This computer-controlled liquid polyurethane dispenser by Max Machinery (Healdsburg, CA) can produce exactly the same polyurethane mix for each shot.
Polyurethanes are used often in medical device applications, and their use continues to grow. But what sustains this growth? Compared with other polymers, polyurethanes often require sophisticated manufacturing processes and are more expensive on a price-per-pound basis. For example, the average flexible PVC compound sells for 85 cents per pound. Polyurethanes sell for 10–20 times that amount. So what is it that motivates medical device developers to use them?
The answer is quite simple: polyurethanes can be used in applications where other materials do not work. Polyurethanes are among the most versatile construction materials that can be formulated for medical devices. They are tough, biocompatible, and hemocompatible. They can be strong elastomers or rigid plastics, and they can be processed using extrusion, injection molding, film blowing, solution dipping, and two-part liquid molding.
Polyurethanes’ unique chemistry gives them this versatility. They are segmented polymers, meaning they have a soft segment that provides flexibility and a hard segment that provides strength. Polyurethanes are made from three basic building blocks: the backbone, the diisocyanate, and the chain extender. The backbone, usually a long chain molecule, provides flexibility to the polymer. The diisocyanate and the chain extender combine to form the hard segment, which acts as a cross-link. It provides the polymer with high tensile strength and high elongation.
Polyurethanes are made from either aromatic or aliphatic diisocyanates. Aromatic diisocyanates contain benzene rings, which create polyurethanes that are generally tougher, stronger, and less costly than the aliphatics. The aromatics generally have tougher hard segments, which are more chemically resistant and give rise to higher tensile strength and elongation than aliphatics. Aliphatic diisocyanates are made with hydrocarbon backbones and contain no benzene rings. Aliphatic polyurethanes make strong polymers but lack the chemical resistance of aromatics. They are more expensive than aromatics and are used primarily in applications that require good light stability. There are thousands of possible combinations of the basic building blocks used to create aromatic and aliphatic polyurethanes, thereby providing device engineers with a myriad of options for their products.
Several types of polyurethane are appropriate for medical applications, including the following:
• Liquid polyurethanes for hollow-fiber devices.
• Polyurethanes for dip-molding.
• Polyurethane coatings.
• Biostable polyurethanes.
• Thermoplastic polyurethanes.
With so many device options and material formulations, the use of polyurethanes in the device industry can only increase. To benefit from their flexibility, manufacturers must know when and how to use these versatile polymers.
Liquid Polyurethanes in Hollow-Fiber Devices
The polyurethane in this hollow-fiber filter by CHF Solutions (Brooklyn Park, MN) separates the membranes to let blood and medicine flow through it without interacting.
Choosing the Material. Hollow-fiber membranes revolutionized the way blood-processing membrane devices, such as hemodialyzers, oxygenators, and hemoconcentrators, are constructed. Before hollow-fiber membranes were introduced, most of these devices used flat-sheet membranes. The flat-sheet membrane devices were difficult to build and had a significant leakage problem. Hollow fibers could be made economically and were structurally reliable. The challenge to device developers was how to integrate them into a membrane device. Manufacturers had to determine how the fibers could be held, thus separating the inside of the fiber membrane from the outside. It is necessary to separate the two sides of the membrane to allow blood to flow on one side of the membrane and a therapeutic medium to flow on the other side. The solution was to encapsulate the membrane ends with a two-part liquid polyurethane. The potting material acts as a structural unit that separates the two sides of the membrane.
The device is placed in a centrifuge and the two-part liquid polyurethane is spun in, encapsulating the individual fibers and separating the membranes. The low viscosity of the polyurethane (500–3000 cP) allows it to easily encapsulate each closely packed fiber. Once inside the potted area, the polyurethane cures into a tough, strong material that bonds both to the case and to the fibers. After the polyurethane is cured, the ends of the bundle are cut off, exposing the blood side of the membrane. Two-part polyurethane systems are available off-the-shelf from several manufacturers, or they can be custom formulated to meet exact requirements.
Two-part polyurethanes are the only option suitable for this application. Other liquid-polymer systems, such as liquid silicone, epoxy, and polyester, have been tried, but none have worked. Initially, two-part silicones were evaluated for potting hollow-fiber devices, but their shortcomings soon became evident. The silicone rubber did not stick to the fibers or to the case. This caused leakage between the blood path and the fluid path. In addition, the silicones were very weak and tore easily. Epoxies and polyesters have low viscosities and bond well, but they are too rigid to allow cutting of the fiber-bundle ends.
Processing Considerations. Two-part polyurethanes used to pot hollow-fiber devices must be conditioned before they are mixed. The process removes dissolved gases and heats the components to a uniform temperature.
To condition the materials, the two components are individually manufactured and packaged in a dry nitrogen environment to exclude moisture. Moisture can react with the active isocyanate groups, compromising the integrity of the cured material.
Nitrogen gas poses a different problem. If the majority of the dissolved nitrogen is not removed, the exothermic reaction during curing will push it out of the solution, creating bubbles. So, vacuum degassing can remove the nitrogen. Optimally, a thin-film degasser should be used. Such a degasser comprises a series of flat plates over which a thin film of polyurethane is passed. Polyurethane in a thin film is more easily degassed than the bulk material, significantly reducing the degassing time required.
This cross section shows how the hollow fibers are held open. Encapsulating the fiber ends with a two-part liquid polyurethane creates a structural unit that separates the membranes from each other.
Consistent and accurate heating of the reactants is paramount to a process that has a high yield and minimal rejects. Ideally, the reactants in a two-part polyurethane are heated, circulated through heat-traced lines to a heated mixing head, and brought back to the main tank. Consistent heating can be further enhanced by using an agitator in the main holding tank. The optimum temperature for the reactants must be determined according to the system use, but generally is around 50°–60°C.
The two components of a polyurethane must be mixed together in the exact stoichiometric amounts determined by the material supplier. If the variations are greater than 1% in the stoichiometric ratio, the final properties of the cured material may differ from specifications. Improper ratios can cause lower tensile strength and create extractable materials that can cause toxicity. To ensure proper ratios, positive-displacement gear pumps are the most reliable for this application. Piston pumps can be used but are not recommended because they do not reliably meter the flow as accurately as gear pumps do. Verification of the actual flow from a gear pump can be determined with a flowmeter, and outputs from the flowmeter can help regulate the gear pump dc-drive stepper motors.
Second only to metering, mixing is a critical step in dispensing polyurethanes for medical applications. Static mixers may be acceptable when the viscosities of the two streams are similar. A 32-element mixer usually will give an acceptable mix. For ratios greater than 2:1, dynamic mixers are a better choice.
A typical dynamic mixer turns at 10,000–15,000 rpm. It can turn the two streams several thousand times as they pass through the mixing head. The resulting polyurethane that is dispensed is clear and homogeneous.
Accuracy in the amount of polyurethane dispensed each time is also very important. The length of time to dispense the required amount of polyurethane must also be considered.
Gear-pump polyurethane dispensers can perform all of these functions. Typically, this type of computer-controlled polyurethane process system can produce exactly the same mix each time. The equipment heats the reactants to within ±1°C, and a vacuum system degasses the reactants. Then, the metering system flows the reactants into a high-speed mixing head, which produces a homogeneous mix from beginning to end of the dispensing operation.
Polyurethanes for Dip Molding
Polyurethanes are also well suited for manufacturing dip-molded devices such as balloons, probe covers, gloves, and condoms. Although natural rubber latex (NRL) is much less expensive, the advantages of using polyurethanes are significant. First and foremost is the lack of extractable chemicals in polyurethanes. Unlike NRL, polyurethanes are pure polymers, meaning that all the ingredients are chemically bound to each other. NRL, however, is loaded with natural products, such as tree proteins and vulcanizing chemicals. These agents may cause dermatitis, allergic reactions, and, in the worse case, anaphylactic shock that can lead to death. Because they wear gloves to perform many of their tasks, healthcare professionals are particularly susceptible to these conditions.
Mechanically, polyurethanes have higher tensile strength (up to 7000 psi), better tear strength (500 psi, ASTM D-624 Die C), and better abrasion resistance (Taber abrasion of 25 mg lost per 1000 cycles) than NRL. Because of these good mechanical properties, gloves made from polyurethanes can be made as thin as 0.005 in. and are resistant to wear and tear. Thinner gloves provide greater tactility and cause less hand fatigue for a user than thicker ones. Polyurethanes are also much more biocompatible than NRL, which fails many of the ISO 10993-1 tests for biocompatibility.1 But properly selected polyurethanes can pass all the tests, including cytotoxicity, hemocompatibility, mutagenicity, and pyrogenicity.
Thermoplastic polyurethanes are suited for most dip-molding applications. They are linear polymers that dissolve easily in solvents like tetrahydrofuran, acetone, and methylethylketone. In some cases, the polyurethane polymer is formed in solvent, as in the case of Lycra. Lycra is a highly elastic polyurethane that is often used to make medical gloves. Lycra is made by chain extending, or solution curing, the polyurethane in solvents such as dimethyformamide or dimethylacetamide. The chain extender is an amine and results in urea linkages in the hard segment. Alternative curatives for use with polyurethanes, all amines produce urea linkages. These linkages are stronger than linkages formed from diols, which are used as curatives in most polyurethanes. The resulting polymers have better recovery after elongation and lower modulus at elongation (200–300 psi at 500%). They are also insoluble in most solvents.
High-strength polymers that have properties suitable for medical devices are not necessarily the best material choices for implanting in the body. For example, heparin coatings are added to a variety of catheters made from PVC, polyamide, and polyurethane to enhance their thromboresistance. Similarly, specially formulated polyurethanes that are suitable for coatings are being used extensively in medical device manufacturing. Again, their versatility makes them suitable for diverse applications. Polyurethane coatings can be hydrophilic, hydrophobic, antimicrobial, nonthrombogenic, drug releasing, or lubricious. The coatings are dissolved in a solvent and are applied to the substrate using dipping or spraying techniques. The chemistry of a polyurethane coating is the same as that of the bulk polyurethane. That materials’ good mechanical properties yield coatings that are tough, strong, and abrasion resistant.
The range of available options for polyurethane coatings continues to grow. For example, several manufacturers are working on drug-eluting stent coatings. Patients having bare-metal stents have a high rate of restenosis, or blockage. Certain drugs that can lower the restenosis rate are incorporated into the polyurethane matrix and are slowly released once the stent is in place. Drug-release rates are controlled by the ratio of the hard-segment to soft-segment content of the polyurethane and the specific chemistry. However, there are limitations on the drugs that can be released from polyurethanes. These limitations include solubility of the drug in the polyurethane matrix and coating solvents, diffusion rate, and chemistry of the drug. Another critical property for these coatings is their ability to stretch and conform to the metal stent when the stent is deployed. The coating must not tear or break.
This Edwards Lifesciences (Irvine, CA) central-venous catheter made of biocompatible polyurethane has good column strength, so it will not kink during insertion.
Polyurethanes used in long-term implants present device developers with special challenges. Among the first uses of polyurethane for implants was pacemaker leads. The polyurethane used was polyether based and generally performed well. However, failures did occur. The traditional polyether-based polyurethane showed signs of deterioration after being in the body for several years. Certain metals, such as cobalt, catalyze such degradation.2 This is particularly true in pacemaker leads where ionic cobalt from the wire catalyzes oxidation. As the polyurethane oxidizes, it loses its physical properties.3
After this discovery, some device companies tried to create biostable polyurethanes, and several approaches have been developed. Some of the formulas include replacing the polyether backbone with a polycarbonate backbone, modifying the end groups of the polymer chain with siloxane, and adding silicone to the backbone. Each approach gives higher resistance to oxidative attack than conventional polyether-based polyurethanes.
The earliest attempt was to substitute a polycarbonate backbone for the traditional polyether backbone. This substitution offers better resistance to direct oxidative attack than the polyether backbone. Although the polycarbonate backbone was an improvement over polyether, it was not completely biostable.4 Implant studies have shown that although the deterioration is slower, it still occurs.
Researchers have also developed polyurethanes that contain silicone at the end of the polymer chain. The silicone end groups are incompatible with the bulk polymer, so they rise to the surface of the device, creating a barrier to oxidative attack. Although the device is made primarily of polyurethane, the surface presented to the body is silicone. Silicone is much more biostable than the underlying polyurethane.5
In another approach, portions of the polyether backbone were replaced with silicone. Samples that substituted up to 80% of the polyether backbone with silicone were evaluated. This approach works quite well to resist oxidation, depending on the level of silicone incorporated into the backbone. Higher levels of silicone reduced the degradation rate significantly.6,7
Whether it is possible to create biostable polyurethanes that can last for several years in the body is still unknown, and companies continue to work toward it. With the diversity of the building blocks the formulator can use, it is likely that the issue of biostability of the polyurethanes will be fully resolved in the future.
Melt-processable, or thermoplastic, polyurethanes are used extensively in medical devices. Thermoplastic polyurethanes are long-chain linear polymers without cross-links. Their linear construction allows the polyurethane to be melted to form parts; the parts then resolidify.
Thermoplastic polyurethanes are made in a batch or continuous process from liquid precursors. In a batch process, the three components—the backbone, diisocyanate, and chain extender—are mixed together, dispensed into trays, and allowed to cure. The cured polyurethane is granulated and then pelletized. Only linear precursors are used to make thermoplastic polyurethanes. In the continuous process, the precursors are metered into an extruder that mixes the ingredients and feeds the mix onto a belt. The strands are fed through a heated tunnel where they are cured. The cured polyurethane strands are granulated and then pelletized. Then, the pellets are used to make films, molded parts, and profile extrusions.
Melt-processable polyurethanes are often used to make catheters, such as over-the-needle IV catheters, central-venous access catheters, and multilumen catheters. In the case of multilumen catheters, very thin walls can be formed between the lumens, which allow the maximum number of lumens while maintaining the minimum OD. Polyurethanes for catheter use must be biocompatible, tough, and strong. They must also have good column strength, which is the property that enables the catheter to be advanced into the body without kinking.
Thermoplastic polyurethanes are also commonly used to injection mold parts. Although this can be a difficult process, their unique properties often make it worthwhile. It is important to keep in mind that some polyurethanes behave like rubber instead of a rigid material. For such thermoplastics, even those with durometers as high as 70 Shore D, the resulting parts have high elongation (160%) and high tensile strength (4200 psi). The parts also have the flexibility of rubber (flexural modulus of 64,000 psi). A typical application is a molded backform that ties together the extension tubes to a multilumen catheter. The photo on page 104 shows an IV catheter. The product includes a molded backform, specialty catheter and needle hubs, oxygen masks, medical tubing, and pliable postsurgical appliances.
Another application that capitalizes on the versatility of thermoplastic polyurethanes is wound dressings. Most wound dressings are composite structures. Polyurethane wound-dressing films are used to make a covering that is impermeable to fluids and bacteria but allows moisture to permeate. The thin outer thermoplastic polyurethane film provides excellent bacterial penetration resistance, yet is permeable to moisture vapor. The absorbent inner layer is produced from open-cell hydrophilic polyurethane foam and absorbs the wound exudate. Low-durometer (70–80 Shore A) polyurethanes, are used in this application because they have 7000-psi tensile strength and 500% elongation in thin sections. They can also be formed into foam and can be made to have different permeability characteristics for different applications. For example, as the population ages, there will be a higher demand for wound dressings that can handle decubitus, or nonhealing, ulcers and skin breakdown.
An example of the versatility of polyurethanes in wound management is the development of asymmetric membranes. These membranes are formed in a phase-inversion process. The top layer is microporous—the pore size is less than 0.7 nm—and the foamlike wound-contact layer has a pore size of 10–100 nm.
Polyurethanes are used in many medical devices, and their usage continues to grow. They are often chosen because they can fulfill product requirements that cannot be met by other biomedical materials. Their biocompatibility and unique chemistry and processing make them ideal for numerous medical applications.
Polyurethanes are used in a myriad of applications, including encapsulants for hollow-fiber devices, dip-molded gloves and balloons, asymmetric membranes, functional coatings, and profile extrusions for catheters. With improvements in biostability continuing and new applications emerging at an increasing rate, polyurethanes are only going to become more common in medical devices.
1. ISO 10993-1:2003, “Biological Evaluation of Medical Devices—Part 1: Evaluation and Testing” (Geneva: International Organization for Standardization, 2003).
2. K Stokes, P Urbanski, and J Upton, “The In Vivo Autooxidation of Polyether Polyurethane by Metal Ions,” Journal of Biomaterials Science, Polymer Ed. 1, no. 3 (1990): 207–230.
3. K Stokes, A Coury, and P Urbanski, “Autooxidative Degradation of Implanted Polyether Polyurethane Devices,” Journal of Biomaterials Applications 1, no. 4 (1987): 411–448.
4. AM Seifalian et al., “Effect of Soft-Segment Chemistry on Polyurethane Biostability during In Vitro Fatigue Loading,” Biomaterials 24, no. 14 (2003): 2549–2557.
5. Z Chen et al., “Interaction of Fibrinogen with Surfaces of End-Group-Modified Polyurethanes: A Surface-Specific Sum-Frequency-Generation Vibrational Spectroscopy Study,” Journal of Biomedical Material Research 2, no. 62 (2002): 254–264.
6. DJ Martin et al., “Polydimethylsiloxane/Polyether-Mixed Macrodiol-Based Polyurethane Elastomers: Biostability,” Biomaterials 20, no. 10 (2000): 1021–1029.
7. A Simmons et al., “Long-Term In Vivo Biostability of Poly(dimethylsiloxane)/Poly(hexamethylene Oxide) Mixed Macrodiol-Based Polyurethane Elastomers,” Biomaterials 24, no. 20 (2004): 4887–4900.
James I. Wright is principal consultant for James I. Wright & Associates (Santa Ana, CA). E-mail him at firstname.lastname@example.org.
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