees F. and used to cut through a Dacron fabric at a speed of about 20 inches per minute. Preferably the cutting parameters are adjusted so that the cut edge is sealed with a thin layer of melted fabric, where the melted area is small enough to remain flexible, and prevent cracking, but thick enough to prevent the fabric from unraveling. The diameter of the bead of melted fabric is preferably between 0.0007 and 0.0070 inches in diameter.
Two edges of a fabric may be sealed together by clamping the edges together to form a lap joint, and then melting the free edge. This may be accomplished with a flame, laser energy, a heated element that contacts the fabric, such as a hot-knife or a heating element that passes near the fabric, or a directed stream of a heated gas such as air. The bead of melted fabric joining the two edges is preferably between 0.0007 and 0.0070 inches in diameter.
The fabric is stitched, sutured, sealed, melted, glued or bonded together to form the desired shape of the implant 100. The preferred method for attaching portions of the fabric together is stitching. The preferred embodiment uses a polypropylene monofilament suture material, with a diameter of approximately 0.005 of an inch. The suture material may range from 0.001 to 0.010 inches in diameter. Larger suture materials may be used at higher stress locations such as where the valve commisures attach to the cuff. The suture material may be of any acceptable implant grade material. Preferably a biocompatible suture material is used such as polypropylene. Nylon and polyethylene are also commonly used suture materials. Other materials or other combinations of materials are possible, including Teflon, flouropolymers, polyimides, metals such as stainless steel, titanium, Kevlar, Nitinol, other shape memory alloys, alloys comprised primarily of a combinations of cobalt, chromium, nickel, and molybdenum such as MP35N. Preferably the sutures are a monofilament design. Multi strand braided or twisted suture materials also may be used. Many suture and stitching patterns are possible and have been described in various texts. The preferred stitching method is using some type of lock stitch, of a design such that if the suture breaks in a portion of its length the entire running length of the suture will resist unraveling. And the suture will still generally perform its function of holding the layers of fabric together.
As mentioned above, the cuff 102 may be manipulated in several ways to form inflation channels 120. In many embodiments, the implant 100 is not provided with separate balloons 111, instead the fabric 106 of the cuff 102 itself can form the inflation channels 100. For example, in one embodiment two fabric tubes of a diameter similar to the desired final diameter of the implant 100 are place coaxial to each other. The two fabric tubes are stitched, fused, glued or otherwise coupled together in a pattern of channels 120 that is suitable for creating the geometry of the inflatable structure 107. In one embodiment the stitching pattern consists of a spiral connecting the two tubes. The spiral channel formed between the sutured areas becomes the inflation channel (see e.g.,
With reference to
In another embodiment, a single fabric tube similar to the final diameter of the prosthetic implant 100 is used. The ends or an end of the tube is turned inside out forming two layers of tube for a short length at one or both ends of the tube. The layers of tube are sewn or otherwise attached together to form a ring shaped inflation channel at the end of the tube in a manner similar to that shown in
If a porous fabric is used for the cuff 102, it may be desired to use a liner (e.g., as shown in
In the preferred embodiment, the fabric inflation channels contain a liner in a form of the balloons 111 as described with reference to
Several embodiments of the inflatable prosthetic implant 100 described above utilize circular or ringed shaped balloon members 111. These balloons 111 can be manufactured using a glass tube bent in a helix. The balloon 111 is then blown inside the tube using methods similar to those used to manufacture balloons for angioplasty. For example, the glass mold may be heated using air, water, steam infared elements and pressure and tension may be applied to blow the balloon to a specific diameter and length. Secondary processes may be added to “set” the balloon's shape by providing a second heating process to hold the balloon as it relaxes and ages. The balloons can be blown from many different materials; Nylon pebax and polyethylene are particularly suitable polymers. The balloon tubing is inserted through the mold, and sealed at one end. A knot tied in the tubing is sufficient for sealing. The other end of the tubing is connected to a pressure source, providing pressure in the range of 80 to 350 psi. The required pressure depends on the material and dimensions of the tubing. The balloon is then heated in a localized area, while tension is optionally applied to either end of the tubing. After the tubing expands to match the inside diameter of the glass mold, the heat source is advanced along the length of the mold, at a rate that allows the tubing to grow to match the inside diameter of the mold. The balloon and mold may then be cooled. One method for cooling is blowing compressed air over the mold. The balloon is then removed from the mold. Optionally a release agent may be used to facilitate this step. Acceptable mold release agents include silicone, Polyvinyl alcohol (PVA) and Polyethylene oxide (PEO) Additionally balloons may be produced by wrapping braiding or weaving a material such as EPTFE over a mandrel to produce a shape desired the material is then bonded to itself by a process such as sintering or gluing.
With reference back to
With reference to
The balloon members 111 are then placed inside each channel formed by the cuff 102. See e.g.
Another problem with an expandable stent based valve prosthesis is that if the stent is over-expanded the valve leaflets may not coapt. This results in a central leak, and an incompetent valve. In one embodiment, the inflatable sealing cuff 176 described above is designed so that if the operator detects a central leak the operator can inflate the cuff to a high pressure causing the stent 172 to decrease in diameter at the prosthetic valves annulus. The operator monitors any regurgant flow using an imaging technique such as echocardiography. Guided by this information the cuff 176 can be inflated to the minimum pressure that eliminates the leak. Using the minimum pressure insures that the maximum possible area is available for blood flow. This technique would allow for a reduction in the initial deployed diameter or a resizing of the structure to properly fit the implantation area.
Non-Inflatable Prosthetic Aortic Valve Implants
A latch or lock mechanism 181 maintains the tension in the wire or locks the distal end to a location near the proximal end. This tension mechanism may be driven from the handle through a tension wire, a hydraulic system, a rotational member to drive a screw. Furthermore the tensioning members may utilize a locking means to maintain the desired circular shape, such as a suture, an adhesive, or a mechanical snap together type lock actuated by the tension wire.
With initial reference to
In the embodiments described above with reference to
In the illustrated embodiment, the cuff 702 contains a spiral channel 704 allowing the delivery of a wire 706, which takes a helical shape after it is inserted into the cuff. The helix extends from the proximal end of the device 700 to the distal end of the valve with the individual coils spaced close together as shown in
The preferred wire material is Nitinol, although many other metals and polymers have suitable properties. Nitinol provides an advantage that its chemistry and thermal history can be used to tune the temperature at which it undergoes a phase change. By adjusting this transition temperature to fall at a temperature just below body temperature the support structure 706 can be delivered (e.g., within the cuff 702) with one set of mechanical properties and after delivery, and after the support structure 706 has equalized in temperature with the body, the support structure 706 assumes a second set of mechanical properties (e.g., shape). Other materials that undergo a phase change near body temperature, such as other shape memory alloys may provide similar benefits.
In one embodiment, the catheter that is attached to the channels 704 in the cuff 702 is preferably in an orientation that allows the wire 706 to be delivered with minimal friction making a minimum number of excessively sharp bends. The catheter may optionally include an inflation portion to allow an inflation media to temporarily act as a support structure during the process of positioning the prosthesis 700.
Leaflet Subassembly
With reference back to the embodiments of
A number of additional advantages result from the use of the implant 100 and the cuff 102 construction utilized therein. For example, for each key area of the cuff 102, the flexibility can be optimized or customized. If desired, the coapting tissue leaflet commissures can be made more or less flexible to allow for more or less deflection to relieve stresses on the tissue at closing or to fine tune the operation of the valve. Similarly, the base radial stiffness of the overall valve 100 structure can be increased or decreased by pressure or inflation media to preserve the roundness and shape of the valve 100.
Attachment of the valve 104 to the cuff 102 can be completed in any number of conventional methods including sewing, ring or sleeve attachments, gluing, welding, interference fits, bonding through mechanical means such as pinching between members. An example of these methods are described in Published Application from Huynh et al 06102944 or Lafrance et al 2003/0027332 or Peredo U.S. Pat. No. 6,409,759, which are hereby incorporated by reference herein. These methods are generally know and accepted in the valve device industry. As mentioned above, the cuff 102 may additionally house an inflation mold where the structure is formed within the body or the cuff made be the mold where the fluid is injected to create the support structure. The valve, whether it is tissue, engineered tissue, mechanical or polymer, may be attached before packaging or in the hospital just before implantation. Some tissue valves are native valves such as pig, horse, cow or native human valves. Most of which are suspended in a fixing solution such as Glutaraldehyde.
Although mechanical heart valves with rigid pivoting occluders or leaflets have the advantage of proven durability through decades of use, they are associated with blood clotting on or around the prosthetic valve. Blood clotting can lead to acute or subacute closure of the valve or associated blood vessel. For this reason, patients with implanted mechanical heart valves remain on anticoagulants for as long as the valve remains implanted. Anticoagulants impart a 3-5% annual risk of significant bleeding and cannot be taken safely by certain individuals.
Besides mechanical heart valves, heart valve prostheses can be constructed with flexible tissue leaflets or polymer leaflets. Prosthetic tissue heart valves can be derived from, for example, porcine heart valves or manufactured from other biological material, such as bovine or equine pericardium. Biological materials in prosthetic heart valves generally have profile and surface characteristics that provide laminar, nonturbulent blood flow. Therefore, intravascular clotting is less likely to occur than with mechanical heart valve prostheses.
Natural tissue valves can be derived from an animal species, typically mammalian, such as human, bovine, porcine canine, seal or kangaroo. These tissues can be obtained from, for example, heart valves, aortic roots, aortic walls, aortic leaflets, pericardial tissue such as pericardial patches, bypass grafts, blood vessels, human umbilical tissue and the like. These natural tissues are typically soft tissues, and generally include collagen containing material. The tissue can be living tissue, decellularized tissue or recellularized tissue.
Tissue can be fixed by crosslinking. Fixation provides mechanical stabilization, for example by preventing enzymatic degradation of the tissue. Glutaraldehyde or formaldehyde is typically used for fixation, but other fixatives can be used, such as other difunctional aldehydes, epoxides, genipin and derivatives thereof. Tissue can be used in either crosslinked or uncrosslinked form, depending on the type of tissue, use and other factors. Generally, if xenograft tissue is used, the tissue is crosslinked and/or decellularized.
The implants 100 can further include synthetic materials, such as polymers and ceramics. Appropriate ceramics include, for example, hydroxyapatite, alumina, graphite and pyrolytic carbon. Appropriate synthetic materials include hydrogels and other synthetic materials that cannot withstand severe dehydration. Heart valve prostheses can include synthetic polymers as well as purified biological polymers. These synthetic polymers can be woven or knitted into a mesh to form a matrix or similar structure. Alternatively, the synthetic polymer materials can be molded or cast into appropriate forms.
Appropriate synthetic polymers include without limitation polyamides (e.g., nylon), polyesters, polystyrenes, polyacrylates, vinyl polymers (e.g., polyethylene, polytetrafluoroethylene, polypropylene and polyvinyl chloride), polycarbonates, polyurethanes, poly dimethyl siloxanes, cellulose acetates, polymethyl methacrylates, ethylene vinyl acetates, polysulfones, nitrocelluloses and similar copolymers. Bioresorbable polymers can also be used such as dextran, hydroxyethyl starch, gelatin, derivatives of gelatin, polyvinylpyrolidone, polyvinyl alcohol, poly[N-(2-hydroxypropyl)methacrylamide], poly (hydroxy acids), poly(epsilon-caprolactone), polylactic acid, polyglycolic acid, poly(dimethyl glycolic acid), poly(hydroxy buterate), and similar copolymers. These synthetic polymeric materials can be woven or knitted into a mesh to form a matrix or substrate. Alternatively, the synthetic polymer materials can be molded or cast into appropriate forms.
Biological polymers can be naturally occurring or produced in vitro by fermentation and the like or by recombinant genetic engineering. Recombinant DNA technology can be used to engineer virtually any polypeptide sequence and then amplify and express the protein in either bacterial or mammalian cells. Purified biological polymers can be appropriately formed into a substrate by techniques such as weaving, knitting, casting, molding, extrusion, cellular alignment and magnetic alignment. Suitable biological polymers include, without limitation, collagen, elastin, silk, keratin, gelatin, polyamino acids, polysaccharides (e.g., cellulose and starch) and copolymers thereof.
A tissue-based valve prosthesis can maintain structural elements, such as leaflets, from its native form and/or structural elements can be incorporated into the prosthesis from the assembly of distinct pieces of tissue. For example, the valve prosthesis can be assembled from a porcine heart valve, from bovine pericardium or from a combination thereof. Porcine tissue valves, for example, the Toronto SPV® valve marketed by St. Jude Medical, Inc. St. Paul, Minn., can be implanted in the patient using the tools described herein. The Toronto SPV® valve is designed for implantation in an aortic heart valve position. See, for example, David et al., J. Heart Valve Dis. 1:244-248 (1992). It will be appreciated by those skilled in the art that the tools of the present invention are applicable to any valve, especially any tissue valve prosthesis, that is adapted for implanting in a patient.
A reinforcement may be placed along the inner surface of the valve commissure supports and/or scallops. In alternative embodiments, the reinforcement is placed on the outer surface of the valve, such as at the valve commissure supports. The reinforcement preferably includes apertures through which the fasteners extend or can be inserted. The reinforcements are thin strips of relatively strong material. The reinforcement can prevent or reduce damage to the prosthesis when the fasteners are inserted and after implantation of the heart valve prosthesis in the patient. The reinforcement, thus, can protect and support the commissure supports from potential damage generated by the presence of the fasteners. In alternative embodiments, the reinforcement is placed on the outside of the aorta such that the fastener pierces the reinforcement after passing through the prosthetic valve.
Tissue valves whether implanted surgically or percutaneously have a risk of calcification after implantation. To prevent or minimize the calcification several treatments have been employed before the tissue is fixed. Some strategies include treating the valves with ethanol, metallic salts, detergents, biophosphonates, coimplants of polymeric controlled release drug delivery systems, and covalent attachment of anticalcifying agents. In the preferred embodiment the valve tissue is treated in 40% to 80% ethanol for 20 to 200 hours before fixation in a buffered glutaraldehyde solution. The ethanol pretreatment may prevent calcification in the valve after implantation and serves to remove cholesterol and phospholipids from the tissue before fixation. (ref Prevention of Bioprosthetic Heart Valve Calcification by Ethanol Preincubation, Vyavahare et al)
Inflation MediaThe inflatable structure 107 can be inflated using any of a variety of inflation media 122, depending upon the desired performance. In general, the inflation media can include a liquid such water or an aqueous based solution, a gas such as CO2, or a hardenable media which may be introduced into the cuff 102 at a first, relatively low viscosity and converted to a second, relatively high viscosity. Viscosity enhancement may be accomplished through any of a variety of known UV initiated or catalyst initiated polymerization reactions, or other chemical systems known in the art. The end point of the viscosity enhancing process may result in a hardness anywhere from a gel to a rigid structure, depending upon the desired performance and durability.
Useful inflation media generally include those formed by the mixing of multiple components and that have a cure time ranging from a few minutes to tens of minutes, preferably from about three and about twenty minutes. Such a material should be biocompatible, exhibit long-term stability (preferably on the order of at least ten years in vivo), pose as little an embolic risk as possible, and exhibit adequate mechanical properties, both pre and post-cure, suitable for service in the cuff of the present invention in vivo. For instance, such a material should have a relatively low viscosity before solidification or curing to facilitate the cuff and channel fill process. A desirable post-cure elastic modulus of such an inflation medium is from about 50 to about 400 psi-balancing the need for the filled body to form an adequate seal in vivo while maintaining clinically relevant kink resistance of the cuff. The inflation media ideally should be radiopaque, both acute and chronic, although this is not absolutely necessary.
Details of compositions suitable for use as an inflation medium in the present invention are described in greater detail in U.S. patent application Ser. No. 09/496,231 to Hubbell et al., filed Feb. 1, 2000 and entitled “Biomaterials Formed by Nucleophilic Addition Reaction to Conjugated Unsaturated Groups” and U.S. patent application Ser. No. 09/586,937 to Hubbell et al., filed Jun. 2, 2000 and entitled “Conjugate Addition Reactions for the Controlled Delivery of Pharmaceutically Active Compounds”. The entirety of each of these patent applications is hereby incorporated herein by reference.
Below is listed one particular three-component medium.
This medium comprises:
(1) polyethylene glycol diacrylate (PEGDA), present in a proportion ranging from about 50 to about 55 weight percent; specifically in a proportion of about 52 weight percent,
(2) pentaerthyritol tetra 3 (mercaptopropionate) (QT) present in a proportion ranging from about 22 to about 27 weight percent; specifically in a proportion of about 24 weight percent, and
(3) glycylglycine buffer present in a proportion ranging from about 22 to about 27 weight percent; specifically in a proportion of about 24 weight percent.
Variations of these components and other formulations as described in copending U.S. patent application Ser. Nos. 09/496,231 and 09/586,937, both to Hubbell et al., may be used as appropriate. In addition, we have found PEGDA having a molecular weight ranging from about 350 to about 850 to be useful; PEGDA having a molecular weight ranging from about 440 to about 560 are particularly useful.
Radiopaque materials as previously discussed may be added to this 3-component system. We have found that adding radiopacifiers such as barium sulfate, tantalum powder, and soluble materials such as iodine compounds to the glycylglycine buffer is useful.
Applicants have found that triethanolamine in phosphate-buffered saline may be used as an alternative to glycylglycine buffer as the third component described above to form an alternative curable gel suitable for use in embodiments of the present invention.
An alternative to these three-component systems is a gel made via polymer precipitation from biocompatible solvents. Examples of such suitable polymers include ethylene vinyl alcohol and cellulose acetate. Examples of such suitable biocompatible solvents include dimethylsulfoxide (DMSO), n-methyl pyrrolidone (NMP) and others. Such polymers and solvents may be used in various combinations as appropriate.
Alternatively, various siloxanes may be used as inflation gels. Examples include hydrophilic siloxanes and polyvinyl siloxanes (such as STAR-VPS from Danville Materials of San Ramon, Calif. and various silicone products such as those manufactured by NuSil, Inc. of Santa Barbara, Calif.).
Other gel systems useful as an inflation medium or material for the present invention include phase change systems that gel upon heating or cooling from their initial liquid or thixotropic state. For example, materials such as n-isopropyl-polyacrylimide (NIPAM), BASF F-127 pluronic polyoxyamer, and polyethylene glycol (PEG) chemistries having molecular weights ranging between about 500 and about 1,200 are suitable.
Effective gels may also comprise thixotropic materials that undergo sufficient shear-thinning so that they may be readily injected through a conduit such as a delivery catheter but yet still are able to become substantially gel-like at zero or low shear rates when present in the various channels and cuffs of the present invention.
In the case of the three-component PEDGA-QT-glycylglycine formulation described above, a careful preparation and delivery protocol should be followed to ensure proper mixing, delivery, and ultimately clinical efficacy. Each of the three components is typically packaged separately in sterile containers such as syringes until the appropriate time for deploying the device. The QT and buffer (typically glycylglycine) are first continuously and thoroughly mixed, typically between their respective syringes for approximately two minutes. PEGDA is then mixed thoroughly with the resulting two-component mixture for approximately three minutes. This resulting three-component mixture is then ready for introduction into the cuff as it will cure into a gel having the desired properties within the next several minutes. Cure times may be tailored by adjusting the formulations, mixing protocol, and other variables according to the requirements of the clinical setting. Details of suitable delivery protocols for these materials are discussed in U.S. patent application Ser. No. 09/917,371 to Chobotov et al.
The post-cure mechanical properties of these gels may be highly tailorable without significant changes to the formulation. For instance, these gels may exhibit moduli of elasticity ranging from tens of psi to several hundred psi; the formulation described above exhibits moduli ranging from about 175 to about 250 psi with an elongation to failure ranging from about 30 to about 50 percent.
It may be helpful to add an inert biocompatible material to the inflation material. In particular, adding a fluid such as saline to the PEGDA-QT-glycylglycine formulation (typically after it has been mixed but before significant curing takes place) lowers the viscosity of the formulation and results in greater ease when injecting the formulation into cuffs and channels without sacrificing the desired physical, chemical, and mechanical properties of the formulation or its clinical efficacy. In the appropriate volume percentages, adding materials such as saline may also reduce the potential for the inflation material such as PEGDA-QT-glycylglycine to pose an embolic risk in case of spillage or leakage. Saline concentrations as a volume percentage of the final saline/three-component formulation combination may range from zero to as high as sixty percent or more; particularly suitable are saline concentrations ranging from about twenty to about forty percent. A saline volume concentration of about thirty percent to be most suitable. Alternatives to saline may include biocompatible liquids, including buffers such as glycylglycine.
In more general terms, it is desirable to use an inflation medium in which each of its components is biocompatible and soluble in blood. A biocompatible inflation medium is desirable so to manage any toxicity risk in the case the inflation medium were inadvertently released into the patient's vasculature. A soluble inflation medium is desirable so to manage any embolism risk if released into the vasculature. Such an inflation medium should not disperse nor gel or solidify if spilled into flowing blood before curing. In the event of a spill, the normal blood flow would then rapidly disperse the components and their concentration would fall below the level required for crosslinking and formation of a solid. These components would then be eliminated by the body through standard pathways without posing an embolic risk to the patient. Among the many possibilities of an inflation medium example in which all of the components are soluble in blood is the combination polyethylene glycol diacrylate, a thiolated polyethyleneamine, and a buffer.
As previously discussed, more than one type of inflation medium, or more than one variant of a single type of inflation medium may be used in a single graft to optimize the graft properties in the region in which it is disposed.
For example, in the cuffs 102 of the various embodiments of the present invention, the inflation material serves as a conformable sealing medium to provide a seal against the lumen wall. Desirable mechanical characteristics for the inflation medium in the proximal and distal cuffs would therefore include a low shear strength so to enable the cuff to deform around any luminal irregularities (such as calcified plaque asperities) and to conform to the luminal profile, as well as a high volumetric compressibility to allow the fill material to expand the cuffs as needed to accommodate any late lumen dilatation and maintain a seal.
Another inflation media that has proven especially useful is an epoxy based two part inflation media, where one part contains the reaction product of epichlorohydrin and bisphenol A, and Butaneddiol diglyceridyl ether. And where one part contains 2,2,4-trimethyl-1, 6-hexanediamine. Whereas the material may have a viscosity of about 100-200 cPs (@ 100 rpm/23 C) but most preferably they may be readily injected through a small lumen to be introduced to the implant from outside the body. The operating temperature range may be from about −55 to about +125 C but would be most advantageous at the body temperature of +37 C. Other properties may include a hardness of about 81 on the Shore D scale and a lap shear strength of 1,700 PSI. An example of this would be EPO-TEK 301 supplied by 14 Fortune Drive Billerica, Mass.
The mixed uncured inflation media preferably has a viscosity less than 2000 cps In one embodiment the epoxy based inflation media has a viscosity of 100-200 cps. In another embodiment the inflation media has a viscosity less than 1000 cps.
In one embodiment the inflation media contains a foaming agent. The foaming inflation media is beneficial because the foaming action can generate pressure within the inflatable portion of the device. Therefore less inflation media needs to be injected. Additionally any pressure loss from the disconnection process is compensated for by the foaming action of the inflation media. Many appropriate foaming medias are possible; one example is a urethane foam.
In another embodiment the balloon or inflation channel may be connected to the catheter on both ends. This allows the balloon to be preinflated with a nonsolidifying material such as a gas or liquid. If a gas is chosen CO2 or helium are likely choices, these gasses are used to inflate intraortic balloon pumps. Preferably the preinflation media is radiopaque so that the balloon position can be determined by angiography. Contrast media typically used in interventional cardiology could be used to add sufficient radiopacity to most liquid preinflation medias. When it is desired to make the implant permanent and exchange the preinflation media for the permanent inflation media, the permanent inflation media is injected into the inflation channel through a first catheter connection. As the permanent inflation media is injected the preinflation media is expelled out a second catheter connection. The catheter connections are positioned in such a way that substantially all of the preinflation media is expelled as the permanent inflation media is injected. In one embodiment an intermediate inflation media is used to prevent entrapment of preinflation media in the permanent inflation media. In one embodiment the intermediate inflation media is a gas and the preinflation media is a liquid. In another embodiment the intermediate inflation media or preinflation media functions as a primer to aid the permanent inflation media to bond to the inner surface of the inflation channel. In another embodiment the preinflation media or the intermediate inflation media serves as a release agent to prevent the permanent inflation media from bonding to the inner surface of the inflation channel.
The permanent inflation media may have a different radiopacity than the preinflation media. A device that is excessively radiopaque tends to obscure other nearby features under angiography. During the preinflation step it may be desirable to visualize the inflation channel clearly, so a very radiopaque inflation media may be chosen. After the device is inflated with the permanent inflation media a less radiopaque inflation media may be preferred. The feature of lesser radiopacity is beneficial for visualization of proper valve function as contrast media is injected into the ventricle or the aorta.
Anchoring Mechanisms
In the embodiments described above, it may be necessary or desirable to incorporate an anchoring mechanism 220 into the cuff 102. The anchoring mechanism 220 can comprise any of a variety of anchors or barbs such as those that have been used extensively on interventional devices, such as grafts for the treatment of abdominal aortic aneurysms, atrial appendage closure devices and filters. Most of the traditional retention mechanisms used for percutaneously implantable valves rely on an interference fit between the implant and the vessel to provide a significant portion of the retention force, or to activate the retention means. However, in the case of a replacement mitral or aortic valve, it can be desirable to minimize the radial force at the valve annulus, because excessive dilation of either annulus may have a detrimental effect on the function of another other valve.
With reference to
In another embodiment, the valve 100 is sutured to the native anatomy. For example, the valve 100 can include a sewing ring configured allow sutures to be easily attached to the implant 100. A percutaneous or minimally invasive sewing device can also be incorporated or used as a secondary procedure. This device would contain at least one needle remotely actuated to attach the valve 100 to the tissue, or to a second device previously implanted at the desired valve location. Other methods may utilize a balloon or other force mechanism to push or pull the suture into position. These needles can be made from metallic or polymer elements or utilize sutures that may be inserted through the anatomy. They would range in diameters from 0.002 inches to about 0.040 inches and may protrude into the anatomy from 0.005 inches to about 0.090 inches depending upon the anatomy.
With reference to
In one embodiment wires similar to, the control wires 230 described in this application serve as guide wires over which the secondary anchoring catheter is delivered. This allows the precise placement of the anchors, staples, sutures etc. relative to the prosthesis, because the anchor catheter will follow the wire right to the desired anchor location. In one embodiment the anchor location is at the valve commisures. In another embodiment the anchor location is at the proximal end of the device. The anchor delivery catheter may consist of a multi lumen tube where one lumen serves to track over the wire and the second lumen or additional lumens deliver the anchor. In one embodiment the anchor is a screw which is actuated with a rotational motion and threaeded through the prosthesis and in to the aortic wall. Other anchor designs described in this application may also be adapted to the anchor delivery catheter.
In another embodiment, an adhesive is used to secure the valve 100 to the tissue. For example, adhesives such as a fibrin glue or cyanoacrylate could be delivered percutaneously or surgically to attach the valve 100 to the tissue. A method for percutaneously delivering an adhesive includes channeling it through a tubular support member, which has openings around its outer surface to allow the adhesive to be released. The adhesive could be used in conjunction with other anchoring methods to ensure that no blood leaks around the valve 100. Adhesion enhancing surfaces can be provided, such as ePTFE patches or jackets, to promote cellular in-growth for long term anchoring.
With reference to
In the embodiment of
In another embodiment, the distal and proximal ends 128, 126 of the implant 100 can be sized to provide an anchor functions For example, as described above with reference to
For an implant 100 that utilizes an hourglass shape as described above, the orientation of the anchoring mechanisms 220 described above can be adapted from radially expandable applications can be reevaluated and reapplied. For example barbs could be placed on the most distal portion 128 of the hourglass shaped structure and the barbs would preferably be oriented approximately parallel to the axial direction. See e.g., Figure
In this embodiment, the hook 266 can be cut from a hypotube 260 of slightly larger inside diameter than the deployment control wire 230 outside diameter. Preferably these diameters are in the range of 0.01 to 0.03 inch. The hook 266 preferably extends from the device at an angle of 10 to 80 degrees, more preferably at an angle of 20 to 45 degrees.
Delivery Catheter
With initial reference to
In one embodiment, the outer diameter of the catheter 300 measures generally about 0.030 inches to 0.200 inches with a wall thickness of the outer tubular member 301 being about 0.005 inches to about 0.060 inches. In another embodiment, the outer diameter ranges from about 0.15 inches to about 0.35 inches or from about 12 French to about 27 French. In this embodiment, the wall thickness of the outer tube 301 is between about 0.005 inches and about 0.030 inches. The overall length of the catheter 300 ranges from about 80 centimeters to about 320 centimeters.
As mentioned above, the catheter 300 includes a connection hub or handle 306 that is configured to allow wires, devices and fluid to pass as will be explained in more detail below. The connection hub 306 is preferably compatible with normal cath-lab components and can utilize a threaded end and a taper fit to maintain seal integrity. The inner diameter of the inner member 305 of the catheter 300 is configured allow for coaxial use to pass items such as guidewires, devices, contrast and other catheters. An inner lining material such as Teflon may be used to reduce friction and improve performance in tortuous curves. Additionally, slippery coatings such as DOW 360, MDX silicone or a hydrophilic coating from BSI Corporation may be added to provide another form of friction reducing elements.
Multidurometer materials in the catheter 300 can help to soften the transition zones and add correct stiffness for pushability. Transition zones may also be achieved through an extrusion process know as bump tubing, where the material inner and outer diameter change during the extrusion process. The entire catheter shafts 301, 305 can be produced in one piece. Another method for producing such a catheter shaft is to bond separate pieces of tubing together by melting or gluing the two components together and forming a single tube with multiple diameters and or stiffness. The application of heat can be applied by laser or heated air that flows over the shaft material or other methods of heat application sufficient to flow the materials together.
With continued reference to
With continued reference to
The inner member 305 material may also consist of stiffening members for transition zones or bump extrusions to reduced diameter and maintain correct pushability. Conventional guidewire passage through the catheter such as “over-the-wire” may be used or technology such as “rapid-exchange” may aid in procedure ease and catheter exchanges. Since multiple devices may be placed in a single catheterization, rapid-exchange may be preferred but not essential. Other features that may aid in ease of use include a slippery coating on the outer and or inner diameter such as mineral oil, MDX (silicone) or a hydrophilic layer to allow easy access to tortuous anatomy, or easier more controlled motion of one portion of the catheter relative to another portion of the catheter. It may be necessary or desirable to utilize a balloon to initiate radial contact of the device to its final position and location. In one embodiment, an inflation lumen and balloon placed distal to the hubis used. This balloon is used to pre-dilate the native valve annulus, vessel or ostium where the valve may be implanted. Elements to transmit signals externally could be imbedded into the catheter 300 for pressure and flow readings or Doppler information. These may include electro-mechanical sensors, such as piezo-electric devices, electrical sensors, wires, pressure portal or lumens or optical fibers.
As mentioned above, delivery of the implant 100 via catheterization of the implantation site can include a mechanism to deploy or expel the implant 100 into the vessel. This mechanism may include a push or pull member to transmit forces to the distal portion of the catheter 300. These forces may be applied externally to the body and utilize a handle at the proximal end of the catheter. Devices to transmit forces to the distal end may also include a rotational member to loosen or tighten, convert a torque into a translational force such as a threaded screw and nut or to add or subtract stiffness to the catheter or device, or to cause the device to assume a specific shape. The handle mechanism may also include a port for hydraulic pressures to be transmitted to the distal portion of the catheter or have the ability to generate hydraulic forces directly with the handle. These forces may include a pushing or pulling transmitted to the device or catheter, an exposure of the device to allow for implantation or to expel the device from the catheter. Further forces may include a radial or longitudinal expansion of the device or catheter to implant or size the location of implantation. The handle may also include connections to electrical signals to monitor information such as pressures, flow rates, temperature and Doppler information.
With reference to
With continued reference to
As will be described in more detail below, the control wires 230 and/or inflation lumen 318 can form part of a deployment mechanism for the implant 100. As the implant is navigated to the site, attachment between the implant 100 and catheter 300 is important. Many detachment mechanisms have been used to deploy devices such as stents and embolic coils through balloon expansion and simple pushable coils expelled from the distal end of a catheter. The implant 100 can utilize many different methods to implant 100 at the selected site such as an expulsion out the end of the catheter, a mechanical release mechanism such as a pin joint, unscrewing the device from the catheter delivery system, a tethered link such as a thread or wire, a fusible link as used in a GDC coil deployment, a cutting tool to sever a attachment of the device from the catheter, a threaded knot to tether the catheter to the device where the as the knot could be untied or cut, a hydraulic mechanism to deploy, expand or fracture a link between the catheter and the device. All above mentioned concepts can be enhanced by the utilization of the flexible tip 312 to allow acute articulation of the device and delivery catheter 300 to gain access to the implantation site.
As will be explained in more detail below, after the implant 100 has been temporarily deployed or positioned, it may be advantageous to recapture or reposition the implant for optimal results. This may include a rotation or translation of the implant 100 or a complete removal and exchange for a different diameter, length or style device. Capture of an implanted device may require a second catheter to reengage the device to remove or reposition to a proper location. This catheter may be constructed from polymer tubing as described above including coils, braids, etc. Additionally there may be a braided section at the distal most portion of the catheter to accept or capture the device for retrieval from the body.
As mentioned above, the guidewire tube 320 preferably extends through the inner sheath 305 and the tip 312. The guidewire tube 320 may have an inside diameter of 0.035 to 0.042 in so that the device is compatible with common 0.035 or 0.038 guide wires. A modified embodiment includes a lumen 0.014 to 0.017 inches in diameter for compatibility with 0.014 in diameter guide wires. In a third embodiment, the guidewire lumen 320 is 0.039 to 0.080 in diameter, so that the device may be delivered over a larger than standard guide wire, or a diagnostic catheter, such as a pig tail catheter. This provides the advantage of a stiffer support to facilitate easier delivery through calcified valves. If a diagnostic catheter is used as a guidewire it may also serve as a port for contrast injection.
The guidewire tube 320 can be made from a lubricious material such as Teflon, polypropolene or a polymer impregnated with Teflon. It may also be coated with a lubricious or hydrophilic coating. The tube 320 can be constructed of multiple layers of material, including a lubricious inner layer and an outer layer to facilitate bonding to other catheter components.
The catheter 300 may be delivered over a guide wire to aid in positioning. The guide wire may pass coaxially through the entire length of the catheter or in modified embodiments may pass coaxially though only a portion of the catheter in a configuration known as rapid exchange. This allows shorter guide wires to be used if devices are to be exchanged out.
In the illustrated embodiment, the catheter 300 comprises the outer catheter shaft 301 and the inner catheter shaft 305 which move relative to one another. In order to minimize the risk of guidewire damage in a rapid exchange design where the catheter must pass through the wall of two sheaths which move relative to one another, a slot feature is desirable. Either the inner or outer elongate tube may contain a longitudinal slot in the area where the guide wire passes from the inner diameter to the outer diameter of the catheter assembly. The other elongate tube preferably contains a hollow pin to engage the slot and prevent the excessive movement of the two elongate members. The guide wire passes through the opening in the hollow pin. The inner diameter of the hollow pin is preferably oriented at an acute angle to the central axis of the catheter.
Another design to enable rapid exchange like performance is for the guide wire to enter the catheter tip through a side hole distal to the location of the prosthetic valve. The guidewire exits the tip of the system near the center of the catheter tip. This design enables the catheter to follow the guide wire across the native valve, while still allowing multiple devices to be exchanged easily on a short length guide wire.
As described above, the internal lumens of the catheter 300 can include the deployment control wires lumens 316, the inflation lumens 320, and an inner sheath 307 that encapsulates these lumens 316, 320. See e.g.,
This catheter configuration advantageously allows the operator to easily switch between the delivery sheath 300 and a recovery sheath (described herein) in the event that the device 100 needs to be recovered, because the delivery sheath 300 can be retracted out of the body over relatively short internal lumens 316, 320, while still maintaining a portion of the lumens 316, 320 outside the catheter so that the operator can manipulate them as necessary.
Because of its shorter length the recovery sheath may not require the exchange hole 650, and it may be possible to locate the internal lumens coaxially within the recovery sheath. However in the preferred embodiment the recovery sheath also includes a hole in a similar location allowing the internal lumens to pass coaxially through the distal portion of the sheath, through the hole, and be located generally parallel to the recovery sheath in the proximal portion.
In one embodiment contrast media is passed through a lumen (e.g., the guidewire tube 320) of the device, and the lumen passes through the prosthetic valve 100. This allows visual evaluation of valve function by angiography, without crossing the valve with an additional device. In the preferred embodiment the lumen crosses the valve while the valve is in the catheter. In the preferred embodiment the lumen also serves as the guidewire tube 320, where the device is delivered over a guide wire. The wire may be removed from the lumen to allow more cross sectional area for contrast injection. The proximal end of the lumen near the handle of the device attaches to a fitting to allow the injection of contrast media with a power injector tool. The inner diameter of the lumen may range from 0.014 to 0.100 inch. The diameter of the lumen may vary along the length of the catheter, for example, Preferably the portion of the lumen which passes through the prosthetic valve is of a minimum possible diameter to allow both sufficient flow and the use of an adequate sized guidewire. This portion is preferably in the range of diameters from 0.014 to 0.080. The portion of the lumen extending along the length of the catheter proximal to the implant may be of larger diameter, the larger diameter allows flow of contrast media at lower pressure gradients, and the corresponding larger outside diameter does not increase the profile of the complete device. This portion of the lumen is preferably in the inside diameter range of 0.035 to 0.100 in. The distal portion of the lumen may contain a diffuser or transition to a larger diameter to minimize the pressure required to inject a sufficient volume of contrast media through the lumen. Multiple exit ports positioned around a nose cone also facilitate the flow of contrast media.
Access for the catheter 300 may be gained through a major artery such as the femoral artery. This access site is particularly appropriate for aortic valve replacement. Alternative access methods may be better suited for other valves. For example the tricuspid valve and possibly the pulmonary valve could best be accessed through the venous system. In this case, access would be gained through either a femoral vein or a jugular vein. The catheter would then be passed into the right atrium through the superior or inferior vena cava. Some embodiment of the current invention utilize a relatively large diameter catheter, which may not be compatable with the diameter of all patients femoral arteries. In these patients it may be desirable to access the common iliac artery or to use a transeptal approach and access the heart through the venous system.
As mentioned above, the catheter 300 includes an atraumatic tip 312 to allow the device to be easily placed through the hemostasis valve of the introducer, and to easily cross the calcified aortic valve. The tip 312 may be cone shaped bullet shaped or hemispherical on the front end. The largest diameter of the tip 312 is preferably approximately the same as the distal portion 309 of the outer sheath 301. The tip 312 preferably steps down to a diameter slightly smaller than the inside diameter of the distal portion 309 of the outer sheath 301, so that the tip can engage the outer sheath 301 and provide a smooth transition. In the illustrated embodiment, the tip 312 is connected to the guide wire tube 320, and the guide wire lumen passes through a portion of the tip 312. The proximal side of the tip 312 also has a cone, bullet or hemispherical shape, so that the tip can easily be retraced back across the deployed valve 100, and into the deployment catheter 300. The tip 312 can be manufactured from a rigid polymer such as polycarbonate, or from a lower durometer material that allows flexibility, such as silicone. Alternatively, the tip 312 may be made from multiple materials with different durometers. For example, the portion of the tip 312 that engages the distal portion 309 of the outer sheath 301 can be manufactured from a rigid material, while the distal and or proximal ends of the tip are manufactured from a lower durometer material.
With reference to
The retractable sheath 340 of
As shown in
As shown in
With reference back to
Connection Between Implant and Inflation Lumens
As described above, in many embodiments, the implant 100 includes an inflatable structure 107, which defines inflation channels 120. In these embodiments, the inflation channels 120 are inflated with inflation media 122 to provide structure to the implant 100. As shown in
In general, in embodiments in which the inflation media 122 is not self sealing the inflation channels 122 will need to be sealed as the inflation lumen 318 is disconnected from the implant 100. Sealing of these lumens could utilize many different techniques known to one skilled in the art. For example, as explained below, the inflation lumen can be placed through a valve, in such a way that it forces the valve into the open position. The valve could be one of a variety of normally closed or one way (check) valves.
For example,
In some embodiments, it is advantageous to configure the deployment catheter 300 and the implant 100 such that the inflation tube 318 cannot disconnected unintentionally. For example, in one embodiment, the inflation tube 318 is connected to a deployment control wire 230 so that the inflation lumen 218 can not be removed from the implant 100 unless the deployment control wire 230 is also disconnected from the implant 100.
After the balloon 111 has been inflated with the desired inflation media and the operator has chosen to disconnect the catheter 300 from the implant 100, the inflation tube 318 is withdrawn past the duckbill valve 430. At this time suction may be applied to remove as much inflation material as possible from the area past the valve 430. A rinse procedure could also be used to remove additional fluid. The inflation tube 318 is then withdrawn past the enlarged ridge 436 and the slit portion of the lock tubing 434. The lock tubing 434 can then be withdrawn from the orifice 440, and the implant 100 is separated from the catheter 300.
Control Wires
As discussed previously above, one advantage of many of the embodiments described herein is that the deployment of the implant 100 can be controlled. In one embodiment, the deployment of the implant is controlled via the use of control wires 230 that can be detachably coupled to the implant. Various mechanisms for detachably coupling the control wires 230 to the implant 10 will now be described.
With initial reference to
With continued reference to
Preferably, three wires 230 are used, but any number between 1 and 10 can provide good results. The diameter of the wire 230 can range from about 0.002 inches to 0.020 inches. The wires 230 can be manufactured from a metal suitable for blood contact such as nitinol, stainless steel or one of many cobalt chrome nickel and/or iron based alloys. The wires 230 can also be made of a polymer that has the desired mechanical properties such as a polyimide. The sheath 472 can be manufactured from the many polymers suitable for blood contact including nylons Teflon PBX polyethylene polypropylene polyimides etc. The sheath 470 is preferably sufficiently rigid in the axial direction to prevent the accidental disconnection of the valve 100, so the dimensions of the sheath depend on the axial stiffness of the material. A polyimide sheath 470 with a 0.026 inches outside diameter and a 0.005 inch thick single wall has proven adequate, while a grillamid nylon sheath with a 0.030 inch outside diameter and a 0.007 inch thick single wall has also proven adequate. Preferably the polymer sheath 470 ranges in outside diameter from about 0.018 inches to 0.040 inches and in wall thickness from about 0.003 inches to about 0.010 inches. Additionally a stainless steel, nitinol or other metallic sheath cab be utilized. In this case, smaller diameters and thinner wall thicknesses are generally desirable. In one embodiment, the stainless steel sheath 470 has a 0.014 outer diameter with an inner diameter of about 0.011 inch and the wire 230 with a 0.009 inches outer diameter. With a metallic sheath 470, the preferred wall is about 0.0005 inch to about 0.0050 inch thick and the preferred outside diameter is about 0.007 inches to about 0.025 inches. The inside diameter of the sheath 470 should provide clearance to move freely over the wire 230. A clearance of 0.00 to about 0.007 inches should provide adequately free motion. A lubricant or hydrophilic coating may be applied to the inside diameter of the sheath 470, or the outside diameter of the wire 230. Different clearances may be required with less lubricious polymers. In addition, extrusion parameters may be adjusted to produce a surface finish on the inner diameter of the tube 470 that optimizes the motion of the sheath 470 relative to the wire 230. With some polymers a rougher surface may result in reduced friction. As mentioned above, the ideal wall thickness of the sheath 470 depends on the strength and stiffness of the particular material selected, but likely ranges between 0.002 and 0.020 inches, single wall thickness.
The proximal end of the deployment control wires 230 preferably contains a lock mechanism (not shown) to prevent the unintended relative motion of the wire relative to the sheath 470. The wires 230 may also be attached to a handle section that allows the relative movement of one wire individually or multiple wires together. In one embodiment the three wires 230 are attached to a ring, equally spaced around the edge of the ring. As the ring is moved proximal or distal relative to the main handle component the implant 100 moves proximal or distal relative to the catheter tip. As the ring is tilted off axis with the axis of the catheter handle, the implant 100 is tilted in a similar direction.
The deployment control mechanism can performs several functions. First as described above, during the initial deployment of the implant 100, it prevents the implant 100 from rotating off axis. Additionally the deployment control mechanism allows the implant 100 to be repositions after it has been removed from the sheath. The wires described above could be used to move the implant 100 proximally and distally.
With reference to
With reference to
The deployment control systems described herein could be used with the cast in place support structure described in this application, or on a self expanding stent structure, or on a inflatable structure as described by herein. The deployment control device may also be used on other non-vascular devices such as stent grafts for aneurysm exclusion or self-expanding stents for treating stenosis.
As described above, the deployment control wires 230 can be used to allow the repositioning of the implant 100 after it has been unsheathed. The deployment control wires 230 are preferably rigid enough to allow the operator to reposition the implant 100 and to prevent the implant 100 from migrating due to the force of blood flow and pressure. Once the implant 100 is inflated, it is desirable for the wires 230 to be flexible and, in one embodiment, as flexible as the tip of a conventional guidewire. This flexibility allows the implant 100 to take the same shape and position that it will take after the wires 230 are removed. This allows both the securement and function of the implant 100 to be tested and evaluated before the operator commits to permanently implanting the implant 100. The increased flexibility is preferably provided in a plane tangent to the generally cylindrical shape defined by the vessel, where the valve 100 is implanted. Therefore, in a preferred embodiment, the control wires 230 will be particularly flexible at the tips allowing the device to be nearly free from forces exerted by the catheter 300, as it would be when disconnected.
Many embodiments of a wire that fulfills the requirements of flexibility and stiffness are possible. In one embodiment, the wires are manufactured to have a flexible tip and a less flexible proximal section. Techniques for manufacturing wires with these properties are widely known to those skilled in the art of guide wire design and manufacture. Techniques include grinding a tapered control wire as shown in
In another embodiment, which is illustrated in
With reference to
Recovery Tools and Techniques
Current valve systems are often deployed through a stent-based mechanism where the valve is sewn to the support structure. In the inflated embodiments described herein, the structure is added to the implant secondarily via the inflation fluid. This allows the user to inflate or pressurize the implant with any number of media including one that will solidify. As such, if the operator desires, the implant 100 can be moved before the inflation media is solidified or depressurization can allow for movement of the implant within the body. Since catheter based devices tend to be small in diameter to reduce trauma to the vessel and allow for easier access to entry, it often difficult to remove devices such as stents once they have been exposed or introduced into the vasculature. However, as will be explained below, a device described herein enables a percutaneous aortic valve to be recovered from the body and reintroduced retrograde to the introducer.
A hemostasis valve (not shown) is preferably attached to the proximal end of the device 500. Also at the proximal end, a flush port and stop-cock can be provided for fluid introduction. In one embodiment, the inner shaft 504 would have a length of about 40 to 60 centimeters and a diameter of about 2 to 18 millimeters. In a modified embodiment, the distal end 508 of the braid section 506 could be attached to end of the outer coaxial sheath 502. This would allow relative motion between the two sheaths 502, 504 and allow the braided section 506 to be inverted upon it self. The braided section 506 can be formed or shaped into a funnel as shown in
In another embodiment the slit tubing is replaced by a fabric cone, where the fabric cone may contain a feature such as a preshaped wire or a balloon to facilitate its opening.
The braided cone 506 can be formed by heat setting or other manners into a cone shape with a free diameter slightly larger than the patients aorta. In another embodiment, the braided cone is manufactured from loops of wire so that the cut ends of the wire are all located at the proximal end of the cone. The wires used to manufacture the cone preferably have a diameter from 0.002 in to 0.020 in. The wires may also be replaced by ribbons having a thickness between 0.002 in and 0.020 in and a width between 0.003 in and 0.030 in. The diameter of the small end of the cone is preferably between 0.007 in and 0.3 in the cone is preferably be capable of collapsing to a diameter small enough to pass through the desired introducer size. The large end of the cone section preferably expands to a diameter similar to or slightly larger than the typical human aorta, or 0.75 in to 1.50 in.
In one embodiment, the separate recovery device 500 is supplied to facilitate the recapture of the implant in the event that the prosthetic valve did not produce the desired result in the patient. To recapture an inflatable aortic implant 100 as describe herein, the delivery catheter 300 for the device would be removed leaving inflation tubes 318 and or deployment control tubes 316 tethered to the implant 100. By inserting the retrieval catheter 500 over these connections the implant 100 is now coaxial to the retrieval system 500 and ready to be removed from the body. By advancing the retrieval catheter 500 over the implant 100 or by pulling the control lines 230, the implant 100 can be retracted into the braided section 506. The implant 100 is now covered and may safely be pulled into the sheath 502 and removed from the body.
Other applications for these recapturing systems may be advantageous for devices such as stents (coronary and peripheral), PFO and ASD closure devices, micro coils and other implantable devices that may need retrieval from the body. Currently snares and other tools are used to drag devices out of the body however, many devices will be hung up on catheters or introducers as they are removed. By creating a basket to protect the device from these events, removal becomes simpler and safer.
Another method for device recovery includes providing a string woven through the prosthetic valve 100. As tension is applied to the string the prosthetic valve 100 collapses back down, to a size small enough to be recovered into the delivery sheath, the introducer or a recovery sheath.
Excision and Debulking Devices
The procedure of implanting a valve preferably begins with enlarging the valve annulus. This could be performed with a simple balloon valvuloplasty. However, in many instances this is not sufficiently. Thus, before a prosthetic valve is replaced in a surgical procedure, the surgeon often modifies or removes the native valve leaflets, and especially any calcification or vegetations in the area As will be explained in more detail below, in order to preserve outflow from the heart, between the time that the native aortic valve is excised or debulked and the time that a prosthetic valve is implanted, a temporary valve 520 (see
In one embodiment, the temporary valve 520 can be configured in a manner similar to the implant 100 described above. In such an embodiment, the temporary valve 520 would be delivered via catheterization technique by delivering a collapsed temporary valve and filling the valve body or cuff with fluid to provide structure or by compressing a valve assembly into a catheter for delivery and introducing the valve by removing a sheath to introduce the device to the targeted implantation site. It is also possible to unroll or unwrap a valve assembly from a catheter for delivery. Any method of delivery will suffice as long as the device can be safely removed once the removal and introduction of the new valve has been completed.
The temporary valve 520 should provide a manner for a catheter to pass across the temporary valve while still maintaining flow. The temporary valve 520 can be delivered with a guidewire advanced through the valve to allow guide wire compatible devices to be easily advanced across the valve. If an umbrella type valve is used blood flows between the device and the wall of the aorta. In this case the guidewire or catheter may pass around the valve rather than through the valve.
A modified method to using a temporary valve is to use a percutaneous bypass procedure. When this procedure is performed it is no longer necessary to maintain the flow through the aortic outflow tract. The aorta may be occluded during the excision step and the debris and fluid from the excised area may be aspirated after or during the excision step. In a percutaneous bypass procedure blood is oxygenated extracorporally and reintroduced into the body. A cardiopelegia solution is used to stop the heart-beat.
With reference to
Many various tools are capable of removing portions of the aortic valve 34 or for removing calcification from the aortic valve 34. Examples of such tools that are known for surgical applications or for percutaneous applications include ultrasonic energy sources such as CUSA, hand tools such as cutters or knives and fluids that may dissolve or soften the tissue and or calcium to be removed. As shown in
In one embodiment, an ultrasound transducer may be positioned near a catheter tip and used as a tool to break up calcium and cause it to release from the valve tissue. This method was used for the surgical repair of calcified aortic valves. Unfortunately, the procedure can also damaged the healthy portions of the leaflets causing aortic insufficiency chronically. Typically, the aortic insufficiency would develop in one to two years. In some patients, the native valve was destroyed during the procedure. As a preparation for valve removal, a percutaneous adaptation of this technique may be appropriate. In addition to the ultrasound catheter, some method of collection the calcified tissue is often required. One method is the embolic protection filter described in this application. Alternatively, suction could be applied to the catheter tip, to remove the small particles. With either method, large nodules of calcium may be released from the native tissue. If the nodules are larger than the catheter they must be broken up before they can be safely removed percutaneously. Preferably, the ultrasound transducer can be manipulated to break up these large nodules into particles small enough that they can be removed. This technology is described in U.S. Pat. Nos. 4,827,911, 4,931,04, 5,015,227, 4,750,488, 4,750,901 and 4,922,902, which are hereby incorporated by reference herein. The frequency range for these devices is often about 10-50 KHz but seems to be optimal at about 35 Khz.
Another tool that can be used to excise the native valve 34 may comprise multiple external energy sources that are focused on the tissue to be removed from different directions. This technique can be used with several energy sources, for example ultrasound energy may be used in this way. Radiation energy may also be used in this way, by a method referred to as a gamma knife.
A heated wire system can also be used to cut the aortic valve out from the annulus. In such an embodiment, the wire may be mounted on a catheter and heated by means such as electric resistance or RF energy. The wire may be manipulated in the area of the valve to be removed, and located by balloons or wires. Wire sizes may range from 0.005-0.100 inches in diameter and are typically made from a Ni-chrome material.
In another embodiment, a laser can be used to cut the calcified tissue apart. The laser energy could be transmitted fiber optically through a catheter and applied to the calcified tissue at the catheter tip. The catheter tip may be manipulated by the operator to direct energy to the site-specific area causing ablation or cutting the tissue and or diseased material. It is important that the laser wavelength is correct and will couple to the material to be affected. There may be a need to adjust the wavelength, rep rate and energy density to customize the removal process.
In yet another embodiment, the calcified valve tissue may be broken up and removed using a cutting balloon, or an inflatable balloon with metal or rigid plastic blades along its length. An example of this is U.S. Pat. No. 5,616,149, which is hereby incorporated by reference herein. As the balloon is expanded the blades are forced into the tissue causing it to break apart. Multiple inflations may be required to create a sufficiently large valve area. In one embodiment the balloon is mounted on a torquable catheter allowing a partially inflated balloon to be torqued scraping tissue away from the valve annulus. This balloon source may be used in the “hot-wire” application above to cut the tissue in a pie shaped pattern before removal or exclusion.
Several of the tools described for removing portions of the aortic valve may remove portions of valve or calcium that are larger than can pass through a catheter. In these cases a catheter with a provision to pulverize and extract the excised material may be needed. In one embodiment the catheter includes a rotating auger near its tip to break up the large particles and feed them back through the catheter shaft. Suction may also be applied to the catheter to prevent smaller particles from exiting the catheter tip. Examples of this may include the RotoBlader device produced by Boston Scientific but may be housed in a catheter to limit the escape of particles down stream.
With continued reference to
In the preferred embodiment, the cutting action is performed by pulling the punch 532 proximally into the cutting edge 542. The punch 532 is coupled to the wire 539, which extends through the catheter and is actuated by a handle 546 provided at the proximal end of the device 530. By cutting in this direction the excised tissue is pulled into the catheter 540, and the wire 539 which transmits the force is loaded in tension. An aspiration function is also incorporated into the lumen 535 into which the excised tissue is pulled. By maintaining a minimal fluid flow out through the catheter lumen 535 the risk of embolic events may also be minimized. A spring (not shown) can be provided at the distal end 552 of the device to pull the punch 532 distally after the wire 539 is released.
For a device such that described above or a DCA device, it is advantageous for the cutting portion of the device to be movable to engage the tissue. A balloon or forced wires which forces the cutting portion against tissue, is traditionally used with a DCA device, however this prevents perfusion. In the illustrated embodiment of
In a modified embodiment, the straps 550 extend axially across the portion of the catheter where the cutting takes place, and attach to an elongate member which is free to move axially relative to the elongate member that is attached to the cutting mechanism. The two elongate members are preferably located coaxially. In one embodiment both elongate members are polymer tubes.
In yet another embodiment, an atherectomy catheter device (not shown) includes a housing at the distal end of a substantially round housing torque cable. A cutter torque cable is disposed within the housing and includes a rotatable and translatable cutter at its distal end. The housing includes a window into which an atheroma protrudes. The cutter severs the atheroma. A nose cone attached to the distal end of the housing collects and stores severed atheroma. A stabilizing member is attached to the exterior of the housing opposite the window. A stabilizing member can be provided and includes a balloon having an inflation lumen disposed within the housing. In a modified embodiment, a mechanical stabilizing member provided and includes a distal end attached to the distal end of the housing or to the nose cone, and a proximal end coupled to a stabilizing cable disposed within a cable lumen of the housing torque cable. The stabilizing cable can be advanced distally to bow the stabilizing member away from the housing and withdrawn proximally to flatten the stabilizing member against the housing, alternately urging the window side of the housing onto the atheroma and allowing it to retreat therefrom.
Another method for removing calcification and vegetation from the valve area is with a pharmacological agent. For example, an agent that dissolves calcium is secreted by osteoblasts. An agent similar to this could be utilized prior to the valve replacement procedure. Alternatively an agent like this could be coated on the valve leaflets or on another portion of the prosthesis so that it slowly elutes over the life of the valve. This would prevent or minimize the calcification that contributes to the deterioration of the valve. The agent could be contained in a polymer coating, in a porous metallic coating, or in the tissue itself
To aid removal or debulking, the calcified tissue may be visualized by echocardiography and or fluoroscopy, ECHO, MRI, CT scan as is known in the art.
With reference back to
Many of the devices described above for removing or cutting the valve commisures could benefit from the use of a centering balloon to locate the catheter in the center of the native annulus while the cutting occurs. The centering balloon could be located proximal or distal to the valve, or balloons could be located both proximal and distal. The balloons could optionally contain perfusion lumens.
In a modified embodiment, the method of enlarging the annulus involves a process of shrinking tissue instead of or in addition to removing tissue. For example, it is possible to shrink collagen type tissue by the application of heat. In such an embodiment, the tissue is preferably heated to a temperature of 50 to 65 C. More preferably the tissue is heated to 55 to 60 C in one embodiment the tissue is heated to a temperature of 59 C. The heating may be accomplished from a variety of energy sources, one particularly advantageous energy source for a percutaneous application is RF energy. Accordingly, a catheter with a heated element on the tip may be used to heat specific portions of the valve.
In one embodiment, the catheter incorporates a needle near the heated portion. The portion of the catheter intended to transfer heat to the leaflet tissue is positioned below the surface of the leaflet. This minimizes the transmission of heat into the bloodstream, while maximizing the transmission of heat to the leaflet tissue.
In another embodiment the heating step is applied by a tool that also dilates the annulus. This tool may be a balloon inflated with a heated solution or a dilation device that contains heating elements, such as those described in this application using deflected straps.
In general, the application of heat is intended to affect the portions of the leaflets nearest to the center of the valve. Excessive shrinking of the outer portion of the valve annulus may cause the effective orifice area to be reduced. Shrinking an area near the tip or free edge of each leaflet will cause the effective orifice area to be increased. It may additionally release the calcium deposits within the valve tissue thus providing a large effective orifice area to implant a new valve.
Procedures for Deploying the Implant
Various procedures and methods for deploying an implant 100 in the aortic position will now be described. In one embodiment, the method generally comprises gaining access to the aorta, most often through the femoral artery. A balloon valvuloplasty may optionally be performed in the case of aortic stenosis, or another method may be used to remove or debulk the native valve as described above. A delivery sheath or catheter is advanced over the aortic arch and past the aortic valve. The outer sheath of the catheter is retracted exposing the valve and cuff. Fluid is used to inflate the valve and a second inflation fluid may be used to partially form the implant. This allows the distal portion of the implant to open to its full diameter. The proximal portion of the implant may be slightly restricted by the deployment control mechanism. In general, the amount that the deployment-control mechanism restricts the diameter of the proximal end of the device depends on the length of the wires extend past the outer sheath, which an be adjusted by the operator. Alternatively, in some embodiments, the implant contains multiple inflation ports to allow the operator to inflate specific areas of the implant different amounts. In another embodiment, burst discs or flow restrictors are used to control the inflation of the proximal portion of the implant 100. The implant is then pulled back into position. The distal ring seats on the ventricular side of the aortic annulus. A balloon may be used to dilate or redilate the device if necessary. At this time, the deployment control wires may act to help separate fused commisures by the same mechanism a cutting balloon can crack fibrous or calcified lesions. Additional casting material may be added to inflate the implant fully. The inflation lumen is then disconnected, and the deployment control wire(s) are then disconnected, and the catheter is withdrawn leaving the device behind. In modified embodiments, these steps may be reversed or their order modified if desired.
The above-describe method generally describes an embodiment for the replacement of the aortic valve. However, similar methods could be used to replace the pulmonary valve or the mitral or tricuspid valves. For example, the pulmonary valve could be accessed through the venous system, either through the femoral vein or the jugular vein. The mitral valve could be accessed through the venous system as described above and then trans-septaly accessing the left atrium from the right atrium. Alternatively, the mitral valve could be accessed through the arterial system as described for the aortic valve, additionally the catheter can be used to pass through the aortic valve and then back up to the mitral valve.
For mitral valve replacement, the implant may require a shorter body length (e.g., 1-4 cm) and would mount in the native mitral valve area. It may be delivered from the right side of the heart from the femoral vein up through the inferior vena cava and into the right atrium. From there, a transeptal-puncture may be made for entry into the left atrium and access to the mitral valve. Once in the left atrium, the implant would be delivered with the valve pointing down to allow flow from the left atrium to the left ventrical. A similar shape would allow the device to be deployed in the left atrium and advanced into the left ventrical. The proximal ring may require inflation to hold the device in the left ventical by creating a diameter difference between the mitral orifice and the proximal cuff diameter. Here as with the aortic replacement, the mitral valve may require partial removal or cutting of the valve or chorde to allow the native valve to be excluded and provide room for the replacement valve to be implanted. This may be achieved by balloon valvuloplasty, cutting techniques such as a cutting balloon or by utilizing a hot-wire or knife to cut slits in the native valve to allow for exclusion. Once the native valve has been prepared for the new valve, the mitral valve orifice may be crossed with the distal portion of the implant and the distal portion may be inflated for proper shape and structure. At this time, the native valve will have been excluded and the replacement valve will be fully operational.
Other methods of mitral replacement would include a transapical delivery where the patent would receive a small puncture in the chest cavity where the operator could access the apex of the heart similar to a ventricular assist device implantation. Once access is gained to the left ventrical, the aortic and mitral valves are a direct pathway for implantation of the replacement valve. In this case, the aortic valve would be delivered with the flow path in the same direction as the catheter. For the mitral valve, the flow path would be against the direction of implantation. Both may still utilize the base of the implant to anchor the device using diameter differences to secure the device. It may be desirable to also use a hook or barb that could protrude from the cuff either passively or actively as the cuff is filled with fluid. The barb could be singular or a plurality of barbs or hooks could be used where the length could be between 1-5 millimeters in length depending upon the tissue composition. Where a longer barb may be required if the tissue is soft or flexible. It may be desired to have shorter lengths if the tissue is a stiffer more fibrous structure where the barbs could hold better.
For the pulmonary and tricuspid valve placement, the operator could access the femoral vein or internal jugular (IJ) vein for insertion of the delivery system. As with the transeptal mitral valve approach the delivery system and device would be introduced either superiorly or inferiorly to the vena cava and to the right atrium and right ventrical where the pulmonary and tricuspid valves are accessible. The femoral approach is preferable due to the acute bends the delivery system would be required to make from a superior or IJ access. Once in the right ventrical the device could be delivered similarly to the aortic method where the cuff utilizes the base of the pulmonary valve for a positive anchor with a diameter difference holding it from migrating distally. It may be desirable to also use a hook or barb that could protrude from the cuff either passively or actively as the cuff is filled with fluid. The barb could be singular or a plurality of barbs or hooks could be used where the length could be between 1-5 millimeters in length depending upon the tissue composition. Where a longer barb may be required if the tissue is soft or flexible. It may be desired to have shorter lengths if the tissue is a stiffer more fibrous structure where the barbs could hold better.
In any placement, the proper valve configuration would be chosen by the performance of each required application. For instance the aortic valve may require a two or three-leaflet valve that will require a high degree of resistance to stress and fatigue due to the high velocities and movement. The pulmonary valve may require a lesser valve due to the more passive nature or the lower pressure that the valve is required to support. Lengths may vary and will be dependant upon the valve and structure surrounding them. A shorter valve (1-4 centimeters) may be required for the mitral but the aortic may allow for a longer valve (1-8 centimeter) where there is more room to work. In any application, the maximum orifice size is generally desired since the cross sectional area helps determine the outflow volume. The aortic cross sectional area may vary from nearly 0.00 square centimeters in a heavily calcified valve to about 5 square centimeters in a healthy valve. Most cases the desire in replacement is to increase a cross sectional area for additional flow.
During the procedure or during patient selection, or follow-up, various imaging techniques can be used. These include fluoroscopy, chest x-ray, CT scan and MRI. In addition, during the procedure or during patient selection, or follow-up, various flows and pressures may be monitored, for example echocardiography may be used to monitor the flow of blood through the relevant chambers and conduits of the heart. Pulmonary wedge pressure, left atrial pressure and left ventricular pressures may all be recorded and monitored. It may be desirable to use a measurement tool to determine the size of valve required or to determine if the anatomy provides enough room to allow implantation of a valve. In the past, marker-wires have been used to measure linear distance and a similar technique could be used in this application to measure a distance such as the distance from a coronary artery to the annulus of the aortic valve. To measure the diameter of a valve, a balloon with a controlled compliance could be used. Ideally, the balloon would be very compliant and inflated with volume control, but a semi compliant balloon could also be used and inflated with a normal interventional cardiology inflation device. The compliance curve of the balloon could then be used to relate the pressure to the diameter. The diameter range of valves in the heart may range from 10-50 mm in diameter and 2-40 mm in length. A similar sizing balloon has been used for sizing septal defects.
In one embodiment, implantation of a prosthetic valve includes the step of dilating the valve after it is positioned and functioning within the native anatomy. If the dilation step is used to replace a balloon valvuloplasty prior to the inflation of the balloon the cuff will minimize the embolization from the dilation. The dilation of the functional implant step may also be used in patients where a valvuloplasty is performed prior to implantation of the device, but where the outflow area is not as large as desired. Certain embodiments of the implantable prosthetic valve include deployment control wires or stiffening wires. If these features are present in the implant at the time of post dilation, then the features may act to concentrate the force from the deployment of the balloon in a mechanism similar to the function of a cutting balloon commonly known in interventional cardiology.
To gain access to the aortic valve the femoral arteries (radial, brachial, carotid) can be used to introduce tools into the vascular system. Once in the arterial conduits, catheters may be advanced to the aortic arch and the native aortic valve. As discussed above, it may be necessary to install a temporary valve to allow gating of the blood flow while the work is being completed on the native valve. This will provide time for the interventional cardiologist to prepare for removal and installation of a new aortic valve. The placement of the temporary valve could be between the native valve and the coronary arteries, or the valve could be placed in a location between the coronary arteries and the location where the great vessels branch off from the aorta or at any other location within the patients aorta. Placing a valve in these non native locations to treat aortic insufficiency has been proven effective in clinical experience by the use of the Huffnagel valve. Placing a temporary valve in these locations has been described by Moulolupos and Boretos. A guidewire or pig-tail catheter may be used to pass a stiffer catheter through the stenotic hole in the aortic valve. It may be necessary to install a filtration device to protect any vessels including the coronary tree from debris as the valve is loosened and removed. This filter may be placed in the region of the aortic valve just before the coronary ostia or distal to the sinus and just before the great vessels. Once through the valve opening a balloon may be passed into the aortic valve to predilate the region and loosen any calcium. This may aid in the removal of the tissue that may be calcified and or fibrosed. The use of a catheter to deliver energy such as ultrasonic, RF, heat or laser may additionally break or loosen the tissue including the calcification in and on the leaflets. There are chemical treatments that have shown some promise in dissolving the calcium such as Corazon Inc. of California (see U.S. Pat. No. 6,755,811). The ultrasonic energy device is described in detail through U.S. Pat. No. 4,827,911 and has a proven track record known as CUSA to remove calcium in a surgical suite from valve tissue. This has shown promise acutely but will denature the collagen tissue and result in a degeneration of the valve tissue remaining in about a year leaving a poorly functioning valve. After a filter has been installed and the valve tissue has been softened, a template may be used to define the area to be removed. This template will define the hole and prevent the removal of healthy tissue. At this time the valve is ready to be removed with adequate time since the temporary valve will be functioning when the native valve is removed. This will be important to not allow the patient to go from aortic stenosis to aortic insufficiency. The removal tool may now be passed through the stenotic valve and begin the removal process of the native valve. As mentioned above and in patents and US applications such as 20040116951 Rosengart there are many ways to remove tissue from this region.
The embodiments described above provide a technique that lends itself well to delivering a catheter based valve removal tool. Through a pushing and pulling force the pin and die set as seen in the drawings will allow the valve to be removed in a controlled manner while leaving the material in a catheter shaft for removal. It is asserted that this is the first that allows the aortic outflow track to be gated or valve temporarily. Though an aortic balloon pump may function as a temporary or supplementary valve in some conditions, the balloon pump is ineffective and dangerous in patients with aortic insufficiency. A removed or partially removed aortic valve constitutes severe aortic stenosis.
As shown in
Alternatively, the temporary valve 520 may be placed so that it acts between the native aortic valve and the coronary arteries although its physical position would likely extend well above the coronary arteries. In this embodiment the inlet side of the temporary valve would seal to the aortic wall just below the coronary arteries. The outlet side of the valve would extend up beyond the coronary arteries. The mid portion of the valve and the outlet side of the valve would have an outside diameter smaller than the inside diameter of the patients aorta. This would allow blood flow from the outlet of the valve, around the outside of the valve back towards the ostia of the coronary arteries. In this embodiment the valve would have a sealing portion on the inlet side of the valve, the sealing portion would have an outside diameter to match the patients aortic root diameter. This diameter would range from about 18 mm to about 38 mm. Multiple sized valves are required to accommodate differing patient anatomies. The sealing portion of the valve may be expandable or compliant to improve sealing and best conform to a wide range of patient anatomies. The length of the sealing portion is limited by the position of the valve and the position of the coronary arteries, the length of the sealing portion may range from about 1 mm to about 5 mm, preferably about 3 mm. The mid and outlet portions of the valve are preferably between 30% and 90% the diameter of the native aorta. This allows sufficient room for blood to flow back around the valve and perfuse the coronary arteries. The valve may also incorporate a secondary retaining mechanism, securing the outlet or mid portion of the valve beyond the coronary arteries
Alternatively, the valve can be replaced by a pump similar to a device designed by Medtronic known as a Hemo Pump, which is placed in the aorta. The pump moves blood out from the ventricle into the aorta, serving the function of both the native aortic valve and the contracting left ventricle. The pump may consist of a screw type pump actuated by a rotating shaft, where the motor is located outside the body. The inlet of the pump located on the distal end of the catheter may optionally be isolated from the outlet of the pump by a balloon. The balloon inflates between the outside diameter of the pump and the inner diameter of the aorta, in a location between the pump inlet and the pump outlet. Alternatively a pump using two occlusion balloons, both between the inlet and the outlet of the pump could isolate an area between the balloons for treatment. The valve removal procedure could take place in this area.
The temporary valve designs described by Moulolupos and Boretos in U.S. Pat. Nos. 3,671,979 and 4,056,854 respectively, include umbrella valve designs that allow the blood to flow in one direction between the valve and the wall of the aorta. The valves prevent flow in an opposite direction as the valve seals against the wall of the aorta. These valves can be attached to a temporary valve catheter and adapted for use with the present invention.
Other valve designs are also possible for a temporary valve including a ball and cage valve, a tilting leaflet valve, bi-leaflet valve a reed type valve, a windsock style valve, a duckbill valve, or a tricuspid valve. In addition to these valves made from synthetic materials including polyurethane or tissue valve may also be utilized. Commonly used in permanent valve replacements valves constructed from bovine pericardium or porcine aortic valves, are adequate. To produce a low profile percutaneous device the preferred embodiment is a thin flexible polymer valve of either a duckbill design or umbrella valve design.
The temporary valve should be placed in such a way that it can be easily removed at the end of the procedure and also in such a way that the operator has access across the valve for performing the remaining steps. A guidewire or catheter lumen placed through the valve or around the valve before the valve is positioned in the body allows the required access for downstream procedures.
Alternatively an inflatable structure may be used. The inflatable structure provides the advantage of improved sealing characteristics with the vessel wall, and the inflatable structure may produce a lower profile device with some valve designs. The inflatable valve structure could be designed to be recoverable using wires as described in previous direct flow disclosures for permanent valve replacement devices, an inflatable prosthetic valve was first described by Block in U.S. Pat. No. 5,554,185 and is also described herein. The inflatable structure preferably inflates to an outside diameter between about 18 mm and about 35 mm.
In another embodiment, the temporary valve structure is a recoverable self-expanding stent. The stent could be a Z-stent formed from wires segments shaped into rings or a coil. Alternatively the Z-stent could be cut from a tube using a process like laser cutting. With a Z-type stent careful design of the stent shape is required to make the stent recoverable. It must be ensured that no crown hangs up on the recovery sheath. One method to accomplish this is to attach each crown to the crown of the next stent segment by welding, fusing or other joining techniques. Or the stent could be braided from wires in a design similar to a Wall Stent as produced by Boston Scientific. The material for the stent is preferably a superelastic material such as nitinol. Alternatively a material with a relatively high yield strength and/or a relatively low modulus of elasticity, such as a cobalt chrome alloy, or titanium, could be used. These non superelastic materials are most appropriate for use in a stent manufactured by a braiding process.
In another embodiment, the structure for the temporary valve 520 consists of an unwrapable structure, similar to the structure described by Yang in U.S. Pat. No. 6,733,525 or as described herein. The structure is delivered in its wrapped position. After the structure is positioned the structure is unwrapped and expanded to its final diameter.
In general, any of a wide variety of valve structures may be utilized for the temporary valve in accordance with the present invention. Since the temporary valve is only intended to remain functional at an intraluminal site for a relatively short period of time (e.g. less than a few hours), the temporary valve of the present invention is not plagued by many of the deficiencies of prior permanent implantable valves (thrombogenicity, efficiency, durability, etc.). Thus, valve design can be selected to minimize the initial crossing profile and optimize removal.
For example, in the example described previously in which a valve is supported by a Z-stent structure, each of the proximal apexes of the stent may be attached to a pull wire, which merge into a common axially moveable control wire which runs the length of the temporary valve deployment catheter. Following transluminal navigation to the desired temporary valve site, an outer sheath may be proximally retracted relative to the control wire, thereby enabling the stent and valve to be deployed from the distal end of the catheter. Following completion of the procedure, the temporary valve may be removed by applying proximal traction to the control wire and/or distal force on the outer sheath. The plurality of control filaments will cause the Z-stent to collapse, as it is drawn back into the tubular sheath.
Thus, the temporary valve of the present invention is preferably permanently attached to its deployment catheter. In this regard, the term “deployment” refers to the conversion of the temporary valve from a reduced cross sectional profile such as for transluminal navigation, to an enlarged cross sectional profile for functioning as a valve in a vascular environment. However, at no time does the valve become detached from the deployment catheter. This eliminates the complexity of snaring or otherwise recapturing the temporary valve, for retraction into a catheter. Alternatively, the present invention may be practiced by the use of a detachable temporary valve, which must be captured prior to removal.
The preferred temporary valve is therefore preferably carried by an elongate flexible catheter body, having a proximal control for advancing the valve into a functional configuration, and retracting the valve into a collapsed configuration for transluminal navigation into or away from the temporary valve site. Activation of the control to retract the valve back into the temporary valve catheter does not necessarily need to preserve the functionality of the valve. Thus, proximal retraction of the valve into the temporary valve catheter may involve a disassembly, stretching, unwinding, or other destruction of the valve if that is desirable to facilitate the step of removing the temporary valve.
Although tissue valves may be used for the temporary valve in accordance with the present invention, due to the short duration of the intended working life of the valve, any of a variety of polymeric valves may be adapted for use in the present context. Polymeric membranes may be configured to mimic the leaflets on a normal heart valve, or may be configured in any of a wide variety of alternative forms, as long as they are moveable between a first, open configuration and a second, closed configuration for permitting blood flow and essentially only a single direction. Thus, polymeric membranes may be formed into any of a wide variety of flapper valves, duck bill valves, or other configurations.
Regardless of the valve leaflet construction, the temporary valve may be supported by an inflatable cuff as has been disclosed elsewhere herein. The temporary valve deployment catheter is provided with an inflation lumen extending between a proximal source of inflation media and a distal point of attachment to the inflatable cuff. Once positioned at the desired site, the temporary valve may be released such as by proximal retraction of an outer delivery sheath. Inflation media may thereafter be expressed from the source to inflate the cuff to enable the valve and provide a seal with the vessel wall. Following the procedure, the inflation media is aspirated out of the cuff by way of the inflation lumen 318 to deflate the cuff, and the temporary valve is withdrawn from the patient.
Alternatively, the temporary valve may take the form of an inflatable balloon, with an inflation cycle which is synchronized to the heart beat so that it is deflated to permit forward flow but inflated to inhibit reverse flow in the artery.
An embolic protection filter may be mounted to the temporary valve or to the temporary valve structure. The filter may be attached to the outlet section of a duckbill type valve. Alternatively the filter may be mounted on its own support structure.
With the temporary valve deployed as shown in
A trapping size of about 35 to 250 micron and may be treated with an anti-thrombogenic coating to prevent clotting. A basket of similar design could be mounted to the catheter shaft of a device designed for percutaneous treatment of a coronary valve, in a case where the valve is approached from a retrograde direction. In an application where the device is placed in an antegrade direction, a larger version of a conventional wire based embolic protection device could be used.
In an application for aortic valve treatment it may be desirable to place the embolic protection very close to the annulus of the valve because the ostia of the coronary arteries are very close to the area being treated. In a balloon, valvuloplasty used as a pretreatment for valve replacement or alone as an independent therapy, the embolic protection filter may be attached to the proximal end of the balloon or to the catheter shaft very near the proximal end of the balloon, specifically within 1 cm of the proximal end of the balloon. The filter could be positioned similarly on a catheter for the delivery of a percutaneous prosthetic valve, this configuration is especially beneficial for a balloon expandable prosthetic valve.
An alternative method of embolic protection applicable to a balloon valvuloplasty or implantation of a percutaneous prosthetic valve by means that prevent flow through the aortic valve is described as follows. Flow is occluded in a position at the treatment site or, preferably beyond the treatment site, in either a retrograde or antegrade direction. The treatment is performed. The treatment site is disengaged from the device. The treatment area is aspirated. Because the flow is prevented by the occlusion the embolic material does not travel. The occlusion is then removed. The preferred embodiment for an aortic application is a valvuloplasty balloon with dual balloons. A larger distal balloon is inflated within the ventricle. The balloon is pulled back so that the aortic outflow is obstructed, the balloon is sized so that it is significantly larger than the aortic valve. The second smaller diameter balloon located immediately proximal to the first balloon is then inflated to dilate the valve annulus. The second balloon is then deflated and the entire area aspirated with an aspiration catheter. The first balloon is then deflated to restore aortic outflow. Alternatively, there may be a tube central to these balloons providing flow while this operation in occurring. This would be a limited by-pass of oxygenated blood around the area being decalcified. During this by-pass, a cutting mechanism may be introduced where as the valve and calcium may be mechanically removed. Examples of cutting mechanisms would include a rotating burr, an oscillating pin and die to punch the material out in segments or ultrasound energy to fragment the material free for aspiration removal. It may be necessary to additionally canulate the coronary arteries to continue flow to these critical vessels.
The system could further contain a perfusion lumen to reintroduce the left ventricular outflow in a location that does not cause the movement of blood in the area of the aortic root. For example blood could be reintroduced in the coronary arteries or in the aortic arch or in the carotid arteries.
It may also be possible to have the filter device 522 mounted to the delivery catheter and actuated by the handle to open and close the filter to the vessel wall. This device would be placed between the aortic valve and the great vessels in the arch. A secondary catheter system could also be used to filter debris from the aorta and delivered from another vessel to the arch. This filter could also be attached to the temporary valve assembly providing filtration protection with valve support as the native valve is removed or decalcified. A filter could also be mounted to the excision tool protecting the down stream vessels from emboli. By protecting each individual vessel such as the carotids, great vessels, and the aorta separately, devices would be required in each of these vessels to protect them from emboli. These filters could be a simple windsox style as seen by EPI (Boston Scientific) and could recover the emboli through a catheter. Other systems for filtration include the Percusurge device sold by Medtronic where balloons protect the area of interest and aspiration withdrawals the emboli.
Filtration devices may be set directly on the calcified aortic valve to prevent any material from escaping. This filtration device may be made from a woven or braided wire such as Nitinol or stainless steel, MP35N, polymeric fibers or other suitable material commonly used in medical devices. The materials may be composed of round, oval or flat ribbon material. This may provide benefits when designing low profile device. These wire would be have cross sectional diameters ranging from 0.001-0.030 inches. These wires may be supported by larger extension wires to hold the filter material open as seen in. The filter may require a support structure such as a stent or series of struts to provide dimensional integrity. This stent structure could be a common Z-stent or an inflatable structure to hold the filter open and sealed to the valve base or vessel wall. The support structure would be expanded or deployed by exposing the device from a sheath or by actively providing a force to move the structure from a beginning shape to a final shape. Housed in a catheter for delivery, the device would be constrained to a small cross section and expand to a larger cross sectional diameter or area as allowed. The deployed device would have a general conical shape with the open large diameter facing down or toward the valve. The opposite end would come together at the catheter and be retrievable by the introduction catheter or a second retrieval catheter to remove any debris captured. These catheters may have a diameter of about 8-24 French. The filtration material could be located inside or outside the support structure depending upon what flow characteristics were required. For instance, if the filter material was located on the outside of the support structure the filter may be in contact with the coronary ostia. It may be more desirable to have the filter material on the inside of the support structure holding it away from the ostia of the coronary arteries. The filtration would trap particles from about 35-250 microns in size and allow adequate flow through the aorta. The distal portion of the filter may have a ring or template at the distal end to allow for a patterned removal of the native aortic valve. The distal portion would fit between the aortic wall of the sinus and the calcium deposits to be removed. The template would provide a pattern that may be traced by a removal tool as described in paragraphs above. By using a template the pattern would be close to the native healthy orifice. An acceptable cross sectional area would be about 2-3 cm2. This would provide adequate room to place a new valve and provide the patient good hemodynamic flow. This template could be as simple as a guide provided by a wire ring or a pattern with three arches as seen in a healthy valve similar to a clover. It may however be simpler to provide a round hole than a complex shape to begin. This template may be above or below the native valve and may require more than one shape and or size.
Another design for a filter is to utilize a braided Nitinol stent that will provide support to the filter material but is still recoverable inside the catheter. In this embodiment, the filter material would be inside the braided structure and the braid would be in contact with the aortic wall. This would provide a seal between the filter device and the vessel directing flow through the filter element and allowing the coronary arteries to be patent
After the temporary valve 520 and embolic protection filter 524 are in place, a debulking or valve removal step is performed as shown in
After wire access has been gained it still may be difficult or impossible to pass some embodiments of the cutting device across the stenossed native valve. If necessary a preliminary cutting step may be performed to enlarge the valve opening sufficiently that a second cutting device may be inserted. In one embodiment the primary cutting device includes a rotating burr centered on a guidewire. The burr is mounted to a small flexible hypotube or solid shaft, which is spun by a motor outside the body. Preferably the hypotube has an inside diameter of 0.014 to 0.040 in and the shaft has a diameter of about 0.010-0.030 inches. The rotating burr preferably has an outside diameter slightly larger than the secondary cutting tool this is preferably in a range of 2 to 6 mm diameter. A similar rotating burr device is marketed by Boston Scientific, under the trade name Rotoblader, for the treatment of stenotic arteries.
A cutting device of this design could also be used to open the calcified valve to the desired diameter. In this case a larger burr may be used ranging in diameter from about 3 to about 9 mm in diameter. A steerable catheter may be required to center the newly enlarged opening in the native anatomy. A steerable catheter may consist of a flexible elongate tube with a pull wire located off center in at least a portion of the elongate tube. When tension is applied to the pullwire, it causes the catheter to bend in the direction to which the wire is offset. Multiple pullwires may be used to allow the catheter to be steered in multiple areas or directions. The catheter my also be manufactured with a preferred bending plane, allowing even a centered pullwire to steer the catheter, and providing more precise control of the catheter shape. The catheter is preferably of an outside diameter between 3 mm and 9 mm.
Several embodiments of cutting devices are possible some of which are describe above. In one embodiment, the cutting device 530 consists of a tool that pushes or pulls a sharpened punch into a die as described with reference to
Alternatively a similar cutting device could be used where the cutting portion consists of a rotating cutter. The cutter is pulled back through the die portion forcing the material into the catheter shaft in a similar manner to the device described above. The rotating edge of the cutter may be sharpened to an edge to minimize embolic material as much as possible or may be serrated to maximize the cutting ability of the device. This device is very similar in function to devices commonly used for DCA or directional coronary atherectomy. Typically DCA devices cut in a push mode, capturing the cut out section in a cavity near the distal tip of the device. The devices described above operate in a pull mode, which allows the cut out material to be evacuated out the catheter shaft or fill a larger area within the catheter shaft. However either cutting device described could be manufactured to operate in a push mode rather than a pull mode. It may be desired to have the helix direction pull the material back or proximally to the catheter handle. This would allow for convenient removal of the debris from the body.
The cutting device may include a device to engage the cutting portion of the device to the tissue. In one embodiment a balloon possibly a perfusion balloon is attached to the non-cutting side of the device. As the balloon is inflated the cutter is moved laterally to engaged into the tissue. To maintain flow out of the heart the balloon inflation and cutting may be accomplished during the time that the aortic valve would be closed. This could be synchronized to the patients heart rate by echocardiography or similar sensing techniques or the patient could be placed on a temporary pace maker and the pacemaker output could be used to time the inflation of the balloon. The inflation media could be a liquid or a gas. A gas such as helium or CO2 would allow the quickest inflation time through a small lumen. Helium would provide an even quicker inflation time than CO2, however CO2 may be better dissolved in the blood in the event of a balloon burst.
Preferably the engagement method allows flow to pass around the engagement device as shown in
Another engagement method that allows flow past the catheter includes a steerable catheter mechanism. The device can be bent in such a way that the window of the cutting device is pushed against the tissue, while this force is opposed by a section of the catheter bushed against tissue in an opposite direction. A DCA device marketed by the company Foxhollow uses this mechanism to engage tissue.
The engagement means may be adjustable to a predetermined range of sizes from the catheter handle. The cutting tool is advanced or retracted into the annulus and successive cuts are made by the operator. The catheter may be rotated slightly between each cut. Once the new annulus is cut out large enough for the engagement means to pass through the annulus the operator knows that the annulus has been enlarged to a size that corresponds with the adjustment, or size of the engagement means. If this is a sufficiently large cross sectional area for adequate flow after the permanent prosthetic valve is implanted, then the cutting device may be removed. If a larger annulus is desired the engagement means may be adjusted or replaced with a larger size, and the process repeated. If the distal end of wire straps are attached at the distal end of the cutting device and the proximal end of the wire straps are attached to the distal end of a sheath mounted over the cutting tool shaft, then the advancement of the sheath will cause the wire straps to bow out and engage the tissue. The distance that the sheath is advanced corresponds to the diameter that the engagement mechanism will pass through. Markings on the cutter shaft show the operator what diameter the engagement means is expanded to. Preferably the engagement means is expandable to at least about 2 cm. This provides an effective orifice area over 3 cm2.
The cutting device 530 may include a lumen for contrast or therapeutic agent injection. The injection of contrast allows the operator to visualize the size and position of the cut out area relative to the aortic root and the ventricle, under fluoroscopy, MRI, NMR or other imaging techniques used in interventional cardiology. The injection of a therapeutic agent may be used to have any desired effect on the heart or ventricle. Certain therapeutic agents such as antibiotics may aid in reducing the risk of endocarditis or in the treatment of a valve damaged by endocarditis. Other therapeutic agents may increase or decrease the heart rate or hearts output as desired by the physician. The diameter of the inflation lumen is preferably between 0.010 and 0.060 in. in diameter.
As the valve 34 is being removed imaging the procedure is important. The operator must be able to visualize the position of the cutout relative to the aortic wall and the aortic root. Two-dimensional imaging techniques such as fluoroscopy need to be performed on multiple axis to allow the cutting procedure to be performed safely. The operator must be careful not to cut through the aortic wall or through the ventricle. Electrical conduction paths near the annulus such as the bundle of his may require special attention and care. The area between the anterior leaflet of the mitral valve and the aortic valve must not be damaged, and the mitral leaflets and chordae must be avoided. To visualize these and other obstacles during the procedure any number of common imaging techniques may be employed either during the procedure or prior to the procedure in a road-mapping step. Echocardiography may be employed in one of several forms to image the necessary areas. TEE or trans esophageal echocardiography may be particularly useful in imaging the valve area before the procedure begins and during the procedure as well. TTE may also be used with the benefit of being less invasive to the patient, but it is limited by a reduced image quality and the fact that the operator's hand must be near the patient's chest. This makes the simultaneous use of fluoroscopy and other imaging techniques unsafe for the operator. During the procedure fluoroscopy or MRI or NMR or similar imaging techniques may be used to visualize the size and shape of the newly cut out opening and the position of the opening relative to all the relevant structures of the native anatomy.
The cutting device 830 could be actuated by a simple lever type handle moving the cutter in either a proximal or a distal direction as the handle is squeezed. In addition the handle could contain a rotational swivel or union that allows the catheter to be rotated while the handle is held in a fixed position. Further, the function of the catheter rotation could be incorporated into the actuation of the handle. The handle mechanism could be designed or adjusted sot that the catheter rotates a predetermined amount each time the cutter is actuated. This could be accomplished with a simple cam and sprag mechanism, or using a stepper motor. The actuation of the cutting mechanism could also be powered electronically or pneumatically to minimize operator fatigue and to prevent the overloading of the device. In this case the operator would simply depress a button to actuate the cutting function. The handle may also contain an aspiration lumen to assist in the removal of debris from within the catheter shaft, and an injection lumen to inject contrast media or a therapeutic agent, or a fluid such as saline. Other means of providing energy to the device include an impact or momentum drive where a high velocity rate would contact the area to be removed providing a high degree of force to the calcific valve. Drive or propulsion methods may include a gaseous discharge or chemical reaction to generate a hydraulic force to drive an object into or through the calcified valve. Other predictable forces may include preloading a spring mechanism and releasing the energy stored to drive an object into or through the calcified valve.
The valve could also be cut out in sections using a laser or heated wire. Severely calcified areas could be broken up with Cavitation Ultrasound energy, prior to removal with a cutting tool or the calcified areas may be broken up with ultrasound and the debris captured in a filter. Similarly a chemical compound could be used to dissolve or break up the calcium.
With reference to
With reference now to
The illustrated embodiment provides a method of implanting a percutaneous prosthetic valve assembly, where the outflow tract is not blocked at any time during the implantation process. In heart failure patients, blocking the aortic output can have serious consequences, such as death. Another less significant problem with blocking aortic output is the contracting ventricle can exert significant pressure on the device, making positioning very difficult, and possibly forcing the device away from the desired location before it is completely deployed or anchored. To overcome this issue in some cases patients have been rapidly paced. By increasing the patients heart rate to such an extent that the heart does not effectively pump blood. This may not be required during the implantation of this inflatable device.
In contrast, devices such as those disclosed in Andersen family of US patents (U.S. Pat. Nos. 5,411,552 6,168,614 6,582,462) result in the complete or nearly complete obstruction of the aortic valve during deployment. For example as a balloon expandable valve structure is expanded the balloon blocks the aortic output. In one embodiment Andersen describes the use of multiple balloons to deploy the valve, as was common with a balloon valvuloplasty. Using multiple balloons would provide a very small path for fluid to flow between the balloons when the balloons are fully inflated to a pressure high enough that they take on their natural generally round cross section. However when the balloons are partially inflated or during the inflation process the multiple balloons conform to and occlude the lumen, resulting in complete or nearly complete blockage of the outflow tract.
The self-expanding valve support structures disclosed in Andersen and Leonhardt (U.S. Pat. Nos. 6,582,462 and 5,957,949) also block aortic outflow as they are deployed. As the sheath is retracted from the distal portion of the device the device opens and begins to conform to the native vessel. The portion of the valve structure designed to seal to the valve annulus or other portion of the native anatomy comes in contact with the native anatomy. At the same time, the proximal portion of the device is still restrained within the deployment catheter, preventing the valve from opening. At this stage of deployment the devices effectively block all aortic output.
The simplest extension of existing technology to allow implantation of a prosthetic valve without blocking flow is the use of a perfusion balloon with a balloon expandable support structure. The perfusion balloon would have a lumen through the balloon large enough to allow significant perfusion through the balloon during deployment. Perfusion balloon technology is well developed, and known. Wasicek et al describe a perfusion balloon catheter is U.S. Pat. No. 6,117,106
Using a self-expanding valve support structure it would be possible to maintain flow past the valve using a tube section placed through the affected valve, outside the self expanding support structure. After the self-expanding support structure is completely deployed the tube section could be withdrawn. The tube section is longer than at least the sealing portion of the self-expanding valve support structure, and preferably attached to an elongate member to allow its withdrawal. Alternatively the tube section could be located inside the valve support structure. In this case the tube section would allow fluid (usually blood) to flow into the deployment catheter. Perfusion holes in the deployment catheter would allow blood to flow out into the native conduit.
Relating to the current inflatable prosthetic valve or cast in place support structure described herein, a different deployment procedure is used which allows outflow to be maintained. This deployment method could also be used with some self-expanding percutaneous valves. The deployment method is described as follows for an aortic valve replacement. The procedure could be easily adapted to any other coronary valve. The deployment catheter is advanced across the aortic valve. The prosthetic valve and inflatable cuff are unsheathed in the ventricle, but remain attached to the deployment control wires. The distal end of the inflatable cuff is inflated. The sheath is retracted far enough that the deployment control wires allow the prosthetic valve to function. The device is then withdrawn across the native valve annulus. The device is then fully inflated. The valve function may be tested using various diagnostic techniques. If the valve function is sufficient the inflation media may be exchanged for the permanent inflation media. The deployment control wires and the inflation lumens are then disconnected and the catheter withdrawn. In this procedure the key to maintaining the outflow tract is the use of deployment control wires. The deployment control wires allow the device to be moved an appreciable distance from the deployment sheath before the device is permanently positioned in the desired location. Other deployment control devices could be used to have similar effect. For example a sheath used as a shear barrier between the retractable sheath and the implant having longitudinal slots could be configured to produce a similar function. It may be desirable to predilate the native valve annulus with a balloon before device implantation. This may allow for a larger effective orifice area to implant the device and precondition the valve area. Secondarily, an additional dilatation may be desired after implantation to ensure the device is apposed to the wall of the annulus and seated properly.
The current percutaneous valve replacement devices do not provide a means for testing the function of the valve before committing to the position of the valve. These devices are deployed at a location and if the location was a wrong location or if the valve does not have a good effect, the valves can-not be removed. The present invention includes a method of valve implantation consisting of the steps of positioning the valve, enabling the valve, testing the function of the valve, and finally deploying the valve.
Relating to the current inflatable prosthetic valve or cast in place cuff, a unique deployment procedure is used, consisting of the steps of position, enable, test, and reposition or deploy. This deployment method could also be adapted to a valve with a self-expanding support structure or to other implantable devices. The deployment method is described as follows for an aortic valve replacement. The procedure could be easily adapted to any other coronary valve. The deployment catheter is advanced across the aortic valve. The prosthetic valve and inflatable cuff are unsheathed in the ventricle, but remain attached to the deployment control wires. The distal end of the inflatable cuff is inflated. The sheath is retracted far enough that the deployment control wires allow the prosthetic valve to function. The device is then withdrawn across the native valve annulus. The device is then fully inflated, enabling the valve to function. The valve function may be tested using various diagnostic techniques. If the valve function, sizing or securement is not sufficient or ideal the valve may be partially deflated, and advanced or retracted, and then reinflated or the valve may be fully deflated and retracted into the deployment catheter or another slightly larger catheter, and removed. Once a valve is positioned, sized and secured acceptably or ideally the inflation media may be exchanged for a permanent inflation media, which may jell, set or cure. The inflation catheters and deployment control wires are then disconnected and the catheter removed, fully deploying the valve.
If the technology from a known self-expanding recoverable stent is adapted to a valve support structure, the stent is only recoverable from a partially deployed state. A self-expanding support structure of a length sufficient only to support and retain the valve would not allow testing of the valve function, until the valve was fully deployed. This is because the proximal portion of the support structure contained within the device would prevent normal function of the valve. A proximal extension of the support structure could be added to act as a deployment control device allowing the valve function to be tested in a configuration where it is still possible to remove or reposition the valve. The proximal extension could be a continuation of the braided or laser cut stent structure, provided that the cell structure is open enough to allow blood flow through the stent. In an aortic valve application the required length of the proximal extension would most likely extend beyond the ostia of the coronary arteries. In this case the shape of the stent structure may be designed to permit unobstructed flow to the coronary arteries or to permit adequate flow to the coronary arteries. Another possibility is to design the proximal extension so that it acts as multiple individual wires. This could be done by laser cutting or by changing a braid pattern. This would also allow the proximal portion of the implant to act as a deployment control device.
A method for recapturing a self-expanding stent is described by Johnson et al in U.S. Pat. No. 5,817,102, as follows.
There is provided an apparatus for deploying a radially self-expanding stent within a body lumen. The apparatus includes a stent confining means for elastically compressing a radially self-expanding stent into a delivery configuration in which the self-expanding stent has a reduced radius along its entire axial length. The apparatus includes an elongate and flexible stent delivery device having a proximal end, a distal end and a distal region near the distal end. The distal region is used in delivering the radially self-expanding stent into a body lumen, and in positioning at a treatment site within the body lumen with the stent surrounding the delivery device along the distal region. The proximal end of the delivery device remains outside of the body. An axial restraining means is disposed along the distal region of the delivery device. A control means is operable associated with the delivery device and the confining means. The control means moves the confining means axially relative to the delivery device toward and away from a confinement position in which the confining means compresses the self-expanding stent into the delivery configuration, and urges the stent into a surface engagement with the axial restraining means. The restraining means, due to the surface engagement, tends to maintain the self-expanding stent axially aligned with the deployment device as the confining means is moved axially away from the confinement position to release the stent for radial self-expansion.
Preferably the stent delivery device is an elongate and flexible length of interior tubing, with a central lumen for accommodating a guidewire. The stent confining means can be an elongate and flexible length of tubing, having a lumen for containing the interior tubing. The second (or outer) tubing surrounds the stent to confine it.
The preferred axial restraining means is a low durometer sleeve surrounding the interior tubing along the distal region. If desired, an adhesive can be applied to an exterior surface of the sleeves. Alternatively, the axial restraining means can consist of several elongate strips disposed along the distal region, with adhesive applied to radially outward surfaces of the strips, if desired.
In either event, so long as the exterior tubing surrounds the stent to radially compress the stent, it also maintains the stent in surface engagement with the sleeve or strips. As the exterior tubing is axially withdrawn to allow part of the stent to radially self-expand, the rest of the stent remains confined against the sleeve or the strips. As a result, the stent does not travel axially with the exterior tubing. Rather, the stent remains substantially fixed in the axial direction with respect to the interior tubing. This structure affords several advantages. First, the interior tubing can be used as a means to positively maintain the radially self-expanding stent in the desired axial position during deployment. The interior tubing can itself be employed as a reliable indicator of stent position, both prior to and during deployment. Further, should the need arise to retract the stent after a partial deployment, the outer tubing can be moved back into the confinement position, without tending to carry the stent along with it.
The current percutaneous valve replacement devices are not removable or repositionable. These devices are deployed at a location and if the location was a wrong location or if the valve does not have a good effect, the valves can not be removed, recaptured or repositioned percutaneously. The present invention includes a method of implantation facilitating percutaneously repositioning, recapturing and/or removing, a prosthetic valve
A balloon expandable support structure is more difficult to make recapturable, repositionable or removable. One method would be to use a shape memory alloy, such as Nitinol. In this case if Nitinol was used it would be in the martensitic phase at body temperature. Martensitic Nitinol is not superelastic, but soft and conformable. It would be somewhat suitable as a balloon expandable support structure material, except the yield strength is very low. This requires relatively thick cross sections to be used. The balloon expandable support structure is deployed in any way desired, such as by the methods described in Andersen. If the location or performance of the valve is not acceptable the support structure may be caused to contract by changing its temperature, causing it to return to its preset “remembered” shape, which in this case is a smaller, radially collapsed shape. The temperature controlling media could be a fluid such as saline, and could be delivered while a catheter or balloon is inserted through the support structure. This would cause the valve and valve support structure to collapse down on the balloon or catheter allowing removal or possibly redeployment. Other shape memory materials are available, and may have more desirable mechanical properties for use as a balloon expandable support structure. In some cases the biocompatibility of these alloys is not known.
It would be possible to construct a self-expanding valve that would be capable of being recaptured. This could be done using technology from recapturable self-expanding stents. Typically these devices are braided from a superelastic or high strength alloy and have relatively low radial strength. As they are pulled back into a sheath they collapse on their diameter and lengthen facilitating recapturability. Not all braided self-expanding structures are recapturable. To our knowledge this technology has not yet been applied to valve support structures.
Relating to the current inflatable prosthetic valve or cast in place support structure, a different deployment procedure is used which allows the device to be repositionable recapturable, and removable. This deployment method could also be used with some self-expanding percutaneous valve support structures. The deployment method is described as follows for an aortic valve replacement. The procedure could be easily adapted to any other coronary valve. The deployment catheter is advanced across the aortic valve. The prosthetic valve and inflatable cuff are unsheathed in the ventricle, but remain attached to the deployment control wires. The distal end of the inflatable cuff is inflated. The sheath is retracted far enough that the deployment control wires allow the prosthetic valve to function. The device is then withdrawn across the native valve annulus. The device is then fully inflated. The valve function may be tested using various diagnostic techniques. If the valve function, sizing or securement is not sufficient or ideal the valve may be partially deflated, and advanced or retracted, and then reinflated or the valve may be fully deflated and retracted into the deployment catheter or another slightly larger catheter, and removed. Once a valve is positioned, sized and secured acceptably or ideally the inflation media may be exchanged for a permanent inflation media which may jell, set or cure. The inflation catheters and deployment control wires are then disconnected and the catheter removed. This deployment method provides many advantages including the ability to reposition recapture and remove the device.
In an alternative delivery method (surgical) transapical access would allow for the device to be placed in a less invasive surgical procedure. This may still be a beating-heart procedure but would limit the access incision area. Through the apex of the heart a tube may be inserted to introduce the device to the aortic valve from a antigrade approach. This would allow the device to be placed and or moved in the same manner previously described in a catheter delivery.
The prosthetic valve with inflatable cuff may also be delivered surgically. The inflatable cuff aids in sealing the valve to the native anatomy. A valve of this design may be placed in any coronary valve position as well as in a vein, lung, ureter, or any area of the body known to benefit from the implantation of a valve or flow control device. In one embodiment the native valve is sutured in place similar to known coronary prosthetic valves. The inflatable cuff is then expanded to form a tight seal with the native anatomy. In another embodiment the valve is placed in the desired location and the valve is expanded. The valve is held in place by physical interference with the native anatomy. The geometry of the implant may be similar to the percutaneous applications for the inflatable prosthetic valve described in previously.
The valve may be further secured by additional methods such as sutures or staples. The surgical procedure may also be performed in a less invasive manner, for example a smaller opening in the atrium or aorta could be used to implant the valve, because the valve attachment process is less critical. In another embodiment the valve may be implanted with a minimally invasive surgical device. A device of this design for an aortic valve application punctures the chest wall and the ventricular wall near the apex of the heart. The device is then advanced across the native valve annulus and implanted in a manner consistent with the percutaneous embodiments of the invention. This procedure may be guided by echocardiography, angiography, thorascopy or any other appropriate visualization method commonly known.
One Step Implantation
By deploying the device at the site in one step the native valve may be excluded while the new valve is being placed. It is conceived that the device may have a shape similar to a tubular hyperbola to exclude the old valve by trapping it under the new structure during deployment. This may aid in patient comfort and safety if the vessel is not occluded during implantation by a balloon deployed stent system. As the sheathed device is delivered via catheter through the vessel past the aortic valve, it may be reveled or exposed by removing the sheath partially or completely and allowing proper placement at or beneath the native valve. Once in the vessel, the device may be moved proximal or distal and the fluid may be introduced to the cuff providing shape and structural integrity. It may be necessary to add or retract the fluid for proper positioning or removal. Once the cuff is positioned properly and the fluid is added creating the structure and sealing the device to the vessel wall, the delivery catheter may be disconnected and removed leaving the now functioning valve device as a permanent implant. The disconnection method may included cutting the attachments, rotating screws, withdrawing or shearing pins, mechanically decoupling interlocked components, electrically separating a fuse joint, removing a trapped cylinder from a tube, fracturing a engineered zone, removing a collecting mechanism to expose a mechanical joint or many other techniques known in the industry.
Two Step Implantation
It may be desirable to implant the valve structure in two steps. It is desirable to attach the valve to the native tissue securely and without leaks. Also it is desirable to avoid blocking the flow of blood for a long period of time. For these reasons it may be desirable to first implant a retention-sealing device as a first step and then as a second step implant the cuff with the valve attached. The retention-sealing device could be a stent like structure expanded in place or a ring shaped support structure where the valve is secondarily attached. The ring shaped structure could utilize the fluid inflation method as mentioned above and could be a separate system and catheter. It could incorporate barbs for anchoring. It could also incorporate a sealing material to help prevent blood from leaking around the valve. The device could incorporate a mechanism to attach the support structure to. The retention mechanism could be a shoulder or a channel that the support registers in. Once in position, the deployment of the valve could take place as mentioned in the One Step Implantation description above
In an alternative embodiment a support structure, such as a stent is delivered in one step and the valve is delivered in a later step. The valve is then attached to the support structure. The support structure may be an expandable scaffold or stent designed to produce a physical interference with the native vessel. The support structure could also use the geometry of the native anatomy as described in other embodiments, for
Deflate Balloons after Anchoring
In another embodiment the balloon inflation step is used to enable the device and the support structure and anchoring device are delivered in a later step. In one embodiment the support structure is a balloon expandable stent. The stent is placed inside the inflated cuff. The stent may also extend proximal or distal from the cuff. More than one stent can be used. Preferably a stent is placed proximal to the valve portion of the implant and a stent is placed distal to the valve portion of the implant, or a portion of the stent extends across the valve. In one embodiment the balloons are left as part of the implant in a deflated state. The balloons are disconnected from the catheter by a mechanism described in this application with the exception that the sealing feature is not required. Other detachment mechanisms are also possible. In another embodiment the balloon is removed from the device after it is deflated. The balloon may be placed in a channel in the cuff and simply retracted after deflation. Alternatively the balloon may be attached to the implant with sutures designed to break as the balloon is inflated. After the balloon is inflated and deflated the balloon can be retracted.
Stent on Device
A method of delivering a valve attached to a cuff as a first step, and delivering an expandable structure as a second step. The structure may be a stent or an unwrapable band, engaged coaxially inside the cuff. The cuff may be positioned using an inflatable cuff, where the cuff remains inflated after the device is disconnected from the catheter. In this case the inflation serves the function of temporary securement and of permanent sealing. Alternatively the cuff may contain a removable balloon. In this embodiment the inflation provides a means of temporary support until the permanent support structure is deployed. Yet another alternative involves a valve and cuff assembly that contains no inflation provision. The cuff is held in place using deployment control wires that are shaped in a way to cause the expansion of the prosthesis. The stent or expandable support structure is then delivered to a position located coaxially within the cuff. The stent is then deployed, securing the device.
Creating Support Structure In Vivo
The present invention includes a method of creating a support structure inside the body of a patient. The preferred embodiment includes manufacturing the support structure by a casting method. In this method fluid is injected into a mold or cuff that is attached to the valve and delivered percutaneously. The fluid then jells hardens or solidifies forming the support structure.
There are other methods of manufacturing a support structure in vivo. In one embodiment the support structure can be assembled from many small solid particles. The particles can be attached to one another by various means, including a thread woven through the particles, in such a way that when the thread is tensioned the thread and the particles form a rigid structure. The particles could be attached to one another by a sintering process, with an adhesive or by another method. The support structure could also be manufactured in place from wire, which is woven and inserted into the shape of a support structure in vivo.
The support structure could also be manufactured in place using a biological reaction such as forming calcium deposits on the appropriate portion of the valve. The support structure could be assembled by nanomachines.
The support structure could also be manufactured from a fluid that solidifies jells or hardens that is not contained inside a mold. The fluid could be applied to an area on the outer surface of the valve or the inner surfaces oft the area where the valve is to be applied, in vivo. The support structure could be manufactured from a material that solidifies hardens or becomes more rigid by the addition of a catalyst, heat, cold or other energy source. The material could be applied to the outer surface of the prosthetic valve before the valve is installed and then activated in vivo. The support structure could be excited or activated by an electronic energy. This source could also be activated by magnets through a suspension fluid that solidifies in a magnetic field.
Attachment of Valve to Non-Structural Element
In the embodiments described above, the valve can be attached only to a nonstructural element. In the preferred embodiment the nonstructural element is the sewing cuff or mold. The support structure is later manufactured within the mold. Other examples of valves permanently attached only to nonstructural elements are possible. A valve could be attached to an unsupported tubular section of fabric. After the fabric graft and valve are positioned in the patient a stent or other support structure could be deployed within the graft anchoring the graft in place. The stent could utilize barbs or fangs to puncture the graft and anchor the devices solidly to the native tissue. The stent could also be placed so that it only partially overlaps the graft. In this way barbs or fangs could be placed that do not puncture the graft. In another embodiment, rigid structural elements such as commissural support posts or barbs or anchors are attached to the cuff and delivered with the nonstructural element.
Radially Moveable and/or Flexible Tissue Supports
In accordance with another aspect of present invention, there is provided an inflatable or formed in place support for a translumenally implantable heart valve, in which a plurality of tissue supports are flexible and/or movable throughout a range in a radial direction. As used herein, a radial direction is a direction which is transverse to the longitudinal axis of the flow path through the valve.
Valve and valve support design preferably accomplish a variety of objectives, including long term durability of the valve. The inflatable valve support of the present embodiment can be optimized in a variety of ways, to enhance valve life. For example, except at its point of attachment to the annulus, the wall and coaptive edges of the tissue leaflet preferably will not contact any structural components of the implant or other tissue of the valve or surrounding environment. Such contact may result in premature wear and ultimately valve failure. In addition, upon valve closure, the tissue supporting elements of the formed in place support preferably allow for a controlled deceleration of the motion of the leaflets. This lessens the stress seen by the connection points of the tissue to the structural elements. In many valve designs, these support elements are referred to as commissural supports.
The most common valve is a three cusp leaflet configuration where support posts extend axially from the base of the valve in a downstream direction to support the tissue, creating a tricuspid valve. Preferably, under pressure, the leaflets will open and close with the stresses being distributed evenly about the structural element. In this tricuspid design, the forces upon closure of the valve are in an upstream axial direction and radially inwardly on the valve. By allowing the commissural supports to flex inward, the forces seen by the connection between the tissue and the support element will drop and the longevity of the valve may be increased. In conventional surgical valves, deceleration or dampening of the closure force is accomplished by a wire formed stent or a polymer cast or machined to distribute the forces evenly about the stent.
Forces experienced by the valve upon valve closure are on the order of about 15 grams per support post, and prosthetic valve testing may be accomplished up to about 45 grams per support post, for a safety factor of about 300 percent. The supports are preferably movable in a radial direction upon closure of the valve through a range, as is discussed below, to dampen the impact stresses on the valve.
Bending of the tissue supports may include not only flexure of the supports but also flexure of the hoop or base of the valve support. This maximizes the distribution of stresses over the structural element, thereby lessening the stress concentrations at any one point or area.
The tissue supports (i.e., commissural supports) on the inflatable valve support of the present arrangement may be provided with a range of radial direction motion in a variety of ways. Referring to
The illustrated inflatable support 107 is configured for supporting a valve having a three cusp leaflet configuration, as has been discussed. Accordingly, the support 107 is provided with three tissue supports 200. Each tissue support 200 comprises a first inflatable strut 202 which is joined with a second inflatable strut 204 at a downstream apex 206. As will be apparent to those of skill in the art, the first inflatable strut 202 and second inflatable strut 204 may be separate components, or may be a unitary tube, which is bent at an angle to form apex 206. As may be seen in
Referring to
As discussed briefly above, each apex 206 is capable of movement in a radial direction through a limited range of motion. As illustrated in
In operation, forward flow of blood (systole) opens the leaflets in a downstream direction and may press the apex 206 against its outer limit of motion which may be contact with the downstream support ring 108a. Upon valve closure, and under diastolic pressure, the apex 206 is forced radially inwardly through the range of motion 218, to provide a spacing which may be seen in
The tissue support 200 may be configured in a variety of ways, to accomplish a range of radial motion. For example, although the tissue support 200 is illustrated in
Alternatively, the tissue support 200 may comprise a non-inflatable component, such as one or two or three or four or more axially-extending support elements. The support elements may be solid elements, such as wire, ribbon, solid rod, or tubing stock which does not require inflation for its structural integrity. Such support elements may be movably connected or rigidly connected to the upstream support ring 108b in any of a variety of ways, depending upon the construction materials and other design choices.
As an independent variable, the tissue support 200, whether inflatable or not, may permit a range of radial motion, either by flexing about a discrete hinge point, or flexing about a force distribution, or by bending, or all of the above. Referring to
Alternatively, referring to
As a further alternative, the tissue supporting strut 202 may be a single element 218 extending in a downstream direction from the upstream support ring 108b, to a distal (downstream) end 216. In this construction the tissue support 200 is only a single element, as opposed to an apex 206 at the junction of a first and second strut. The upstream end of the tissue support 200 may be connected to the upstream support ring 108b or to an inflatable tissue support. As illustrated in
The tissue support 200 may extend in an axial direction such that the apex 206 or downstream end 216 is positioned approximately at the level of the downstream support ring 108a, as has been illustrated, for example, in
Alternatively, the apex 206 or downstream end 216 may be positioned downstream from the downstream support ring 108a, as illustrated in
In any of the foregoing embodiments, the tissue support 200 may be connected with respect to the downstream support ring 108a in any of a variety of ways, such as through the use of sutures, glue, welding, or other tethered structures. The connection between the tissue support 200 and the downstream support ring 108a may either be rigid, or permit a degree of radial flexibility as has been discussed.
In one implementation, the tissue support 200 is secured with respect to the downstream support ring 108a using a bioabsorbable suture or adhesive, which will maintain the structural orientation of the valve during implantation, inflation and curing of the inflation media. After a period of from about a few hours to 2 or 3 or more days, depending upon the inflation media and the suture materials, the connection between the tissue support 200 and the downstream support ring 108a would dissipate, allowing the tissue support 200 to move radially throughout its predetermined range of motion.
In any of the foregoing embodiments, the geometry of the tissue support 200 may take any of a variety of forms, depending upon the desired performance characteristics. For example, although illustrated in
The cross sectional dimension of the 108a and 108b rings may measure about 2.0 mm to about 4.0 mm in diameter but may also measure 1.0 mm in diameter where the apex inflation channels may measure about 0.7 mm to about 3.0 mm in diameter but preferably about 2.0 mm in diameter. Within these inflation channels are also housed valving systems that allow for pressurization without leakage or passage of fluid in a single direction. The two valves at each end of the inflation channel are utilized to fill and exchange fluids such as saline, contrast and inflation media. The length of this inflation channel 311 may vary depending upon the size of the device and the complexity of the geometry but measures about 10 to 30 cm in length and has a diameter of about 2 to 4 mm with a wall thickness of about 0.0002 to 0.010 inches. The inflation channel material may be blown using heat and pressure from materials such as nylon, polyethylene, Pebax, polypropylene or other common materials that will maintain pressurization. The fluids that are introduced are used to create the support structure where without them the implant is an undefined fabric and tissue assembly. In one embodiment the inflation channels 311 are filled with saline and contrast for radiopaque visualization under fluoroscopy. This fluid is introduced from the proximal end of the catheter with the aid of an inflation device such as an endoflator or other means to pressurize fluid in a controlled manner. This fluid is transferred from the proximal end of the catheter through two inflation tubes 306 which are connected to the implant at the end of each inflation channel 311. With reference to
The end valve system 301 consists of a tubular section 312 with a soft seal 304 and spherical ball 303 to create a sealing mechanism 313. The tubular section 312 is about 0.5 to 2 cm in length and has an outer diameter of about 0.010 to 0.090 inches with a wall thickness of 0.005 to 0.040 inches. The material may include a host of polymers such as nylon, polyethylene, Pebax, polypropylene or other common materials such as stainless steel, Nitinol or other metallic materials used in medical devices. The soft seal material may be introduced as a liquid silicone or other material where a curing occurs thus allowing for a through hole to be constructed by coring or blanking a central lumen through the seal material. The soft seal 304 is adhered to the inner diameter of the wall of the tubular member 312 with a through hole for fluid flow. The spherical ball 303 is allowed to move within the inner diameter of the tubular member 312 where it seats at one end sealing pressure within the inflation channels and is moved the other direction with the introduction of the inflation tube 306 but not allowed to migrate too far as a stop ring or ball stopper 305 retains the spherical ball 303 from moving into the inflation channel 311. As the inflation tube 306 is screwed into the inflation channel check valve (i.e., end valve) 301 the spherical ball 303 is moved into an open position to allow for fluid communication between the inflation channel 311 and the inflation tube 306. When disconnected the ball 303 is allowed to move against the soft seal 304 and halt any fluid communication external to the inflation channel 311 leaving the implant pressurized. Additional embodiments may utilize a spring mechanism to return the ball to a sealed position and other shapes of sealing devices may be used rather than a spherical ball. A duck-bill style sealing mechanism or flap valve would additionally suffice to halt fluid leakage and provide a closed system to the implant.
The various methods and techniques described above provide a number of ways to carry out the invention. Of course, it is to be understood that not necessarily all objectives or advantages described may be achieved in accordance with any particular embodiment described herein. Thus, for example, those skilled in the art will recognize that the methods may be performed in a manner that achieves or optimizes one advantage or group of advantages as taught herein without necessarily achieving other objectives or advantages as may be taught or suggested herein.
Furthermore, the skilled artisan will recognize the interchangeability of various features from different embodiments disclosed herein. Similarly, the various features and steps discussed above, as well as other known equivalents for each such feature or step, can be mixed and matched by one of ordinary skill in this art to perform methods in accordance with principles described herein. Additionally, the methods which is described and illustrated herein is not limited to the exact sequence of acts described, nor is it necessarily limited to the practice of all of the acts set forth. Other sequences of events or acts, or less than all of the events, or simultaneous occurrence of the events, may be utilized in practicing the embodiments of the invention.
Although the invention has been disclosed in the context of certain embodiments and examples, it will be understood by those skilled in the art that the invention extends beyond the specifically disclosed embodiments to other alternative embodiments and/or uses and obvious modifications and equivalents thereof. Accordingly, the invention is not intended to be limited by the specific disclosures of preferred embodiments herein
Claims
1-25. (canceled)
26. A method of implanting a prosthetic heart valve, the method comprising the steps of:
- translumenally advancing a catheter carrying a prosthetic valve that comprises an inflatable support structure to a position proximate a native valve of a patient, wherein the inflatable support structure comprises a directional valve positioned within the inflatable support structure;
- inflating at least a first chamber of the inflatable support structure with a contrast media;
- visualizing the prosthetic valve under fluoroscopy;
- displacing the contrast media in the first chamber of the inflatable support structure with a new inflation media through the directional valve; and
- removing the catheter from the patient, leaving the prosthetic valve and the inflated first chamber of the inflatable support structure within the patient.
27. The method of claim 26, further comprising allowing the new inflation media to solidify within the first chamber of the inflatable support structure.
28. The method of claim 26, further comprising proximally retracting the valve after the first chamber is at least partially inflated.
29. The method of implanting a prosthetic valve as in claim 26, additionally comprising the step of removing the native valve prior to the removing the catheter step.
30. The method of claim 26, wherein the inflatable support structure comprises an inflation channel having a first end and a second end, and the directional valve resides between the first end and the second end of the inflation channel.
31. The method of claim 30, wherein inflating at least a first chamber comprises filling the contrast media from the first end of the inflation channel, and displacing the contrast media comprises filling the new inflation media from the second end of the inflation channel.
32. The method of claim 31, wherein the contrast media is pushed out from the first end of the inflation channel by the new inflation media.
33. The method of claim 32, wherein the new inflation media and the contrast media move in a single direction during the step of displacing the contrast media.
34. The method of claim 30, wherein the first chamber is at least partially inflated with the contrast media.
Type: Application
Filed: Oct 30, 2018
Publication Date: May 30, 2019
Inventors: Gordon B. Bishop (Santa Rosa, CA), Do Uong (Santa Rosa, CA), Randall T. Lashinski (Santa Rosa, CA)
Application Number: 16/175,175
