Patent Application
20090326508
This invention relates to percutaneously delivering oxygen and/or other gases to tissue for the treatment of cardiovascular disease and/or for other treatment purposes. Among other things, the method and apparatus disclosed herein may be used to reduce the risks of ischemic events during an angioplasty procedure and/or a plaque removal procedure, to improve healing of hypoxic tissues, and/or to slow down restenosis after vascular interventions.
A percutaneous transluminal angioplasty (PTA) of blood vessels, including the coronary arteries (PTCA), is a very common procedure to reduce vessel narrowing (i.e., stenosis) that obstructs blood flow to tissue, especially human organs. The angioplasty procedure typically involves inflating a balloon within the constricted region of the blood vessel so as to re-open the blood vessel. The success rates of coronary angioplasty procedures are typically inversely related to (i) the extent of the vascular disease, and (ii) the patient's intolerance to myocardial ischemia (i.e., blood flow obstruction) during the temporary blood vessel occlusion which is associated with a PTA procedure.
More particularly, one of the principle limitations of a coronary angioplasty procedure is the complete obstruction of blood flow during the inflation of the angioplasty balloon. After a short period of balloon occlusion, patients experience myocardial ischemia due to the interruption of oxygenated blood to the myocardium. Myocardial ischemia is usually indicated by angina pectoris and/or cardiac arrhythmias.
In the past, several perfusion balloon catheters have been developed to overcome the problem of total blood flow obstruction during percutaneous coronary interventions. By way of example but not limitation, U.S. Pat. No. 4,944,745 (Sograd) discloses a perfusion balloon catheter that allows passive perfusion of blood through a catheter whose balloon is obstructing blood flow. U.S. Pat. No. 4,909,252 (Goldberger) discloses a perfusion balloon catheter with a central opening which allows blood flow through the catheter when the balloon is fully inflated. U.S. Pat. No. 5,087,247 (Horn et al.) discloses a balloon perfusion catheter with an elongated flexible perfusion shaft, with multiple openings proximal and distal to the balloon, in order to permit blood flow through an artery during balloon inflation. International Patent Publication No. WO 9732626 (Cox et al.) discloses an inflatable balloon envelope allowing blood passage during inflation of the device.
While such perfusion balloon catheters permit some continued blood flow while their balloons are inflated, they are nonetheless limited to a flow rate which is something less than the normal flow rate of the blood passing through the vessel. In other words, perfusion balloon catheters can provide, at best, only some fraction of the normal flow rate which existed in the blood vessel prior to insertion of the catheter and inflation of the balloon. Thus, when perfusion balloon catheters are placed into relatively small arteries (e.g., the coronary arteries) which already have modest flow rates, the further reduction of an already-low flow rate is frequently clinically unacceptable. The inadequacies of the perfusion balloon catheter were characterized in a publication by Ferrari et al. (Coronary Artery Disease, 1997) who conclude their studies with the statement that in “high-risk patients dependent on adequate coronary perfusion, autoperfusion balloons are not able to provide sufficient distal coronary blood flow during balloon inflation”.
Insufficient blood flow distal to an inflated balloon causes ischemia and hence hypoxia (i.e., oxygen deprivation) in tissue (e.g., the end organs) because the oxygenation of tissue previously supplied with blood is reduced.
For this reason, angioplasty in the coronary arteries is a relatively high risk procedure in patients who require dilatation of the unprotected trunk of the left main coronary artery. Tan et al. (Circulation, 2001) concluded that although percutaneous balloon interventions are a generally accepted treatment modality for coronary artery disease, left main PTCA procedures remain a high risk procedure for the patient.
Another limitation of a coronary angioplasty is restenosis. Restenosis after a PTCA procedure has been successfully inhibited by ionizing radiation therapy (i.e., brachytherapy) applied prior to, or shortly after, angioplasty. Thus, vascular brachytherapy using radioactive sources has become a new treatment option to prevent restenosis. More particularly, radioactive stents disclosed in U.S. Pat. No. 5,059,166 (Fischell et al.) and/or radioactive catheters disclosed in U.S. Pat. No. 5,199,939 (Dake et al.) have been used to minimize or eliminate neointimal hyperplasia after angioplasty. However, the logistical complexities of using radiation sources in coronary arteries, and radiation safety issues, have prompted researchers to improve the irradiation technology. To this end, U.S. Pat. No. 5,951,458 (Hastings et al.) discloses a radiation catheter that releases oxidizing agents such as H2O2 to prevent restenosis after a cardiovascular intervention. The method described by Hastings et al. helps to reduce the radiation doses, or treatment times, necessary to prevent restenosis.
Oxygenated perfluorocarbon (PFC) emulsions have been used to treat ischemic and hypoxic disorders. Oxygen-transferable PFC emulsions became known as artificial blood substitutes more than twenty years ago. By way of example but not limitation, in U.S. Pat. No. 3,958,014 (Watanabe et al.) and U.S. Pat. No. 4,252,827 (Yokoyama et al.), perfluorocarbon (PFC) emulsions are disclosed that have a small PFC “particle” size of 0.02 microns to 0.25 microns, and which were injected into the bloodstream. Additionally, U.S. Pat. No. 4,445,500 (Osterholm) teaches that oxygenated perfluorocarbon (PFC) emulsions can be injected into the cerebrospinal pathway to improve aerobic respiration of tissue. Furthermore, U.S. Pat. No. 4,795,423 (Osterholm) discloses an intraocular perfusion with perfluorinated substances to treat ischemic retinopathy.
Unfortunately, clinical experience has shown that the current approaches for using PFCs to oxygenate tissue are highly problematic. More particularly, and as will hereinafter be discussed in further detail, the current approaches for using perfluorocarbons (PFCs) prevent the use of “pure” PFC solutions and, instead, require the use of PFC emulsions. These emulsions themselves introduce a whole new set of problems which effectively limit the clinical use of PFCs in the bloodstream.
More particularly, it has been found that a pure perfluorocarbon (PFC) solution, with or without a “passenger” gas (e.g., oxygen), cannot be safely injected directly into the arterial or venous bloodstream, e.g., using a standard intravenous (IV) line or syringe. This is because introducing pure PFC solutions in this manner creates dangerous (and potentially fatal) embolisms in the bloodstream. These embolisms are created due to the fact that the PFCs are hydrophobic and are not soluble in blood. Thus, when a pure PFC solution is injected directly into the bloodstream (e.g., for hyperoxic medical therapy), the PFC tends to aggregate into relatively large bodies (or “particles”) within the bloodstream. These relatively large aggregations of PFC tend to create embolisms in the bloodstream. For this reason, introducing pure PFCs (with or without a “passenger” gas) directly into the bloodstream, without the provision of some sort of PFC-dispersing mechanism, is not feasible due to the creation of dangerous embolisms.
Furthermore, it is not possible to eliminate the problematic PFC aggregations by simply diluting the PFC with another liquid prior to its introduction into the bloodstream, because the PFCs are not easily soluble in biocompatible fluids (e.g., the PFCs are insoluble in saline). Thus, the PFC tends to re-aggregate even when it is diluted with another liquid, so that the problematic PFC aggregations remain.
As a result, and as noted above, emulsifying agents (such as egg yolk, phospholipids, Pluronic-F68 and other emulsifiers) have been added to the PFC prior to the injection of the PFC into the bloodstream, whereby to “break up” the PFC particles and minimize aggregations of the PFC within the bloodstream. See, for example, U.S. Pat. No. 3,958,014 (Watanabe et al.), U.S. Pat. No. 4,252,827 (Yokoyama et al.), U.S. Pat. No. 4,445,500 (Osterholm) and U.S. Pat. No. 4,795,423 (Osterholm). Thus, with the prior art approach, emulsifying agents are used as a PFC-dispersing mechanism to break up the PFC and prevent the problematic PFC aggregations which can lead to embolisms.
However, clinical studies in humans evaluating such PFC emulsions (e.g., Fluosol and others) have shown that the use of these emulsions, infused into blood with the PFC for hyperoxic therapy, can cause respiratory insufficiency and pulmonary edema (Wall T C et al., Circulation 1994), most likely due to fluid overload and subsequent congestive heart failure. Thus, PFC emulsions can be considered as PFC “particles” (i.e., aggregations) that are accompanied by large quantities of another therapeutic agent (i.e., the emulsifier) which serves to emulsify (i.e., disperse) the pure PFC within the bloodstream. However, these large quantities of additional therapeutic agent (i.e., the emulsifier) in turn significantly increase intravascular volumes and thereby induce unwanted side effects such as respiratory insufficiency and pulmonary edema.
In addition, PFC emulsions are capable of uploading and releasing, per unit of volume, far less oxygen than a pure PFC solution. Thus, where emulsions are added to the PFC in order to avoid the creation of embolisms, it is generally necessary to provide additional systemic oxygenation to the patient via the lung (e.g., by breathing 100% oxygen) so as to create a sufficiently therapeutic oxygen tension of the PFC emulsions (Kim H W et al., Artificial Organs, Vol. 28, No. 9 2004). However, such intensive systemic oxygenation is normally to be avoided clinically, due to the adverse affects of elevated oxygen concentration on the lungs (e.g., oxygen toxicity) (Kim H W et al., Artificial Organs, Vol. 28, No. 9 2004).
Moreover, the use of emulsions to disperse the PFC in blood can also cause allergic reactions in the patient. Mattrey et al. showed that PFC emulsions can cause allergic reactions (Mattrey R F et al., Radiology 1987). More particularly, in an investigation of Fluosol-DA 20% as a contrast agent using Pluoronic-F68 and others as emulsifiers for PFC in humans, it was reported that Fluosol-DA 20% caused allergic reactions which are most likely triggered by complement activation of the substance Pluoronic-F68 (Mattrey R F et al., Radiology 1987). Since pure PFCs are chemically inert and contain no emulsifiers, no allergic reactions are to be expected when using pure PFCs in the blood; thus it has been concluded that it is the presence of the emulsifiers which triggers the allergic reaction in the patient.
For these reasons, using oxygenated PFCs in conjunction with emulsifiers to prevent hypoxia has not heretofore been clinically successful.
Thus it will be seen that pure PFCs (with or without a “passenger” gas) cannot be introduced directly into the bloodstream without also providing some PFC-dispersing mechanism to prevent embolisms. However, it will also be seen that the prior art approach of using emulsions as the PFC-dispersing mechanism for the PFC introduces a whole new set of problems which effectively limit the clinical use of PFCs in the bloodstream.
For these reasons, prior art PFC systems for delivering oxygen to tissue have not heretofore been clinically successful.
The present invention provides a radically new (i.e., non-emulsifier) PFC-dispersing mechanism to permit the introduction of a pure PFC solution in the bloodstream while preventing the formation of large, embolism-inducing PFC aggregations in the bloodstream.
More particularly, the present invention employs a carefully constructed porous membrane (which may also be referred to as a porous substrate) to safely dispense pure, chemically inert PFCs directly into the bloodstream at sufficiently low rates, and in sufficiently small bodies, as to prevent the creation of the aforementioned large PFC aggregations which lead to embolisms.
This carefully constructed porous membrane may be mounted on, and/or disposed within and/or otherwise carried by, a catheter or wire or other intravascular device or structure (e.g., an atherectomy device, a stent, etc.); a pure PFC solution loaded into the porous membrane; and the catheter or wire or other intravascular device or structure advanced into the vascular system of the patient so that the porous membrane is located at a selected site within the bloodstream; whereupon the porous membrane will act as a PFC-dispersing mechanism to dispense the pure PFC solution directly into the bloodstream—in a carefully controlled, highly dispersed manner—so that micro-, nano-, and subnano-sized quantities of PFC molecules safely enter the bloodstream, without the occurrence of large, embolism-inducing PFC aggregations. The pure PFC solution preferably carries a sizable quantity of therapeutic gas (e.g., oxygen) therein, so that the gas-rich (e.g., oxygen-rich) PFC solution can deliver the therapeutic gas to downstream tissue (e.g., for oxygenation purposes.
An important aspect of the present invention is that the porous membrane must be carefully constructed so as to permit the gas-rich (e.g., oxygen-rich) PFC to enter the bloodstream at the appropriate rate. In fact, it has been discovered that it is important to form the porous membrane with a porosity which permits the gas-rich PFC to disperse into the bloodstream in very small volumes, and at a highly controlled rate which is both (i) sufficiently high to provide therapeutic benefit to the patient by the delivery of adequate quantities of therapeutic gas (e.g., oxygen) molecules to tissue, and (ii) sufficiently low so as to avoid the creation of embolisms in the bloodstream, even when using pure PFC solutions.
In practice, it has been discovered that, for a catheter or wire or other intravascular device or structure (e.g., atherectomy device, stent, etc.) placed into an artery having a typical rate of blood flow, forming the porous membrane with a porosity in the range of 0.001-200 microns, and preferably in the range of 20-200 microns, permits appropriate dispersion of the gas-rich PFC into the bloodstream without inducing embolisms.
It has been discovered that a pore size of greater than 200 microns can increase the likelihood of creating embolisms in the bloodstream.
It has also been discovered that a pore size which is too small (e.g., less than 20 microns) can make it difficult to deliver enough gas molecules to a site to provide certain therapeutic benefits. Thus, for example, where it is desired to provide oxygenation therapy in larger diameter blood vessels, it may not be desirable to use a pore size of less than 20 microns, since this may not provide enough oxygen molecules to the downstream tissue. However, where the oxygenation therapy is to be provided in smaller diameter vessels, or where some other, non-oxygenation therapy is to be provided to the patient, smaller quantities of therapeutic gas molecules may be adequate, in which case smaller pore sizes (e.g., 0.001 microns) may be satisfactory.
It has been discovered that, for oxygenation therapy, a pore size of 20-200 microns provides excellent therapeutic benefits while still preventing the creation of embolisms.
The present invention may also utilize the aforementioned porous membrane (which may also be referred to as a porous substrate) to deliver pharmacological agents to tissue, with the porous membrane regulating the rate of delivery so as to avoid overdosing or underdosing of the pharmacological agent.
In one preferred form of the invention, there is provided a system comprising:
a hollow tube having a distal end, a proximal end, and a lumen extending between the distal end and the proximal end;
at least a portion of the tube comprising a porous membrane; and
a pharmacological agent incorporated in the porous membrane;
wherein the porous membrane has a porosity such that:
In another preferred form of the invention, there is provided a system comprising:
a medical wire;
at least a portion of the medical wire comprising a porous membrane; and
a pharmacological agent incorporated in the porous membrane;
wherein the porous membrane has a porosity such that:
In another preferred form of the invention, there is provided a method for treating a patient, comprising:
providing:
loading the pharmacological agent into the porous membrane; and
positioning the tube in the vascular system of the patient so that porous membrane is exposed to blood;
wherein the porous membrane has a porosity such that:
In another preferred form of the invention, there is provided a method for treating a patient, comprising:
providing:
loading the pharmacological agent into the porous membrane; and
positioning the medical wire in the vascular system of the patient so that porous membrane is exposed to blood;
wherein the porous membrane has a porosity such that:
In another preferred form of the invention, there is provided an intravascular treatment device comprising:
an intravascular device having a distal end and a proximal end;
at least a portion of the intravascular device comprising a porous membrane; and
a pharmacological agent incorporated in the porous membrane;
wherein the porous membrane has a porosity such that:
In another preferred form of the invention, there is provided a method for treating a patient, comprising:
providing:
loading the pharmacological agent into the porous membrane; and
positioning the intravascular device in the vascular system of the patient so that porous membrane is exposed to blood;
wherein the porous membrane has a porosity such that:
In another preferred form of the invention, there is provided a intravascular treatment device comprising:
an intravascular device having a distal end and a proximal end; and
at least a portion of the intravascular device comprising a porous membrane;
wherein the porous membrane has a porosity in the range of 0.001-200 microns, in order that when a pharmacological agent is introduced to the porous membrane:
These and other objects, features and advantages of the present invention will be more fully disclosed in, or rendered obvious by, the following detailed description of the preferred embodiments of the invention, which is to be considered together with the accompanying drawings wherein like numbers refer to like parts, and further wherein:
As noted above, it has been found that a pure PFC solution, with or without a “passenger” gas saturation, cannot be safely injected directly into the arterial or venous bloodstream, e.g., using a standard intravenous (IV) line or syringe. This is because introducing pure PFC solutions in this manner creates dangerous embolisms in the blood. These embolisms are created due to the fact that the PFCs are hydrophobic and are not soluble in blood. Thus, when the PFC is injected directly into the bloodstream of a patient, the PFC tends to aggregate into relatively large bodies (or “particles”) within the bloodstream. These relatively large aggregations of PFC tend to create embolisms in the bloodstream. For this reason, introducing pure PFCs (with or without a “passenger” gas) directly into the bloodstream of the patient, without using some sort of PFC-dispersing mechanism to “break up” the large PFC aggregations (or “particles”), is not feasible due to the creation of embolisms.
However, as also noted above, the use of emulsifiers as the PFC-dispersing mechanism introduces a whole new set of problems. Among other things, the use of emulsifiers as the PFC-dispersing mechanism can cause respiratory insufficiency and pulmonary edema, require the use of additional systemic oxygenation via the lung (with the associated risk of oxygen toxicity), and may cause allergic reactions.
Thus, a new PFC-dispersing mechanism is needed in order to permit a pure PFC solution (with or without “passenger” gas) to be safely and efficaciously introduced into the bloodstream.
The present invention provides a radically new (i.e., non-emulsifier) PFC-dispersing mechanism to permit the introduction of a pure PFC solution in the bloodstream while preventing the formation of large, embolism-inducing PFC aggregations in the bloodstream.
More particularly, the present invention employs a carefully constructed porous membrane (which may also be referred to as a porous substrate) to safely dispense pure, chemically inert PFCs directly into the bloodstream at sufficiently low rates, and in sufficiently small bodies, to prevent the creation of the aforementioned large PFC aggregations which lead to embolisms.
This carefully constructed porous membrane may be mounted on a catheter or wire or other device or intravascular structure (e.g., an atherectomy device, a stent, etc.); a pure PFC solution loaded into the porous membrane; and the catheter or wire or other intravascular device structure advanced into the vascular system of the patient so that the porous membrane is located at a selected site within the bloodstream; whereupon the porous membrane will act as a PFC-dispersing mechanism to dispense the pure PFC solution directly into the bloodstream—in a carefully controlled, highly dispersed manner—so that micro-, nano-, and subnano-sized quantities of PFC molecules safely enter the bloodstream, without the occurrence of large, embolism-inducing PFC aggregations. The pure PFC solution carries a sizable quantity of therapeutic gas (e.g., oxygen) therein, so that the gas-rich (e.g., oxygen-rich) PFC solution can deliver the therapeutic gas to downstream tissue (e.g., for oxygenation purposes).
An important aspect of the present invention is that the porous membrane must be carefully constructed so as to permit the gas-rich (e.g., oxygen-rich) PFC to enter the bloodstream at the appropriate rate. In fact, it has been discovered that it is important to form the porous membrane with a porosity which permits the gas-rich PFC to disperse into the bloodstream in very small volumes, and at a highly controlled rate which is both (i) sufficiently high to provide therapeutic benefit to the patient by the delivery of adequate quantities of therapeutic gas (e.g., oxygen) molecules to tissue, and (ii) sufficiently low so as to avoid the creation of embolisms in the bloodstream, even when using pure PFC solutions.
In practice, it has been discovered that, for a catheter or wire or other intravascular device or structure (e.g., atherectomy device, stent, etc.) placed into an artery having a typical rate of blood flow, forming the porous membrane with a porosity in the range of 0.001-200 microns, and preferably in the range of 20-200 microns, permits appropriate dispersion of the gas-rich PFC into the bloodstream without inducing embolisms.
It has also been discovered that a pore size of greater than 200 microns can increase the likelihood of creating embolisms in the bloodstream.
It has also been discovered that a pore size which is too small (e.g., less than 20 microns) can make it difficult to deliver enough gas molecules to a site to provide certain therapeutic benefits. Thus, for example, where it is desired to provide oxygenation therapy in larger diameter blood vessels, it may not be desirable to use a pore size of less than 20 microns, since this may not provide enough oxygen molecules to the downstream tissue. However, where the oxygenation therapy is to be provided in smaller blood vessels, or where some other, non-oxygenation therapy is to be provided to the patient, smaller quantities of therapeutic gas molecules may be adequate, in which case smaller pore sizes (e.g., 0.001 microns) may be satisfactory.
It has been discovered that, for oxygenation therapy, a pore size of 20-200 microns provides excellent therapeutic benefits while still preventing the creation of embolisms.
The present invention uses pure perfluorocarbon (PFC) as a media for delivering therapeutic gas molecules (e.g., O2, NO, CO, etc., or any combination thereof) to cells at a target site. Although any PFC media may be used, PFO (perfluoro-n-octane) is preferred. As used herein, the term “pure PFC” is intended to mean a PFC solution with or without a gas therein, but which does not include emulsifiers therewith. Thus, the term “pure PFC” as used herein is intended to mean non-emulsified PFC. Furthermore, wherever the term “PFC” is used herein, it is intended to refer to pure (i.e., non-emulsified) PFC, unless it is otherwise stated.
The PFC is loaded with the desired therapeutic gas molecules (i.e., the “passenger” gas) until a certain percentage of saturation is achieved (preferably 100%). Preferably, the gas-rich (e.g., oxygen-rich) PFC is produced under normobaric or hyperbaric conditions at a production facility and then stored in a vial until use (e.g., until the gas-rich PFC is loaded into the porous membrane in the operating room).
As noted above, with prior art approaches, introducing a pure PFC solution (with or without gas molecules) directly into the bloodstream (e.g., via a needle) is not clinically acceptable due to the creation of dangerous embolisms. As also noted above, it is not practical to dilute the PFC with another liquid prior to injection, so as to reduce PFC aggregations in the bloodstream, due to the insoluble nature of the PFCs. Furthermore, as also noted above, it is not practical to use emulsifiers to disperse the PFCs within the bloodstream, since the use of emulsifiers can lead to problems of high fluid volume, less efficient oxygen delivery and possible allergic reactions.
The invention described herein overcomes these problems by dispensing a pure, chemically-inert PFC solution (with “passenger” therapeutic gas molecules carried therein) directly into the bloodstream, using a porous membrane (also sometimes referred to as a porous substrate) as a PFC-dispersing mechanism. The porous membrane dispenses the pure PFC solution directly into the bloodstream in a carefully controlled, highly dispersed manner so that micro-, nano-, and subnano-sized quantities of the PFC molecules enter the bloodstream. These tiny quantities of PFC molecules are small enough to avoid the creation of dangerous embolisms in the bloodstream.
Therefore, the present invention provides a unique approach for solving the aforementioned problems associated with prior art PFC delivery and makes it possible—for the first time—to clinically use a pure (i.e., non-emulsified) PFC solution to deliver a therapeutic gas (e.g., oxygen) to treat a medical condition (e.g., to prevent ischemia).
More particularly, the present invention provides a safe and effective way to deliver a gas-rich (e.g., oxygen-rich) PFC solution directly into the bloodstream, without the creation of embolisms, by loading the gas-rich PFC into a porous membrane which is part of a catheter or wire or other intravascular device or structure (e.g., atherectomy device, stent, etc.). The porous membrane is specifically constructed so that the PFCs elute out of the porous membrane, and are dispersed into the bloodstream, in a highly controlled manner, at a reproducible rate, and in small enough volumes, to avoid the creation of dangerous embolisms. This makes it practical, for the first time, to introduce a pure (i.e., non-emulsified) PFC solution directly into the bloodstream, without the risk of embolisms.
To this end, the porous membrane is formed out of a suitable porous material, e.g., Teflon, polyethylene, polyethylene terephthalate, nylon, silicon, cellulose acetate, etc. The porous material has a porosity which permits the gas-rich (e.g., oxygen-rich) PFC to be loaded into the porous membrane outside of the body and then, once the porous membrane is positioned in the bloodstream, to automatically disperse out of the porous material and into the bloodstream in very small volumes, and at a highly controlled rate which is both (i) sufficiently high to provide therapeutic benefit to the patient by the delivery of adequate quantities of therapeutic gas molecules to tissue, and (ii) sufficiently low so as to avoid the creation of fluid overload and/or embolisms in the bloodstream, even when using pure PFCs.
In practice, for oxygenation applications, forming the porous membrane with a porosity in the range of 20-200 microns has been found to permit appropriate dispersion of the oxygen-rich PFC into the bloodstream to adequately oxygenate tissue without causing embolisms. It has been found, however, that a pore size of >200 microns will tend to increase the likelihood of embolisms. Reducing the pore size of the porous membrane to the range of 0.001-20 microns further decreases the size of the PFC particles and hence further reduces the possibility of embolisms. However, it is believed that less PFC can be uploaded (per unit of membrane surface area, per unit of time) when the substrate is nanoporous and, in oxygenation applications, it may be necessary to use larger (e.g., 20-200 micron) pore sizes when the PFC is to be used to oxygenate tissue in larger diameter blood vessels. However, smaller pore sizes (e.g., 0.001-20 microns) may still be satisfactory when the PFC is being used to oxygenate tissue from within smaller diameter blood vessels, or when the therapeutic gas is something other than oxygen.
Further, it is believed that less PFC can be held in a porous membrane with smaller pore sizes than with a porous membrane with larger pore sizes. Thus, where substantial quantities of gas must be delivered to the tissue, and where it is desired to use smaller pore sizes, the overall surface area of the porous membrane may need to be increased, and/or the thickness of the porous membrane may need to be increased, in order to provide an adequate quantity of the therapeutic gas to the tissue.
Thus, a porous membrane formed with an appropriate pore size can be used to dispense the gas-rich (oxygen-rich) PFC into the bloodstream while limiting the size of the PFC aggregations within the bloodstream. For oxygenation applications, a pore size of 20-200 microns has been found to provide excellent therapeutic benefit while still preventing the creation of embolisms.
As noted above, the porous membrane (i.e., the porous substrate) preferably comprises an appropriate polymer. Teflon, polyethylene, polyethylene terephthalate, nylon, silicone, and cellulose acetate, etc. may all be used to form the porous membrane. The porous membrane preferably comprises a hydrophobic material which binds the hydrophobic perfluorocarbon (PFC) solution non-covalently via London forces (named after Fritz London, the German-American physicist). London forces are exhibited by non-polar molecules because electron density moves about a molecule probabilistically. The London forces become stronger with larger amounts of surface contact. Greater surface area contact means closer interaction between different molecules. A porous membrane with a porosity of between 0.001 and 200 microns, and preferably between 20 and 200 microns, offers a sufficient surface area, and is therefore ideal, for PFC applications where the PFC is to be released into the bloodstream in relatively small (i.e., non-embolism-causing) aggregations.
If the hydrophobic (non-polar) porous membrane is brought into contact with the hydrophobic (non-polar) perfluorocarbon (PFC) solution, the contact angle (e.g., wettability) of the pores of the porous membrane is 0°, which means that the PFC solution will be taken up by the porous membrane. In contrast, when the hydrophobic porous membrane is brought into contact with water or saline, the contact angle (e.g., wettability) is about 120°. The water or saline solution will therefore not be taken up by the hydrophobic porous membrane, and the perfluorocarbon (PFC) solution will not be diluted by other fluids, e.g., the water or saline solution.
The carefully-selected porosity of the hydrophobic polymer substrate (i.e., the porous membrane) allows the pure perfluorocarbon (PFC) solution to disperse into the bloodstream in PFC “particles” of micro, nano and sub-nano sizes. Forming the porous membrane out of polymers with a pore size of 0.001-200 microns, and preferably 20-200 microns, provides an effective incorporation of the gas-rich (oxygen-rich) perfluorocarbon (PFC) solution into the porous membrane, and provides a safe and effective rate of dispersion of the PFC solution into blood. A pore size above 200 microns increases the aggregation of the perfluorocarbon (PFC) molecules into the large aggregates that increase the likelihood of creating dangerous embolisms in blood. Therefore, a pore size above 200 microns is generally not preferred in the present invention.
Due to the construction of the porous membrane, predominantly nano- and micro-sized PFC aggregates (or “particles”) are dispersed from the surface of the porous membrane into the bloodstream. In order to achieve a sufficient amount of oxygen delivery into blood so as to create a substantial hyperoxia in the blood for hyperoxic therapy, a sufficient amount of the nano- and micro-PFC particles have to be released from the surface of a catheter or wire or other intravascular device or structure (e.g., atherectomy device, stent, etc.) introduced into the bloodstream. Of course, many different catheter configurations, or wire configurations, or intravascular device or structure configurations, are possible, and many different porous membrane lengths (and/or surface areas) and porosities are possible, so it should be appreciated that variations and combinations of length (and/or surface area)/porosity/thickness may be employed in order to achieve the desired degree of gas deployment without the creation of embolisms. Furthermore, it should be appreciated that many different degrees of gas deployment may be desirable, depending on the therapeutic gas therapy which is to be effected (e.g., oxygenation or otherwise), the size of the blood vessel involved (e.g., larger or smaller), the quantity of tissue to be treated (e.g., oxygenated), etc.
In animal studies using a porous membrane to dispense a pure PFC solution carrying oxygen molecules, the actual pore size of the porous membrane was set to a mean size of 100 microns (range 20-200 microns). Animal studies in rabbits and pigs, studying the safety and efficacy of a catheter comprising a polymer membrane having a mean pore size of 100 microns (range 20-200 microns), clearly indicated that pores in the range of 20-200 microns are capable of delivering sufficient oxygen-rich perfluorocarbon (PFC) particles to blood so as to provide effective hyperoxic therapy. Moreover, in two different animal models of rabbits and pigs, no embolization of the PFC particles was detected in any of the studied animals. Pathology of pig hearts revealed that no perfluorocarbon (PFC) particles could be detected in the small arterioles and capillaries of the heart muscle (i.e., vessels of the end organ), and thus it was concluded that no embolization of the PFC particles had occurred during use of the inventive catheter in blood.
The amount (i.e., the quantity of molecules) of uptake of the gas-rich PFC solution into the porous membrane, and the amount (i.e., the quantity of molecules) of release of the gas-rich PFC solution from the porous membrane into the bloodstream generally depends on the length, the thickness and the porosity of the substrate membrane. The rate of release of the gas-rich PFC solution from the porous membrane into the bloodstream generally depends on the pore size of the porous membrane. Therefore, in order to induce adequate hyperoxic therapy with the present invention, e.g., elevating the oxygen tension of the blood for hyperoxic therapy without inducing embolisms, the pore size of the substrate (i.e., porous membrane) should preferably be in the range of 20-200 microns for blood vessels of a typical size.
The pore size required to achieve the desired rate of dispersion is effectively determined by the size of the PFC molecules, and is not dependent upon the type or concentration of the therapeutic gas molecules which are bound to the PFC. Thus, a catheter having a porous membrane with a porosity of 0.001-200 microns can be used to safely deliver PFC carrying substantially any therapeutic gas molecule (e.g., O2, NO, CO, etc., or any combination thereof), at substantially any percentage of saturation (up to 100% saturation).
As noted above, in practice, it has been found that the pore size of the porous membrane governs the rate of release of the PFC from the catheter. Furthermore, it has been found that the surface area (i.e., length and circumference) and thickness of the porous membrane, together with the pore size, governs the total volume of PFC which may be carried by the device (and hence the total volume of the therapeutic “passenger” gas which may be carried by the medical device).
In one preferred form of the present invention, the porous membrane comprises multiple layers, with the multiple layers being deployed one on top of another.
And in one preferred form of the present invention, the porous membrane comprises multiple layers, with the porosity of the layers varying from one another. More particularly, in one preferred form of the present invention, the innermost layers of the porous membrane (i.e., those lying closest to the center axis of the catheter or wire or other intravascular device or structure) comprise relatively large pore sizes so as to accommodate relatively large amounts of PFC and so as to release that PFC to the outermost layers of the porous membrane as rapidly as the PFC may be accepted by the outermost layers of the porous membrane. At the same time, however, it is preferred that the outermost layers of the porous membrane (i.e., those contacting the bloodstream) be provided with smaller pore sizes (e.g., in the range of 0.001-200 microns, and preferably in the range of 20-200 microns) so as to control the rate of release of the PFC from the catheter in order to avoid the creation of dangerous embolisms.
At the time of use, the catheter (or wire or other intravascular device or structure) is immersed in a vial of gas-rich PFC so that its porous membrane is loaded with the gas-rich PFC, similar to how a sponge is loaded with water. The catheter (or wire or other intravascular device or structure) is then inserted into the vascular system of the patient. Due to the carefully selected porosity of the porous membrane, the gas-rich PFC then elutes out of the porous membrane and disperses into the patient's bloodstream at a rate which limits aggregations of the gas-rich (e.g., oxygen-rich) PFC within the bloodstream to a relatively small size, e.g., 0.001-200 microns. This controlled dispersion of the gas-rich PFC from the porous membrane into the bloodstream prevents embolisms from occurring while still providing sufficient quantities of the therapeutic gas (e.g., oxygen) molecules to provide the desired treatment to the patient. In other words, the porous membrane is carefully engineered so as to elute the gas-rich (e.g., oxygen-rich) PFC at a rate which effectively disperses the PFC in the bloodstream so as to avoid the creation of embolisms. Thus, the present invention permits the direct introduction of pure PFC solutions into the bloodstream, without requiring the use of emulsifiers to avoid the creation of embolisms (and hence without the aforementioned disadvantages associated with the use of emulsifiers).
As the gas-rich PFC travels downstream, most of the gas molecules remain attached to the PFC. Some of the gas molecules may also be released from the PFC into the blood. The gas molecules which are released from the PFC into the blood may or may not be picked up by various blood components (e.g., hemoglobin).
At the target tissue site, the gas (e.g., oxygen) molecules bound to the PFC are released to the cells of the patient's tissue. It will be appreciated that the manner in which the gas molecules are released from the PFC is dependent upon both the hemodynamics of the blood environment and time, in much the same way that oxygen molecules are normally released from the blood components of the patient.
More particularly, the gas-rich PFC enters the target tissue region. Due to the fact that the gas (e.g., oxygen) concentration (“tension”) in the cells is lower than the gas (e.g., oxygen) concentration (“tension”) in the capillary blood, the gas-rich PFC releases the therapeutic gas (e.g., oxygen) molecules. The therapeutic (e.g., oxygen) molecules can then enter the cells of the patient's tissue.
At the target tissue site, the PFC molecules are also available to pick up waste materials (e.g., gases such as carbon dioxide) and carry those waste materials away from the target site, in essentially the same manner that hemoglobin carries away waste materials from the cells. More particularly, the carbon dioxide (CO2) level increases after cellular activity, and therefore the CO2 concentration (“tension”) in the cells is higher than the CO2 concentration (“tension”) in the capillary blood. As a result, the CO2 molecules move from the cells into the capillary blood and become attached to the “gas-poor” PFC, which has previously given up its “passenger” gas (e.g., oxygen) to the cells. The PFC, now loaded with CO2, enters the venous bloodstream and is transported to the lungs, where the CO2 is expelled from the body.
It should also be appreciated that the PFC solution incorporated in the porous membrane need not necessarily carry a therapeutic gas. More particularly, where the primary concern is to remove waste materials (e.g., carbon dioxide) from tissue, the PFC solution loaded into the porous membrane may not be loaded with, or at least may not be completely saturated with, a therapeutic gas. In this case, the gas-poor PFC solution (which is still released safely from the porous membrane without the creation of embolisms) can pick up waste materials (e.g., carbon dioxide) at the tissue and carry it downstream for purging (e.g., by the lungs).
The present invention may be incorporated in various medical devices, in the form of various embodiments, according to the therapy which is to be provided to the patient.
More particularly, in one form of the present invention, there is provided a therapeutic gas delivery apparatus (e.g., a catheter or wire or other intravascular device or structure) for the treatment of disorders (e.g., cardiovascular diseases) that allows the local diffusion of a gas-rich (e.g., oxygen-rich) PFC solution into blood (and/or tissue), whereby to deliver that gas to the blood (and/or tissue). The invention is characterized by a porous membrane which is part of an appropriate medical device, with the porous membrane being impregnated with a gas-loaded (e.g., O2, NO, CO, etc.) perfluorocarbon (PFC) solution, e.g., by the application of a heating or cooling solution, or by utilizing heating or cooling apparatus such as resistance heaters, thermoelectric heaters and/or coolers, etc.).
The release kinetics of the PFC solution from the porous membrane may be modulated by controlled temperature changes of the environment. In other words, the rate of release of the PFC solution from the porous membrane may be modulated by heating or cooling the porous membrane with a warm or cold PFC solution. The preferred cooling temperature is 30° C.-35° C. and the preferred heating temperature is 40° C.-42° C.
The PFC-impregnated porous membrane is preferably sealed in a protective housing made of plastic or metal, allowing the medical device to be pre-loaded with the gas-rich (e.g., oxygen-rich) PFC solution and then stored without the loss of the therapeutic gas and/or the gas-carrying PFC solution.
One of the goals of the present invention is to improve oxygen supply to ischemic organs during an angioplasty procedure. For instance, the present invention may be used to prolong balloon inflation times during high-risk PTCA procedures such as balloon or stent treatment of the trunk of the left main coronary artery. Moreover, the present invention may be used to reduce the extent of acute or subacute myocardial infarction and ischemic stroke. The gas-rich PFC prevents cell death by providing oxygen and other gases, such as nitric oxide (NO) and/or carbon monoxide (CO), thereby preventing excessive inflammation of an organ's tissue. This can be particularly true in infarctions with massive inflammation occurring as a response to tissue damage, where adding small amounts of nitric oxide (NO) and/or carbon monoxide (CO) to oxygen may reduce the negative effects of inflammatory cells such as neutrophils and macrophages. In other words, where infarctions have massive inflammation, providing a PFC solution rich in oxygen and smaller amounts of nitric oxide (NO) and/or carbon monoxide (CO) can have substantial therapeutic benefit. In addition, cooling the treated tissues by injecting a cold fluid (e.g., a fluid having a temperature between 30° C.-35° C.) through the catheter helps to reduce tissue damage in the brain and in the heart in the presence of an ischemic event, thus improving myocardial or cerebral tissue salvage and reducing the risks of infarction.
Alternatively, in the event of a standard percutaneous coronary intervention procedure in patients without serious ischemia or an infarction, fluid temperatures in the range of 40° C.-42° C. may be utilized to increase the release kinetics of the therapeutic gases and to prevent restenosis after angioplasty.
Furthermore, in another embodiment characterized by a setting of cardiac arrest, the present invention may be used to oxygenate the body via the endovascular approach while chest compressions are performed. Thus, the body will be oxygenated without a ventilation of the lung during the resuscitation.
Furthermore, the invention disclosed herein may be utilized to reduce restenosis following an angioplasty procedure.
The invention disclosed herein presents a novel approach for an angioplasty procedure (including a stent implantation) by improving not only the acute safety of the procedure but also the long-term outcome of the procedure.
In a similar manner, the present invention also may be used to prolong procedure times for plaque removal procedure times with atherectomy devices, for example, atherectomy devices that use mechanical blades or laser energy as a means to extract or ablate atherosclerotic plaque within an artery.
A major aspect of the present invention is the local delivery of oxygen (or other therapeutic gases) into blood (and/or tissue) via a perfluorocarbon (PFC) solution delivered via a percutaneously deliverable device. In addition, with the present invention, the local delivery of oxygen (or other therapeutic gases) can be achieved without requiring the use of software, electronic equipment, or mechanical pumping equipment or hardware (e.g., pumps, chambers, computers, bubble detectors, etc.). The gas-rich (e.g., oxygen-rich) PFC is released to the target area from a porous membrane carried by a catheter (e.g., a tube catheter, a balloon catheter, a perfusion balloon catheter, etc.) or a wire (e.g., a coronary wire, a guidewire, etc.) or other intravascular device or structure (e.g., an atherectomy device, a stent, etc.).
The apparatus presented herein allows for the local diffusion of an oxygen-rich PFC solution into hypoxic target tissues, where oxygen is safely released from the PFC into the bloodstream and increases the oxygen tension of the target tissue.
A porous membrane is used to releasably hold the gas-rich PFC on the device. Preferably the porous membrane comprises a polymer. During the manufacture of the porous membrane polymer, the porosity of the basic polymer material is induced in the range of 0.001-200 microns, and preferably in the range of 20-200 microns.
The porous membrane may be formed as an integral part of an appropriate medical device, or it may be securely attached to the medical device, or it may be securely attached to another component which is itself attached to the medical device.
In addition, a surface or portion of the catheter or wire or other intravascular device or structure which itself comes in contact with the bloodstream may be manufactured (e.g., etched or chemically treated) so as to induce the desired porosity on such surface or portion of the catheter or wire or other intravascular device or structure, so as to create the desired porosity in the range of 0.001-200 microns in order to releasably hold and safely disperse the gas-rich (e.g., oxygen-rich) PFC. It should be noted that a catheter and/or wire and/or other intravascular device or structure with a surface so treated so as to create the desired porosity may also be configured so as to further incorporate a porous membrane(s) within one or more lumens of the catheter or wire or other intravascular device or structure.
It is disclosed herein that the microporous material is carried by a medical device, and the medical device is impregnated with a gas-rich (e.g., oxygen-rich) PFC solution. Perfusion channels carrying liquids around the medical device may also be provided to allow the perfusion of warm and/or cold liquids so as to modulate the release of the gas-rich (e.g., oxygen-rich) PFC from the porous membrane. These induced local temperature changes modulate (i.e., increase or decrease) the rate of release of the PFC solution from the porous membrane, whereby to modulate the rate of delivery of the therapeutic gas (e.g., oxygen) molecules to the tissue.
Polymer tubes formed out of a porous structure and/or incorporating a porous material, and impregnated with oxygenated perfluorocarbon (PFC) solutions, may be used to supplement oxygen delivery to the blood during a cardiopulmonary bypass procedure.
Modified stent delivery catheters, (e.g., balloon catheters with a pre-mounted stent), and/or perfusion balloon catheters, and/or wires (e.g., cardiac wires, guidewires, etc.) and/or arterial plaque-removing atherectomy devices, and/or other intravascular devices or structures are all among the preferred embodiments of the invention. The porous membrane may be dispersed substantially anywhere on the medical device, including on an outer surface of the device, an interior surface of the device, and on the outer surface of any balloon carried by those devices. Endovascular stents themselves may also be coated with a thin film porous membrane which incorporates the gas-rich (e.g., oxygen-rich) perfluorocarbon (PFC) solution.
For restenosis prevention, the local delivery of a oxygenated perfluorocarbon (PFC) solution may be combined with the application of ionizing radiation or low energy ultraviolet light so as to increase the production of oxygen free radicals in the target cell of an arterial wall. The effect of increased oxygen free radical production on the proliferating target cell in the arterial wall is DNA damage, which will cause a reduction of restenosis formation.
A therapeutic device that provides local tissue oxygenation may also be applied to other fields of vascular medicine. By way of example but not limitation, wound healing of skin in patients with peripheral occlusive arterial disease and impaired blood flow in the lower limb organs may be significantly improved with the local delivery of an oxygenated perfluorocarbon (PFC) solution via a skin patch placed onto the ischemic skin, where the skin patch includes the porous membrane therein. These oxygenated tissue patches promote the growth of new blood vessels into the area of ischemia, for instance in surgically-opened wounds. Gangrenes of the lower limb due to arteriosclerosis may be reduced in size through the use of the present invention.
By way of example but not limitation, the tissue patch carrying the gas-rich PFC solution may be incorporated in a bandage or other wound dressing.
In addition, these skin patches can deliver therapeutic gases in addition to oxygen.
Furthermore, these oxygenated tissue patches can be used to oxygenate tissue other than skin. By way of example but not limitation, these oxygenated tissue patches can be used to topically apply a gas-rich (e.g., oxygen-rich) PFC to internal tissues (e.g., the intestines), whereby to supply a therapeutic gas (e.g., oxygen) to such tissues.
Thus, in one form of the present invention, there is provided a tissue patch for delivering a therapeutic gas to tissue, wherein the tissue patch comprises a porous membrane which is impregnated with a gas-rich PFC solution.
In another preferred embodiment of the present invention, the porous membrane is located on the surface of a balloon of an angioplasty catheter. The porous membrane comprises a porous polymer, preferably at a thickness of between 0.5-4 mm (and most preferably at a thickness of between 0.6-1.4 mm). Among other things, the thickness of the porous membrane may be constrained by the inner diameter of the guiding catheter used, i.e., in the case of coronary and cerebal artery guiding catheters, the limit of membrane thickness might typically be in the range of 1-2 mm. The porous membrane can be integrated into the balloon, and/or into the catheter shaft structure, or can be wrapped around the balloon and/or the catheter shaft structure. The thin film porous polymer membrane is impregnated with an oxygenated perfluorocarbon (PFC) solution. The porous membrane is preferably sealed within a housing so as to prevent premature release of the gas-rich PFC (and/or the therapeutic gas itself). Prior to the intended angioplasty procedure, the housing is removed from the medical device, and the medical device is then advanced into the bloodstream. At the target site, the porous membrane may be brought into contact with the vessel wall. The release kinetics of the gas-rich PFC may also be modified by changes of local temperature between about 0 degrees Celsius and 50 degrees Celsius, e.g., by the injection of cold and/or warm fluids via the guiding catheter prior to inflation of the balloon. The oxygen enters the blood vessel wall by diffusion. Direct contact of the medical device with the target tissue typically improves oxygen delivery. The local increase in oxygen molecules creates an excess of oxygen free radicals when either (i) ionizing radiation with beta-particle emitters (such as Sr-90/Y-90 or P-32) is applied to the target area, or (ii) ultraviolet light is applied to the target area. A simultaneous application of the oxygenated PFC solution with vessel irradiation (using ionizing radiation or ultraviolet light) is the preferred treatment modality for restenosis prevention.
The oxygen saturation of an end organ increases with improved oxygenated blood flow. Therefore, in another embodiment of the present invention, the oxygenated PFC is released from a perfusion balloon catheter. The perfusion balloon catheter provides for the flow of blood from the proximal end of the occluding balloon into the vascular bed distal to the occluding balloon (i.e., blockage), and thus increases the distribution of the oxygenated perfluorocarbon (PFC) solution to the end organ. Perfusion of blood through the occluded balloon is permitted, and the blood is oxygenated at the proximal end of the balloon, upstream of the balloon (i.e., upstream of the blood flow blockage), so that the oxygenated blood can flow past the balloon to the tissue.
In yet another embodiment of the present invention, the oxygenated PFC is delivered from a porous membrane which is part of a flexible coronary wire or other medical wire device or structure. In a preferred embodiment, the metallic wire is in the form of a flexible hypo-tube, whereby the wire has a lumen that extends from its proximal end through to the distal tip. The porous membrane, which carries the gas-rich PFC solution, is configured such that the porous membrane is positioned inside the lumen, and can extend within a portion of the lumen or from the proximal end all the way to the distal tip of the metallic wire. For example, the porous membrane could be modified to form a thread-like structure or structures. The gas-rich PFC solution is introduced into the porous membrane from the proximal end of the metallic wire through a delivery mechanism (e.g., including but not limited to a syringe) in the appropriate dose or dosages. In addition, the gas-rich PFC solution could be introduced into the porous membrane by a means providing for continuous delivery, which can be a powered device (e.g., including but not limited to an infusion pump) or a passive device (e.g., including but not limited to a gravity-fed drip, much like how a intravenous solution is infused from an IV bag). Then, as the porous membrane is loaded with the liquid oxygen carrier (i.e., the gas-rich PFC) at the proximal end of the metallic wire, capillary action enables the absorption of the gas-rich PFC from the proximal portion of the modified porous membrane through to the distal end of the modified porous membrane, much like dipping the proximal end of a strip of dry facial tissue into water and watching the water being absorbed up into the tissue to the distal end of the strip of tissue. Then, at the tip of the metallic wire where the distal end of porous membrane terminates, the release kinetics of the gas-rich PFC into the bloodstream (as described herein) draws the gas-rich PFC from the porous membrane at the tip of the flexible hypo-tube wire and into the bloodstream. Thus, the gas-rich PFC can be impregnated throughout the length of the modified porous membrane contained within the internal lumen of the wire and then be dispersed out the tip of the flexible metallic wire and into the blood stream, either in dosages or via a continuous flow, at rates which are both (i) sufficiently high to provide therapeutic benefit of the delivery of adequately high therapeutic gas molecules to tissue, and (ii) sufficiently low so as to avoid the creation of fluid overload and/or large particle embolisms in the bloodstream.
In a similar manner, the membrane carrying the liquid perfluorocarbon (PFC) solution can be modified such that the oxygen carrier membrane forms a tube around a retrievable metallic core that is positioned within the wire lumen. The wire containing the core within it can be advanced beyond the lesion (i.e., coronary obstruction) in the distal coronary artery. Then, as the core holding the porous membrane tube within the wire is held at a fixed position, the wire can be retracted an appropriate distance to expose the tube-shaped carrier membrane so as to allow the tube carrying the oxygen source (i.e., the oxygen-rich PFC solution) to dwell in the bloodstream. Thereafter, in all of the aforementioned flexible wire embodiments, a conventional balloon catheter can be advanced over the wire to a treatment zone preferentially proximal to the oxygen delivery source. These aforementioned embodiments permit prolonged balloon inflation as a result of allowing simultaneous oxygen delivery distal to the lesion during balloon inflation, thus eliminating the risk of myocardial ischemia during balloon inflation. Additionally, supplemental oxygen can continue to be delivered after balloon inflation. It should be noted that it may be desirable to advance a balloon catheter over the wire to a treatment zone distal to the oxygen delivery source, depending on the anatomical structure of the blood vessels, so as to allow dispersement of gas-rich PFC before and after balloon inflation.
In yet another embodiment, the distal tip of a coronary wire is coated with the porous membrane carrying the liquid perfluorocarbon (PFC) solution. Alternatively, the porous membrane carrying the gas-rich PFC solution is modified such that the porous membrane forms a tube around the wire. The wire is placed in the distal coronary artery, and the porous membrane is allowed to dwell in the bloodstream so as to dispense the gas-rich PFC solution into the bloodstream. Thereafter, a conventional balloon catheter can be advanced over the wire to a treatment zone, which may be proximal or distal to where the gas-rich PFC was released. If desired, the tubular porous membrane can be withdrawn from the wire prior to advancing the balloon catheter over the wire, or the balloon catheter can be advanced over the tubular porous membrane. In either case, this approach permits prolonged balloon inflation without inducing myocardial ischemia.
In some cases it may be preferably to place the porous membrane inside of the lumen of a guidewire. For oxygen delivery, a length of ePTFE “tubing” is placed inside the wire lumen. Because the wire may be designed with an 0.014″ outer diameter (which is a typical maximum outer diameter of a coronary wire), it may not always be possible to place the ePTFE tube on the outside of the wire in the case where an angioplasty catheter is to be advanced over the wire. Note that a huge clinical advantage can be obtained where a catheter is to be advanced over an oxygen-delivering wire, thus simultaneously providing balloon dilatation and oxygenation delivery distal to the obstruction.
In yet another embodiment of the present invention, the wire is porous. The wire is impregnated with the gas-rich PFC at its distal tip or along its length.
In yet another embodiment of the present invention, the distal tip of the wire forms a plastic thread which is tightly connected to the metallic portion of the wire.
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In practice, for a catheter placed into an artery having a typical blood flow, forming the porous membrane with a porosity in the range of 0.001-200 microns has been found to permit appropriate dispersion of the gas-rich PFC into the bloodstream. However, it has also been found that a pore size of >200 microns will increase the likelihood of embolisms. Thus, it is desired to keep the pore size in the range of 0.001-200 microns. This pore size tends to limit gas-rich PFC aggregations within the bloodstream to a very small size, e.g., 0.001-200 microns, which has been found to provide therapeutic benefit while still preventing the creation of embolisms. For oxygenation applications, the porous membrane preferably has a pore size in the range of 20-200 microns.
The pore size required to achieve the desired rate and volume of PFC dispersion is effectively determined by the size of the PFC molecules, and is not dependent upon the type or concentration of the therapeutic gas molecules which are bound to the PFC. Thus, a catheter having a porous membrane with a porosity of 0.001-200 microns can be used to deliver PFCs carrying substantially any therapeutic molecule (e.g., O2, NO, CO, etc., or any combination thereof), at substantially any percentage of saturation.
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As the gas-rich PFC travels downstream, most of the gas molecules remain attached to the PFC molecules. Some of the gas molecules, however, may also be released from the PFC molecules into the blood. The gas molecules which are released from the PFC molecules into the blood may or may not be picked up by hemoglobin or other blood components.
At the target tissue site, the gas molecules bound to the PFC are released to the cells. It will be appreciated that the manner in which the gas molecules are released from the PFC is dependent upon both the hemodynamics of the blood environment and time, in much the same way that oxygen is normally released from hemoglobin.
More particularly, gas rich PFC enters the target tissue region. Due to the fact that oxygen tension in the cells is lower than the oxygen tension in the capillary blood, the oxygen-rich PFC releases its oxygen molecules. The oxygen molecules can then enter the cells.
At the target site, PFC molecules are also available to pick up waste materials (e.g., gases such as CO2) and carry them away from the target site, in essentially the same manner that hemoglobin carries away waste materials from cells. More particularly, the CO2 level increases in a cell after cellular activity, and therefore the CO2 tension in the cells is higher than the CO2 tension in the capillary blood. The CO2 molecules move from the cell into the capillary blood and become attached to the “gas-poor” PFC (which has previously given up its oxygen). The PFC, now loaded with CO2, enters the venous bloodstream and is transported to the lungs, at which time the CO2 is expelled.
It should also be appreciated that the PFC solution incorporated in the porous membrane need not necessarily carry a therapeutic gas. More particularly, where the primary concern is to remove waste materials (e.g., carbon dioxide) from tissue, the PFC solution loaded into the porous membrane may not be loaded with, or at least may not be completely saturated with, a therapeutic gas. In this case, the gas-poor PFC solution (which is still released safely from the porous membrane without the creation of embolisms) can pick up waste materials (e.g., carbon dioxide) at the tissue and carry it downstream for purging (e.g., at the lungs).
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Significantly, in one preferred form of the present invention, catheter 100 can be placed into a blood vessel and left to dwell there for several minutes before balloon inflation, whereby to permit the tissue downstream of the lesion to be pre-conditioned with a supply of PFC-delivered oxygen. As a result, when the balloon is subsequently inflated, the patient can tolerate “standard” balloon inflation times with less or no pain. In addition, longer periods of balloon inflation can be achieved with less risk of ischemia, less risk of tissue damage, and less risk of arrhythmias that otherwise could result due to hypoxia.
Furthermore, after balloon deflation, the catheter can be maintained in position within the blood vessel so as to continue to deliver oxygen-rich PFC to the downstream tissue and remove waste materials (e.g. CO2), so as to extend the therapeutic event.
If desired, balloon 140 (and preferentially a so-called “Rapid Exchange”, or stent delivery, balloon) may be omitted from shaft 105 of catheter 100.
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In another preferred construction of the present invention, porous membrane 110 may be applied to the walls of balloon 140, in order to deliver oxygen (or another gas) directly to the walls of blood vessel 130. See, for example,
The present invention may be incorporated in still other embodiments.
Thus, for example,
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In one preferred construction, and as shown in
In another preferred construction, porous membrane 415 extends along only a portion of lumen 410. More particularly, in this alternative construction, porous membrane 415 extends from the distal tip of medical wire 400 back along a portion of the length of lumen 410. In this construction, the gas-rich PFC solution can be introduced into the proximal end of lumen 410 (either before or after medical wire 400 is deployed in the patient), whereupon porous membrane 415 will “wick” the gas-rich PFC solution down the remainder of lumen 410 to the distal tip of porous membrane 415, where it is released into the bloodstream of the patient.
The invention described herein consists of a gas (e.g., O2, NO, CO, etc., or a combination of these gases) delivery source for local rescue of ischemic tissue. The invention consists of porous polymer membranes being part of a medical device from which a liquid gas carrier (i.e., the gas-rich PFC) is locally or systemically released. The porous membrane impregnated with the liquid gas carrier may be a part of a tube, a balloon, a perfusion balloon, a stent, and a wire. The porous membrane is preferably sealed with a removable housing to allow storage of the medical device.
The foregoing discussion discloses, among other things, a system for delivering oxygen and/or other gases to tissue using a gas-rich perfluorocarbon (PFC) solution releasably incorporated into a porous membrane which is disposed on an intravascular device.
The following discusses further aspects of the present invention, including how the porous membrane can be used to deliver additional therapeutic agents (e.g., pharmacological agents) to tissue.
Pharmacological agents can be easily over-dosed or under-dosed when injected into the bloodstream to treat arteriosclerosis and/or other forms of coronary artery disease. Over-dosing may result in toxic reactions of a non-target organ, potentially leading to organ failure. Under-dosing may result in a limited drug response or no response, which may lead to the progression of the disease with no beneficial therapeutic effects. Under-dosing easily occurs when the pharmacological agent is injected into the patient's bloodstream and the agent is then diluted as it passes into side branches of the circulatory system. This prevents the pharmacological agent from reaching the target area with a sufficient dosage for the desired therapeutic effect.
Many researchers have previously studied the mechanism of the restenotic process after percutaneous coronary interventions (PCI). The major step by which restenosis (i.e., the repeated narrowing of a vessel lumen) occurs is a repair stimulus of the vascular injury induced during the re-opening of an artery, for instance by means of a balloon inflation. Significantly, after a PCI procedure, the injured vessel wall is more susceptible to the intrusion of pharmacologically active agents than a non-injured vessel wall. Thus, it is not surprising that, after a PCI procedure, the mere injection of a pharmacological agent into the bloodstream can reduce restenosis rates.
Recently, Albrecht et al. published a paper (Invest. Radiol. 2007; 42: 579-585) indicating that the injection of a mixture of the pharmacologically active agent paclitaxel and a contrast agent prevented restenosis after a PCI procedure in pigs. A mixture of contrast agent with paclitaxel was also disclosed in a publication by Speck et al. (ES2289721T). However, under-dosing in the target area, and over-dosing in the non-target area (e.g., the remaining organs of a body), is likely to occur because the injection is systemic and does not limit the therapeutic effect of the drug to only the target area. Therefore, the systemic injection of a pharmacological agent into the circulatory system of the patient is not the preferred method of therapy to prevent restenosis.
Scheller et al. disclosed a paclitaxel-coated balloon (EP Patent No. 1857127). This system is intended to restrict drug delivery to the target area. However, it has been found that particles of the drug coating can be mechanically scraped off the balloon during advancement of the drug-coated balloon through a tight stenosis (i.e., lumen narrowing). In a recent publication discussing use of the Scheller drug-coated balloon in clinical applications, it was shown that only 20% of the paclitaxel mounted on the balloon surface was actually taken up by the target vessel wall. Thus, a disadvantage of the Scheller approach is the significant loss of the therapeutic drug during advancement of the balloon to the target lesion. This can result in under-dosage of the desired drug.
Dommke et al. published a technique for enhancing the local concentration of a pharmacological agent in blood for the reduction of restenosis (Thromb. Haemost. 2007; 98:674-680). The Dommke device employed two balloons to occlude the vessel on either side of the treated restenosis zone, with the pharmacological agent being injected out of the catheter and into the treatment zone between the two inflated balloons. Although this device increases the local concentration of the pharmacological agent in the bloodstream at the target area, it can cause ischemia due to the vessel occlusion from the two balloons. Therefore, the device disclosed by Dommke, while capable of reducing restenosis rates, is not desirable due to the occurrence of ischemia of the heart.
The present invention provides a novel method and apparatus for the controlled delivery of therapeutic agents (e.g., pharmacological agents) in an intravascular approach so as to treat coronary artery disease, among other disorders. In the coronary artery disease application, the novel method and apparatus is configured to achieve regression of the size of arteriosclerotic plaques and to prevent restenosis after percutaneous coronary interventions (PCI), including angioplasty. Moreover, the novel device (i.e., catheter) is preferably specifically designed for local drug treatment of multiple atherosclerotic lesions along the vasculature, e.g., along the length of an injured vessel, during one single drug application procedure.
The novel catheter may also be configured so that it can be used as a stent delivery system by which the stent is placed to complete any type of catheter revascularization of a stenotic artery. In one preferred construction, the porous membrane is disposed proximal to the stent-setting balloon. After the stent has been placed, the catheter is moved distally into the periphery of the vascular bed, so that the porous membrane is disposed substantially adjacent to the just-placed stent or just upstream of the just-placed stent, and then the catheter remains in this position within the vessel for 2-20 minutes in order to allow complete drug elution from the porous membrane into the bloodstream and the vessel wall.
The method and apparatus of the present invention is preferably also configured to facilitate the introduction of a therapeutic agent (e.g., a pharmacological agent) into the bloodstream so that it will reach the target area of blood vessel with the desired dosage and without creating a temporary vessel occlusion, whereby to reduce ischemia of downstream organs.
In one preferred form of the present invention, the apparatus comprises an intravascular device (e.g., a catheter) which includes a porous membrane which incorporates a lipophilic pharmacological agent such that the application of mechanical stress to the intravascular device does not easily remove the pharmacological agent from the porous membrane. Thus, with this form of the invention, the catheter can move through tight spaces within the vascular system of the patient without concern that the lipophilic pharmacological agent will be mechanically “stripped off” the catheter due to engagement of the porous membrane with the side walls of the blood vessels.
In this form of the invention, the porous membrane is a membrane of the type disclosed above, except that it is adapted to release a lipophilic pharmacological agent instead of the gas-rich perfluorocarbon (PFC) solution (which is also highly lipophilic). To this end, the porous membrane is formed with an appropriate porosity such that, for the particular pharmacological agent which is to be delivered, the rate of elution of the pharmacological agent from the porous membrane matches the desired rate of dosage for the pharmacological agent. The intravascular device is preferably configured to remain in the blood vessel for a period of about 2-20 minutes to release the desired amount of the lipophilic pharmacological agent to the target area, e.g., the site of a previous PCI injury, with the lipophilic pharmacological agent preferably being released exclusively from the porous membrane carried by the catheter.
In one form of the present invention, the porous membrane may be disposed on the catheter shaft but not on the surface of the balloon. In another form of the present invention, the porous membrane may be disposed on both the catheter shaft and on the balloon surface. In other words, the surface of the balloon may or may not comprise a porous membrane, as desired. It is even possible that the porous membrane may be provided on the balloon without being provided on the catheter shaft. However, in this respect it should be appreciated that it is generally preferred to place the porous membrane on at least the catheter shaft in order to incorporate and deliver a sufficient dose of the pharmacological agent.
The porous surface (i.e., porous membrane) may be part of any catheter construction as long as the porous membrane is present to bind the lipophilic pharmacological agents by London Forces.
The length of the porous membrane coating disposed on the catheter shaft may vary in accordance with various factors, e.g., the length of the vessel which is to be treated by local drug delivery, the dose of the pharmacological agent which is desired, etc.
The catheter shaft may consist of a multi-layer of different porous membrane polymers to increase the amount of uploaded and releasable drug. In other words, the porous membrane may be formed as a series of layers, one on top of another, and each of the layers may be formed out of identical or different polymers, and/or each of the layers may have different thicknesses, and/or each of the layers may have different porosities, etc.
In the situation where a coronary artery is to be treated, the length of the catheter shaft coating (i.e., porous membrane) may be, by way of example, 3-12 cm, according to the length of the diseased coronary artery. Alternatively, other lengths of porous membrane may be used. In the situation where a peripheral artery is to be treated, the length of the catheter shaft coating (which delivers the lipophilic drug) may vary, by way of example, between 5 and 60 cm. Alternatively, other lengths of porous membrane may be used.
The kinetics by which the lipophilic drug is released from the porous catheter surface depend on the porosity of the porous membrane and the nature of the lipophilic pharmacological agent. The kinetics by which the lipophilic drug is delivered to tissue depend on the blood flow characteristics around the catheter. In other words, a distinction should be recognized between (i) the rate of drug release from the catheter (which is governed by the London Forces reversibly binding the lipophilic drug to the pores of the porous membrane), and (ii) the rate of drug distribution to the tissue (which is governed by blood flow). Thus, by way of example, the greater the blood flow around the porous membrane, the greater the rate of distribution of the eluted drugs to the tissue. Additionally, if the blood flow around the catheter is turbulent, then the lipophilic drug (released from the porous membrane) is distributed more rapidly, e.g., within minutes after the catheter enters the bloodstream. If the porous membrane section of the catheter is withdrawn into a guiding catheter, the release of the lipophilic drug from the porous surface is slowed down because blood flow is reduced while a portion of the porous membrane or all of the porous membrane resides within the tube of the guiding catheter. Therefore, and significantly, the release of the lipophilic pharmaceutical agent can be modified by pushing the porous membrane section of the catheter out of a guiding catheter into the bloodstream and pulling it back into the guiding catheter.
The release of the lipophilic drug from the porous membrane can be further modified by injecting fluids into the guiding catheter while all or a portion of the porous membrane section is located within the tube of the guiding catheter. If these fluids reduce the environmental temperature around the porous membrane section of the catheter, the lipophilic drug will remain longer in the pores of the substrate. However, when the temperature is increased around the porous membrane section of the catheter, then the release of the lipophilic drug will be increased.
In another embodiment of this invention, the porous membrane may be located on the surface of the balloon of a balloon catheter. This porous membrane may be configured as a multi-layer of polymers so as to increase the amount of drug to be uploaded and delivered. When the porous membrane is placed on the balloon, and when the balloon is thereafter inflated, the porous membrane is stretched and the pores of the substrate change their conformity and configuration. This change in pore size of the substrate provokes changes in the adhesion of the London Forces that reversibly bind the lipophilic drug to the pores of the substrate, and hence the drug is released more quickly from the porous membrane into the blood stream. In this respect it should be appreciated that increased lipophilic drug elution is due to London Forces, not mechanical ejection.
If the lipophilic drug is brought into close contact with the injured vessel wall, then the blood-borne elements (e.g., blood substitutes like leucocytes, macrophages) transport the drug into the vessel wall to support the healing process. Blood borne elements carry important information to start repair mechanisms and blood cascades. By way of example but not limitation, blood borne elements are responsible for and able to drive the growth factor movement from the blood borne elements to the tissue. Thus, the movement of the growth factors of the blood-borne elements which enter the vessel wall and start re-establishing regeneration of the lacerated tissue will also drag the lipophilic drug into the vessel wall. Since lipophilic drugs are known to be easily taken up by human cells, the action of the drug is uniform in the vessel wall and yet localized to the site of vessel injury.
In the preferred forms of the present invention, the apparatus comprises an intravascular device (e.g., a catheter) which includes a porous membrane which incorporates a lipophilic pharmacological agent. The porous membrane is of the type disclosed above, except that it is adapted for the controlled release of the lipophilic pharmacological agent instead of the gas-rich perfluorocarbon (PFC) solution. To this end, the porous membrane is formed with an appropriate porosity such that, for the particular pharmacological agent which is to be delivered, the rate of elution of the pharmacological agent from the porous membrane matches the desired rate of dosage for the pharmacological agent.
More particularly, a porous device surface (e.g., a porous membrane carried on a catheter surface) incorporates a lipohilic pharmacological agent and binds the substance reversibly to the porous membrane by the same London forces described above with respect to PFC. The lipophilic pharmacological agent may be a statin (such a simastatin, cerivastatin, lovastatin, pravastatin, etc.), a mitose-inhibiter such as paclitaxel, or an immunosuppressant such as sirolimus, tacrolimus, pimecrolimus, zotarolimus, etc. The rate of release of the lipophilic pharmacological agent from the porous membrane into the bloodstream is controlled by the pore size of the porous membrane. More specifically, the rate of release of the lipophilic pharmacological agent from the porous membrane is regulated by substantially the same release mechanisms discussed above with respect to PFC, except that the pore size is instead coordinated with the characteristics of the specific lipophilic pharmacological agent in order to achieve the desired rate of release, which is preferably in the range between 1 μg/mm2 to 100 μg/mm2
The rate of release of the pharmacological agent from the porous membrane may also be affected by temperature and/or the local fluid dynamics surrounding the porous membrane. Thus, the intravascular device may also include structure for modifying temperature (e.g., a heated or cooled fluid flush) and/or modifying local fluid dynamics (e.g., a chemically-influencing fluid solution).
The porous construction of the membrane carried by the intravascular device incorporates the pharmacological agent in such a way that the agent cannot be dislodged, or otherwise lost, from the intravascular device due to engagement with vascular structure (e.g., when moved through small diameter vessels and/or a tight stenosis). The profile of the intravascular device with porous membrane preferentially is sized to enable placement of the porous membrane in close proximity to the location of the vessel wall that has been treated during the PCI procedure.
In one preferred construction, the intravascular device comprises a balloon catheter, with the porous membrane being disposed distal to, or proximal to, the balloon. After treatment (balloon inflation) of a stenosis, the uncoated balloon portion of the catheter is moved distal to the area of the vessel which has been treated. The treatment of stenosis (PCI) can be repeated several times and can be combined with a stent implantation. After the PCI procedure has been completed, however, the catheter is not removed from the body. The deflated balloon portion of the catheter is located more proximally, beyond the treated area of the blood vessel, allowing the porous membrane, which is located distal to the balloon, to reside in closer proximity to the treated vessel wall tissue. The catheter remains temporarily within the bloodstream in the vessel for about 2-20 minutes. During this time, the pharmacological agent elutes from the porous membrane at the appropriate rate of release and makes its way downstream to the treated tissue. Significantly, the balloon of the catheter remains deflated during this local drug delivery. Since the balloon typically is inflated during PCI under high atmospheric pressure (often multiple times), it is stretched and thus increases its natural profile while dilating the vessel wall. This used, higher-profile balloon helps reduce blood flow proximally of the target area (i.e., the site of the previous PCI procedure). In other words, the increased profile of the deflated balloon on the catheter reduces blood flow in the target area or areas of the previous PCI procedure, without completely occluding the vessel and obstructing blood flow. This reduction in blood flow is sufficient for the pharmacological agent to elute from the porous membrane and dwell in the target area of the previous PCI procedure (i.e., the site or sites of the vessel injury), which enhances absorption of the lipophilic drug by the injured vessel wall while the catheter is indwelling during the above mentioned 2-20 minute time period. Importantly, this preferred construction does not require inflating a balloon at the site of the vessel injury to a diameter greater than the inner diameter of the vessel, thus having to directly contact the vessel wall in order to “press” or “push” the drug into the vessel wall. This is a significant advantage of this preferred construction, since (i) it eliminates the risk of further vessel wall injury from additional balloon dilatation, or dilatations, for the purposes of drug delivery, especially when predilatation with regular PCI balloon catheters (i.e., balloon catheters without drug delivery capabilities) is performed at the lesion site(s), (ii) less precision is required for catheter placement compared to drug eluting balloons, which must be carefully positioned to avoid missing the target site, (iii) there is less risk of ischemia and/or arrhythmias since balloon inflation is avoided and therefore the blood vessel is not occluded during drug delivery, (iv) a single catheter incorporating a standard balloon and the porous membrane can treat multiple lesion sites that are located proximal to the deflated balloon and in close proximity to and/or downstream from the membrane, and (v) the porous membrane can be constructed to deliver a wider range of types of drugs and/or drug delivery rates and dosages.
First, the lipophilic pharmacological agent is dissolved in alcohol (e.g., methanol). Then, the alcohol-pharmacological agent mixture is incorporated in the porous membrane by dipping or immersing the intravascular device, or, alternatively, the portion of the intravascular device incorporating the porous membrane, into the mixture of alcohol and pharmacological agent. Thereafter, the intravascular device, or the portion incorporating the porous membrane, is removed from the mixture and air-dried so as to allow the alcohol to dissipate from the porous membrane. At this point, only the lipophilic pharmacological agent remains in the pores of the porous membrane. The rate and quantity of the uptake of the lipophilic pharmacological agent into the porous membrane depends upon (i) the pore size of the porous membrane, (ii) the concentration of the pharmacological agent in the mixture, and (iii) the molecular weight of the pharmacological agent. Once the porous membrane is loaded with lipophilic pharmacological agent, the intravascular device is packaged, sterilized and subsequently stored at room temperature (21° C.) until clinical use. Alternatively, the porous membrane may be loaded with the aforementioned alcohol-lipophilic pharmaceutical agent mixture after the intravascular device incorporating the porous membrane is packaged and sterilized. In this case, the sterilized intravascular device incorporating the porous membrane may be removed from its packaging using a standard sterile technique. The porous membrane may then be immersed or dipped one or more times into the aforementioned alcohol-lipophilic pharmaceutical agent mixture in order to load the porous membrane immediately prior to clinical use.
When the intravascular device enters the bloodstream, the pharmacological agent begins to elute from the porous membrane.
Significantly, since the bloodstream is at body temperature (37° C.), the difference in temperatures between the porous membrane and the bloodstream increases the rate of release of the pharmacological agent from the porous membrane.
A further increase in temperature from 37° C. to 40° C. can further increase the rate of release of the pharmacological agent into the bloodstream.
In a preferred approach, the pharmaceutical agent is carried downstream from the porous membrane to the target treatment area of the vessel prior to performing a percutaneous coronary intervention (PCI). In this case, blood flow is reduced by the narrowed vessel at the site of the untreated lesion. This reduction of blood flow results in greater dwell time of the released pharmaceutical agent at the target treatment area prior to PCI balloon dilatation, thus allowing the pharmaceutical agent to penetrate the tissue and therefore pre-treat the target tissue prior to balloon inflation and/or stent delivery. Then, the interventional device may be advanced within the blood vessel to the point where the balloon on the interventional device is placed across the lesion. The balloon may then be inflated to dilate the lesion to restore more normal blood flow.
Upon deflation of the balloon, the interventional device may thereafter be further advanced within the blood vessel past the treatment area. As described above, the now-enlarged deflated balloon serves to restrict the flow of blood carrying the pharmaceutical agent to allow further penetration of the pharmaceutical agent into the treated target tissue.
It should be appreciated that the present invention allows pre-PCI and/or post-PCI drug delivery using a single interventional device without totally occluding the blood vessel during therapeutic drug delivery, thus significantly reducing the perioperative risks of ischemia, arrhythmias, or myocardial infarction during therapeutic drug delivery.
As noted above, the intravascular device may be configured to comprise structure for modifying local fluid dynamics. More particularly, the intravascular device may be surrounded with a tube or guiding catheter filled with a modulating fluid which can be used to modify local fluid dynamics.
At the target area, the fluid may be injected from the intravascular device through the surrounding tube and into the bloodstream. The injection of this modulating fluid at the site of the treated vessel changes the fluid dynamics surrounding the porous membrane and therefore increases the rate of release of the lipophilic pharmacological agent into the bloodstream.
It is to be understood that the present invention is by no means limited to the particular constructions herein disclosed and/or shown in the drawings, but also comprises any modifications or equivalents within the scope of the invention.
This patent application: (i) is a continuation-in-part of pending prior U.S. patent application Ser. No. 12/321,964, filed Jan. 27, 2009 by Christoph Hehrlein et al. for DELIVERY SOURCE OF OXYGEN (Attorney's Docket No. OXIRA-1 CON); (ii) is a continuation-in-part of pending prior U.S. patent application Ser. No. 12/008,130, filed Jan. 9, 2008 by Christoph Hehrlein et al. for METHOD AND APPARATUS FOR DELIVERING OXYGEN AND/OR OTHER GASES TO TISSUE (Attorney's Docket No. OXIRA-5); and (iii) claims benefit of pending prior U.S. Provisional Patent Application Ser. No. 61/128,965, filed May 27, 2008 by Michael Braun et al. for METHOD AND APPARATUS FOR DELIVERING OXYGEN AND/OR OTHER GASES AND/OR PHARMACOLOGICAL AGENTS TO TISSUE (Attorney's Docket No. OXIRA-6 PROV). The three above-identified patent applications are hereby incorporated herein by reference.
| Number | Date | Country | |
|---|---|---|---|
| Parent | 12321964 | Jan 2009 | US |
| Child | 12472759 | US | |
| Parent | 12008130 | Jan 2008 | US |
| Child | 12321964 | US |
