|Publication number||US20040142016 A1|
|Application number||US 10/751,706|
|Publication date||Jul 22, 2004|
|Filing date||Jan 5, 2004|
|Priority date||Jun 26, 2001|
|Also published as||CA2451255A1, CA2451255C, EP1401511A2, US8101196, US20030021762, US20110274821, WO2003000305A2, WO2003000305A3|
|Publication number||10751706, 751706, US 2004/0142016 A1, US 2004/142016 A1, US 20040142016 A1, US 20040142016A1, US 2004142016 A1, US 2004142016A1, US-A1-20040142016, US-A1-2004142016, US2004/0142016A1, US2004/142016A1, US20040142016 A1, US20040142016A1, US2004142016 A1, US2004142016A1|
|Inventors||Ajay Luthra, Shivpal Sandhu, John Hudson|
|Original Assignee||Biointeractions, Ltd., University Of Reading|
|Export Citation||BiBTeX, EndNote, RefMan|
|Patent Citations (11), Referenced by (11), Classifications (33)|
|External Links: USPTO, USPTO Assignment, Espacenet|
 The present application is a continuation-in-part of U.S. application Ser. No. 10/179,453, filed Jun. 25, 2002, which claims priority to U.S. provisional patent application serial No. 60/301,176, entitled “Polysaccharide Biomaterials and Methods of Use Thereof”, filed Jun. 26, 2001, all of which applications are claimed as priority documents and which are hereby incorporated herein by reference.
 Further, certain other commonly owned and assigned patent applications are hereby incorporated by reference herein: U.S. patent Ser. No. 10/179,453, filed Jun. 26, 2002; Ser. No. 10/467,950, filed Aug. 15, 2003; and No. 60/451,333, filed Feb. 28, 2003.
 Some aspects of the invention relate to the field of making and using polysaccharides and polymers of polysaccharides as layers on medical devices.
 Many synthetic materials have been medically used in the body, including polyester (e.g., DACRON™), polyethylene (e.g., milk jugs), and fluorocarbons (e.g., TEFLON™), and metals. A patient's body responds by treating a synthetic material as an invader, although it responds only mildly in some medical applications; for example, a metal hip implant is generally well tolerated. One common response to an implant is called a foreign body response in which the body forms a capsule of cells around the material; the body's response to a splinter is a foreign body response. When synthetic material is used as an artificial blood vessel, for example, the blood that flows through the artificial blood vessel reacts with the synthetic material. The reaction can cause clots to form that flow downstream that may eventually become stuck in a smaller vessel; if this happens in the brain, it is called a stroke. The blood clot can also grow on the inside of the vessel and block or severely restrict the blood flow. Blood's clotting mechanism is highly reactive and, despite years of medical research, no implantable blood-contacting synthetic material has yet been found that does not cause blood to react. Synthetic, as used herein, means not naturally found in nature and does not refer to the process whereby an object is made.
 Tubes made of polyester or fluorocarbons are currently used as large diameter blood vessels. The blood clots onto the interior walls and reduces the inside diameter of the vessel but the blood flow is not unduly reduced. The blood clot on the tubular wall serves as a protective layer that elicits very little reaction from blood flowing through the vessel. This approach, however, does not work for small diameter blood vessels because the small diameter tubes are blocked when the blood clots onto the walls.
 There are no currently known materials and techniques for manufacturing small diameter vascular grafts made of synthetic materials. Unfortunately, there is a great need for such grafts. One example is the condition called deep vein thrombosis wherein a patient's veins become blocked. The blood drains poorly from the leg and amputation can result. Unlike heart bypass surgeries where the patient often has some blood vessels that can be harvested from elsewhere in the body and sewn into place, there are few choices for replacing the long veins of the leg.
 One area of synthetic biomaterials research has focused on hydrogels, materials that have a high water content and are soft and slippery. Soft contact lenses are examples of hydrogels. The materials used to make a pure hydrogel might all dissolve in water but the hydrogel itself does not dissolve in water because the materials are cross-linked; in other words, the individual molecular chains are linked together like the strands in a net or a spider's web. Hydrogels tend to elicit a milder foreign body response than other synthetic materials. Some hydrogel biomaterials that are currently considered to be commercially useful are made from polyethylene glycol (PEG), hyaluronic acid, and alginates. Although hydrogels tend to elicit less blood clotting than other synthetic materials, hydrogels have not previously been successfully used to make a small diameter vascular graft.
 Scientists have also tried to use heparin to coat the inside of vascular grafts made of synthetic materials. Heparin is a molecule that belongs to a group of molecules called polysaccharides that are polymers made from combinations of sugar monomers. There are many sugars; glucose and sucrose (table sugar) are two examples. Polysaccharides are naturally-occurring polymers. The terms polysaccharide and mucopolysaccharide are used interchangeably herein. Polymers are molecules built up by the repetition of smaller units that are sometimes called monomers. Polymers are typically made by special chemical schemes that make the monomers chemically react with each other to form molecular chains that can range in length from short to very long molecules. Polymers can be assembled into larger materials; for example, many polymers may be linked together to form a hydrogel.
 Heparin is a polysaccharide polymer with an important property: it interferes with key molecules in the blood clotting mechanism such that the blood will not clot. Coating the inside of a synthetic material tube with heparin tends to increase the amount of time that the tube remains open to blood flow but, to date, small diameter vascular grafts coated with heparin have failed to resist blockage by blood clots for a medically useful length of time.
 Heparin has been applied to materials in many ways. General strategies include letting it naturally stick to a surface (termed adsorption), making a charge-charge bond with the surface (e.g., an ionic bond), and attaching it via an even stronger, more permanent chemical linkage such as a covalent bond. Heparin has been applied as a thin coating of polymers adsorbed to a surface by dipping the surface into a solution of heparin or drying the heparin onto the surface. Heparin has a negative charge and has been exposed to surfaces that have a positive charge so that it remains there via a charge-charge interaction. Photoactivated chemical groups have been put onto heparin so that the heparin is put close to the surface, the surface is bathed in light, and the photoactive groups make permanent covalent chemical bonds between the heparin polymer and the surface. Similarly, heparin has been chemically attached to monomers that have then been reacted with the surface.
 Patent families and patent applications that describe the use of heparin include U.S. Pat. No. 6,127,348, which include descriptions of cross-linked alginate and certain other polysaccharides as compositions useful for inhibiting fibrosis. U.S. Pat. No. 6,121,027 includes descriptions of decorating heparin with a photoactive cross linking chemical group. Application PCT GB9701173 and U.S. Pat. No. 6,096,798 include descriptions of heparins with monomers used to make polymers. U.S. Pat. Nos. 5,763,504 and 5,462,976 include descriptions of glycosaminoglycans derivatized with photoactive groups and cross-linked thereby. U.S. Pat. No. 6,060,582 includes descriptions of macromers with a water soluble region, a biodegradable region, and at least two free-radical polymerizable regions. Other patents include descriptions of a polysaccharide reacted with other polymers, decorated with a polymerizable group, and/or reacted to form a coating on a surface; including U.S. Pat. Nos. 5,993,890; 5,945,457; 5,877,263; 5,855,618; 5,846,530; 5,837,747; 5,783,570; 5,776,184; 5,763,504; 5,741,881; 5,741,551; 5,728,751; 5,583,213; 5,512,329; 5,462,976; 5,344,455; 5,183,872; 4,987,181; 4,331,697; 4,239,664; 4,082,727; and European patents 049,828 A1 & B1.
 Despite many years of research in the areas of polysaccharides, hydrogels, and blood-contacting materials, the need for better implantable synthetic materials that cause little or no unfavorable reaction from a patient's body remains acute. In particular, there is a great need for a medically useful small diameter vascular graft made of synthetic materials.
 Materials and methods disclosed herein are provided to meet all of these needs by providing synthetic materials that successfully combine the advantages of polysaccharides and hydrogels in a medically useful manner so that they may be used for devices, including small diameter vascular grafts. Other embodiments include hydrogels made of polysaccharides and materials and methods for making such hydrogels, as well as products that incorporate such hydrogels. Another embodiment is a hydrogel made of a polysaccharide, for example heparin. Another embodiment is a hydrogel made by polymerizing heparin macromers. Another embodiment is a small diameter vascular graft made by polymerizing heparin macromers around a tube, for example a polyester tube.
 Another embodiment is a medical apparatus having polysaccharide macromers polymerized into a three-dimensional crosslinked hydrogel that makes a hollow cylinder; with the cylinder being formed during polymerization of the polysaccharide macromers. A hollow cylinder is essentially equivalent to a tubular structure and may be made out of any material, e.g., metals, plastics, ceramics, having a variety of properties, including e.g., rigid, compliant, and elastic.
 Another embodiment is a biocompatible encapsulation for an inert medical device. The encapsulation has polysaccharide macromers polymerized into a three-dimensional crosslinked hydrogel that encapsulates the medical device.
 Another embodiment is a medical apparatus made of a material that has a heparin macromers polymerized into a three-dimensional crosslinked hydrogel that forms a hollow cylinder that is not covalently bonded to another material.
 Another embodiment is a pollyvinylpyrrolidone macromer. Another embodiment is a biocompatible coating system that has polyvinylpyrrolidone macromers polymerized into a three-dimensional crosslinked material that contacts a medical device and thereby forms the coating.
 Another embodiment is a polysaccharide polymer of at least two polysaccharide macromers polymerized together. The polymer is preferably in an isolatable form. The macromers may have polymerizable moieties such as polyhydroxyethylmethylacrylates, methyl methacrylates, methacrylates, acrylates, photopolymerizable monomers, monomers with hydroxyl groups, monomers with glycerol groups, monomers with polyoxyalkylene ether groups, monomers with polypropylene oxide groups, monomers with vinyl groups, monomers with zwitterionic groups, monomers with silicone groups, monomers having sulphate groups, monomers having sulphonate groups, and heparin monomer.
 Another embodiment is a method of making a medical apparatus from a material that includes a plurality of polysaccharide macromers by polymerizing the macromers into a three-dimensional crosslinked hydrogel that defines a hollow cylinder, wherein the cylinder is formed during polymerization of the polysaccharide macromers.
 Another embodiment is a method of encapsulating an inert medical device by polymerizing a plurality of polysaccharide macromers into a three-dimensional crosslinked hydrogel that encapsulates the medical device.
 Another embodiment is a method of making a medical apparatus by polymerizing heparin macromers into a three-dimensional crosslinked hydrogel and thereby making a hollow cylinder having an exterior, wherein the cylinder is formed during polymerization of the heparin macromers and the exterior is not covalently bonded to another material.
 Another embodiment is a method of making polyvinylpyrrolidone polymers from polyvinylpyrrolidone macromers. Another embodiment is coating a medical device with polyvinylpyrrolidone macromers by polymerizing the macromers into a three-dimensional crosslinked polyvinylpyrrolidone material, and applying a coating that includes the crosslinked polyvinylpyrrolidone material onto the medical device.
 Another embodiment is a method of making a polysaccharide polymer by obtaining or making polymerizable polysaccharide macromers, synthetically polymerizing the macromers with each other to form a group of polymers having an average length of at least two macromers per polymer, and isolating the polymers.
 Another embodiment is a method of making a coating on a medical device by providing a group of polysaccharide polymers having an average length of at least two macromers per polymer, putting the polymers in a solvent to make a mixture, and contacting the medical device with the mixture.
 Certain embodiments include a material made of a hydrogel that preferably has at least 5% polymerized polysaccharide macromers by dry weight. The hydrogel is preferably covalently cross-linked such that the hydrogel remains intact in water and preferably contains at least 30% water by total weight when hydrated. The polysaccharide macromers are polymerizable while in a solution or in a suspension. Normal polymerization techniques, including free-radical, addition, and condensation polymerization, may be used to polymerize the polysaccharide macromers.
 One product is a tubular member with its inner wall and outer wall covered with a hydrogel as described herein, i.e., the tubular member is “encapsulated” by the hydrogel. A preferred macromer formulation is made from heparin, the term heparin including all molecular weights of heparin, heparan sulfate, heparan sulfate proteoglycans, fragments thereof, and/or derivatives thereof. The preferred embodiment of the heparin hydrogel is at least 80% heparin by dry weight. The tubular member preferably has a diameter of less than approximately 6.0 mm when the hydrogel is hydrated and blood is flowing though the tubular member. The tubular member may be a simple plastic extrusion or a stent, but the preferred embodiment is a knitted or woven fabric substrate. The fabric substrate is preferably pre-coated to enhance the integrity and adhesion of the encapsulant and/or improve the non-thrombogenic or anti-thrombogenic properties or both of the encapsulant.
 In an embodiment of the tubular member, the tubular member is pre-coated with a very thin layer of the hydrogel containing non-thrombogenic or anti-thrombogenic properties or both, the layer being applied to cover the components of the tubular member. In the case of the knitted or woven tube, these are the individual strands from which the fabric is made. The tubular member preferably has a low porosity such that blood leakage is not a paramount concern.
 In another embodiment, a porous tubular member, such as one made from a fabric, is used as the tubular member. The fabric tubular member is pre-coated with a polymeric material in order to prevent blood leakage. The coated fabric tubular member is then further coated with the hydrogel containing non-thrombogenic or anti-thrombogenic properties or both. This structure imparts an extremely thin, flexible, and compliant wall that can serve as a vascular prosthesis, especially in the context of a small diameter vascular graft.
 Another embodiment is a tissue engineering matrix made from a polysaccharide hydrogel. A tissue engineering matrix is, for example, a three-dimensional material that serves as a scaffold for cellular invasion or a nerve growth matrix. Examples of tissue engineering matrices include matrices for making cartilaginous body parts such as ears or joint cartilage; ligaments; scaffolds for breast tissue invasion; liver matrices; and tissue engineered blood vessels.
 The present inventors have also recognized that there is a need to use better organic solvents to dissolve polysaccharides, including heparin. The use of better organic solvents allows scientists to use chemistries and chemical techniques that are more powerful than those that are conventionally used. These techniques improve the cost, quality, and efficiency of conventional techniques for making materials from polysaccharides and enable better materials to be made.
 Embodiments include the use of low dielectric organic solvents and/or low boiling-point solvents for polysaccharide chemistries to make derivatives of polysaccharides, including attaching monomers to polysaccharides to make polysaccharide macromers and the use of these improved solvents for making polysaccharide hydrogels from polysaccharide macromers or polysaccharide polymers. Further, some embodiments include steps for using salts to decomplex the quaternary ammonium-heparin complex.
 Another embodiment is a method of making a polysaccharide macromer, for example from heparin. The polysaccharide is reacted with a quaternary ammonium salt to form a polysaccharide-quaternary ammonium salt complex and then dissolved in an organic solvent with a dielectric constant less than the dielectric constant of DMSO and/or in an organic solvent with a boiling point less than DMSO. The polysaccharide-quaternary ammonium salt complex may be reacted with a chemical such as a monomer to form useful derivatives. The polysaccharide-quaternary ammonium salt complex may then be treated with another salt to remove the quaternary ammonium salt.
 Certain embodiments optionally include steps of using a vacuum to remove organic solvent from the polysaccharide, derivatized polysaccharide, or complexes of the polysaccharide. The vacuum removal is preferably performed at room temperature without adding heat. Alternatively, heat may be applied to evaporate the solvent, preferably enough heat to raise the temperature of the solvent to its boiling point without denaturing the heparin such that its biological activity is substantially reduced, a temperature that may vary according to the solvent used but typically being a temperature of less than approximately 100 degrees Centigrade and preferably less than 70 degree Centigrade. Alternatively, a mix of vacuum and heat may be used.
 A preferred embodiment uses an organic solvent that has a boiling point at atmospheric pressure and a dielectric constant that are less than conventionally used organic solvents. A more preferred embodiment has a boiling point of less than approximately 115 degrees Centigrade and a dielectric constant that is less than that of DMSO. A more preferred embodiment uses an organic solvent that has a boiling point of less than approximately 70° C. at atmospheric pressure and a dielectric constant that is less than that of DMSO.
 Certain embodiments include polymerizing the polysaccharide macromer in an organic solvent to make a polymer of at least two macromers. The macromers may be the same or different, to thereby make a homopolymer or a copolymer. Some embodiments include making a hydrogel from the polysaccharide macromer and/or polysaccharide macromer, preferably in an organic solvent.
FIG. 1 is a longitudinal sectional view of an embodiment of the invention having a fabric graft encapsulated within a hydrogel;
FIG. 2 is a side elevational view of the structure depicted in FIG. 1;
FIG. 3 depicts a longitudinal sectional view of an alternative embodiment of the invention having a fabric graft coated with a hydrogel;
FIG. 4 depicts a longitudinal sectional view of another alternative embodiment of the invention having a double-coated fabric graft; and
FIG. 5 depicts a longitudinal sectional view of another alternative embodiment of the invention like the embodiment of FIG. 4 except that the fabric graft tube has been inverted.
FIG. 6 depicts a partial view of an alternate embodiment of the invention having a plastic surface coated with a hydrogel.
FIG. 7 depicts a cross-sectional view of a hydrogel being formed on a mandrel.
FIG. 8 is a sectional view of FIG. 7.
FIG. 9 depicts a reaction scheme for making and using polysaccharide macromers.
FIG. 10 depicts an alternative reaction scheme for making and using polysaccharide macromers.
 Synthetic materials implanted into the soft tissue of a patient elicit a range of undesirable reactions that are typically categorized as acute inflammation, chronic inflammation, the formation of granulation tissue, foreign body reaction, and fibrosis (Ratner et al., 165-173, Biomaterials Science, 1996 Academic Press). Other reactions are possible, such as an immune system response or systemic toxicity and hypersensitivity. At a materials-blood interface, the blood coagulation mechanism can be activated to cause local clotting and downstream events such as complement activation or generation of blood clots (Ratner et al., 193-199). These reactions impact how synthetic articles, including vascular grafts, are designed and manufactured.
 Current medical practices for the use of synthetic articles as vascular grafts are essentially restricted to the use of polyester and polytetrafluoroethylene tubes as replacements for large-diameter blood vessels to treat patients with certain types of vascular disease.
 Vascular disease takes many forms, but one of the most common is stenosis, where an artery becomes narrowed by build up of plaque. Atheroscherotic Stenosis is a condition where the artery becomes hardened and less flexible by the process of calcification. The condition is often accompanied by a build up of tissue or plaque on the inside wall of the lumen of the artery. This build up causes the lumen of the artery to narrow and restrict the passage of blood.
 Total occlusion of an artery may occur when a clot of blood (a thrombus) lodges inside an artery, at a narrowing. Such occlusions can prove fatal if the artery is in a critical position; for example, coronary thrombosis causing a heart attack, or cerebral thrombosis causing a stroke.
 Even if there is not total occlusion, a restriction of the flow of blood can cause severe problems in limbs and organs downstream from the stenosis, due to starvation of the oxygen and nutrients supplied by the blood. A very common example is the reduction of flow in the lower leg, which may lead to claudication and eventually to gangrene and loss of a limb.
 A higher profile example of stenosis is observed when the disease affects the coronary arteries, which supply blood to the muscle of the heart. Stenosis, or occlusion due to thrombosis, can lead to an infarction due to parts of the muscle tissue of the heart dying (necrosis). That is, a myocardial infarction or heart attack.
 Another common form of vascular pathology is an aneurysm. An aneurysm is a condition where the wall of the artery weakens and dilates to form a balloon-like swelling. An aneurysm is a weakening of the artery wall leading to dilation. This dilation can develop so that the arterial wall is too thin and weak to withstand the pressure of the blood. A burst aneurysm causes severe hemorrhage, and can be fatal. One example is the so-called AAA (triple A) or abdominal aortic aneurysm; rupture of an artery having this defect is almost always fatal.
 Conventional vascular surgery techniques have made the surgical replacement of diseased arteries commonplace. The use of autologous grafts where a non-essential vein from a person's own blood vessels is used as a replacement artery is the oldest form of vascular grafting, and is still used today especially for coronary bypass procedures. A patient, however, may not have enough autologous donor vessels to fill the needs of the replacement surgery. Further, it is desirable that there be an alternative to the loss of a functioning blood vessel. It was the development in the 1960s of synthetic fabric prostheses, which led to the range of products available to the vascular surgeon today.
 Conventional synthetic vascular grafts are woven or fabric seamless fabric tubes, which are used as a direct replacement for a section of diseased artery. Various materials have been tried, but the most successful is polyester, which is the only material now used by clinicians for fabric grafts. Polyester is a very bio-stable material and, although slightly thrombogenic, is reasonably well tolerated in use for larger diameter arteries (approximately 6-mm diameter and above).
 When implanted, a vascular graft causes the body to react and generate a blood clot layer around its inner perimeter. This blood clot may be accompanied by tissue growth. If such tissue growth is not securely anchored in place, there is a danger of a loose blood/tissue clot or embolus being formed. The flow of blood may carry the embolus downstream until it reaches a narrower part of the artery where it may cause a blockage or occlusion.
 In order to ensure that tissue growth remains anchored to the graft, fabric grafts are made slightly porous so that the tissue grows into the pores of the fabric and is firmly attached. This has an added advantage that the tissue encapsulates the fabric graft and covers it, using the graft as scaffolding for new growth. In general, grafts with higher porosities, and especially flexible and elastic fabric grafts, heal better and generally perform better than low porosity woven grafts. It is because of this effect that the so-called “velour” grafts have been developed. Velour grafts have textile filaments raised up on the surface of the fabric.
 The need for porosity creates an initial problem for the surgeon since a high porosity graft will leak if simply implanted with no special preparation. After implantation, tissue growth fills in the pores in the fabric and therefore renders the graft “blood tight”. The problem is to make the graft blood tight during the initial surgery and for the immediate few hours afterwards. The original approach taken was to “pre clot” the graft before implantation. A small amount of blood was taken from the patient before the operation, and the graft was soaked in this blood so as to fill all the pores of the fabric with clots. The porous structure ensured that the clotted blood was firmly attached. This procedure was, however, time consuming, and it was and is very difficult to properly pre-clot those materials with high porosities. Because of these problems, the soft and very compliant fabric grafts, which heal better than stiff woven grafts, were not considered suitable for areas of high pressure such as thoracic arteries near the heart.
 Subsequent to the development of fabric grafts, a new development was introduced, which was the invention of a graft made from polytetrafluoroethylene (PTFE). PTFE is a material that is generally well tolerated in the body. A chief advantage with a PTFE graft is that it does not require pre clotting. However, PTFE material does not heal as well as warp fabric grafts. Modern PTFE prostheses are made of expanded material in order to imitate the cellular structure of fabric grafts, but this is only partially satisfactory.
 Fabric grafts are now available which do not require pre-clotting. They are coated with a bio-absorbable material such as gelatin or collagen. The coating material is gradually adsorbed and tissue grows to replace it. Such grafts are not only simpler to use because there is no requirement for pre-clotting, but they also tend to give better healing and performance.
 All of these conventionally used grafts have a common disadvantage in that they cannot be used for small diameter applications. Synthetic vascular grafts with a diameter of less than approximately 6 mm will not remain functional over a clinically significant period of time.
 The reason for the failure of conventional small diameter vascular grafts is thought to be related to the flow rate and type of blood flow within the graft. With a large inner diameter graft, there is a high volume of blood passing though the vessel. Any small tissue growth on the inside wall is insignificant and does not disrupt flow. With small inner diameter grafts, a small tissue build up is more significant in relation to the overall diameter of the vessel. A small build up will cause flow turbulence, which, because of the already low volume flow, tends to cause even more tissue growth, which leads to stenosis.
 Conventional medical clinicians urgently need a small diameter synthetic graft to cope with a range of small caliber replacement requirements. The only procedure conventionally available is an autologous transplant of the saphenous vein in the leg. The problem with this approach is that the amount of graft material available is very limited. In addition, the removal of the saphenous vein causes severe discomfort to the patient. One of the most important of these small diameter applications is the replacement or bypass of coronary arteries in cases of coronary stenosis. Another use is for the replacement of stenosed or occluded infragenticulate arteries.
 In order to replace the smaller caliber vessels using a small diameter vascular graft, the implant should be highly biocompatible and not thrombogenic. The biocompatibility should be such that little tissue growth is stimulated and the blood and body accept the prosthesis in a similar way that they accept the natural artery.
 Many polymers have been tested for the small diameter vascular graft application. A polymer is a single molecule built up by the repeated reaction of molecules that may be referred to as monomers. Monomers are molecules that may be reacted with other monomers to make a polymer. For example, a methylmethacrylate monomer represented by “A” may be used to make polymethylmethacrylate, which would be the polymer “AAAAA”. A polymer made of monomers A and B could have a random structure of ABAAABAABA and could be called a polymer or a copolymer of A and B. Polymer AAAAA and BBBBB could be joined to form copolymer AAAAABBBBB. A molecule that is derivatized is a molecules that has been chemically changed; a molecules that has been decorated is a molecules that has had a chemical unit attached.
 A macromer, as used herein, is a monomer or a polymer that is polymerizable and is a convenient term for referring to monomers or polymers that have been decorated with a monomer or used as decorations. Polysaccharide macromers include multiple polysaccharides that are polymerizable. For example, several macromers may be polymerized together to form a larger polymerizable group that is a macromer. Or several polymers, e.g., polysaccharides, may be joined together and decorated with polymerizable groups to form a macromer. Thus a heparin macromer is a heparin molecule that is polymerizable, for example after being decorated with a monomer. The term polymer, as used herein, includes oligomers and chains of at least two monomers in length. Polymers can be assembled into larger materials; for example, many polymers may be linked together to form a hydrogel.
 A variety of hydrogels are blood compatible. Examples that are commercially useful are made from polyhydroxyethylmethylacrylate (PHEMA), polyacrylamides, polyacrylic acid, N-vinyl-2-pyrrolidone (NVP), methacrylic acid, methyl methacrylate, and maleic anhydrides, each of which have been proven to be polymerizable from monomers. Further examples of biocompatible materials include hydrogels made from polyvinyl alcohol, methacrylates in general, acrylates in general, polyethylene glycol (PEG), hyaluronic acid, and alginates.
 The term hydrogel, as used herein, is a cross-linked material that can absorb or imbibe a water and is produced by the cross linking of one or more monomers or polymers. The cross-links in a hydrogel may be the result of covalent bonding or of association bonds, for example hydrogen bonds, charge-charge interactions, or strong van der Waals interactions between chains (Ratner et al., pages 60-64). A hydrogel cannot be suspended in water nor does it dissolve in water; instead, it remains intact water. For example, a material can shrink or swell and still remain intact without dissolving. For example, a contact lens shrinks and swells in water but does not dissolve therein.
 Hydrogels have been used for a myriad of applications, such as artificial tendon materials, wound-healing bioadhesives, wound dressings, artificial kidney membranes, articular cartilage, knee cartilage replacement, artificial skin, maxillofacial and sexual organ reconstruction, tissue engineering scaffolds, and vocal cord replacement materials. There are many types of hydrogels known to those skilled in these arts, such as aerogels, xerogels, equilibrium-swollen hydrogels, solvent-activated hydrogels, and swelling-controlled relapse hydrogels (e.g., Ratner et al., page 60-64). These types of hydrogels may be used with certain embodiments as described herein. Moreover, all of the uses for hydrogels described in this application may be used in embodiments described herein.
 Polysaccharides are polymers made from monomers that are sugars. Alginate is a polysaccharide. Glycosaminoglycans are a subcategory of polysaccharides that are made from repeats of disaccharide units. Glycosaminoglycans include hyaluronic acid, chondroitin sulfate, dermatan sulfate, heparan sulfate, heparin, chitin, chitosan, and keratan sulfate. Polysaccharides and glycosaminoglycans can also be found in a related proteoglycan form; proteoglycans are polysaccharides unified with a protein. Methods, compositions, and uses as described herein for polysaccharides are applicable to proteoglycans, glycosaminoglycans, and natural or synthetic derivatives or fragments thereof.
 Heparin is a polysaccharide polymer with an important property: it interferes with key molecules in the blood clotting mechanism such that the blood will not clot. Coating the inside of a tube with heparin tends to increase the amount of time that the tube remains open to blood flow but, to date, small diameter vascular grafts coated with heparin have failed to resist blockage by blood coats for a clinically useful length of time. Many researchers have tried to use heparin to coat the inside of vascular grafts. Heparin is a molecule that belongs to a group of molecules called polysaccharides that are molecules made from combinations of smaller molecules called sugars. Polysaccharides are naturally-occurring polymers but certain embodiments are directed to synthetic, derivitized, man-made, and semi-synthetic polysaccharides.
 Heparin has been applied to materials in many ways. General strategies have included adsorption, making charge-charge bonds with the surface, covalent immobilization, and release from a surface. Photoactivatable chemical groups have been put onto heparin so that if heparin is put close to the surface and the surface is bathed in light to make the photoactive groups make permanent covalent chemical bonds with the surface. Similarly, heparin has been reacted with polymerizable monomers that have then been reacted to achieve covalent bonding to the surface.
 Despite the great amount of research effort that has been expended in the field of biomaterials, including research with hydrogels, heparin, and polysaccharides, there is a continued need for improved biomaterials. Biomaterials that may be used in blood-contacting applications are especially required.
 Embodiments are provided that include an improved biomaterial that successfully combines the advantages of polysaccharides and hydrogels. One embodiment is a hydrogel made of polysaccharides. Another embodiment is a linear or multi-armed polysaccharide or polyvinylpyrrolidone polymer that is absorbable to a surface. Multi-armed means a soluble polymer that is branched or cross-linked. The hydrogels and the coatings may be used for the many applications for which a hydrogel may be used, e.g., as already described. There are other uses for hydrogels and the linear or multi-armed polymers that include coatings for stents, catheter coatings, cardiac valves or leaflets, cartilage replacement, replacement knee cartilage, organ scaffolds, lumbar disks, cell encapsulation, wound healing, nerve guides or tubes, and postoperative adhesions. A hydrogel may be used to make, coat, or encapsulate such devices. A coating may be used to such devices to improve their performance. Another hydrogel use is for a wound dressing for large, shallow wounds on animals so that a scab does not form over the wound but prevents blood clotting at the surface, thereby preventing scar formation.
 The hydrogel may be made by mixing polysaccharide macromers, e.g., heparin macromer, with other macromers or monomers to make a mixture. The mixture is poured into a mold and polymerized. After polymerization, the mold is removed and the polymerized macromers/monomers are hydrated. Various shapes, e.g., sheets, tubes, spheres, rods, may be formed by using suitable molds.
 The resultant shape is not covalently or otherwise bonded to other materials: the exterior and luminal surfaces are “free”. The free surfaces are not attached to other surfaces. A free surface may be decorated with moieties, e.g., drugs, polymers, and other agents. Such decorations do not cause the free surface to thereby be attached to other materials. Further, the resultant shape is formed during the polymerization process. This shape-forming process is distinct from processes that build up a coating on the inside of, e.g., a tube, in part, because the coating on the tube is essentially inseparable from the tube, especially if it is ionically or covalently bonded thereto.
 Moreover, the prior art methods of applying a coating to a tube and building up the coating is not equivalent to the present process. First, the prior art coating procedure does not create the tubular shape. But polymerizing macromers into a tubular shape during the polymerization process does create the tubular shape. Making a shape during a polymerization process is difficult because the polymerization reaction must be effective enough to make a solid material, with the effectiveness depending on polymerization variables known to those skilled in these arts. For example, an effective process requires using a macromer that can be provided in sufficient concentration. Not all macromers are sufficiently soluble to be present in solution with a high enough concentration to make a solid. The polymerization mixture must have be crosslinkable for crosslinks to form. The kinetics of the polymerizable groups must be suitable.
 Many prior art processes have not overcome these limitations and instead dry polymerizable groups onto a surface and then crosslinking them. The drying step results in drastically different structures than those made with polymerization from a solution (or a melt). If proteins or polysaccharides are dried, they aggregate and form clumps on the molecular level. In contrast, polymerization from a solution gives structures that are not aggregated but instead have a network of unaggregated polysaccharides. Since the macromers are not aggregated, the density of materials polymerized from solution can be lower. Moreover, a true polymerization process may take place wherein the polymerizable groups react with each other to form a polymeric backbone. In contrast, dried solutions have little mobility and the polymerizable groups react with the chemical structures closest to them instead of reacting with other polymerizable groups.
 In short, chemical crosslinking is not equivalent to polymerization. The present materials and methods provide for polymerization methods and for polymerized materials as opposed to coatings built up on surfaces, gelled structures, and merely aggregated, chemically crosslinked, or surface-immobilized materials. The advantages of polymerization are numerous and well known to ordinary artisans.
 One advantage of a polysaccharide covalently polymerized with a hydrogel is that the polysaccharide may be stably incorporated into the hydrogel so that it is not released over time. This stability is useful for a long-term implant because the hydrogel would otherwise dissipate over time and eventually fail. The heparin hydrogels are hypothesized to function by reversibly binding antithrombin III. The antithrombin III is bound by the heparin and thereby changes its shape so that it reacts with and inactivates both thrombin and Factor X(a), which are key enzymes required for blood to clot. The antithrombin III is hypothesized to stay on the heparin hydrogel temporarily so that it attaches, reacts with thrombin and Factor X(a), and departs back into the bloodstream so that a new antithrombin III molecule may be bound to the heparin.
 Another advantage of the polysaccharide hydrogel is that it can be made as a thick film. Thick films may be handled by surgeons, grasped with tools such as forceps, punctured by sutures, and suffer scratches and damages to their surface without losing their favorable blood-contacting properties. The thickness of the film and the three-dimensional structure of the film allows it to suffer minor damage while continuing to cover the surface with heparin molecules. In contrast, damage to a thin coating or a synthetic material that has been merely reacted with a polysaccharide can entirely remove the polysaccharide and expose the underlying material to the body. For example, a surgeon that uses forceps to firmly grasp a plastic tube covered with a layer of heparin that has been reacted with the tube's surface might accidentally scratch the tube and remove the heparin thereby exposing the underlying plastic material of the tube. In contrast, a surgeon might accidentally scratch a plastic tube encapsulated with a thick film heparin coating but would not thereby expose the underlying plastic because a scratch in the thick film would expose only more of the heparin hydrogel.
 Further, the thickness of the hydrogel film is hypothesized to minimize blood contact with the synthetic material that the thick film is encapsulating. Blood or its components must penetrate through the entire thickness of the hydrogel prior to reacting with the encapsulated synthetic material. In contrast, a thin coating, especially a coating of a few molecules' thickness, presents a shorter distance between the blood and the encapsulated synthetic material. This distance is important because the efficiency of surface chemistry reactions used for conventional techniques is hypothesized to typically provide a surface coverage of less than 100%, i.e., not every space on a surface coated with heparin is completely covered with a heparin molecule. In contrast, a thick film provides more than 100% coverage because any molecule that would react with the surface must pass through a thick coating that has a thickness of many molecules.
 One embodiment is a tubular member encapsulated by a polysaccharide hydrogel. The encapsulation may be achieved by a number of processes, one such process being placing tubular member into a mold and charging the mold with the desired formulation of polysaccharide macromers and polymerized by conventional techniques at room or at elevated temperature and/or by electromagnetic radiation. A material that encapsulates a member is termed an encapsulant.
 Certain embodiments involve polymerization processes, including polymerization of polysaccharide macromers. The polysaccharide macromers may be polymerized using conventional techniques, for example using initiators, carriers, accelerators, retardants, viscosifiers, and/or cross-linkers. The polysaccharide macromers may also be copolymerized with and other monomers and/or macromers and/or polymerizable polymers. In another embodiment, a tubular member is covered with a mixture of polysaccharide macromers and/or monomers and/or polymers. The polymerizable groups are subsequently polymerized to form a crosslinked hydrogel. It may be desired to placed the tubular member on a rotating mandrel and apply the mixture so as to ensure that the tubular member is encapsulated uniformly.
 The polysaccharide macromers and polymers made thereof may also be copolymerized, blended, mixed, and/or cross-linked with other monomers and/or macromers, and/or polymers including engineering polymers, blood compatible polymers, hydrogel polymers, natural polymers (e.g., deoxyribonucleic acid, polysaccharides, and proteins and bioactive fragments thereof) and/or fillers. The type of initiation is not limited and may include thermal, X-ray, ultraviolet, infrared, visible light, free radical, addition, sonic, and condensation initiation.
 Monomers, i.e., for mixing with the polysaccharide macromers, can include, but are not limited to, monomers with hydroxyl groups (e.g., hydroxyethyl methacrylate), monomers with glycerol groups (e.g., glycerol monomethacrylate, glycerol dimethacrylate, glycerol trimethacrylate), monomers with polyoxyalkylene ether groups (e.g., polyethylene glycol methacrylate, polypropylene glycol methacrylate), monomers with vinyl groups (e.g., N-vinyl pyrrolidone), monomers with zwitterionic groups (e.g., 2-methacryloyloxyethyl-2-(trimethyl ammonium) phosphate, monomers with silicone groups (e.g., methacryloxypropyl tris (trismethyl-siloxy) silane and other silicone methacrylate or acrylates), monomers having sulphate groups (e.g., vinyl sulphonic acid), monomers having sulphonate groups (e.g., ammonium sulphatoethyl methacrylate), heparin monomer as cited in the patent PCT GB9701173 and U.S. Pat. No. 6,096,798, which are hereby incorporated herein by reference.
 Polymers for mixing, blending, and/or copolymerization with polysaccharide macromers include derivatized polymers, for example, derivatized polyoxyalkylene ether groups (e.g., polyethylene oxide terminating in hydroxyl group, carboxylic groups and/or isocyanate groups and polypropylene oxide terminating in hydroxyl group, amino groups, carboxylic groups and/or isocyanate groups), polyvinyl pyrrolidone functionalized with methacrylate groups, methacrylate terminating dimethysiloxone, vinyl terminating dimethylsiloxone, polyurethane terminating in isocyanate, polyester terminating in isocyanate and also other polymers that can be derivatized with methacrylate, acrylate, isocyanate, carboxylic acid, amino, hydroxyl and/or vinyl groups.
 Polymers mixed, blended and/or copolymerized with polysaccharide macromers may also be used to enhance the viscosity of the mixture formulation for the application to the rotary mandrel and hence the polymerization of the mixture formulation. This technique may be used to give the encapsulated tubular member a homogenous and smooth surface, a feature that enhances vascular prosthesis biocompatibility. The encapsulated tubular member and hence the biocompatible vascular prosthesis has a smooth surface containing non-thrombogenic or anti-thrombogenic properties or both and preferably has a water content ranging from about 30% to about 90%.
 In one embodiment, a tubular member is encapsulated by the hydrogel, e.g., as in FIGS. 1 and 2, the hydrogel having cross-linked polymers made from polysaccharide macromers and copolymerized with monomers from at least three of these classes, as discussed in patent application PCT GB97 01173, U.S. Pat. No. 6,096,798: (a) monomers having sulphate groups, (b) monomers having sulphonate groups, (c) monomers having sulphamate groups, (d) monomers having polyoxyalkylene ether groups and (e) monomers having zwitterionic groups. The polysaccharide macromers are preferably heparin macromers. Hydrogel encapsulation of the tubular member is performed by placing the tubular member into a mold, adding a macromer and/or monomer formulation and then polymerizing to make a hydrogel. The monomer constituents may vary from 10% to 90% by weight and are preferably polymerized with a bifunctional monomer, e.g., ethylene glycol dimethacrylate. This formulation provides the prosthesis with a smooth surface, prevents the leakage of blood, is non-thrombogenic and/or anti-thrombogenic, and has a water content ranging from 30% to 90% when hydrated. When hydrated, the biocompatible vascular prosthesis is soft and pliable so it will not compromise the mechanical properties of the prosthesis.
FIG. 1 depicts a longitudinal cross section of a small diameter vascular graft 10. FIG. 2 shows an end view of a radial cross section of FIG. 1. The loops of synthetic material 12 are completely encapsulated within hydrogel 14. This embodiment is suitable for porous members that allow passage of the macromer and/or monomer constituents through the pores of the tubular member, e.g., a fabric, in order to provide binding between the inner and outer faces of the encapsulant.
 An alternative embodiment uses approximately the same materials and methods but a tubular blood vessel member 13 is used which has sufficiently low porosity so that blood leakage is not a consideration. e.g., a tightly knitted or woven fabric, or a plastic extrusion. With this kind of tubular member, the basic porosity is low. Referring to FIG. 3, instead of encapsulating tubular member 13, it is covered with a hydrogel layer 14 of polysaccharide macromer and/or monomers. In the case of a fabric tubular member, the hydrogel coats the individual yams and fibers. At the same time, the hydrogel may be used to cover the pores completely or partially. The loops of knitted fabric 13 are coated by hydrogel 14. This embodiment is simpler to make than the encapsulation process as the covering is applied by dip process, spraying technique or by other conventional process.
 Another embodiment uses approximately these same materials and methods but uses a porous tubular member 15, such as one made from a fabric, which is pre-coated (a primary coating layer 16) with a polymeric material in order to prevent blood leakage, as the layer provides a strong “leak proof” security layer (FIG. 4). The leak proof layer is essentially impermeable to blood, meaning that it generally prevents flow of blood in surgical applications. Alternatively, the barrier may be permeable to blood, and control the flow of blood. Primary coating layer 16 may be made from a number of flexible polymers, e.g., silicone polymer, polypropylene, polyester, polyurethane, polytetrafluoroethylene (PTFE) or elastomeric polymer such as silicone rubber. The total composite is then further coated with hydrogel 14 containing polysaccharides. This structure imparts an extremely thin, flexible, and compliant wall to vascular prosthesis 10. Each of the coating processes may be applied by dip process, spraying technique or by other conventional coating process. FIG. 4 depicts a longitudinal cross section of a knitted tubular member coated according to this embodiment. The fabric loop of porous tubular member 15 is inside primary coating 16, which in turn is inside hydrogel 14.
 Referring to FIG. 5, another embodiment takes advantage of the manufacturing processes of woven vascular prostheses. A tubular member made from knitted fabric has a different surface texture on the internal face and the external face. The technical face of the fabric tends to be smoother than the technical back. In the case of the knitted tubular member, the technical back is on the internal face of the fabric, whereas the technical face is on the external face. When used as a vascular prosthesis, the knitted tubular member can be inverted so that the smoother face is on the internal face of the prosthesis. In addition, the internal face has very shallow grooves, which run longitudinally. These grooves may help to smooth the flow of the blood and reduce turbulence.
FIG. 5 shows a schematic longitudinal cross section of a porous tubular member 15 made of a knitted material that is coated with the fabric inverted. The fabric loop of the porous tubular member 15 is inside the primary coating 16, which in turn is inside the hydrogel 14. A tubular vascular vessel made from knitted fabric has a different surface texture on the internal face and the external face. The technical face of the fabric tends to be smoother than the technical back. In the case of the knitted tubular member, the technical back of the fabric is on the internal face of the vessel, whereas the technical face is on the external face of the vessel. The fabric loop of the porous tubular member 15 is inside the primary coating 16, which in turn is inside the hydrogel 14.
 A most preferred tubular member is a warp knitted or weft knitted seamless fabric tube. In the warp knitted form, the most preferred structure is reverse locknit, but tricot can also be used. The fabric tube can also be weft knitted or woven. The tube may be continuous or it may be bifurcated in order to fulfill the needs of a graft designed to replace the Aorto-Iliac bifurcation.
FIG. 6 shows a portion of a vascular graft 10 with surface 18 with a thick film of hydrogel 24. The thickness of film 24 protects surface 18 from being exposed as a result of damage caused by handling the vascular graft 10. The film is preferably at least 25 μm thick, more preferably from about 5 to about 1500 μm thick, and even more preferably about 500 to about 800 μm thick.
 Alternatively, a very thin coating or hydrogel may be applied to a graft or other structure. The coating is applied, e.g., by spraying, so that the coating or hydrogel is present on the fibers of the graft. The interstices between the fibers are not coated by this technique. For example, a wire mesh stent may be coated with a hydrogel without the hydrogel filling the interstices of the mesh. Such a hydrogel or coating is preferably less than about 10 μm in thickness.
FIG. 7 depicts vascular graft 10 being made by molding hydrogel 14 around mandrel 20 and within outer mold 22 while entrapping synthetic material 12 in annulus 26 created by mandrel 20 and outer mold 22. The monomers and/or macromers used to make hydrogel 14 are poured into annulus 26 and then polymerized around synthetic material 12. FIG. 8 depicts cross-section 8-8 of FIG. 7; hydrogel 14 is shown in phantom lines.
 A tubular member may also be a plastic tubular extrusion of a polymeric material. Such materials could be, for example, silicone polymers, polyester, polyurethane, polypropylene, silicone elastomer, polytetrafluoroethylene (PTFE) or other suitable materials. The tubular member may also be porous or have a defined permeability, e.g., as a hollow tube fiber with a defined molecular weight cut-off. Porosity/permeability may be incorporated into a tube by the selection of the material or by adding pores or holes, e.g., by lasers, and punctures.
 Another embodiment is a method of making a vascular graft. A first polysaccharide hydrogel tube is made and introduced inside of a tubular member that has an exterior side and an inner side that faces the lumen of the tubular member. A second polysaccharide hydrogel tube is made and introduced around the outside of the tubular member to form a sandwich of a tubular member between two hydrogel tubes. The two hydrogel tubes are then treated to form one unit. One suitable treatment is to swell the hydrogel tubes in a solvent to bring them in to contact with each other and then to chemically react them. Examples of suitable chemical reactions include polymerization, polymeric cross-linking with ultraviolet, heat, or sonic initiators, and chemical cross-linking with gluteraldehyde, or diisocyantes. The cross-linking agents may be present in the tubes prior to swelling or may be introduced with the solvent. Suitable solvents include aqueous solvents, organic solvents, and low boiling point and/or low dielectric constant organic solvents. One option for forming one unit of the hydrogels is to use a tubular member that is shorter than the two hydrogel tubes so that the two hydrogel tubes are joined around the tubular member. Another option for forming one unit of the hydrogels is to use a tubular member that is porous so that the hydrogels are forced into the pores during swelling so that the two hydrogel tubes contact each other through the pores. As a result, the hydrogels may become cross-linked through the pores.
 Polymers and hydrogels as described herein, whether encapsulating or coating the tubular member, may if desired, incorporate and slowly release growth factors, thrombolytic drugs, thrombotic drugs, enzymes, restenosis-preventing drugs, inhibitors and other agents used to treat diseased tissues. Further, gene therapy delivery may be performed by complexing gene therapy victors to the polymer or hydrogel, e.g., by complexing DNA and a polysaccharide together with a positively charged ion or polymer. Other functions, uses, applications, formulations, and technologies of hydrogels known those skilled in the art may be used to create further alternative embodiments.
 Another embodiment is related to making a heparin macromer by reacting heparin with a quaternary ammonium salt to form a heparin-quaternary ammonium salt complex; dissolving the heparin-quaternary ammonium salt complex in an organic solvent with a dielectric constant less than the dielectric constant of dimethylsulfoxide; and decorating the heparin in the heparin-quaternary ammonium salt complex with a polymerizable monomer. Further, a vacuum removal step of removing the organic solvent with a vacuum may be used, preferably at room temperature.
 An embodiment is a method of making a heparin macromer by reacting heparin with a quaternary ammonium salt to form a heparin-quaternary ammonium salt complex; dissolving the heparin-quaternary ammonium salt complex in an organic solvent that has a boiling point of less than 190 degree Centigrade at atmospheric pressure; and decorating the heparin in the heparin-quaternary ammonium salt complex with a polymerizable monomer. The step of dissolving the heparin-quaternary ammonium salt complex in an organic solvent is alternatively performed with an organic solvent with a boiling point of less than 114 degrees Centigrade at atmospheric pressure.
 Certain embodiments include an optional additional step of decomplexing the heparin quaternary ammonium salt from the heparin-quaternary ammonium salt complex by mixing the heparin-quaternary ammonium salt complex with a salt that is not a quaternary ammonium salt.
 Another embodiment is a method of making a heparin polymer by reacting heparin with a quaternary ammonium salt to form a heparin-quaternary ammonium salt complex; dissolving the heparin-quaternary ammonium salt complex in an organic solvent that has a boiling point of less than 190 degree Centigrade at atmospheric pressure; decorating the heparin in the heparin-quaternary ammonium salt complex with a polymerizable monomer; and polymerizing the monomer to make a polymer.
 Another embodiment is a method of making a heparin hydrogel by reacting heparin with a quaternary ammonium salt to form a heparin-quaternary ammonium salt complex; dissolving the heparin-quaternary ammonium salt complex in an organic solvent with a dielectric constant less than the dielectric constant of dimethylsulfoxide; decorating the heparin in the heparin-quaternary ammonium salt complex with a polymerizable monomer to make a heparin macromer; and polymerizing the heparin macromer to form a polymer and cross-linking the polymers to form a hydrogel.
 Another embodiment is a method of making a polysaccharide macromer, the method comprising: reacting a polysaccharide with a quaternary ammonium salt to form a polysaccharide-quaternary ammonium salt complex; dissolving the polysaccharide-quaternary ammonium salt complex in an organic solvent that has a boiling point of less than 190 degree Centigrade at atmospheric pressure; and decorating the polysaccharide in the polysaccharide-quaternary ammonium salt complex with a polymerizable monomer. The step of dissolving the polysaccharide-quaternary ammonium salt complex in an organic solvent may also be performed with an organic solvent with a boiling point of less than 114 degrees Centigrade at atmospheric pressure.
 Another embodiment is a method of making a material from a polysaccharide by reacting a polysaccharide with a quaternary ammonium salt to form a polysaccharide-quaternary ammonium salt complex; dissolving the polysaccharide-quaternary ammonium salt complex in an organic solvent that has a boiling point of less than 190 degree Centigrade at atmospheric pressure; and decorating the polysaccharide in the polysaccharide-quaternary ammonium salt complex with a polymerizable monomer to make a polysaccharide macromer; decorating the polysaccharide in the polysaccharide-quaternary ammonium salt complex with a polymerizable monomer to make a polysaccharide macromer. Moreover there may be included a step of polymerizing the polysaccharide macromer to make a polysaccharide polymer. There may further be included a step of polymerizing the polysaccharide macromer to form a hydrogel.
 Another embodiment is a material for use in a medical context, the material including a hydrogel made of a material including polymers, the polymers including heparin polymers made of polymerizable heparin macromers, the hydrogel having covalently cross-linked polymers such that the hydrogel remains intact in water. The heparin macromers may be macromers that are polymerizable while in a solution or in a suspension. The hydrogel may include polymerizable heparin macromers polymerizable in aqueous solvent and/or polymerizable heparin macromers are polymerizable in organic solvent. The amount of heparin in the heparin hydrogel may be least 1% as measured by dividing the dry weight of heparin macromers by the total dry weight of the hydrogel. And the amount of water in the heparin hydrogel may be at least 5% as measured by dividing the weight of water in the hydrogel by the total weight of the hydrated hydrogel and is preferably in the range of 10%-90% and more preferably 60%-80% water.
 Another embodiment is a material for use in a medical context, the material including a hydrogel made of a material including polymers, the polymers including polysaccharide polymers made of polymerizable polysaccharide macromers, the hydrogel having covalently cross-linked polymers such that the hydrogel remains intact in water, and the polysaccharide macromers being macromers that are polymerizable while in a solution or in a suspension.
 Another embodiment is a material for use in a medical context, the material including a hydrogel including polymers, the polymers including heparin polymers made of polymerizable heparin macromers, the hydrogel having covalently cross-linked polymers such that the hydrogel remains intact in water, the heparin macromers being heparin molecules decorated with a monomer chosen from the group consisting of monomers polymerizable by free-radical polymerization, monomers polymerizable by addition polymerization, and monomers polymerizable by condensation polymerization.
 Another embodiment is a material for use in a medical context, the material including a hydrogel including polymers, the polymers including polymers made of polysaccharide macromers, the hydrogel being covalently cross-linked such that the hydrogel remains intact in water, the polysaccharide macromers being polysaccharides decorated with a monomer chosen from the group consisting of monomers polymerizable by free-radical polymerization, monomers polymerizable by addition polymerization, and monomers polymerizable by condensation polymerization.
 Another embodiment is a vessel for use in a medical context, the vessel comprising: a tubular member with an inside wall defined by an inner diameter and an outside wall defined by an outside diameter joined by a thickness, a portion of the cylinder having its inner wall and outer wall covered with a hydrogel, the hydrogel including polymers, the polymers including heparin polymers made of polymerizable heparin macromers, with the hydrogel having covalently cross-linked polymers such that the hydrogel remains intact in water.
 The hydrogel for the small diameter vascular graft and other such vessels is preferably at least several μm thick, and more preferably is at least about 50 μm thick, and even more preferably about 500-800 μm thick. The vessel's polysaccharide macromers may be polysaccharide molecules decorated with a monomer chosen from the group consisting of monomers polymerizable by free-radical polymerization, monomers polymerizable by addition polymerization, and monomers polymerizable by condensation polymerization. In contrast to certain conventional processes, the polysaccharide macromers may be macromers that are polymerizable while in a solution or in a suspension.
 Another embodiment is a vessel wherein the diameter of the minimum cross-sectional area available for blood flow through the vessel after the vessel is covered with hydrogel is less than approximately 6.0 mm. The vessel may be a fabric vessel that has pores and the hydrogel is continuous through a portion of the pores of the fabric vessel.
 Another embodiment is a coated stent. Another embodiment is a tissue engineering matrix. Another embodiment is a medical device covered to make a biomaterial covering around the device.
 The medical device may be coated on one surface, or a portion thereof. Alternatively, the device may be completed coated on all exterior surfaces. An encapsulated device is coated on all surfaces with an essentially continuous material. If the material is a hydrogel, the hydrogel is preferably crosslinked so as to have an increased strength. A continuous hydrogel is distinguished from a thin coating because the coating is attached or adsorbed to the coated surface and is stable so long as that attachment is maintained. But a hydrogel forms a coherent structure that has stability independent of its attachment to the encapsulated surface. The term encapsulated, as used herein, means to cover. In the case of an encapsulated medical device, the covering is essentially total. In the case of a hydrogel encapsulating a tube, the inside and outside of the tube are covered. The ends of the tube are not necessarily covered.
 Embodiments include coated devices, for example, medical devices or components thereof, and includes medical devices that contact a bodily fluid. For example, springs, wires, guide wires, pacemaker leads, stents, implants, antennae, sensors, glucose sensors, tubing, blood bypass tubing, syringes, catheters, i.v. bags, needles, oxygen tubing, ventricular assist device components, and trochars.
 Improved Processes for Preparing Polysaccharides and Uses Thereof
 Synthetic materials have gained wide-ranging acceptance as suitable materials for medical devices. The extent of their application has extended from simple disposable devices, like syringes, blood bags, catheters, also products like extracorporeal devices, artificial blood vessels, stents, stent grafts to complex artificial organs, such as kidneys, lung, liver, heart assist devices and implantable devices. These medical devices are required to have the appropriate functional properties, durability, and biological safety.
 There is now emerging an additional requirement for medical devices, especially implantable devices, to have improved biocompatibility with the biological environment, with reduced or no tissue rejection or reaction. Anti-thrombogenicity is a biocompatibility property that is important in many cases. Further, it is often desirable to make medical devices with a lubricious exterior. Coatings may be used to impart lubricity and biocompatibility. Polysaccharide coatings, in particular, can impart these properties. Heparin is a natural polysaccharide with desirable properties.
 Conventional techniques for applying a heparin coating to medical devices consist primarily of two routes: (1) complexing heparin with a polymer for attaching a mucopolysaccharide (e.g., heparin) to a surface, and, alternatively, (2) chemical modification to introduce groups onto the mucopolysaccharides or the surface to make them hydrophilic, zwitterionic and/or charged, (e.g., anionic or cationic). Route (1) has been achieved by: (I) blending or attaching an organic-soluble polymer to heparin so that the heparin goes into solution in organic solvent; alternatively by (II) treating heparins to dissolve them into an organic solvent and using the organic solution to coat a medical device, alternatively by (III) electrostatic binding of heparin to a surface, and alternatively by (IV) chemically linking the heparin to the surface.
 To cite a few examples of these methods, Pusineri et al disclose in U.S. Pat. No. 4,469,827 polymer compositions containing quaternized amino groups that ionically bind to heparin. Hsu discloses ionic heparin coating in U.S. Pat. No. 5,047,020, where alkylbenzl ammonium cations are used to complex with heparin. U.S. Pat. No. 5,541,167, describes a coating composition consisting of stearyldimetylbenzyl ammonium heparin complex with antifoaming agents. Hsu et al in U.S. Pat. No. 5,417,969 inform processes for coating the surface of polyvinylchloride with an organic solvent soluble solution of heparin complexed with an organic cation. EP 0 769 503 A2, patent application discloses a heparin complex coating that contains stable ionic bonding and where reduction in anticoagulant activity is minimized. The preferred quaternary ion is alkyldimethylammonium. Yokota and et al have described (in EP 0 781 566), an organic soluble heparin complex coating for medical devices, where the cation consists of a quaternary phosphonium moiety. In another disclosure, U.S. Pat. No. 5,270,046, a monomer is formed containing quaternary ammonium groups, which are complexed with heparin. The monomeric complexed heparin is polymerized with other monomers. In general, however, these coatings fail under prolonged use in physiological conditions. The reason for failure is that ions generally decomplex from the heparin, causing the release of the heparin and the cation.
 Polysaccharide Processing by Formation of Polysaccharide Complexes with Cations
 In contrast to the other approaches, materials and methods are set forth herein that allow for improved processing of polysaccharides. In particular, the complexation of polysaccharides with cations to make them soluble in organic solvents enables for improved chemical processes to be used for polysaccharide derivitization and reaction. Polysaccharides carry negative charges that can be neutralized, at least partially, by selected cations so that the solubility of the polysaccharide complex is different than the solubility of the polysaccharide by itself. An association of cations with a polysaccharide may be referred to as a complexed polysaccharide or a polysaccharide complex. Natural, synthetic, and derivitized polysaccharides may all be complexed with cations and dissolved in an organic solvent. The polysaccharide part of the organic soluble complex can subsequently undergo a variety of chemical reactions. After completion of the chemical reaction, the polysaccharide may be de-complexed. The term soluble traditionally refers a substance that can be dissolved into molecular units. In the case of polymers and biological macromolecules, however, it is appreciated that dissolution may be more complex. Scientists refer to a polymer as being soluble in a solvent when the polymer is substantially removed from the solid phase and dispersed when exposed to that particular solvent.
 A polysaccharide complex may be formed by reacting a polysaccharide with a quaternary salt, e.g., a quaternary ammonium salt, to form a polysaccharide-quaternary ammonium salt complex. A quaternary ammonium salt has four substituents on the nitrogen so that the resulting moiety has a positive charge. Examples of quaternary ammonium salts include, e.g., cetyltrimethylammonium chloride, dodecyldimethylbenzylammonium chloride, benzalkonium chloride, didecyldimethylammonium chloride, benzethonium chloride, hexyl trimethyl ammonium, decyl trimethyl ammonium, lauryl trimethyl ammonium, myristyl trimethyl ammonium, cetyl trimethyl ammonium, stearyl trimethyl ammonium, didecyl dimethyl ammonium, dilauryl dimethyl ammonium, and distearyl dimethyl ammonium.
 Decomplexing a quaternary salt from a polysaccharide complex may be accomplished by mixing the polysaccharide complex with a salt that is not a quaternary ammonium salt. For example, a sodium chloride solution may be used for decomplexation. Decomplexation may be accomplished at room temperature, or other any suitable temperatures. As shown in, e.g., Example 14, exposure to a saturated NaCl solution for about a minute may be used to perform the decomplexation.
 Examples of organic solvents include dimethylformamide, dimethylacetamide, dimethyl sulfoxide, hexamethylphosphoric triamide, formic acid, acetonitrile, methanol, ethanol, acetone, acetic acid, dichloromethane, pyridine, and formamide. Some solvents for forming a polysaccharide complex or for reacting a polysaccharide or polysaccharide complex are solvents with a boiling point of less than approximately 115 degrees Centigrade and a dielectric constant that is less than that of DMSO. Other embodiments use an organic solvent that has a boiling point of less than approximately 70° C. at atmospheric pressure and a dielectric constant that is less than that of DMSO. These solvent properties have certain advantages; for example, they allow for expedited heat-and-vacuum removal of the solvents with minimal denaturation of the polysaccharides. Certain mebodiemtns inlcludea mixture of at least 50% by volume of a particular organic solvent or a particular combination of organic solvents.
 Polysaccharide complexes may be formed with a native polysaccharide or with a polysaccharide that previously has undergone chemical reaction(s) to decorate it with a functional group or otherwise modify the native polysaccharide; examples of such reactions are set forth in FIGS. 9 and 10. FIGS. 9 and 10 show schemes that involve formation of polysaccharide complexes at various points in the chemical processing of a polysaccharide from a first chemical state to a second chemical state. Referring to FIG. 9, which depicts scheme I, certain embodiments of chemical reactions involving complexed polysaccharides are set forth. FIG. 9 depicts the complexing of a polysaccharide with cations to make a polysaccharide complex that comprises the polysaccharide and multiple cations. The polysaccharide complex is soluble in an organic solvent as a result of the formation of the complex. A molecule that is an organic solvent has at least one carbon in its structure. The organic-solvent soluble complex can then undergo chemical reaction in an organic solvent to chemically modify the polysaccharide. A chemically modified polysaccharide that has undergone modification in an organic solvent is abbreviated as O-MPSAC, i.e., organic-solvent reaction modified polysaccharide. Many chemical processes are feasible only in organic solvent so that the processes could not take place if the polysaccharide were not complexed. Provided herein are numerous examples of a reaction of a complexed polysaccharide in organic solvent, e.g., as in the Examples. An O-MPSAC may be decomplexed from its cations and is typically water soluble.
 Referring to FIG. 10, which depicts scheme II, additional embodiments of chemical reactions involving complexed polysaccharides are set forth. A polysaccharide undergoes chemical reaction and is soluble in water or mixtures of water with other solvents, e.g., DMSO. A polysaccharide derivitized in water is herein referred to as a W-MPSAC, i.e., a polysaccharide modifiable in a solvent of water or water mixed with other solvents. The W-MPSAC is then complexed with a cationic moiety to form a complex that is organic soluble, and which can undergo further chemical reaction. After the further reactions, the W-MPSAC may then be de-complexed.
 Reaction of Polysaccharides with Surfaces
 Polysaccharides may be attached to a surface using the materials and methods set forth herein. A polysaccharide may be, for example, natural, synthetic, derivitized, or processed according the methods described herein. The surface, the polysaccharide, or both the surface and the polysaccharide may be modified to accomplish the polysaccharide attachment.
 The surface that is reacted with a polysaccharide may be any surface, and may be a surface of a medical device. The surface may be of any material, for example, polymeric, metallic, natural, synthetic, or biological. Surfaces of a medical device that contact a bodily fluid are suitable for modification with polysaccharides using materials and methods described herein. Such surfaces are essentially any surface that is exposed to a tissue after implantation in a body or patient. Such surfaces also include extracorporeal devices, e.g., catheters, trochars, pumps, blood-contacting devices, stents.
 Modifying a polysaccharide and/or a surface may involve processing the polysaccharide and/or surface to associate a functional group with the polysaccharide and/or surface. The functional group is subsequently involved in a chemical reaction to thereby attach to polysaccharide and surface to each other. The functional groups could be all of a single type, or could be a variety of types. The functional group may be introduced by adding new chemical moieties to the modified material, or the material may be modified to remove some chemical moieties so as to create the functional groups.
 The modification of a surface or a polysaccharide is often advantageous. For example, new functional groups can be introduced that are capable of forming covalent bonds with other functional groups. Or the new functional groups may allow for more efficient chemistries to be used. Or the new functional groups may allow for a greater degree of bonding. For example, a polymeric surface may be modified to express primary amines. The primary amines may then be reacted with carboxyl groups on a polysaccharide to make covalent bonds. In some cases the reaction between the amines and carboxyls can be efficiently performed only in an organic solvent.
 Certain embodiments involve the chemical reaction between a polysaccharide complex and a surface in the presence of organic solvent. Referring to FIGS. 9 and 10, for example, the polysaccharide complex may have a functional group on the polysaccharide that is capable of reacting with a functional group on the surface in the presence of an organic solvent to thereby form a bond. After reaction of the polysaccharide complex with the surface, the polysaccharide is de-complexed, and the polysaccharide remains bonded to the surface, e.g., by a covalent bond. Thus, an O-MPSAC may be chemically reacted to a surface; with the chemical reaction in step 2 of FIG. 9 denoting the binding of the polysaccharide to the surface. Or a W-MPSAC may be reacted with a surface, with the chemical reaction of step 3 of FIG. 10 denoting the binding of the polysaccharide to the surface. Thus, the O-MPSAC or W-MPSAC may be complexed as needed to complete the reaction and decomplexed to leave the polysaccharide in association with the surface without the presence of the complexing cations.
 Certain embodiments involve the chemical reaction between a polysaccharide complex and a surface in the presence of water, DMSO, or a comparable solvent. Referring to FIGS. 9 and 10, for example, the polysaccharide may have a functional group that is capable of reacting with a functional group on the surface in the presence of a such a solvent to thereby form a bond. After reaction of the polysaccharide complex with the surface, the polysaccharide is bonded to the surface, e.g., by a covalent bond. Thus an O-MPSAC may be chemically reacted to a surface after it has been de-complexed, as shown in step 3 of FIG. 9. And a W-MPSAC may be reacted with a surface, with the binding of the polysaccharide to the surface being denoted as the first chemical reaction in step 1 of FIG. 10. Alternatively, the W-MPSAC may be reacted with the surface after it has been de-complexed. Thus the O-MPSAC or W-MPSAC may be complexed as needed to complete a reaction and decomplexed to leave a polysaccharide in association with the surface without the presence of the complexing cations.
 A chemical reaction between a polysaccharide complex, e.g., O-MPSAC or W-MPSAC, and a surface may take place as a result of the activation of the surface, the polysaccharide, or a combination of both. For example, a surface that has been derivitized with electrophilic functional groups is thereby activated to react with nucleophiles on a polysaccharide. Or a polysaccharide may be activated with electrophilic functional groups to thereby react with nucleophiles on the surface. Or a monomer may be attached to both a surface and a polysaccharide so that they polymerize together when polymerization of the monomers is initiated. Such reactions may be performed in an organic solvent or other solvents, with the polysaccharide being complexed or decomplexed as needed to make the polysaccharide in the solvent.
 Chemical Reactions Involving Polysaccharides and/or Polysaccharide Complexes
 A tremendous variety of chemical reactions are available for use with the various embodiments of polysaccharide and polysaccharide complexes set forth herein. Polysaccharides often have hydroxyls and carboxyls in their native states; these chemical groups are particularly suitable for further modification. An advantage of certain chemistries, e.g., some photoactivation chemistries, is that reactions may be undertaken with other chemical groups besides hydroxyls and carbosyls. Some chemical reactions may involve reacting a polysaccharide or polysaccharide complex with a surface, other polysaccharides, polysaccharide complexes, polymers, or other molecules. Such reactions may include, for example, a free radical process, a photo initiated reaction, an electrophilic-nucleophilic reaction, and Sn2 substitution, an elimination reaction, chemical coupling reaction to various polymers, and other reactions that are known to those of skill in these arts. These reactions may be performed with the polysaccharides in complexed or uncomplexed form, as needed to adapt to the appropriate solvent and chemical reaction conditions.
 For example, a polysaccharide may incorporate or be decorated with functional groups. The functional groups may be used to enhance or enable the reaction of the polysaccharide with other functional groups. Functional groups could be, for example, groups that are commonly found to be useful in chemical reactions in the life sciences, e.g., amines, carboxyls, hydroxyls, and sulfhydryls. Functional groups could be, for example, groups that are commonly useful in the polymer sciences, e.g., monomers, free radical monomers, photoinitiated monomers, thermally initiated monomers, chemically initiated monomers, acrylates, and methacrylates. Functional groups could be, for example, groups that are reactable with nucleophiles, e.g., isocyanates, epoxides, sulfhydryls, succinimide esters, and N-hydroxy succinimide esters. Functional groups could be, for example, groups that are photoactivatable, e.g., azides. Other examples of functional groups are maleimides, carbodiimides, carbonyl diimidazole, and acid anhydrides.
 For example, the reaction depicted in FIG. 9 may involve the incorporation of a functional group to the O-MPSAC that can undergo a further chemical reaction by a photo initiated reaction. Or, for example, the second chemical reaction in FIG. 10 may include chemically attaching the W-MPSAC to the surface of a material that contains reactive species. Or the second chemical reaction in FIG. 10 may involve the incorporation of a functional group to the W-MPSAC that can undergo a further chemical reaction by free radical process or by a photo initiated reaction or chemical coupling reaction to various polymers. The first chemical reaction in FIG. 10 could alternatively involve the incorporation of a functional group to the polysaccharide that can undergo a further chemical reaction by free radical process or by photo initiated reaction. The first chemical reaction in FIG. 10 could also involve chemical coupling reactions to various other chemical moieties.
 Other embodiments involve activating a surface by decorating the surface with functional groups that are reactive towards functional groups on a polysaccharide or polysaccharide complex. For example, a surface may be activated with a photoactivatable or polymerizable group. Examples include azide, methacrylate, isocyanate, and epoxide. Alternatively, a polysaccharide or polysaccharide complex may be activated by decorating the polysaccharide with a functional group that is reactive towards a functional group on a surface. For example, a polysaccharide may be activated with a photoactivatable or polymerizable group; examples include azide, methacrylate, isocyanate, and epoxide. Alternative functional groups are, e.g., amines, hydroxyls, carboxyls, sulfhydryls, isocyanates, carbodiimides, carbonyl diimidazoles, and acid anhydrides. The reactions may be performed in organic and/or non-organic solvents, with the polysaccharides being complexed or uncomplexed as needed for the particular chemical reaction.
 Some embodiments relate to the use of photoactivatable groups. Photoactivatable groups respond to specific applied external stimuli to undergo active specie generation with resultant covalent bonding to a nearby chemical structure. Photoactivatable groups retain covalent bonds unchanged under conditions of storage but, upon activation by an external energy source, form covalent bonds with other molecules. Photoactivatable groups may generate active species such as free radicals and particularly nitrenes, carbenes, and excited states of ketones upon absorption of electromagnetic energy. Photoactivatable groups can be chosen to be responsive to various portions of the electromagnetic spectrum, e.g., ultraviolet and visible portions of the spectrum.
 Photoactivatable groups may be, e.g., aryl ketones are preferred, such as acetophenone, benzophenone, anthraquinone, anthrone, and anthrone-like heterocycles (e.g., heterocyclic analogs of anthrone), or their substituted (e.g., ring substituted) derivatives. Examples of aryl ketones include heterocyclic derivatives of anthrone, including acridone, xanthone, and thioxanthone, and their ring substituted derivatives.
 The azides constitute a class of photoactivatable groups and include derivatives based on arylazides such as phenyl azide and particularly 4-fluoro-3-nitrophenyl azide, acyl azides, azido formates (e.g., ethyl azidoformate), phenyl azidoformate, sulfonyl azides (e.g., benzenesulfonyl azide), and phosphoryl azides (e.g., diphenyl phosphoryl azide and diethyl phosphoryl azide). Also, diazo compounds are contemplated, including diazoalkanes (e.g., diazomethane and diphenyldiazomethane), diazoketones (e.g., diazoacetophenone and 1-trifluoromethyl-1-diazo-2-pentanon), diazoacetates (e.g., t-butyl diazoacetate and phenyl diazoacetate), and beta-keto-alpha-diazoacetates (e.g., t-butyl alpha diazoacetoacetate). Other photoactivatable groups include the diazirines (e.g., 3-trifluoromethyl-3-phenyldiazirine), and ketenes (e.g., as ketene and diphenylketene).
 Formation of Polysaccharide Polymers
 Polysaccharide polymers are made of at least two polysaccharides that have been synthetically reacted to form the polysaccharide polymer. Polysaccharide polymers may also include monomers, oligomers, polymers, or macromers. For example, polyesters, polyamines, polyethylenes, polyethylene oxides, and polyurethanes may be incorporated in to the polysaccharide polymers. Such polysaccharide-polymers may include functional groups that are capable of further reaction, for example, monomers, electrophiles, photoactivatable, or photopolymerizable groups. The polysaccharide-polymers may be branched and/or cross-linked. Some embodiments have a molecular weight of at least about 25,000, 100,000, 250,000, or 500,000. Other embodiments have a molecular weight in the range of about 50,000 to about 5,000,000; a person of ordinary skill will in the art will understand that all molecular weights and ranges within the explicit ranges are contemplated, and are within this disclosure. Embodiments include a molecular weight range of about 200,000 to about 2,000,000, from 300,000 to 1,000,000, and from 400,000 to 800,000. Other embodiments have a molecular weight outside of these explicit ranges. A plurality of polymers typically have a range of molecular weights that are distributed about an average molecular weight for that group of polymers. The molecular weight ranges herein may be directed to average molecular weights for a plurality of polymers. Thus, for some embodiments, a plurality of polymers having an average molecular weight that falls within a range is contemplated.
 Chemical reactions, use of functional groups, and solvents may all be chosen as needed to form the desired structure. O-MPSAC and/or W-MPSAC may be used in such reactions to form polysaccharide-polymers. The polysaccharides may be complexed or decomplexed as needed to perform the chemical reactions. Specific examples are presented below.
 Polysaccharide-polymers made of a plurality of polysaccharide macromers are distinct from a polysaccharide that has been derivitized with a functional group to form a macromer. The polymers are made of more than one macromer. The polysaccharide-polymers have a synthetic polymeric backbone formed by the polymerization of the polymerizable groups. The polysaccharide-polymers have a higher molecular weight than the macromers because they are the product of joining at least two, and possibly multiple, macromers. The increased molecular weight is a factor that affects their solubility and their stability when adsorbed or otherwise attached to a surface. The polysaccharide-polymers may have a linear, branched (multi-armed), or cross-linked structure, with the structure being dependent upon the reaction conditions.
 A polymer such as a polysaccharide-polymer has different properties than the unpolymerized starting material, or a typical polysaccharide; these properties can be, e.g., molecular weight, presence or degree of branching, cross-linking, and the type of bond between the polysaccharides. These properties may be manipulated to achieve a polymer that has desired adsorptive properties. An unbranched, uncrosslinked polymer is typically poorly adsorptive. A multi-armed crosslinked or branched polymer, however, is highly adsorptive and stable. But a polymer that is too highly crosslinked may become insoluble such that it comes out of solution. An insoluble polymer may be difficult to use in coating techniques.
 Selection of Molecular Weight, Cross-linking, and Branching of Polymers and Polysaccharide Polymers
 A suitable degree of branching or cross-linking that is appropriate is determinable by empirical processes. Techniques for readily determining the molecular weight of large crosslinked polymers are not readily available. Therefore, it is usually necessary to perform empirical procedures to develop multi-armed polymers with appropriate branching or crosslink densities. For example, a 100,000 molecular weight polysaccharide that has been decorated with between two to ten polymerizable groups can be placed into solution in five samples that vary in concentration by a factor of ten. Each sample is exposed to initiating conditions to react all of the polymerizable groups. A surface or portion thereof that is to be coated, e.g., a wire or tube, is exposed to the samples for a set time, such as between about 2-10 hours. The surface is removed, rinsed in aqueous solution, and tested for adsorption. The sample that causes the highest amount of adsorption is identified and the procedure is repeated with a new concentration range built around the initially selected solution. Samples wherein the polymers fall out of solution are rejected as having polymers that are too highly crosslinked. A range of parameters may be varied to ascertain desirable conditions, including, for example, starting polymeric molecular weight, number of polymerizable groups per polymer, and solution concentrations. This optimization procedure is applicable for making polymeric, linear, high-molecular weight, and multi-armed polysaccharides as well as linear, multi-armed, and high molecular weight vinylpyrrolidones (as described below). Other approaches can be used by a person of ordinary skill in the art based on the disclosure herein, such as by varying the number of solutions in the empirical procedure or other reasonable modification in the empirical procedure.
 Polysaccharide Complexes in a Layer on a Surface
 Polysaccharide complexes may be used as part of a process for creating a layer on a surface or selected portion thereof. A layer refers to a coating, hydrogel, or other material that is in intimate association with a surface or another layer. Thus a layer could be a coating of molecules on a surface or a hydrogel or other three-dimensional structure that is associated with a surface. Polysaccharide complexes may be part of the process of preparing a polysaccharide that is subsequently layered onto a surface. Alternatively, a polysaccharide complex may be coated onto a surface and the surface may subsequently be decomplexed, leaving the polysaccharide associated with the coating. Polysaccharide polymers may be used in a layer, with the polymers being in a complexed or uncomplexed form. Alternatively, polysaccharides may be used in a layer in a complexed on uncomplexed form. Components of a layer may include, e.g., a polysaccharide, polysaccharide complex, non-polysaccharide polymers, and combinations thereof.
 Certain embodiments are layers of polysaccharide and non-polysaccharide polymers, e.g., branched polymers. A polysaccharide may be combined with other polysaccharides or polymers to make a modified polysaccharide polymer. The polymer may be added by covalent bonding, cross-linking, ionic attraction, or by the creation of other associations. The addition of polymers to the polysaccharide may be advantageously used, e.g., to increase it molecular weight, adsorptive proprieties.
 Layers may be of uniform, non-uniform, or continuously varying thicknesses. Layers may be of any suitable thickness. For example, a layer may have a thickness of the order of less than a micron, a micron, tens of microns, hundreds of microns, a millimeter, or other thicknesses. Layers may fully or partially cover a surface. Multiple layers may be built up on a surface. All of the layers could comprise a polysaccharide, e.g., a modified polysaccharide, or polysaccharide-polymer. Alternatively, layer(s) that comprise a polysaccharide could be interspersed or otherwise mixed with other materials, objects, or layers of other materials.
 Layers may be formed using, e.g., a spray process, a dip process, or other processes known in the art. A dip process entails dipping an object into a solution or melt of the layer components and removing the object. A spray process entails preparing a solution or melt of the layer components and spraying them onto an object. A mold may be used to prepare a layer, especially a thick hydrogel or coating. The components of a layer at least partially solidify on the object to create the layer. Solvents that evaporate at low temperatures and/or pressures can are advantageous in that they can typically be efficiently removed.
 Layers may be associated with a surface by a variety of physical or chemical processes. Certain embodiments are directed to forming layers on surfaces that are at least a part of a medical device. Adsorption, ionic attraction, and covalent bonding are possible mechanisms that can be used by choosing suitable functional groups. Certain embodiments entail associating a layer with a surface and subsequently stabilizing the layer by activating functional groups that are in the layer or on the surface. Examples of such functional groups are thermally initiatable monomers, photoinitiatable monomers, photoactivatable groups, methacrylates, isocyanates, and epoxides. Alternatively, the layer may be simultaneously formed and stabilized; for example, by allowing functional groups on a surface to react with functional groups on a component of a layer. The simultaneous processes could involve immobilization of a polysaccharide or polysaccharide complex to a surface. The simultaneous processes could further involve reactions between functional groups on the components with each and with the surface; for example, polysaccharide complexes comprising polymerizable monomers could be reacted with each other and simultaneously with a surface decorated with suitable monomers.
 Additional Embodiments Related to Polysaccharide Complexation
 An O-MPSAC may be chemically reacted with a surface of a medical device that contains reactive functional groups. The surface's reactive functional groups are able to react with the mucopolysaccharide, which is followed by decomplexation, leaving the mucopolysaccharide bound to the surface. This process is exemplified by forming an organic solvent soluble heparin complex, which can be reacted with but not limited to, isocyanate or epoxide groups, which have been incorporated to the surface of the medical device. The isocyanate or epoxide groups are able to react with free hydroxyl of the heparin to form urethane or ether linkages, respectively. In another manner the isocyanate or epoxide groups are able to react with the free amino groups of the heparin to form urea or substituted amine linkage. On completion of the surface reaction the heparin is decomplexed with salt solution, leaving the heparin chemically bound to the surface in its active form. In the case of heparin, the complexed, chemically modified, and de-complexed heparin is produced in its active anti-thrombogenic form. The organic soluble heparin complex can be applied to the medical device by, e.g., dip coating, spray coating or any other coating process. Alternatively, the surface could be activated with photoactivatable groups that are triggered to react with the polysaccharides.
 Another embodiment is to chemically modify a mucopolysaccharide component of a organic solvent soluble complex (see FIG. 9), to produce chemically activated O-MPSAC, which is activated to be able to undergo further chemical reactions. This process is exemplified by forming an organic solvent soluble heparin complex, which can be reacted with, e.g., isocyanatoethyl methacrylate or methacryloyl chloride. The isocyanatoethyl methacrylate or methacryloyl chloride is able to react with free hydroxyl of the heparin to form methacrylate urethane or methacrylate ester linkage with the heparin. In another manner the isocyanatoethyl methacrylate or methacryloyl chloride is able to react with the free amino groups of the heparin to form methacrylate urea or methacrylate amide linkage with the heparin. Other functional groups may be substituted for methacrylates, e.g, polymerizable or photoactivatable groups.
 The complexed methacrylate heparin of the form described above can undergo free radical polymerization with other macromers and/or monomers, either in solid state, gel state, in solution, in emulsion, or in suspension. The final polymer, where the heparin complex is attached to the polymer backbone, can be used to coat medical devices. The heparin is then decomplexed with salt solution, leaving the active heparin coated onto the medical device. Other polysaccharides or proteoglycans may be substituted for heparin. And other polymerizable groups may be substituted for methacrylates.
 A complexed methacrylated heparin macromer may be decomplexed with salt solution, giving a methacrylated heparin macromer that can undergo polymerization with other macromers and/or monomers, by e.g., solution, emulsion, or suspension polymerization. The final polymer, where the active heparin is attached to the polymer backbone, can be used to coat medical devices. Thus the macromers may be polymerized to form larger macromers, polymers, and three-dimensional structures. Or they may be polymerized to form a three-dimensional hydrogel.
 Another embodiment is a chemically modified O-MPSAC, which is able to undergo photochemical reactions. This embodiment is exemplified by forming an organic solvent soluble heparin complex, which can be chemically modified to contain photochemical reactive groups that are able to link to the surface of medical devices. The photochemical reactive groups may be photopolymerizable or photoactivatable, e.g., allyl, vinyl, acrylates, methacrylates, azides, nitrenes, carbenes and excited states of ketones, diazo, azo compounds and peroxy compounds, and such groups as are cited in and those cited in WO 90/00887, which is hereby incorporated herein by reference. An advantage of this method is that photochemical reactive groups may be bound to the heparin complex in organic solvents, and the resultant product may then be dissolved in organic solvent and coated onto the medical device. Then, by applying the appropriate electromagnetic radiation to carry out the photochemical reaction, the photoactivated groups bond to the surface to form the desired coating. Photoactivated bonding is followed by decomplexation, leaving the active heparin bound to the surface.
 Another embodiment comprises of the chemical modification of the mucopolysaccharide in water or DMSO or another equivalent solvent or solvent mixtures. This is exemplified by the reaction of the activated imidazole carbonate of polyethyleneglycol methacrylate (see WO 97/41164, hereby incorporated by reference herein) with heparin to form the heparin-polyethyleneglycol methacrylate. The heparin-polyethyleneglycol methacrylate is then complexed (W-MPSAC, Scheme 2). The W-MPSAC can then undergo further chemical reaction in an organic solvent, e.g., by free radical process or by photo-initiated reaction. Alternatively, the heparin-polyethyleneglycol methacrylate can undergo chemical reactions by free radical process or by photo initiated reaction without the complexation step. The resultant product may then be complexed to form an organic soluble heparin complex.
 An advantage of the modified polymers and the polysaccharide macromers is that they may be reacted with a surface while in solution. For example, modified polymers with polymerizable groups may be polymerized while in solution. Or photoactivatable groups or electrophilic groups may be activated while the polymer or macromer is in solution. These approaches are particularly effective when organic solvents are used because many chemical reactions are much more efficient in organic solvents as compared to aqueous solvents, e.g., electrophile-nucleophile reactions. The present disclosure provides numerous techniques for bringing polysaccharides and modified polymers into solution so that an effective reaction with the surface may be performed.
 Polymers of Vinylpyrrolidone
 Another embodiment is a polymer (or macropolymer) of vinylpyrrolidone. Vinylpyrrolidone polymers or oligomers may be decorated with a polymerizable group using techniques known to those of ordinary skill in these arts to make a vinylpyrrolidone macromer that is polymerizable. The vinylpyrrolidone macromer may be polymerized to make a three-dimensional cross-linked structure. Or the vinylpyrrolidone macromer may be polymerized to make a larger polymer. Further, cross-links may be incorporated into the larger polymer. As discussed above, structures that are branched, cross-linked, and/or have a relatively high MW have advantageous properties for use as a coating.
 The vinylpyrrolidone polymer (or macropolymer) is useful for making lubricious coatings. The coatings may be made on essentially any surface for a medical device. The coatings may be made, for example, by drying a solution of vinylpyrrolidone polymer, multi-armed polymer, or macromer onto an object. The macromer may subsequently be polymerized. Alternatively, the object may be covered with the macromer and the macromer polymerized to make a coating or encapsulating membrane. The vinylpyrrolidone macropolymer or the macromer may be combined with other monomers/polymer/macromers, especially those described herein.
 5 g sodium heparin (Celsus Laboratories, Inc, USP lyophilized from porcine intestinal mucosa) was dissolved in 80 ml de-ionized water and allowed to stir for 1 hour in a 250 ml beaker.
 8 g benzalkonium chloride (Aldrich Chemical Company, Inc) was allowed to dissolve in 80 ml de-ionized water with gentle warming (40-50° C.) on a magnetic stirrer hotplate for 1 hour and then allowed to reach room temperature.
 To the above vigorously stirred solution of sodium heparin, the benzalkonium chloride solution was added. A white precipitate immediately formed and the suspension was further stirred for 1 minute. The precipitate was filtered through a Whatman qualitative filter paper (grade 1).
 The white precipitate was collected from the filter paper and re-suspended in 400 ml de-ionized water and allowed to stir for 20 minutes. The suspension was filtered as above and re-suspended in 400 ml de-ionized water and filtered again. The precipitate was once again suspended in 400 ml de-ionized water and then poured into a dialysis membrane Cellu Sep:MWCO 3,500 and dialyzed against 10 L de-ionized water for a minimum of 16 hours.
 The precipitate was collected and dried on a glass dish in a vacuum oven at 60° C. for 12 hours.
 Dry gray-yellow crystals were obtained with a yield of 10 g.
 3 g 4, 4′-methylenebis (phenyl isocyanate) (MDI) (Aldrich Chemical Company, Inc) was dissolved in 100 ml anhydrous tetrahydrofuran (THF). Polyurethane tubing was dipped in to the above solution for 30 seconds and was then allowed to stir-dry at 60° C. for 1 hour.
 5 g of dry crystals of complexed heparin from Example 1 were dissolved in 100 ml anhydrous dichloromethane (DCM). The MDI coated polyurethane tubing was dipped into the DCM solution of complexed heparin for 30 seconds and then allowed to air-dry for 2 hours.
 The tubing was dyed with 0.075% w/v pH 8.5 toludine blue aqueous solution for 30 seconds and washed with de-ionized water. A very faint purple color due to the complexed heparin was observed. The tubing was then immersed in 25% w/v solution of sodium chloride at 40° C. for 30 minutes. The tubing was washed with de-ionized water and again dyed. An intense dark purple color due to complexed heparin was observed.
 In a similar experiment where the polyurethane tubing was not initially coated with MDI and was de-complexed in sodium chloride solution (as above) no purple coloration due to the heparin was observed.
 1. Heparin Methacrylate (Methacryloyl Chloride)
 5 g of dry crystals of complexed heparin from Example 1 were dissolved in 100 ml anhydrous DCM in a 250 ml quickfit conical flask. To this was added 0.1265 g (1.25×10−3 moles) triethylamine.
 Methacryloyl chloride (Aldrich Chemical Company, Inc) was distilled under reduced pressure to obtain a very pure sample. 0.1306 g (1.25×10−3 moles) of the above distilled methacryloyl chloride was dissolved in 30 ml anhydrous DCM and placed in a stoppered quickfit pressure equalizing funnel above the vigorously stirred solution of DCM containing the complexed heparin. The methacryloyl chloride solution was added drop-wise over a period of 30 minutes to the complexed heparin solution in DCM.
 The DCM was rotary evaporated and the complexed heparin methacrylate was dried in a vacuum oven at 40° C. for 2 hours.
 Complexed methacrylate was characterized by 1H and 13C.
 0.2 g of the above complexed heparin methacrylate was dissolved in 10 ml 2-hydroxyethylmethacrylate and to this was added 0.02 g ethylene gluol dimethacrylate and 0.02 g 2,2′-azobis (2,4-dimethylvaleronitrile) (Dupont). The above clear solution was degassed for 30 minutes.
 The above polymerization mixture was poured into a polypropylene concave mold and then a polypropylene convex mold was placed onto the concave mold allowing the excess solution to overflow, thereby uniformly filling the space between the concave and convex molds. The sealed molds were then heated to a temperature of 65° C. for 4 hours and then at 110° C. for 1 hour.
 The molds were cooled and opened to obtain a clear dehydrated rigid hydrogel. These were then hydrated in de-ionized water for 10 hours, after which they were placed in boiling solution of 25% w/v sodium chloride aqueous solution for 1 hour and then equilibrated in de-ionized water. The hydrogel was dyed (as in Example 2) and a uniform intense dark purple coloration was observed throughout the hydrogel.
 In a similar experiment where no methacrylate of the heparin was formed, a hydrogel of the complexed heparin alone was formed. After boiling in 25% w/v sodium chloride aqueous solution and equilibrating in water the hydrogel was dyed. Dark purple precipitated particles of heparin could be observed on the surface of the hydrogel and were easily washed away with de-ionized water, leaving a blue coloration to the hydrogel. This coloration is identical to the hydrogels formed without any complexed heparin.
 2. Heparin Methacrylate (Isocyanatoethylmethacrylate)
 5 g of dry crystals of complexed heparin from Example 1 were dissolved in 100 ml anhydrous DCM in a 250 ml thick-walled glass bottle with cap. To this was added 0.194 g (1.25×10−3 moles) 2-isocyanatoethylmethacrylate (Aldrich Chemical Company, Inc) and 0.05 g dibutyltin dilaurate (Aldrich Chemical Company, Inc). The cap was screwed on tight and the solution was stirred for 16 hours at 40° C.
 The DCM was rotary-evaporated off and the product dried in a vacuum at 40° C. for 2 hours.
 As in Example 3(1), hydrogels were made and the heparin was decomplexed and dyed. Again, a uniform intense dark purple coloration was observed throughout the hydrogel whereas the complexed heparin with no methacrylate coupling heparin particles precipitated on the surface of the hydrogel and were easily washed away with de-ionized water.
 5 g of Heparin methacrylate from example 3 i) was dissolved in 100 ml of 2-propanol. To this was added 20 g methoxypolyethlyeneglycol 2000 methacrylate (MPEG 2000 MA) (Inspec U.K.) and 3 g of 2-hydroxyethylmethacrylate.
 A 250 ml, 3-necked reaction vessel equipped with stirrer, thermometer, condenser and nitrogen inlet tube was charged with 100 ml 2-propanol. To the 2-propanol was added 5 g of heparin methacrylate [from example 3 (i)], 20 g methoxypolyethyleneglycol 2000 methacrylate (MPEG 2000 MA) (Inspec U.K.) and 3 g 2-hydroxyethylmethoxylate. (HEMA) The 250 ml, 3 necked reaction vessel was placed in a silicone oil bath at 120° C. and the 2-propanol was stirred gently and nitrogen was bubbled through the solution (100 cm3/min).
 When the temperature in the 250 ml, 3-necked reaction vessel reached 75° C., 0.25 g 2,-2′-azobis (2, 4-dimethylvaleronitrile) was added and the stirrer speed was increased to 750 rpm. After approximately 15 minutes a very viscous solution was obtained and the reaction was allowed to continue for 30 minutes, periodically added 30 ml 2-proponal to dilute and reduce the viscosity of the solution. In total four 30 ml aliquots of 2-propanol was added.
 The polymer was cooled down to room temperature and the isopropanol rotary evaporated off and the polymer dried in a vacuum oven at 40° C. for 8 hours.
 GPC (Gel permeation chromatography) showed that the average molecular weight of the polymer was approximately 400,000 when using polyethylene glycols as standards.
 3 g MDI was dissolved in 100 ml of anhydrous THF. A 150 mm long and 2 mm diameter polyurethane tube was dipped into the above solution for 30 seconds and was then allowed to air dry at 60° C. for 1 hour.
 5 g of heparin copolymer from Example 4 was dissolved in anhydrous DCM. The MDI coated polyurethane tubing was dipped into the DCM solution of copolymer for 30 seconds and was then allowed to air at 60° C. for 30 minutes and then at room temperature for 16 hours.
 The tubing was immersed in 25% w/v sodium chloride solution at 40° C. for 30 minutes. The tubing was washed with de-ionized water and dried with toludine blue (example 2). A homogenous intense dark purple coloration was obtained on the tubing. In addition the polyurethane tubing was completely wetted and very lubricious. The lubricity did not diminish even when it was rubbed between forefingers and thumb 20 times. The tubing was then immersed in phosphate buffered saline at 50° C. for 16 hrs and then rubbed between the forefinger and thumb. Again there was no observable discharges in lubricity.
 As in Example 4, 5 g heparin methacrylate was copolymerized, except that instead of MPEG 2000 MA and HEMA, 20 g N-vinylpyrrolidone was used as the co-monomer in 2-propanol. The polymerization was allowed to continue for 1 hour.
 The average molecular weight of the copolymer was approximately 300,000 as determined by GPC.
 As in Example 5 a 150 mm long and 2 mm diameter polyurethane tubing was dipped into a 3% w/v MDI in THF solution and air dried at 60° C. for 1 hour.
 The tubing was then dipped in a 5% w/v heparin—vinylpyrrolidone copolymer in DCM for 30 seconds and then air dried at 60° C. for 2 hours.
 The tubing was then immersed in 25% w/v sodium chloride solution at 40° C. for 30 minutes. The tubing was washed with de-ionized water and dyed with toludine blue. A homogenous intense dark purple colorization was obtained. The polyurethane tubing was completely wetted and extremely lubricious. The lubricity was equivalent to that obtained for in Example 5.
 5 g heparin methacrylate was copolymerized with 20 g acrylic acid in 2-proponal. The reaction conditions were the same as in Example 6. The average molecular weight was determined to be 400,000 using GPC and polyacrylic acids as standards.
 As in Example 5 a 150 mm long and 2 mm diameter polyurethane tubing was dipped into a 3% w/v MDI in THF solution and air dried at 60° C. for 1 hour.
 The tubing was dipped into a 5% w/v heparin-acrylic acid copolymer in methanol/dimethylacetamide solution (90:10) for 30 seconds and the air dried at 60° C. for 3 hours.
 The lubricity of the surface of the polyurethane tubing was very similar to that in Example 5 and 6 after washing in 25% w/v sodium chloride solution (as performed in Example 5 and 6).
 XPS confirmed the presence of heparin on the surface as SO4 groups be detached on the surface of the polyurethane tubing.
 PVC tubing 150 mm long 2 mm in diameter were coated with the heparin copolymers synthesized in Example 4, 6 and 7. The coating conditions for the respective heparin copolymers were identical to the Examples 5, 6 and 7 respectively.
 In all cases a very durable lubricious coating (when wet) was obtained and the presence of heparin on the surface of the PVC was detected either using toludine blue or XPS.
 A polyurethane tube (150 mm long, 2 mm diameter) was dipped into a 2% w/v solution of poly [1,4-phenylenediisocyanate-co-poly (1,4-butanediol)] diisocyanate (Aldrich Chemical Co.) in anhydrous THF and allowed to air dry at 60° C. for 2 hours.
 2 g of heparin complex (from Example 1), 2 g polyethylene oxide (M. W. 100,00) (Aldrich Chemical Co.) and 0.25 g MDI were dissolved in 100 ml anhydrous DCM. The polyurethane tubing was dipped with the above solution for 30 seconds and allowed to air dry at 60° C. for 1 hour and then at 22° C. for 16 hours.
 The polyurethane tubing was immersed in 25% w/v sodium chloride solution at 40° C. for 30 minutes and then washed with de-ionized water and then dyed with toludine blue. An intense homogenous dark purple color developed on the tubing and when wet the tubing was highly lubricious. The lubricity was comparable to in Example 5, 6 and 7.
 A similar experiment to the one in Example 9 was conducted with polyvinylpyrrolidone (M. W. 1,300,000) (Aldrich Chemical Co.) with all the experimental conditions the same. Again the results showed that heparin was present on the surface and the polyurethane tubing was highly lubricious and comparable to the coating in Example 9.
 A PTFE tube measuring 3.8 mm O.D; length 8 cm was placed on a rotatable mandrel. Placed over the PTFE was a knitted fabric tube measuring 4.0 mm I.D; length of 7 cm.
 A methacrylated polyvinylpyrrolidone (MPVP) was made by dissolving 6 g of polyvinylpyrrolidone (PVP) in dichloromethane. A solution of isocyanato ethyl methacrylate (0.8 g dissolved in 20 ml of dichloromethane) was added drop wise to the stirred solution of PVP. The reaction was allowed to proceed for a further 2 hours. Dichloromethane was then rotary evaporated and the MPVP formed was used in the vascular graft formulation as per the following: Hydoxyethyl Methacrylate 18 g; Heparin Methacrylate (as from example 3) 1 g; MPVP 0.6 g; Ethylene glycol dimethacrylate 0.08 g; thermal initiator 0.12 g (Vazo 52).
 The vascular graft formulation was then added drop wise onto the rotating knitted fabric until a homogeneous viscous film was formed, which totally encapsulated the fabric. Then the graft was exposed to UV light from a medium pressure mercury arch lamp for 10 minutes; then the graft was placed in a vacuum oven for 3 hours at 70° C.
 The heparin contained in the graft was de-complexed with saturated NaCl solution. The final vascular graft had a hydrogel that totally encapsulated the fabric, which had a water content of 50%. Heparin activity was measured by antithrombin binding assay and was found to be 5 units per 100 mm2.
 This example shows that vascular grafts made as described herein are effective for use with patients. A 28 day animal study was carried out on two pigs. The vascular grafts as made according to Example 11 were implanted in the carotid artery of each animal using aseptic surgical techniques known to those skilled in the art. The carotid arties were visualized at day nine, and day 28 using a medical X-ray arteriogram. The arteriogram showed that the grafts were patent with endothelial cells, were clean, and showed no evidence of thrombus at both days.
 This Example describes the production of an azide-polysaccharide complex, which is exemplified by the production of a heparin-azide complex.
 5 g of sodium heparin, end-aminated (containing a terminal amine group; Celsus Laboratories, Inc., Cincinnati, Ohio) was complexed with benzalkonium chloride and purified as in Example 1, except that the product was dried by freeze-drying with a yield of 10 g. The complexed heparin (10 g) was dissolved in anhydrous dichloromethane (100 ml). 0.3 g of 4-azidophenyl isothiocyanate (Aldrich Chemical Company, Inc.) was dissolved in 50 ml anhydrous dichloromethane, and this solution was added drop-wise to the solution containing the heparin complex. After the addition had been completed, the solution was kept in the dark to prevent the reaction of the azide group. Infrared spectroscopy showed the disappearance of the isothiocyanate at 2000 cm−1 due to the reaction of the isothiocyanate with the amine group of the heparin complex.
 This example shows how an azide-polysaccharide complex can be used to coat a variety of surfaces, e.g., by dipping or spraying. The process is exemplified using tubing and stents. From Example 13, the azide-heparin complex dichloromethane solution was used to coat a variety of surfaces by dipping or spraying.
 In particular, a polyurethane tube was dipped into the azide-heparin complex dichloromethane solution, dried, and then exposed to UV light from a medium pressure mercury arc lamp for one minute. The polyurethane tubing was then immersed into a saturated sodium chloride solution for one minute to de-complex the heparin from the benzalkonium chloride. Any heparin not covalently bonded to the surface via the azide group was thereby dissolved in the saturated sodium chloride solution.
 The polyurethane tubing was then washed with de-ionized water and dyed in toluidine blue. A homogenous deep purple color was observed indicative of heparin being bound to the surface.
 Similarly, polyethylene, polypropylene, polyamide and polyester tubing was coated with azide-heparin as above and then exposed to UV. The heparin was then de-complexed in saturated sodium chloride solution and dyed in toluidine blue. The appearance of a homogenous deep purple coloration showed that heparin was chemically bound to the surface.
 Expanded PTFE tubing was sprayed with the above azide-heparin complex solution and then processed similarly. A homogenous but lighter purple coloration was observed after dying with toluidine blue. However, when the process was repeated onto the same PTFE tubing, a deep-purple coloration was observed after dying in toluidine blue.
 Stents were coated with a variety of different polymer coatings that are used as drug delivery systems, such as polyethylene-vinyl acetate, polyethylene-vinyl alcohol, polyurethane and polydimethyl siloxane. These stents were then sprayed with the azide-heparin complex from a solution of dichloromethane and isopropanol (50:50). The sprayed stents were exposed to UV light and then decomplexed, washed and dried.
 The stents were then crimped on to a Percutaneous Transluminal Coronary Angioplasty (PTCA) catheter, which was then expanded. Light microscopy showed that the polymers contained no cracks or any imperfection upon expansion of the stents. Dying the stents with toluidine blue gave a homogenous deep purple coloration indicative of chemically linked heparin.
 This example shows that an azido-polymer, e.g., an azide-polysaccharide complex, can be used in combination with a second polymer, e.g., a hydrophilic polymer, to create a coating on surface that includes the azido-polymer and the second polymer. 0.5 g of polyethylene oxide (MW 4,000,000) was added to a 150 ml solution of the azide-heparin complex in dichloromethane from Example 13, and allowed to dissolve.
 A polyurethane tubing was dipped coated with the above solution and then exposed to UV light for one minute. The heparin was de-complexed, as described previously, and the tubing was washed with de-ionized water and dyed with toluidine blue. The stained stent had a homogenous deep-purple coloration and was found to be very lubricious. The above procedure was repeated with polyvinylpyrrolidone (5 g; MW=1,000,000) and it was again found that heparin was chemically linked to the surface, which was also lubricious.
 The association of polyvinylpyrrolidone or polyethylene oxide with a surface gives the surface a lubricious character. The lubricious character of the surfaces therefore showed that polyethylene oxide or polyvinylpyrrolidone was associated with the surface.
 This example shows that an azido-polymer can be complexed with monomers or macromers to make a layer of a hydrogel on a surface. 20 g hydroxyethylmethacrylate (HEMA) HEMA, 10 g MPEG2000MA (PEG with MW 2000 having a methyl on one end and a methacrylate on the other end), 0.3 g ethylene glycol dimethacrylate, and 0.2 g benzoin methyl ether (an UV initiator) were added to a 50 ml solution of azide-heparin complex in dichloromethane (from Example 13).
 This solution was then dip coated onto a polyurethane tubing and exposed to UV light for 5 minutes. Then, the heparin was de-complexed (as described previously), washed and exposed to toluidine blue dye. The tubing surface was found to stain with a homogenous deep-purple coloration. Scanning electron microscopy (SEM) showed that a layer of about 10 microns in thickness was present. The presence and deep color of the stain after the washing step indicated that the heparin was present in a high concentration in the layer. Moreover, the SEM results showed that a three-dimensional structure was built up on the surface, indicating the monomeric groups were polymerized to form a hydrophilic polymer.
 The embodiments set forth herein are exemplary of the inventions and not intended to be limiting in scope. Patents, patent applications, and articles mentioned in this application are hereby incorporated herein by reference.
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|U.S. Classification||424/423, 514/54|
|International Classification||A61L27/00, A61L33/08, A61L33/10, A61L27/28, A61L27/20, A61L27/50, A61L27/34, A61L33/00, A61L27/52, C08B37/00|
|Cooperative Classification||C08B37/0075, C08B37/0069, C08B37/0063, A61L33/0064, A61L33/0011, A61L27/52, A61L27/507, A61L27/34, A61L27/20, Y10S514/944, Y10S514/953, Y10S514/822|
|European Classification||A61L27/34, C08B37/00P2D, A61L33/00H2, A61L27/50E, C08B37/00P2G, A61L27/52, A61L33/00H7, C08B37/00P2, A61L27/20|